Biomaterials Science: An Introduction to Materials in Medicine [4 ed.] 012816137X, 9780128161371

The revised edition of the renowned and bestselling title is the most comprehensive single text on all aspects of biomat

6,754 980 44MB

English Pages 1616 [1651] Year 2020

Report DMCA / Copyright

DOWNLOAD FILE

Polecaj historie

Biomaterials Science: An Introduction to Materials in Medicine [4 ed.]
 012816137X, 9780128161371

Table of contents :
Biomaterials Science
Copyright
List of Contributors
Preface
How to Use this Book
1.1.1 - Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor
Biomaterials and Biomaterials Science
Key Definitions
The Expansion of the Biomaterials Field
Examples of Today’s Biomaterials Applications
Heart Valve Prostheses
Total Hip Replacement Prostheses
Dental Implants
Intraocular Lenses
Ventricular Assist Devices
Characteristics of Biomaterials Science
Multidisciplinary
Diverse Materials Are Used
Biomaterials to Devices to Markets and Medicine
Magnitude of the Field
Success and Failure
Subjects Integral to Biomaterials Science
Toxicology
Biocompatibility
Inflammation and Healing
Functional Tissue Structure and Pathobiology
Dependence on Specific Anatomical Sites of Implantation
Mechanical Requirements and Physical Performance Requirements
Industrial Involvement
Risk/Benefit and Corporate Realities
Ethics
Regulation
Biomaterials Literature
Biomaterials Societies
Summary
1.1.2 - A History of Biomaterials
Biomaterials Before World War II
Before Civilization
Dental Implants in Early Civilizations
Sutures Dating Back Thousands of Years
Artificial Hearts and Organ Perfusion
Contact Lenses
Basic Concepts of Biocompatibility
World War II to the Modern Era: The Surgeon/Physician-Hero
Intraocular Lenses
Hip and Knee Prostheses
Dental Implants
The Artificial Kidney
The Artificial Heart
Breast Implants
Vascular Grafts
Stents
Pacemakers
Heart Valves
Pyrolytic Carbon
Drug Delivery and Controlled Release
Designed Biomaterials
Silicones
Polyurethanes
Teflon
Hydrogels
Poly(Ethylene Glycol)
Poly(Lactic-Glycolic Acid)
Hydroxyapatite
Titanium
Bioglass
The Contemporary Era (Modern Biology and Modern Materials)
Conclusions
1.2.1 - Introduction: Properties of Materials—the Palette of the Biomaterials Engineer
1.2.2 - The Nature of Matter and Materials
Introduction
Atoms and Molecules
Molecular Assemblies
Surfaces
Conclusion
1.2.3 - Bulk Properties of Materials
Introduction
Mechanical Variables and Mechanical Properties
Five Types of Mechanical Loading
From External Loads to Internal Loads and Stresses
Linear and Nonlinear Relationship, Elastic and Plastic Behavior
Pseudoelastic, Hyperelastic, and Viscoelastic Materials
Common Mechanical Properties of Isotropic Materials
Elastic Properties
Yield Strength and Ductility
Strength and Failure
Hardness
Resilience
Toughness
Fracture Toughness and Fatigue Strength
Generalized Hooke's Law and Anisotropy of Materials
Loading Modes, Stress States, and Mohr's Circle
Plane-Stress and Plane-Strain Simplification
Trajectories of Tensile and Compressive Stress Lines
Other Bulk Properties
Thermal Properties
Optical Properties
Piezoelectric Properties
Electrochemical Properties
Chapter Questions
Solution
Solution
Solution
Solution
Solution
Solution
Solution
Solution
Solution
1.2.4 - Surface Properties and Surface Characterization of Biomaterials
Introduction
General Surface Considerations and Definitions
What Surface Properties Are We Interested in?
Surface Analysis Techniques: Principles and Methods
Sample Preparation
Surface Analysis General Comments
Contact Angle Methods
Electron Spectroscopy for Chemical Analysis
Secondary Ion Mass Spectrometry
Scanning Electron Microscopy
Infrared Spectroscopy
Scanning Tunneling Microscopy (STM), Atomic Force Microscopy (AFM), and the Scanning Probe Microscopies (SPMs)
Newer Methods
Studies With Surface Methods
Platelet Consumption and Surface Composition
Contact-Angle Correlations
Contamination of Intraocular Lenses
Titanium
SIMS for Adsorbed Protein Identification and Quantification
Poly(Glycolic Acid) Degradation Studied by SIMS
MultiTechnique Characterization of Adsorbed Peptides and Proteins
Conclusions
Chapter Questions
1.2.5 - Role of Water in Biomaterials
Water: The Special Molecule
Melting Point and Boiling Point
Density and Surface Tension
Specific Heat and Latent Heats of Fusion and Evaporation
Water as a Solvent
Water: Structure
Water: Significance for Biomaterials
Hydrophobic Effect, Liposomes, and Micelles
Hydrogels
Protein Adsorption
Life
Chapter Exercises
Suggested External Reading
1.3.1 - The Materials Side of the Biomaterials Relationship
1.3.2 - Polymers: Basic Principles
Introduction
The Polymer Molecule
Molecular Structure of Single Polymer Molecules
Chemical Structure of Single Polymer Molecules
Copolymers
Determination of Chemical Composition
Tacticity
Molecular Mass
The Molecular Mass Distribution and Its Averages
Characterizing the Molecular Mass Distribution
Connecting Physical Behavior With Chemical Characteristics
Physical States of Linear Polymers
The Rubbery State
The Glassy State
The Semicrystalline State
The Physical Behavior of Linear and Amorphous Polymers
The Physical Behavior of Other Physical States
Characterizing a Polymer's Physical State and Behavior
Measuring the Transition Temperatures Between States
Interactions With Water
Measuring the Hydrophilicity of Polymer Materials
Degradation Characteristics
Polymer Synthesis
Polymerization Mechanisms
Using Synthesis Conditions to Build the Desired Polymer
Case Studies
The Present and the Future
Further Reading
1.3.2A - Polyurethanes
Introduction
Anatomy of a Polyurethane Molecule
The Physical Properties of Polyurethanes
Thermosets
Thermoplastic Elastomers
Polyurethane Synthesis
Precursors
Synthesis Reactions
Tailoring Polyurethane Behavior
Concluding Remarks
Chapter Exercises
1.3.2B - Silicones
Chemical Structure and Nomenclature
Preparation
Silicone Polymers
Polymerization and Polycondensation
Physicochemical Properties
Types, Properties, and Preparation of Silicone Materials
Silicone Elastomers
Elastomer Filler
Processing of Silicone Elastomers
Silicone Gels
Silicone Adhesives
Silicone Film-in-Place, Fast-Cure Elastomers
Biocompatibility of Silicones
Biodurability of Silicones
Medical Applications
Siliconization
Extracorporeal Equipment
Medical Inserts and Implants
Catheters, Drains, and Shunts
Aesthetic Implants
Conclusion
Chapter Questions
Chapter Answers
Question 1
Question 2
Question 3
Question 4
Question 5
Question 6
Question 7
1.3.2C. - Fluorinated Biomaterials
Introduction
Distinguishing the Different Fluoropolymers
Polytetrafluoroethylene
Fluorinated Ethylene Propylene
Polyvinylidene Fluoride
Fluoropolymer Melt Processing
Original Gore-Tex and Generic Equivalents
Surfaces Modified by Fluorination Treatments (Grainger and Stewart, 2001)
Biomedical Applications
Fluorinated Material Biological Response
PTFE (Teflon) Mesh and Fabric Vascular Implants
ePTFE and Teflon Soft Tissue Repair Meshes
ePTFE Vascular Implants
Arteriovenous ePTFE Grafts for Dialysis Access
Multilumen Catheters
Guiding Catheters
PTFE Catheter Introducers
Perfluorocarbon Liquids and Emulsions as Oxygen-Carrying Blood Substitutes
Fluorinated Liquids in the Eye as Experimental Vitreous Substitutes
Fluorinated (Meth)Acrylates and (Meth)Acrylated Perfluoroalkyl Silicones as Cross-Linked Polymer Cores for Soft Contact Lenses
Fluorinated Materials as Antifouling Coatings for Intraocular Lenses
PTFE Paste Injectable Bulking Agent
Ligament Replacement
Sutures
Summary
Glossary
References
Chapter Exercises
1.3.2D - The Organic Matrix of Restorative Composites and Adhesives
Introduction—Historical Perspective
The Monomer Matrix—Conventional Systems
Dimethacrylates (Base and Diluent Monomers) Used in Commercial Composites
Adhesive Monomers
The Monomer Matrix—Novel Systems
Lower Stress Resin Systems
Low-Shrinkage Materials
Network Modulation
Fast Polymerizing Monomers
Antimicrobial Resins
Enhanced Chemical Stability
Enhanced Toughness
Hydrophobic Resins
Silane Coupling Agents
Chapter Exercises
1.3.2E - Hydrogels
Introduction
Classification and Basic Structures of Hydrogels
Synthesis of Hydrogels
Swelling Behavior of Hydrogels
Determination of Structural Characteristics
Biomedical Hydrogels
Acrylic Hydrogels
Poly(Vinyl Alcohol) (PVA) Hydrogels
Poly(Ethylene Glycol) (PEG) Hydrogels
Degradable Hydrogels
Star Polymer and Dendrimer Hydrogels
Self-Assembled Hydrogel Structures
“Smart” or “Intelligent,” Stimuli-Responsive Hydrogels and Their Applications
pH-Sensitive Hydrogels
pH-Responsive Complexation Hydrogels
Temperature-Sensitive Hydrogels
Affinity Hydrogels
Biomedical Applications of Hydrogels
Contact Lenses
Blood-Contacting Hydrogels
Drug Delivery From Hydrogels
Targeted Drug Delivery From Hydrogels
Tissue Engineering Scaffolds From Hydrogels
Miscellaneous Biomedical Applications of Hydrogels
1.3.2F - Degradable and Resorbable Polymers
Introduction
Brief History of Degradable Polymers
Definition of Degradation, Erosion, Bulk, and Surface Processes
Degradable Polymer Properties
Polymer Backbone Functionality
Polyanhydrides
Poly(Ortho Esters)
Polyesters and Polycarbonates
Polymer Architecture
Polymerization Routes
Molecular Weight
Morphology
Relative Hydrophobicity versus Hydrophilicity
Degradation Routes and Kinetics
Hydrolytic Degradation
Surface Erosion
Bulk Degradation
Photodegradation
Enzymatic Degradation
Polymer Design and Processing
Lifetime—How Long Does the Biomaterial Need to Function?
Location—Where Will the Biomaterial Perform Its Task?
Mechanical Properties—What Mechanical Properties Are Required for the Task?
Delivery—How Will the Biomaterial Reach the Required Site?
Composites—When Should a Composite Be Used and How Will Additives Affect Degradation?
Shape—How Will the Material Be Shaped and How Does Shape Affect Degradation Kinetics?
Sterilization—Will Degradation Properties Be the Same After Sterilization?
Performance Metrics
Worked Examples
Question
Solution
Question
Solution
Question
Solution
Case Studies on Degradable Polymers Used in Medicine
Chapter Exercises
1.3.2G - Applications of “Smart Polymers” as Biomaterials
Introduction
Smart Polymers in Solution
Smart Polymer–Protein Bioconjugates
Site-Specific Smart Polymer Bioconjugates
Smart Polymers on Surfaces
Smart Polymer Hydrogels
Stimuli-Responsive Polymer Micelles and Carriers
Conclusions
References
1.3.3 - Metals: Basic Principles
Introduction
Medical Devices and Metals in the Body
The Major Alloy Systems (Ti, NiTi, CoCrMo, SS, Pt, Au, Mg, Ag)
Metal Processing
Processing–Structure–Properties–Performance Paradigm
Structure of Metals and Alloys
Electronic and Atomic Structure: Crystal Structures
Alloying, Microstructure, and Phase Diagrams
Defects in Crystals
Point Defects
Line Defects
Area Defects
Volume Defects
Bulk Mechanical Properties of Metallic Biomaterials
Elastic and Plastic Deformation of Metals
Strength of Metals and Strengthening Mechanisms
Strengthening Mechanisms: Alloying
Strengthening Mechanisms: Cold Working
Strengthening Mechanisms: Grain Size
Strengthening Mechanisms: Precipitation Strengthening
Fracture of Metals
Fatigue of Metals
Surfaces of Metals: Oxide Films and Passivity
High-Field, Low-Temperature Oxide Film Growth
Introduction to Metallic Corrosion
Electrochemical Reactions (Oxidation and Reduction) and the Nernst Equation
The Principal Reduction Reaction in Biomaterials (Oxygen Reduction)
Polarizable and Nonpolarizable Electrodes
Pourbaix Diagrams (Electrode Potentials vs. pH)
Electrochemical Currents (Evans Diagrams)
Electrochemical Impedance Spectroscopy (an Introduction)
Resistive (Faradaic) and Capacitive (Non-Faradaic) Behavior
Basic Impedance Concepts
Semiconducting Oxide Impedance (Mott–Schottky Analysis)
References
Questions
1.3.3A - Titanium Alloys, Including Nitinol
Introduction
Biocompatibility
Biocompatible Titanium Alloys
Recent Efforts in Fabrication Processes
Mechanical Properties of Titanium Alloys
Elastic Modulus
Wear Resistance
Fatigue Behavior
Effects of Interstitial Atoms on Mechanical Properties
Surface Modification of Titanium Alloys
Recent Efforts in the Anodization Process
Effects of Anodization on Corrosion and Surface Mechanical Properties
Coloring Methods for Titanium Alloys
Conclusions
Chapter Exercise 1
Chapter Exercise 2
Chapter Exercise 3
1.3.3B - Stainless Steels
Overview
History
Composition and Types
Structure
Structure, Composition, and Processing Effects on Mechanical Properties
Corrosion
Summary
1.3.3C - CoCr Alloys
Introduction
Microstructure, Mechanical Properties, and Manufacturing of CoCr Alloys
3D Printing of CoCr Alloys
Bio-Tribocorrosion of CoCr Alloys
Application of CoCr Alloys in Biomedical Devices
Properties Leading to Biocompatibility of CoCr Alloys and Their Applications
Corrosion Resistance
CoCr Alloys in Biological Environments
Clinical Concerns Related to Metal Ion Release From CoCr Alloys
Conclusions
Questions
1.3.3D - Biodegradable Metals
Introduction
General Considerations of Corrosion Design of Biodegradable Metals
General Ideas on the Influence of Alloying Elements, Corrosion Behavior, and the Biocompatibility of Zn and Mg
Iron-Based Biodegradable Metals
Introduction to Fe-Based Implants
Modifications to Accelerate the Corrosion Rate of Fe-Based Biodegradable Metals
The Proposed Degradation Process of Fe-Based Biodegradable Metals
Biocompatibility Evaluations
Current Perspective on Fe-Based Degradable Implants
Zinc-Based Biodegradable Metals
Introduction to Zn-Based Implants
Zn-Based Materials Under Investigation
The Proposed Degradation Process of Zn-Based Biodegradable Metals
Biocompatibility of Dissolved Zn Corrosion Products
Current Perspective on Zn-Based Degradable Implants
Magnesium-Based Biodegradable Metals
Introduction to Mg-Based Implants
Impact of Alloying Elements on Mg Processing and Microstructure
Current Models of the Corrosion Process In Vitro
Situation In Vivo: Tissue Perfusion, pH, and the Issue of Gas Formation
Methods to Measure Mg-Based Implant Corrosion In Vitro and In Vivo
Preclinical and Clinical Observations for Mg-Based Biodegradable Metals
Orthopedic Devices Based on MgYREZr Alloy (Magnezix)
Orthopedic Devices Based on MgCaZn Alloy (Resomet)
Orthopedic Devices Based on Pure Mg
Current Perspective on Mg-Based Degradable Implants
Summary
Acknowledgment
References
1.3.4 - Ceramics, Glasses, and Glass-Ceramics: Basic Principles
Introduction
Nearly-Bioinert Ceramics
Alumina and Zirconia Ceramics
Bioactive Ceramics and Glasses
Bioactive Ceramics
Porous Calcium Phosphate Ceramics
Calcium Phosphate Cements
Bioglass and Bioactive Glass
Bioglass Granules
Bioactive Glass Composites and Putties for Bone Repair
Porous Bioactive Glasses
Wound Healing
Bioactive Glass in Toothpaste
Glasses for Cancer Therapy
Glass-Ceramics
Summary
Chapter Questions
Questions with answers
1.3.4A - Natural and Synthetic Hydroxyapatites
Introduction
Synthesis of Hydroxyapatite Ceramics
Characterization of Hydroxyapatite Ceramics
Physicochemical Characterization
In Vitro and In Vivo Characterization
Clinical Use of Hydroxyapatite Ceramics
1.3.4B - Structural Ceramic Oxides
Introduction
Structural Ceramic Oxides
Aluminum Oxide (Alumina)
Zirconia
Yttria and Magnesia-Stabilized Zirconias
Zirconia-Toughened Alumina
A History of These Structural Materials in Medical Devices
Properties in General
Questions
1.3.5 - Carbon Biomaterials
Introduction
Carbon Biomaterials
Diamond and Diamond-Like Carbon
Diamond
Diamond-Like Carbon
Pyrolytic Carbon
Hexagonally Bonded Carbon
Graphite
Fullerenes
Carbon Nanotubes
Graphene-Based Materials
Other Hexagonally Bonded Carbons
.Graphene quantum dots (GQD) are 0D materials (2–20nm) with a crystalline form of carbon containing sp2 hybridized atoms. These ...
.Carbon fibers (CF) are a 3D material (diameter: 5–10μm) with a crystalline form of carbon containing sp2 hybridized atoms. Thes...
.Carbon nanofibers (CNF) are noncontinuous 1D materials with a crystalline form of carbon containing sp2 hybridized atoms. CNF c...
.Graphene nanoribbons (GNR) are 1D materials with a crystalline form of carbon containing sp2 hybridized atoms. GNR are commonly...
Other Carbon Biomaterials
Carbon Dots
Glassy Carbon
Activated Charcoal
Biomedical Applications of Carbon Biomaterials
Drug Delivery
Phototherapy and Imaging
Biosensors
Antimicrobial Therapy
Cardiovascular Applications
Long-Term Implants
Mechanical Heart Valves
Vascular Stents
Ventricular Assist Devices
Tissue-Engineering Approaches
Orthopedic Applications
Long-Term Implants
Tissue-Engineering Approaches
Dental Applications
Neurological Applications
Ophthalmologic Applications
Contact Lenses
Catheters
Guidewires
Other Biomedical Applications
Safety of Carbon Biomaterials: Short Considerations
Summary
Chapter Questions and Answers
1.3.6 - Natural Materials
Introduction to Natural Materials
Natural Based-Biomaterials Exploring Structural Molecules
Extracellular Matrix-Based Biomaterials
Proteins
Glycosaminoglycans
Blood Derivatives as a Source of Bioinstructive Materials
Multifunctional Biomaterials Based on DNA
Dynamic Hydrogels Exploring Supramolecular Chemistry
Reversible Hydrogels Based on Supramolecular Cross-Linking of Polymeric Precursors
Hydrogels Based on Natural Supramolecular Self-Assembly
Soft Nanocomposite Smart Materials
Stimuli-Responsive Soft Nanocomposites
Future Perspectives
Questions
1.3.6A - Processed Tissues
Introduction
Cryopreservation and Vitrification
Tissue Cross-Linking
Decellularization
Decellularization Methods
Quality of Decellularization
Post-decellularization Processing and Modifications
Milling for ECM Powder and Partial Enzymatic Digestion for Hydrogel Formation
Cross-Linking
Applications of Decellularized ECM
Scaffold-Based Therapies
Whole Organ Recellularization
Powder and Injectable Decellularized ECM Therapies
Tissue-Specific In Vitro Models of the Native Microenvironment
Current Challenges and Future Directions for Decellularized Tissues
Conclusion
Acknowledgments
Questions
1.3.6B - Use of Extracellular Matrix Proteins and Natural Materials in Bioengineering
Introduction
Collagens
Elastin, Elastic Fibers, and Elastin-Like Peptides
Proteoglycans and Glycosaminoglycans
Alginates
Chitosan
Fibrin
Manufacturing Approaches Utilizing Natural Materials
Human Recombinant ECM Protein Production
Purification of Recombinant ECM Proteins
3D Bioprinting
Electrospinning of ECM Proteins and Natural Materials
Summary
1.3.7 - Composites
Introduction
Matrix and Reinforcement in Composites
Matrix Materials
Reinforcements
Nonporous and Porous Composites
Properties of Composites
Major Influencing Factors
Geometry and Size of the Dispersed Phase and Its Distribution in Composite
Fiber Arrangement
Interfaces in Composites
Mechanical Properties of Composites
Tensile Properties of Fibrous Composites
Compressive Properties of Fibrous (CF/PEEK) Composites: A New Perspective
Rosen’s Microbuckling Model and the Contradictions
A First-Principles-Based Compressive Microbuckling Model
Selected Results
Medical Applications of Composites
Biomedical Composites in Orthopedic Applications
Biomedical Composites in Dental Applications
Biomedical Composites for Tissue Engineering
Chapter Questions
1.3.8A - Microparticles
Introduction
Why Size Matters
Materials for the Synthesis of Microparticles
Natural Polymers
Synthetic Polymers
Nonpolymeric Materials
Microparticle Preparation
Characterization of Microparticles
Drug Release Mechanisms
Biomedical Applications of Microparticles
Drug Delivery
Radiotherapy
Other Applications
Concluding Remarks
Chapter Exercises
1.3.8B - Nanoparticles
Chapter Objectives
Introduction
Categories of NPs
Polymeric NPs
Lipid-Based NPs
Inorganic NPs
Bio-Inspired NPs
Hybrid NPs
Characterization of NPs
Size
Surface Charge
Morphology
Biocompatibility
In Vitro Toxicity
Hemocompatibility
In Vivo Toxicity
Drug Delivery Applications of NPs
Drug Loading
Covalent Bonding (Prodrug)
Noncovalent Encapsulation
Systemic Barriers Against Drug Delivery
Approaches to Overcome Systemic Barriers
Long-Circulating NPs
Targeted Drug Delivery
Tumor Penetration
Stimuli-Responsive Drug Delivery
Clinical Development
Nucleic Acid Delivery Applications of NPs
Intracellular Barriers Against Nucleic Acid Delivery
Strategies to Overcome Intracellular Barriers
Nucleic Acid Condensation and Cellular Internalization
Endosomal Escape
Stimuli-Responsive NPs for Intracellular Gene Release
Nuclear Transport
Clinical Development
Diagnostic/Theranostic Applications of NPs
In Vitro Diagnosis
In Vivo Imaging
Theranostics
Imaging-Guided Surgery
Conclusion
Chapter Assessment Questions
1.4.1 - Introduction to Materials Processing for Biomaterials
1.4.2 - Physicochemical Surface Modification of Materials Used in Medicine
Introduction
General Principles
Thin Surface Modifications
Delamination Resistance
Surface Rearrangement
Surface Analysis
Manufacturability and Commercializability
Methods for Modifying the Surfaces of Materials
Chemical Reaction
Surface Grafting: Radiation Grafting, Photografting, and Newer Methods
RFGD Plasma Deposition and Other Plasma Gas Processes
The Nature of the Plasma Environment
The Apparatus to Generate Plasmas for Surface Modification
RFGD Plasmas for the Immobilization of Molecules
High-Temperature and High-Energy Plasma Treatments
Specific Chemical Reactions for Forming Surface Grafts
Silanization
Ion Beam Implantation
Langmuir–Blodgett Deposition
Self-Assembled Monolayers
Layer-By-Layer Deposition and Multilayer Polyelectrolyte Deposition
Surface-Modifying Additives
Conversion Coatings
Parylene Coating
Laser Methods
Patterning
References
Conclusions
1.4.3A - Nonfouling Surfaces
Introduction
Background and Mechanism
Nonfouling Materials and Methods
Conclusions and Perspectives
1.4.3B - Nonthrombogenic Treatments and Strategies
Introduction
Historical
Criteria for Nonthrombogenicity
Inert Materials
Hydrogels
Polyethylene Glycol (PEG) Immobilization
Albumin Coating and Alkylation
Zwitterionic Group/Phospholipid-Mimicking Surfaces
Surface-Modifying Additives (SMAs)
Fluorination
Heparin-Like Materials
Self-Assembled Surface Layers
Active Materials
Heparinization
Ionically Bound Heparin and Controlled-Release Systems
Covalently Bound Heparin
Thrombin Inhibition Without Heparin
Immobilization of Antiplatelet Agents
Immobilization of Fibrinolytic Agents
Use of Endothelial Cells and RGD Peptides
Strategies to Lower the Thrombogenicity of Metals
Summary
1.4.4 - Surface-Immobilized Biomolecules
Introduction
Patterned Surface Compositions
Immobilized Biomolecules and Their Uses
Immobilized Cell Ligands and Cells
Immobilization Methods
Conclusions
References
1.4.5 - Surface Patterning
Introduction
Common Concerns In Biomolecular Surface Patterning
Resolution
Throughput
Contrast
Bioactivity
Shelf-Life and Durability
Patterning Techniques
Direct-Write Patterning
Writing With A Stylus
Printing With Inkjets, Quills, and Pins
Dip-Pen Nanolithography
Nanoshaving and Nanografting
Writing With Beams
Direct-Write Photolithography
Electron Beam Lithography
Focused Ion Beam Lithography
Writing With Fields
Electric Field
Magnetic Field
Patterning With Masks
Photolithography With Masks
Deposition/Etching With Masks
Patterning With Masters
Imprinting With a Master
Printing With a Stamp
Microcontact Printing: Use of Protruding Features of a Stamp
Microfluidic Patterning: Use of Void Features of a Stamp
Patterning by Self-Assembly of Polymers and Colloids
Block Copolymer Self-Assembly
Nanosphere Lithography
Magnetic Self-Assembly
Dynamic Patterning
Three-Dimensional Printing
Conclusions
1.4.6 - Medical Fibers and Biotextiles
Introduction
Fiber-Forming Polymers
Characteristics of Fiber-Forming Polymers
Natural and Synthetic Polymers for Biotextile Production
Medical Fibers and Production Methods
Introduction to Textile Fibers
Melt Extrusion
Wet/Gel Spinning
Electrospinning
Electrospinning Process and Spinning Parameter Optimization
Materials Selection for Electrospinning
Coelectrospinning
Centrifugal Electrospinning
Hydrogel Fiber Spinning
Surface Functionalization
Textile Structures
Woven Textiles
Knitted Textiles
Braided Textiles
Nonwoven Textiles
Finishing and Surface Coating
Applications of Medical Fibers and Biotextiles
Biotextiles of General Surgery
Meshes and Sutures: Design and Materials
Barbed and Drug-Eluting Sutures
Cardiovascular Applications of Biotextiles
Design Criteria for Vascular Prostheses
Woven Versus Knitted Structure
Examples of Cardiovascular Biotextiles
Endovascular Stent Grafts
Knitted Textile Structures as Sewing Rings
Orthopedic Applications of Biotextiles
Ligament and Tendon Replacement With Woven and Braided Biotextiles
Fiber Reinforcement in Bone Graft Cement
Biotextiles as Wound Dressings and Skin Grafts
Wound Dressings and Hemostats
Skin Grafting for Burn Injuries
Applications of Electrospun Fibers
Wound Dressing
Musculoskeletal Tissue Engineering
Neural Tissue Engineering
Nanofibers for Cardiovascular Repair
Nanofibers for Local Drug Delivery
Future Directions
Chapter Study Questions
1.4.7 - Textured and Porous Biomaterials
Introduction
Importance of Texture and Porosity in Facilitating Biomaterial Integration
Textured Devices Promote Healing and Restore Organ Function
Porosity to Promote Tissue Ingrowth
Biomaterials for Tissue Engineering
Fabrication Methods for Biomimetic Nanoscale Texture
Electrospinning
Self-Assembly of Nanoscale Features
Thermally Induced Phase Separation
Grooves and Micropatterns
Fabrication Methods for Micro- and Macroscale Architectural Features
Interconnected Spherical Macropores by Porogen Methods
Nonspherical Architectural Patterning
Combining Multiple Fabrication Methods
Macroporous, Nanofibrous Tissue-Engineering Scaffolds
Multiphasic Scaffolds
3D Printed Scaffolds
Injectable Tissue-Engineering Scaffolds
Surface Modification of Biomaterial Constructs
Summary and Future Perspectives
Chapter Exercises
1.4.8 - Biomedical Applications of Additive Manufacturing
Introduction
3D Printing Modalities
Vat Photolithography
Material Jetting
Material Extrusion
Powder Bed Fusion
Binder Jetting
Sheet Lamination
Directed Deposition
Bioprinting
Bioprinting Approaches
Bioink Design Parameters
Biofabrication Window
Biomaterials for Bioprinting
Medical Applications of 3D Printing
Surgical Planning and Medical Training
Fabrication of Complex Implants
Personalized Drug Delivery Systems
Summary
Chapter Review Questions
2.1.1 - Introduction to Biology and Medicine—Key Concepts in the Use of Biomaterials in Surgery and Medical Devices
2.1.2 - Adsorbed Proteins on Biomaterials
Introduction
Examples of the Effects of Adhesion Proteins on Cellular Interactions With Materials
The Effects of Preadsorption With Purified Adhesion Proteins
Depletion Studies
Inhibition of Receptor Activity With Antibodies
The Adsorption Behavior of Proteins at Solid–Liquid Interfaces
Adsorption Transforms the Interface
Rapid Adsorption Kinetics and Irreversibility
The Monolayer Model
Competitive Adsorption of Proteins to Surfaces From Protein Mixtures
Molecular Spreading Events: Conformational and Biological Changes in Adsorbed Proteins
Physicochemical Studies of Conformational Changes
Changes in Biological Properties of Adsorbed Proteins
The Importance of Adsorbed Proteins in Biomaterials
Surface Chemistries Highly Resistant to Protein Adsorption
Concluding Remarks
Chapter Solutions to Problems
Chapter Solutions to Problems
Protein Monolayer Calculation
2.1.3 - Cells and Surfaces in Vitro
Introduction
A Basic Overview of Cell Culture
Primary Culture
Cell Lines
Characteristics of Cultured Cells
Understanding Cell–Substrate Interactions
Surfaces for Cell Culture
Process of Cell Attachment in Vitro
Commercial and Experimental Modifications of Culture Surfaces
Dynamic Control of Cell Culture Surfaces
Investigating Cell–Substrate Interactions
Cell Response to Substrate Chemistry
Micrometer-Scale Chemical Patterns
Nonfouling Surfaces in Cell Culture
Chemical Patterning for the Coculture of Cells
High-Throughput Screening
Nanometer-Scale Chemical Patterning
Cell Response to Substrate Topography
Micrometer-Scale Topography
Nanometer-Scale Topography
High-Throughput Screening of Surface Topography
Cell Response to Substrate Elasticity
Cell Response to Mechanical Deformation (Strain)
Comparison and Evaluation of Substrate Cues
Chemistry and Topography
Chemistry and Strain
Topography and Strain
Organ-on-a-Chip 3D culture
Summary
2.1.4 - Functional Tissue Architecture, Homeostasis, and Responses to Injury
Tissue Constituents, Organization, and Integration
The Essential Role of Cells
Parenchyma and Stroma
Vascular Supply: Tissue Perfusion
Extracellular Matrix (See also Chapter 2.1.5)
Organ Structure
Cell and Tissue Differentiation, Phenotype, and Maintenance
Structure–Function Correlation
Stem Cells
Cellular Differentiation and Gene Expression
Tissue Homeostasis
Cell Turnover
Matrix Remodeling
Cell and Tissue Injury, Adaptation, and Other Responses (Fig. 2.1.4.12)
Cell Regeneration and Proliferation
Reversible versus Irreversible Injury
Adaptation
Hyperplasia Can Be Physiologic or Pathologic
Atrophy, Proteasomes, and Autophagy
Metaplasia
Neoplasia
Causes of Cell Injury
Hypoxia and Ischemia
Toxic Injury and Trauma
Infection and Inflammation
Pathogenesis of Cell Injury
Ischemia-Reperfusion Injury
Cell Death
Necrosis
Apoptosis
Response to Tissue Injury and Biomaterials
Inflammation and Innate Immunity
Macrophage Recruitment and Polarization
Regeneration Versus Fibrosis (Scar)
Growth Factors
Vascular Response
Wound Healing in the Presence of Biomaterials
Complications and Defective Wound Healing
2.1.5 - The Extracellular Matrix and Cell–Biomaterial Interactions
Introduction
Extracellular Matrices
Properties of the Extracellular Matrix
Collagens and Elastin
Fibronectin
Laminins
Proteoglycans, Glycosaminoglycans, and Hyaluronic Acid
Growth Factor Sequestering Proteins and Motifs
ECM Remodeling and Proteolysis
Integrins and Adhesion Receptors
Cell–Biomaterial Interactions
Cell Interactions With Adsorbed Proteins on Biomaterials
Engineered Receptor-Targeting Peptide Sequences for Cell Adhesion
Engineered MMP-Sensitive Peptide Sequences for ECM Remodeling and Proteolysis
Engineered Peptide Fibers That Mimic the ECM Structure
Summary
Chapter Exercises
2.1.6 - Effects of Mechanical Forces on Cells and Tissues
Introduction
Molecular Mechanisms of Cellular Mechanotransduction
Focal Adhesion and Mechanosensing at the ECM–Biomaterial Interface
Cytoskeletal Mechanotransduction
Nuclear Mechanotransduction
Techniques for Studying Mechanical Interactions of Cells
Shear Stress
Mechanical Stretch
Substrate Stiffness
Micro- and Nanopatterning
Mechanical Forces in the Vascular System
Effect of Shear Stress on Blood Vessels
Effect of Cyclic Strain on Blood Vessels
Bone and Cartilage
Summary
2.2.1 - Introduction to Biological Responses to Materials
2.2.2 - Inflammation, Wound Healing, the Foreign-Body Response, and Alternative Tissue Responses
Biocompatibility and Implantation
Sequence of the Host Response Following Implantation of Medical Devices
Wound Healing
Host Response to Implanted Biomaterials
Tissue Remodeling and Biomaterial Integration—Alternative Tissue Responses
Cellular and Molecular Mediators of Constructive Remodeling and Tissue Restoration
Strategies to Control Host Responses
2.2.3 - Innate and Adaptive Immunity: The Immune Response to Foreign Materials
Overview
Innate Immunity
First Barriers Against Danger
Complement System
Pattern Recognition by the Innate Immune System
Cells of the Innate Immune System
Antigen Uptake, Processing, and Presentation
Costimulatory Molecules
Chemokines and Cytokines
Adaptive Immunity
Components of Adaptive Immunity
Humoral Immunity
Cell-Mediated Immunity
Cytotoxic T Cells
Helper T Cells
Recognition in Adaptive Immunity
B Cell and Antibody Recognition
T Cell Recognition
Effector Pathways in Adaptive Immunity
Immunological Memory
Overview of the Immune Response to Pathogens
Overview of Immune Regulation and Tolerance
Intersection of Biomaterials and Immunology
Chapter Exercises
Innate Immunity
Adaptive Immunity
2.2.4 - The Complement System
Introduction
Classical Pathway
Lectin Pathway
Alternative Pathway
Membrane Attack Complex
Control Mechanisms
Complement Receptors
Measurement of Complement Activation
Complement–Coagulation System Interactions
Clinical Correlates
Summary and Future Directions
Chapter Questions
2.2.5 - Systemic and Immune Toxicity of Implanted Materials
Basic Principles of Systemic Distribution and Toxicity of Biomaterial Constituents
Metals and Metal Alloy Toxicity
Hypersensitivity and Immunotoxicity
Organ Localization of Inflammatory and Immune Responses to Device Materials
Summary and Conclusions
Chapter Exercises
2.2.6 - Blood Coagulation and Blood–Material Interactions
Introduction
Platelet Adhesion and the Blood Coagulation Cascade—An Overview
Cellular Composition of Blood
Erythrocytes (Red Cells)
Leukocytes (White Cells)
Platelets
Platelet Adhesion
Platelet Aggregation
Platelet Release Reaction
Platelet Coagulant Activity
Platelet Consumption
Coagulation
Mechanisms of Coagulation
Control Mechanisms
Fibrinolysis
Complement
Blood–Material Interactions
Overview
Platelet–Material Interactions
Contact Activation of the Blood Coagulation Cascade
Approaches to Improve the Blood Compatibility of Artificial Materials
Conclusions
Chapter Exercise Questions
Question 1
Question 2
Question 3
2.2.7 - Tumorigenesis and Biomaterials
General Concepts
Association of Implants With Human and Animal Tumors
Pathobiology of Foreign Body Tumorigenesis
Stem Cell Therapies and Tumorigenesis
Conclusions
References
2.2.8 - Biofilms, Biomaterials, and Device-Related Infections
Introduction
Bacterial Biofilms
What Are Biofilms and Why Are They Problematic?
The Biofilm Microenvironment
Antibiotic and Antimicrobial Tolerance of Bacteria in Biofilms
Biofilms and the Immune Response
Bacterial Adhesion
The Process of Bacterial Adhesion to Surfaces
DLVO Theory
Thermodynamic Model
Influence of Material Properties on Bacterial Adhesion
Surface Free Energy (Wettability)
Roughness
Environment Factors Influence Bacterial Adhesion
Device-Related Infection
Major Medical Devices, Materials, and Pathogens
Evidence for Biofilms on Devices
Control of Biofilm Formation
Antimicrobial Approaches: Biomaterials With Antimicrobial Properties
Biomaterials Releasing Bioactive Molecules
Antibiotics
Silver
Low-Dose Nitric Oxide
Intrinsically Bioactive Biomaterials: Cationic Materials
Natural Cationic Polymers
.Chitosan is a polysaccharide composed of randomly distributed N-acetylglucosamine and d-glucosamine having low toxicity toward ...
.Antimicrobial peptides (AMPs) are produced as part of the first line of defense in innate immunity system. Typical AMPs are sma...
Synthetic Cationic Polymers
Antifouling Approaches: Biomaterials That Repel Microbes
Hydrophilic Materials Based on Polyethylene Glycol
Superhydrophobic Materials
Materials With Nano/Microscale Surface Texture
Biomaterials Affecting Biofilm Architecture
Biomaterials Modified With QS-Quenching Enzymes
Biofilm Matrix-Degrading Enzymes
Methods for Testing Antibacterial and Antifouling Properties of Biomaterials
Conclusions
Chapter Questions
2.3.1 - How Well Will It Work Introduction to Testing Biomaterials
2.3.2 - The Concept and Assessment of Biocompatibility
Biocompatibility Today
Toxicology
The Products of Extrinsic Organisms Colonizing the Biomaterial
Mechanical Effects
Cell–Biomaterial Interactions
Summary of Ideas to This Point
New Developments Are Changing the Paradigm of Biocompatibility
Clinical Significance of Biocompatibility
Conclusions
2.3.3 - In Vitro Assessment of Cell and Tissue Compatibility
Introduction
Background Concepts
Use of Medical Device/Biomaterial Chemical Composition and Their Extracts for Toxicological Risk Assessment and In Vitro Testing...
In Vitro Assays to Assess Cell and Tissue Compatibility in Medical Device/Biomaterial Evaluation for Regulatory Purposes
In Vitro Tests for Genotoxicity, Carcinogenicity, and Reproductive Toxicity: ISO 10993-3
In Vitro Tests for Interactions with Blood: ISO 10993-4
In Vitro Tests for Cytotoxicity: ISO 10993-5
Application-Specific In Vitro Assays Considered in Proof-of-Concept Testing
Future Challenges in In Vitro Assessment of Cell and Tissue Compatibility
Summary Remarks
Chapter Questions
2.3.4 - In Vivo Assessment of Tissue Compatibility
Introduction
Selection of in Vivo Tests According to Intended Use
Biomaterial and Device Perspectives in In Vivo Testing
Specific Biological Properties Assessed by In Vivo Tests
Sensitization, Irritation, and Intracutaneous (Intradermal) Reactivity
Systemic Toxicity: Acute, Subacute, and Subchronic Toxicity
Genotoxicity
Implantation
Hemocompatibility
Chronic Toxicity
Carcinogenicity
Reproductive and Developmental Toxicity
Biodegradation
Immune Responses
Selection of Animal Models for In Vivo Tests
Future Perspectives on In Vivo Medical Device Testing
2.3.5 - Evaluation of Blood–Materials Interactions
Introduction
Background and Principles of Blood–Materials Interactions Assessment
What Is Blood Compatibility?
Why Measure Blood Compatibility?
What Is Thrombogenicity?
Key Considerations for BMI Assessment
Blood: A Fragile Fluid That Is Readily Compromised
Flow: Blood Interactions Dictated by Shear and Mass Transport
Surfaces: Actively Studied, but Least Well Defined, of the BMI Variables
Blood Interaction Times With Materials and Devices
Evaluation of BMI
In Vitro Tests
In Vivo Tests of BMI
In Vivo Evaluation of Devices
Contemporary Concepts in BMI Evaluation
Examples of BMI Evaluation
What Materials Are Blood Compatible?
Conclusions
References
2.3.6 - Animal Surgery and Care of Animals
Introduction
Ethical and Regulatory Overview
Governmental Regulations
United States Department of Agriculture
Public Health Service
Food and Drug Administration
Institutional Responsibilities
Institutional Animal Care and Use Committee
Attending Veterinarian
Principal Investigator
Surgical Facility Design
Model Selection
Cardiovascular Devices
Heart Valve Replacement
Ventricular Assist Devices
Orthopedic Devices
Bone Defect Models
Vascular
Ophthalmology
Skin
Animal Management and Care of Animals
Rodent
Animal Selection and Preoperative Preparation
General Anesthesia
Analgesia
Ruminants (Sheep, Goats, Calves)
Animal Selection and Preoperative Preparation
Brief Restraint
General Anesthesia
Analgesia
Rabbit
Animal Selection and Preoperative Preparation
Brief Procedures
General Anesthesia
Analgesia
Swine
Animal Selection and Preoperative Preparation
Brief Restraint
General Anesthesia
Analgesia
Chapter Study Questions
2.4.1 - Introduction: The Body Fights Back–Degradation of Materials in the Biological Environment
2.4.2- Chemical and Biochemical Degradation of Polymers Intended to Be Biostable
Introduction
Polymer Degradation Processes
Preimplant Degradation
Postimplant Degradation Forces
Hydrolytic Biodegradation
Structures of Hydrolyzable Polymers
Host-Induced Hydrolytic Processes
Hydrolysis: Preclinical and Clinical Experience
Polymers Containing Hydrolyzable Pendant Groups
Oxidative Biodegradation
Oxidation Reaction Mechanisms and Polymer Structures
Direct Oxidation by Host
Stress Cracking
Device- or Environment-Mediated Oxidation
Chemical Structure Strategies to Combat Oxidation
Oxidative Degradation Induced by External Environment
Emerging Long-Term Elastomer Applications
Polyurethanes
Hydrocarbon Elastomers
Conclusions
Chapter Questions
2.4.3 - Metallic Degradation and the Biological Environment
Introduction
The Severe Biological Environment (Fatigue, Tribology, Corrosion, and Biology)
Basic Corrosion of Passive Oxide-Covered Alloys
Tribological Aspects of Metal-Hard Contact Degradation
Metal-on-Metal (Hard) Surface Mechanics
Clinically Observed Mechanically Assisted Crevice Corrosion (Fretting Crevice Corrosion) In Vivo
Mechanically Assisted Corrosion Basics for CoCrMo and Ti–6Al–4V Alloys
Tribocorrosion Layer and Surface Damage on Metallic Biomaterials Surfaces
Biology and Corrosion: Additional Insights
Reduction Reactions Affect Cells
Reactive Oxygen Species May Enhance Corrosion Reactions
Summary
Acknowledgments
References
2.4.4 - Degradative Effects of the Biological Environment on Ceramic Biomaterials
Introduction
Reactivity of Bioceramics
Factors Influencing the Degradation of Bioceramics
Reactivity and Degradation of Natural Apatites
Evolution in the Use of Bioceramics for Bone Repair
Bioceramic Interactions With the Biological Environment
Inert Ceramics: First-Generation Bioceramics
Resorbable and Bioactive Ceramics: Second-Generation Bioceramics
Third-Generation Ceramics
Summary and Future Perspectives
2.4.5 - Pathological Calcification of Biomaterials
The Spectrum of Pathologic Biomaterial and Medical Device Calcification
Bioprosthetic Heart Valves
Transcatheter (or Percutaneous) Cardiac Valve Replacements
Polymeric Heart Valves and Blood Pump Bladders
Breast Implants
Intrauterine Contraceptive Devices
Urinary Stents and Prostheses
Intraocular and Soft Contact Lenses and Scleral Buckles
Assessment of Biomaterial Calcification
Morphologic Evaluation
Chemical Assessment
Mechanisms of Biomaterial Calcification
Regulation of Pathologic Calcification
Role of Biological Factors
Role of Biomaterial Factors
Role of Biomechanical Factors
Experimental Models for Biomaterial Calcification
Role of Cells
Role of Collagen and Elastin
Role of Glutaraldehyde
Role of Immunologic Factors
Prevention of Calcification
Inhibitors of Hydroxyapatite Formation
Bisphosphonates
Trivalent Metal Ions
Calcium Diffusion Inhibitor
Removal/Modification of Calcifiable Material
Surfactants
Alcohol Treatments
Glutaraldehyde Neutralization
Decellularization
Modification of Glutaraldehyde Fixation and Other Tissue Fixatives
Alternative Materials
Design Considerations and Selection of Materials to Avoid Calcification
Conclusions
References
2.5.1 - Introduction to Applications of Biomaterials
2.5.2A - Cardiovascular Medical Devices: Heart Valves, Pacemakers and Defibrillators, Mechanical Circulatory Support, and Other Intracardiac Devices
Introduction
Heart Valve Function and Valvular Heart Disease
Surgical Bioprosthetic and Mechanical Heart Valves
Percutaneous Transcatheter Valves and Other Devices
Cardiac Arrhythmias
Cardiac Pacemakers
Implantable Cardioverter-Defibrillators
Complications of Pacemakers and ICDs
Congestive Heart Failure
Cardiopulmonary Bypass
Percutaneous Mechanical Circulatory Support Devices
Durable Ventricular Assist Devices and Total Artificial Hearts
Atrial Septal Defects and Other Intracardiac Defects
Closure Devices
Atrial Fibrillation
Left Atrial Appendage Occlusion Devices
2.5.2B - Cardiovascular Medical Devices: Stents, Grafts, Stent-Grafts and Other Endovascular Devices
Key Concepts in Vascular Structure and Function
Architecture of the Circulation
Vascular Pathology
Vascular Injury and Healing
Thrombosis
Atherosclerosis
Aneurysms and Dissections
Vascular Devices and Biomaterials
Angioplasty and Endovascular Stents
Vascular Grafts
Endovascular Stent-Grafts
Other Vascular Devices
Endovascular Catheters
Diagnostic Catheters
Therapeutic Catheters
Endovascular Coils
Vascular Filters
Vascular Closure Devices (VCDs)
Unintended Embolic Biomaterials
Ex Vivo Evaluation
Conclusions
2.5.3 - Extracorporeal Artificial Organs and Therapeutic Devices
Introduction
Extracorporeal Respiratory Support
Pulmonary Disease—Incidence, Causes, and Mortality
Extracorporeal Membrane Oxygenation (ECMO)
Alternative Extracorporeal Gas Exchange Devices
Oxygenator Biocompatibility Challenges: Coagulation and Inflammation
Surface Coatings
Nitric Oxide Surface Flux
Renal Replacement Therapies and Therapeutic Apheresis
Introduction
Renal Replacement Therapy
Function of the Kidney
Treatment of Renal Failure
Peritoneal Dialysis
Hemodialysis
Dialyzer Materials and Coatings
Coagulation and Inflammation During Hemodialysis
Extracorporeal Hemofiltration
Hemoperfusion
Therapeutic Apheresis
Plasmapheresis
Plasma Separation
Plasma Exchange
Plasma Treatment
Sorbent Dialysis
Blood Pumps in Extracorporeal Circulation
Roller Pumps
Summary
Chapter Exercises
2.5.4 - Orthopedic Applications
Biomaterials Development: A History of Total Hip Arthroplasty
Current Biomaterials in Total Arthroplasty
Orthopedic Biomaterials: Clinical Concerns
Orthopedic Biomaterial Wear
Orthopedic Biomaterial Corrosion
Fretting Corrosion or Mechanically Assisted Crevice Corrosion (MACC)
Implant Debris Types: Particles and Ions
Particulate Debris
Metal Ions (Soluble Debris)
Local Tissue Effects of Wear and Corrosion
Remote and Systemic Effects of Wear and Corrosion
Hypersensitivity
Carcinogenesis
Preventive Strategies and Future Directions
Chapter Study Questions
2.5.5 - Dental Applications
Overview
Unique Needs in Developing Biomaterials for DOC Procedures
Restorative Materials
Dental Implants
Criteria for Successful Implant Function
Osseointegration and Accelerating Healing and Attachment to Tissue
Surface Topology and Chemistry
Mechanical Parameters and Implant Design
Materials Used in Dental Implants
Metals
Ceramics
Future Directions
Tissue Engineering in Dentistry
Need for Tissue Engineering in Dentistry
Materials for Engineering DOC Tissue Structure and Function
DOC Tissue-Engineering Applications
Teeth
Temporomandibular Joint
Oral Mucosa
Salivary Glands
Bone and Periodontium
Summary and How Experience From Dental Biomaterials has Brought Value to Other Areas of Biomaterials
Chapter Exercises
2.5.6 - Ophthalmologic Applications: Introduction
Overview of the Anatomy of the Eye
Eye-related Conditions and Statistics
Considerations for Ophthalmic Materials
Biomaterials: Contact Lenses
Contact Lens Materials
Hard Contact Lenses
Soft Hydrogel Contact Lenses
Silicone Hydrogel Contact Lenses
Surface Modification
Contact Lens Solutions
Intraocular Lens Implants
Introduction to Intraocular Lens Implants, the Optics of the Eye, and Cataracts
IOL Biomaterials and Design
IOLs With Variations of Optical Function
Multifocal IOLs
Accommodative IOLs
Adjustable-Power IOLs
Summary and Future of IOLs
Glaucoma Drainage Devices
Aqueous Humor Production and Drainage
New-generation Microinvasive Glaucoma Surgery (MIGS) Implantation Devices
The Glaukos iStent Series
Summary
Retinal Implants
Epiretinal Devices
Argus II
The Intelligent Retinal Implant System II (IRIS II)
EPI-RET3 Retinal Implant System
Subretinal Devices
Alpha IMS/AMS
Photovoltaic Retinal Implant (PRIMA) Bionic Vision System
Suprachoroidal Devices
Bionic Vision Australia (BVA) Team
Suprachoroidal–Transretinal Stimulation (STS)
Conclusions and Future Directions
2.5.7 - Bioelectronic Neural Implants
Introduction
Bioelectronic Devices
Electrode Materials
Factors That Influence Materials Selection
Conducting/Capacitive Materials
Insulating Materials
Equivalent Circuit Models
Technologies
Battery, IPG
Leads and Interconnects
Electrode Contacts
Applications
Research
Rehabilitation
Sensory Restoration
Visual
Tactile
Auditory
Genitourinary, Bladder Dysfunction
Motor Function
Brain–Computer Interface
Bioelectronic Medicine, “Electroceuticals”
Regeneration
Failure Modes
Mechanical
Materials
Biological
Biomaterial-Based Strategies to Enable Neural Implants
Micromotion and Tissue Mechanics
Antioxidative Strategies
Conclusions and Future Directions
2.5.8 - Burn Dressings and Skin Substitutes
Burn Wounds
Surgical Planning for Wound Care
Ideal Properties of Dressings and Skin Substitutes
Topical Microbial Management
Negative-Pressure Dressings
Degradable Polymers
Temporary Skin Substitutes
Permanent Skin Substitutes
Cost Considerations
Regulatory Considerations
Conclusions and Future Directions
Chapter Exercises
2.5.9 - Description and Definition of Adhesives, and Related Terminology
Introduction
Description and Definition of Adhesives, and Related Terminology
The Logic of Adhesion Procedures
Hard-Tissue Adhesives: Bone and Tooth Cements
Autopolymerizing PMMA Bone Cement
Historical Background
Mechanism of Setting of PMMA/MMA Dough
Mechanism of “Bonding” or Grouting
Alternative Bone Cements: Calcium Phosphate
Classical and Modern Dental-Bonding Cements: Conventional Acid–Base Cements
Polyelectrolyte Cements: Zinc Polycarboxylates and Glass Ionomers
Acid-Etch Bonding to Enamel
Chemistry of Etchants, Primers, and Bonding Agents
Hybrid-Layer Creation Via A Three-Stage Approach: Etch, Prime, Bond
Aging and Stability of the Bonded Interface
Inhibitors for the Preservation of the Hybrid Interfacial Zone Between Adhesives and Human Dentin
Soft-Tissue Adhesives and Sealants
Performance Requirements
Historical Overview
The Relationship Between Soft-Tissue Adhesion and Drug Delivery
Cyanoacrylate Esters
Chemistry
Performance
Fibrin Sealants
Formulation, Presentation, and Setting Processes
Advantages and Applications
Bioadhesives
Hydrogel Sealants
New Research Directions: Biomimetic Approaches
Sutures
Genesis and Common Uses
Description of Surgical Sutures
Surgical Gut Sutures
Silk Sutures
Polyester Sutures
Nylon Sutures
Polypropylene Sutures
Ultrahigh-Molecular-Weight Polyethylene (UHMWPE) Sutures
Stainless Steel Sutures
Synthetic Absorbable Sutures
Monomers and Preparation of Polymers
Poly(Glycolic Acid) (PGA)
Poly(Dioxanone) (PDO) Sutures
High-Glycolide Copolymeric Sutures
Dyes
Coatings
Needles and Attachment
Packaging
Physical Properties
In Vitro and In Vivo Performance
Newer Trends and Future Developments
2.5.10 - Biomaterials for Immunoengineering
Use of Biomaterials in Vaccine Development
Introduction
Biomaterials for Improving Vaccine Efficacy
Use of Biomaterials to Adjuvant the Immune System
Use of Biomaterials to Improve Delivery of Antigen to APCs
Activation of B Cells and Humoral Immunity
Overview of the B Cell Activation Process
Biomaterial Design for Enhancing the Humoral Response
Biomaterials for Alternative Vaccine Administration Routes
Biomaterials for Improved Vaccine Manufacturing and Accessibility
Conclusion/Future Directions
Use of Biomaterials in T Cell Modulation
Introduction
Biomaterials for Targeting and Modulation of T Cell Therapies
Biomaterials for Enhanced T Cell Manufacturing
Conclusions/Future Directions
Use of Biomaterials to Induce Tolerance
Introduction
Induction of Tolerance in Autoimmune Disorders
T Cell Anergy and Deletion Through Incomplete Dendritic Cell Activation
Elevation of Treg Activity to Induce Tolerance
Suppression of B Cell Activation
Concluding Remarks
Exercises
2.5.11 - Biomaterials-Based Model Systems to Study Tumor–Microenvironment Interactions
Introduction
Biological Design Considerations
Tissue Dimensionality
Transport Phenomena and Interstitial Pressure
ECM Physicochemical Properties
Immunological Changes
Biomaterials to Study the Tumor Microenvironment
Natural Biomaterials
Protein-Based Materials
Carbohydrate-Based Materials
Cell- and Tissue-Derived Materials
Synthetic Biomaterials
Synthetic Hydrogels
Polyesters
Composite Materials
Applications of Biomaterials-Based Tumor Models
Analyzing the Effect of Tissue Dimensionality
Modeling Tumor–Stroma Interactions
Platforms to Interrogate Cell–ECM Interactions
Dynamic Materials Systems for Studies of Mechanical Memory
Analyzing the Effect of Local and Systemic Transport Phenomena
Metastasis
Conclusions
Chapter Exercises With (Guided) Solutions
2.5.12 - Drug Delivery Systems
History of DDS Development
General Considerations in DDS Design
Routes of Drug Delivery
DDS Biomaterials Design Considerations
Biomaterials Used in DDSs
DDS Biomaterial Properties
Degradation
Surface Properties
Mechanics
DDSs to Improve Drug Pharmacokinetics
Pharmacokinetics
Dosage and Distribution Control
Controlling Drug Release Kinetics
DDSs to Improve Drug Solubility
Colloidal DDSs
Noncolloidal DDSs
Biomaterial DDSs Can Enhance Drug Stability
Small Molecule Drugs
Protein/Peptide Drugs
Nucleic Acid Drugs
DDS Design to Overcome Biological Barriers
Epithelial Barriers
Parenteral Administration
Transdermal DDSs
Mucosal DDSs
Oral DDSs
Endothelial Barriers
Biomaterial DDSs for Drug Targeting
Passive Targeting
Active Targeting
Antibodies
Proteins
Peptides
Aptamers
Carbohydrates
Small Molecules
Regulatory and Intellectual Property Considerations for DDSs
Regulation
Intellectual Property
Final Remarks
Chapter Review Questions
2.5.13 - Responsive Polymers in the Fabrication of Enzyme-Based Biosensors
Introduction
Classic Biosensor System
Bioreceptor (Recognition Layer)
Physicochemical Transducers
Computer Processing
Types of Enzymatic Glucose Biosensors
Electrochemical Biosensors
Amperometric Biosensors
Conductometric Biosensors
Impedimetric Biosensors
Potentiometric Biosensors
Optical Biosensors
Piezoelectric Biosensors
Thermal Biosensors
Roles of Responsive Polymers in Enzymatic Biosensors
Passive Roles (Physical Support)
Covalent Linkage
Cross-Linking
Entrapment
Encapsulation
Active Roles
Redox Mediators
Chromogenic Agents
Fast Ion Conductors
Fluorescence Probes
Integrating Responsive Polymers With Enzymes
Physicochemical Conjugation With CNTs
Active Site Conjugation Using Boric Acid
Molecular Wiring
Covalent Conjugation
Integrating Responsive Polymers With Transducers
Interface Engineering
Systems Integration
Microfabrication
Three-Dimensional (3-D) Bioprinting
Future Outlook
2.6.1 - Rebuilding Humans Using Biology and Biomaterials
2.6.2 - Overview of Tissue Engineering Concepts and Applications
General Introduction
History of Tissue Engineering
Goals of Tissue Engineering and Classification
Goals of Tissue Engineering
Classification of Tissue-Engineering Approaches
Components of Tissue Engineering
The Cell
Materials
Biological Factors
Scaffold Design
Integration of Multiple Factors
Models for Tissue Engineering
Bioreactors
Organoids
In Vivo Models
Applications of Tissue Engineering
Transplantation
Replacing/Regenerating Target Organs
Drug Delivery
Disease Models and Therapy
Organ-On-a-Chip Systems
Current Challenges and Opportunities
Cell Source
Vascularization
Tissue Maturation
In Vivo Integration
FDA Regulations for Clinical Translation
Gene Editing and CRISPR
Future Perspectives
2.6.3 - Tissue Engineering Scaffolds
Introduction
Scaffold Design Criteria
Scaffold Applications
Cell Delivery
Drug and Biomolecule Delivery
Scaffold Materials
Polycondensation Polymers
Ring-Opening Polymerization
Click Reactions
Polyaddition Polymers
Ionic Polymerization
Free Radical Polymerization
Biological Polymers
Composites and Additives
Scaffold Fabrication Techniques
Rapid Prototyping
Electrospinning/Electrospraying
Superstructure Engineering
Solvent Casting, Particulate/Porogen Leaching
Freeze-Drying
Phase Separation
Gas Foaming/Supercritical Fluid Processing
Scaffold Characterization Techniques
Cell-Incorporated Scaffolds
Conclusions
Chapter Exercises
Chapter Exercise Answers
2.6.4 - Micromechanical Design Criteria for Tissue-Engineering Biomaterials
Introduction
Cell–Matrix Interactions and Mechanotransduction
Focal Adhesion
Roles of Focal Adhesion Maturation and Stress Fiber Formation in Mechanotransduction
Important Mechanotransduction Molecular Pathways for Design of Scaffolds
Direct Transmission of Forces to the Nucleus
Design Considerations for Scaffolds to Regulate Tissue Development
Local Stiffness
Surface Topography
Fibrous Scaffolds
Multicellular Interactions
Mechanical Stimulation
Effects of Combined Mechanical Stimuli
Implications for Future Materials Design
Conclusion
2.6.5 - Tendon Tissue-Engineering Scaffolds
Introduction
Native Adult Tendon Properties
Mechanical Properties
Extracellular Matrix Composition and Molecular Arrangement
Tendon Cells
Scaffold Design Goals
Immediate or Early Return to Load-Bearing Function
Guidance Cues to Induce Tenogenic Cell Behaviors
Fabrication Methods
Spinning
Textile Technologies
Gelation
Freeze Drying
Decellularization
Postfabrication Modifications
Delivery of Bioactive Molecules (e.g., Drugs and Growth Factors)
Polymer Selection and Scaffold Designs
Natural Plant- and Animal-Derived Polymers
Alginates
Chitin/Chitosan
Collagen
Gelatin
Silk
Synthetic Polymers
Poly(Glycolic Acid) and Poly(Lactic Acid)
Poly(ε-Caprolactone)
Other Synthetic Polymers
Future Directions: Developmental Biology-Inspired Strategies
Summary
2.6.6 - Bone Tissue Engineering
Introduction
Bone Biology
Types of Bone Tissue
Cells Involved
Osteoblasts
Bone Lining Cells
Osteocytes
Osteoclasts
Bone Tissue Development
Intramembranous Ossification
Endochondral Ossification
Bone Tissue Engineering
Bone Grafts
Autograft
Allograft
Bone Graft Substitutes
Allograft-Based Substitutes
Natural Polymer-Based Substitutes
Synthetic Polymer-Based Substitutes
Ceramic-Based Substitutes
Cell-Based Substitutes
Growth Factor-Based Substitutes
Composite Substitutes
Porosity in Bone Graft Substitutes
Dimension in Bone Graft Substitutes
Sintered Microspheres
Nanofibers
In Vitro Culture Techniques for Bone Graft Substitutes
Conclusion
2.6.7 - Biomaterials for Cardiovascular Tissue Engineering
Introduction
Endothelial Cells
Cardiac Muscle
Heart Valves
Blood Vessels
Scaffold Materials
Protein Hydrogels
Decellularized Tissues
Synthetic Polymers
Synthetic Hydrogels
Conclusions
2.6.8 - Soft Tissue Engineering
Introduction
Properties of Soft Tissues
Common Biomaterials Used for Soft Tissue Engineering
Synthetic Polymers
Natural Polymers
Decellularized Tissues
Soft Tissue Engineering Applications: Adipose, Gastrointestinal, and Skin
Adipose Tissue Engineering
Anatomy and Physiology
Design Criteria for Adipose Tissue Engineering
Commercially Available and Clinically Tested Biomaterials
Novel Materials and Technologies
Challenges
Gastrointestinal Tissue Engineering
Anatomy and Physiology
Gastrointestinal Disorders and the Need for Tissue Engineering
Design Criteria for Engineered Gastrointestinal Tissues
Gastrointestinal Soft Tissue-Engineering Strategies
Challenges and Future Goals
Tissue-Engineered Skin: Future Goals of Skin Substitutes
Anatomy and Physiology
Design Criteria
Skin Substitute Technology
Challenges
Conclusions
Chapter Exercises
3.1.1 - Introduction: Biomaterials in Medical Devices
3.1.2 - Total Product Lifecycle for Biomaterial-Based Medical Devices
Chapter Questions for the Student
3.1.3 - Safety and Risk Considerations in Medical Device Development
Introduction
Absence of Toxicity Is Not Evidence of Safety
Assessing the Continuum of Biological Risk in Performance
Assessing the Contribution of Secondary Processes to Biological Risk
Assessing Biological Risk of Aging Biomaterials in the Aging Patient
Summary and Study Guide
Chapter Study Guide
3.1.4 - Sterilization and Disinfection of Biomaterials for Medical Devices
Introduction
Radiation-Based Techniques
Safety Considerations
Principles of Action and Efficacy
Gamma Sterilization
Electron Beam Sterilization
X-Ray Sterilization
Application Considerations
R&D, Pilot, and Low-Volume Technologies
Material Considerations for Radiation Sterilization
Biologics and Human-Based Tissue: Compatibility With Radiation Sterilization
Chemical Techniques
Safety Considerations
Principles of Action and Efficacy
Ethylene Oxide Sterilization
Sterilization by Oxidation: Hydrogen Peroxide or Ozone
Physicochemical Methods: Gas Plasma
Material Considerations for Chemical Sterilization
R&D, Pilot, and Low-Volume Technologies
Pharmaceuticals and Biologics: Compatibility With EO Sterilization
Thermal Techniques
Safety Considerations
Principles of Action and Efficacy
Dry Heat Sterilization
Steam Sterilization and Disinfection
Application Considerations
Materials Development Considerations for Sterility
Safety Testing and Validation After Sterilization
Patient Safety: FDA Recall Classifications: Class I, Class II, and Class III
Biological Safety Verification
Maintaining Sterility: Packaging and Shelf Life
Sourcing, Quality Systems, and Manufacturing Controls
Sterilization Standards
Summary and Future Challenges
Chapter Exercises
3.1.5 - Verification and Validation: From Bench to Human Studies
Introduction: Focusing on Commercial Medical Device Development
Starting a Medical Device Project
Design Controls for Medical Device Development
Verification of Medical Device Design
Types of Verification Testing
Validation of Medical Device Design
Verification Versus Validation
Concluding Remarks: Design Transfer Beyond Human Studies
Chapter Study Questions
3.1.6 - Commercial Considerations in Medical Device Development
Introduction
Traditional Model of Product Development
Determining Market Opportunity
Medical Device Reimbursement
Securing Intellectual Property and Funding
Intellectual Property
Securing Funding
Commercial Operations: Sales and Marketing
Summary
Student Questions
3.1.7 - Regulatory Constraints for Medical Products Using Biomaterials
Introduction and History in the United States
Global Premarket Assessment Methods
Premarket Assessment Requirements
Premarket Clearance and Approvals
Manufacturing and Material Supplier Controls
Postmarketing Management of Risk and Product Performance
Registration, Device Listing, Licenses
Summary and Study Guide
Chapter Questions: True or False
3.1.8 - Role of Standards for Testing and Performance Requirements of Biomaterials
Introduction: What Is a Standard
Reference Materials
Reference Data
Documentary Standards
Documentary Standards: Voluntary, Consensus
Who Writes Documentary Standards?
How Are Documentary Standards Developed?
Applications of Documentary Standards
Accelerating the Regulatory Process
Specificity versus Universality
A Standard Test Method Does Not Necessarily Define the Best Measurement
Clinical Relevance
Measurement Assurance
Interlaboratory Comparison Studies
Looking Ahead
Conclusion
Homework Questions
Answer Key for Homework Questions
3.1.9 - Medical Device Failure—Implant Retrieval, Evaluation, and Failure Analysis
Overview and Definitions
Medical Implants
Implant Retrieval
Postmarket Surveillance
Goals for Implant Retrieval and Evaluation and Failure Mode Analysis
Medical Surveillance and the Role of Retrieval Analysis in Device Development
Chapter Exercises (With Answers)
3.1.10 - Legal Concepts for Biomaterials Engineers
Introduction
Employment Agreements
Confidentiality and Materials Use Agreements
Intellectual Property: Patents, Trade Secrets, and Freedom to Operate
Contract Negotiation, Performance, and Compliance
Sponsored Research Agreements
License Agreements
Litigation
Conclusion
3.1.11 - Moral and Ethical Issues in the Development of Biomaterials and Medical Products
Introduction
Selected Approaches to Ethical Reasoning
The Utilitarian Approach
The Rights Approach
The Justice Approach
The Virtue Approach
Safety
Animal Testing
Human Testing
Research Integrity
Conflict of Interest
Emerging Ethical Issues in Medical Product Development
Ethical Issues in Stem Cell Research
Gene Editing
Cost and Access to Medical Products
Conclusions
A - Properties of Biological Fluids
B - Properties of Soft Materials
C - Chemical Composition of Metals and Ceramics Used for Implants
D - The Biomaterials Literature
E - Assessment of Cell and Matrix Components in Tissues
Light Microscopy
Special Staining
Immunohistochemical Staining
In Situ Hybridization
Electron Microscopy
Special Techniques
Index
A
B
C
D
E
F
G
H
I
J
K
L
M
N
O
P
Q
R
S
T
U
V
W
X
Y
Z

Citation preview

Biomaterials Science An Introduction to Materials in Medicine

Fourth Edition

Edited By

William R. Wagner

Guigen Zhang

Distinguished Professor of Surgery Bioengineering & Chemical Engineering University of Pittsburgh Director McGowan Institute for Regenerative Medicine Pittsburgh, PA, United States

Professor and F Joseph Halcomb III, M.D. Endowed Chair, Chair of the F. Joseph Halcomb III, M.D. ­Department of Biomedical Engineering University of Kentucky Lexington, KY, United States

Shelly E. Sakiyama-Elbert Professor and Department Chair Fletcher Stuckey Pratt Chair in Engineering Department of Biomedical Engineering The University of Texas at Austin Austin, TX, United States

Founding Editors

Buddy D. Ratner, Allan S. Hoffman Frederick J. Schoen, Jack E. Lemons

Michael J. Yaszemski Krehbiel Family Endowed Professor of Orthopedics and Biomedical Engineering, Mayo Clinic Rochester, MN, United States

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 © 2020 Elsevier Inc. All rights reserved. Chapter 11: Silicones: Copyright © 2020 DuPont, Published by 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: 978-0-12-816137-1 For information on all Academic Press publications visit our website at https://www.elsevier.com/books-and-journals Publisher: Matthew Deans Acquisitions Editor: Sabrina Webber Editorial Project Manager: Naomi Robertson Production Project Manager: Surya Narayanan Jayachandran Cover Designer: Alan Studholme Typeset by TNQ Technologies

List of Contributors

Abhinav Acharya School for Engineering of Matter, Transport and Energy, Biodesign Center for Immunotherapy, Vaccines and Virotherapy, Arizona State University, Tempe, AZ, United States Marian A. Ackun-Farmmer Department of Biomedical Engineering, University of Rochester, Rochester, NY, United States; Department of Orthopaedics and Center for Musculoskeletal Research, University of Rochester, Rochester, NY, United States John R. Aggas Center for Bioelectronics, Biosensors and Biochips (C3B), Department of Biomedical Engineering, Texas A&M University, College Station, TX, United States; Department of Biomedical Engineering, Texas A&M University, College Station, TX, United States Phillip J. Andersen Andersen Metallurgical, LLC, Madison, WI, United States James M. Anderson Department of Pathology, Case Western Reserve University, Cleveland, OH, United States Kristi Anseth Department of Chemical and Biological Engineering, University of Colorado, Boulder, CO, United States; Biofrontiers Institute, University of Colorado, Boulder, CO, United States Paul A. Archer Parker H. Petit Institute for Bioengineering and Bioscience, Georgia Institute of Technology, Atlanta, GA, United States; School of Chemical and Biomolecular Engineering, Georgia Institute of Technology, Atlanta, GA, United States Nureddin Ashammakhi Center for Minimally Invasive Therapeutics (C-MIT), California NanoSystems Institute (CNSI), University of California-Los Angeles, Los Angeles, CA, United States

Jose D. Avila W. M. Keck Biomedical Materials Research Laboratory, Washington State University, Pullman, WA, United States Julia E. Babensee Wallace H. Coulter Department of Biomedical Engineering, Georgia Institute of Technology and Emory University, Atlanta, GA, United States Stephen f. Badylak University of Pittsburgh, Pittsburgh, PA, United States Kiheon Baek Department of Biomedical Engineering, University of Rochester, Rochester, NY, United States; Center for Musculoskeletal Research, University of Rochester Medical Center, Rochester, NY, United States Aaron B. Baker Department of Biomedical Engineering, University of Texas at Austin, Austin, TX, United States Syeda Mahwish Bakht 3B’s Research Group, I3Bs–Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, Guimarães, Portugal; ICVS/3B’s–PT Government Associate Laboratory, Braga/Guimarães, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal Amit Bandyopadhyay W. M. Keck Biomedical Materials Research Laboratory, Washington State University, Pullman, WA, United States Aaron Barchowsky Department of Environmental and Occupational Health, University of Pittsburgh, Pittsburgh, PA, United States

xi

xii   List of Contributors

Garrett Bass Departments of Chemistry, Mechanical Engineering and Materials Science, Biomedical Engineering and Orthopaedic Surgery, Duke University, Durham, NC, United States Matthew L. Becker Departments of Chemistry, Mechanical Engineering and Materials Science, Biomedical Engineering and Orthopaedic Surgery, Duke University, Durham, NC, United States Sarah Miho Van Belleghem University of Maryland College Park; NIH/NIBIB Center for Engineering Complex Tissues Danielle S.W. Benoit Department of Biomedical Engineering, University of Rochester, Rochester, NY, United States; Department of Orthopaedics and Center for Musculoskeletal Research, University of Rochester, Rochester, NY, United States; Materials Science Program, University of Rochester, Rochester, NY, United States; Department of Chemical Engineering, University of Rochester, Rochester, NY, United States; Department of Biomedical Genetics and Center for Oral Biology, University of Rochester, Rochester, NY, United States; Translational Biomedical Science Program, University of Rochester, Rochester, NY, United States Arne Biesiekierski School of Engineering, RMIT University, Bundoora, VIC, Australia Kristen L. Billiar Biomedical Engineering Department, Worcester Polytechnic Institute, Worcester, MA, United States Susmita Bose W. M. Keck Biomedical Materials Research Laboratory, Washington State University, Pullman, WA, United States Christopher Bowman Department of Chemical and Biological Engineering, University of Colorado, Boulder, CO, United States Steven Boyce Department of Surgery, University of Cincinnati and Shriners Hospitals for Children – Cincinnati, Cincinnati, OH, United States Bryan N. Brown Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA, United States; McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, United States

Justin L. Brown Department of Biomedical Engineering, The Pennsylvania State University, University Park, PA, United States Jeffrey R. Capadona Department of Biomedical Engineering, Case Western Reserve University, Cleveland, OH, United States; Advanced Platform Technology Center, Rehabilitation Research and Development, Louis Stokes Cleveland VA Medical Center, Cleveland, OH, United States David G. Castner University of Washington, Seattle, WA, United States Calvin Chang Department of Biomedical Engineering, Johns Hopkins University, Baltimore, MD, United States; Translational Tissue Engineering Center, Johns Hopkins School of Medicine, Baltimore, MD, United States; Institute for NanoBioTechnology, Johns Hopkins University, Baltimore, MD, United States Philip Chang Department of Surgery, University of Cincinnati and Shriners Hospitals for Children – Cincinnati, Cincinnati, OH, United States Ashutosh Chilkoti Department of Biomedical Engineering, Duke University, Durham, NC, United States Karen L. Christman Department of Bioengineering, Sanford Consortium of Regenerative Medicine, University of California San Diego, La Jolla, CA, United States Sangwon Chung Department of Textile Engineering, Chemistry & Science, North Carolina State University, Raleigh, NC, United States; Biomedical Engineering, Joint Department of Biomedical Engineering, North Carolina State University and University of North Carolina at Chapel Hill, Chapel Hill, NC, United States Kelly P. Coleman Medtronic, Physiological Research Laboratories, Minneapolis, MN, United States Dan Conway Department of Biomedical Engineering, Virginia Commonwealth University, Richmond, VA, United States Keith E. Cook Department of Biomedical Engineering, Carnegie Mellon University, Pittsburgh, PA, United States

  List of Contributors

Stuart L. Cooper William G. Lowrie Department of Chemical and Biomolecular Engineering, The Ohio State University, Columbus, OH, United States

Pedro Esbrit Chemical Department of Pharmaceutical Sciences, Faculty of Pharmacy, Universidad Complutense of Madrid, Spain

Elizabeth Cosgriff-Hernandez The University of Texas at Austin, Austin, TX, United States

Suzanne G. Eskin Wallace H. Coulter Department of Biomedical Engineering, Georgia Institute of Technology, Atlanta, GA, United States

Arthur J. Coury Northeastern University, Boston, MA, United States Joseph D. Criscione Department of Biomedical Engineering, University of Rochester, Rochester, NY, United States Heidi Culver Department of Chemical and Biological Engineering, University of Colorado, Boulder, CO, United States

Michael Y. Esmail Tufts Comparative Medicine Services, Tufts University, Boston, MA, United States Jack L. Ferracane Department of Restorative Dentistry, Oregon Health & Science University, Portland, OR 97201

Jim Curtis DuPont Health Care Solutions, Midland, MI, United States

Claudia Fischbach Nancy E. and Peter C. Meinig School of Biomedical Engineering, Cornell University, Ithaca, NY, United States; Kavli Institute at Cornell for Nanoscale Science, Cornell University, Ithaca, NY, United States

Feiyang Deng Department of Biomedical Engineering, University of Rochester, Rochester, NY, United States

Gary Fischman, PhD Future Strategy Solutions, LLC, Gambrills, MD, United States

Prachi Dhavalikar The University of Texas at Austin, Austin, TX, United States

John P. Fisher University of Maryland College Park; NIH/NIBIB Center for Engineering Complex Tissues

Luis Diaz-Gomez Department of Bioengineering, Rice University, Houston, TX, United States

Iolanda Francolini Department of Chemistry, Sapienza University of Rome, Rome, Italy

Rui M.A. Domingues 3B’s Research Group, I3Bs–Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, Guimarães, Portugal; ICVS/3B’s–PT Government Associate Laboratory, Braga/Guimarães, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal

Steven J. Frey School of Chemical and Biomolecular Engineering, Georgia Institute of Technology, Atlanta, GA, United States

Elaine Duncan Paladin Medical, Inc., Stillwater, MN, United States; Joseph Halcomb III, MD. Department of Biomedical Engineering, College of Engineering, University of Kentucky, Lexington, KY, United States Pamela Duran Department of Bioengineering, Sanford Consortium of Regenerative Medicine, University of California San Diego, La Jolla, CA, United States

Akhilesh K. Gaharwar Texas A&M University, College Station, TX, United States Andrés J. García George W. Woodruff School of Mechanical Engineering, Georgia Institute of Technology, Atlanta, GA, United States; Parker H. Petit Institute of Bioengineering and Biological Science, Georgia Institute of Technology, Atlanta, GA, United States Iain R. Gibson Institute of Medical Sciences, School of Medicine, Medical Sciences and Nutrition, University of Aberdeen, Aberdeen, United Kingdom

xiii

xiv   List of Contributors

Jeremy L. Gilbert Department of Bioengineering, Clemson University, Charleston, SC, United States Brian Ginn Secant Group, Telford, PA, United States Zachary E. Goldblatt Biomedical Engineering Department, Worcester Polytechnic Institute, Worcester, MA, United States Seth J. Goldenberg Veeva Systems, Pleasanton, CA, United States Manuela E. Gomes 3B’s Research Group, I3Bs–Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, Guimarães, Portugal; ICVS/3B’s–PT Government Associate Laboratory, Braga/Guimarães, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal Manuel Gómez-Florit 3B’s Research Group, I3Bs–Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, Guimarães, Portugal; ICVS/3B’s–PT Government Associate Laboratory, Braga/Guimarães, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal Inês C. Gonçalves i3S - Instituto de Inovação e Investigação em Saúde, Universidade do Porto, Rua Alfredo Allen 208, Porto, Portugal; INEB - Instituto de Engenharia Biomédica, Universidade do Porto, Rua Alfredo Allen 208, Porto, Portugal Maud B. Gorbet Department of Systems Design Engineering, University of Waterloo, ON, Canada David W. Grainger Department of Biomedical Engineering, University of Utah, Salt Lake City, UT, United States; Department of Pharmaceutics and Pharmaceutical Chemistry, University of Utah, Salt Lake City, UT, United States

Miles Grody Miles Grody Law, Potomac, MD, United States Teja Guda Department of Biomedical Engineering, University of Texas at San Antonio, San Antonio, TX, United States Scott A. Guelcher Department of Chemical and Biomolecular Engineering, Vanderbilt University, Nashville, TN, United States Anthony Guiseppi-Elie Center for Bioelectronics, Biosensors and Biochips (C3B), Department of Biomedical Engineering, Texas A&M University, College Station, TX, United States; Department of Biomedical Engineering, Texas A&M University, College Station, TX, United States; ABTECH Scientific, Inc., Biotechnology Research Park, Richmond, VA, United States S. Adam Hacking Laboratory for Musculoskeletal Research and Innovation, Department of Orthopaedic Surgery, Massachusetts General Hospital and Harvard Medical School, Boston, MA, United States Nadim James Hallab Department of Orthopedic Surgery, Rush University Medical Center, Chicago, IL, United States Luanne Hall-Stoodley Department of Microbial Infection and Immunity, The Ohio State University, Columbus, OH, United States Stephen R. Hanson Division of Biomedical Engineering, School of Medicine, Oregon Health & Science University, Portland, OR, United States Woojin M. Han George W. Woodruff School of Mechanical Engineering, Georgia Institute of Technology, Atlanta, GA, United States; Parker H. Petit Institute of Bioengineering and Biological Science, Georgia Institute of Technology, Atlanta, GA, United States Melinda K. Harman Department of Bioengineering, Clemson University, Clemson, SC, United States Roger Harrington Medical Products Development Director, Medtronic, Boston, MA, United States

  List of Contributors

Martin J. Haschak Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA, United States; McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, United States Daniel E. Heath William G. Lowrie Department of Chemical and Biomolecular Engineering, The Ohio State University, Columbus, OH, United States Emily Anne Hicks Department of Chemical Engineering, McMaster University, Hamilton, ON, Canada Ryan T. Hill Center for Biologically Inspired Materials and Material Systems, Duke University, Durham, NC, United States Allan S. Hoffman Bioengineering and Chemical Engineering, University of Washington, Seattle, WA, United States Thomas A. Horbett Bioengineering and Chemical Engineering, University of Washington, Seattle, WA, United States Jeffrey A. Hubbell The University of Chicago, Chicago, IL, United States Rasim Ipek Department of Mechanical Engineering, Ege University, Izmir, Turkey Joshua J. Jacobs Department of Orthopedic Surgery, Rush University Medical Center, Chicago, IL, United States Young C. Jang School of Biological Sciences, Georgia Institute of Technology, Atlanta, GA, United States; Parker H. Petit Institute of Bioengineering and Biological Science, Georgia Institute of Technology, Atlanta, GA, United States Shaoyi Jiang Departments of Chemical Engineering and Bioengineering, Seattle, WA, United States Richard J. Johnson BioPhia Consulting, Lake Forest, IL, United States Julian R. Jones Department of Materials, Imperial College London, South Kensington Campus, London, United Kingdom

Vickie Y. Jo Department of Pathology, Brigham and Women’s Hospital and Harvard Medical School, Boston, MA, United States Ravi S. Kane School of Chemical and Biomolecular Engineering, Georgia Institute of Technology, Atlanta, GA, United States David L. Kaplan Department of Biomedical Engineering, Tufts University, Medford, MA, United States Ronit Kar The University of Texas at Austin, Austin, TX, United States Benjamin George Keselowsky Department of Biomedical Engineering, University of Florida, Gainesville, FL, United States Ali Khademhosseini Center for Minimally Invasive Therapeutics (C-MIT), California NanoSystems Institute (CNSI), University of California-Los Angeles, Los Angeles, CA, United States Yu Seon Kim Department of Bioengineering, Rice University, Houston, TX, United States Martin W. King Department of Textile Engineering, Chemistry & Science, North Carolina State University, Raleigh, NC, United States Daniel S. Kohane Laboratory for Biomaterials and Drug Delivery, Department of Anesthesiology, Division of Critical Care Medicine, Boston Children’s Hospital, Harvard Medical School, Boston, MA, United States David H. Kohn Departments of Biologic and Materials Sciences, and Biomedical Engineering, The University of Michigan, Ann Arbor, MI, United States Liisa T. Kuhn Department of Biomedical Engineering, University of Connecticut Health Center, Farmington, CT, United States

xv

xvi   List of Contributors

Mangesh Kulkarni Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA, United States; McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, United States Catherine K. Kuo Department of Biomedical Engineering, University of Rochester, Rochester, NY, United States; Center for Musculoskeletal Research, University of Rochester Medical Center, Rochester, NY, United States; Department of Orthopaedics, University of Rochester Medical Center, Rochester, NY, United States

Jack E. Lemons Schools of Dentistry, Medicine and Engineering, University of Alabama at Birmingham, Birmingham, AL, United States Robert J. Levy Department of Pediatrics, The Childrens’ Hospital of Philadelphia, The Perelman School of Medicine at the University of Pennsylvania, Philadelphia, PA, United States Gregory M. Lewerenz Medtronic, Physiological Research Laboratories, Minneapolis, MN, United States

Angela Lai Department of Biomedical Engineering, Carnegie Mellon University, Pittsburgh, PA, United States

Jamal S. Lewis Biomedical Engineering, University of California Davis, CA, United States

Bryron Lambert Sterilization Science, Abbott Vascular, Temecula, CA, United States

Simone Liebscher Department of Women’s Health, Research Institute for Women’s Health, Eberhard Karls University Tübingen, Tübingen, Germany

Ziyang Lan The University of Texas at Austin, Austin, TX, United States Robert A. Latour Bioengineering Department, Clemson University, Clemson, SC, United States Cato T. Laurencin Department of Biomedical Engineering, The Pennsylvania State University, University Park, PA, United States Bryan K. Lawson, M.D. Department of Orthopaedic Surgery, Mike O’Callaghan Military Medical Center, Nellis AFB, NV, United States Shannon Lee Layland Department of Women’s Health, Research Institute for Women’s Health, Eberhard Karls University Tübingen, Tübingen, Germany Jae Sung Lee Department of Orthopedics and Rehabilitation, University of Wisconsin–Madison, Madison, WI, United States David Lee-Parritz Department of Environmental and Population Health, Tufts University Cummings School of Veterinary Medicine, North Grafton, MA, United States Ying Lei Biomedical Engineering Department, Worcester Polytechnic Institute, Worcester, MA, United States

Chien-Chi Lin Department of Biomedical Engineering, Indiana University-Purdue University Indianapolis, Indianapolis, IN, United States Natalie K. Livingston Department of Biomedical Engineering, Johns Hopkins University, Baltimore, MD, United States; Translational Tissue Engineering Center, Johns Hopkins School of Medicine, Baltimore, MD, United States; Institute for NanoBioTechnology, Johns Hopkins University, Baltimore, MD, United States Yang Li Laboratory for Biomaterials and Drug Delivery, Department of Anesthesiology, Division of Critical Care Medicine, Boston Children’s Hospital, Harvard Medical School, Boston, MA, United States Yuncang Li School of Engineering, RMIT University, Bundoora, VIC, Australia Helen Lu Department of Biomedical Engineering, Columbia University, New York, NY, United States Laura Macdougall Department of Chemical and Biological Engineering, University of Colorado, Boulder, CO, United States; Biofrontiers Institute, University of Colorado, Boulder, CO, United States

  List of Contributors

Bhushan Mahadik University of Maryland College Park; NIH/NIBIB Center for Engineering Complex Tissues

Lei Mei Department of Biomedical Engineering, University of Texas at Austin, Austin, TX, United States

Sachin Mamidwar Orthogen, LLC, Springfield, NJ, United States

Bárbara B. Mendes 3B’s Research Group, I3Bs–Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, Guimarães, Portugal; ICVS/3B’s–PT Government Associate Laboratory, Braga/Guimarães, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal

Margaret P. Manspeaker Parker H. Petit Institute for Bioengineering and Bioscience, Georgia Institute of Technology, Atlanta, GA, United States; School of Chemical and Biomolecular Engineering, Georgia Institute of Technology, Atlanta, GA, United States Hai-Quan Mao Department of Biomedical Engineering, Johns Hopkins University, Baltimore, MD, United States; Translational Tissue Engineering Center, Johns Hopkins School of Medicine, Baltimore, MD, United States; Institute for NanoBioTechnology, Johns Hopkins University, Baltimore, MD, United States; Department of Materials Science and Engineering, Johns Hopkins University, Baltimore, MD, United States

Antonios G. Mikos Department of Bioengineering, Rice University, Houston, TX, United States Richard N. Mitchell Department of Pathology/Brigham and Women’s Hospital and Harvard Medical School, Boston, MA, United States

Peter X. Ma Department of Biologic and Materials Science, School of Dentistry, University of Michigan, Ann Arbor, MI, United States

Indranath Mitra W. M. Keck Biomedical Materials Research Laboratory, Washington State University, Pullman, WA, United States

Tyler Marcet Department of Biomedical Engineering, Tufts University, Medford, MA, United States

Ben Muirhead Department of Chemical Engineering, McMaster University, Hamilton, ON, Canada

Jeffrey Martin President and Principal Consultant, Sterilization and Quality System Consulting LLC, Dallas, TX, United States

Khurram Munir School of Engineering, RMIT University, Bundoora, VIC, Australia

M. Cristina L. Martins INEB - Institute of Biomedical Engineering, University of Porto, Porto, Portugal Sally L. McArthur Bioengineering Research Group, Swinburne University of Technology, Melbourne, VIC, Australia; Biomedical Manufacturing, CSIRO Manufacturing, Melbourne, VIC, Australia Meghan McGill Department of Biomedical Engineering, Tufts University, Medford, MA, United States Larry V. McIntire Wallace H. Coulter Department of Biomedical Engineering, Georgia Institute of Technology, Atlanta, GA, United States

William L. Murphy Department of Orthopedics and Rehabilitation, University of Wisconsin–Madison, Madison, WI, United States; Department of Biomedical Engineering, University of Wisconsin–Madison, Madison, WI, United States Phong K. Nguyen Department of Biomedical Engineering, University of Rochester, Rochester, NY, United States; Center for Musculoskeletal Research, University of Rochester Medical Center, Rochester, NY, United States Alexis L. Nolfi Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA, United States; McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, United States

xvii

xviii   List of Contributors

Clyde Overby Department of Biomedical Engineering, University of Rochester, Rochester, NY, United States; Department of Orthopaedics and Center for Musculoskeletal Research, University of Rochester, Rochester, NY, United States Sertan Ozan School of Engineering, RMIT University, Bundoora, VIC, Australia; Department of Mechanical Engineering, Yozgat Bozok University, Yozgat, Turkey Robert F. Padera Department of Pathology/Brigham and Women’s Hospital and Harvard Medical School, Boston, MA, United States Hannah A. Pearce Department of Bioengineering, Rice University, Houston, TX, United States Nicholas A. Peppas Department of Chemical Engineering and Biomedical Engineering, Pediatrics, Surgery and Molecular Pharmaceutics and Drug Delivery, The University of Texas, Austin, TX, United States Andreia T. Pereira i3S - Instituto de Inovação e Investigação em Saúde, Universidade do Porto, Rua Alfredo Allen 208, Porto, Portugal; INEB - Instituto de Engenharia Biomédica, Universidade do Porto, Rua Alfredo Allen 208, Porto, Portugal; Graduate Program in Areas of Basic and Applied Biology, Instituto de Ciências Biomédicas Abel Salazar, Universidade do Porto, Rua Jorge de Viterbo Ferreira 228, Porto, Portugal Carmem S. Pfeifer Biomaterials and Biomechanics, Oregon Health & Science University, Portland, OR, 97201 Artur M. Pinto LEPABE – Laboratory of Process Engineering, Environment, Biotechnology and Energy, Faculty of Engineering, University of Porto, Portugal Nicole R. Raia Department of Biomedical Engineering, Tufts University, Medford, MA, United States Edward A. Rankin Medtronic, Physiological Research Laboratories, Minneapolis, MN, United States Buddy D. Ratner Bioengineering and Chemical Engineering, Director of University of Washington Engineered Biomaterials (UWEB), Seattle, WA, United States

Maria Vallet Regi Chemical Department of Pharmaceutical Sciences, Faculty of Pharmacy, Universidad Complutense of Madrid, Spain Rui L. Reis 3B’s Research Group, I3Bs–Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, Guimarães, Portugal; ICVS/3B’s–PT Government Associate Laboratory, Braga/Guimarães, Portugal; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal Alastair Campbell Ritchie Department of Mechanical, Materials and Manufacturing Engineering, University of Nottingham, Nottingham, United Kingdom Shelly E. Sakiyama-Elbert Department of Biomedical Engineering, The University of Texas at Austin, Austin, TX, United States Karim Salhadar The University of Texas at Austin, Austin, TX, United States Antonio J. Salinas Chemical Department of Pharmaceutical Sciences, Faculty of Pharmacy, Universidad Complutense of Madrid, Spain Katja Schenke-Layland Department of Women’s Health, Research Institute for Women’s Health, Eberhard Karls University Tübingen, Tübingen, Germany; Natural and Medical Sciences Institute (NMI) at the University of Tübingen, Reutlingen, Germany; Cluster of Excellence iFIT (EXC 2180) “ImageGuided and Functionally Instructed Tumor Therapies”, Eberhard Karls University Tübingen, Tübingen, Germany; Department of Medicine/Cardiology, Cardiovascular Research Laboratories (CVRL), University of California (UCLA), Los Angeles, CA, United States Frederick J. Schoen Department of Pathology/Brigham and Women’s Hospital and Harvard Medical School, Boston, MA, United States Brittany E. Schutrum Nancy E. and Peter C. Meinig School of Biomedical Engineering, Cornell University, Ithaca, NY, United States

  List of Contributors

Michael V. Sefton Department of Chemical Engineering and Applied Chemistry, Institute of Biomaterials and Biomedical Engineering, University of Toronto, ON, Canada Michael A. Seidman Department of Pathology/Brigham and Women’s Hospital and Harvard Medical School, Boston, MA, United States Darshan S. Shah, M.D. Department of Orthopaedic Surgery, San Antonio Military Medical Center, Ft Sam Houston, TX, United States Heather Sheardown Department of Chemical Engineering, McMaster University, Hamilton, ON, Canada Andrew J. Shoffstall Department of Biomedical Engineering, Case Western Reserve University, Cleveland, OH, United States; Advanced Platform Technology Center, Rehabilitation Research and Development, Louis Stokes Cleveland VA Medical Center, Cleveland, OH, United States Carl G. Simon, Jr. Biosystems Biomaterials Division, National Institute of Standards & Technology, Gaithersburg, MD, United States Josh Simon Spiral Medical Development, Lansdale, PA, United States Kenneth R. Sims Jr. Department of Biomedical Engineering, University of Rochester, Rochester, NY, United States; Translational Biomedical Science Program, University of Rochester, Rochester, NY, United States Steven M. Slack† Benjamin Slavin Translational Tissue Engineering Center, Johns Hopkins School of Medicine, Baltimore, MD, United States; Department of Plastic and Reconstructive Surgery, Johns Hopkins School of Medicine, Baltimore, MD, United States Kirstie Lane Snodderly University of Maryland College Park; NIH/NIBIB Center for Engineering Complex Tissues Patrick S. Stayton Bioengineering, University of Washington, Seattle, WA, United States

†Deceased.

Stephanie D. Steichen DuPont Health Care Solutions, Midland, MI, United States Paul Stoodley Department of Microbial Infection and Immunity, The Ohio State University, Columbus, OH, United States; Departments of Orthopaedics and Microbiology, The Ohio State University, Columbus, OH, United States; Campus Imaging and Microscopy Facility, Office of Research, The Ohio State University, Columbus, OH, United States; National Centre for Advanced Tribology, Mechanical Engineering, University of Southampton, Southampton, United Kingdom W. Benton Swanson Department of Biologic and Materials Science, School of Dentistry, University of Michigan, Ann Arbor, MI, United States Hobey Tam Department of Bioengineering, Rhodes Engineering Research Center, Clemson University, Clemson, SC, United States Aftab Tayab Department of Chemical Engineering, McMaster University, Hamilton, ON, Canada Susan N. Thomas Parker H. Petit Institute for Bioengineering and Bioscience, Georgia Institute of Technology, Atlanta, GA, United States; George W. Woodruff School of Mechanical Engineering, Georgia Institute of Technology, Atlanta, GA, United States Kellen D. Traxel W. M. Keck Biomedical Materials Research Laboratory, Washington State University, Pullman, WA, United States Rocky S. Tuan Institute for Tissue Engineering and Regenerative Medicine, The Chinese University of Hong Kong, Hong Kong, SAR, China Erik I. Tucker Division of Biomedical Engineering, School of Medicine, Oregon Health & Science University, Portland, OR, United States Rei Ukita Department of Biomedical Engineering, Carnegie Mellon University, Pittsburgh, PA, United States

xix

xx   List of Contributors

Austin Veith Department of Biomedical Engineering, University of Texas at Austin, Austin, TX, United States

Michael F. Wolf Medtronic, Corporate Science and Technology, Minneapolis, MN, United States

Sarah E. Vidal Yucha Department of Biomedical Engineering, Tufts University, Medford, MA, United States

Zhicheng Yao Translational Tissue Engineering Center, Johns Hopkins School of Medicine, Baltimore, MD, United States; Institute for NanoBioTechnology, Johns Hopkins University, Baltimore, MD, United States; Department of Materials Science and Engineering, Johns Hopkins University, Baltimore, MD, United States

Christopher Viney School of Engineering, University of California at Merced, Merced, CA, United States Naren Vyavahare Department of Bioengineering, Rhodes Engineering Research Center, Clemson University, Clemson, SC, United States William R. Wagner Departments of Surgery, Bioengineering & Chemical Engineering, McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, United States Min Wang Department of Mechanical Engineering, The University of Hong Kong, Pok Fu Lam, Hong Kong, China Raymond M. Wang Department of Bioengineering, Sanford Consortium of Regenerative Medicine, University of California San Diego, La Jolla, CA, United States Petra Warner Department of Surgery, University of Cincinnati and Shriners Hospitals for Children – Cincinnati, Cincinnati, OH, United States Cuie Wen School of Engineering, RMIT University, Bundoora, VIC, Australia Jennifer L. West Duke University, Durham, NC, United States Matthew A. Whitman Nancy E. and Peter C. Meinig School of Biomedical Engineering, Cornell University, Ithaca, NY, United States Frank Witte Department of Prosthodontics, Geriatric Dentistry and Craniomandibular Disorders, Berlin, Germany

Michael Yaszemski, M.D. Department of Orthopaedic Surgery, Mayo Clinic, Rochester, MN, United States Michael J. Yaszemski Orthopaedic Surgery and Biomedical Engineering, Mayo Clinic College of Medicine, Rochester, MN, United States Lichen Yin Jiangsu Key Laboratory for Carbon-Based Functional Materials and Devices, Institute of Functional Nano and Soft Materials (FUNSOM), the Collaborative Innovation Center of Suzhou Nano Science and Technology, Soochow University, Suzhou, Jiangsu, China Guigen Zhang F Joseph Halcomb III, M.D. Department of Biomedical Engineering, University of Kentucky, Lexington, KY, United States Peng Zhang Department of Chemical Engineering, University of Washington, Seattle, WA, United States Zhiyuan Zhong Biomedical Polymers Laboratory and Jiangsu Key Laboratory of Advanced Functional Polymer Design and Application, College of Chemistry, Chemical Engineering and Materials Science, Soochow University, Suzhou, Jiangsu, China Nicholas P. Ziats, Ph.D. Pathology, Biomedical Engineering & Anatomy Vice Chair of Pathology for Academic Affairs Case Western Reserve University Department of Pathology Cleveland, Ohio

Preface

Biomaterials Science: An Introduction to Materials in Medicine was launched as an educational project to provide a nascent biomaterials community with an authoritative tool for training and education. Conceptual background material and a broad overview of applications were both envisioned as being integral to the textbook. In the late 1980s, biomaterials was in transition from an emerging field to a respected discipline embracing convergence (i.e., the merging of expertise from different “silos,” such as provided by engineers, biomedical scientists, and physicians)—the biomaterials community was in need of an integrated, comprehensive, and definitive educational resource. This rationale for launching the textbook has been validated by the success of the first three editions of Biomaterials Science: An Introduction to Materials in Medicine. With well over 33,000 print copies, numerous online readers through institutional subscriptions, and the thousands who have bought the e-book version, the text has made consequential contributions to the biomaterials education of students and researchers around the world. The previous editions have been widely adopted for classroom use and serve as a reference resource for thousands of biomaterials professionals. Biomaterials Science: An Introduction to Materials in Medicine strives for a balanced view of the biomaterials field. When this project was first launched, monographs available at that time did articulately address biomaterials, but they tended to strongly emphasize the authors’ areas of expertise, while only superficially addressing other important subjects outside of their intellectual sphere. Balanced presentation means appropriate representation of: • hard and soft biomaterials, with coverage of all major material classes both synthetic and biologically derived; •  the common application areas, including orthopedic, cardiovascular, ophthalmologic, dental, and emerging applications; • a balance of fundamental biological and materials science concepts, contemporary medical/clinical concerns, and regulatory/commercial/societal issues that reflect the complex environment in which biomaterials are developed and used; • coverage of the past, present, and future of biomaterials, their applications, and key challenges that lie ahead. In this way, the reader can embrace the broad field, absorb the unifying principles common to all materials in contact with biological systems, and gain a solid appreciation for the special significance of the word biomaterial.

Biomaterials Science: An Introduction to Materials in Medicine, fourth edition, strives for curricular cohesion. By integrating the experience of many leaders in the biomaterials field, we endeavor to present a balanced yet comprehensive perspective on an exciting and rapidly evolving discipline— and present this information in a graphically attractive and readable format, intended to be useful as an educational resource to a broad array of students, teachers, and practicing professionals. With this fourth edition a new set of four editors has taken on the challenge of moving the textbook forward. It is a humbling task to take over editor responsibilities from the four pioneers and visionaries in the field who launched the text and thoughtfully guided what became the quintessential biomaterials textbook through three editions. The new editor team recognized that the book could not physically grow any further than the size of the third edition (as any student who has carried that edition for a semester in a backpack would appreciate). Thus a balanced approach had to be implemented to both grow and cut back. Furthermore, with broad adoption of online access, the opportunity to place some supplementary content online was utilized (e.g., end-of-chapter questions and a further reading section). First and foremost, the new editors wished to retain continuity with previous editions of the text and changed little where revision was not required. Moderate updates were made in many chapters, often keeping existing author teams and enlisting new coauthors. However, where a fresh direction seemed appropriate for a chapter given the state of the science, or the introduction of new chapters was indicated for where the field had expanded, new author teams were recruited. Overall there are 17 new chapters and 43 chapters with a new set of authors. To balance this growth, some topics were consolidated. A change the reader may notice in content is that the editors have elected to reduce the depth of emphasis on tissue engineering and some device topics, believing that there are now excellent texts devoted exclusively to the maturing tissue engineering field. With medical devices, the desire was not only to cover the critical application areas in overview and to include major new advances, but also to recognize that readers will be able to find fieldspecific, in-depth device coverage elsewhere. Finally, particular attention was given to the third section of the book, where guidance is provided that seeks to prepare the reader for the development pathway beyond the demonstration xxi

xxii       Preface

of a proof-of-concept. Here our hope was that the disconnect between the vast biomaterials scientific literature and the reality of what devices are used clinically and how they are implemented can be appreciated. Such understanding is critical if scientific advances are to ultimately reach the target patient population. The founding editors are still very much a part of this textbook. They have provided wise counsel as the new editors developed the plan for the fourth edition, but also authored revisions in several chapters from previous editions. Much of the excellent content and perspective that has made this text so useful to the community has come from the writings of the founding editors. We have sought to preserve and update this valued content. Acknowledgments and thanks are in order. First, the Society for Biomaterials (SFB), sponsor and inspiration for this book, is a model of “multidisciplinary cultural diversity.” Composed of engineers, physicians, scientists, veterinarians, industrialists, inventors, regulators, attorneys, educators, ethicists, and others, the SFB provides the nidus for the intellectually exciting, humanistically relevant, and economically important biomaterials field. As was the case for all of the previous editions, the editors recognize the importance of the SFB by donating all royalties from sales of this volume to the Society, to directly support education and career development related to biomaterials. For information on the SFB, visit the SFB website at http://www.biomaterials.org/. We offer special thanks to those who have generously invested time, energy, experience, and intelligence to author the chapters of this textbook. The over 100 scientists, physicians, engineers, and industry leaders who contributed their expertise and perspectives are clearly the backbone of

this work, and they deserve high praise—their efforts will strongly impact the education of the next generation of biomaterials scientists. It is only with such a distinguished group of authors that this text can provide the needed balance, scope, and perspective. We also pay respect and homage to biomaterials pioneers who have contributed to this or previous editions, but have since passed on; their contributions and collegiality are remembered and will be missed. The organizational skills, experience, and encouragement of the staff at Elsevier Publishers have led this fourth edition from a broad and complex challenge to a valuable volume— a tangible resource for the community. Thank you, Elsevier, for this contribution to the field of biomaterials. The biomaterials field, since its inception in the 1950s, has been ripe with opportunity, intellectual stimulation, compassion, creativity, and rich collaboration. In this field we look to the horizons where the new ideas from science, technology, and medicine arise. Importantly, we strive to improve the survival and quality of life for billions through biomaterials-based medical devices and treatments. We, the new editors, together with the founding editors hope that the biomaterials textbook you are now reading will stimulate you as much as it has us. William R. Wagner Shelly E. Sakiyama-Elbert Guigen Zhang Michael J. Yaszemski Founding Editors: Buddy D. Ratner Allan S. Hoffman Frederick J. Schoen Jack E. Lemons March, 2020

How to Use this Book

Biomaterials Science: An Introduction to Materials in ­Medicine, fourth edition, was conceived as a learning tool to “compatibilize” through common language and fundamental principles, a number of independent communities (basic sciences, engineering, medicine, dentistry, industry, regulatory, legal, etc.). Although the book has 98 chapters, 4 a­ppendices, there is a logic of organization and curriculum that should make the book straightforward to use in an academic course or as a reference work. A guiding principle in assembling this multiauthor, multidisciplinary textbook is that fundamental and translational progress in the field of biomaterials necessitates integration of concepts and tools from the full spectrum of the physical sciences, engineering, clinical medicine, biology, and life sciences. Indeed, the discipline of biomaterials utilizes a convergence of multidisciplinary elements to enable the development of specific diagnostic or therapeutic devices—i.e., using biomaterials science and technology to create and implement real medical devices and other products that solve clinical problems and improve patient outcomes. Nevertheless, the editors believe (and the book has been assembled so) that a physician should be able to pick up the textbook and glean a baseline knowledge of the science, engineering, and commercialization aspects of biomaterials. A chemist could use this book to appreciate the biology behind biomaterials, the physiology associated with medical devices, and the applications in medicine. An engineer hired by a medical device company might learn the basic science underlying the technological development and details on medical applications. Similarly, for other disciplines that interface with biomaterials, this book can guide the reader through diverse but related topics that are generally not found in one volume. The textbook has well over 100 authors. The field of biomaterials is so diverse in subject matter that a guiding principle has been that no one author can write it all—let us use the experience and wisdom of acknowledged masters of each subject to communicate the best information to

the reader. But, to prevent this book from appearing to be simply a collection of review papers, considerable editorial effort has gone into ensuring logical curricular organization, continuity of ideas, and extensive cross-referencing between chapters. Biomaterials Science: An Introduction to Materials in Medicine, fourth edition, is divided into three parts: Materials Science and Engineering, Biology and Medicine, and The Medical Product Life Cycle. Sections then serve to subdivide each of these three parts (for example, under the part called “Materials Science and Engineering” there are sections on properties of materials and classes of materials used in medicine). And finally, within these sections can be found chapters, for example, on the major types of materials that are used in medicine (hydrogels, polyurethanes, titanium alloys, etc.). Each section begins with an introduction by one of the editors that will guide the reader through the chapters, giving cohesion to the sections and highlighting key issues that are worthy of attention. Finally, there are appendices that tabulate useful data and information. For most chapters, exercises are provided for classroom use and for self-testing. These can be accessed via the companion website at https://www.elsevier.com/booksand-journals/book-companion/9780128161371. This site makes it possible to update problems and add new ones, and provides other resource materials, including a full artwork catalog and downloadable images used in the text. For instructors, solutions to the end-of-chapter exercises are provided on an Instructor Site at http://textbooks.elsevier.c om/web/Manuals.aspx?isbn=9780128161371. We hope that the textbook organization, the extensive editorial effort, and the expertly authored chapters will serve their intended purpose—to guide the reader into and through this complex field of biomaterials science. The editors always appreciate feedback and commentary—contact information is provided for them. And now it is time to delve into the rich world of biomaterials science.

xxiii

SE C TI ON 1 .1

Overview of Biomaterials

1.1.1

Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor BUDDY D. RATNER 1 , ALLAN S. HOFFMAN 2 , FREDERICK J. SCHOEN 3 , JACK E. LEMONS 4 , WILLIAM R. WAGNER 5 , SHELLY E. SAKIYAMA-ELBERT 6 , GUIGEN ZHANG 7 , MICHAEL J. YASZEMSKI 8 1Bioengineering

and Chemical Engineering, Director of University of Washington Engineered Biomaterials (UWEB), Seattle, WA, United States 2Bioengineering 3Department

and Chemical Engineering, University of Washington, Seattle, WA, United States

of Pathology, Brigham and Women’s Hospital and Harvard Medical School, Boston, MA,

United States 4Schools

of Dentistry, Medicine and Engineering, University of Alabama at Birmingham, Birmingham, AL, United States 5Departments

of Surgery, Bioengineering & Chemical Engineering, McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, United States 6Department

of Biomedical Engineering, The University of Texas at Austin, Austin, TX, United States

7F

Joseph Halcomb III, M.D. Department of Biomedical Engineering, University of Kentucky, Lexington, KY, United States 8Orthopaedic

Surgery and Biomedical Engineering, Mayo Clinic College of Medicine, Rochester, MN,

United States

Biomaterials and Biomaterials Science Biomaterials Science: An Introduction to Materials in Medicine, fourth edition, addresses the design, fabrication, testing, applications, and performance as well as nontechnical considerations integral to the translation of synthetic and natural materials that are used in a wide variety of implants, devices, and process equipment that contact biological systems. These materials are referred to as biomaterials. The compelling, human side to biomaterials is that millions of lives are saved, and the quality of life is improved for millions more.

The field of biomaterials is some 70–80 years old at the time of publication of this fourth edition. It significantly impacts human health, the economy, and many scientific fields. Biomaterials and the medical devices comprised of them are now commonly used as prostheses in cardiovascular, orthopedic, dental, ophthalmological, and reconstructive surgery, and in other interventions such as surgical sutures, bioadhesives, and controlled drug release devices. The compelling, human side to biomaterials is that millions of lives have been saved, and the quality of life improved for millions more, based on devices fabricated from biomaterials. The biomaterials field has seen accelerating growth since the 3

4 SEC T I O N 1 .1     Overview of Biomaterials

• Figure 1.1.1.1  The path from the basic science of biomaterials, to a medical device, to clinical application.

first medical devices that were based on accepted medical and scientific principles made their way into human usage in the late 1940s and early 1950s. And the growth of the field is ensured, with the aging population, the increasing standard of living in developing countries, and the growing ability to address previously untreatable medical conditions. Biomaterials science addresses both therapeutics and diagnostics. It encompasses basic sciences (biology, chemistry, physics) and engineering and medicine. The translation of biomaterials science to clinically important medical devices is dependent on: (1) sound engineering design; (2) testing in vitro, in animals and in humans; (3) clinical realities; and (4) the involvement of industry permitting product development and commercialization. Fig. 1.1.1.1 schematically illustrates the path from scientific development to the clinic. Biomaterials science, in its modern incarnation, is an example of the emerging convergence paradigm that pushes multidisciplinary collaboration among experts and multidisciplinary integration of concepts and practice (Sharp and Langer, 2011). Not only biomaterials, but also bioinformatics, synthetic biology, computational biology, nanobiology, systems biology, molecular biology, and other forefront fields depend on convergence for their continued progress. This textbook aims to introduce these diverse multidisciplinary elements, particularly focusing on interrelationships rather than disciplinary boundaries, to systematize the biomaterials subject into a cohesive curriculum—a true convergence. The title of this textbook, Biomaterials Science: An Introduction to Materials in Medicine, is accurate and descriptive. The intent of this work is: (1) to focus on the scientific and engineering fundamentals behind biomaterials and their applications; (2) to provide sufficient background knowledge to guide the reader to a clear understanding and appreciation of the clinical context where biomaterials are applied; and (3) to highlight the opportunities and challenges in the field. Every chapter in this text can serve as a portal to an extensive contemporary literature that expands on the basic ideas presented here. The magnitude of the biomaterials endeavor, its broadly integrated multidisciplinary scope, and examples of biomaterials applications will be revealed in this introductory chapter and throughout the book. The common thread in biomaterials is the physical and chemical interactions between complex biological systems and synthetic or modified natural materials.

Although biomaterials are primarily used for medical applications (the focus of this text), they are also used to grow cells in culture, to assay for blood proteins in the clinical laboratory, in processing equipment for biotechnological applications, for implants to regulate fertility in cattle, in diagnostic gene arrays, in the aquaculture of oysters, and for investigational cell–silicon “neuronal computers.” How do we reconcile these diverse uses of materials into one field? The common thread is the physical and chemical interactions between complex biological systems and synthetic materials or modified natural materials. In medical applications, biomaterials are rarely used as isolated materials, but are more commonly integrated into devices or implants, and complex devices may use multiple biomaterials, often selected from several classes (e.g., metal and polymer). Chemically pure titanium can be called a biomaterial, but shaped (machined) titanium in conjunction with ultrahigh molecular weight polyethylene becomes the device, a hip prosthesis. Although this is a text on biomaterials, it will quickly become apparent that the subject cannot be explored without also considering biomedical devices and the biological response to them. Indeed, both the material and the device impact the recipient (patient) and, as we will see, the patient’s host tissue impacts the device. These interactions can lead to device success or, where there is inappropriate choice of biomaterials or poor device design, device failure. Moreover, specific patient characteristics may influence the propensity to failure (e.g., obesity increasing the likelihood of fracture or excessive wear of a hip joint prosthesis, or clotting of a mechanical heart valve in a patient with a genetic mutation that causes hyper-coagulability). Furthermore, a biomaterial must always be considered in the context of its final fabricated, sterilized form. For example, when a polyurethane elastomer is cast from a solvent onto a mold to form the pump bladder of a heart assist device, it can elicit different blood reactions compared to when injection molding is used to form the same device. A hemodialysis system serving as an artificial kidney requires materials that must function in contact with a patient’s blood, and also exhibit appropriate membrane permeability and mass transport characteristics. Much fabrication technology is applied to convert the biomaterials of the hemodialysis system (polysulfone, silicone rubber, polyethylene) into the complex apparatus that is used for blood purification. Due to space limitations and the materials focus of this work, many aspects of medical device design are not addressed in this book. Consider the example of the hemodialysis system. This textbook will overview membrane materials and

CHAPTER 1.1.1   Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor

their biocompatibility; there will be little coverage of mass transport through membranes, the burst strength of membranes, dialysate water purification, pumps, flow systems, and monitoring electronics. Other books and articles cover these topics in detail, and chapter authors provide references useful to learn more about topics not explicitly covered. 

5

the foreign-body reaction (Chapter 2.2.2). This leads us to consider a widely used definition of biocompatibility: “Biocompatibility” is the ability of a material to perform with an appropriate host response in a specific application. WILLIAMS (1987).

“Biocompatibility” is the ability of a material to perform with an appropriate host response in a specific application.

Key Definitions The words “biomaterial” and “biocompatibility” have already been used in this introduction without formal definition. A few definitions and descriptions are in order, and will be expanded upon in this and subsequent chapters. A definition of “biomaterial” endorsed by a consensus of experts in the field is: A biomaterial is a nonviable material used in a medical device, intended to interact with biological systems. WILLIAMS (1987).

A biomaterial is a nonviable material used in a medical device, intended to interact with biological systems.

Although biomaterials are most often applied to meet a therapeutic or diagnostic medical need, if the word “medical” is removed, this definition becomes broader and can encompass the wide range of applications already suggested. If the word “nonviable” is removed, the definition becomes even more general, and can address many new tissue-engineering and hybrid artificial organ applications where living cells are used. “Biomaterials science” is the study (from the physical and/or biological perspective) of materials with special reference to their interaction with the biological environment. Traditionally, emphasis in the biomaterials field has been on synthesis, characterization, and the host–material interactions biology. Yet, most biomaterials (which meet the special criteria of biocompatibility—see Chapters 2.3.2 and 2.3.4) induce a nonspecific biological reaction that we refer to as

Examples of “appropriate host responses” include resistance to blood clotting, resistance to bacterial colonization, and normal, uncomplicated healing. Examples of “specific applications” include a hemodialysis membrane, a urinary catheter, or a hip joint replacement prosthesis. Note that the hemodialysis membrane might be in contact with the patient’s blood for 5 h (and outside the body), the catheter may be inserted for a week (inside the body, and designed to be easily removed), and the hip joint may be in place for the life of the patient (deeply implanted and meant to be long-term). This general concept of biocompatibility has been extended to tissue engineering, in which in vitro and in vivo processes are harnessed by careful selection of cells, materials, and metabolic and biomechanical conditions to regenerate functional tissues. Ideas central to biocompatibility are elaborated upon in Ratner (2011) and Chapter 2.3.2. In the discussion of these definitions, we are introduced to considerations that set biomaterials apart from most materials explored in materials science. Table 1.1.1.1 lists a few applications for biomaterials in the body. It includes many classes of materials used as biomaterials. Note that metals, ceramics, polymers, glasses, carbons, and natural and composite materials are listed. Such materials are used as molded or machined parts, coatings, fibers, films, membranes, foams, fabrics, and particulates. Table 1.1.1.1 also gives estimates of the specific device global market size and, where available, an estimate of the number of medical devices utilized annually. The human impact, and the size of the commercial market for biomaterials and the broad array of medical devices, is impressive (Tables 1.1.1.1 and 1.1.1.2). 

TABLE 1.1.1.1    Key Applications of Synthetic Materials and Modified Natural Materials in Medicine

Biomaterials Used

Number/Year—Global (or Global Market in US$)

Joint replacements (hip, knee, and shoulder)

Titanium, CoCr, polyethylene, alumina, zirconia

4,000,000 ($16B)

Trauma fixation devices (plates, screws, pins, and rods)

Titanium, stainless steel, CoCr, polyether ether ketone, poly(lactic acid) (PLA)

1,500,000 ($5.5B)

Spine disks and fusion hardware

Nitinol, titanium, polyether ether ketone, stainless steel

1,100,000 ($8.5B)

Bone defect repair

Calcium phosphates, human bone products

($4.5B)

Bone cement (fixation)

Polymethyl methacrylate (PMMA), glass polyalkenoate (ionomer), calcium phosphate cements

($1.1B)

Cartilage, tendon, or ligament repair and replacement

Decellularized porcine tissue, poly(lactide) and metallic fixation devices, collagen, hyaluronic acid lubricants

($8.6B)

Dental implant-tooth fixation

Titanium, zirconium

10,000,000 ($4B)

Application Skeletal System

(Continued)

6 SEC T I O N 1 .1     Overview of Biomaterials

TABLE 1.1.1.1    Key Applications of Synthetic Materials and Modified Natural Materials in Medicine—cont’d

Application

Biomaterials Used

Number/Year—Global (or Global Market in US$)

Cardiovascular System Vascular grafts, patches, and endovascular devices (stent grafts)

Dacron, expanded poly(tetrafluoroethylene), Nitinol, CoCr, stainless steel, fixed tissue

($2.5B)

Heart valves: mechanical and bioprosthetic (transcatheter and traditional)

Dacron, carbon, CoCr, fixed bovine and porcine tissue, stainless steel, Nitinol

600,000 ($5.5B)

Pacemakers

Titanium, polyurethane

1,000,000 ($6.5B)

Implantable defibrillators

Titanium, polyurethane

300,000 ($9.0B)

Stents: coronary, peripheral vasculature, and nonvascular

Stainless steel, Nitinol, CoCr, Pt, tantalum, Mg alloys, poly(styrene-b-isobutylene-b-styrene), poly(n-butyl methacrylate), polyethylene-co-vinyl acetate, phosphoryl choline containing block copolymers, poly(lactic-coglycolic acid), PLA

5,000,000 ($10.6B)

Catheters: cardiovascular, urologic, and others

Polytetrafluoroethylene (PTFE), poly(vinyl chloride), silicone, polyurethane

($28B)

Cardiac assist devices (acute and chronic)

Titanium alloy, polycarbonate, PTFE, poly(ethylene terephthalate), stainless steel

($1.7B)

Hemodialysis

Polysulfone, modified cellulose, polyacrylonitrile, polycarbonate, silicone, polyvinylchloride Polymethylpentene, polypropylene, polysiloxane, poly(vinyl chloride), polycarbonate

2,000,000 patients ($12B)

Collagen, cadaver skin, alginate, polyurethane, carboxymethylcellulose, nylon, silicone

($1.3B)

Contact lens

PMMA, polyhydroxyethylmethacrylate (PHEMA), polyvinyl alcohol, polyvinyl pyrrolidone, silicone (polydimethyl siloxane [PDMS])

($7.5B)

Intraocular lens

PMMA, PDMS, polyacrylate-PMMA, PHEMA

25,000,000 ($4.5B)

Glaucoma drains

Silicone, polypropylene, cross-linked collagen, stainless steel

($500M)

Cochlear prostheses

Platinum, platinum–iridium, PDMS, titanium, aluminum oxide

45,000 ($2.7B)

Breast implants

PDMS

3,600,000 ($1.2B)

Hernia and body wall repair meshes

Polypropylene, polyester, expanded PTFE, decellularized porcine/bovine tissue

($4.2B)

Sutures

Silk, nylon, poly(glycolic acid), PLA, polydioxanone, polyester copolymers, polypropylene, PTFE, processed bovine tissue

($3.9B)

Blood bags

Poly(vinyl chloride)

($170M)

Ear tubes (tympanostomy)

Silicone, PTFE

1,500,000 ($70M)

Intrauterine device

Polyethylene, copper, stainless steel, PDMS

168,000,000 ($2.9B)

Organs

Blood oxygenator Skin substitute (chronic wounds, burns)

($300M)

Ophthalmologic

Other

Data compiled from multiple sources—these numbers should be considered rough estimates that are changing with growing markets and new technologies. Where only US numbers were available, world usage was estimated at approximately 2.5× US. B, Billion; M, million.

CHAPTER 1.1.1   Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor

7

TABLE   The Medical Device Global Market by Segment With Projected Compound Annual Growth Rate (CAGR) 1.1.1.2  ($ Millions)

Segments

2016

2017

2022

CAGR 2017–2022 (%)

Drug delivery devices

200,072

207,814

243,367

3.2

Urology and renal

75,378

82,668

109,003

5.7

In vitro diagnostics

66,143

72,816

99,357

6.4

Orthopedics and spine

65,756

72,086

99,559

6.7

Imaging devices

41,194

45,816

64,282

7.0

Cardiovascular devices

25,384

29,658

45,260

8.8

Endoscopy

9,573

10,372

13,693

5.7

Total

483,500

521,230

674,521

5.3

Source: BCC Research.

The Expansion of the Biomaterials Field Biomaterials research and development have been stimulated and guided by advances in cell and molecular biology, pathology, clinical medicine and dentistry, chemistry, materials science, and engineering. The biomaterials community has been the major contributor to the understanding of the interactions of materials with the physiological environment (often referred to as the biointerface). The development of biomaterials for medical and dental applications has evolved with time, as new concepts and understandings are applied to offer a broadening repertoire of choices to meet device design objectives (Fig. 1.1.1.2). Early applications of biomaterials sought to achieve a suitable combination of functional properties to adequately meet the design needs for the medical device under development.

Generally, this would involve the layering of biocompatibility concerns from host–material interactions on top of those more readily understood physical and chemical requirements. For instance, for a mechanical cardiac valve, materials could be selected and integrated to provide the functional response in an altering pressure flow field, resistance to cyclic mechanical wear, and suturability. In these early applications, industrial materials were typically taken off the shelf, i.e. “medical grade” biomaterials were not yet available. Nevertheless from the array of industrially available materials that might meet these requirements, considerations of blood and tissue compatibility would be included. Pioneers in the device field effectively applied empirical approaches to arrive at materials that could meet both the traditional (nonbiological) design requirements and

Industrial material adaptation Select off-theshelf material for desired device properties

Design of passive materials Design material to meet desired physical and chemical properties for device

Design of bioactive & degradable materials Design material for a particular biological response, including drug delivery and temporary scaffolding

Self-assembling materials Micro- and nano-scale assembly to integrate functionalities, allow in situ assembly or disassembly; controlled particulate formation

Constructive remodeling materials Designed to facilitate a tissue response that results in functional tissue and a healing response not typical with synthetic materials

• Figure 1.1.1.2  The growing palette of biomaterials. Generally moving with time from the 1940s adapta-

tion of industrially available materials for early medical devices to the present, the breadth of described biomaterials continues to grow. In device development a biomaterial may be selected to leverage recent progress. However, it is important to note that major advances in the medical device field continue to be made with materials that could be considered first generation. The growing palette provides the design engineer with more tools to optimize device functionality in concert with other concerns such as manufacturability, regulatory burden, and economic considerations.

8 SEC T I O N 1 .1     Overview of Biomaterials

exhibit adequate levels of biocompatibility. Materials would generally be selected because they were tolerable (i.e., they elicited minimal response from the host tissues), and this would be consistent with biocompatibility for many applications (see Chapter 2.3.2). While the understanding of biomaterials science has evolved substantially from these early days, it is important to recognize that industrially repurposed materials continue to be utilized in many widely used medical devices today, including poly(tetrafluoroethylene) and poly(ethylene terephthalate), from which virtually all synthetic vascular grafts are made, stainless steel, cobalt– chromium alloys and titanium alloys, from which many orthopedic devices are constructed, and the polyurethanes and polysiloxanes that are utilized in a broad array of catheters and medical tubing. Furthermore, industrially adapted materials continue to be the biomaterials of choice for many revolutionary new devices introduced in recent years such as many of the components related to neurostimulatory devices and structural heart repair. In this early period it would also occasionally be noted where an off-the-shelf material or class of materials might not fully achieve the target device design objectives and novel materials would be designed or refined specifically for a biomedical purpose. As highlighted in Chapter 1.1.2, biomaterials scientists would, for instance, create polyurethanes with segments selected for the purpose of improving blood biocompatibility. Hydrogels would be synthesized for soft tissue applications. Pyrolytic carbon, originally developed in the 1960s as a coating material for nuclear fuel particles, was studied and tuned for what is now in wide use in modified compositions to coat components of mechanical cardiac valves. These designed materials broadened the biomaterials palette, but the materials were still designed to be passive in achieving biocompatibility. As with early adapted industrial materials, these types of biomaterials continue to play an important part in device design and active research continues to seek to develop materials that are better suited for specific device applications while still targeting a passive, bioinert posture. For instance, efforts to find more degradation-resistant polymers for challenging device applications are ongoing (Chapter 2.4.2). As knowledge of biological interactions with materials evolved, new types of biomaterials were developed with the intention of eliciting a controlled reaction with the tissues they contacted to induce a desired therapeutic effect. In the 1980s, these bioactive materials were in clinical use in orthopedic and dental surgeries as various compositions of bioactive glasses and ceramics (Hench and Pollak, 2002, Chapter 1.3.4), in controlled localized drug release applications such as the Norplant hormone-loaded contraceptive formulation (Meckstroth and Darney, 2001), and in the attachment of the anticoagulant heparin to the surfaces of membrane oxygenators with various modification strategies (Chapter 1.4.3B). Vascular stents have also been profoundly impacted by the implementation of a bioactive approach, with the application of polymer coatings that release antiproliferative agents and markedly reduce a major failure mechanism of tissue overgrowth and vessel occlusion (Chapter 2.5.2B).

Bioactive biomaterial development has also included the synthesis of resorbable polymeric biomaterials, with rates of degradation that could be tailored to the requirements of a desired application (Chapter 1.3.2F). Thus the discrete interface between the implant site and the host tissue could be eliminated in the long term, because the foreign material would ultimately be degraded to soluble, nontoxic products by the host. Many groups continue to develop new biodegradable polymers designed with defined objectives in strength, flexibility, a chemical composition conducive to tissue development, and a degradation rate consistent with the specific application. Degradable materials have been integral to the tissue-engineering paradigm where a scaffold, alone or in combination with cells and drugs, may provide for the generation of functional tissue. This paradigm as applied to the engineering of a cardiac valve is presented in Fig. 1.1.1.3 and is covered in the chapters of Section 2.6. Tissue-engineering approaches often leverage degradable biomaterials scaffolds, drug-releasing biomaterials, and in some cases utilize specific cell receptor–ligand interactions or enzymatic degradability to build bioactivity into the biomaterials scaffold (Chapters 1.4.4 and 1.4.5). A characteristic of bioactive biomaterials development over the past several decades has been the leverage of fundamental knowledge from molecular biology. As this knowledge base has grown, biomaterials scientists and engineers have translated the understanding of biomolecular interactions to engineer biological interactions with designed materials. An early example of this was the application of knowledge of the adhesion peptide sequences from proteins such as fibronectin (e.g., Arg-Glu-Asp-Val) to engineer peptide-modified surfaces that would support specific types of cell adhesion (Hubbell et  al., 1991). Polymeric materials with other novel properties such as shape-memory and programmable and interactive surfaces that control the cellular microenvironment are areas of development (Chapter 1.3.2G). In addition to having implications for medical applications, such engineered smart biomaterials systems have been used to advance our understanding of molecular biology principles, for instance, in elucidating the roles of substrate stiffness, ligand density, and three-dimensional culture in mammalian cell behavior. The need for maximally effective pharmacologic dosing regimens and minimization of systemic toxicities has stimulated development of innovative particulate systems for targeted drug delivery and gene therapy (Chapter 1.3.8). Such systems may also provide the basis for targeted imaging or the combination of targeted imaging and therapeutic delivery, representing the growing field of theranostics. This focus area is experiencing a great deal of research attention at present by the biomaterials community and many of the approaches involve the production of nano- and microscale particulates using the principles of self-assembling materials. Several factors are driving this effort to design better biomaterials-based approaches: advances in protein and nucleic acid-based drugs (which cannot be taken in classical pill form, have high cost, and are labile); a better understanding of transport mechanisms systemically, within

CHAPTER 1.1.1   Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor

9

• Figure 1.1.1.3  An example of various tissue-engineering paradigms applied to a cardiac valve. Current

approaches for many tissues, including heart valves, may involve scaffolds, cells, and drugs. These components may be used alone or in combination. For each component, different material types, drugs, or cells may be used. Again, alone or in combination, bioreactors may be employed to allow some level of tissue construct maturation, or the body may serve as the bioreactor. (From Jana, S., Tefft, B.J., Spoon, D.B., Simari, R.D., 2014. Scaffolds for tissue engineering of cardiac valves. Acta Biomater. 10, 2877–2893.)

specific tissues or tumors, and intracellularly; and an increasing ability to create precise structures at macromolecular scales through controlled polymerization techniques and specific orthogonal reaction schemes. Self-assembled biomaterials have allowed design at the nanoscale to protect drug payloads that are released with an appropriate signal (e.g., pH, radiation, intracellular environmental cues) upon delivery at the desired site through the surface presentation of targeting moieties and molecules that reduce fouling and improve pharmacokinetics. Self-assembled biomaterials have also found broad interest as injectable networks for the creation of scaffolds and depots in regenerative medicine, drug delivery, and immunoengineering (Sahoo et  al., 2018). Here, by switching out functional groups on larger self-assembling molecules, the specific bioactivity and mechanical properties of an assembled network may be designed. Interest is also growing in the structural and functional tailoring of self-assembled two-dimensional organic biomaterials to open new opportunities by controlling biomaterials at the nanoscale for highly specific, spatially orchestrated biological interactions (Zhang et al., 2019).

A final category in the growing repertoire of biomaterials is what could be termed constructive remodeling materials. Such materials have been designed or processed to facilitate a healing response that does not follow the classic foreign-body response found with most synthetic tissues, but is characterized by remodeling of the tissue with minimal scarring. This response has been observed and well documented in biomaterials derived from animal-based tissues that have been decellularized to reduce immunogenicity, but which have not been chemically cross-linked. Chemical cross-linking is a critical part of maintaining structural viability for many biomaterials (e.g., bioprosthetic cardiac valves from bovine or porcine source tissues), but in this approach a tissue, such as porcine bladder or dermis, is meant to be degraded and remodeled at the implantation site, and in this remodeling process, be replaced with host tissue that is functional as opposed to fibrotic. These types of materials are described in Chapter 1.3.6, and the role of the immune response with these types of natural materials versus most synthetic materials or crosslinked natural materials is addressed in Chapter 2.2.2. As the biomaterials community better understands the mechanisms

10 SEC T I O N 1 .1     Overview of Biomaterials

• Figure 1.1.1.4  Prosthetic heart valves. Left: A bileaflet tilting disk mechanical heart valve (St. Jude Medical

Inc., St. Paul, MN). Right: A bioprosthetic (xenograft) tissue heart valve (Hancock valve, Medtronic Inc., MN).

by which constructive remodeling can be achieved, this knowledge is being applied in materials designed to orchestrate specific interactions with the immune system to moderate the host response. More broadly, biomaterials are being developed in the area of immunoengineering (Chapter 2.5.10) to impact the immune system in applications related to immunization, cancer, infection, and autoimmune diseases. 

Examples of Today’s Biomaterials Applications Five examples of biomaterials applications now follow to illustrate important ideas. The specific devices discussed were chosen because they are widely used in humans with good success. However, key limitations with these biomaterial devices are also highlighted. Each of these examples is also discussed in detail in later chapters in this textbook.

Heart Valve Prostheses Diseases and degeneration of the heart valves often make surgical repair or replacement necessary. The natural heart valve opens and closes over 40 million times per year, and can require replacement due to disease or wear. Approximately 4,500,000 replacement valves are implanted each year worldwide, because of acquired damage to the natural valve and congenital heart anomalies. There are many types of heart valve prostheses, and they are fabricated from carbons, metals, elastomers, plastics, fabrics, and animal or human tissues chemically pretreated to reduce their immunologic reactivity, and to enhance durability. Fig. 1.1.1.4 shows a bileaflet tilting disk mechanical heart valve and a bioprosthetic (porcine xenograft) tissue heart valve, two of the most widely used designs. Generally, as soon as the valve is implanted, cardiac function is restored to near normal levels, and the patient shows rapid improvement. In spite of the overall success seen with replacement heart valves, there are problems, many of them specific to a certain type of valve; they include



Figure 1.1.1.5 A hip prosthesis. Microplasty titanium alloy femoral stem, Biolox alumina–zirconia ceramic femoral head, and ultrahigh molecular weight polyethylene acetabular cup infused with vitamin E antioxidant. (Image courtesy of Biomet, Inc.)

induction of blood clots, predominantly with mechanical valves (sometimes shed into the bloodstream as emboli and creating an ongoing risk for stroke, thus necessitating long-term therapy with potentially dangerous anticoagulant drugs), degeneration of valve tissue leaflets, mechanical failure, and infection. This biomaterial application continues to be an active area of innovation, most recently with growing clinical use of valve designs implanted through a catheter and in the advancement of tissueengineering approaches for valve replacement (Zhang et  al., 2019). Heart valve substitutes are discussed in Chapter 2.5.3A. 

Total Hip Replacement Prostheses The human hip joint is subjected to high levels of mechanical stress and receives considerable abuse in the course of normal and extraordinary activity. It is not surprising that after 50 or more years of cyclic mechanical stress or because of degenerative or rheumatoid disease, the natural joint wears out, leading to loss of mobility and sometimes confinement to a wheelchair. Hip joint prostheses are fabricated from a variety of materials, including titanium, stainless steel, special high-strength alloys, ceramics, composites, and ultrahigh molecular weight polyethylene. Replacement hip joints (Fig. 1.1.1.5) are implanted

CHAPTER 1.1.1   Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor

11

• Figure 1.1.1.6  Schematic images of early dental root form implants and a photograph of several designs used in clinical practice.

• Figure 1.1.1.7  Two styles of multipiece intraocular lenses.

in more than 300,000 humans each year in the United States alone. With some types of replacement hip joints and surgical procedures that use a polymeric cement, ambulatory function is restored within days after surgery. For other types, a healing-in period is required for integration between bone and the implant before the joint can bear the full weight of the body. In most cases, good function is restored. Even athletic activities are possible, although those activities that subject the repaired joint to high stress are generally not advisable. After 10–15 years, many of these implants fail by loosening, which usually necessitates another operation (a revision procedure). Metal-on-metal implants also experience problems of corrosion and adverse responses to released metal ions. Artificial hip joint prostheses are discussed in Chapter 2.5.4. 

Dental Implants The development of root form designs of titanium implants (Fig. 1.1.1.6) by Per-Ingvar Brånemark revolutionized dental implantology (Carlsson et al., 1986). These devices form

an implanted artificial tooth anchor upon which a crown is affixed and are implanted in 5,000,000 people each year in the United States alone, according to the American Dental Association. A special requirement of a material in this application is the ability to form a tight seal against bacterial invasion where the implant traverses the gingiva (gum). Other issues are associated with the rapidly growing junctional epithelium inhibiting regrowth of the slower growing bone. Also, in normal physiology, the tooth is connected to the jaw by the periodontal ligament and is not directly attached to the jawbone. One of the primary advantages originally cited for the titanium implant was its osseous integration with the bone of the jaw. In recent years, however, this attachment has been more accurately described as a tight apposition or mechanical fit, and not true bonding. Loss of tissue support leading to loosening remains an occasional problem, along with infection and issues associated with the mechanical properties of unalloyed titanium that is subjected to long-term cyclic loading. Dental implants are discussed in Chapter 2.5.5. 

Intraocular Lenses Implants to replace lenses in the eye that have clouded due to cataracts are called intraocular lenses (IOLs). They have been fabricated from a variety of transparent materials, including poly(methyl methacrylate), silicone elastomer, soft acrylic polymers, and hydrogels (Fig. 1.1.1.7). By the age of 75, more than 50% of the population suffers from cataracts severe enough to warrant IOL implantation. This now translates to an estimated 25 million implants worldwide as cataract treatment is expanding rapidly in healthcare systems with growing economies and aging populations. Good vision is generally restored almost immediately after the lens is inserted, and the success rate with this device is high. IOL surgical procedures are well developed, and implantation is most often performed on an outpatient

12 SEC T I O N 1 .1     Overview of Biomaterials



Figure 1.1.1.8  A commonly utilized chronic ventricular assist device (VAD). (A) Continuous flow pump with associated inflow/outflow grafts and electrical drive line (Heartmate II device). (B) Schematic of VAD implanted as a left ventricular assist device with associated external power source. (C) A catheter-based rotary blood pump from Impella for acute cardiac support. The tip of the device is placed in the ventricle where openings allow blood flow into the lumen of the device. Blood is pumped with a microaxial rotary motor and expelled through catheter openings in the region of the device that rests in the aorta. (Images A and B obtained with permission from Thoratec Corporation. Image C obtained from https://mms.busines swire.com/media/20190831005001/en/740858/5/SmartAssist_Press_release_image.jpg?download=1.)

basis. Observations of implanted lenses through the cornea using a microscope to directly study the implants show that inflammatory cells such as macrophages migrate to the surface of the lenses after implantation. Thus the conventional healing pathway is seen with these devices, similar to that observed with materials implanted in other sites in the body. Outgrowth of cells onto the IOL from the posterior lens capsule, stimulated by the presence of the IOL, can cloud vision, and this is a significant complication. IOLs are discussed in Chapter 2.5.6.

Ventricular Assist Devices Nearly 5,000,000 Americans are living with seriously failing hearts (congestive heart failure), and 300,000 individuals will die each year from this disease. According to the American Heart Association, 50,000–100,000 of these individuals might benefit from cardiac transplantation or mechanical circulatory support. Since the available pool of donor hearts for transplantation is only ∼3500 per year (United States), effective and safe mechanical cardiac assist or replacement seems like a desirable option. Ventricular assist devices (VADs) have evolved from a daring experimental concept, the mechanical total heart, to a life-prolonging tool (see Chapter 2.5.2A). A number of devices have received regulatory approval and designs have evolved from bulky pulsatile pumps with chambers and opposing valves (mimicking the human heart) to much smaller rotary devices compatible with a broader array of patients. VADs are now used in multiple ways: to maintain (“bridge”) a patient with a failing heart while awaiting a donor organ, as a permanent source of support for patients

not destined for a heart transplant, and as temporary support where cardiac function is at risk due to a procedure, or when a weakened heart is expected to recover in the short term and device support can be withdrawn. A commonly utilized rotary VAD for extended cardiac support and a catheter-based device used for temporary support are illustrated in Fig. 1.1.1.8. Recipients of VADs designed for chronic support can have considerable mobility and freedom, with most being discharged from the hospital setting. Despite patients being supported in some cases for several years, there remains an elevated risk for device-related infection (particularly in the region where the control and power line crosses the skin) and stroke related to the embolization of clots formed within the device. Furthermore, although VAD therapy that sends a patient home with a device may be economically more efficient and provide better outcomes than an extended period in an intensive care unit, the therapy remains expensive and is not feasible for broad application in the health systems of many countries. Can so expensive an innovation be made available to the wide patient base that could benefit from them? Wider adoption and increased entries into the market are reducing costs, but the fact of profound global disparities in medical device adoption remains for VADs and other complex and new device technologies. Developing approaches for best practices in medical device management are an area of interest and study within the World Health Organization and the reader is referred to an ongoing series of publications and forums from this source that consider how to best translate medical device advances across disparate healthcare economies. These five cases, only a small fraction of the important medical devices that could have been used as examples,

CHAPTER 1.1.1   Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor

spotlight central ideas and themes relevant to most medical devices interfacing with the human biology. A few generalizations are: • Implantation in hundreds of thousands of patients with good success is noted. • A broad range of synthetic materials of varying properties are used. • Most anatomical sites can be interfaced with a device. • The normal response by which the body responds to foreign bodies is observed. • Problems, concerns, unexplained observations, or unintended consequences may be noted for each device. • Most device complications are related to biomaterials–tissue interactions. • Companies are manufacturing devices and bringing value to shareholders (and patients). • Regulatory agencies are carefully assessing device performance, and making policy decisions to monitor the device industry, ensure quality, and protect the patient. • Ethical and societal issues are associated with each device. These ideas are relevant to nearly all medical devices. As you work through this text, consider how these ideas impact the specific topic you are studying. 

Necessary Steps

Now that we have defined key terms and reviewed specific examples highlighting successes and also complications, we can examine core characteristics of the field of biomaterials.

Multidisciplinary More than any other field of contemporary technology, biomaterials science brings together teams of researchers from diverse academic and industrial backgrounds, who characteristically speak different “languages” yet must clearly communicate and integrate complex concepts and data. Fig. 1.1.1.9 lists some of the disciplines and key steps that are encountered in the progression from identifying the need for a biomaterial or device to its testing, regulation, manufacture, sale, and implantation. 

Diverse Materials Are Used The biomaterials scientist must have an appreciation of materials science, including polymers, metals, ceramics, glasses, composites, and biological materials. This may range from an impressive command of the theory and practice of the field demonstrated by the professional materials scientist to a general understanding of the properties of materials that should be possessed by the physician or biologist investigator involved in biomaterials-related research. A wide range of materials is routinely used in medical devices (Table 1.1.1.1), and no one researcher will be comfortable synthesizing, processing, characterizing, designing, and fabricating with all these materials. Thus specialization is common and appropriate. However, a broad appreciation of the properties and applications of these materials,

Facilitator

Identify a need • Treat a condition • Replace an organ • Cosmetic

• Physician/Dentist • Researcher • Inventor • Entrepreneur

Device Design

• Physician (consults) • Engineer (implements)

Materials Synthesis

• Ceramicist • Metallurgist • Polymer Chemist

Materials testing • Mechanical properties • Toxicology • Bioreaction to the material protein interactions cell activation tissue reaction • Biostability mechanical chemical

• Bioengineer • Mechanical Engineer • Biochemist • Cell biologist • Veterinary surgeon

Fabrication

• Engineer • Machinist

Sterilization and Packaging

• Bioengineer • Industrial Designer

Device Testing Toxicology In Vitro Biointeraction Animal Testing

• Bioengineer • Veterinary surgeon • Physician/Dentist

Regulatory pre-market approval limited clinical study clinical trials long-term follow-up

• Regulatory Specialist • Regulatory Agency • Legislators

Clinical Use

• Physician • Dentist • Optometrist

Explant Analysis explant registry pathological examination testing to understand failure

• Pathologist • Bioengineer

Characteristics of Biomaterials Science

13

• Figure 1.1.1.9  The path from an identified need to a clinical product, and some of the disciplines that facilitate this developmental process.

the palette from which the biomaterials scientist “creates” medical devices, is a hallmark of professionals in the field. There is a tendency to group biomaterials and researchers into the “hard tissue replacement” camp, typically represented by those involved in orthopedic and dental materials, and the “soft tissue replacement” camp, frequently associated with cardiovascular implants and general plastic surgery materials.

14 SEC T I O N 1 .1     Overview of Biomaterials

TABLE   Cardiovascular Device Market by Region and With Projected Compound Annual Growth Rate (CAGR) 1.1.1.3  ($ Millions)

Region

2016

2017

2020

CAGR 2017–2022 (%)

North America

9,811

11,329

16,543

7.9

Europe

7,293

8,458

12,084

7.4

Asia

5,610

6,709

11,406

11.2

Rest of World

2,670

3,162

5,218

10.6

Total

25,384

29,658

45,260

8.8

Source: BCC Research.

Hard tissue biomaterials researchers are thought to focus on metals and ceramics, while soft tissue biomaterials researchers are considered polymer experts. In practice, this division is artificial: a heart valve may be fabricated from polymers, metals, and carbons (and processed tissue). A hip joint will also be composed of metals and polymers (and sometimes ceramics), and may be interfaced with the body via a polymeric bone cement. There is a need for a general understanding of all classes of materials and the common conceptual theme of their interaction with the biological milieu. This text provides a background to the important classes of materials, hard and soft, and their interactions with the biological environment. 

Biomaterials to Devices to Markets and Medicine Thomas Edison once said that he would only invent things that people would buy. In an interesting way, this idea is central to biomaterials device development. The process of biomaterials/medical device innovation is driven by clinical need: an informed engineer, patient, or physician defines a need and then initiates an invention. However, someone must test and manufacture the device, and shepherd it though the complex and expensive development process, which includes rigorous regulatory requirements. This “someone” is generally a company, and a company exists (by law) to return value to its shareholders. Fig. 1.1.1.9 illustrates multidisciplinary interactions in biomaterials, and shows the progression in the development of a biomaterial or device. It provides a perspective on how different disciplines work together, starting from the identification of a need for a biomaterial through development, manufacture, implantation, and (possibly) removal from the patient. Note that the development process for medical devices is very different from that for drugs. There are insightful reference works available to help understand this specialized device commercialization process (Yock et al., 2015) and this pathway is the general focus of Part 3 of this text. 

Magnitude of the Field The magnitude of the medical device field expresses both the magnitude of the need and a sizable and growing commercial market (Table 1.1.1.2). Of particular note is how

various medical devices are seeing growth occurring in more recently expanding economies. This is exemplified by the cardiovascular device market (Table 1.1.1.3) and reflects economic growth, extending lifespans, and (unfortunately) increased cardiovascular disease burden in these regions. Consider four commonly used biomaterial devices: a contact lens; a hip joint; a hydrocephalous drainage shunt; and a heart valve. Let us examine these devices in the contexts of human needs and commercial markets. The contact lens offers improved vision and, some will argue, a cosmetic enhancement. The hip joint offers mobility to the patient who would otherwise need a cane or crutch or be confined to a bed or wheelchair. The hydrocephalus shunt will allow an infant to survive without brain damage. The heart valve offers a longer life with improved quality of life. A disposable contact lens may sell for less than $0.50, and the hip joint, hydrocephalus shunt, and heart valve may sell for $1000–$5000 each. Each year there will be hundreds of millions of contact lenses purchased worldwide, 4,500,000 heart valves, 160,000 hydrocephalus shunts, and 1,400,000 total artificial hip prostheses. Here are the issues for consideration: (1) the large number of devices (an expression of both human needs and commercial markets); (2) medical significance (cosmetic to life saving); and (3) commercial potential (who will manufacture it and why, for example, what is the market for the hydrocephalus shunt?). Always, human needs and economic issues color this field we call “biomaterials science.” Medical practice, market forces, and bioethics come into play almost every day. Lysaght and O’Laughlin (2000) estimated the magnitude and economic scope of the contemporary organ replacement enterprise to be much larger than was generally recognized. In the year 2000, the lives of over 20 ­million patients were sustained, supported, or significantly improved by functional organ replacement. The impacted population grows at over 10% per year. Worldwide, firstyear and follow-up costs of organ replacement and prostheses exceed US$300 billion per year and represent between 7% and 8% of total worldwide healthcare spending. In the United States, the costs of therapies enabled by organ replacement technology exceed 2% of the gross national product. The costs are also impressive when reduced to the

CHAPTER 1.1.1   Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor

needs of the individual patient. For example, the cost of an implanted heart valve is roughly $5000. The surgery to implant the device entails a hospital bill and first-year follow-up costs upward of at least $40,000. Reoperation for replacing a failed valve will have these same costs. Reoperations for failed valves now exceed 10% of all valve operations. Global expenditures on medical devices by category are summarized in Table 1.1.1.2.

Success and Failure Most biomaterials and medical devices perform satisfactorily, improving the quality of life for the recipient or saving lives. However, no artificial construct is perfect. All manufactured devices have a failure rate. Also, all humans are different, with differing ethnicities, ages, genetics, gender, body chemistries, living environments, and degrees of physical activity. Furthermore, physicians implant or use these devices with varying degrees of skill. The other side to the medical device success story is that there are problems, compromises, complications, and unintended consequences that often occur with medical devices. Central issues for the biomaterials scientist, manufacturer, patient, physician, and attorney are: (1) is the design competent and optimal; (2) who should be responsible when devices perform “with an inappropriate host response”; and (3) what are the risk/benefit or cost/benefit ratios for the implant or therapy? Some examples may clarify these issues. Clearly, heart valve disease is a serious medical problem. Patients with diseased aortic heart valves have a 50% chance of dying within 3 years. Surgical replacement of the diseased valve leads to an expected survival of 10 years in 70% of the cases. However, of these patients whose longevity and quality of life have clearly been enhanced, approximately 60% will suffer a serious valve-related complication within 10 years after the operation. Another example involves VADs. A clinical trial called Randomized Evaluation of Mechanical Assistance for the Treatment of Congestive Heart Failure (REMATCH) led to the following important statistics (Rose et al., 2001). Patients with an implanted Heartmate VAD had a 52% chance of surviving for 1 year, compared with a 25% survival rate for patients who took medication. Survival for 2 years in patients with the Heartmate was 23% versus 8% in the medication group. Also, the VAD enhanced the quality of life for the patients—they felt better, were less depressed, and were mobile. Importantly, patients participating in the REMATCH trial were not eligible for a heart transplant. In the cases of the heart valve and the VAD, clinical complications possibly associated with less than stellar biomaterials performance do not preclude widespread clinical acceptance. Biomaterials science: • multidisciplinary; • multibiomaterial; • clinical need driven; • substantial world market; and • risk/benefit issues.

15

Thus, these five characteristics of biomaterials science: • multidisciplinary • multibiomaterial • clinical need driven • substantial world market, and • risk/benefit issues color all aspects of the field. In addition, there are certain unique subjects that are particularly prominent in our field and help delineate the biomaterials endeavor as a unique field of science and engineering. Let us review a few of these. 

Subjects Integral to Biomaterials Science Toxicology A biomaterial should not be toxic, unless it is specifically engineered for such requirements (e.g., a “smart” drug delivery system that targets cancer cells with a toxic drug). Since the nontoxic requirement is the norm, toxicology for biomaterials has evolved into a sophisticated science. It deals with the substances that migrate out of biomaterials or result from their degradation. For example, for polymers, many low molecular weight “leachables” exhibit some level of physiologic activity and cell toxicity. It is reasonable to say that a biomaterial should not give off anything from its mass unless it is specifically designed to do so. Toxicology also deals with methods to evaluate how well this design criterion is met when a new biomaterial is under development. Chapter 2.2.5 provides an overview of methods in biomaterials toxicology. Implications of toxicity are addressed in Chapters 2.3.2–2.3.4. 

Biocompatibility The understanding and measurement of biocompatibility are unique to biomaterials science. Unfortunately, we do not have precise definitions or accurate measurements of biocompatibility. More often than not, biocompatibility is defined in terms of performance or success at a specific task. Thus for a patient who is doing well with an implanted large diameter Dacron fabric vascular prosthesis, few would argue that this prosthesis is not “biocompatible.” However, the prosthesis probably did not recellularize, and may also generate a small amount of surface-bound clot that may embolize, usually with little clinical consequence. This operational definition of biocompatible (“the patient is alive and not experiencing complications, so it must be biocompatible”) offers us little insight in designing new or improved vascular prostheses. It is probable that biocompatibility may have to be specifically defined for applications in soft tissue, hard tissue, and the cardiovascular system (blood compatibility, Chapters 2.2.6 and 2.3.5). In fact, biocompatibility may have to be uniquely defined for each application. The problems and meanings of biocompatibility will be explored and expanded upon throughout this textbook; in particular see Chapters 2.3.1 and 2.3.2. 

Inflammation and Healing Specialized biological mechanisms are triggered when a material or device interfaces with the body. Injury to tissue will

16 SEC T I O N 1 .1     Overview of Biomaterials

stimulate the well-defined inflammatory reaction sequence that ultimately leads to healing. Healing can be normal (physiological) or abnormal (pathological). Where a foreign body (e.g., an implant) is present in the wound site (the surgical incision), the reaction sequence is referred to as the “foreign-body reaction” (Chapters 2.2.2 and 2.3.4). The normal response of the body will be modulated because of the solid implant. Furthermore, this reaction will differ in intensity and duration, depending upon the anatomical site involved. An understanding of how a foreign object shifts the normal inflammatory reaction sequence is an important concern for the biomaterials scientist, and how some classes of materials may avoid this reaction is an area of growing interest, as already noted. 

Functional Tissue Structure and Pathobiology Biomaterials-based medical devices are implanted into almost all tissues and organs. Tissues and organs vary widely in cell composition, morphological organization, vascularization, and innervation. Implantation of a biomaterial into bone, liver, or heart will have special physiological consequences. Therefore key principles governing the structure of normal (and abnormal) cells, tissues, and organs are important to biomaterials researchers. Also, techniques by which the structure and function of normal and abnormal tissue are studied must be mastered. In addition, fundamental mechanisms leading to abnormal cell, tissue, and organ structures (i.e., diseases and other pathologies) are critical considerations to biomaterials researchers (see Chapters 2.1.4 and 2.1.5). 

Dependence on Specific Anatomical Sites of Implantation Consideration of the anatomical site of an implant is essential. An intraocular lens may be implanted into the lens capsule or the anterior chamber of the eye. A hip joint will be implanted in bone across an articulating joint space. A prosthetic heart valve will be sutured into cardiac muscle and will contact both soft tissue and blood. A catheter may be placed in an artery, a vein, or the urinary tract. Each of these sites challenges the biomedical device designer with special requirements for anatomy, physiology, geometry, size, mechanical properties, and bioresponses. 

Mechanical Requirements and Physical Performance Requirements Each biomaterial and device has mechanical and performance requirements originating from the need to perform a physiological function. These requirements can be divided into three categories: mechanical performance, mechanical durability, and physical properties. First, consider mechanical performance. A hip prosthesis must be strong and rigid. A tendon material must be strong and flexible. A tissue heart valve leaflet must be flexible and tough. A dialysis membrane must be strong and flexible, but not elastomeric. An articular cartilage substitute must be soft and elastomeric. One significant example of a controlled micromechanical interface is the contact zone between a synthetic biomaterial (titanium, tantalum, alumina, zirconia, and hydroxyapatites) and oral bones. Microstrain

magnitudes have been controlled through macro/micro/ nanosurface topographies and construct designs to be within the microstrain limits of bone. The result has been decades of chemical stability, now called osseous integration. Then, we must address mechanical durability. A catheter may only have to perform for 3 days. A bone plate may fulfill its function in 6 months or longer. A leaflet in a heart valve must flex 60 times per minute without tearing for the lifetime of the patient (realistically, at least for 10 or more years). A hip joint must not fail under heavy loads for 20 years or more. Finally, the bulk physical properties impact other aspects of performance as they meet the physical and/or mechanical demands of the medical devices for which they are designed. In addition to mechanical durability, the dialysis membrane has a specified permeability, the acetabular cup of the hip joint must have high lubricity, and the intraocular lens has transparency and refraction requirements. To meet these requirements, design principles are borrowed from physics, chemistry, mechanical engineering, chemical engineering, and materials science. 

Industrial Involvement A significant basic research effort is now under way, primarily at universities, to understand how biomaterials function and how to optimize them. At the same time, companies are producing implants for use in humans and, appropriate to the mission of a company, earning profits on the sale of medical devices. Thus although we are now only learning about the fundamentals of biointeraction, we manufacture millions of devices for implantation in humans. How is this dichotomy explained? Basically, as a result of 50 or more years of experience we now have a set of materials that perform satisfactorily in the body. The medical practitioner can use them with reasonable confidence, and the performance in the patient is largely acceptable (generally considerably better than other alternatives). Though the devices and materials are far from perfect, the complications associated with the devices are fewer or of less impact than the complications of the original diseases. 

Risk/Benefit and Corporate Realities A risk/benefit analysis must be considered in developing new devices and improving existing devices. Central to biomaterials science is the desire to alleviate suffering and death, and also the desire to improve the quality of life for patients. These considerations are convoluted with the excitement of new scientific ideas, the corporate imperative to turn a profit, and the mandate of the regulatory agencies to protect the public. Indeed, although failure of biomaterials and medical devices is common, benefit to risk ratio in individual cases is often high, and despite a device complication, a patient may have had a markedly improved outcome (enhanced survival and/or quality of life) over the natural history of the disease. The acceptable risk varies with different types of medical devices. Moreover, the acceptable risk of devices that sustain life (e.g., heart valve, defibrillator, cardiac assist device, hemodialysis device/access graft, hydrocephalus shunt) is greater than that of devices that alleviate pain/disability or enhance function (e.g., hip joint, drug delivery device, intraocular lens, intrauterine contraceptive device). Then consider the

CHAPTER 1.1.1   Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor

acceptable risk for devices that have only cosmetic application (e.g., collagen injections, breast implants). Obviously, ethical concerns enter into the risk/benefit picture. Remember that companies have large investments in the development, manufacture, quality control, clinical testing, regulatory clearance, and distribution of medical devices. How much of an advantage (for the company and the patient) will be realized in introducing an improved device? The improved device may indeed work better for the patient. However, the company will incur a large expense (development and regulatory costs) that will be perceived by the stockholders as reduced profits. The development of a new or improved device, as well as offering benefits, entails risks that months or years after introduction some unforeseen complication will compromise the device. Product liability issues are a major concern to manufacturers. Consider questions about the ethics of withholding improved devices from people who could benefit from them because of development costs and regulatory hurdles, the market share advantages of having a better product, and the gargantuan costs (possibly nonrecoverable) of introducing a new product into the medical marketplace. If companies did not have the profit incentive, would there be any medical devices, let alone improved ones, available for clinical application? From the biomaterials industry we see specialized, essential contributions to our field. Industry deals well with technologies such as packaging, sterilization, storage, distribution, quality control, and analysis. These subjects are grounded in specialized technologies, often ignored in academic communities, but having the potential to generate stimulating research questions. Also, many companies support in-house basic research laboratories, and contribute in important ways to the fundamental study of biomaterials science. 

Ethics A wide range of ethical considerations impact biomaterials science. Some key ethical questions in biomaterials science are summarized in Table 1.1.1.4. Typical of ethical questions, an absolute answer may be difficult to come by. Some articles have addressed ethical questions in biomaterials and debated the important points (Saha and Saha, 1987; Schiedermayer and Shapiro, 1989; Merryman, 2008). Chapter 3.1.11 introduces moral and ethical issues related to biomaterials and medical devices. 

Regulation The consumer (the patient) and the physician demand safe medical devices. To prevent inadequately tested devices and materials from coming on the market, and to screen out those clearly unqualified to produce biomaterials, the US government has evolved a complex regulatory system administered by the US Food and Drug Administration. In the United States, medical device regulatory requirements were introduced less than 50 years ago (with the 1976 Medical Device Amendments legislation). Most nations of the world have similar medical device regulatory bodies. The International Standards Organization has introduced international standards for the world community. Obviously, a substantial base of biomaterials knowledge went into establishing these standards. The costs to comply with the standards and to implement materials, biological, and

17

TABLE   Ethical Concerns Relevant to Biomaterials 1.1.1.4  Science

Animals Is the animal model relevant to human physiology? Specifically, is the experiment well designed and the outcome sufficiently important so that the data obtained will justify the suffering and sacrifice of the life of a living creature?

Human Subjects How should human subject research be conducted to minimize negative outcomes to the patient and offer a reasonable risk/benefit ratio? How can we best ensure informed consent?

Industrial Involvement Companies fund much biomaterials research and also own proprietary biomaterials. How can the needs of the patient be best balanced with the financial goals of a company? Consider that someone must manufacture devices—these would not be available if a company did not choose to manufacture them.

Researchers Since researchers often stand to benefit financially from a successful biomedical device, and sometimes even have devices named after them, how can investigator bias be minimized in biomaterials research?

Patients For life-sustaining devices, what is the trade-off between sustaining life and the quality of life with the device for the patient? Should the patient be permitted to “pull the plug” if the quality of life is not satisfactory?

Regulatory Agencies With so many unanswered questions about the basic science of biomaterials, do government regulatory agencies have sufficient information to define adequate tests for materials and devices and to properly regulate biomaterials?

clinical testing are enormous. Introducing a new biomedical device to the market requires a regulatory investment of tens of millions of dollars. Are the regulations and standards truly addressing the safety issues? Is the cost of regulation inflating the cost of healthcare and preventing improved devices from reaching those who need them? Under this regulation topic, we see the intersection of all the players in the biomaterials community: government, industry, scientists, physicians, and patients. The answers are not simple, but the problems must be addressed every day. Part 3 of this text contains several chapters that expand on standards and regulatory concerns to provide a whole-spectrum view, including issues related to device life cycle, safety and risk, sterilization and disinfection, verification and validation, commercialization, and legal concepts. 

Biomaterials Literature Over the past 70 years, the field of biomaterials has evolved from individual physicians innovating to save the lives of their patients to the science-grounded multidisciplinary endeavor we see today. In 1950, there were no biomaterials or medical device journals, and few books. Concurrent with the evolution

18 SEC T I O N 1 .1     Overview of Biomaterials



Figure 1.1.1.10 The cover of the program book for the 1975 International Biomaterials Symposium, Clemson, South Carolina.

of the discipline, a literature has also developed addressing basic science, applied science, engineering, medicine, and commercial issues. A bibliography is provided in Appendix D to highlight some of the key reference works and technical journals in the biomaterials field. As might be expected, these journals stem from many disciplines and technical societies. 

Biomaterials Societies The biomaterials field evolved from individual researchers and clinicians who intellectually associated their efforts with established disciplines such as medicine, chemistry, chemical engineering, materials science, or mechanical engineering, to a modern field called “biomaterials.” This evolution was paralleled by a growing sense of professionalism and the formation of biomaterials societies as homes for the profession to develop in. Probably the first biomaterials-related society was the American Society for Artificial Internal Organs. Founded in 1954, this group of visionaries established a platform to consider the development of devices such as the artificial kidney and the artificial heart. A Division of Interdisciplinary Studies, the administrative home of a nascent biomaterials effort, was established at Clemson University, Clemson, South Carolina, in 1969. Clemson began organizing annual International Biomaterials Symposia (IBS) in 1969. The first of these symposia was titled “Use of Ceramics in Surgical Implants.” About 100 scientists and physicians attended, and 17 papers were presented. Between 1969 and 1975, seven IBS were held at Clemson. The cover of the 1975 International Biomaterials Symposium program is shown in Fig. 1.1.1.10. The Society for Biomaterials (SFB) was chartered in San Antonio, Texas, in 1974 with 205 charter members from across the United States and nine other countries—a truly

international society. One unique feature of the SFB founding members was that they included clinicians, engineers, chemists, and biologists. Their common interest, biomaterials, was the engaging focus for the multidisciplinary participants. Because of the founding of the SFB in 1974, the seventh IBS in 1975 was also known as the world’s inaugural meeting of the SFB. Table 1.1.1.5 lists the timeline of events leading to the start of the SFB annual meetings and the quadrennial World Biomaterials Congress (WBC), as well as the establishment of Clemson Awards and the honorary status of “Fellow, Biomaterials Science and Engineering.” The Canadian Biomaterials Society/Société Canadienne des Biomatériaux was established in 1973, a year earlier than the SFB. The European Society for Biomaterials was founded in 1975, and the Japanese Society for Biomaterials was formed in 1978. To promote international communication and cooperation, these four societies decided to establish an International Liaison Committee of Societies for Biomaterials (ILC) in 1980 to organize a WBC every 4 years and the first WBC was held in Baden, Vienna, Austria, that year. Six more international societies for biomaterials were established afterward: the Society for Biomaterials & Artificial Organs (India) in 1986, the Australasian Society for Biomaterials and Tissue Engineering in 1989, the Korean Society for Biomaterials in 1996, the Chinese Taipei Society for Biomaterials and Controlled Release in 1997, the Latin American Society for Biomaterials and Artificial Organs in 1998, and the Chinese Society for Biomaterials in 2011. In 1997, the constituent societies renamed the ILC to the International Union of Societies for Biomaterials Science and Engineering. Aside from these societies, there are other groups. For example, the Controlled Release Society is a group strongly rooted in biomaterials for drug delivery, and the Tissue Engineering and Regenerative Medicine International Society is also very active in biomaterials-related research.

CHAPTER 1.1.1   Introduction to Biomaterials Science: An Evolving, Multidisciplinary Endeavor

TABLE   Timeline for Development of the Society 1.1.1.5  for Biomaterials and Other International

Biomaterials Organizations

19

master certain key material from many fields of science, technology, engineering, and medicine to be competent and conversant in this profession. The reward for mastering this volume of material is immersion in an intellectually stimulating endeavor that advances a new basic science of biointeraction and contributes to reducing human suffering.

References

FBSE, Fellow, Biomaterials Science and Engineering; IBS, International Biomaterials Symposia; ILC, International Liaison Committee of Societies for Biomaterials; IUSBSE, International Union of Societies for Biomaterials Science and Engineering; SC, South Carolina; SFB, Society for Biomaterials; WBC, World Biomaterials Congress.

The development of biomaterials professionalism and a sense of identity for the biomaterials field can be attributed to these societies and the researchers who organized and led them. 

Summary This chapter provides a broad overview of the biomaterials field. It offers a vantage point from which the reader can gain a perspective to see how the subthemes fit into the larger whole. Biomaterials science may be the most multidisciplinary of all the sciences. Consequently, biomaterials scientists must

Carlsson, L., Rostlund, T., Albrektsson, B., Albrektsson, T., Branemark, P.I., 1986. Osseointegration of titanium implants. Acta Orthop. Scand. 57, 285–289. Hench, L.L., Pollak, J.M., 2002. Third-generation biomedical materials. Science 295, 1014–1017. Hubbell, J.A., Massia, S.P., Desai, N.P., Drumheller, P.D., 1991. Endothelial cell-selective materials for tissue engineering in the vascular graft via a new receptor. Biotechnology 9 (6), 568–572. Jana, S., Tefft, B.J., Spoon, D.B., Simari, R.D., 2014. Scaffolds for tissue engineering of cardiac valves. Acta Biomater. 10, 2877–2893. Lysaght, M.J., O’Laughlin, J., 2000. The demographic scope and economic magnitude of contemporary organ replacement therapies. Am. Soc. Artif. Intern. Organs J. 46, 515–521. Merryman, W.D., 2008. Development of a tissue engineered heart valve for pediatrics: a case study in bioengineering ethics. Sci. Eng. Ethics 14, 93–101. Meckstroth, K.R., Darney, P.D., 2001. Implant contraception. Semin. Reprod. Med. 19, 339. Ratner, B.D., 2011. The biocompatibility manifesto: biocompatibility for the twenty-first century. J. Cardiovasc. Translat. Res. 4 (5), 523–527. Rose, E.A., Gelijns, A.C., Moskowitz, A.J., Heitjan, D.F., Stevenson, L.W., Dembitsky, W., Long, J.W., Ascheim, D.D., Tierney, A.R., Levitan, R.G., Watson, J.T., Ronan, N.S., Shapiro, P.A., Lazar, R.M., Miller, L.W., Gupta, L., Frazier, O.H., Desvigne-Nickens, P., Oz, M.C., Poirier, V.L., Meier, P., 2001. Long-term use of a left ventricular assist device for end-stage heart failure. N. Engl. J. Med. 345, 1435–1443. Saha, S., Saha, P., 1987. Bioethics and applied biomaterials. J. Biomed. Mater. Res. Appl. Biomater 21, 181–190. Sahoo, J.K., VandenBerg, M.A., Webber, M.J., 2018. Injectable network biomaterials via molecular or colloidal self-assembly. Adv. Drug Deliv. Rev. 127, 185–207. Schiedermayer, D.L., Shapiro, R.S., 1989. The artificial heart as a bridge to transplant: ethical and legal issues at the bedside. J. Heart Transplant. 8, 471–473. Sharp, P.A., Langer, R., 2011. Promoting convergence in biomedical science. Science 333 (6042), 527. Williams, D.F., 1987. In: Definitions in biomaterials. Proceedings of a Consensus Conference of the European Society for Biomaterials. (Vol. 4). New York: Elsevier, Chester, England, March 3–5 1986, . Yock, P.G., Zenios, S., Makower, J., Brinton, T.J., Kumar, U.N., Watkins, F.T.J, Denend, L., Krummel, T.M., Kurihara, C.Q. (Eds.), 2015. Biodesign: The Process of Innovating Medical Technologies. Cambridge: Cambridge University Press. Zhang, B.L., Bianco, R.W., Schoen, F.J., 2019. Preclinical assessment of cardiac valve substitutes: current status and considerations for engineered tissue heart valves. Front. Cardiovasc. Med. 6, 72. Zhang, X., Gong, C., Akakuru, O.U., Su, Z., Wu, A., Wei, G., 2019. The design and biomedical applications of self-assembled two-dimensional organic biomaterials. Chem. Soc. Rev. 48 (23), 5564–5595.

1.1.2

A History of Biomaterials BUDDY D. RATNER 1 , GUIGEN ZHANG 2 1Bioengineering

and Chemical Engineering, Director of University of Washington Engineered Biomaterials (UWEB), Seattle, WA, United States 2F

Joseph Halcomb III, M.D. Department of Biomedical Engineering, University of Kentucky, Lexington, KY, United States

History consists of a series of accumulated imaginative inventions. Voltaire At the turn of the third decade of the 21st century, biomaterials are still widely used throughout medicine, dentistry, and biotechnology. Just about 70 years ago, biomaterials as we think of them today did not exist. The word “biomaterial” was not used. There were no medical device manufacturers (except for external prosthetics such as limbs, fracture fixation devices, glass eyes, and dental fillings and devices), no formalized regulatory approval processes, no understanding of biocompatibility, and certainly no academic courses on biomaterials. Yet, crude biomaterials have been used, generally with poor to mixed results, throughout history. This chapter will broadly trace the history of biomaterials from the earliest days of human civilization to the third decade of the 21st century. It is convenient to organize the history of biomaterials into four eras: prehistory; the era of the surgeon-hero; designed biomaterials/engineered devices; and the contemporary era taking us into the new millennium. The emphasis of this chapter will be on the experiments and studies that set the foundation for the field we call biomaterials, largely between 1920 and 1980.

Biomaterials Before World War II Before Civilization The introduction of nonbiological materials into the human body took place throughout history. The remains of a human found near Kennewick, Washington, USA (often referred to as the “Kennewick Man”), were dated (with some controversy) to be 9000 years old. This individual, described by archeologists as a tall, healthy, active person, wandered through the region now known as southern Washington with a spear point embedded in his hip. It had apparently healed in and did not significantly impede

his activity. This unintended implant illustrates the body’s capacity to deal with implanted foreign materials. The spear point has little resemblance to modern biomaterials, but it was a “tolerated” foreign material implant. Another example of the introduction of foreign material into the skin, dated at over 5000 years ago, is the tattoo. The carbon particles and other substances probably elicited a classic foreign-body reaction. Moreover, one of the first surgical textbooks dated around 600 BC (Willyard, 2016) documented, arguably the first written record of skin-graft techniques, a method for repairing torn earlobes with skin from the cheek and the reconstruction of the nose from a flap of forehead skin. 

Dental Implants in Early Civilizations Unlike the spear point described earlier, dental implants were devised as implants and used early in history. The Mayan people fashioned nacre teeth from sea shells in roughly 600 AD, and apparently achieved what we now refer to as osseointegration (see Chapter 2.5.5), basically a seamless integration into the bone (Bobbio, 1972). Similarly, in France, a wrought iron dental implant in a corpse was dated to 200 AD (Crubezy et al., 1998). This implant, too, was described as properly osseointegrated. There was no materials science, biological understanding, or medicine behind these procedures. Still, their success (and longevity) is impressive and highlights two points: the adaptive nature of the human body and the pressing drive, even in prehistoric times, to restore the lost functions of physiologic/anatomic parts of the body with an implant. 

Sutures Dating Back Thousands of Years There is loose evidence that sutures may have been used even in the Neolithic period. Large wounds were closed early in history primarily by one of two methods—cautery or sutures. Linen sutures were used by the early Egyptians. 21

22 SEC T I O N 1 .1     Overview of Biomaterials

Catgut was used in the Middle Ages in Europe. In South Africa and India, the heads of large, biting ants clamped wound edges together. Metallic sutures are first mentioned in early Greek literature. Galen of Pergamon (ca.130–200 AD) described ligatures of gold wire. In 1816, Philip Physick, University of Pennsylvania Professor of Surgery, suggested the use of lead wire sutures, noting little reaction. J. Marion Sims of Alabama had a jeweler fabricate sutures of silver wire, and in 1849 performed many successful operations with this metal. Consider the problems that must have been experienced with sutures in times with no knowledge of sterilization, toxicology, immunological reaction to extraneous biological materials, inflammation, and biodegradation. Yet sutures were a relatively common fabricated or manufactured biomaterial for thousands of years. 

Artificial Hearts and Organ Perfusion In the fourth century BC, Aristotle called the heart the most important organ in the body. Galen proposed that veins connected the liver to the heart to circulate “vital spirits throughout the body via the arteries.” English physician William Harvey in 1628 espoused a relatively modern view of heart function when he wrote “the heart’s one role is the transmission of the blood and its propulsion, by means of the arteries, to the extremities everywhere.” With the appreciation of the heart as a pump, it was a logical idea to think of replacing the heart with an artificial pump. In 1812, the French physiologist Le Gallois expressed his idea that organs could be kept alive by pumping blood through them. A number of experiments on organ perfusion with pumps were performed from 1828 to 1868. In 1881, Étienne-Jules Marey, a brilliant scientist and thinker who published and invented in photographic technology, motion studies, and physiology, described an artificial heart device (Fig. 1.1.2.1), primarily oriented to studying the beating of the heart. In 1938, aviator (and engineer) Charles Lindbergh and surgeon (and Nobel Prize winner) Alexis Carrel wrote a visionary book, The Culture of Organs. They addressed issues of pump design (referred to as the Lindbergh pump), sterility, blood damage, nutritional needs of perfused organs, and mechanics. This book is a seminal document in the history of artificial organs. In the mid-1950s, Dr. Paul Winchell, better known as a ventriloquist, patented an artificial heart. In 1957, Dr. Willem Kolff and a team of scientists tested the artificial heart in animals. More modern conceptions of the artificial heart (and left ventricular assist device) are presented in the following sections and in Chapter 2.5.2A and B). 

Contact Lenses Leonardo DaVinci, in the year 1508, developed the concept of contact lenses. Rene Descartes was credited with the idea of the corneal contact lens in 1632 and Sir John F.W. Herschel suggested in 1827 that a glass lens could protect the eye. Adolf Gaston Eugen Fick (nephew of Adolf Eugen Fick

• Figure 1.1.2.1  An artificial heart by Étienne-Jules Marey, Paris, 1881.

of Fick’s law of diffusion fame) was an optometrist by profession. One of his inventions (roughly 1860) was a glass contact lens, possibly the first contact lens offering real success. He experimented on both animals and humans with contact lenses. In the period from 1936 to 1948, plastic contact lenses were developed, primarily of poly(methyl methacrylate).

 Basic Concepts of Biocompatibility Most implants prior to 1950 had a low probability of success, because of a poor understanding of biocompatibility and sterilization. As will be elaborated upon throughout this book, factors that contribute to biocompatibility include the chemistry of the implant, leachables, shape, mechanics, and design. Early studies, especially with metals, focused primarily on ideas from chemistry to explain the observed bioreaction. Possibly the first study assessing the in vivo bioreactivity of implant materials was performed by H.S. Levert (1829). Gold, silver, lead, and platinum specimens were studied in dogs, and platinum, in particular, was found to be well tolerated. In 1886, bone fixation plates of nickel-plated sheet steel with nickel-plated screws were studied. In 1924, A. Zierold published a study on tissue reaction to various

CHAPTER 1.1.2   A History of Biomaterials

materials in dogs. Iron and steel were found to corrode rapidly, leading to resorption of adjacent bone. Copper, magnesium, aluminum alloy, zinc, and nickel discolored the surrounding tissue. Gold, silver, lead, and aluminum were tolerated, but were inadequate mechanically. Stellite, a Co–Cr–Mo alloy, was well tolerated and strong. In 1926, M. Large noted inertness displayed by 18-8 stainless steel containing molybdenum. By 1929 Vitallium alloy (65% Co–30% Cr–5% Mo) was developed and used with success in dentistry. In 1947, J. Cotton of the United Kingdom discussed the possible use of titanium and alloys for medical implants. The history of plastics as implantation materials does not extend as far back as metals, simply because there were few plastics prior to the 1940s. What is possibly the first paper on the implantation of a modern synthetic polymer, nylon, as a suture appeared in 1941. Papers on the implantation of cellophane, a polymer made from plant sources, were published as early as 1939, documenting it being used as a wrap for blood vessels. The response to this implant was described as a “marked fibrotic reaction.” In the early 1940s papers appeared discussing the reaction to implanted poly(methyl methacrylate) and nylon. The first paper on polyethylene as a synthetic implant material was published in 1947 (Ingraham et al.). The paper pointed out that polyethylene production using a new high-pressure polymerization technique began in 1936. This process enabled the production of polyethylene free of initiator fragments and other additives. Ingraham et al. demonstrated good results on implantation (i.e., a mild foreign-body reaction), and attributed these results to the high purity of the polymer they used. A 1949 paper commented on the fact that additives to many plastics had a tendency to “sweat out,” and this might be responsible for the strong biological reaction to those plastics (LeVeen and Barberio, 1949). They found a vigorous foreign-body reaction to cellophane, Lucite, and nylon, but an extremely mild reaction to “a new plastic,” Teflon. The authors incisively concluded: “Whether the tissue reaction is due to the dissolution of traces of the unpolymerized chemical used in plastics manufacture or actually to the solution of an infinitesimal amount of the plastic itself cannot be determined.” The possibility that cellulose might trigger the severe reaction by activating the complement system could not have been imagined, because the complement system had not yet been discovered. 

World War II to the Modern Era: The Surgeon/Physician-Hero During World War I, and particularly at the end of the war, newly developed high-performance metal, ceramic, and especially polymeric materials, transitioned from wartime restriction to peacetime availability. The possibilities for using these durable, inert materials immediately intrigued surgeons with needs to replace diseased or damaged body parts. Materials, originally manufactured for airplanes,

23

automobiles, clocks, and radios, were taken “off the shelf ” by surgeons and applied to medical problems. These early biomaterials included silicones, polyurethanes, Teflon, nylon, methacrylates, titanium, and stainless steel. A historical context helps us appreciate the contribution made primarily by medical and dental practitioners. Just after World War II, there was little precedent for surgeons to collaborate with scientists and engineers. Medical and dental practitioners of this era felt it was appropriate to invent (improvise) where the life or functionality of their patient was at stake. Also, there was minimal government regulatory activity, and human subject protections as we know them today were nonexistent (see Chapters 3.1.7 and 3.1.11). The physician was implicitly entrusted with the life and health of the patient and had much more freedom than is seen today to take heroic action when other options were exhausted. These medical practitioners had read about the post-World War II marvels of materials science. Looking at a patient open on the operating table, they could imagine replacements, bridges, conduits, and even organ systems based on such materials. Many materials were tried on the spur of the moment. Some fortuitously succeeded. These were high-risk trials, but usually they took place where other options were not available. The term “surgeon-hero” seems justified, since the surgeon often had a life (or a quality of life) at stake and was willing to take a huge technological and professional leap to repair the individual. This laissez faire biomaterials era quickly led to a new order characterized by scientific/engineering input, government quality controls, and a sharing of decisions prior to attempting high-risk, novel procedures. Still, a foundation of ideas and materials for the biomaterials field was built by courageous, fiercely committed, creative individuals, and it is important to look at this foundation to understand many of the attitudes, trends, and materials common today. The regulatory climate in the United States in the 1950s was strikingly different from today. This can be appreciated in this recollection from Willem Kolff about a pump oxygenator he made and brought with him from Holland to the Cleveland Clinic (Kolff, 1998): Before allowing Dr. Effler and Dr. Groves to apply the pump oxygenator clinically to human babies, I insisted they do 10 consecutive, successful operations in baby dogs. The chests were opened, the dogs were connected to a heart-lung machine to maintain the circulation, the right ventricles were opened, a cut was made in the interventricular septa, the septa holes were closed, the right ventricles were closed, the tubes were removed, and the chests were closed.

Intraocular Lenses Sir Harold Ridley, MD (1906–2001) (Fig. 1.1.2.2), inventor of the plastic intraocular lens, made early, accurate observations of biological reaction to implants consistent with currently accepted ideas of biocompatibility. After World War II, he had the opportunity to examine aviators

24 SEC T I O N 1 .1     Overview of Biomaterials

• Figure 1.1.2.2  Left: Sir Harold Ridley, inventor of the intraocular lens, knighted by Queen Elizabeth II for his achievement. Right: Shards from the canopy of the Spitfire airplane were the inspiration leading to intraocular lenses. (Image by Bryan Fury75 at fr.wikipedia [GFDL (www.gnu.org/copyleft/fdl.html), from Wikimedia Commons].)

who were unintentionally implanted in their eyes with shards of plastic from shattered canopies in Spitfire and Hurricane fighter planes (Fig. 1.1.2.2). Most of these flyers had plastic fragments years after the war. The conventional wisdom at that time was that the human body would not tolerate implanted foreign objects, especially in the eye— the body’s reaction to a splinter or a bullet was cited as examples of the difficulty of implanting materials in the body. The eye is an interesting implant site, because you can look in through a transparent window to observe the reaction. When Ridley did so, he noted that the shards had healed in place with no further reaction. They were, by his standard, tolerated by the eye. Today, we would describe this stable healing without significant ongoing inflammation or irritation as “biocompatible.” This is an early observation of “biocompatibility” in humans, perhaps the first, using criteria similar to those accepted today. Based on this observation, Ridley traced down the source of the plastic domes, ICI Perspex poly(methyl methacrylate). He used this material to fabricate implant lenses (intraocular lenses) that were found, after some experimentation, to function reasonably in humans as replacements for surgically removed natural lenses that had been clouded by cataracts. The first implantation in a human was on November 29, 1949. For many years, Ridley was the center of fierce controversy because he challenged the dogma that spoke against implanting foreign materials in eyes—it is hard to believe that the implantation of a biomaterial would provoke such an outcry. Because of this controversy, this industry did not instantly arise—it was the early 1980s before intraocular lenses became a major force in the biomedical device market. Ridley’s insightful observation, creativity, persistence, and surgical talent in the late 1940s evolved into an industry that presently puts more than 7,000,000 of these lenses annually in humans. Through all of human history, cataracts meant blindness or a surgical procedure that left the recipient needing thick, unaesthetic spectacle lenses that poorly corrected the vision. Ridley’s concept, using a plastic material found to be “biocompatible,” changed the course of history and substantially

improved the quality of life for millions of individuals with cataracts. Harold Ridley’s story is elaborated upon in an obituary (Apple and Trivedi, 2002).

 Hip and Knee Prostheses The first hip replacement was probably performed in 1891 by a German surgeon, Theodore Gluck, using a cemented ivory ball. This procedure was not successful. Numerous attempts were made between 1920 and 1950 to develop a hip replacement prosthesis. Surgeon M.N. Smith-Petersen, in 1925, explored the use of a glass hemisphere to fit over the ball of the hip joint. This failed due to poor durability. Chrome–cobalt alloys and stainless steel offered improvements in mechanical properties, and many variants of these were explored. In 1938, the Judet brothers of Paris, Robert and Jean, explored an acrylic surface for hip procedures, but it had a tendency to loosen. The idea of using fast-setting dental acrylics to glue prosthetics to bone was developed by Dr. Edward J. Haboush in 1953. In 1956, McKee and Watson-Farrar developed a “total” hip with an acetabular cup of metal that was cemented in place. Metal-on-metal wear products probably led to high complication rates. It was John Charnley (1911–82) (Fig. 1.1.2.3), working at an isolated tuberculosis sanatorium in Wrightington, Manchester, England, who invented the first really successful hip joint prosthesis. The femoral stem, ball head, and plastic acetabular cup proved to be a reasonable solution to the problem of damaged joint replacement. In 1958, Dr. Charnley used a Teflon acetabular cup with poor outcomes due to wear debris. By 1961 he was using a high molecular weight polyethylene cup, and was achieving much higher success rates. Interestingly, Charnley learned of high molecular weight polyethylene from a salesman selling novel plastic gears to one of his technicians. Dr. Dennis Smith contributed in an important way to the development of the hip prosthesis by introducing Dr. Charnley to poly(methyl methacrylate) cements, developed in the dental community, and optimizing those cements for hip replacement use. Total knee replacements borrowed elements of the hip

CHAPTER 1.1.2   A History of Biomaterials

25

• Figure 1.1.2.3  Left: Sir John Charnley. Right: The original Charnley hip prosthesis. (Hip prosthesis photo courtesy of the South Australian Medical Heritage Society, Inc.)

prosthesis technology, and successful results were obtained in the period 1968–72 with surgeons Frank Gunston and John Insall leading the way. 

Dental Implants Some “prehistory” of dental implants has just been described. In 1809, Maggiolo implanted a gold post anchor into fresh extraction sockets. After allowing this to heal, he affixed to it a tooth. This has remarkable similarity to modern dental implant procedures. In 1887, this procedure was used with a platinum post. Gold and platinum gave poor long-term results, and so this procedure was never widely adopted. In 1937, Venable used surgical Vitallium and Co–Cr–Mo alloy for such implants. Also around 1937, Strock at Harvard used a screw-type implant of Vitallium, and this may be the first successful dental implant. A number of developments in surgical procedure and implant design (for example, the endosteal blade implant) then took place. In 1952, a fortuitous discovery was made. Per-Ingvar Brånemark, an orthopedic surgeon at the University of Lund, Sweden, was implanting an experimental cage device in rabbit bone for observing healing reactions. The cage was a titanium cylinder that screwed into the bone. After completing the experiment that lasted several months, he tried to remove the titanium device and found it tightly integrated in the bone (Brånemark et al., 1964). Dr. Brånemark named the phenomenon “osseointegration,” and explored the application of titanium implants to surgical and dental procedures. He also developed low-impact surgical protocols for tooth

implantation that reduced tissue necrosis and enhanced the probability of good outcomes. Most dental implants and many other orthopedic implants are now made of titanium and its alloys. 

The Artificial Kidney Kidney failure, through most of history, was a sentence to an unpleasant death lasting over a period of about a month. In 1910, at Johns Hopkins University, the first attempts to remove toxins from blood were made by John Jacob Abel. The experiments were with rabbit blood, and it was not possible to perform this procedure on humans. In 1943, in Nazi-occupied Holland, Willem Kolff (Fig. 1.1.2.4), a physician just beginning his career at that time, built a drum dialyzer system from a 100-L tank, wood slats, and 130 feet of cellulose sausage casing tubing as the dialysis membrane. Some successes were seen in saving lives where prior to this there was only one unpleasant outcome to kidney failure. Kolff took his ideas to the United States and in 1960, at the Cleveland Clinic, developed a “washing machine artificial kidney” (Fig. 1.1.2.5) taking advantage of Maytag washing machines purchased from Sears. Major advances in kidney dialysis were made by Dr. Belding Scribner at the University of Washington (Fig. 1.1.2.6). Scribner devised a method to routinely access the bloodstream for dialysis treatments. Prior to this, after just a few treatments, access sites to the blood were used up and further dialysis was not possible. After seeing the potential of dialysis to help patients, but only acutely, Scribner tells the

26 SEC T I O N 1 .1     Overview of Biomaterials

and is said to be responsible for more than a million patients being alive today. Interestingly, Dr. Scribner refused to patent his invention because of its importance to medical care. Additional important contributions to the artificial kidney were made by chemical engineering Professor Les Babb (University of Washington) who, working with Scribner, improved dialysis performance and invented a proportioning mixer for the dialysate fluid. The first dialysis center was opened in Seattle making use of these important technological advances (Fig. 1.1.2.6). The early experience with dialyzing patients where there were not enough dialyzers to meet the demand also made important contributions to bioethics associated with medical devices (Blagg, 1998). 

The Artificial Heart

• Figure 1.1.2.4  Dr. Willem Kolff at age 92. (Photograph by B. Ratner.)

Willem Kolff was also a pioneer in the development of the artificial heart. He implanted the first artificial heart in the Western hemisphere in a dog in 1957 (a Russian artificial heart was implanted in a dog in the late 1930s). The Kolff artificial heart was made of a thermosetting poly(vinyl chloride) cast inside hollow molds to prevent seams. In 1953, the heart–lung machine was invented by John Gibbon, but this was useful only for acute treatment, such as during open heart surgery. In 1964, the National Heart and Lung Institute of the National Institutes of Health set a goal of a total artificial heart by 1970. Dr. Michael DeBakey implanted a left ventricular assist device in a human in 1966, and Dr. Denton Cooley and Dr. William Hall implanted a polyurethane total artificial heart in 1969. In the period 1982–85, Dr. William DeVries implanted a number of Jarvik hearts based upon designs originated by Drs. Clifford Kwan-Gett and Donald Lyman—patients lived up to 620 days on the Jarvik 7 device. 

Breast Implants

• Figure 1.1.2.5  Willem Kolff (center) and the washing machine artificial kidney.

story of waking up in the middle of the night with an idea to gain easy access to the blood—a shunt implanted between an artery and vein that emerged through the skin as a “U.” Through the exposed portion of the shunt, blood access could be readily achieved. When Dr. Scribner heard about the new plastic, Teflon, he envisioned how to get the blood out of and into the blood vessels. His device, built with the assistance of Wayne Quinton (Fig. 1.1.2.6), used Teflon tubes to access the vessels, a Dacron sewing cuff through the skin, and a silicone rubber tube for blood flow. The Quinton–Scribner shunt made chronic dialysis possible,

The breast implant evolved to address the poor results achieved with direct injection of substances into the breast for augmentation. In fact, in the 1960s, California and Utah classified use of silicone injections as a criminal offense. In the 1950s, poly(vinyl alcohol) sponges were implanted as breast prostheses, but results with these were also poor. University of Texas plastic surgeons Thomas Cronin and Frank Gerow invented the first silicone breast implant in the early 1960s, a silicone shell filled with silicone gel. Many variations of this device have been tried over the years, including cladding the device with polyurethane foam (the Natural Y implant). This variant of the breast implant was fraught with problems. However, the basic silicone rubber–silicone gel breast implant was generally acceptable in performance (Bondurant et al., 1999). 

Vascular Grafts Surgeons have long needed methods and materials to repair damaged and diseased blood vessels. Early in the century,

CHAPTER 1.1.2   A History of Biomaterials

(A)

(B)

27

(C)

• Figure 1.1.2.6  (A) Belding Scribner; (B) Wayne Quinton; (C) plaque commemorating the original location in Seattle of the world’s first artificial kidney center. ((A) Courtesy of Dr. Eli Friedman. (B) Photo by B. Ratner.)

Dr. Alexis Carrel developed methods to anastomose (suture) blood vessels, an achievement for which he won the Nobel Prize in medicine in 1912. In 1942 Blackmore used Vitallium metal tubes to bridge arterial defects in war-wounded soldiers. Columbia University surgical intern Arthur Voorhees (1922–92), in 1947, noticed during a postmortem that tissue had grown around a silk suture left inside a lab animal. This observation stimulated the idea that a cloth tube might also heal by being populated by the tissues of the body. Perhaps such a healing reaction in a tube could be used to replace an artery? His first experimental vascular grafts were sewn from a silk handkerchief and then parachute fabric (Vinyon N), using his wife’s sewing machine. The first human implant of a prosthetic vascular graft was in 1952. The patient lived many years after this procedure, inspiring many surgeons to copy the procedure. By 1954, another paper was published establishing the clear benefit of a porous (fabric) tube over a solid polyethylene tube (Egdahl et al., 1954). In 1958, the following technique was described in a textbook on vascular surgery (Rob, 1958): “The Terylene, Orlon or nylon cloth is bought from a draper’s shop and cut with pinking shears to the required shape. It is then sewn with thread of similar material into a tube and sterilized by autoclaving before use.” 

Stents Partially occluded coronary arteries lead to angina, diminished heart functionality, and eventually, when the artery occludes (i.e., myocardial infarction), death of a localized portion of the heart muscle. Bypass operations take a section of vein from another part of the body and replace the occluded coronary artery with a clean conduit—this is major surgery, hard on the patient, and expensive. Synthetic vascular grafts in the 3 mm diameter size that is appropriate to the human coronary artery anatomy will thrombose, and thus cannot be used. Another option is percutaneous transluminal coronary angioplasty (PTCA). In this procedure, a balloon is threaded on a catheter into the coronary artery and then inflated to open the lumen of the occluding vessel.

• Figure 1.1.2.7  Dr. Julio Palmaz, inventor of the coronary artery stent. (Photograph by B. Ratner.)

However, in many cases the coronary artery can spasm and close from the trauma of the procedure. The invention of the coronary artery stent, an expandable mesh structure that holds the lumen open after PTCA, was revolutionary in the treatment of coronary occlusive disease. In his own words, Dr. Julio Palmaz (Fig. 1.1.2.7) describes the origins and history of the cardiovascular stent: I was at a meeting of the Society of Cardiovascular and Interventional Radiology in February 1978 when a visiting lecturer, Doctor Andreas Gruntzig from Switzerland, was presenting his preliminary experience with coronary balloon angioplasty. As you know, in 1978 the mainstay therapy of coronary heart disease was surgical bypass. Doctor Gruntzig showed his promising new technique to open up coronary atherosclerotic blockages without the need for open chest surgery, using his own plastic balloon catheters. During his

28 SEC T I O N 1 .1     Overview of Biomaterials

presentation, he made it clear that in a third of the cases, the treated vessel closed back after initial opening with the angioplasty balloon because of elastic recoil or delamination of the vessel wall layers. This required standby surgery facilities and personnel, in case acute closure after balloon angioplasty prompted emergency coronary bypass. Gruntzig’s description of the problem of vessel reclosure elicited in my mind the idea of using some sort of support, such as used in mine tunnels or in oil well drilling. Since the coronary balloon goes in small (folded like an umbrella) and is inflated to about 3–4 times its initial diameter, my idealistic support device needed to go in small and expand at the site of blockage with the balloon. I thought one way to solve this was a malleable, tubular, crisscross mesh. I went back home in the Bay Area and started making crude prototypes with copper wire and lead solder, which I first tested in rubber tubes mimicking arteries. I called the device a BEIS or balloonexpandable intravascular graft. However, the reviewers of my first submitted paper wanted to call it a stent. When I looked the word up, I found out that it derives from Charles Stent, a British dentist who died at the turn of the century. Stent invented a wax material to make dental molds for dentures. This material was later used by plastic surgeons to keep tissues in place, while healing after surgery. The word “stent” was then generically used for any device intended to keep tissues in place while healing. I made the early experimental device of stainless steel wire soldered with silver. These were materials I thought would be appropriate for initial laboratory animal testing. To carry on with my project I moved to the University of Texas Health Science Center in San Antonio (UTHSCSA). From 1983– 1986 I performed mainly bench and animal testing that showed the promise of the technique and the potential applications it had in many areas of vascular surgery and cardiology. With a UTHSCSA pathologist, Doctor Fermin Tio, we observed our first microscopic specimen of implanted stents in awe. After weeks to months after implantation by catheterization under X-ray guidance, the stent had remained open, carrying blood flow. The metal mesh was covered with translucent, glistening tissue similar to the lining of a normal vessel. The question remained whether the same would happen in atherosclerotic vessels. We tested this question in the atherosclerotic rabbit model and to our surprise, the new tissue free of atherosclerotic plaque encapsulated the stent wires, despite the fact that the animals were still on a high cholesterol diet. Eventually, a large sponsor (Johnson & Johnson) adopted the project and clinical trials were instituted under the scrutiny of the Food and Drug Administration. Coronary artery stenting is now performed in well over 1.5 million procedures per year. 

Pacemakers In London in 1788, Charles Kite wrote “An Essay Upon the Recovery of the Apparently Dead,” where he discussed

• Figure 1.1.2.8 The Albert Hyman Model II portable pacemaker, c. 1932–33. (Courtesy of the NASPE-Heart Rhythm Society History Project (www.Ep-History.org).)

electrical discharges to the chest for heart resuscitation. In the period 1820–80 it was already known that electric shocks could modulate the heartbeat (and, of course, consider the Frankenstein story from that era). The invention of the portable pacemaker, hardly portable by modern standards, may have taken place almost simultaneously in two groups in 1930–31—Dr. Albert S. Hyman (USA) (Fig. 1.1.2.8) and Dr. Mark C. Lidwill (working in Australia with physicist Major Edgar Booth). Canadian electrical engineer, John Hopps, while conducting research on hypothermia in 1949, invented an early cardiac pacemaker. Hopps’ discovery was if a cooled heart stopped beating, it could be electrically restarted. This led to Hopps’ invention of a vacuum tube cardiac pacemaker in 1950. Paul M. Zoll developed a pacemaker in conjunction with the Electrodyne Company in 1952. The device was about the size of a small microwave oven, was powered with external current, and stimulated the heart using electrodes placed on the chest—this therapy caused pain and burns, although it could pace the heart. In the period 1957–58, Earl E. Bakken, founder of Medtronics, Inc., developed the first wearable transistorized (external) pacemaker at the request of heart surgeon Dr. C. Walton Lillehei. Bakken quickly produced a prototype that Lillehei used on children with postsurgery heart block. Medtronic commercially produced this wearable, transistorized unit as the 5800 pacemaker. In 1959 the first fully implantable pacemaker was developed by engineer Wilson Greatbatch and cardiologist W.M. Chardack. He used two Texas Instruments transistors, a technical innovation that permitted small size and low power drain. The pacemaker was encased in epoxy to inhibit body fluids from inactivating it.

CHAPTER 1.1.2   A History of Biomaterials

29

 Pyrolytic Carbon Pyrolytic carbon (PyC) is a manmade material not found in nature. The term “pyrolytic” is derived from “pyrolysis”— a thermal decomposition process. In other words, PyC is formed from the thermal decomposition of hydrocarbons such as propane, propylene, acetylene, and methane in the absence of oxygen. The use of PyC as leaflets for blood contact application was an accidental event, according to the story told by Professor Robert E. Baier of the University at Buffalo in 2016 (Baier, 2016). Here are his own words:

• Figure 1.1.2.9  The Hufnagel heart valve consisting of a poly(methyl methacrylate) tube and nylon ball. (United States Federal Government image in the public domain.)

 Heart Valves The development of the prosthetic heart valve paralleled developments in cardiac surgery. Until the heart could be stopped and blood flow diverted, the replacement of a valve would be challenging. Charles Hufnagel, in 1952, implanted a valve consisting of a poly(methyl methacrylate) tube and nylon ball in a beating heart (Fig. 1.1.2.9). This was a heroic operation and basically unsuccessful, but an operation that inspired cardiac surgeons to consider that valve prostheses might be possible. The 1953 development of the heart–lung machine by Gibbon allowed the next stage in the evolution of the prosthetic heart valve to take place. In 1960, a mitral valve replacement was performed in a human by surgeon Albert Starr, using a valve design consisting of a silicone ball and poly(methyl methacrylate) cage (later replaced by a stainless-steel cage). The valve was invented by engineer Lowell Edwards. The heart valve was based on a design for a bottle stopper invented in 1858. Starr was quoted as saying: “Let’s make a valve that works and not worry about its looks,” referring to its design that was radically different from the leaflet valve that nature evolved in mammals. Prior to the Starr–Edwards valve, no human had lived with a prosthetic heart valve longer than 3 months. The Starr–Edwards valve was found to permit good patient survival. The major issues in valve development in that era were thrombosis and durability. In 1969, Warren Hancock started the development of the first leaflet tissue heart valve based upon glutaraldehyde-treated pig valves, and his company and valve were acquired by Johnson & Johnson in 1979.

It was in late 1967, early 1968, that a strange finding emerged from the pioneering “Gott Ring” studies in canine vena cavae being done at Johns Hopkins Hospital, in Baltimore, Maryland. Classical inorganic materials scientist, Dr. Jack Bokros, of San Diego’s General Atomic Corp., had become “accidentally” a co-worker of Vincent Gott through Bill Ellis, who had read an abstract in Carbon by Gott et al.: “The Anticlot Properties of Graphite Coatings on Artificial Heart Valves” (Gott et al., 1964) and informed Bokros that colloidal graphite (carbon) coatings were being used as a “base” for blood anti-coagulant on Dr. Vincent Gott’s short-ring implants, to good effect. Ellis was a bit offended by the choice of a commercial product which was much less pure than what had recently been developed at General Atomic, and so samples of the newly developed PyC at General Atomic Corp. were sent to Dr. Gott as “positive” (clotprovoking) controls. I was one of the “boys in the back room,” providing surface analyses of all proposed new blood-contact materials. Early measurements predicted that these PyC rings would immediately cause both thrombosis and coagulation of slow-flowing dog blood. We were all surprised when the naked PyC rings stayed clean for 2-hours (most everything else had clotted solid by then), and then were amazed when the dogs were brought back after 2-weeks “at the farm!” The rings were still clean, clearly violating all that we had considered ample predictive data to the contrary. Carrying the tests further, I was able to show that the PyC material uniquely bound one of the blood’s proteins in a configuration that expressed an outermost “Critical Surface Tension” in a zone prior identified as triggering the least thrombosis and coagulation. It was in that zone, and specifically for the PyC, that Surgeon Eugene Bernstein and Fred Schoen had shown that such PyC leaflets on a pioneering centrifugal blood pump least distorted attached blood platelets and thus did not trigger viscous metamorphosis and thrombus growth. Our colleague, Dr. Emery Nyilas, then working at AVCO-Everett Corporation near Boston, showed that the “heat of adsorption” was minimal during key blood protein trials (carried out in the middle of the night to avoid traffic vibrations) by micro-calorimetry. Drs. Andrade and Kim subsequently showed that the mode of protein adsorption and not the quantity of protein adsorbed on foreign surfaces was a key, consistent with Nyilas’ micro-calorimetric studies. PyC adsorbed a layer of blood proteins rapidly without the expected denaturing of proteins on blood contacting surfaces which generally triggered the clotting cascade.

30 SEC T I O N 1 .1     Overview of Biomaterials

Designed Biomaterials In contrast to the biomaterials of the surgeon-hero era, when largely off-the-shelf materials were used to fabricate medical devices, the 1960s on saw the development of materials designed specifically for biomaterials applications. Here are some key classes of materials and their evolution from commodity materials to engineered/synthesized biomaterials.

Silicones • Figure 1.1.2.10  An in vivo flow cell made of pyrolytic carbon, postcanine implantation.

Many others joined in, General Atomic became Gulf General Atomic which spun-off CarboMedics which inspired Medical Carbon Research Institute—and a large fraction of the world’s synthetic heart valves have since been rendered from those pioneering contributions between academia and industry. The recently announced sale of St. Jude Medical to Abbott Laboratories for $25 Billion is predominantly owing to the initial success of St. Jude with these PyC heart valves, and subsequent copying of the technology. Fig. 1.1.2.10 is a contemporary photo of a PyC “flow cell” that had been implanted in a 27 kg dog by Dr. Gott, to establish the details of thromboresistance now shown in over 15 million successful valve implants. 

Drug Delivery and Controlled Release Throughout most of history drugs were administered orally or by hypodermic syringe. In general, there was no effort to modulate the rate of uptake of the drug into the body. In 1949, Dale Wurster invented what is now known as the Wurster process that permitted pills and tablets to be encapsulated and therefore slow their release rate. However, modern ideas of controlled release can be traced to a medical doctor, Judah Folkman. Dr. Folkman noted that dyes penetrated deeply into silicone rubber, and he surmised from this that drugs might do the same. He sealed isoproterenol (a drug used to treat heart block) into silicone tubes, and implanted these into the hearts of dogs (Folkman and Long, 1964). He noted the delayed release and later applied this same idea to delivering a birth control steroid. He donated this development, patent free, to the World Population Council. An entrepreneur and chemist, Alejandro Zaffaroni, heard of the Folkman work and launched a company in 1970, Alza (originally called Pharmetrics), to develop these ideas for the pharmaceutical industry. The company developed families of new polymers for controlled release, and also novel delivery strategies. Alza was a leader in launching this new field that is so important today; further details on the field of controlled release are provided in an excellent historical overview (Hoffman, 2008). 

Although the class of polymers known as silicones has been explored for many years, it was not until the early 1940s that Eugene Rochow of General Electric pioneered the scale-up and manufacture of commercial silicones via the reaction of methyl chloride with silicone in the presence of catalysts. In Rochow’s 1946 book, The Chemistry of Silicones (John Wiley & Sons, Publishers), he comments anecdotally on the low toxicity of silicones, but did not propose medical applications. Possibly the first report of silicones for implantation was by Lahey (1946) (see also Chapter 1.3.2B). The potential for medical uses of these materials was realized shortly after this. In a 1954 book on silicones, McGregor has a whole chapter titled “Physiological Response to Silicones.” Toxicological studies were cited suggesting to McGregor that the quantities of silicones that humans might take into their bodies should be “entirely harmless.” He mentions, without citation, the application of silicone rubber in artificial kidneys. Silicone-coated rubber grids were also used to support a dialysis membrane (Skeggs and Leonards, 1948). Many other early applications of silicones in medicine are cited in Chapter 1.3.2B. 

Polyurethanes Polyurethane was invented by Otto Bayer and colleagues in Germany in 1937. The chemistry of polyurethanes intrinsically offered a wide range of synthetic options leading to hard plastics, flexible films, or elastomers (Chapter 1.3.2A). Interestingly, this was the first class of polymers to exhibit rubber elasticity without covalent cross-linking. As early as 1959, polyurethanes were explored for biomedical applications, specifically heart valves (Akutsu et al., 1959). In the mid-1960s a class of segmented polyurethanes was developed that showed both good biocompatibility and outstanding flex life in biological solutions at 37°C (Boretos and Pierce, 1967). Sold under the name Biomer by Ethicon and based upon DuPont Lycra, these segmented polyurethanes comprised the pump diaphragms of the Jarvik 7 hearts that were implanted in seven humans. 

Teflon DuPont chemist Roy Plunkett discovered a remarkably inert polymer, Teflon (polytetrafluoroethylene) (PTFE), in 1938. William L. Gore and his wife, Vieve, started a company in 1958 to apply Teflon for wire insulation. In 1969, their

CHAPTER 1.1.2   A History of Biomaterials

son Bob found that Teflon, if heated and stretched, forms a porous membrane with attractive physical and chemical properties. Bill Gore tells the story that, on a chairlift at a ski resort, he pulled from his parka pocket a piece of porous Teflon tubing to show to his fellow ski lift passenger. The skier was a physician and asked for a specimen to try as a vascular prosthesis. Now, Goretex porous Teflon and similar expanded PTFEs are the leading synthetic vascular grafts, and are also used in numerous other applications in surgery and biotechnology (Chapters 1.3.2C and 2.5.2B). 

Hydrogels Hydrogels have been found in nature since life on earth evolved. Bacterial biofilms, hydrated extracellular matrix components, and plant structures are ubiquitous, waterswollen motifs in nature. Gelatin and agar were also known and used for various applications early in human history. But the modern history of hydrogels as a class of materials designed for medical applications can be accurately traced. In 1936, DuPont scientists published a paper on recently synthesized methacrylic polymers. In this paper, poly(2hydroxyethyl methacrylate) (polyHEMA) was mentioned. It was briefly described as a hard, brittle, glassy polymer, and clearly was not considered of importance. After that paper, polyHEMA was essentially forgotten until 1960. Wichterle and Lim published a paper in Nature describing the polymerization of HEMA monomer and a cross-linking agent in the presence of water and other solvents (Wichterle and Lim, 1960). Instead of a brittle polymer, they obtained a soft, water-swollen, elastic, clear gel. Wichterle went on

31

to develop an apparatus (built originally from a children’s construction set; Fig. 1.1.2.11) for centrifugally casting the hydrogel into contact lenses of the appropriate refractive power. This innovation led to the soft contact lens industry, and to the modern field of biomedical hydrogels as we know them today. Interest and applications for hydrogels have steadily grown over the years, and these are described in detail in Chapter 1.3.2E. Important early applications included acrylamide gels for electrophoresis, poly(vinyl alcohol) porous sponges (Ivalon) as implants, many hydrogel formulations as soft contact lenses, and alginate gels for cell encapsulation. 

Poly(Ethylene Glycol) Poly(ethylene glycol) (PEG), also called poly(ethylene oxide), in its high molecular weight form, can be categorized as a hydrogel, especially when the chains are crosslinked. However, PEG has many other applications and implementations. It is so widely used today that its history is best discussed in its own section. The low reactivity of PEG with living organisms has been known since at least 1944, when it was examined as a possible vehicle for intravenously administering fat-­soluble hormones (Friedman, 1944). In the mid-1970s, Frank Davis and colleagues (Abuchowski et al., 1977) discovered that if PEG chains were attached to enzymes and proteins, they would a have a much longer functional residence time in  vivo than biomolecules that were not PEGylated. Professor Edward Merrill of MIT, based upon what he called “various bits of evidence” from the literature, concluded

• Figure 1.1.2.11  Left: Otto Wichterle (1913–98). (Wikipedia.) Right: The centrifugal casting apparatus Wichterle used to create the first soft, hydrogel contact lenses. (Photograph by Jan Suchy, Wikipedia public domain.)

32 SEC T I O N 1 .1     Overview of Biomaterials

that surface-immobilized PEG would resist protein and cell pickup. The experimental results from his research group in the early 1980s bore out this conclusion (Merrill, 1992). The application of PEGs to a wide range of biomedical problems was significantly accelerated by the synthetic chemistry developments of Dr. Milton Harris while at the University of Alabama, Huntsville. 

end of World War II, titanium metallurgy methods and titanium materials made their way from military application to peacetime uses. By 1940, satisfactory results had already been achieved with titanium implants (Bothe et al., 1940). The major breakthrough in the use of titanium for bony tissue implants was the Brånemark discovery of osseointegration, described earlier in the section on dental implants. 

Poly(Lactic-Glycolic Acid)

Bioglass

Although originally discovered in 1833, the anionic polymerization from the cyclic lactide monomer in the early 1960s made creating materials with mechanical properties comparable to Dacron possible. The first publication on the application of poly(lactic acid) in medicine may be by Kulkarni et  al. (1966). This group demonstrated that the polymer degraded slowly after implantation in guinea pigs or rats, and was well tolerated by the organisms. Cutright et  al. (1971) were the first to apply this polymer for orthopedic fixation. Poly(glycolic acid) and copolymers of lactic and glycolic acid were subsequently developed. Early clinical applications of polymers in this family were for sutures, based upon the work of Joe Frazza and Ed Schmitt at David & Geck, Inc (Frazza and Schmitt, 1971). The glycolic acid/ lactic acid polymers have also been widely applied for controlled release of drugs and proteins. Professor Robert Langer’s group at MIT was the leader in developing these polymers in the form of porous scaffolds for tissue engineering (Langer and Vacanti, 1993). 

Bioglass is important to biomaterials as one of the first completely synthetic materials that seamlessly bonds to bone. It was developed by Professor Larry Hench and colleagues. In 1967 Hench was Assistant Professor at the University of Florida. At that time his work focused on glass materials and their interaction with nuclear radiation. In August of that year, he shared a bus ride to an Army Materials Conference in Sagamore, New York, with a US Army colonel who had just returned from Vietnam where he was in charge of supplies to 15 MASH units. This colonel was not particularly interested in the radiation resistance of glass. Rather, he challenged Hench with the following: hundreds of limbs a week in Vietnam were being amputated because the body was found to reject the metals and polymer materials used to repair the body. “If you can make a material that will resist gamma rays why not make a material the body won’t resist?” Hench returned from the conference and wrote a proposal to the US Army Medical R and D Command. In October 1969 the project was funded to test the hypothesis that silicate-based glasses and glass-ceramics containing critical amounts of Ca and P ions would not be rejected by bone. In November 1969 Hench made small rectangles of what he called 45S5 glass (44.5 wt% SiO2), and Ted Greenlee, Assistant Professor of Orthopedic Surgery at the University of Florida, implanted them in rat femurs at the VA Hospital in Gainesville. Six weeks later Greenlee called: “Larry, what are those samples you gave me? They will not come out of the bone. I have pulled on them, I have pushed on them, I have cracked the bone and they are still bonded in place.” Bioglass was born, and with the first composition studied! Later studies by Hench using surface analysis equipment showed that the surface of the bioglass, in biological fluids, transformed from a silicate-rich composition to a phosphate-rich structure, possibly hydroxyapatite (Clark et al., 1976). 

Hydroxyapatite Hydroxyapatite is one of the most widely studied materials for healing in bone. It is a naturally occurring mineral, a component of bone, and a synthesized material with wide application in medicine. Hydroxyapatite can be easily made as a powder. One of the first papers to describe biomedical applications of this material was by Levitt et al. (1969), in which they hot pressed the hydroxyapatite power into useful shapes for biological experimentation. From this early appreciation of the materials science aspect of a natural biomineral, a literature of thousands of papers has evolved. In fact, the nacre implant described in the prehistory section may owe its effectiveness to hydroxyapatite—it has been shown that the calcium carbonate of nacre can transform in phosphate solutions to hydroxyapatite (Ni and Ratner, 2003). 

Titanium In 1791, William Gregor, a Cornish amateur chemist, used a magnet to extract the ore that we now know as ilmenite from a local river. He then extracted the iron from this black powder with hydrochloric acid, and was left with a residue that was the impure oxide of titanium. After 1932, a process developed by William Kroll permitted the commercial extraction of titanium from mineral sources. At the

The Contemporary Era (Modern Biology and Modern Materials) It is probable that the modern era in the history of biomaterials, biomaterials engineered to control specific biological reactions, was ushered in by rapid developments in modern biology. As illustrated in Fig. 1.1.1.2, major advances in the medical device field continue to be made with materials that could be considered first generation with the incorporation of concepts and approaches from the later generations. In the 1960s, when the field of

CHAPTER 1.1.2   A History of Biomaterials

biomaterials was laying down its foundational principles and ideas, concepts such as cell-surface receptors, growth factors, nuclear control of protein expression and phenotype, cell attachment proteins, stem cells, and gene delivery were either controversial observations or not yet discovered. Thus pioneers in the field could not have designed materials with these ideas in mind. It is to the credit of the biomaterials community that it has been quick to embrace and exploit new ideas from biology. Similarly, new ideas from materials science such as phase separation, anodization, self-assembly, surface modification, and surface analysis were quickly assimilated into the biomaterial scientists’ toolbox and vocabulary. A few of the important ideas in biomaterials literature that set the stage for the biomaterials science we see today are useful to list: • Protein adsorption • Biospecific biomaterials • Nonfouling materials • Healing and the foreign-body reaction • Controlled release (also, programmed release and release on demand) • Tissue engineering • Immunoengineering • Regenerative materials • Nanotechnology Since these topics are addressed later in some detail in Biomaterials Science: An Introduction to Materials in Medicine, fourth edition, they will not be expanded upon in this history section. Still, it is important to appreciate the intellectual leadership of many researchers who promoted these ideas that comprise modern biomaterials—this is part of a recent history of biomaterials that will someday be completed. We practice biomaterials today immersed within an evolving history. 

Conclusions Biomaterials have progressed from adventurous practices by surgeon-heroes, sometimes working with engineers, to a field dominated by engineers, chemists, and physicists, to our modern era innovations with biologists and bioengineers as the key players. At the moment you are reading the fourth edition of Biomaterials Science: An Introduction to Materials in Medicine, many individuals who were biomaterials pioneers in the formative days of the field are well into their eighth or ninth decades of life. A number of leaders of biomaterials, pioneers who spearheaded the field with vision, creativity, and integrity, have passed away. Yet, since biomaterials is relatively a new field, the first-hand accounts of its roots are still available. We encourage readers of this book to document their conversations with pioneers of the field (many of whom still attend biomaterials conferences), so that the exciting stories and accidental discoveries that led to the successful and intellectually stimulating field we see today are not lost.

33

References Abuchowski, A., McCoy, J.R., Palczuk, N.C., van Es, T., Davis, F.F., 1977. Effect of covalent attachment of polyethylene glycol on immunogenicity and circulating life of bovine liver catalase. J. Biol. Chem. 252 (11), 3582–3586. Akutsu, T., Dreyer, B., Kolff, W.J., 1959. Polyurethane artificial heart valves in animals. J. Appl. Physiol. 14, 1045–1048. Apple, D.J., Trivedi, R.H., 2002. Sir Nicholas Harold Ridley, Kt, MD, FRCS, FRS. Arch. Ophthalmol. 120 (9), 1198–1202. Baier, R.E., 2016. Surprising blood Compatibility of pyrolytic carbon!. Biomater. Forum 38 (4), 4–5. Blagg, C., 1998. Development of ethical concepts in dialysis: Seattle in the 1960s. Nephrology 4 (4), 235–238. Bobbio, A., 1972. The first endosseous alloplastic implant in the history of man. Bull. Hist. Dent. 20, 1–6. Bondurant, S., Ernster, V., Herdman, R. (Eds.), 1999. Safety of Silicone Breast Implants. National Academies Press, Washington DC. Boretos, J.W., Pierce, W.S., 1967. Segmented polyurethane: a new elastomer for biomedical applications. Science 158, 1481– 1482. Bothe, R.T., Beaton, L.E., Davenport, H.A., 1940. Reaction of bone to multiple metallic implants. Surg. Gynecol. Obstet. 71, 598– 602. Branemark, P.I., Breine, U., Johansson, B., Roylance, P.J., Röckert, H., Yoffey, J.M., 1964. Regeneration of bone marrow. Acta Anat. 59, 1–46. Clark, A.E., Hench, L.L., Paschall, H.A., 1976. The influence of surface chemistry on implant interface histology: a theoretical basis for implant materials selection. J. Biomed. Mater. Res. 10, 161–177. Crubezy, E., Murail, P., Girard, L., Bernadou, J.-P., 1998. False teeth of the Roman world. Nature 391, 29. Cutright, D.E., Hunsuck, E.E., Beasley, J.D., 1971. Fracture reduction using a biodegradable material, polylactic acid. J. Oral Surg. 29, 393–397. Egdahl, R.H., Hume, D.M., Schlang, H.A., 1954. Plastic venous prostheses. Surg. Forum 5, 235–241. Folkman, J., Long, D.M., 1964. The use of silicone rubber as a carrier for prolonged drug therapy. J. Surg. Res. 4, 139–142. Frazza, E., Schmitt, E., 1971. A new absorbable suture. J. Biomed. Mater. Res. 5 (2), 43–58. Friedman, M., 1944. A vehicle for the intravenous administration of fat soluble hormones. J. Lab. Clin. Med. 29, 530–531. Gott, V.L., Whiffen, J.D., Dutton, R.C., Koepke, D.E., Daggett, R.L., Young, W.P., 1964. The anticlot properties of graphite coating on artificial heart valves. Carbon 1, 378. Hoffman, A., 2008. The origins and evolution of “controlled” drug delivery systems. J. Control. Release 132 (3), 153–163. Ingraham, F.D., Alexander Jr., E., Matson, D.D., 1947. Polyethylene, a new synthetic plastic for use in surgery. J. Am. Med. Assoc. 135 (2), 82–87. Kolff, W.J., 1998. Early years of artificial organs at the Cleveland Clinic Part II: open heart surgery and artificial hearts. ASAIO J. 44 (3), 123–128. Kulkarni, R.K., Pani, K.C., Neuman, C., Leonard, F., 1966. Polylactic acid for surgical implants. Arch. Surg. 93, 839–843. Lahey, F.H., 1946. Comments (discussion) made following the speech “Results from using Vitallium tubes in biliary surgery,” by Pearse H. E. before the American Surgical Association, Hot Springs, VA. Ann. Surg. 124, 1027. Langer, R., Vacanti, J., 1993. Tissue engineering. Science 260, 920–926.

34 SEC T I O N 1 .1     Overview of Biomaterials

LeVeen, H.H., Barberio, J.R., 1949. Tissue reaction to plastics used in surgery with special reference to Teflon. Ann. Surg. 129 (1), 74–84. Levitt, S.R., Crayton, P.H., Monroe, E.A., Condrate, R.A., 1969. Forming methods for apatite prostheses. J. Biomed. Mater. Res. 3, 683–684. McGregor, R.R., 1954. Silicones and Their Uses. McGraw Hill Book Company, Inc., New York. Merrill, E.W., 1992. Poly(ethylene oxide) and blood contact. In: Harris, J.M. (Ed.), Poly(ethylene Glycol) Chemistry: Biotechnical and Biomedical Applications. Plenum Press, New York, pp. 199–220.

Ni, M., Ratner, B.D., 2003. Nacre surface transformation to hydroxyapatite in a phosphate buffer solution. Biomaterials 24, 4323–4331. Rob, C., 1958. Vascular surgery. In: Gillis, L. (Ed.), Modern Trends in Surgical Materials. Butterworth & Co, London, pp. 175–185. Skeggs, L.T., Leonards, J.R., 1948. Studies on an artificial kidney: preliminary results with a new type of continuous dialyzer. Science 108, 212. Wichterle, O., Lim, D., 1960. Hydrophilic gels for biological use. Nature 185, 117–118. Willyard, C., 2016. Timeline: regrowing the body. Nature 540, S50.

S E C TI ON 1 .2

Properties of Biomaterials

1.2.1

Introduction: Properties of Materials— the Palette of the Biomaterials Engineer JACK E. LEMONS 1 , GUIGEN ZHANG 2 1Schools

of Dentistry, Medicine and Engineering, University of Alabama at Birmingham, Birmingham, AL, United States 2F

Joseph Halcomb III, M.D. Department of Biomedical Engineering, University of Kentucky, Lexington, KY, United States

T

he platform, or palette, upon which the biomaterials engineer arranges information into parts for subsequent blending has expanded and evolved significantly in content over the past several decades. The depth and breadth of what is now included on this palette go well beyond expectations expressed by the founding members of the Society for Biomaterials in the late 1960s and early 1970s. The Science of Biomaterials has expanded from dealing with the fundamental aspects of the physical, mechanical, chemical, electrical, electrochemical, thermal, optical, and biological (compatibility) properties of biomaterials of natural or synthetic origin, to the design principles of tissue/cell/molecule-synthesized or -engineered biomaterials and nanomaterials. Because of that, the methods for measuring and analyzing materials’ properties have also evolved. Following the recognition of the need by the pioneers of the biomaterials discipline in the 1960s, one focus of the science of biomaterials has been the fundamental structure versus property relationships leading to in  vivo biocompatibility. These relationships, and the supporting scientific information, have changed with time and experience, especially as the biological and clinical disciplines have also expanded and evolved. For example, considerations for biocompatibility are very different for biomaterials listed within biotolerant, surface bioactive, and biodegradable categories. This shift of emphasis is reflected in the progression of content of the first four editions of this book. For

example, initial considerations focused on materials were based primarily on substances of the metallics, ceramics, and polymerics available within various biomedical applications. Thus the emphasis in the first edition was on materials of natural and synthetic origin, including metals, ceramics, polymers, and composites, and the underlying science leading to biomedical applications. The second and third editions represent the transitions from combination products to the new areas of bioactives and biodegradables and tissueengineered constructs. As an integrated, comprehensive, and authoritative text, this fourth edition reflects a broader range of biomaterials as well as basic properties of new classes of biomaterials that possess more of a resemblance to complex biological constructs formed by nature and constantly remodeled by biology in response to biophysical demands and biomechanical environments. Considering relationships between biomaterial and biological systems (the interface) and the dynamics of change from nanoseconds to years, we now better understand many mechanisms of interaction at the dimensions and concentrations used to describe interactions of atoms and molecules. It is also realized that all biomaterial and hostenvironment interactions play a role in the broader aspects of biocompatibility, especially the functionality and longevity of implant devices. In this regard, Part 1 on “Materials Science and Engineering” emphasizes the more basic information on the bulk and surface properties of synthetic and natural origin biomaterials. Critical aspects of constitution 35

36 SEC T I O N 1 . 2    Properties of Biomaterials

(chemistry) and structure (nano-, micro-, and macrodimension) relationships are presented as related to properties of implant systems. These basic considerations include the nature of the matters (Chapter 1.2.2), bulk (Chapter 1.2.3), and surface (Chapter 1.2.4) properties, and the role of water in biomaterials (Chapter 1.2.5). The science of interactions of synthetic and natural substances with water is recognized as one of the key aspects of surgical implant biocompatibilities. This is especially important for the evolution of the

discipline to include new-generation biomaterials needed for future implant applications. In short, the content of this Part 1 is broadly applicable to all parts of the fourth edition. Therefore students are advised to always consider the basic principles as provided in this section. This has been recognized as critical to the education of a specialist in biomaterials science leading to the selection of biomaterials for medical treatments utilizing all types of implant devices.

1.2.2

The Nature of Matter and Materials BUDDY D. RATNER Bioengineering and Chemical Engineering, Director of University of Washington Engineered Biomaterials (UWEB), Seattle, WA, United States

Introduction Biomaterials are materials.

Biomaterials are materials. What are materials and how are they structured? This is the subject of this chapter, a lead into subsequent chapters with discussions of the bulk (including mechanical) properties of materials, surface properties, and the role of water (since biomaterials most commonly function in an aqueous environment, and that environment can alter both the nature of the material and the interactions that occur with the material). This chapter also considers fundamental atomic and molecular interactions that underlie subsequent chapters addressing specific classes of materials relevant to biomaterials (polymers, metals, ceramics, natural materials, and particulates). 

Atoms and Molecules The key to understanding matter is to understand attractive and interactive forces between atoms.

This “Nature of Matter” section aims to communicate an understanding of the basic structure of materials that will drive their properties—both the mechanical properties important for specific applications (strong, elastic, ductile, permeable, etc.), and the surface properties that will mediate reactions with the external biological environment. The key to understanding matter is to understand attractive and interactive forces between atoms. Argon is a gas at room temperature—it must be cooled to extremely low temperatures to transition it into liquid form. An argon atom interacts (attracts) very, very weakly with another argon atom—so at room temperature, thermal fluctuations that randomly propel the atoms exceed attractive forces that might result in coalescence to a solid material.

A titanium atom strongly interacts with another titanium atom. Extremely high temperatures are required to vaporize titanium and liberate those atoms from each other. The understanding of matter is an appreciation of interactive forces between atoms. What holds those atoms and molecules together to make a strong nylon fiber or a cell membrane, or a hard, brittle hydroxyapatite ceramic, or a sheet of gold, or a drop of water? Even in the early 18th century, Isaac Newton was pondering this issue: “There are therefore Agents in Nature able to make the Particles of Bodies stick together by very strong Attractions.” Entropy consideration would say these molecules and atoms should “fly apart” to increase randomness. However, there is an energy term contributing to the stability of the ensemble leading to a negative Gibbs free energy, which, according to the second law of thermodynamics, should make such solids energetically favorable (of course, we intuitively know this). Thus we must examine this energy term. We know of just four attractive forces in this universe: • Gravitational • Weak nuclear • Strong nuclear • Electromagnetic Gravity holds us to the surface of the planet Earth (a massive body), but the gravitational potential energy of two argon atoms is only about 10−52 J, 30 orders of magnitude weaker than is observed for intermolecular forces. The weak nuclear force and the strong nuclear force are only significant over 10–4 nm—but molecular dimensions are 5 × 10−1 nm. So these forces do not explain what holds atoms together. This leaves, by default, electromagnetic forces (positive charge attracts negative charge). Electromagnetic forces have appropriate magnitudes and distance dependencies to justify why atoms interact. Interactions can be weak, leading to liquids, or stronger, leading to solids. 37

38 SEC T I O N 1 . 2    Properties of Biomaterials

TABLE 1.2.2.1    Forces That Hold Atoms Together

Interatomic Force

Explanation

Relative Strength

Examples

Van der Waals interactions

Transient fluctuations in the spatial localization of electron clouds surrounding atoms lead to transient positive and negative charges, and consequent interactive forces in molecules would seem to have no permanent polarity

Weak

Argon at cryogenic temperatures Polyethylene (the forces that hold the chains together to make a solid)

Ionic

Atoms with a permanent positive (+) charge attract atoms with a permanent negative (−) charge

Very strong

NaCl CaCl2

Hydrogen (H) bonding

The interaction of a covalently bound hydrogen with an electronegative atom, such as oxygen or fluorine

Medium

Water ice Nylon (the forces that hold the chains together to make a strong, high-melting point solid)

Metallic

The attractive force between a “sea” of positively charged atoms and delocalized electrons

Medium–strong

Gold Titanium metal

Covalent

A sharing of electrons between two atoms

Strong

The carbon–carbon bond Cross-links in a polyacrylamide hydrogel

_

+

_

+

_

+ Na+

Cl–

(A)

(B)

(C)

Na+

Cl–



Figure 1.2.2.1 (A) Consider the electron clouds (charge density in space) of two atoms or molecules, both without permanent dipole moments. (B) Electron clouds are continuously in motion and can shift to one side of the atom or molecule; therefore, at any moment, the atoms or molecules can create a “fluctuating instantaneous dipole.” (C) The “fluctuating instantaneous dipole” in one molecule electrostatically induces such an “instantaneous dipole” in the next molecule.

Na N a+ Cl–

Cll– C Na+

• Figure 1.2.2.2  The unit cell of a sodium chloride crystal illustrating the

Electromagnetic forces manifest themselves in a number of ways. The types of interactions usually observed between atoms (all explained by electrostatics) are summarized in Table 1.2.2.1. We consider here van der Waals forces (also called induction or dispersion forces), ionic forces, hydrogen bonding, metallic forces, and covalent interactions. Van der Waals or dispersion forces rationalize the interaction of two atoms or molecules, each without a dipole (no plus or minus faces to the molecules). For example, argon atoms can be liquefied at low temperature. Why should this happen? Why should argon atoms want to interact with each other enough to form a liquid? Fig. 1.2.2.1 explains the origin of dispersive forces. Such forces are important, as they dictate the properties of many materials (for example, some polymers such as polyethylene, which has no obvious dipole), but they also explain why the lipids in cell membranes assemble into the bilayer structure described later in this textbook. A typical van der Waals interactive force (for example,

plus–minus electrostatic interactions.

CH4 ⋯ CH4) is about 9 kJ/mol. The Ar ⋯ Ar interactive force is approximately 1 kJ/mol. Ionic forces are probably the easiest of the intermolecular forces to understand. Fig. 1.2.2.2 illustrates a unit cell of a sodium chloride crystal. The + and − charges are arrayed to achieve the closest interaction of opposite charges and the furthest separation of similar charges. This unit cell can be repeated over and over in space, and the forces that hold it together are the electrostatic interaction of a permanent + charge and a permanent − charge. Typical ionic bond strengths (for example, NaCl) are about 770 kJ/mol. Hydrogen bonding interactions are also straightforward to appreciate as electrostatic interactions. An electronegative element such as oxygen (it demands electrons) can distort the binding electron cloud from the hydrogen nucleus leaving the hydrogen (just a proton and electron) with less electron and thus more plus-charged proton.

CHAPTER 1.2.2   The Nature of Matter and Materials

– +



H O H

– +













H O

– Mg +

– –



– –



– –

Mg + – –



H

• Figure 1.2.2.3  A hydrogen bond between two water molecules.















Mg +

– –





Mg +





– –

– –

























39

– –

– –

– –



– –

Mg + –

• Figure 1.2.2.4  Metallic bonding in magnesium. The 12 electrons from each Mg atom are shared among positively charged nuclear cores (the single + charge on each magnesium atom in the figure is simply intended to indicate there is some degree of positive charge on each magnesium nuclear core).



Figure 1.2.2.5 Covalent bonding along a section of polyethylene chain. Carbons share pairs of electrons with each other, and each hydrogen shares an electron pair with carbon.

This somewhat positive charge will, in turn, then interact with an electronegative oxygen (Fig. 1.2.2.3). Typical hydrogen bond strengths (for example, O–H ⋯ H) are about 20 kJ/mol. Metallic bonding is explained by a delocalized “sea” of valence electrons with positively charged nuclear cores dispersed within it (Fig. 1.2.2.4). A single metallic bond is rarely discussed. The total interactive strength is realized through the multiplicity of the plus–minus interactions. The strength of this interaction can be expressed by heats of sublimation. For example, at 25°C, aluminum will have a heat of sublimation of 325 kJ/mol, while titanium will be about 475 kJ/mol. Covalent bonds are relatively strong bonds associated with the sharing of pairs of electrons between atoms (Fig. 1.2.2.5). Typical covalent bond strengths (for example, C–C) are about 350 kJ/mol. 

Molecular Assemblies Atoms can combine in defined ratios to form molecules (usually they combine with covalent bonds), or they can form cohesive assemblies of atoms (think of gold and metallic interactive forces, for example). Thus materials can be made of atoms or of molecules (i.e., covalently joined atoms). The difference between the dense, lubricious plastics used in orthopedics, the soft, elastic materials of catheters, and the hard, strong metals of a hip joint is associated with how those atoms and molecules are organized (due to attractive and interactive forces) in materials. Metals used in biomaterials applications can be strong, rigid, and brittle, or flexible and ductile. Again, the difference is largely how the atoms making up the metal are organized, and how strongly their atoms interact.

Molecules also organize or assemble. The widely varying properties of polymers are due to molecular organization. The assembly of lipid molecules to make a cell membrane or a microparticulate for drug delivery is another example of this organization. A key concept in appreciating the properties of materials is hierarchical structures. The smallest size scale that we need consider here in materials is atoms, typically about 0.2 nm in diameter. Atoms combine to form molecules with dimensions ranging from 1  to 100 nm (some large macromolecules). Molecules may assemble or order to form supramolecular structures with dimensions up to 1000 nm or more. These supramolecular structures may themselves organize in bundles, fibers, or larger assemblies with dimensions reaching into the range visible to the human eye. This concept of hierarchical structure is illustrated in Fig. 1.2.2.6, using collagen protein as the example. The single α-chains comprising the collagen triple helix would break under tension with an application of nanograms of force. On the other hand, the collagen fibers in a hierarchical structure such as a tendon can support many kilograms of force. Such hierarchical structures are noted frequently in both materials science and in biology.

 Surfaces As assemblies of atoms and/or molecules form, within the bulk of the material, each unit is uniformly “bathed” in a field of attractive forces of the types described in Table 1.2.2.1. However, those structural units that are at the surface are pulled upon asymmetrically by just the units beneath them. This asymmetric attraction distorts the electron distributions of the surface atoms or molecules, and gives rise to the phenomenon of surface energy, an excess energy associated with this imbalance. For this reason, surfaces always have unique reactivities and properties. This idea will be expanded upon in Chapter 1.2.4.

40 SEC T I O N 1 . 2    Properties of Biomaterials

Collagen molecules (triple helices) Collagen fibrils

Collagen fibers

-chains

0.5 m

• Figure 1.2.2.6  Collagen fibers make up many structures in the body (tendons, for example). Such ana-

tomical structures as tendons are comprised of collagen fibrils, formed of aligned bundles of collagen triple helices that are themselves made up of single collagen protein chains (α-chains). The α-chains are constructed of joined amino acid units, and the amino acids are molecules of carbon, oxygen, nitrogen, and hydrogen atoms in defined ratios and orders. (Illustration from Becker, W., 2002. The World of the Cell. Reprinted with permission of Pearson Education, Inc.)

 Conclusion In this section we reviewed the transition from chemistry to matter. Interacting assemblies of atoms and molecules comprise matter. Without matter, we cannot have biomaterials. Matter exists because of electrostatic forces—­ positive and negative charges, in all cases, hold atoms together. The strength of those interactions, associated with the magnitude of the charge on each atom, and the environment the atoms are in (water, air, etc.) ultimately dictate the properties of matter (a soft gel, a hard metal, etc.). Now that we have a general idea what “matter” is, we can take these concepts from physics and chemistry and

bring them to a consideration of the mechanical properties of materials, the surface properties of materials, and then into the specifics of polymers, metals, ceramics, and other types of materials.

Further Reading Barton, A., 1997. States of Matter, States of Mind. Institute of Physics Publishing, Bristol, UK. Becker, W.M., 2003. World of the cell. In: Kleinsmith, L.J., Hardin, J. (Eds.), fifth ed. Pearson Education, Inc., Upper Saddle River, NJ. Holden, A., 1992. The Nature of Solids. Dover Publications Inc, New York, NY.

1.2.3

Bulk Properties of Materials GUIGEN ZHANG 1 , CHRISTOPHER VINEY 2 1F

Joseph Halcomb III, M.D. Department of Biomedical Engineering, University of Kentucky, Lexington, KY, United States 2School

of Engineering, University of California at Merced, Merced, CA, United States

Introduction When describing a material, the term “bulk properties” is often used to differentiate, either intentionally or unintentionally, from the term “surface properties.” The importance of surfaces for biomaterials science is highlighted briefly in Chapter 1.2.2 and in great detail in Chapter 1.2.4. While it is important to know that the success or failure of many biomaterials depends on the physical and chemical characteristics of their surface because the surface properties dictate interactions between a material and its environment (thereby determining whether a permanently implanted material will be tolerated or rejected), it does not suggest that bulk properties of biomaterials are any less important. In fact, for almost all biomedical applications, either as short-term degradable applications or long-term structural or load-bearing applications, the bulk properties must meet the physical and/or mechanical demands of these applications over the desired time period, even if the surface properties are deemed to facilitate biocompatible material–tissue interactions. The requirements for biomaterials to exhibit certain bulk characteristics are multifaceted, including maintaining physical shapes, carrying mechanical loads, and possessing certain desirable electrochemical behavior, optic index, and/or thermal characteristics, among others. As discussed in Chapter 1.2.2, the differences in bulk properties of many materials such as metals, polymers, and ceramics are the results of different types of interatomic or intermolecular forces that hold atoms and molecules together. For example, metals and alloys are typically characterized by metallic bonds, ceramics are held together by ionic bonds, and polymers are predominately formed by covalent bonds. The differences in these bonding mechanisms and energies dictate the different bulk properties of these materials. This chapter provides an overview of some of the useful concepts related to the bulk properties of these materials, with an emphasis on mechanical properties. For more in-depth information, the reader is encouraged to consult books on materials science and strength of materials. 

Mechanical Variables and Mechanical Properties The mechanical properties of a material refer to the characteristic values of a material under various mechanical loading conditions. By characteristic values, we refer to quantitative measures of mechanical variables at which certain transitions in the material would occur. For example, stress and strain are two mechanical variables that are commonly referred to for a material, but they are not properties of the material. Only the stresses that cause certain structural changes, such as yielding or breaking, etc. will be regarded as properties. This can be likened to the fact that the temperature of water is a physical variable, but we do consider certain temperatures, at which transitions occur in the state of water, properties such as the freezing point or boiling point of water. Taking a cylindrical rod with uniform cross-section area (often called prismatic rod) as an example of a material structure, stress (σ) is defined as force per unit area, obtainable by dividing the applied force (F) by the cross-section area (A) of the rod, σ = F/A. Stress determined this way is commonly regarded as the nominal stress because the actual value of the area (cross-section) to which the load is applied will change as the sample deforms in response to the load. Nominal stress is often synonymously referred to as engineering stress. Strain (ε) is deformation per unit length, obtainable by dividing the deformation (δ) by the length of the rod (L), ε = δ/L. By the same argument, strain determined this way is regarded as the nominal strains or engineering strain because the actual value of the length will change as the sample deforms in response to the load. Stresses and strains determined with reference to the actual cross-section area (A) and length (L) are initially known as true stresses and true strains, respectively. Unfortunately, because A and L change continuously during the course of a mechanical test, true stresses and strains can be difficult to measure in practice. Therefore, nominal stress and strain are commonly used instead of true stress and strain. 41

42 SEC T I O N 1 . 2    Properties of Biomaterials

Stress (or strain) can be either normal or shear stress (or strain). Normal stress is determined by considering only the normal component of an arbitrary applied force, and shear stress by considering the tangent component of the force. The SI units of stress are newtons per square meter (N/m2) or pascals (Pa), and strain is dimensionless.

Five Types of Mechanical Loading A mechanical loading situation in a structure can be represented by a combination of several mechanical loads. In general, there are five basic types of loads as illustrated in Fig. 1.2.3.1 (Beer et al., 2014). When a pair of forces (F–F) with opposite directions is applied to a cylindrical rod along its axial direction, the situation is called axial loading. Under axial loading, we are more concerned with the deformations along the axial direction. If the axial load is making the rod longer, it is regarded as tensile loading, and if the axial load is making the rod shorter, it is compressive loading. When a pair of shear forces (V–V) with opposite directions

(A)



(B)

(C)

(D)

(E)

Figure 1.2.3.1  Five basic types of mechanical loading situations: (A) tension, (B) compression, (C) shearing, (D) torsion, and (E) bending.

is applied to the rod in a transverse direction, the loading condition is regarded as shearing loading with which we are mainly concerned about the transverse shear stress, strain, and deformation. When a pair of torques (T–T) with opposite directions is applied to the rod with respect to its axial orientation, the rod is regarded as under torsion loading in which we are concerned with twisting stress, strain, and angle. Finally, when a pair of moments (M–M) with opposite directions is applied to the rod, it is regarded as under bending. In a bending situation, we are more concerned about bending-induced tensile, compressive, and shear stresses and strains, as well as flexure deformation. Note that the term load is used here to represent all these forces, torques, and moments. Because of that, these loads will have different units, for instance, tensile, compressive, and shearing forces carry the units of newtons (N), and torque and moment have the units of newtons × meter (Nm). An actual structure is likely subjected to a combined loading situation consisting of several or all of these five types of basic loading. Very often, we impose certain mechanical constraints to the structure such that we can focus on stresses, strains, and deformations under the mechanical load and ignore other mechanical consequences such as translation and rotation, motion (linear and rotational velocity) and acceleration, among others. 

From External Loads to Internal Loads and Stresses In analyzing the mechanical properties of a material structure, the external loading condition (P) is first considered to determine the equivalent forces, torques, and moments at selected internal section planes. In a simplified twodimensional (2D) view for example, in section plane 1 (Fig. 1.2.3.2A), the equivalent forces may consist of a compressive force (F), a shear force (V), and a bending moment

• Figure 1.2.3.2  Two-dimensional illustration of how external mechanical loads are converted to equivalent

internal forces, torques, and moments on different section planes (A), and internal stresses in a 2D (B) and 3D sense (C).

CHAPTER 1.2.3   Bulk Properties of Materials

(M), but in section plane 2, an additional torque will have to be considered aside from the compressive and shear forces and moment. With these sectional forces, moments, and torques, the resulting internal stresses and strains as well as deformations can be obtained. As depicted by the three 2D stress elements shown in Fig. 1.2.3.2B, stress elements 1 and 3 are taken at the upper and lower surfaces of the neck region, where only tensile stress exists on stress element 1 and compressive stress on stress element 3. On a stress element isolated from an interior location (e.g., stress element 2), however, a much more complex stress state could exist. In a three-dimensional (3D) sense, we will have to deal with stress cubes on which six stress components would exist including σx, σy, σz, τxy, τyz, and τxz. 

Linear and Nonlinear Relationship, Elastic and Plastic Behavior In dealing with materials' mechanical properties, we always encounter terms like linear and nonlinear relationship, and elastic and plastic behavior, and very often we tend to (wrongly) associate a linear relationship with elastic behavior and a nonlinear relationship with plastic behavior. The terms “linear” and “nonlinear” are often used to refer to a relationship between two variables such as in a load–displacement curve or in a stress–strain curve. When these relationships are straight lines, we call them linear, and when they are curved lines, we call them nonlinear. The terms “elastic” and “plastic,” however, refer to a material's deformational behavior, specifically the ability (or inability) to regain the original shape after the removal of loads. When a material structure regains its original shape (all deformations vanish) after the removal of loads, we consider it deforming with elastic behavior, and when it does not regain its original shape, we call it deforming with plastic behavior. A linear relationship is not necessarily always associated with elastic behavior and a nonlinear relationship plastic behavior. In fact, these terms are often used in combinations to describe a certain mechanical property or behavior (Zhang, 2017). For example, a material structure can exhibit elastic behavior with either linear or nonlinear relationship between load and displacement. As illustrated in

43

Fig. 1.2.3.3, linear elastic behavior (Fig. 1.2.3.3A) is one in which the load–displacement curve is a straight line (hence linear) and the loading and unloading curves follow the same trace such that the displacement will vanish (hence elastic) when the applied load is removed. Nonlinear elastic behavior (Fig. 1.2.3.3B) describes a curved load–displacement relationship, but the displacement induced by loading will vanish after the removal of the load. As illustrated in Fig. 1.2.3.3C, the linear elastic behavior of a material typically occurs at the beginning of the loading stage when the induced displacement is extremely small. When the displacement increases slightly, the material may exhibit nonlinear but elastic behavior. This behavior is sometimes referred to as geometric nonlinearity. When the displacement increases further, the material will reach an elastic yielding zone where it still exhibits elastic behavior but with reduced rigidity. As the displacement continues to increase the material will enter a plastic yielding zone in which the displacement or deformation will keep increasing without needing any further increase in loads. Fig. 1.2.3.3D shows the corresponding stress–strain curve derived conceptually from the load–displacement curve shown in Fig. 1.2.3.3C by dividing the load by the cross-section area and the displacement by the length. It presents a different look at the mechanical properties from a pure material’s perspective by minimizing geometric-related influences. By convention, the stress is plotted vertically and the strain is plotted horizontally. When the material behaves elastically, in both a linear and nonlinear manner, the stress– strain curve typically exhibits a straight-line relationship with a constant slope E1 (which is commonly known as the modulus of elasticity or Young's modulus; more discussion in the next section). When the stress in the material reaches its elastic yielding point σy it will start to exhibit reduced elastic property with decreased slope E2, where E2  SO4− 2 > S2 O3− 2 > H2 PO4− > F − > Cl − > Br − > NO3− > I − > ClO4− > SCN −

Ions to the left of the series are referred to as kosmotropes (“order makers”)—they generally precipitate proteins

79

from solution and inhibit denaturation. Ions on the right side of the series are considered chaotropes (“disorder makers”), and these are more denaturing. Note that chloride is roughly at the center of the series, and thus might be expected to induce little change in proteins and water. Though earlier theories on the Hofmeister series suggested that ions to the left enhanced water structure while those to the right destructured or disordered water, this concept is not fully supported by recent experiments. The Hofmeister effect is impacted by the degree of hydration of ions in water, which cation is involved, and specific interactions of ions with solutes. Review articles that discuss the complexities of the Hofmeister effect and consider contemporary experimental and theoretical work are available (Zhang and Cremer, 2010; Paschek and Ludwig, 2011). The “swelling” water entrained in gels (hydrogels, see Chapter 1.3.2E) is thought to be in three possible forms. Different nomenclatures are used to describe these forms, but basically three states have been proposed: (1) free water (similar to bulk water); (2) tightly bound water (more structured and with limited mobility); and (3) intermediate water (with characteristics of both free and bound water) (Jhon and Andrade, 1973; Akaike et  al., 1979). Below, when the hydrophobic effect is discussed, other possibilities will be suggested. Evidence for different forms of water in gels continues to accumulate (Sekine and Ikeda-Fukazawa, 2009) and implications for biomaterials and biocompatibility have been discussed (Tsuruta, 2010). The nature of the water in the gel may also impact on the diffusion of molecules through it, blood interactions, and its performance as a cell-support in tissue engineering and regenerative medicine. The water that swells a hydrogel is impacted by the polymer chains at a nanoscale, because of its close proximity to the polymer that comprises the mesh of the hydrogel. When macroscopic pores are introduced into the hydrogel, the polymer will impact the water in a different fashion, i.e., the surface of the pores will interact with water in a manner similar to surface effects discussed in the next paragraph; in the interior of the pore (away from the pore wall), the water will have an organization more similar to bulk water. When a solid surface or biomaterial disrupts the continuum structure of water, the water near the solid surface will adopt a new organization to achieve free-energy minimization for the total system. Since all our biomaterials will first see water before proteins or cells ever diffuse or transport to the surface, the nature of this surface-water layer may be, from a biomaterials-science standpoint, the most important event driving biointeractions at interfaces. The nature of water in proximity to surfaces may be the primary driver for interactions between biomaterials and biological systems. There is ample evidence from many analytical techniques, and also from computational methods, that water organization is altered close to a surface compared to the bulk. Most experimental data suggest this difference persists over a length of one to four water molecules (i.e., about 0.3–1.2 nm), before reverting to a structure

80 SEC T I O N 1 . 2    Properties of Biomaterials

• Figure 1.2.5.4  Snapshots from molecular simulations of a helical leucine-lysine (LK) peptide adsorbing

onto (A) a negatively charged hydrophilic surface [self-assembled monolayer (SAM) with COOH/COO− surface groups], and (B) a hydrophobic surface (SAM with nonpolar CH3 surface groups) (from studies conducted by Collier et  al., 2012). Color coding: greenish-blue surface atoms and “sticks” in the LK peptide = C, white surface atoms and sticks in the peptide = H, red surface atoms and sticks in the LK peptide = O, dark blue sticks in the peptide = N, yellow spheres in solution = Na+ ions, greenish-blue spheres in solution = Cl− ions, red dotted lines = hydrogen bonds formed by water molecules. Molecular images created using VMD software (Humphrey et al., 1996).

indistinguishable from the bulk water. There is much research on the adsorption of the first layer of water on materials. Clearly, the organization of this first layer will dictate the structure of subsequent layers. On many closepacked metal crystal surfaces (for example, Pt, Ni, Pd) results suggest that a water bilayer exists with a structure analogous to ice (Hodgson and Haq, 2009). This may have led to the somewhat misleading term for interfacial water, “ice-like water.” Although the water structure is different from the bulk liquid at all surfaces, the specific ice-like organization is predominantly seen at these close-packed metal crystal surfaces. Above the water bilayer directly in contact with the metal, water structure is altered for a few molecular layers until it becomes indistinguishable from bulk water. Other studies demonstrate differences between hydrophilic and hydrophobic surfaces as to their interactions with water (Howell et  al., 2010). Both types of surfaces have a substantial effect on water organization, but in very different ways. Hydrophilic surfaces have functional groups that are attractive to water molecules by forming hydrogen bonds with them (e.g., surfaces with polar functional groups like O–H or N–H), or strong electrostatic interactions (e.g., charged functional groups), thus resulting in a closely bound layer of water over the surface. In contrast to this, for hydrophobic interfaces, which lack the ability to form hydrogen bonds or strong electrostatic interactions with water molecules, there appears to be a low water density zone, sometimes called a depletion zone, a few angstroms thick over the surface. Fig. 1.2.5.4 presents results from molecular dynamics simulations that illustrate these effects. In this figure, the presence of hydrogen bonds

formed by the water molecules over a negatively charged hydrophilic surface (Fig. 1.2.5.4A) and a hydrophobic surface (Fig. 1.2.5.4B) are indicated by the red dotted lines. Over the hydrophilic surface, hydrogen bonds are clearly formed with either the negatively charged surface functional groups or the tightly adsorbed counter-ions with these hydrogen bonds primarily oriented perpendicular to the surface. However, over the hydrophobic surface, which does not have the capability to form hydrogen bonds with the interfacial water molecules, there is a distinct gap a few angstroms thick directly over the surface (i.e., a depletion zone). The hydrogen bonds formed by the water molecules directly above this layer are primarily oriented parallel to the surface, indicating hydrogen bonding only between the water molecules themselves. The resulting hydrophobic effect is also clearly shown in this molecular model, with nonpolar side-chains of the adsorbing peptide being excluded by the surrounding water and tightly adsorbed to the hydrophobic surface. There are hundreds of recent studies on the water–solid interface, mostly in physical chemistry journals. A comprehensive review of this complex and still controversial subject would be impossible in this textbook. However, the takehome message is relatively straightforward—the presence of an interface alters the water structure adjacent to it. 

Water: Significance for Biomaterials When a protein or a cell approaches a biomaterial, it interacts with the surface water first. This final section will briefly review some implications of this surface water for biomaterials.

CHAPTER 1.2.5   Role of Water in Biomaterials

81

 Hydrogels The structured water within hydrogels has been discussed. Water structure associated with hydrogels has been implicated in their interactions with blood (Garcia et al., 1980; Tanaka and Mochizuki, 2004). Water structure in hydrogels is also thought to be important in the performance of hydrogel contact lenses, specifically the rate of lens dehydration (Maldonado-Codina and Efron, 2005).

Protein Adsorption



Figure 1.2.5.5 Lipid molecules with polar head groups (white) and hydrophobic tails (brown), when placed in water, organize themselves to minimize surface area contact between the hydrophobic tail (typically comprised of methylene units, –CH2-) and water. This minimization of contact area, depending on precise conditions, can lead to micelles, liposomes, or bilayer sheets. (Modified from a public domain image on Wikimedia Commons.)

Hydrophobic Effect, Liposomes, and Micelles If we place a drop of oil under water it will round up to a sphere as it floats to the water surface. The common explanation for this is that because oil molecules do not hydrogen bond with water molecules the oil cannot strongly interact with the bulk water phase. Another way to say this is that the oil disrupts the bulk (continuum) structure of water, and the water molecules at the oil interface then have to restructure to a new (more ordered) organization. This decrease in water entropy is energetically unfavorable according to the second law of thermodynamics (i.e., it results in an increase in the free energy of the system). Therefore, the oil minimizes its interfacial area with the water by coalescing into a sphere (a sphere has the lowest surface area for a given volume). This is called the hydrophobic effect (Tanford, 1978; Widom et  al., 2003). It is not driven by the oil–water molecular interactions, but rather by the necessity to minimize the more structured (lower entropy, higher free energy) organization of water molecules. If we now take a surfactant molecule comprised of an “oily” segment and a polar (water-loving) segment, by shielding the oily phase from the water with the hydrophilic head groups, contact between oil and water is minimized. Fig. 1.2.5.5 illustrates some hydrophobiceffect-driven supramolecular aggregate structures that lead to this free-energy minimization. Liposomes and block copolymer micelles are widely used in drug delivery, where the hydrophobic region (micelle core or liposome bilayer) or the aqueous center (liposome) can carry a hydrophobic drug or hydrophilic drug, respectively (see Chapters 1.3.8A and 1.3.8B). Bilayer sheets are used in biosensors to orient and stabilize receptor proteins.

Protein adsorption is discussed in detail in Chapter 2.1.2, and is critically important for understanding the performance of biomaterials. A question commonly posed is: “Why do proteins bind rapidly and tenaciously to almost all surfaces?” A model that can explain this considers structured water at interfaces. All surfaces will organize water structure differently from the bulk structure; this organization almost always gives more structured (lower entropy, higher free energy) water. If a protein can displace the ordered water in binding to the surface, the entropy of the system will increase, and thus the free energy will decrease as the water molecules are released and gain freedom in bulk water. This is probably the driving force for protein adsorption at most interfaces. There are some specially engineered surfaces referred to as “nonfouling” or “protein-resistant” (see Chapter 1.4.3A). These surfaces may resist protein adsorption by binding or structuring water so strongly that the protein molecule cannot “melt” or displace the organized or tightly bound water, and thus there is no driving force for adsorption.

 Life Living systems self-assemble from smaller molecular units. For example, think about the organizational processes in going from an egg, to a fetus, to a mature creature. Much of this assembly is driven by hydrophobic interactions (i.e., entropically by water structure). Another area where water structure has a major impact on life involves DNA and its unique binding of water (Khesbak et al., 2011). Also, consider enzymes that are so essential to life. A substrate molecule enters the active site of an enzyme by displacing water molecules. The unique mechanical properties of cartilage under compression can be modeled by considering water organization. In fact, the average human is 57% by weight water, on a molar basis by far the major component in the human body. Thus we can well appreciate, as Henderson surmised in 1913, that water is essential to this phenomenon we call life. As you work through this textbook, think about how the phenomena you are reading about may be driven or controlled by water that comprises such a large fraction of all biological systems. 

References Akaike, T., Sakurai, Y., Kosuge, K., Kuwana, K., Katoh, A., et al., 1979. Study on the interaction between plasma proteins and synthetic polymers by circular dichroism. ACS Polym. Prepr. 20 (1), 581–584.

82 SEC T I O N 1 . 2    Properties of Biomaterials

Collier, G., Vellore, N.A., Yancey, J.A., Stuart, S.J., Latour, R.A., 2012. Comparison between empirical protein force fields for the simulation of the adsorption behavior of structured LK peptides on functionalized surfaces. Biointerphases 7 (1), 1–19 article 24. Garcia, C., Anderson, J.M., Barenberg, S.A., 1980. Hemocompatibility: effect of structured water. Trans. Am. Soc. Artif. Intern. Organs 26, 294–298. Henderson, L.J., 1913. The Fitness of the Environment. The Macmillan Company, New York, NY. Hodgson, A., Haq, S., 2009. Water adsorption and the wetting of metal surfaces. Surf. Sci. Rep. 64 (9), 381–451. Howell, C., Maul, R., Wenzel, W., Koelsch, P., 2010. Interactions of hydrophobic and hydrophilic self-assembled monolayers with water as probed by sum-frequency-generation spectroscopy. Chem. Phys. Lett. 494 (4–6), 193–197. Humphrey, W., Dalke, A., Schulten, K., 1996. Vmd – visual molecular dynamics. J. Mol. Graph. 14 (1), 33–38. Jhon, M.S., Andrade, J.D., 1973. Water and hydrogels. J. Biomed. Mater. Res. 7, 509–522. Keutsch, F.N., Saykally, R.J., 2001. Water clusters: untangling the mysteries of the liquid, one molecule at a time. Proc. Natl. Acad. Sci. 98 (19), 10533–10540. Khesbak, H., Savchuk, O., Tsushima, S., Fahmy, K., 2011. The role of water H-bond imbalances in B-DNA substrate transitions and peptide recognition revealed by time-resolved FTIR spectroscopy. J. Am. Chem. Soc. 133 (15), 5834–5842.

Maldonado-Codina, C., Efron, N., 2005. An investigation of the discrete and continuum models of water behavior in hydrogel contact lenses. Eye Contact Lens 31 (6), 270–278. Marcus, Y., 2009. Effect of ions on the structure of water: structure making and breaking. Chem. Rev. 109, 1346–1370. Paschek, D., Ludwig, R., 2011. Specific ion effects on water structure and dynamics beyond the first hydration shell. Angew. Chem. 50 (2), 352–353. Sekine, Y., Ikeda-Fukazawa, T., 2009. Structural changes of water in a hydrogel during dehydration. 034501 J. Chem. Phys. 130 (3), 034501. Smith, J.D., Cappa, C.D., Wilson, K.R., Cohen, R.C., Geissler, P.L., et al., 2005. Unified description of temperature-dependent hydrogen-bond rearrangements in liquid water. Proc. Natl. Acad. Sci. U. S. A 102 (40), 14171–14174. Tanaka, M., Mochizuki, A., 2004. Effect of water structure on blood compatibility: thermal analysis of water in poly (meth) acrylate. J. Biomed. Mater. Res. 68A, 684–695. Tanford, C., 1978. The hydrophobic effect and the organization of living matter. Science 200 (4345), 1012–1018. Tsuruta, T., 2010. On the role of water molecules in the interface between biological systems and polymers. J. Biomater. Sci. Polym. Ed. 21 (14), 1831–1848. Widom, B., Bhimalapuram, P., Koga, K., 2003. The hydrophobic effect. Phys. Chem. Chem. Phys. 5 (15), 3085. Zhang, Y., Cremer, P.S., 2010. Chemistry of Hofmeister anions and osmolytes. Annu. Rev. Phys. Chem. 61 (1), 63–83.

Chapter Exercises Consider having two types of homogeneous surface chemistries: one with hydrophilic hydroxyl (-OH) groups and the other with hydrophobic methyl (-CH3) groups. (a) Which surface would more strongly adsorb a protein in air? (b) Which surface would more strongly adsorb a protein in water? Answers: (a) In air, the –OH surface would more strongly adsorb the protein due to the fact that it would be able to form the strongest type of van der Waals (vdW) interactions (i.e., hydrogen bonds) with the polar groups of the protein, as well as forming moderately strong dipole-induced dipole vdW interactions with the nonpolar groups of the protein. In comparison, the -CH3 surface would only be able to form the moderate and very weak vdW interactions

with the protein (i.e., induced dipole–dipole and induced dipole–-induced dipole interactions, respectively). (b) In water, the -CH3 surface would more strongly adsorb the protein due to the ability of water molecules to form hydrogen bonds with the –OH groups of the surface, which will directly compete with the protein, thus weakening the –OH surface’s bonding with the protein. In contrast, the -CH3 surface would form very strong hydrophobic interactions with the hydrophobic domains of the protein due to the thermodynamically unfavorable interactions between the -CH3 groups of the surface and the surrounding water.

Suggested External Reading Voet, D., Voet, J.G., Pratt, C.W., 2016. Fundamentals of Biochemistry: Life at the Molecular Level, fifth edition, Chapter 2: Physical and Chemical Properties of Water, John Wiley & Sons, pp. 23–41.

82.e1

S E C TI ON 1 . 3

Classes of Materials Used in Medicine

1.3.1

The Materials Side of the Biomaterials Relationship WILLIAM R. WAGNER Departments of Surgery, Bioengineering & Chemical Engineering, McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, United States

W

hen seeking to cover the “Classes of Materials Used in Medicine,” as this section of the book states, a quick overview of the landscape, together with some caveats and definitions, is in order. What are these materials that will be brought to bear as biomaterials in devices? Materials employed as biomaterials will need to serve many needs: replacement of physiologic functions lost to trauma, disease, or congenital conditions (e.g., cardiac valves); sensing (e.g., ultrasound contrast agents); delivery of materials within the body (e.g., stent-delivery catheters); temporary mechanical support (e.g., sutures); or extracorporeal processes (e.g., dialyzers). It is not surprising that the array of applied biomaterials is very broad and growing. However, this chapter will not consider specific classes of materials that are used daily in the clinic to meet some of the needs just outlined. Specifically, consider biomedical devices that seek to replace a given physiologic function and the analogous situation with an automobile. For an automobile, tires that no longer grip on wet roads or a battery that cannot start the vehicle elicit a trip to the mechanic’s shop and replacement of these parts with “devices” that are fully capable of restoring those functions, or perhaps even exceeding the capabilities and duration of the original equipment manufacturer. The parts are designed to be fully compatible and are readily available for a quick and unambiguous replacement procedure. The situation in the clinic from a needs perspective is similar, but unfortunately society has not advanced to the state where off-the-shelf replacement parts of compositional and functional equivalence are available. Arguably, the best source for such materials in many cases is from the patient: autograft materials. Muscle flaps moved by plastic surgeons for reconstructive procedures, saphenous veins and arteries used for arterial bypass, and skin harvested and mechanically processed to increase its

coverage area for burn treatment all represent current goldstandard materials used to meet clinical needs for replacement tissues. Which of these autograft materials are best selected for specific conditions and how they are isolated and processed is the focus of surgical textbooks, but may only be peripherally mentioned here. Yet the reader should be aware that biomaterials and medical device technology falls usually well short of addressing clinical needs as effectively as autografts may. This is true despite the downsides associated with autografting, such as donor site morbidity. Similar to autografts, allografts (from organ and tissue donation) fill critical needs for tissue replacement where autografts and other material-based approaches are not possible or fall short. For instance, the best option for patients in end-stage cardiac, kidney, liver, and pulmonary failure is currently to receive an allograft. Despite the morbidity associated with immunosuppressant therapy, and issues associated with a limited supply, the quality of life and survival for allograft recipients exceeds similar populations that would be supported by the best current medical support devices for heart, lung, or kidney. The considerations, particularly immunology, surrounding allograft materials are beyond the scope of this text, but the target populations for important classes of medical devices are similar. For both auto- and allografts used in the manner already described, these materials are utilized in a minimally processed state. Particularly for allografts and certainly for xenograft materials, the application of processing methods, from decellularization to cryopreservation and cross-linking, bring these materials into the realm of biomaterials that are considered in this text. With the preceding discussion the need for extensive consideration of surgical principles as well as broad coverage of allografting, immunology, and immunosuppressive therapy has been taken off the table. At the same time, two critically 83

84 SEC T I O N 1 . 3     Classes of Materials Used in Medicine

important classes of materials used in medicine have been noted, but largely dismissed. Having recognized this, the focus for Section 1.3 is to cover the remaining major classes of materials used in medicine. Studying biomaterials requires the study of the interface between the material and the biological, and in this section the material side of this relationship is covered. To fulfill the design requirements of diverse medical devices, diverse materials are needed that can meet mechanical, chemical, and biological requirements. For some applications it may be that very different material types are being considered due to different design considerations. For instance, in articulating joint replacement devices, some parts have been made from polymer, metal, or ceramic in different designs. In vascular stents a variety of metals have been used, but with recent consideration of degradable stents, both degradable polymeric stents from poly(lactide) and degradable metallic stents made of a magnesium alloy have been tested clinically. Understanding what different material classes and subclasses offer in terms of properties and how structure relates to function is essential knowledge for the biomaterials expert. The chapters here are generally grouped into categories of polymers, metals, ceramics, natural materials, composites, and particulates with subchapters devoted to particularly important material types. An major trend in the biomaterials community over the past several decades has been the harnessing of advances in molecular biology to design materials with specific biological functionalities. Some of these advances will be covered in this section (Chapter 1.3.2G) and also in select chapters later in the book. More recently there has been a dramatic increase in research focused on particulate biomaterials,

where biological knowledge is critically leveraged to impart targeting potential, responsiveness to conditions at a targeted site (e.g., pH), and to orchestrate specific interactions at a cellular level upon arrival. Traditional bulk materials such as polymers, ceramics, and metals may be used as part of these designs, but self-assembly of macromolecules is also often a feature. Chapters 1.3.8A and 1.3.8B seek to summarize this rapidly advancing area in the biomaterials community. Finally, the use of biologically derived materials has been of growing interest and clinical impact. These materials range from highly purified single components that may be from plant or animal sources, to engineered natural materials where functionality such as photocross-linking activity may be added, to materials derived from the extracellular matrix or tissues with processing for degradation resistance (fixation) or intended remodeling (decellularized tissues). The subsections of Chapter 1.3.6 seek to cover this space. While the theme for this section is firmly on the materials side of the biomaterials relationship, as the field progresses the knowledge bases that inform either side of the material and biological interface are merging, and with some materials the design and control of the biological response has become the critical issue in engineering the intended effect. At the same time, the reality in today’s clinics and hospitals is that when autografts and allografts are not available or when devices are needed for other purposes beyond tissue replacement, the vast majority of medical devices remain comprised of materials adopted from the broader materials community to meet the principal design objectives. There remains the need for understanding and working with the fundamentals of materials science.

1.3.2

Polymers: Basic Principles GARRETT BASS 1 , MATTHEW L. BECKER 1 , DANIEL E. HEATH 2 , STUART L. COOPER 2 1Departments

of Chemistry, Mechanical Engineering and Materials Science, Biomedical Engineering and Orthopaedic Surgery, Duke University, Durham, NC, United States 2William

G. Lowrie Department of Chemical and Biomolecular Engineering, The Ohio State University, Columbus, OH, United States

P

olymers represent the largest class of materials used in medicine. They have a range of unique properties making them useful in biomaterial applications, including orthopedics, dental, hard and soft tissue replacements, drug delivery, and cardiovascular devices. This chapter explains the basic principles of polymer science, illustrates how polymer materials can be specifically designed to fill needs in the biomaterials field, and provides examples of how this class of materials is currently used in medical applications. Structure–property relationships, including molecular architecture, molecular mass, and chemical composition related to the physical and chemical properties of the macroscopic material, are included. A biomaterials scientist aware of structure–property relationships can engineer a polymer system for a specific need.

Introduction Polymer materials possess an array of unique properties that make them useful in a variety of biomaterial applications such as orthopedics, dental, hard and soft tissue replacements, drug delivery and cardiovascular devices. In fact, polymers represent the largest class of materials used in medicine. This chapter introduces the basic principles in polymer science, illustrates how polymer materials can be specifically designed to fill needs in the biomaterials field, and provides examples of how this class of materials is currently used in medical applications. The central idea of this chapter is structure–property relationships, which means how molecular characteristics such as molecular architecture, molecular mass, and chemical composition are directly related to the physical and chemical properties of the macroscopic material. It also includes a brief introduction to polymer synthesis, presents advantages and disadvantages of these methods, and outlines recent developments in postpolymerization functionalization. For instance, polymer scientists in other fields have been able to

exploit structure–property relationships to create nonstick coatings, pressure-sensitive adhesives, and the penetrationresistant materials used in bulletproof vests. A biomaterials scientist aware of structure–property relationships can rationally engineer a polymer system for a specific need. 

The Polymer Molecule Molecular Structure of Single Polymer Molecules The hallmark of polymer molecules is high molecular mass. A single polymer molecule could have a molecular mass of 2,000,000 Da compared to a water molecule, which has a molecular mass of 18 Da. Furthermore, polymer molecules are organized into very interesting architectures. Common architectures of polymer molecules are shown in Fig. 1.3.2.1. The simplest is the linear chain where there is a single molecular backbone. When linear chains of two different compositions (e.g., polymers A and B) are linked together, the resultant polymer is called an A–B block copolymer. If another chain is added to the second chain, it may be called an A–B–C triblock copolymer, or more simply an A–B–C block copolymer. Differences in monomer reactivity can also lead to random or gradient copolymers where the chain sequence of monomer addition is altered. Branched structures are also possible where a central polymer backbone has smaller side chains extending from it. Branches can occur due to undesired side reactions during synthesis, or can be purposefully incorporated into the molecular structure. The type and extent of branching introduces significant changes in properties to a polymer system. Star-shaped polymers consist of a multifunctional core molecule from which at least three polymer chains (arms) radiate. These arms can be chemically identical (homostars) or different (heteroarm stars). Star polymers have been widely used in biomedical applications such as drug and 85

86 SEC T I O N 1 . 3     Classes of Materials Used in Medicine

end. The repeat unit itself consists of the polymer backbone, a series of carbon–carbon single bonds, while the hydrogen atoms and methyl group are pendant groups. The repeat unit can be controlled through choice of synthetic method and plays a large role in the macroscopic behavior of the polymer. Fig. 1.3.2.3 shows the repeat units of several synthetic and natural polymers commonly used in the biomaterials field. 

Copolymers • Figure 1.3.2.1 Three polymer architectures commonly seen.

gene delivery, tissue engineering, diagnostics, and antifouling biomaterials. The intensified interest in star polymers is attributed to their unique topological structures and physical/chemical properties. Polymer brushes are a branched polymer system in which chains of polymer molecules are attached along a linear polymer backbone in such a way that the graft density of the polymers is sufficiently high to force stretching of the tethered chains away from the backbone. Some examples of naturally occurring polymer brushes include extracellular polysaccharides on bacterial surfaces and the proteoglycans of cartilage. Dendrimers are monodisperse, precisely branched macromolecules in which all bonds emerge radially from a central focal point with a regular branching pattern and with repeat units that each contribute a branch point. These structures exhibit a high degree of molecular uniformity, precise molecular mass, tunable size, shape, and number of functional groups. So far we have discussed polymer systems in which there are discrete polymer molecules. However, what would happen if you took a linear polymer molecule and covalently bonded it to the backbone of another linear chain? A few of these bonding events would produce a branched structure as noted previously. If you repeated this act many times over you would eventually link all of the polymer chains together into one very large network polymer. This is also possible by using small di- or trifunctional “cross-linker” molecules to react with pendant or terminal reactive groups on linear polymer chains, eventually yielding a network polymer. 

Chemical Structure of Single Polymer Molecules If you were able to see the individual atoms making up a polymer molecule, you would notice the same basic structure repeats over and over again. This structure is called the repeat unit of a polymer molecule. Etymologically, polymer comes from the Greek “poly,” meaning many, and “meros,” meaning part—many parts. In polymer molecules the “mer” is the repeat unit. Fig. 1.3.2.2 shows a schematic of a linear polypropylene molecule. The polymer chain is composed of many –(–CH2–CHCH3–)– repeat units covalently linked end to

Sometimes it is advantageous to synthesize copolymers— polymers containing more than one chemically distinct repeat unit. For instance, a researcher may synthesize a polymer that contains repeat units “A” and “B.” As shown in Fig. 1.3.2.4, there are many different ways the repeat units could be organized. Random copolymers occur when the “A” and “B” repeat units have no order in the backbone; however, alternating, block, gradient, and graft copolymers are also possible and the arrangement of repeat units affects the physical behavior of the biomaterial. The repeat units of two random copolymers commonly used in the biomaterials field are shown in Fig. 1.3.2.5. 

Determination of Chemical Composition A researcher will often need to verify the chemical structure of polymers or determine the composition of copolymer systems. Two common techniques a scientist would use are nuclear magnetic resonance (NMR) spectroscopy and infrared (IR) spectroscopy. NMR is an analytical technique that exploits the magnetic moments associated with isotopes that contain an odd number of protons and/or neutrons. These atoms have an intrinsic nuclear magnetic moment and angular momentum, in other words a nonzero nuclear spin, while all nuclides with even numbers of both have a total spin of zero. The most commonly used nuclei are 1H, 13C, and 19F although isotopes of many other elements can be studied by high-field NMR spectroscopy as well. In a typical experiment, these nuclei are excited to a higher energy state through a burst of radiofrequency radiation. The nuclei relax to a lower energy state, which is measured as an electric signal. Fourier transform analysis is used to convert this time domain electrical signal into the frequency domain. Due to electron shielding, protons attached to different structural units will display chemical shifts, meaning their resonances in the NMR spectrum will be at different frequencies. Through analysis of the resonance placement, splitting patterns, and intensity, the chemical structure of molecules can be determined. IR spectroscopy is also used to determine the chemical composition of polymers. In a typical experiment, a sample of interest is irradiated with IR radiation. The sample absorbs certain wavelengths, resulting in specific molecular motion or vibrations (such as C–H stretching). The IR spectrum is created by plotting absorbance versus wavelength. Like NMR, analysis of the spectra can lead to the verification of a polymer's composition.

CHAPTER 1.3.2   Polymers: Basic Principles

H

H

H

Me H

H

H

Me H

H

H

Me H

H

H

Me H

H

H

Me H

H

Me

H

H

Me

H

H

87

Isotactic PP H

H

H

Me Me

H

H

H

H

H

H

Me Me

H

H

H

H

Me Me

H

H

H

H

Me Me

H

H

H

H

H

Syndiotactic PP Me

H

H

H

H

H

H

Me H

H

Me H

H

H

H

H

H

Me

Atactic PP

• Figure 1.3.2.2  Polypropylene repeat unit and its different tactic isomers. In the schematic Me indicates a methyl group.

Often the chemical composition of the polymer surface is different from the bulk. For a medical implant the surface composition is highly important, since it will interface with the biological environment. To probe the surface composition, X-ray photoelectron spectroscopy—also known as electron spectroscopy for chemical analysis—is a common technique. A sample is bombarded with X-rays, which results in the ejection of inner shell electrons from the atoms displayed on the material surface. The kinetic energies of the ejected electrons are measured and interpreted into information about the chemical composition of the surface (for more information on surface characterization techniques, see Chapter 1.2.4). 

synthesis, a polymer chemist would normally produce an atactic version of PP, one where the methyl group is randomly located in front of and behind the polymer backbone. However, when a special catalyst is used during synthesis a chemist can produce isotactic PP, where all the methyl groups are located on one side of the “stretched-out” polymer backbone, or syndiotactic PP, where the methyl groups regularly alternate from side to side. As will be discussed later, tacticity can drastically affect the physical behavior of the polymer system, largely by affecting the ability of the polymer molecules to crystallize. Tactic isomers occur whenever an atom in the polymer backbone has the capacity to form tetrahedral bonding and is bonded to four different chemical groups. Such atoms (generally carbon) are referred to as asymmetric.

Tacticity

 Molecular Mass The Molecular Mass Distribution and Its Averages

Tacticity describes the stereochemistry of the repeat units in polymer chains. To illustrate the discussion on tacticity, let us consider a molecule of polypropylene (PP). When stretched into its planar zigzag form, as seen in Fig. 1.3.2.2, you can see that sometimes the methyl groups are all on one side of the backbone, sometimes they alternate from side to side, and sometimes they are randomly distributed. During routine

During polymerization, polymer chains are built up from monomers to a desired molecular mass. Polymers with identical composition but different molecular mass (different chain length) may exhibit different physical properties. The number of monomer repeat units in each polymer chain is called the degree of polymerization (DP).

88 SEC T I O N 1 . 3     Classes of Materials Used in Medicine

Polyethylene (PE)

Polytetrafluoroethylene (PTFE)

Polyvinylidene fluoride (PVDF)

Polyvinylchloride (PVC)

Polypropylene (PP)

Polybutylene (PB)

Polyisobutylene (PIB)

Poly(lactic acid) (PLA)

Poly(acrylic acid) (PAA)

Poly(acrylamide) (PA)

Poly(hydroxyethyl methacrylate) (PHEMA)

Polydimethylsiloxane (PDMS)

Poly(ether ether ketone) (PEEK)

Cellulose

Poly(methyl methacrylate) (PMMA)

Poly(ethylene terephthalate) (PETE)

Nylon 6,6

• Figure 1.3.2.3  Repeat units of common polymer biomaterials.

- - - - - AAAAAAAAAAAAA- - - - - Homopolymer -----AABABBBAABBAB-----Random copolymer

-----ABABABABABABA-----Alternating copolymer

-----AAAABBBBBAAAA-----Block copolymer B B B B B B -----AAAAAAAAAAAAA-----Graft copolymer



Figure 1.3.2.4 Various copolymers formed by varying the arrangement of two monomers.

For a homopolymer, there is only one type of monomeric unit and the number average degree of polymerization is given by Mn = M0 × DP, where Mn is the number average molecular mass and M0 is the molecular mass (or formal weight) of the monomer unit.

However, in a traditional free radical or condensation polymerization synthesis, each polymer chain will not have the same DP. For instance, in the free radical polymerization of polyethylene (PE), one polymer chain may add 3000 monomers, a second may add 4500 monomers, and a third may only add 1500. Therefore most polymer systems have a distribution of molecular masses. Since polymer materials are made from molecules with a variety of molecular mass, it is incorrect to talk about the molecular mass of a polymer system. Instead, polymer systems are described by average values of molecular mass. The two most commonly used averages are the number average molecular mass (Mn) and the weight average molecular weight (Mw). The mathematical definitions of these averages are supplied in Eqs. (1.3.2.1) and (1.3.2.2), where Ni is the number of molecules with “i” repeat units and Mi is the molecular weight of a polymer chain with “i”

CHAPTER 1.3.2   Polymers: Basic Principles

3RO\ WHWUDIOXRURHWK\OHQHFRKH[DIOXRURSURS\OHQH  UDQGRPFRSRO\PHU

89

3RO\ ODFWLFFRJO\FROLFDFLG  UDQGRPFRSRO\PHU

• Figure 1.3.2.5  Repeat units of two common copolymer biomaterials.

repeat units. Since Mw is calculated using the square of the molecular weight it is almost always greater than or equal to Mn: Ðm >1(1.3.2.1) ∑

Ni Mi i (1.3.2.2) Mn = ∑ Ni i



Ni M2i Mw = ∑ Ni Mi i

i

The ratio of Mw to Mn is the molecular mass distribution (Ðm) (Eq. 1.3.2.3), which is a measurement of the breadth of the chain length populations. If the Ðm of a polymer system is unity (1), the number average and weight average molecular masses are identical, meaning the polymer sample is monodisperse (all chains have the same degree of polymerization). This level of precision is only found in dendrimers and biopolymers. For most condensation polymers, the Ðm is approximately 2: Mw Ðm = (1.3.2.3) Mn The higher the average molecular mass, the stronger a polymer material will be (up to a point). However, the melt/ solution viscosity increases with increasing average molecular mass, making the material more difficult to process. Often a molecular mass range exists within which most desired physical behaviors are achieved, yet the material is still easily processed. Generally, condensation polymerizations are suitable for a number average molecular mass range from 25,000 to 50,000 Da, while addition polymerizations are preferred for values from about 50,000 Da up to hundreds of thousands.

 Characterizing the Molecular Mass Distribution Through understanding the importance of molecular mass and its mass distribution, polymer scientists have developed many methods for measuring average molecular mass values. For instance, Mn can be determined through techniques such as end-group analysis by NMR, vapor pressure lowering, and freezing point depression. In the following paragraphs we will first discuss common methods historically used to measure Mn

(osmotic pressure) and Mw (light scattering). Although osmotic pressure and light scattering are powerful and useful techniques in the field of polymer science, they have been replaced in recent years by size exclusion chromatography (SEC), which gives much more detail about the molecular mass distribution, and does so much quicker. From the molecular mass distribution obtained by SEC, the first and second moments of the distribution, Mn and Mw, are readily determined. In osmotic pressure experiments, a dilute polymer solution is separated from pure solvent by a semipermeable membrane through which solvent can freely pass but which excludes polymer. The activity of the pure solvent differs from that of the solvent molecules in the solution phase resulting in a thermodynamic driving force—the osmotic pressure (π or P) causes molecules of pure solvent to diffuse across a membrane (which is permeable only to solvent) and into a compartment containing a solute (e.g., polymer) in solution. This causes the fluid level in the solution compartment to rise, resulting in a hydrostatic pressure head. Equilibrium is reached when the pressure head exactly offsets the osmotic pressure. By measuring the rise in fluid level, the osmotic pressure can be calculated and related to the number average molecular weight. A number of polymer solutions are prepared, each with a distinct concentration. The osmotic pressure is measured for each solution and plotted against concentration, the concentration is then extrapolated to zero, and the number average molecular weight can be determined from the intercept of the regression line. Light scattering is a common technique to determine the Mw of a polymer system. In dilute solutions, the scattering of light is directly proportional to the number of molecules. The scattered intensity is observed at a distance r and angle Q from the incident beam, and is related to the size or weight average molecular weight of the molecule. The light-scattering behavior of a series of polymer solutions with varying concentrations is measured and the data are extrapolated to zero concentration to determine Mw. Once the Mn and Mw values are determined, the Đm of the polymer system can be calculated. SEC is the most commonly used molecular mass characterization technique in modern polymer science laboratories. Unlike osmotic pressure and light scattering, which provide average values of molecular mass, SEC provides the entire molecular mass distribution from which all the desired average values can be calculated. Different average values can be defined using SEC, depending on the statistical method applied. In practice, two averages are commonly used, representing the weighted mean taken with the mole fraction and the weight fraction. In an SEC experiment, a dilute polymer solution flows through a column packed with solid

90 SEC T I O N 1 . 3     Classes of Materials Used in Medicine

separation media that contains pores on the same size scale as the polymer molecules. As the polymer molecules flow though the column, smaller chains diffuse into the pores and thus must travel a further distance, while larger polymer chains move through the column faster since they are excluded from the pore structure, resulting in higher molecular mass material eluting from the column first, followed by lower molecular mass species. The concentration of the effluent stream is monitored and a plot of detector signal (which is proportional to polymer concentration) versus retention time can be generated. The use of narrow molecular mass distribution standards allows the user to correlate retention time with molecular mass, although the actual correlation is with the hydrodynamic volume of the known standards. If the relationship between molar mass and the hydrodynamic volume changes (i.e., the polymer is not the same shape in solution, i.e., branched, cyclic, as the standard, or is of a differing chemical makeup), then the calibration for mass is in error.

 Connecting Physical Behavior With Chemical Characteristics We have discussed the key characteristics of polymer molecules (molecular architecture, chemical composition, tacticity, and molecular mass). Now we will relate these molecular characteristics to macroscopic properties, and illustrate how these characteristics can be manipulated to create a polymer system with the desired behavior. We will focus on the tensile properties, hydrophilicity, and biodegradability of polymer systems.

Physical States of Linear Polymers When designing a biomaterial, physical properties are key features. For instance, if you were creating a cement for use in loadbearing bones (tibia or femur) you would have to ensure the material is both strong enough to act as a cement, but not so brittle that it would fail due to low fracture toughness. The physical properties of a polymer stem from the intermolecular interactions occurring between individual polymer molecules; thus the molecular characteristics we have discussed up to this point are of extreme importance. As you will see in the following text, the four most fundamental molecular characteristics of polymer chains that determine the physical behavior of a polymer are chain stiffness, chain composition or polarity, chain architecture or regularity, and molecular mass. They determine two important temperatures that characterize polymer molecules: Tg, the glass transition temperature, and Tm, the crystalline melting temperature. Tm is only present when there are crystalline regions in a polymer. In Fig. 1.3.2.2, we see polypropylene molecules extended into planar zigzags. Although this is a convenient way to draw polymer molecules, this type of extended structure is rarely seen in nature. More often, polymer molecules are found as unorganized and three-dimensional structures called a random coil. In an amorphous structure, each random coil is highly entangled with its neighbors. Polymers in the rubbery state or the glassy state have this amorphous

molecular arrangement. Under certain conditions, some polymers will arrange themselves into highly organized crystalline domains resulting in a semicrystalline material. Each of these states will now be explored in more depth. 

The Rubbery State Rubbery polymers are amorphous. However, the random coils have enough thermal energy for rotation to occur around single bonds. If you were able to look at polymer molecules in the melt state, you would see that each random coil is continuously changing shape (conformation). This molecular motion becomes more intense as the thermal energy in the system increases. Macroscopically these materials are soft, flexible, and extensible, due to the molecular motion available to the molecules. 

The Glassy State As the polymer system is cooled, rotation around bonds becomes hindered due to energy barriers created by a segment of a chain having to move (rotate) past a neighboring segment. As the temperature drops, the rate of segmental motion in a polymer chain becomes slower and slower, and the chain gets stiffer and stiffer. As the system approaches the Tg the interpenetrated random coils become frozen in space. This is called the glassy state. A material below its Tg is called a glass because it is hard, stiff, and brittle. Molecules in the glassy state can no longer rearrange themselves under applied stress, so deformation results in straining the secondary interactions between molecules. The opposite occurs when an amorphous polymer is heated: the amorphous region goes from hard and glassy, to “leathery,” to rubbery, and if the material is not cross-linked, it will eventually flow as a viscous fluid and can be processed into shapes. 

RAMEN NOODLES: GLASSY VERSUS RUBBERY POLYMER CHAINS If you are reading this you are probably a college student, which means you are very familiar with Ramen noodles. A package of noodles can be cooked in about 10 min. When you first remove the noodles from the package, you will notice that the individual noodles are rather random in shape and intertwined with their neighbors. Furthermore, before boiling, these noodles are very rigid and fixed in relation to one another. This food product is a pretty good example of a polymer in the glassy state, where each noodle represents one polymer molecule. As the noodles are boiled, they become flexible and can easily change their shape and slip around; however, they retain their intertwined character. These noodles are a good example of a polymer material in the rubbery state.

The Semicrystalline State All polymer systems form glasses at sufficiently low temperatures. However, as a melt is cooled, certain polymers have the ability to pack into a regular lattice, leading to the formation of stable crystalline domains. In PE, these stable crystalline domains are formed by chains in the planar zigzag conformation, while the crystalline chains

CHAPTER 1.3.2   Polymers: Basic Principles

91

TABLE 1.3.2.1    Physical Properties and Equilibrium Water Absorption of Common Polymeric Biomaterials

Material

Tensile Modulus (GPa)

Tensile Strength (MPa)

Elongation at Break (%)

Water Absorption (%)

Polyethylene

0.8–2.2

30–40

130–500

0.001–0.02

Poly(methyl methacrylate)

3–4.8

38–80

2.5–6

0.1–0.4

Polytetrafluoroethylene

1–2

15–40

250–550

0.1–0.5

Polylactide

3.4

53

4.1

365 nm) can be used to degrade biomaterials in the presence of cells. Photodegradable biomaterials have predominantly been used in tissue engineering, allowing for precise spatiotemporal control over material properties, capturing extracellular matrix heterogeneity in a synthetic scaffold (Brown and Anseth, 2017). Furthermore, photodegradation has been applied for release of biomolecules from a network, as well as spatially controlled photopatterning (Kloxin et al., 2010; DeForest and Anseth, 2011; Arakawa et al., 2017). In this field, a common method involves incorporation of photolabile cross-linkers into a network, especially hydrogels. When irradiated with a specific wavelength, the crosslinked molecules absorb light causing bond rearrangement and cleavage of the cross-links. The reactions depend on the intensity and wavelength of light, as well as the photophysical properties of the photodegradable linker. First-order kinetics

are often relevant to describe the rate of cleavage through a material with uniform illumination, Eq. (1.3.2F.9): dC dt

= − kC , k =

ΦεI NA hv 

(1.3.2F.9)

Here, C is the concentration of absorbing species and the kinetic rate constant k is a product of the quantum yield (Φ) and absorptivity (ε) of the photoactive species at the wavelength used along with the light intensity ( I ), light frequency (v ), Planck’s constant (h ), and Avogadro’s number (NA ). For elastomers, the rate of cleavage of a photolabile cross-linker can be related to a network’s storage modulus (G ′) and crosslink density ( ρx ), which can be used to predict changes in material properties and mass loss with time, Eq. (1.3.2F.10): G′ ′

G0

=

ρx ρx , 0

= e − kt 

(1.3.2F.10) 

Enzymatic Degradation Several synthetic polymers are susceptible to enzymatic degradation. For example, amorphous PLA can be degraded by proteinase K, while both amorphous and crystalline PCL can be degraded by lipases of various origins (Liu et  al., 2000, 2019) (Fig. 1.3.2F.9). Enzymatic degradation of synthetic polymers can occur by either surface or bulk mechanisms, depending on the location and stability of the acting enzyme. However, due to the limited water accessibility of most hydrophobic polymers, surface erosion is often the dominant degradation mechanism, described in an earlier section of this chapter. The degradation for a hydrophilic polymer, permitting the diffusion of the

CHAPTER 1.3.2F   Degradable and Resorbable Polymers

(A)

177

(B)

• Figure 1.3.2F.9  Scanning electron micrographs of esterase degraded PCL films after 48 h; (A) control PCL film (no enzyme present), (B) Lipase degraded film, Magnification 150x. (Adapted from Liu et al. (2019)).

enzyme to the interior of the polymer, can also be mediated by a surface erosion mechanism if the rate of enzymatic polymer bond cleavage is faster than the rate of enzyme diffusion. Bulk degradation of synthetic polymers induced by enzymatic activity may occur under two conditions: (1) the enzyme is able to infiltrate and distribute uniformly throughout the bulk of the polymer, and (2) the rate of enzymatic bond cleavage is slower than the diffusion of the enzyme. Typically, hydrogels containing enzymatically cleavable units in their polymer backbone fall into this category of degradation. Under these assumptions, Michaelis–Menten enzymatic kinetics are commonly used to predict the enzymatic degradation rates of synthetic polymers via a bulk degradation mechanism: v0 = vmax

[S] KM + [S]



(1.3.2F.11)

Here, v0 and vmax are the initial and maximum reaction rate of degradation, respectively. [S] is the degradable polymeric substrate concentration, and KM is the Michaelis– Menten constant. In Eq. (1.3.2F.11), vmax can also be expressed as: vmax = kcat [E] (1.3.2F.12) where kcat is the catalytic constant describing the rate of enzymatic degradation of polymer bonds and [E] is the concentration of enzyme catalytic sites. As in other enzymatic reactions, both kcat and KM are important parameters in characterizing the enzymatic degradation of polymers. While kcat represents the sensitivity of an enzyme for a specific polymeric substrate, KM is the substrate concentration needed to achieve a half-maximum enzyme velocity. Factors affecting these parameters will also determine the rate of polymer degradation. While simple Michaelis–Menten kinetics can be used to describe enzymatic polymer bond cleavage, they do not provide information regarding the mass loss behavior of the polymers. Sophisticated mathematical models are often required to correlate microscopic enzymatic bond cleavage to macroscopic polymer mass loss, and it depends on the structure and connectivity of the polymer. For example, to describe the mass

loss of cross-linked hydrogels containing enzymatic cleavable substrate (e.g., PCL), a statistical-kinetic model integrating the structural information of the hydrogels with the enzyme degradation kinetics is required (Rice et al., 2006).  Orthogonal Stimuli-Labile Strategies While most degradable polymer systems rely on single degradation pathways, materials capable of undergoing multiple routes of degradation have emerged in recent years. For example, release of model therapeutics from hydrogels with orthogonal stimuli-labile linkers was demonstrated. The degradation routes were reduction of disulfide bonds, enzymatic cleavage, photodegradation, or combinations thereof, leading to logic-based, controlled release (Ruskowitz et al., 2019). Adopting such a strategy to include light, enzyme, or other stimuli-degradable linkages within a hydrolytically degradable scaffold will enable development of other tailorable scaffolds for precision medicine applications.

Polymer Design and Processing The chemical structure (e.g., backbone functional groups, crystallinity, and architecture) of a polymer significantly influences its degradation mechanism and rate. Because of this, it is often desirable to use well-studied degradable polymers (e.g., PLA, PLGA) for the synthesis of biomaterial devices, because their breakdown is better understood in vivo and they are used in many FDA-approved applications, giving precedent for use in other applications. As biomaterials are required for a growing number of clinical applications, from synthetic replacements for biological tissue to materials as diagnostics (Langer and Tirrell, 2004), more sophisticated polymers are continually being developed. However, in designing a new degradable biomaterial, one must first understand the unique performance requirements and constraints for the intended application, including lifetime, location, mechanical properties, and delivery method (Table 1.3.2F.3). Then, different polymer-processing strategies can be applied to meet the design requirements, in terms of both degradation rate and overall performance. This section covers the design and processing considerations that are especially important for degradable biomaterials, providing examples of degradable systems that are commercially available or are being explored in academia.

178 SEC T I O N 1 . 3     Classes of Materials Used in Medicine

TABLE 1.3.2F.3   Design Criteria of Different Degradable Biomaterial Application With Polymer Examples

Lifetime/ Biodegradation Time

Location in the Body

Mechanical Properties

Mechanism of Delivery

Degradable Polymer Examples

More than 10 years

Bone joint, high mechanical load

High modulus, good fracture toughness

Surgery

Poly(lactic acid), poly(glycolic acid), poly(caprolactone)

Allow improved healing process without secondary surgery/reduce inflammation

1 week–1 month

Organ, skin

High tensile strength

Surgery

Poly(trimethylene carbonate), polydioxanone

Drug Delivery

Releases drug to a target site

Dependent on location/delivery short term/ long term

Digestive track, blood stream, tumor

Not critical

Oral/ injection

Poly(ortho-esters), polyphosphazenes

Adhesives

Wound healing, hemorrhage control, drug carriers, closure of pulmonary leaks

1 week–Years

Skin

High adhesion strength

Surgery

Poly(lactic-coglycolic acid), degradable hydrogels

Application

Function

Bone Fixatives

Provide support to the joint, alleviate pain/arthritis

Sutures

Lifetime—How Long Does the Biomaterial Need to Function?

Location—Where Will the Biomaterial Perform Its Task?

For any given device, the time that a biomaterial is required to perform a desired function must be known. Some materials are only required for a short period of time, on the order of days to week, serving a temporary function that allows for rapid tissue healing (e.g., sutures). Other biomaterials are required to last longer and degrade slowly over months and years as regrowing tissue infiltrates the scaffold (e.g., some bone fixation devices). In drug delivery, an important aspect is to deliver a drug at a constant rate over a predictable amount of time and at a dose that can be tuned for the patient (Sershen and West, 2002). The lifetime of a biomaterial can be controlled by varying different aspects of the polymer design (e.g., backbone functionality, porosity, shape). For example, diabetic patients require daily intravenous insulin injections, which could lead to “peak-and-valley-shaped” drug-concentration profiles in the body and cause gastrointestinal side effects and patient discomfort (Bratlie et al., 2012). Through the use of biodegradable polymeric nanoparticles (e.g., poly(3hydroxybutyrate-co-3-hydroxyhexanoate) embedded in a thermosensitive, injectable hydrogels), the release rate of insulin can be tightly controlled, a strategy with significant potential to meet the therapeutic needs of diabetic patients (Peng et al., 2013) (Fig. 1.3.2F.10). 

The body is made up of a range of different environments (e.g., different pH levels, high vs. low moisture environments, varied enzyme concentrations) that need to be considered collectively when designing a functional polymeric biomaterial for any given setting. For example, when designing gastric devices (e.g., weight loss balloons, ingestible diagnostic devices, oral drug delivery formulations) the biomaterial needs to reach its target area intact after traveling through areas with vastly different pH and enzyme profiles (Knipe et al., 2015). For example, Zhang et  al. designed a gastric-retentive device from a supramolecular hydrogel (e.g., poly(acryloyl 6-aminocaproic acid) and poly(methacrylic acid-co-ethyl acrylate) system) that could respond to changes in pH through protonation and deprotonation of carboxylic acid side chains (Zhang et al., 2015). The elastic nature of the polymer enabled packing of the structure into a gelatin capsule to allow for oral delivery. The material was designed to change shape as it moved through the digestive tract, retaining its extended shape in the gastric environment to perform its function (e.g., release of a drug or delivery of a gastric balloon) before degrading at neutral pH once it moved into the small intestines, preventing intestinal obstructions (Fig. 1.3.2F.11). This application demonstrates how the application of degradable polymers can increase the safety and efficacy of biomaterials. 

CHAPTER 1.3.2F   Degradable and Resorbable Polymers

179

• Figure 1.3.2F.10  (A) In vitro release of free insulin (circle), free insulin-loaded hydrogel (square) and nanoparticle-

loaded hydrogel (triangle) in PBS (pH 7.4) at 37°C, (B) Blood glucose level-time curve after subcutaneous injection into mice of blank blank hydrogel (circle), free insulin loaded hydrogel (square) and nanoparticle-loaded hydrogels (triangle) to male diabetic rats. Showing the nanogels had a remarkable retaining effect enabling the therapeutic effects to last 5 days. Data presented as mean ± SD (n = 5). (Adapted from Peng et al. (2013)).

Mechanical Properties—What Mechanical Properties Are Required for the Task? The mechanical properties required for the application should be determined based on both the location of the device and the timescale of degradation required (Engelberg and Kohn, 1991). While nondegradable devices have stable and predictable mechanical properties throughout their duration of use, degradable devices will exhibit a decrease in most mechanical properties (e.g., strength, modulus) as the polymer degrades. In some cases, especially those that involve implantation in a load-bearing site, it is critical to match the breakdown of the polymer mechanical properties with the biological repair time. For example, sutures must maintain sufficient tensile strength to keep the wound closed until it heals. The polymer used to make sutures should also allow doctors to easily stitch and tie the sutures without breaking. For orthopedic applications, polymeric devices with initially high mechanical strength and good fracture toughness, comparable to the mechanical properties of the native tissue, are required to support load bearing. Ideally, gradual stress transfer will occur as the biomaterial degrades and the tissue heals (Brauer et al., 2008) (Fig. 1.3.2F.12). 

Delivery—How Will the Biomaterial Reach the Required Site? The delivery mechanism by which the biomaterial will reach its intended site without degrading should also be carefully considered when designing degradable polymeric biomaterials. Popular modes of delivery are injection, surgical implantation, or ingestion. Injections are commonly used to deliver shear thinning hydrogels that serve as drug and cell delivery vehicles (Bakaic et  al., 2015). Injections minimize invasive surgery, leading to faster recovery times and reduced infection rates and costs. Furthermore, the material injected can fill voids and pack defects specific to the patient, without needing to preshape the material, which leads to improved comfort and efficacy. A range of different approaches is available to make degradable polymers injectable, including use of covalent adaptable

networks (e.g., oxime and hydrazine chemistries) or use of in situ-forming devices that rely on thermoresponsive chemistries (Kretlow et al., 2007), photopolymerization (Rydholm et al., 2005), or solvent-exchange strategies (Parent et al., 2013). Covalent Adaptable Networks Covalent adaptable networks are structures that incorporate exchangeable chemical bonds that undergo cleavage and reformation in response to an external stimulus (e.g., mechanical force, heat, base). An active species undergoes a reaction that results in bond exchange and the formation of new active species, which subsequently undergoes an additional exchange reaction (Bowman and Kloxin, 2012). Covalent adaptable chemistries used in polymeric biomaterials include boronate esters (Smithmyer et al., 2018), Diels–Alder (Kalaoglu-Altan et al., 2017), imines (i.e., oxime and hydrazone) (Boehnke et al., 2015), and thioester chemistries (Brown et al., 2018). The presence of these dynamic chemistries can lead to depolymerization or dynamic changes in mechanical properties. In tissue-engineering applications, covalent adaptable hydrogel networks have been synthesized to more closely mimic tissue viscoelasticity and support matrix deposition by encapsulated cells. For example, hydrazones and their network reorganization allow stress relaxation in a hydrogel with encapsulated, proliferating chondrocytes. The relaxation can be tuned to foster the deposition of a collagen and proteoglycan-rich matrix by the encapsulated cells (Richardson et al., 2019).

Injectable, in situ-forming devices or implants that rely on solvent-exchange strategies have been designed such that they set into their shape inside the body by relying on insolubility of the polymer in water. Specifically, degradable polymers (e.g., PLGA) are dissolved in a biocompatible solvent, such as N-methyl-2-pyrollidone or dimethylsulfoxide, before injection. Upon injection, solvent exchange occurs, increasing the concentration of water near the polymer as the solvent diffuses away from the injection site. When the surrounding concentration of water is sufficiently high, phase inversion occurs, causing the polymer to precipitate. The shape that the polymer takes depends on the concentration of polymer in the solvent and the miscibility of the solvent with water. The rate of in situ formation and the structure formed by the polymer, which are

180 SEC T I O N 1 . 3     Classes of Materials Used in Medicine

CHAPTER 1.3.2F   Degradable and Resorbable Polymers

181



Figure 1.3.2F.11 A pH-responsive supramolecular polymer gel as an enteric elastomer for use in gastric devices. (A) Supramolecular polymer gel network. Structures in yellow, synthesized poly(acryloyl 6-aminocaproic acid) (PA6ACA); structures in purple, linear poly(methacrylic acid-co-ethyl acrylate) (EUDRAGIT L); red, inter-polymer hydrogen bonds. (B) Folding of the ring into a standard gelatin 000 capsule by using the elasticity of the polymer gel. (C) Escape from the capsule and recovery to the ring shape after dissolution of the gelatin capsule in simulated gastric fluid at 37°C. (D) Recovery of the ring shape after delivery of an encapsulated ringshaped device through the esophagus and dissolution of the gelatin capsule in the stomach of a Yorkshire pig. (E) Schematic representation of the delivery and gastric retention of a ring-shaped device. (F) Schematic illustration of safe passage of PCL arcs through the small and large intestine on dissociation of the ring-shaped device as a result of the total or partial dissolution of enteric elastomer linkers. (G) X-ray image of a ring-shaped device residing in the gastric cavity of a Yorkshire pig. (H) X-ray image of four PCL arcs passing through the intestine after dissolution of the enteric elastomer linker. For visualization purposes, six to ten radio-opaque stainless steel beads (1mm diameter) were incorporated in every PCL arm. The total bead mass was 200 mg and the weight of the whole device (with iron beads) was 1,000 mg. Figure adapted from Zhang et al. (2015).

Composites—When Should a Composite Be Used and How Will Additives Affect Degradation?

• Figure 1.3.2F.12  The mechanical properties (e.g., modulus) of degrad-

able scaffolds decreases over time. In orthopedic applications, where the scaffold is expected to be load bearing, the mechanical properties as a function of time need to be characterized. In well-designed materials, gradual stress transfer will occur as the tissue re-grows.

dictated by polymer concentration and solvent identity, determine the burst and sustained release kinetics of drugs from the in situ-formed device (Parent et al., 2013). While noninvasive strategies are preferred for many applications, other applications, such as joint replacements, still require open surgery to get the biomaterial to the correct site in the body. The joint replacement is often a prefabricated, synthetic joint shaped to match the native joint. Alternatively, bone cement, which cures in situ to take the shape of the cavity, can be used. Both delivery techniques are invasive processes and pose a high risk of infection, causing postoperative pain and numerous side effects for the patient (ter Boo et al., 2015). To help prevent infections, prosthetic joints and implants have been designed with degradable polymer coatings that deliver a local dose of antibiotics, thereby reducing the changes of infection and localizing the antibiotics to a specific region of the body. A clinical example is ETN PROtect, where gentamicin-loaded PDLA (Fuchs et  al., 2011; Metsemakers et al., 2015) is coated on intramedullary nails to reduce infection. Clinical studies with this degradable polymer coating revealed no deep wound infections, good fracture healing, and increased weight-bearing capacity after 6 months compared to the nail alone (Fig. 1.3.2F.13). These results were found even in patients with complex fractures and multiple traumas. 

If the desired mechanical properties of a temporary support device or erosion kinetics of a controlled release system are not accessible through changing the degradable polymer composition alone, composites with nondegradable materials can be used. For example, low concentrations of natural materials (e.g., ceramic fibers or particles) are often embedded within degradable polymers used to increase the mechanical properties of degradable scaffolds for orthopedic applications (Rezwan et al., 2006). The effect of nondegradable additives on the final mechanical properties, and degradation, of the device/materials is dependent on both the loading density and hydrophobicity of the additive. For example, low loadings ( CP-Ti > Co-Cr-Mo > 316L (b) CP-Ti > AZ91D > Co-Cr-Mo > 316L (c) 316L > AZ91D > CP-Ti > Co-Cr-Mo (d) Co-Cr-Mo > 316L > CP-Ti > AZ91D 8. What are the advantages of using β-type Ti alloys over α-type Ti alloys for orthopedic applications? 9. Which of the following sets of properties should loadbearing implant materials have? (a) High elastic modulus, high fatigue strength, high tensile strength, high wear resistance, and high corrosion resistance (b) Low elastic modulus, high fatigue strength, high tensile strength, high wear resistance, high corrosion resistance (c)  High elastic modulus, high fatigue strength, high tensile strength, low wear resistance, high corrosion resistance 

Chapter Exercise 2 1. Explain the term “osseointegration” in metallic biomaterials. 2. Which of the following biomaterials exhibit the lowest wear resistance? (a) CP-Ti (Grade 1) (b) Co–20Cr–15W–10Ni (c) Stainless steel (316L) 3. What is the effect of interstitial oxygen atoms on the mechanical properties of Ti alloys? 4. Which of the following Ti alloys exhibit better coldforming ability?

(a) α-type Ti alloys (b) β-type Ti alloys

5. Why is fatigue strength important for metallic biomaterials? 6. Why do surface properties of Ti alloys differ from their bulk properties? 7. Which of the following contains information on critical anodization parameters? (a) Applied potential, type of electrolyte, time, temperature (b) Laser pulse intensity, laser energy, exposure time 8. What are the most commonly used coloring methods for Ti alloys? 9. What are the key steps involved in the additive manufacturing of Ti structures? 

Chapter Exercise 3 1. Which of the following materials exhibit shape memory property? (a) CP-Ti (Grade 1) (b) NiTi alloy (c) Stainless steel (316 SS) (d) Co-Cr-Mo (ASTM-F75) 2. Which of the following techniques can fabricate complex Ti geometrical shapes? (a) Powder metallurgy (b) Additive manufacturing (c) Casting 3. What are the complications of using stainless steels for orthopedic applications? 4. Why cannot α-type Ti alloys be strengthened by heat treatment? 5. Porous Ti alloys with the desired porosity ratio and pore size may be produced in the required geometric shape without the use of any space-holder material via: (a) Casting (b) Selective laser sintering (c) Powder metallurgy 6. Which of the following is an additive manufacturing process? (a) Cold rolling (b) Hot isostatic pressing (c) Electron beam melting 7. Co-based alloys have higher wear resistances than Ti alloys? (a) True (b) False 8. Which of the following methods is commonly used to grow oxide nanotubes on Ti alloys? (a) Forging (b) Anodization (c) Additive manufacturing 9. What are the key strategies to avoid stress shielding in orthopedic surgeries?

248.e1

1.3.3B

Stainless Steels PHILLIP J. ANDERSEN Andersen Metallurgical, LLC, Madison, WI, United States

Overview Stainless steel refers to the family of iron-based alloys having appreciable concentrations of nickel and chromium alloying elements. The term stainless derives from the corrosion resistance afforded by the protective surface oxide layer attributable to these alloying elements. Specific stainless steels are useful for biomedical applications due to their desirable shaping, joining, and mechanical properties, acceptable corrosion resistance, durability, and manufacturing economics. 

History Iron-based tools and weapons have been used since the end of the Bronze Age and the beginning of the historical record. Iron is a particularly advantageous material because material hardness and tensile strength are readily controlled by appropriate heating and quenching. A major disadvantage of iron is its propensity to react with oxygen to corrode (rust) and lose desirable mechanical properties. Humans recognized these characteristics for more than two millennia and this limited use of iron for many applications (prohibited it for biomedical use). Discovery within the past century, showing that adding appreciable amounts of chromium and nickel to iron produced a new alloy (now known as stainless steel), was an extraordinary technological development enabling many new uses of iron-based alloys in almost every area of human endeavor, especially biomedical. While there are a large number of different types of stainless steels available commercially, only a few of these alloys have properties that enable their use as biomaterials for implantable devices. Use of stainless-steel implants began in the 1920s and 1930s and applications for stainless steels expanded as new medical procedures and material refinements were developed. 

Composition and Types Stainless steel is an iron-based alloy that contains an appreciable percentage of chromium and nickel. There

are three main classes of stainless steels: austenitic, ferritic, and martensitic. These classes are differentiated by their microstructure. Austenitic stainless steels have austenite (gamma-phase iron) as their primary phase and contain chromium and nickel (sometimes manganese and nitrogen also). The 300-series of stainless steels, i.e., Types 301 to 304, Type 316 and Type 347, are those most commonly used for biomedical purposes because of their toughness, durability, ability to be joined by welding, corrosion resistance, and manufacturing economics (ESPI Metals, 2019). The basic composition of austenitic stainless steels is exemplified by Type 302 stainless steel. This alloy is composed of iron with 18% chromium and 8% nickel added. Austenitic steels are amenable to work hardening but cannot be hardened by heat treatment. In addition to favorable corrosion resistance and mechanical properties, austenitic stainless steel is desirable for implanted devices because it is nonmagnetic and thus has minimal or no interactions with the intense magnetic fields used in magnetic resonance imaging (MRI). The most familiar austenitic stainless steel used for surgical implants is Type 304, which contains 18%–20% chromium and 8%–10% nickel (National Specialty Alloys, 2011). Ferritic and martensitic stainless steels, as well as precipitation-hardened and cast stainless steels, lack the desirable corrosion and mechanical properties of austenitic stainless steels. For these reasons, they are not commonly used for biomedical applications and will not be discussed further. Composition of stainless steels used for biomedical applications is closely controlled because under some processing conditions, it is possible to form different phases within the austenitic stainless-steel alloys, e.g., delta ferrite (which is magnetic). Modest amounts of heating and material displacement can occur for delta ferrite phases when exposed to the large magnetic fields accompanying MRI. Such localized heating or material motion in implants in patients exposed to high magnetic fields during MRI is deleterious, hence the need to control the composition of stainless-steel alloys used for implants.  249

250 SEC T I O N 1 . 3     Classes of Materials Used in Medicine

Structure Metals, including stainless steel alloys, are characterized by a high percentage of crystallinity. The constituent crystalline structures in metals have one of three basic types of unit cells: face-centered cubic, body-centered cubic, or hexagonal close packed. Stainless steels used for implants are face-centered cubic (Fig. 1.2.3.2, Chapter 1.2.3). Table 1.3.3B.1 shows the ranges of chemical compositions specified for several common implantable stainless steels. Chromium is present in these alloys and its primary purpose is to form a protective Cr2O3 surface layer (passive film) that confers corrosion resistance. Since chromium stabilizes the ferrite (body-centered cubic) phase, other alloying elements are added to stabilize the desired austenite phase. This objective is attained by addition of nickel, manganese, and nitrogen. Nitrogen added as an alloying element to stainless steel also increases mechanical strength and corrosion resistance. Molybdenum additions have a beneficial impact on the pitting corrosion resistance of stainless steels. Carbon content in stainless-steel alloys must be limited to small percentages (the “L” in 316L designates low carbon) to prevent formation of chromium carbides. Formation of these carbides can result in a phenomenon known as sensitization. If sufficient carbon is available, chromium carbides can form when austenitic stainless steels are held at temperatures in the range of 450–815°C. The time required to form these carbides depends on temperature. Under select time and hightemperature conditions, carbides tend to form preferentially along grain boundaries, leaving the adjacent areas with depleted chromium levels. Such chromiumdepleted areas then become prone to corrosion and this can lead to macroscopic material (implant) failure. Slow cooling after welding is a classic means by which sensitization can occur when time and cooling rates are unmanaged. Examination of the chemical compositions of surgically implantable stainless steels (Table 1.3.3B.1) shows variability with time and alloy selection. While 316 and 316L stainless steels have been used successfully for implants for many decades, stronger and more corrosion-resistant alloys have subsequently been developed. The alloys, e.g., Rex 734 (also known as Ortron 90) and 22-13-5, were developed in the 1980s. They contain higher levels of chromium, manganese, and nitrogen compared to 316L, leading to improved mechanical properties and enhanced corrosion resistance. Recent concerns for adverse (allergic) patient responses to nickel (see Orthopedic Applications, Chapter 2.5.4 for additional discussion of this topic) have led to the development of stainless steels that are essentially nickel free. One example is BioDur 108 (BioDur, 2008), a stainless-steel alloy that uses manganese and nitrogen instead of nickel to stabilize the austenite phase. The nitrogen level in this alloy is much greater than the level found in the Type-300 family of stainless steels. 

Structure, Composition, and Processing Effects on Mechanical Properties The mechanical properties of metals and alloys depend on their structure, chemical composition, and processing history. To understand how these factors influence mechanical properties, consider the mechanisms involved when permanent (plastic) deformation occurs. Plastic deformation results from mechanical stress-induced movement of atoms within the crystalline structure of a metal. These movements result in irregularly distributed atoms that cause defects or disturbed regions in the crystal lattice. These disturbances are known as “dislocations.” Subsequent straininduced movement of atoms in the presence of dislocations is typically more difficult, and in this event the mechanical properties of the material change. Typical changes accompanying increased dislocation density include increases in yield strength and decreases in ductility. Chemical composition of an alloy influences mechanical properties by a process known as solid solution strengthening. Metallic alloying elements in stainless steels, i.e., chromium, nickel, etc., replace iron atoms at random locations within the crystal structure. Since the replacement atoms differ in size compared to iron atoms, the crystal lattice can be distorted. This distortion makes material deformation via dislocation movement more difficult and this also increases yield strength as well as generally decreases ductility. Smaller atoms, such as nitrogen, fit within the gaps between the larger metallic atoms and for this reason they are often referred to as interstitials. Additions of these chemical species leads to substantial strength increases due to interactions between solute atoms and dislocations. Processing also affects the mechanical properties of stainless steels. While processes are commonly used to alter mechanical properties of other iron-based alloys, e.g., aging reactions during heat treatment, such processes are not applicable to austenitic stainless steels. Instead, material strengthening mechanisms such as work hardening, also known as “cold working,” are used. Work hardening involves repeated mechanical stresses applied to an unheated material. Work hardening occurs by mechanical stress-induced introduction of increasing amounts of dislocations. As the percentage of dislocations increases, motion of these dislocations becomes more difficult as more dislocations interact within the material. As the material continues deforming by the applied mechanical stress, increasing quantities of dislocations develop within the grains of the alloy. If the deformation process takes place at high temperatures typical of the large-scale processes used to produce metal products, or if the material is heated above its recrystallization temperature after deformation, then the dislocations are removed by formation of new, annealed grains (recrystallization). Deformation at lower temperatures does not eliminate dislocations. As the density of dislocations within the material increases, they interact causing increasing difficulty for further dislocation motion. This leads to increases

TABLE 1.3.3B.1   Compositions of Common Implantable Stainless Steels (Weight Percent)

Alloy

Cr

Ni

Mn

Mo

C

N

Nb

V

Si

Cu

P

S

316L ASTM F138, ISO 5832-1

17–19

13–15

10%) and hydrolysis was claimed to produce environmental stress cracking. The cracked coating allegedly provided a pathway for bacteria to travel from the vagina into the uterus along the filaments, resulting in significant pelvic inflammatory disease (Hudson and Crugnola, 1987). Degradation of a poly(arylamide) intended for orthopedic use (the fiber-reinforced polyamide from m-xylylene

CHAPTER 2.4.2   Chemical and Biochemical Degradation of Polymers Intended to Be Biostable

diamine and adipic acid) was also shown in a rabbit implant study. Although the material provoked a foreign body reaction comparable to a polyethylene control, surface pitting associated with resolving macrophages was noted at 4 weeks and became more pronounced by 12 weeks. This result was not predicted, since poly(arylamides) are very resistant to solvents and heat (Finck et al., 1994). Polyamides with long aliphatic hydrocarbon chain segments [e.g., poly(dodecanamide), Nylon 12] (Duke Extrusion.com, Copyright, 2019) such as that used for intraarterial balloon catheters are more hydrolytically stable than shorter-chain nylons, and correspondingly degrade slower in vivo. Poly(Alkyl Cyanoacrylates). This class of polymers used as tissue adhesives is noteworthy as a rare case in which carbon–carbon bonds are cleaved by hydrolysis (Fig. 2.4.2.1). This occurs because the methylene (–CH2–) hydrogen in the polymer is highly activated inductively by electronwithdrawing neighboring groups. Formation of the polymer adhesive from monomers is base catalyzed, with adsorbed water on the adherend being basic enough to initiate the reaction. Catalysts for equilibrium reactions affect the reverse, as well as the forward, reaction. Therefore, water associated with tissue can induce poly(cyanoacrylate) hydrolysis by a “reverse Knoevenagel” reaction (Fig. 2.4.2.1). More basic conditions and (as suggested by in  vitro cell culture or implant studies) enzymatic processes are much more effective at inducing hydrolysis. In chick embryo liver culture (a rich source of a variety of enzymes), poly(methyl cyanoacrylate) degraded much faster than in cell culture medium alone. In animal implants, poly(methyl cyanoacrylate) was extensively degraded within 4–6 months (Kopecek and Ulbrich, 1983). Higher alkyl (e.g., butyl) homologs degraded slower than the methyl homolog, and were less cytotoxic (Hegyeli, 1973; Vauthier et  al., 2003). Octyl cyanoacrylate polymer, introduced to the device field as a dermal tissue adhesive (Singer et al., 2008), promises to be the most stable cyanoacrylate device composition to date, based on increased hydrophobicity. Preclinical and clinical studies have yielded promising results with the use of octyl cyanoacrylate as a surgical sealant (Barbarini Ferraz et  al., 2009; Carr, 2011), and, more recently, a composition containing the monomers octyl cyanoacrylate and butyl lactoyl cyanoacrylate was FDA (US Food and Drug Administration) approved as a blood vessel anastomotic sealant (US FDA, 2010). 

925

over decades of use, to provide reliable, stable service, with hydrolysis rates being so slow as to be inconsequential. Another polymer system with a hydrocarbon backbone, poly(methyl acrylate-co-2-hydroxyethyl acrylate) also contains hydrolyzable ester side groups. This polymer, which forms hydrogels in an aqueous environment, has been used as a “scleral buckling” device for retinal detachment surgery. Basically, the dry polymer, shaped as a band or ring, placed as a “belt” around the sclera, expands through hydration to create an indentation in the zone of the retinal detachment to reestablish retinal contact. The device is left in place as a permanent implant (or “exoplant” as it is sometimes called because it is external to the sclera) (Braunstein and Winnick, 2002). This hydrogel device, introduced into clinical practice in the 1980s (Refojo and Leong, 1981; Colthurst et al., 2000), apparently performed satisfactorily for years as an approved product. However, in the 1990s, reports of longterm complications of these hydrogel scleral buckles began to surface (Hwang and Lim, 1997; Roldan-Pallares et  al., 1999). The hydrogel structures resumed swelling, sometimes with fragmentation, after maintaining stable dimensions for years. One report described a difficult explantation 13 years after implantation (Braunstein and Winnick, 2002). An article described three unique complications from hydrogel scleral buckle use: orbital cellulitis mimicry; fornical shortening with orbital prosthetic intolerance; and orbital pseudotumor (Bernardino et al., 2006). Those buckles remained in place for 7–15 years, with a mean time of 10.7 years. In a study of 15 patients with 17 scleral buckles, all reported complications within 4–14 years (Figs. 2.4.2.4–2.4.2.6). Removal of the buckles was technically difficult and postoperative complications were significant, although immediate palliative relief was experienced after surgery. Pressure applied to the eye by the swelling has led to blindness and loss of the eyeball. Hydrogel scleral buckles are no longer used in retinal surgery (Watt, 2001).

Polymers Containing Hydrolyzable Pendant Groups Certain polymers intended for long-term implantation consist of biostable main chain sequences and hydrolyzable pendant groups. Poly(methyl methacrylate) (PMMA) used in bone cements and intraocular lenses is an example of a hydrophobic polymer with a stable hydrocarbon main chain and hydrolyzable ester side groups. It has been proven,

• Figure 2.4.2.4  A subconjunctival and subpalpebral hydrogel explant

interferes with ocular motility, especially on attempted gaze up and to the right. (From Kearney, J.J., Lahey, J.M., Borirakchanyavat, S., Schwartz, D.M., Wilson, D., et  al., 2004. Complications of hydrogel explants used in scleral buckling surgery. Am. J. Ophthalmol. 137 (1), 96–100.)

926 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Very little speculation has been provided in published articles about the mechanism of failure of acrylate scleral buckling devices other than that chemical degradation has occurred (Roldan-Pallares et  al., 1999). I suggest that a likely mechanism involves hydrolysis of the ester side groups enhanced by the hydrophilic nature of the polymer (as contrasted to hydrophobic polymers such as PMMA). Hydrolysis of either of the two acrylate esters in the polymer chain provides an acrylic acid moiety. Linear poly(acrylic acid) is fully water soluble, and each hydrolytic event renders the polymer more hydrophilic and subject to enhanced swelling. This process is slow but inexorable in the case of the scleral buckling device. The valuable lesson is that devices with intrinsically susceptible groups can eventually degrade by predictable mechanisms. This may take longer than is required for pivotal preclinical qualification studies (typically two-year animal implants). In fact, Bernardino et  al. (2006), in discussing long-term hydrogel implants, warn: “Patients with newer uses of hydrogel, such as orbital expanders, should also be observed for long-term



Figure 2.4.2.5 A washed 4-mm hydrogel explant showing multiple fragments of a swollen explant and a 3-mm silicone sponge explant for comparison. The sponge was removed at the same surgical sitting as the hydrogel explant. (From Kearney, J.J., Lahey, J.M., Borirakchanyavat, S., Schwartz, D.M., Wilson, D., et al., 2004. Complications of hydrogel explants used in scleral buckling surgery. Am. J. Ophthalmol. 137 (1), 96–100.)

complications.” If late degradation is suspected, therefore, accelerated aging studies should be performed in vitro with correlations made to in vivo studies. Such efforts give valuable, if not completely trustworthy, information (see Section Polymer Degradation Processes). 

Oxidative Biodegradation Oxidation Reaction Mechanisms and Polymer Structures While much is known about the structures and reaction products of polymers susceptible to oxidative biodegradation, confirmation of the individual reaction steps has not yet been demonstrated analytically. Still, mechanistic inferences are possible from extensive knowledge of physiological oxidation processes and polymer oxidation in vitro. The polymer oxidation processes to be discussed may be consistent with a homolytic chain reaction or a heterolytic mechanism. Species such as carbonyl, hydroxyl, and chain scission products are detectable in these types of processes. Classic initiation, propagation, and termination events for homolysis and ionic heterolytic processes are detailed in Fig. 2.4.2.7. The principles of polymer degradation resistance stated in the section on hydrolyzable polymers (e.g., group frequency, crystallinity, hydrophobicity) are valid for predicting relative oxidation resistance of polymers, except where particularly oxidation-susceptible groups are present. Sites favored for initial oxidative attack, consistent with a homolytic or heterolytic pathway, are those that allow abstraction of an atom or ion and provide resonance stabilization of the resultant radical or ion. Fig. 2.4.2.8 provides a selection of readily oxidized groups and the atom at which initial attack occurs. In Fig. 2.4.2.9, examples of radical and ion stabilization by resonance in ether and branched hydrocarbon structures are provided. Peroxy, carbonyl, and other radical intermediates are stabilized by similar resonance delocalization of electrons from the elements C, O, H, or N. Within Fig. 2.4.2.9, a coupling reaction which can lead to polymer cross-linking as a consequence of oxidative degradation is shown. Fig. 2.4.2.10

• Figure 2.4.2.6  (Left panel) Computed tomography scan showing a hydrogel implant in the anterior orbit of the right eye. (Right panel) Magnetic resonance image of the left orbit shows an expanded hydrogel explant. (From Kearney, J.J., Lahey, J.M., Borirakchanyavat, S., Schwartz, D.M., Wilson, D., et al., 2004. Complications of hydrogel explants used in scleral buckling surgery. Am. J. Ophthalmol. 137 (1), 96–100.)

CHAPTER 2.4.2   Chemical and Biochemical Degradation of Polymers Intended to Be Biostable

incorporates the points made in the previous three figures and illustrates the general sequence of oxidative degradation of an ether function by cleavage at the carbon–oxygen bond which would lead to chain scission of polyether-containing polymers. Two general categories of oxidative biodegradation, based on the source of initiation of the process, are direct oxidation by the host and external environment-mediated oxidation. 

927

Direct Oxidation by Host In these circumstances, host-generated molecular species effect or potentiate oxidative processes directly on the polymer. Current thinking, based on solid analytical evidence, is that such reactive molecules are derived from activated phagocytic cells responding to the injury and the properties of the foreign body at the implant site (Zhao et  al., 1991). The two major types of these cells are the neutrophils

• Figure 2.4.2.7  Proposed homolytic chain reaction and heterolytic oxidation mechanisms.

928 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment



Figure 2.4.2.8 Readily oxidizable functional groups (* is the site of homolysis or heterolysis).

(polymorphonuclear leukocytes, PMNs) and the monocytes. The latter, which are found in circulation, can differentiate upon attachment to tissue and divide into macrophage and foreign-body giant cell (FBGC) phenotypes (see Chapter 2.2.2) (Ziats, 1988). Much work is under way to elucidate the sequence of events leading to phagocytic oxidation. Certain important processes of wound healing in the presence of biologically derived foreign bodies such as bacteria and parasites have shown some relevance to biomaterial implants (Northup, 1987). Neutrophils, responding to chemical mediators at the wound site, mount a powerful but transient chemical attack within the first few days of injury (Test and Weiss, 1986; Northup, 1987). Chemically susceptible biomaterials may be affected if they are in close apposition to the wound site (Sutherland et  al., 1993). Activated macrophages subsequently multiply and subside within days at a benign wound site or in weeks if stimulants such as toxins or particulates are released at the site. Their fusion products (FBGCs) can survive for months to years on the implant surface. Macrophages also remain resident in the collagenous foreign-body capsule for extended periods. While it is recognized that the complete mechanism of cellular attack and oxidation of biomaterials is as yet

• Figure 2.4.2.9  (A) Resonance stabilization of ether and hydrocarbon radicals; (B) resonance stabilization of ether and hydrocarbon cations.

CHAPTER 2.4.2   Chemical and Biochemical Degradation of Polymers Intended to Be Biostable

unconfirmed, the following discussion attempts to provide logical biological pathways to powerful oxidants capable of producing known degradation products. Both PMNs and macrophages metabolize oxygen to form − a superoxide anion ( O2 • ). This intermediate can undergo transformation to more powerful oxidants, and conceivably can initiate homolytic reactions on the polymer. Superoxide

• Figure 2.4.2.10  Pathways for oxidative fragmentation of polyethers.

929

dismutase (SOD), a ubiquitous peroxidase enzyme, can catalyze the conversion of superoxide to hydrogen peroxide which, in the presence of myeloperoxidase (MPO) derived from PMNs, is converted to hypochlorous acid (HOCl). A potent biomaterial oxidant in its own right (Coury et  al., 1987), hypochlorite (ClO−)2 can oxidize free amine functionality (e.g., in proteins) to chloramines which can perform as long-lived sources of chlorine oxidant (Test and Weiss, 1986; Figs. 2.4.2.11 and 2.4.2.12). Hypochlorite can oxidize other substituted nitrogen functional groups (amides, ureas, urethanes, etc.) with potential chain cleavage of these groups. The following paragraphs describe potential cooperative reactions involving acquired peroxidase and free ferrous ions. Macrophages contain essentially no MPO, so their hydrogen peroxide is not normally converted to HOCl. However, PMN-derived MPO can bind securely to foreign-body surfaces (Locksley et al., 1982), and serve as a catalyst reservoir for macrophage- or FBGC-derived HOCl production. If free ferrous ion, which is normally present in negligible quantities in the host, is released to the implant site by hemolysis or other injury, it can catalyze the formation of the powerfully oxidizing hydroxyl radical via the Haber–Weiss cycle (Klebanoff, 1982; Fig. 2.4.2.10). Fig. 2.4.2.11 shows radical and ionic intermediates of HOCl that may initiate biomaterial oxidation. Fig. 2.4.2.12 is a diagram showing a leukocyte phagocytic process that employs endogenous MPO catalysis of HOCl formation. In a more general sense, the MPO may come from within or outside of the cell. The foregoing discussion of sources of direct oxidation is focused primarily on acute implant periods in which bursts of PMN activity followed by macrophage activity normally resolve within weeks. However, since the foreign body subsequently remains implanted, a sustained if futile attempt to phagocytose an implanted device provides a prolonged release of chemicals onto the biomaterial. This phenomenon, called exocytosis, occurs over months to possibly years (Zhao et  al., 1990), and results primarily from the macrophage-FBGC line. It can contribute to long-term chemical degradation of the polymer. The oxidation processes induced by phagocytes are the result of oxidants produced by the general foreign-body response (innate immune response), not direct receptor–ligand catalysis by oxidase enzyme (adaptive immune response). Attempts to degrade oxidatively susceptible polymers by direct contact with oxidase enzymes have produced shortrange or limited effects (Sutherland et al., 1993; Santerre et al., 1994). Macrophages mediate other processes, such as fibrous capsule formation, around the device. Their release of cellular regulatory factors stimulates fibroblasts to populate the implant site and produce the collagenous sheath. Any knowledge of the effects of factors such as fibroblasts or fibrous capsules on rates and mechanisms of polymer degradation is, at this time, incompletely understood. 

930 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

• Figure 2.4.2.11  Generation of potential oxidants by phagocytic processes.

• Figure 2.4.2.12  Hypochlorous acid: formation and potential reaction intermediates.

CHAPTER 2.4.2   Chemical and Biochemical Degradation of Polymers Intended to Be Biostable

Stress Cracking An important category of host-induced biodegradation with an oxidative component is stress cracking as manifested in poly(ether urethane) elastomer implants. It differs from classic environmental stress cracking (ESC), which involves a susceptible material at a critical level of stress in a medium that may permeate and swell the polymer but does not dissolve the polymer. Classic ESC is not accompanied by significant chemical degradation (Stokes, 1988). In contrast, stress cracking of polyurethanes is characterized by surface attack of the polymer, and by chemical changes induced by relatively specific in vivo or in vitro oxidizing conditions. Conditions relevant to stress cracking of certain poly(ether urethane) compositions are stated in Table 2.4.2.3.  Experimental information on the stress cracking of poly(ether urethanes) and poly(ether urethane ureas) (e.g., TABLE   Characteristics of Poly(Ether Urethanes) That 2.4.2.3  Cracked In Vivo Components contained residual processing and/or applied mechanical stresses/strains Components were exposed to a medium of viable cellular and extracellular body constituents Polymers had oxidatively susceptible (aliphatic ether) groups Analysis of polymers showed surface oxidation products

931

the polymers described in Fig. 2.4.2.3) has provided insights which may be valid for these and other compositions that can be oxidized, for example, polypropylene (Liebert et al., 1976; Altman et al., 1986) or polyethylenes (Wasserbauer et al., 1990; Zhao et al., 1995). Poly(ether urethane) compositions containing low polyether content, which are resistant to degradation in vivo, are used as connectors (also called “headers”), insulators, and adhesives for cardiac pacemakers and neurological stimulators (Fig. 2.4.2.13). They have performed with high reliability in chronic clinical applications since 1975. Certain poly(ether urethane) pacing leads have displayed surface cracks in their insulation after residence times in  vivo of months to years. These cracks are directly related in frequency and depth to the amount of residual stress (Figs. 2.4.2.14 and 2.4.2.15) and the ether (soft segment) content of the polyurethane (Coury et al., 1987; Martin et al., 2001). Morphologically, the rough-walled cracks display regular patterns predominately normal to the force vectors, occasionally with “tie fibers” bridging the gaps, indicative of ductile rather than brittle fracture (Figs. 2.4.2.16 and 2.4.2.17). Infrared (IR) analysis indicates that oxidation does not take place detectably in the bulk, but only on the surface where extensive loss of ether functionality (as seen in the ether IR stretch, 1110 cm−1) and enhanced absorption in the hydroxyl and carbonyl regions are observed (Stokes et  al., 1987). Possible mechanisms for the oxidative degradation of ethers are

• Figure 2.4.2.13  Activation of phagocyte redox metabolism: chemiluminigenic probing with luminol and

lucigenin. (From Allen, R.C., 1991. Activation of Phagocyte Redox Metabolism: Chemiluminigenic Probing with Luminol and Lucigenin. Drawing provided (Personal communication).)

932 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

• Figure 2.4.2.17  Stress crack pattern (frosting) near tight ligature (×14). •

Figure 2.4.2.14 Cardiac pacemaker with polyurethane lead insulation, tine, and connector. (Courtesy of Medtronic, Inc.)



Figure 2.4.2.18 Single stress crack in pacemaker lead tubing with rough walls and “tie fibers” indicative of ductile fracture (×700).



Figure 2.4.2.15 Pellethane 2463-80A pacemaker lead tubing with high applied radial stress showing total breach.

• Figure 2.4.2.16  Pellethane 2363-80A pacemaker lead tubing showing “frosting” due to stress from tight ligature.

presented in Fig. 2.4.2.18. The participation of molecular oxygen in the degradation mechanism is supported by studies which showed that poly(ether urethane urea) degradation in  vitro correlated with oxygen diffusion into the polymer bulk after surface oxidation was initiated by hydrogen peroxide/cobalt chloride (Schubert et al., 1997a,b). In a seminal study, Zhao et  al. (1990) placed polyurethane tubing under strain in cages permeable to fluids and cells (therefore under high initial stress, which was subject to subsequent stress relaxation), and implanted them in rats. In certain cases, antiinflammatory steroids or cytotoxic polymers were coimplanted in the cages. Implants of up to 15 weeks were retrieved. The only prestressed samples to crack were those that did not reside in the cages with the cytotoxic coimplants. The authors concluded that adherent cells caused the stress cracking, and cell necrosis or deactivation inhibited crack induction. Subsequently, viable phagocytic cells were implicated as a cause of crack initiation in vivo (Zhao et al., 1991). By removing adherent foreign-body giant cells after 10-week implantation of a curved poly(ether urethane urea) film in a wire cage, exposed foreign-body cell “footprints” showed localized surface cracking in the order of several microns deep and wide.

CHAPTER 2.4.2   Chemical and Biochemical Degradation of Polymers Intended to Be Biostable

Adjacent areas of polymer which were devoid of attached cells were not cracked. Owing to relatively low stresses in the implanted film, deep crack propagation was not observed. In vitro studies of strained (Stokes, 1988) and unstrained poly(ether urethane) films (Phua et al., 1987; Ratner et al., 1988; Bouvier et al., 1991; Wiggins et al., 2003) using oxidants, enzymes, etc., have sought to duplicate in vivo stress cracking. Although some surface chemical degradation with products similar to those seen in vivo was demonstrated, stress crack morphology was most closely matched in vitro in two studies. A test which involves immersing stressed poly(ether urethane) tubing in a medium of glass wool, hydrogen peroxide, and cobalt chloride produces cracks which duplicate those produced in  vivo, but with rate acceleration of up to seven times (Zhao et  al., 1995). These investigators also showed that human plasma proteins, particularly alpha2-macroglobulin and ceruloplasmin, enhance in vitro stress cracking by oxidants in patterns morphologically similar to those observed in vivo (Zhao et al., 1993). The potential of macrophages to contribute to stress cracking of poly(ether urethanes) was verified in an in vitro study which succeeded in potentiating macrophage oxidative effects with ferrous chloride, and inhibiting them with the antiinflammatory steroid dexamethasone (Casas et  al., 1999). In another study, comparable crack patterns were produced when specimens of stressed tubing in rats were compared with those incubated with PMNs in culture (Sutherland et al., 1993). Moreover, this study revealed a difference in chemical degradation products with time of implants which correlated with products from oxidants generated primarily by PMNs (HOCl) and macrophages (ONOO−). Early implant times, activated PMNs, and HOCl caused a preferential decrease in the urethane oxygen stretch peak, while longer implant times and ONOO− caused selective loss of the aliphatic ether stretch peak (by infrared spectroscopy). Taken together, the foregoing observations are consistent with a hypothesized two-step mechanism for stress cracking in vivo. In the first step, surface oxidation induces very shallow, brittle microcracks. The second step involves propagation of the cracks in which specific body fluid components act on the formed cracks to enhance their depth and width, without inducing major detectable bulk chemical reactions. Should this hypothesis prove correct, the term “oxidation-initiated stress cracking” would be reasonably descriptive. The above description of stress cracking has generally considered static stress, such as that entrained in polymers during the cooling of molten parts, or the assembly of components. Dynamic stresses and strains such as those occurring during cooling with stretching as with extruded tubing or by the operation of diaphragm or bladder heart pumps or artificial joints can cause related cracking in areas of high flex. The cracking has been purported to increase with time of device operation, but to display only minor surface chemical changes (Tomita et al., 1999; Wu et al., 1999). The stress cracking related to entrained stress in the polymer has been controlled by reducing residual stress,

933

• Figure 2.4.2.19  Random crack pattern of Pellethane 2363-80A lead insulation caused by metal ion-induced oxidation (×480).

isolating the polymer from cell contact (Tang et al., 1994), protecting the polymer from stress-cracking media, or using stress crack-resistant polymers (e.g., in the case of polyurethanes, ether-free compositions) (Coury et al., 1990; Takahara et al., 1994; Tanzi et al., 1997), and use of antioxidants such as hindered phenols (e.g., vitamin E, Monsanto Santowhite powder) (Schubert et al., 1997a,b). Stress cracking is next compared with another type of degradation, metal ion-induced oxidation. Particularly alpha-2-macroglobulin and ceruloplasmin, enhance in vitro stress cracking by oxidants in patterns morphologically similar to those observed in vivo (Zhao et al., 1993). 

Device- or Environment-Mediated Oxidation Metal Ion-Induced Oxidation. A process of oxidative degradation that has, thus far, primarily been reported clinically for poly(ether urethane) pacemaker leads, requires, as does stress cracking, a very specific set of conditions. The enabling variables and fracture morphology are quite different from stress cracking, although oxidative degradation products are similar. Biodegradation of implanted devices through stress cracking always occurs on polymer surfaces exposed to living cells and provides characteristic rough-walled fissures (indicative of ductile fracture) oriented perpendicular to the stress vector (Figs. 2.4.2.14–2.4.2.17). Metal ion-induced oxidation initiates on the enclosed inner surfaces of pacing lead insulation near corroded metallic components and their entrapped corrosion products. Smooth crack walls and microscopically random crack orientation is indicative of brittle fracture (Figs. 2.4.2.19 and 2.4.2.20). Macroscopically, crack patterns that track metal component configurations may be present (Fig. 2.4.2.21). Degradation products that may be found deeper in the bulk than with stress cracking are again indicative of brittle fracture. This phenomenon, called metal ion-induced oxidation (MIO), has been confirmed by in  vitro studies, in which poly(ether urethanes) were aged in metal ion solutions of

934 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

TABLE   Effect of Metal Ion Oxidation Potential 2.4.2.4  on Properties of Poly(Ether Urethane)

(Pellethane 2363-80A)a

• Figure 2.4.2.20  Smooth crack wall indicative of brittle fracture caused by metal ion-induced oxidation (×830).

Aqueous Solution

Standard Oxidation Potential

Change in Tensile Strength (%)

Change in Elongation (%)

PtCl2

~ +1.2

−87

−77

AgNO3

+0.799

−54

−42

FeCl3

+0.771

−79

−10

Cu2Cl2

+0.521

−6

+11

Cu2(OAc)2

+0.153

−11

+22

Ni(OAc)2

−0.250

−5

+13

Co(OAc)2

−0.277

+1

+13

aConditions:

0.1 M solutions/90°C/35 days versus controls aged in deionized water; ASTM (D-1708) microtensile specimens; specimens were tested wet.

TABLE   Effect of Ether Content of Poly(Ether 2.4.2.5  Urehane) on Susceptibility to Metal Ion-

Induced Oxidationa

• Figure 2.4.2.21  Crack pattern on inner lumen of poly(ether urethane)

lead insulation tracking coil imprint indicative of metal ion-induced oxidation (×100).

different standard oxidation potentials. Above an oxidation potential of about +0.77, chemical degradation was severe. Below that oxidation potential, changes in the polymer that are characteristic of simple plasticization were seen (Coury et al., 1987, Table 2.4.2.4). This technique also showed that metal ion-induced oxidation was proportional to the ether content of the polyurethane (Table 2.4.2.5). The effect of various metals on oxidation in  vitro and in vivo has also been studied. Different metallic components of pacing lead conductors were sealed in poly(ether urethane) (Dow Pellethane 2363-80A) tubing and immersed in 3% hydrogen peroxide at 37°C for up to 6 months (Stokes et al., 1987) or implanted in rabbits for up to 2 years (Stokes et al., 1990). Both techniques resulted in corroded metals and degraded tubing lumen surfaces under certain conditions within 30 days. In particular, the in vivo interaction of body fluids with cobalt and its alloys resulted in oxidative cracking of the polymer. The metal ion-induced oxidation process clearly involves corrosion of metallic elements to their ions, and subsequent oxidation of the polymer. In operating devices, the metal ion may be formed by solvation, galvanic or electrolytic corrosion, or chemical or biochemical oxidation (Fig. 2.4.2.22).

Poly(Ether Urethane)

Polyether Content

Change in Tensile Strength (%)

Pellethane 2363-80A

High

−54

−42

Pellethane 2363-55D

Low

−23

−10

Model segmented polyurethane

None

+9

+3

Change in Elongation (%)

aConditions:

0.1 M AgNO3/90°C/35 days versus controls aged in deionized water; ASTM (D-1708) microtensile specimens.

• Figure 2.4.2.22  Formation of metal ion from metal.

In turn, these metal ions develop oxidation potentials that may well be enhanced in body fluids over their standard half-cell potentials. As strong oxidants, they produce intermediates or attack the polymer to initiate the chain reaction (Fig. 2.4.2.23). Metal ion-induced oxidation is, therefore,

CHAPTER 2.4.2   Chemical and Biochemical Degradation of Polymers Intended to Be Biostable

• Figure 2.4.2.23  Initiation of oxidation pathways by metal ions.

the result of a highly complex interaction of the device, the polymer, and the body. Should metal ion-induced oxidation, specifically, be a possibility in an implanted device, several approaches are available to control this problem. They are not universally applicable, however, and should be incorporated only if functionality and biocompatibility are retained. Potentially useful techniques include using corrosion-resistant metals, “flushing” corrosive ions away from the susceptible polymer, isolating the metals and polymer from electrolyte solutions, incorporating appropriate antioxidants, and using oxidation-resistant polymers. 

Chemical Structure Strategies to Combat Oxidation To address oxidation susceptibility, polyurethane elastomers with enhanced oxidation stability have been developed. They should be ester-free and ether-free or deficient, and generally contain unconventional soft segments, including, for example, hydrogenated polybutadiene, polyisobutylene, polydimethylsiloxane (subject to hydrolytic consideration, above), polycarbonate, and dimerized fatty acid derivatives (Coury et al., 1990; Pinchuk et al., 1991; Takahara et al., 1991, 1994; Kato et al., 1995; Mathur et al., 1997; Hernandez et al., 2007; Kang et al., 2010; Desai et al., 2011). In implant tests, they have shown reduced tendency to stress crack oxidatively, and some have shown high resistance to metal ion oxidants in vitro. Early attempts to stabilize polyurethanes by laminating more biostable polymers (such as silicone rubbers) to tissue-facing surfaces have met with limited success in dynamic applications, due to the delamination tendencies Pinchuk (1992). Other approaches to stabilizing polyurethanes to oxidative attack in situ have involved the use of surface modifying (“blooming”) macromolecules (SMMs) (Santerre et al., 2000), surface-modifying end groups (SMEs) (Ward et al., 1995, 1998), and, more recently, self-assembling monolayer end groups (SAMEs) (Ward, 2008). SMMs, typically fluorocarbon-based polymers, are blended with the polyurethanes during processing, and migrate to the surface prior to implantation. SMEs are moieties (typically polysiloxane) bonded to polyurethane as end groups. SAMEs are formed from molecules appended to the backbone of polymers

935

during synthesis. Such molecules consist of three components: a chemically reactive group for conjugation to the polymer as it is being made; typically, a hydrophobic spacer chain, to provide self-assembly at the surface of the formed article; and a head group from a variety of surface chemistries to provide specific bioactivity. One potential function is to stabilize the polymer against biodegradation (Ward, 2008). The covalently modified polyurethanes may be used in bulk or as additives to conventional polyurethanes. Both approaches have provided enhanced in  vivo stability for polyurethane implants; however, the long-term effects of these treatments are not, as yet, confirmed. SMMs have been covalently modified with bioactive agents, such as antioxidants, to provide further degradation resistance (Ernsting et al., 2002). All of the “barrier” strategies to protecting polyurethanes described above appear to have validity, at least for protecting polyurethanes in the short term. The long-term (multiyear) benefits of these approaches remain to be seen in light of issues such as surface dynamics, interfacial interactions, and coating durability. All of these polyurethane modifications, while potentially providing enhanced resistance to biodegradation, still allow susceptibility to attack by biological components, often at a slow rate. 

Oxidative Degradation Induced by External Environment Under very limited circumstances, the body can transmit electromagnetic radiation that may affect the integrity of implanted polymers. For example, the cornea and vitreous humor of the eye, as well as superficial skin layers, allow the passage of long-wave (320–400 nm) “ultraviolet A” radiation. Absorption of ultraviolet radiation causes electron excitation that can lead to photooxidative degradation. This process has been suggested in the breakdown of polypropylene components of intraocular lenses (Altman et al., 1986; Jongebloed and Worst, 1986). In maxillofacial exo- and very likely endoprostheses, elastomers may undergo undesirable changes in color and physical properties as a consequence of exposure to natural sunlight-frequency radiation (Craig et al., 1980). Photooxidation mechanisms involving the aromatic units and the urethane of aromatic poly(ether urethanes) or poly(ester urethanes) are shown in Fig. 2.4.2.24. Antioxidants and ultraviolet absorbers provide limited protection for these materials. 

Emerging Long-Term Elastomer Applications Polyurethanes Three polyurethane families are now elaborated, the first for its broad-based popularity as an alternative to poly(ether

936 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

in the body buffer, with poly(carbonate urethane) integrity also susceptible to a combination of mechanical stress and vigorous oxidizing conditions (Faré et  al., 1999; Labow et al., 2002b). A 3-year implantation study indicated detectable degradation (Seifalian et  al., 2003). Only long-term implant studies (e.g., 5 years or greater) would ascertain the acceptability of poly(carbonate urethanes) or, for that matter, other new polymers having potentially susceptible groups. Unfortunately, financial constraints dictate that these extended studies would only be feasible as postmarket surveillance. One of the most promising polyurethane families for use as tough, biostable polyurethane elastomers is in development by Boston Scientific Corporation, Formerly Cardiac Pacemakers, Inc., now Abbott Corporation (Desai et  al., 2014, 2015). This polyurethane family employs poly(isobutylene diol) as a major soft segment component along with a lesser, reinforcing amount of poly(ether diol) soft segment and conventional methylene dianiline diisocyanate/short chain diol hard segment. Studies predictive of high biostability have been executed (Cozzens et al., 2010). 

Hydrocarbon Elastomers • Figure 2.4.2.24  Photooxidative reactions of aromatic polyurethanes.

(A) Formation of quinone-imide from aromatic polyurethane. (B) Photolytic cleavage of urethane link. (From Coury, A.J., Slaikeu, P.C., Cahalan, P.T., Stokes, K.B., Hobot, C.M., 1988. Factors and interactions affecting the performance of polyurethane elastomers in medical devices. J. Biomater. Appl. 3, 130–179. (B) From Brauman, S.K., Mayorga, G.D., Heller, J., 1981. Light stability and discoloration of segmented polyether urethanes. Ann. Biomed. Eng. 9, 45–58.)

urethanes), sometimes perhaps, unwarranted. The second is one that uses a large proportion of hydrocarbon soft segment with a lesser amount of polyether soft segment. The third elastomer family comprises polymers and copolymers that are completely hydrocarbon-based. Poly(carbonate urethanes) are specially elaborated in this section because their in vivo use has both oxidative and hydrolytic implications. Their mechanical properties rival the best poly(ether urethanes) and are favorite compositions for short-term implantable catheters with consideration for chronic use (Wright, 2006). In some studies, they have shown superior oxidation resistance to poly(ether urethanes) in several studies (Mathur et al., 1997; Tanzi et al., 1997). In other studies, however, these compositions were shown to be susceptible to oxidation (Christenson et  al., 2004). Additionally, in aqueous media in vitro and in vivo, slow degradation attributable to simple hydrolysis was also detected (Zhang et al., 1997). The body fluid environment provides a relatively stable long-term hydrolytic medium, generally less subject to cellular “respiratory bursts” that strongly enhance oxidative processes. Although phagocytic processes may also produce hydrolytic enzymes, their effects on synthetic polymers are specific and limited (Labow et al., 2002a). Simple hydrolysis, therefore, may be expected to take place continuously,

In 1987, Carl McMillin published a pioneering article on a set of elastomers intended for long-term flex fatigue resistance in medical devices such as artificial heart diaphragms (McMillin, 1987a,b). For simple flex life in an in  vitro setup, Hexsyn [cross-linked poly(1-hexene)] rubber stood out. It was used in early artificial heart designs under the name of “Bion” rubber. A communication from Dr. McMillin indicated that its lack of adequate tear strength inhibited its successful use, but it may be a very valuable material for dynamic use not subject to cuts and abrasions. A final very promising hydrocarbon rubber enjoying clinical use is the triblock copolymer, poly(styrene-isobutylenestyrene) (SIBS), spearheaded by (Pinchuk et al., 2007). It is reported to be stable in the body without degradation, long term, and causes a minimal host response. Its early use was as a drug carrier for the Boston Scientific TAXUS coronary artery stent. An advanced current clinical application is as a glaucoma drainage conduit (Pinchuk et al., 2015). The previous two uses were basically static implants. For dynamic applications, it is noted that SIBS is not as tough as polyurethanes and has not been reported for pacemaker lead insulation. In animal studies for heart valve leaflets, the leaflets failed, from “material failure and calcification” (Wang et al., 2010). However, subsequent studies of these copolymers, in cross-linked form to enhance toughness, have shown indications of greater durability while maintaining favorable host response for synthetic heart valves (Sheriff et al., 2015). 

Conclusions Polymers that are carefully chosen for use in implanted devices generally serve effectively for their intended lifetimes if they are properly selected, processed, and

CHAPTER 2.4.2   Chemical and Biochemical Degradation of Polymers Intended to Be Biostable

device–material–host interactions are adequately addressed. In certain limited circumstances, excessive hydrolytic or oxidative biodegradation occurs. This may be induced by direct attack by the host or by nonbiological factors in the environment surrounding the implant. With susceptible polymers, protective measures can be taken to ensure extended efficacy, although new, biodegradation-resistant polymers that are on the horizon may require less protection. Knowledge of biodegradation mechanisms and the employment of appropriate countermeasures such as proper material selection or modification, optimal component design, protection from environmental attack, and careful processing, handling, and implantation procedures will promote continued progress in the development of polymers as longterm implantable biomaterials.

Acknowledgments The author is very grateful to Dr. R.C. Allen for providing the drawing on activated phagocyte redox metabolism. For their technical advice and contributions, I sincerely thank James Anderson, Ken Stokes, Jonathan Sears, John Eaton, Allan Hoffman, John Mahoney, Maurice Kreevoy, Grace Picciolo, Buddy Ratner, Bob Ward, SuPing Lyu, Darrel Untereker, and Len Pinchuk. For the preparation of the original manuscript, I am deeply indebted to my “computer wizard,” Mrs. Jayne McCaughey. For help in updating original literature sources, I thank Ms. Mari Ferentinos.

References Abbott, 2019. Product Performance Report, first ed., pp. 162–194. Cardiac Arrythmia and Heart Failure. Allen, R.C., 1991. Activation of Phagocyte Redox Metabolism: Chemiluminigenic Probing with Luminol and Lucigenin. Drawing provided. Altman, J.J., Gorn, R.A., Craft, J., Albert, D.M., 1986. The breakdown of polypropylene in the human eye: is it clinically significant? Ann. Ophthalmol. 18, 182–185. ASTM D1708 – 10, 2010. Standard Test Method for Tensile Properties of Plastics by Use of Microtensile Specimens. http://www.ast m.org/Standards/D1708.htm. Barbarini Ferraz, L.C., Schellini, S.A., Wludarski, S.L., Padovani, R., Selva, D., et al., 2009. Extraocular muscle fixation to porous polyethylene orbital implants using 2-octyl cyanoacrylate. Eur. J. Ophthalmol. 19 (4), 527–529. Bernardino, C.R., Mihora, L.D., Fav, A.M., Rubin, P.A., 2006. Orbital complications of hydrogel scleral buckles. Ophthalmic Plast. Reconstr. Surg. 22 (3), 206–208. Blais, P., 1990. Letter to the editor. J. Appl. Biomater. 1, 197. Blanchet, T.A., Burroughs, B.R., 2001. Numerical oxidation model for gamma radiation-sterilized UHMWPE: consideration of dosedepth profile. J. Biomed. Mater. Res. 58 (6), 684–693. Bloch, B., Hastings, G., 1972. Plastics Materials in Surgery, second ed. Charles C. Thomas, Springfield, IL, pp. 97–98. Booth, A.E., February 1995. Industrial sterilization technologies: new and old trends shape manufacturer choices. Med. Dev. Diagn. Ind. 64–72.

937

Bouvier, M., Chawla, A.S., Hinberg, L., 1991. In vitro of a poly(ether urethane) by trypsin. J. Biomed. Mater. Res. 25, 773–789. Brauman, S.K., Mayorga, G.D., Heller, J., 1981. Light stability and discoloration of segmented polyether urethanes. Ann. Biomed. Eng. 9, 45–58. Braunstein, R.A., Winnick, M., 2002. Complications of Mira-gel: pseudotumor. Arch. Ophthalmol. 120, 228–229. Cardia, G., Regina, G., 1989. Degenerative Dacron graft changes: is there a biological component in this textile defect? A case report. Vasc. Surg. 23 (3), 245–247. Carr, J., 2011. The intracorporeal use of 2-octyl cyanoacrylate resin to control air leaks after lung resection. Eur. J. Cardiothorac. Surg. 39 (4), 579–583. Casas, J., Donovan, M., Schroeder, P., Stokes, K., Untereker, D., 1999. In vitro modulation of macrophage phenotype and inhibition of polymer degradation by dexamethasone in a human macrophage/Fe/stress system. J. Biomed. Mater. Res. 46, 475–484. Cauich-Rodríguez, J.V., Chan-Chan, L.H., Hernandez-Sánchez, F., Cervantes-Uc, J.M., March 27, 2013. Degradation of polyurethanes for cardiovascular applications. In: Pignatello, R. (Ed.), Advances in Biomaterials Science and Biomedical Applications. IntechOpen. Chaffin, K.A., Buckalew, A.J., Schley, J.L., Chen, X., et  al., 2012. Influence of water on the structure and properties of PDMScontaining multiblock polyurethanes. Macromolecules 45 (22), 9110–9120. Chen, L.E., Seaber, A.V., Urbaniak, 1993. Comparison of 10-0 polypropylene with 10-0 nylon sutures in rat arterial anastomoses. Microsurgery 14 (5), 328–333. Christenson, E.M., Anderson, J.M., Hiltner, A., 2004. Oxidative mechanisms of poly(carbonate urethane) and poly(ether urethane) biodegradation: In vivo and in vitro correlations. J. Biomed. Mater. Res. Part A. 70, 245–255. Cipriani, E.,P., Bracco, P., Kurtz, S.M., Costa, L., Zanetti, M., 2013. In-vivo degradation of poly(carbonate-urethane) based spine implants. Polym. Degrad. Stab. 98 (6), 1225–1235. Colthurst, M.J., Williams, R.L., Hiscott, P.S., Grierson, I., 2000. Biomaterials used in the posterior segment of the eye. Biomaterials 21, 649–665. Coury, A.J., Slaikeu, P.C., Cahalan, P.T., Stokes, K.B., 1987. Medical applications of implantable polyurethanes: current issues. Prog. Rubber Plast. Technol. 3 (4), 24–37. Coury, A.J., Slaikeu, P.C., Cahalan, P.T., Stokes, K.B., Hobot, C.M., 1988. Factors and interactions affecting the performance of polyurethane elastomers in medical devices. J. Biomater. Appl. 3, 130–179. Coury, A.J., Hobot, C.M., Slaikeu, P.C., Stokes, K.B., Cahalan, P.T., 1990. A new family of implantable biostable polyurethanes. In: Trans. 16th Annual Meeting Soc. for Biomater., May 20–23, p. 158. Coury, A.J., 1999. Biostable polymers as durable scaffolds for tissue engineered vascular prostheses. In: Zilla, P., Greisler, H. (Eds.), Tissue Engineering of Vascular Prosthetic Grafts, vol. 43. R.G. Landes Company, Austin, TX, pp. 469–480. Cozzens, D., Ojha, U., Kulkarni, P., Faust, R., Desai, S., 2010. Long term in vitro biostability of segmented polyisobutylene-based segmented, thermoplastic polyurethanes. J. Biomed. Mater. Res. 95 (3), 774–782. Craig, R.G., Koran, A., Yus, R., April 1980. Elastomers for maxillofacial applications. Biomaterials 1, 112–117. Crowninshield, R., Muratoglu, O., 2008. How have new sterilization techniques and new forms of polyethylene influenced wear in total joint replacement? J. Am. Acad. Orthop. Surg. 16 (Suppl. 1), S80–S85. Daly, B.M., Yin, J., 1998. Subsurface oxidation of polyethylene. J. Biomed. Mater. Res. 42, 523–529.

938 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Desai, S., Boden, M., January 6, 2015. U.S. Patent 8,927,660, Crosslinkable Polyisobutylene-Based Polymers and Medical Devices Containing the Same. Desai, S., Boden, M., DeRoche, S., Foster, A., Reddy, S., March 3, 2011. US Patent Application 20110054580. Polyisobutylene Urethane Urea and Urethane/Urea Copolymers and Medical Leads Containing the Same. Desai, S., Boden, M., DeRoche, S., et al., December 2, 2014. U.S. Patent 8,903,507, Polyisobutylene Urethane, Urea and Urethane/ Urea Copolymers and Medical Devices Containing the Same. Duke Extrusion.com (Copyright, 2019). https://www.dukeextrusion. com/materials/nylon-medical-extrusion. Ernsting, M.J., Santerre, J.P., Labow, R.S., 2002. Surface modification of a polycarbonate-urethane using a Vitamin E derivatized fluoroalkyl surface modifier. In: Trans. 28th Annual Meeting Soc. Biomater., April 24–27, p. 16. Faré, S., Petrini, P., Motta, A., Cigada, A., Tanzi, M.C., 1999. Synergistic efforts of oxidative environments and mechanical stress on in  vitro stability of polyetherurethanes and polycarbonateurethanes. J. Biomed. Mater. Res. 45, 62–74. Finck, K.M., Grosse-Siestrup, C., Bisson, S., Rinck, M., Gross, U., 1994. Experimental in  vivo degradation of polyarylamide. In: Trans. 20th Annual Meeting Soc. for Biomater., April 5–9, p. 210. Furman, B., Li, S., 1995. The effect of long-term shelf life aging of ultra high molecular weight polyethylene. In: Trans. 21st Annual Meeting Soc. for Biomater., March 18–22, p. 114. Graiver, D., Farminer, K.W., Narayan, R., 2003. A review of the fate and effects of silicones in the environment. J. Polym. Environ. 11, 129–136. Griessbach, E.F., Lehmann, R.G., 1999. Degradation of polydimethylsiloxane fluids in the environment–a review. Chemosphere 38 (6), 1461–1468. Greisser, H.J., Gengenbach, T.R., Chatelier, R.C., 1994. Longterm changes in the surface composition of polymers intended for biomedical applications. Trans. 20th Annual Meeting Soc. for Biomater. 19, 5–9. Gumargalieva, K.Z., Moiseev, Y.V., Daurova, T.T., Voronkova, O.S., 1982. Effect of infections on the degradation of polyethylene terephthalate implants. Biomaterials 3 (3), 177–180. Hagmar, L., Welinder, H., Mikoczy, Z., 1993. Cancer incidence and mortality in the Swedish polyurethane foam manufacturing industry. Br. J. Ind. Med. 50, 537–543. Hauser, R.G., Abdelhadi, R.H., McGriff, D.M., Kallinen Retel, L., 2013. Failure of a novel silicone-polyurethane copolymer (Optim™) to prevent implantable cardioverter-defibrillator lead insulation abrasions. Europace 15 (2), 278–283. Hegyeli, A., 1973. Use of organ cultures to evaluate biodegradation of polymer implant materials. J. Biomed. Mater. Res. 7, 205–214. Hernandez, R., Weksler, J., Padsalgikar, A., Runt, J., 2007. In vitro oxidation of high polydimethylsiloxane content bio-medical polyurethanes: correlation with the microstructure. J. Biomed. Mater. Res. 87 (2), 546–556. Hoffman, A., 1999. Personal Communication. Hudson, J., Crugnola, A., 1987. The in vivo biodegradation of nylon 6 utilized in a particular IUD. J. Biomater. Appl. 1, 487–501. Hwang, K.I., Lim, J.I., 1997. Hydrogel explant fragmentation 10 years after scleral buckling surgery. Arch. Ophthalmol. 115, 1205–1206. Jongebloed, W.L., Worst, J.F.G., 1986. Degradation of polypropylene in the human eye: a SEM-study. Doc. Ophthalmol. 64, 143–152. Kang, J., Erdodi, G., Brendel, C.M., Ely, D., Kennedy, J.P., 2010. Polyisobutylene-based polyurethanes. V. Oxidative- hydrolytic stability and biocompatibility. J. Polym. Sci. A 48 (10), 2194–2203.

Kato, Y.P., Dereume, J.P., Kontges, H., Frid, N., Martin, J.B., et al., 1995. Preliminary mechanical evaluation of a novel endoluminal graft. In: Trans. 21st Annual Meeting Soc. for Biomater., March 18–22, p. 81. Kearney, J.J., Lahey, J.M., Borirakchanyavat, S., Schwartz, D.M., Wilson, D., et al., 2004. Complications of hydrogel explants used in scleral buckling surgery. Am. J. Ophthalmol. 137 (1), 96–100. Klebanoff, S., 1982. Iodination catalyzed by the xanthine oxidase system: role of hydroxyl radicals. Biochemistry 21, 4110–4116. Kopecek, J., Ulbrich, K., 1983. Biodegradation of biomedical polymers. Prog. Polym. Sci. 9, 1–58. Kurtz, S.M., Rimnac, C.M., Hozack, W.J., Turner, J., Marcolongo, M., et al., 2005. In vivo degradation of polyethylene liners after gamma sterilization in air. J. Bone Jt. Surg. 87, 815–823. Labow, R.S., Erfle, D.J., Santerre, J.P., 1995. Neutrophil- mediated degradation of segmented polyurethanes. Biomaterials 16, 51–59. Labow, R.S., Tang, Y., McCloskey, C.B., Santerre, J.P., 2002a. The effect of oxidation on the enzyme-catalyzed hydrolytic degradation of polyurethanes. Can. J. Biomater. Sci. Polym. Ed. 13 (6), 651–665. Labow, R.S., Meek, E., Matherson, L.A., Santerre, J.P., 2002b. Human macrophage-mediated biodegradation of polyurethanes: assessment of candidate enzyme activities. Biomaterials 23 (19), 3969–3975. Lamba, K., Woodhouse, K.A., Cooper, S.L., Lelah, M.D., 1998. Polyurethanes in Biomedical Applications. CRC Press, Boca Raton. Liebert, T.C., Chartoff, R.P., Cosgrove, S.L., McCuskey, R.S., 1976. Subcutaneous implants of polypropylene filaments. J. Biomed. Mater. Res. 10 (6), 939–951. Ling, M.T.K., Westphal, S.P., Qin, S., Ding, S., Woo, L., 1998. Medical plastics failures from heterogeneous contamination. Med. Plast. Biomater. 5 (2), 45–49. Locksley, R., Wilson, C., Klebanoff, S., May 1982. Role of endogenous and acquired peroxidase in the toxoplasmacidal activity of murine and human mononuclear phagocytes. J. Clin. Investig. 69, 1099–1111. Lyu, S., Untereker, D., 2009. Degradability of polymers for implantable medical devices. Int. J. Mol. Sci. 10 (9), 4033–4065. Martin, D.J., Poole Warren, L.A., Gunatillake, P.A., McCarthy, S.J., Meijs, G.F., et  al., 2001. New methods for the assessment of in  vitro and in  vivo stress cracking in biomedical polyurethanes. Biomaterials 22 (9), 973–978. Mathur, A.B., Collier, T.O., Kao, W.J., Wiggins, M., Schubert, M.A., et al., 1997. In vivo biocompatibility and biostability of modified polyurethanes. J. Biomed. Mater. Res. 36, 246–257. McKellop, H., Yeom, B., Campbell, P., Salovey, R., 1995. Radiation induced oxidation of machined or molded UHMWPE after seventeen years. In: Trans. 21st Annual Meeting Soc. Biomater., March 18–22, p. 54. McMillin, C., 1987a. Development of tests to evaluate candidate elastomers for artificial heart diaphragms. Artif. Organs 11 (5), 395–404. McMillin, C., 1987b. Characterization of Hexsyn, a polyolefin rubber. J. Biomater. Appl. 2 (1), 3–100. Medel, F.J., Kurtz, S.M., Hozack, W.J., Parvizi, J., Purtill, J.J., et al., 2009. Gamma inert sterilization: a solution to polyethylene oxidation? J. Bone Jt. Surg. 91, 839–849. Northup, S., 1987. Strategies for biological testing of biomaterials. J. Biomater. Appl. 2, 132–147. Oppenheimer, E.T., Willhite, M., Danishefsky, I., Stout, A.P., January 1961. Observations on the effects of powdered polymer in the carcinogenic process. Cancer Res. 21, 132–134.

CHAPTER 2.4.2   Chemical and Biochemical Degradation of Polymers Intended to Be Biostable

Padsalgikar, A., Gallagher, G., Cosgriff-Hernandez, Runt, J., 2015. Polyurethanes in cardiac device leads – effect of morphology on performance. PU Mag. 12 (2). Parmar, A., January 24, 2012. Who makes the secret sauce in St. Jude's Optim technology? MedCity News. https://medcitynews. com/2012/01/who-makes-the-secret-sauce-in-st-judes-highlytouted-optim-technology-another-mn-company/. Phua, S.K., Castillo, E., Anderson, J.M., Hiltner, A., 1987. Biodegradation of a polyurethane in  vitro. J. Biomed. Mater. Res. 21, 231–246. Pinchuk, L., Esquivel, M.C., Martin, J.B., Wilson, G.J., 1991. Corethane: A new replacement for polyether urethanes for longterm implant applications. In: Trans. 17th Annual Meeting Soc. Biomater., May 1–5, p. 98. Pinchuk, L., Wilson, G., Barry, J., Schoephoerster, R., Parel, J.-M., Kennedy, J., 2007. Medical applications of poly(styrene-blockisobutylene-block-styrene) (“SIBS”). Biomaterials 2–13 Accepted, September. Pinchuk, L., Riss, I., Batlle, J.F., et al., August 2015. The development of a micro-shunt made from poly(styrene-block-isobutylene-blockstyrene) to treat glaucoma. J. Biomed. Mater. Res. Part B 1–11. Pinchuk, L., September 15, 1992. Adhesiveless Bonding of Silicone Rubber to Polyurethanes and the Use of Bonded Materials. US Patent, 5. 147, 725. Pinchuk, L., 1994. A review of the biostability and carcinogenicity of polyurethanes in medicine and the new generation of 'biostable' polyurethanes. J. Biomater. Sci. Polym. Ed. 6 (3), 225–267. Pitt, C.G., 1992. Non-microbial degradation of polyesters: mechanisms and modifications. In: Vert, M., Feijin, J., Albertson, A., Scott, G., Chiellini, E. (Eds.), Biodegradable Polymers and Plastics. R. Soc. Chem., Cambridge, UK, pp. 1–19. Portnoy, R., 1997. Clear, radiation-tolerant autoclavable polypropylene. Med. Plast. Biomater. 4 (1), 40–48. Ratner, B.D., Gladhill, K.W., Horbett, T.A., 1988. Analysis of in vitro enzymatic and oxidative degradation of polyurethanes. J. Biomed. Mater. Res. 22, 509–527. Refojo, M.F., Leong, F.L., 1981. Poly(methylacrylate-co- hydroxyethyl acrylate) hydrogel implant material of strength and softness. J. Biomed. Mater. Res. 15, 497–509. Rimnac, C., Pruitt, L., 2008. How do material properties influence wear and fracture mechanisms? Am. Acad. Orthop. Surg. 16 (Suppl. 1), S94–S100. Roldan-Pallares, M., del Castillo, J.L., Awad-El Susi, S., Refojo, M.F., 1999. Long-term complications of silicone and hydrogel explants in retinal reattachment surgery. Arch. Ophthalmol. 177, 197–201. Santerre, J.P., Labow, R.S., Duguay, D.G., Erfle, D., Adams, G.A., 1994. Biodegradation evaluation of polyether- and polyesterurethanes with oxidative and hydrolytic enzymes. J. Biomed. Mater. Res. 28, 1187–1199. Santerre, J.P., Meek, E., Tang, Y.W., Labow, R.S., 2000. Use of fluorinated surface modifying macromolecules to inhibit the degradation of polycarbonate-urethanes by human macrophages. In: Trans. 6th World Biomaterials Congress, p. 77. Schnabel, W., 1981. Polymer Degradation Principles and Practical Applications. Macmillan, New York, NY. pp. 15–17, 179–185. Schoen, F.J., 1987. Biomaterial-associated infection, neoplasia and calcification. Clinicopathologic features and pathophysiologic concepts. Trans. Am. Soc. Artif. Intern. Organs 33 (1), 8–18. Schubert, M.A., Wiggins, M.J., Anderson, J.M., Hiltner, A., 1997a. Comparison of two antioxidants for poly(etherurethane urea) in an accelerated in vitro biodegradation system. J. Biomed. Mater. Res. 34, 493–505.

939

Schubert, M.A., Wiggins, M.J., Anderson, J.M., Hiltner, A., 1997b. Role of oxygen in biodegradation of poly(etherurethane urea) elastomers. J. Biomed. Mater. Res. 34, 519–530. Seifalian, A.M., Salacinski, H.J., Tiwari, A., Edwards, E., Bowald, S., et al., 2003. In vivo biostability of a poly(carbonate urea) urethane graft. Biomaterials 24 (14), 2549–2557. Shen, F.W., Yu, Y.J., McKellop, H., 1999. Potential errors in FTIR measurement of oxidation in ultra-high molecular weight polyethylene implants. J. Biomed. Mater. Res. 48, 203–210. Sheriff, J., Claiborne, T.E., Tran, P.T., Pinchuk, M., Slepian, M.J., Bluestein, D., et al., 2015. Physical characterization and platelet interactions under shear flows of a novel thermoset polyisobutylene-based copolymer. ACS Appl. Mater. Interfaces 7 (39), 22058–22066. Simmons, A., Hyvarinen, J., O'Dell, R., Martin, D., et  al., 2004. Long-term in  vivo biostability of poly(dimethylsiloxane)/ poly(hexamethylene oxide) mixed macrodiol-based polyurethane elastomers. Biomaterials 25, 4887–4900. Singer, A.J., Quinn, J.V., Hollander, J.E., 2008. The cyanoacrylate topical skin adhesives. Am. J. Emerg. Med. 26 (4), 490–496. Smith, R., Oliver, C., Williams, D.F., 1987. The enzymatic degradation of polymers in vitro. J. Biomed. Mater. Res. 21, 991–1003. Snow, J., Harasaki, H., Kasick, J., Whalen, R., Kiraly, R., et  al., 1981. Promising results with a new textured surface intrathoracic variable volume device for LVAS. Trans. Am. Soc. Artif. Intern. Organs XXVII, 485–489. St. Jude Medical, Inc., 2012. Optim® Insulation, A New Material for a New Generation of Cardiac Leads. Product Information. http://www.sjmprofessional.com/Products/US/CRT-Systems/ Optim-Insulation.aspx. Stokes, K., Coury, A., Urbanski, P., April 1987. Autooxidative degradation of implanted polyether polyurethane devices. J. Biomater. Appl. 1, 412–448. Stokes, K., Urbanski, P., Upton, J., 1990. The in vivo autooxidation of polyether polyurethane by metal ions. J. Biomater. Sci. Polym. Ed. 1 (3), 207–230. Stokes, K., October 1988. Polyether polyurethanes: biostable or not? J. Biomater. Appl. 3, 228–259. Sutherland, K., Mahoney, J.R., II, Coury, A.J., Eaton, J.W., 1993. Degradation of biomaterials by phagocyte-derived oxidants. J. Clin. Investig. 92, 2360–2367. Szycher, M., Siciliano, A., 1991. An assessment of 2,4-TDA formation from Surgitek polyurethane foam under stimulated physiological conditions. J. Biomater. Appl. 5, 323–336. Takahara, A., Coury, A.J., Hergenrother, R.W., Cooper, S.L., 1991. Effect of soft segment chemistry on the biostability of segmented polyurethanes. I. In vitro oxidation. J. Biomed. Mater. Res. 25, 341–356. Takahara, A., Coury, A.J., Cooper, S.L., 1994. Molecular design of biologically stable polyurethanes. In: Trans. 20th Annual Meeting Soc. Biomater., April 5–9, p. 44. Tang, W.W., Santerre, J.P., Labow, R.S., Waghray, G., Taylor, D., 1994. The use of surface modifying macromolecules to inhibit biodegradation of segmented polyurethanes. In: Trans. 20th Annual Meeting Soc. Biomater., April 5–9, p. 62. Tanzi, M.C., Mantovani, D., Petrini, P., Guidoin, R., Laroche, G., 1997. Chemical stability of polyether urethanes versus polycarbonate urethanes. J. Biomed. Mater. Res. 36, 550–559. Test, S., Weiss, S., 1986. The generation of utilization of chlorinated oxidants by human neutrophils. Adv. Free Radical Biol. Med. 2, 91–116. Thomas, C., August 23, 2012. ICD Warning: Defective Defibrillator Leads Recalled, Heart Sisters for Women with Heart Disease. http s://myheartsisters.org/2012/08/23/icd-warning-defective-defibrillator-leads-recalled/.

940 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Tomita, N., Kitakura, T., Onmori, N., Ikada, Y., Aoyama, E., 1999. Prevention of fatigue cracks in ultra-high molecular weight polyethylene joint components by the addition of vitamin E. J. Biomed. Mater. Res. 48, 474–478. US FDA (Food and Drug Administration) Document: Medical Devices, 2010. Device Approvals: http://www.fda.gov/medicaldevices/productsandmedicalprocedures/deviceapprovalsandclearances/recently-approveddevices/ucm215106.htm2010. US FDA, February 6, 2019. Breast Implant Associated-Anaplastic Large Cell Lymphoma (BIA-ALCL) – Letter to Health Care Providers. https://www.fda.gov/medical-devices/lettershealth-care-providers/breast-implant-associated-anaplastic-largecell-lymphoma-bia-alcl-letter-health-care-providers. Vauthier, C., Dubernet, C., Fattal, E., Pinto-Alphandary, H., Couvreur, P., 2003. Poly (alkylcyanoacrylates) as biodegradable materials for biomedical applications. Adv. Drug Deliv. Rev. 55 (4), 519–548. Vinard, E., Eloy, R., Descotes, J., Brudon, J.R., Giudicelli, H., et al., 1991. Human vascular graft failure and frequency of infection. J. Biomed. Mater. Res. 25, 499–513. Wang, Q., McGoron, A.J., Bianco, R.W., Kato, Y., Pinchuk, L., Schoephoerster, R.T., 2010. In-vivo assessment of a novel polymer (SIBS) trileaflet heart valve. J. Heart Valve Dis. 19 (4), 499–505. Ward, R.S., White, K.A., Gill, R.S., Wolcott, C.A., 1995. Development of biostable thermoplastic polyurethanes with oligomeric polydimethylsiloxane end groups. In: Trans. 21st Annual Meeting Soc. Biomater., March 18–22, p. 268. Ward, R.S., Tian, Y., White, K.A., 1998. Improved polymer biostability via oligomeric end groups incorporated during synthesis. Polym. Mater. Sci. Eng. 79, 526–527. Ward, R.S., 2008. New horizons for biomedical polymers. Med. Des. Tech. 19 (5), 26–28 30–31. Wasserbauer, R., Beranova, M., Vancurova, D., Dolezel, B., January 1990. Biodegradation of polyethylene foils by bacterial and liver homogenates. Biomaterials 11, 36–40. Watt, D.R., 2001. Miragel Sponge Complications. www.retinadoc.com/scripts/retina.pl?function=viewquestions&;forum= retina, 11/07/012001. Weaver, K.D., Sauer, W.L., Beals, N.B., 1995. Sterilization induced effects on UHMWPE oxidation and fatigue strength. In: Trans. 21st Annual Meeting Soc. Biomater., March 18–22, p. 114.

Wiggins, M.J., Anderson, J.M., Hiltner, A., 2003. Biodegradation of polyurethane under fatigue loading. J. Biomed. Mater. Res. Part A 65A (4), 524–535. Williams, D.F., 1982. Review: biodegradation of surgical polymers. J. Mater. Sci. 17, 1233–1237. Williams, D.F., 1989. Definitions in Biomaterials. Amsterdam Elsevier. Wright, J.I., March 1, 2006. Using Polyurethanes in Medical Applications, Medical Device and Diagnostics Industry. QMed. https://www.mddionline.com/using-polyurethanes-medical-applications. Wu, L., Weisberg, D.M., Runt, J., Felder III, G., Snyder, A.J., et al., 1999. An investigation of the in vivo stability of poly(ether urethane urea) blood sacs. J. Biomed. Mater. Res. 44, 371–380. Zaikov, G.E., 1985. Quantitative aspects of polymer degradation in the living body. JMS Rev. Macromol. Chem. Phys. C25 (4), 551–597. Zhang, Z., Marois, Y., Guidoin, R., Bull, P., Marois, M., et al., 1997. Vascugraft® polyurethane arterial prosthesis as femoro-popliteal and femoro-peroneal bypass in humans: pathological, structural and chemical analyses of four excised grafts. Biomaterials 18, 113–124. Zhao, Q., Agger, M., Fitzpatrick, M., Anderson, J., Hiltner, A., et al., 1990. Cellular interactions with biomaterials: In  vivo cracking of pre-stressed pellethane 2363-80A. J. Biomed. Mater. Res. 24, 621–637. Zhao, Q., Topham, N., Anderson, J.M., Hiltner, A., Lodoen, G., et al., 1991. Foreign-body giant cells and polyurethane biostability: In  vivo correlation of cell adhesion and surface cracking. J. Biomed. Mater. Res. 25, 177–183. Zhao, Q.H., McNally, A.K., Rubin, K.R., Renier, M., Wu, Z., et al., 1993. Human plasma α2-macroglobulin promotes in vitro oxidative stress cracking of Pellethane 2363-80A. Biomed. Mater. Res. 27, 379–389. Zhao, Q., Casas-Bejar, C., Urbanski, P., Stokes, K., 1995. Glass wool-H2O2/COCl2 for in vitro evaluation of biodegradative stress cracking in polyurethane elastomers. J. Biomed. Mater. Res. 29, 467–475. Ziats, N., Miller, K., Anderson, J., January 1988. In vitro and in vivo interactions of cells with Biomaterials. Biomaterials 9, 5–13.

Chapter Questions •  Hydrophobic polymer biomaterials designed both to degrade and to be stable in vivo may produce sub-micron particles during use. What are issues with the host responses that are common to the degradation of both types of polymers? • Consider the degradation of materials commonly used in medicine that do not have well-defined breakdown mechanisms. Some examples include poly(ethylene glycol), hydroxyapatite, and some polysaccharides. How does the body deal with these common materials? • A new class of biomaterials is now under development that degrades on cue. The cue might be thermal, photonic, or enzymatic. Ingenious chemical design principles are being applied to create such materials, but how might the body react to the products generated by a sudden breakdown of the structure? • Learn about new strategies to stabilize materials against degradation, for example, vitamin E loading of orthopedic polymers, and incorporation of poly-isobutylene segments into elastomers.

• Endovascular stents are among the most widely used of all medical devices (Chapter 2.5.2B). A new generation of biodegradable stents is expected to have a huge impact on cardiovascular therapies. Consider how biodegradable poly(lactic acid) or magnesium or iron will perform in the complex intravascular environment. • For a medical device intended for years of service, especially a device where failure can lead to death, how can we test and qualify the device for the expected period of service? Are there useful in vitro tests? Are there relevant and justified animal models? • Henry Petroski and other authors have discussed the important role of failure in advancing engineering design. Consider medical device failure, past and present, associated with degradation, and how these unintended complications will lead to better medical devices. A few examples include the degradation of polyurethane pacemaker leads, the breakdown of a protective sheath on the tailstring of the Dalkon Shield IUD, and the wear debris associated with the oxidation of ultrahigh-molecularweight polyethylene in hip prostheses.

940.e1

2.4.3

Metallic Degradation and the Biological Environment JEREMY L. GILBERT Department of Bioengineering, Clemson University, Charleston, SC, United States

Introduction Metallic biomaterials experience degradation in the biological environment with a rate, form, and extent of damage accumulation that depends on the alloy, its surface structure and properties, and the mechanical, chemical, and biological processes ongoing adjacent to the alloy surface. Some alloys (e.g., magnesium alloys) degrade rapidly and are under consideration as “biodegradable” alloys where the implant is designed to be temporary and is removed by electrochemical processes over time, while other alloys (e.g., stainless steel, titanium, and cobalt–chromium– molybdenum alloys) are not intended to degrade and are meant to survive for the life of the patient. Such distinctions are important to understand and consider when making medical devices out of medical alloys. Alloys intended for degradation, including magnesium, zinc, and others, are rapidly corroding metals, while those intended for lifetime performance in the human body are highly corrosion resistant. The materials factors dictating such differences in performance include the basic concepts of electrochemistry where both the thermodynamics of oxidation (i.e., the driving force for oxidation) and the kinetics of oxidation governed primarily by the presence of passive oxide films on the surface are important. Generally, metal alloys are used in the body to provide mechanical support and to resist fatigue, wear, and corrosion processes in various medical device applications. In some circumstances, metals serve an electrical or electrochemical role (e.g., as an electrode). When considering degradation of metals in the body, the processes of most concern typically relate to tribology (wear processes), corrosion, and fatigue (Gilbert, 2017). This chapter will focus on the tribology– corrosion–biology interactions that may arise in a range of medical device applications where metals are used. The biological system affects the potential forms of metal alloy degradation (tribology and corrosion) and is

variable with location (skeletal tissue, cardiovascular, etc.). The effects of degradation on the biological system and the effects of the biological system on degradation are both important and central to the complex interplay that often results in failure of a medical device. The basic mechanisms of degradation of metal surfaces will be discussed and the conjoint interaction between processes, and the interplay between the biological system and the degrading alloy will be discussed in this chapter. 

The Severe Biological Environment (Fatigue, Tribology, Corrosion, and Biology) Metallic biomaterials are used in a wide range of applications where high cyclic stresses, tribological interactions, and exposure to body solutions, biological molecules, tissues, and cells are all present. This combination of factors is most easily recognized in total hip prostheses, where tribology, corrosion, and fatigue are experienced in a biological environment that can be inflamed and reactive. Typical hip implants used in the recent past (Fig. 2.4.3.1) have been comprised primarily of Ti–6Al–4V hip stems, CoCrMo necks and heads, CoCrMo acetabular liners, and Ti–6Al–4V acetabular shells. The combination of harsh body environment, comprised of local biological solutions, inflammation and inflammatory cells, and a range of proteins, enzymes, cytokines, etc., the large mechanical stresses due to activities of daily living, and the wear and corrosion mechanisms that arise can lead to large-scale, severe tribologically based and corrosion-based damage. In most cases these two combine to form a conjoint degradation mechanism known as mechanically assisted corrosion (MAC) and tribocorrosion (Gilbert et al., 1993; Swaminathan and Gilbert, 2012; Cao and Mischler, 2018; Mischler and Debaud, 1998). 941

942 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Each aspect of this system (i.e., body environment, mechanical loading, tribology) can result in failure of the metal but it is more often the case that conjoint mechanisms arise leading to failure. That is, wear and corrosion and biological interactions often combine to result in more

complex and severe failure modes than any one process might otherwise cause. Understanding such system-based failure modes is critical to understanding metallic biomaterials-based failures. 

Basic Corrosion of Passive Oxide-Covered Alloys

• Figure 2.4.3.1  Several examples of metallic total hip prostheses that have been used in the recent past. These include implants with modular head/neck tapers, neck/stem tapers, modular body tapers, and acetabular tapers. The tapered metal–metal junctions in each of these locations have been known to be susceptible to mechanically assisted corrosion (Goldberg et al., 2002).

It is well known that all metallic biomaterials exhibit some form of corrosion process, whether passive dissolution of ions through the passive film, or more aggressive corrosion damage that is linked to wear or crevice geometries or both. The basic corrosion interactions observed (Fig. 2.4.3.2) in vivo consist of oxidation and reduction reactions, where the oxidation processes take metal (zero valence) and increase its valence to make ions (cations) or oxides (or other solid oxidation products). These oxidation reactions occur at the passive oxide film surface, which is typically only a few nanometers thick. One of the potential reduction processes thought to be present takes water and oxygen and makes hydroxide ions and/or reactive oxygen intermediates (Fig. 2.4.3.2). However, there are many other biologically based molecules that are susceptible to redox processes and very little is known about these reactions, how the presence of oxidizing metals affects these species, or how biochemical species may alter oxidation of the metal or alteration of the oxide film. Little is known about such specific biological species and their role in corrosion of metallic biomaterials. The oxidation and reduction reactions are electrically connected through the metal and complete the circuit through

• Figure 2.4.3.2  General summary of oxidation and reduction reactions that comprise corrosion of metallic biomaterials. Oxidation reactions, typically through the oxide film-covered surface, can form ions (in solution) or solid products like oxides or phosphates. Reduction reactions take oxygen and water in solution and make hydroxide ions and other reactive oxygen intermediates (e.g., reactive oxygen species [ROS]). (Used with permission from Gilbert, J.L., Kubacki, G.W., 2015. Oxidative stress, inflammation and the corrosion of metallic biomaterials: corrosion causes biology and biology causes corrosion. In: Dziubla, T.D., Butterfield, D.A. (Eds.), Oxidative Stress and Biomaterials. Elsevier Press (Chapter 3).)

CHAPTER 2.4.3   Metallic Degradation and the Biological Environment

the solution, resulting in currents (electronic and ionic) through both phases. Corrosion, in and of itself, does not define failure of the metallic implant. Failure, clinically, is more appropriately defined as the need to remove the device (metal) due to a clinically defined failure mode (e.g., pain, loosening, infection, adverse biological reaction, etc.). Thus while corrosion may be associated with failure, it is likely that other biologically based processes (e.g., adverse local tissue reactions) lead to the clinical conditions for revision. It is often difficult to determine the causal relationships between materials-based damage modes and biological processes leading to inflammation, pain, and ultimately revision. However, corrosion is a nearly universal observation in metallic biomaterials (i.e., all metals corrode to some extent). It is often the case that these metals, even though they are corroding, can perform their function and not lead to clinically significant failure modes. Indeed, biocompatibility is not the absence of degradation and even with such degradation processes (which are inevitable), the biomaterials can perform their function with an appropriate host response, making them biocompatible. The biological environment adjacent to the metallic biomaterial may change with time from implantation, where highly inflammatory wound-healing mechanisms are at play, to chronic inflammatory conditions often associated with ongoing degradation processes. That is, there are feedback

943

systems that may develop between the immune system response and the degradation mechanisms that can accelerate degradation of the metal and/or amplify the immune reaction. Such positive feedback systems and conditions are currently under active investigation for metallic biomaterial interactions with the living system (FDA, 2019). 

Tribological Aspects of Metal-Hard Contact Degradation From a tribological perspective, the contact and movement of metallic biomaterial surfaces relative to other surfaces can lead to wear processes of the metal (oxide) surface that include adhesive, abrasive (second and third body), and fatigue wear mechanisms (discussed in Chapter 1.2.4 of this text). Examples of CoCrMo alloy surfaces from retrieved total hip metalon-metal implants, Fig. 2.4.3.3, show examples of abrasive, adhesive, and fatigue wear damage. When metals make contact with other metals or other hard contacting surfaces, the tribological interaction has a number of unique elements that require special consideration. This includes the fact that hard second-phase particles (e.g., oxides and carbides) may become free and serve as third body wear particles. These may roll or slide across the interface and imprint or scratch the surface (Fig. 2.4.3.3A). Adhesive wear interactions between CoCrMo and Ti–6Al–4V (Fig. 2.4.3.3B) are also possible where the

(A)

(B)

(C)

(D)

• Figure 2.4.3.3  (A) Evidence of abrasive wear damage on a CoCrMo head-bearing surface. (B) Adhesive

wear of Ti–6Al–4V debris (darker gray) onto CoCrMo femoral head (lighter gray). (C) and (D) Fatigue wear of CoCrMo metal-on-metal acetabular-bearing surfaces from retrieved total hip replacements. Note, evidence of abrasive wear (scratching) and of rolling hard particle imprinting in (A). In (B), the Ti alloy debris is a reacted oxide debris that came from entrainment of Ti alloy into the bearing surface from other components of the implant. In (C) and (D), fatigue wear arose from the high cyclic contact stresses near the edge of the acetabular component, the debris oxidized and partially refilled the region where the fatigue fracture occurred.

944 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

• Figure 2.4.3.4  Schematic of an asperity–asperity contact between two hard surfaces. While the nominal stresses that arise may fall within the elastic range, the stresses at the asperity contact points rapidly reach the hardness or surface yield strength of the material and plastically deform.

titanium alloy and its degradation products can adhere to the CoCrMo surface in an oxidized form. Fatigue wear results from high cyclic contact stresses inducing fatigue fracture of the metal in the subsurface region and loss of alloy particles from the surface (Fig. 2.4.3.3C,D). The remnant holes can be partially refilled with oxidized tribocorrosion debris. In all wear cases that breach the oxide film in the body, there is a corresponding oxidation reaction that accompanies the wear damage resulting in repassivation of the oxide. Thus wear of these metals in the body is always associated with corrosion processes. It is important to note that not all tribological interactions of a metal surface result in breaching of the oxide film. It is only when the metal surface is abraded such that plastic deformation of the surface occurs that oxide films are disrupted. Thus wear against materials softer than the metal will not typically induce oxide film disruption. However, if there are hard particles (as hard as or harder than the metal surface) embedded within softer (polymeric) materials, then such third body particles can induce oxide film disruption and repassivation reactions on the alloy surface. The three main nonbiodegradable alloy systems (titanium, cobalt–chromium–molybdenum, and stainless steel) are all passive oxide film-covered alloys. As such, contact of the metal surface involves these few nanometer-thick oxides and the immediately adjacent metallic grains contacting the opposing surface in an asperity–asperity contact condition. Even the smallest (nanometer or smaller) topographies give rise to localized asperities of the metal oxide surface, which will influence the contact conditions across the interface

(Fig. 2.4.3.4). Thus the stress developed at these asperities quickly rises to the level of the yield stress (or hardness) of the surface and can induce localized plastic deformation. While the overall stress distribution may be nominally elastic, these local asperities experience plastic deformation and may cause disruption of the oxide film on these surfaces due to the local plastic deformation at and near the contacting asperities. 

Metal-on-Metal (Hard) Surface Mechanics In many circumstances, metal surfaces within the body come into contact with other hard surfaces (metals, ceramics, etc.). These metal–hard contact regions have several unique mechanical aspects to them that need to be clearly understood in terms of the contacting conditions and the associated corrosion and tribocorrosion processes that may arise. First, it has been generally acknowledged that the contact mechanics of metal–hard surfaces can be, to first approximation, modeled in terms of Hertzian elastic contact where geometrically smooth surfaces are in contact (e.g., sphere on sphere, flat on sphere, etc.). For two elastic spheres of radius R1 and R2, in contact with a force, F, there is elastic deformation of both spheres to develop a circular contact area, A, with radius of contact, a, where a is a=

(

3FR 4E *

)1 / 3

(2.4.3.1)

CHAPTER 2.4.3   Metallic Degradation and the Biological Environment

where the effective radius, R, is 1 R

=

1 R1

+

1

(2.4.3.2)

R2

and the reduced modulus for the two materials is 1 E*

=

1 − ν21 E1

+

1 − ν22 E2

(2.4.3.3)

The average stress, σ = F/A, and the stress across the contact is distributed as (

σ = σ max 1 −

( )2 ) 1 / 2 r a

and σ max =

3F 2A

where A = πa2

(2.4.3.4) It is commonly thought that such analyses apply to metal–metal, or metal–hard contacts, and while the overall stress interaction is typically elastic when metals are in contact, there is an important role played by asperities (smallscale protruding features on the metal and other surfaces, see Fig. 2.4.3.4) that modify this conceptual framework. As shown in Fig. 2.4.3.4, when a hard sphere with asperities comes into contact with a metal surface also with small asperities present, the overall geometric deformation follows Hertzian contact; however, the local stresses at the peak asperities are more severe. At these asperities, the stresses quickly rise to the hardness (H) or yield stress of the metal surface and induce plastic deformation (dislocation motion) at and near the surface (Popov, 2010). When considering metal–hard surfaces, the true area of contact is dictated not by the Hertzian analysis, but rather by this asperity–asperity yielding process. This is shown schematically in Fig. 2.4.3.4B (bottom), where a smooth flat hypothetical surface is brought into contact with a random asperity-based surface. In this case, the highest asperities contact first and the contact stress in these asperities rises to the yield point, deforming the asperities plastically and allowing the next highest asperities to come into contact. This continues, with each asperity quickly reaching the yield stress of the surface, until the true area of contact is just the amount of asperity area needed to carry the applied load, F. That is, the true area of contact is given by A = F/H, where H is the hardness of the alloy that is yielding. Typically, the true area of contact of metal–hard surfaces is only a small fraction of the nominal (or even the Hertzian-based) contact area. The distribution of asperity heights in the surface can be defined by the cumulative probability (CP) function as the fraction of the area at a particular height, z, that is occupied by the material of the surface. This distribution varies from 0 to 1. The differential probability is simply the derivative of the CP and shows the height-based distribution of asperities. This differential probability function describes the area contact of two hard asperity–asperity contacting conditions (Fig. 2.4.3.4).

945

When tribological events occur (e.g., sliding of the two surfaces), it is the interaction of asperities that dominates the overall wear processes with the highest contacting regions experiencing damage and degradation. The oxide films on these interacting asperities will be disrupted by plastic deformation processes and the sliding will expose fresh metal to oxidation (releasing ions and repassivating the surface) and to generation of oxidized particles of debris. 

Clinically Observed Mechanically Assisted Crevice Corrosion (Fretting Crevice Corrosion) In Vivo One of the most studied and clearly identified degradation mechanisms associated with metallic biomaterials is a process known as mechanically assisted crevice corrosion (MACC, see Fig. 2.4.3.5 for a description of the complex multifactorial processes associated with this conjoint degradation mechanism). This term arose from the study of modular tapered junctions in total hip implants in the 1990s up to the present time. It is a degradation process that combines mechanical processes, particularly fretting, with corrosion processes and often involves crevice-like geometries where the local solution chemistry can be dramatically altered from the bulk environment by reactions at the surface. The biological system, its reaction to the alloy degradation, and its affect upon the degradation is also a critical aspect of this conjoint process. MACC was first identified in modular tapered total hip replacements where the head of the prosthesis is attached to the stem by way of a conical tapered section called a modular taper junction. The surgeon assembles these parts during implantation surgery by impacting the head on to the neck. The junction between these two parts results in a crevice geometry into which body fluids can penetrate and fretting within the junction causes oxide abrasion/disruption at the asperity–asperity contacts leading to a tribocorrosion process. Such fretting crevice corrosion processes are also observed in other orthopedic implants, fracture fixation devices, dental implants, and any other cases where metal– metal contact is possible. The asperity–asperity nature of metal–metal junctions, described earlier, results in only a few percent of the nominal taper contact area being in actual contact, and the remaining space results in a crevice geometry into which solution can penetrate. During loading, high cyclic stresses induce fretting motion, asperity– asperity sliding, and fretting crevice corrosion. Modular tapers often consist of different alloys coming into contact and concerns have been raised about the possibility of galvanic effects between, for example, Ti–6Al– 4V and CoCrMo. However, these concerns are overstated. Typically, galvanic interactions occur between two alloys with significantly different electrode potentials (open circuit potential, OCP), where the more active (negative) potential alloy will experience an increased rate of corrosion, while the more noble (positive) potential alloy will be protected

946 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

• Figure 2.4.3.5  Schematic description of the interdisciplinary nature of mechanically assisted corrosion of

metallic biomaterials. Knowledge across a range of disciplines is required to understand the in vivo interactions of metal surfaces, wear and corrosion-based degradation, and the biological environment. It includes materials science, bulk and surface mechanics, electrical and transport phenomena, and electrochemistry, all within a biological system that is both reactive to the degradation and contributing to the accelerated attack possible. ALTR, Adverse local tissue reaction; ROS, reactive oxygen species; H, hardness; E, Modulus; ef, fracture strain; ys, yield stress. (Used with permission from Gilbert, J.L., Sivan, S., Mali, S., 2015. Corrosion of modular tapers in total joint replacements: a critical assessment of design, materials, surfaces structure, mechanics, electrochemistry and biology. In: Greenwald, K., Lemons, M. (Eds.), ASTM Special Technical Publication on Implant Modularity, STP 1591. ASTM Int, 192–223.)

and will see a decreased rate of corrosion. Titanium and CoCrMo alloys, with passive oxide films intact, have very similar OCPs, thus little galvanic potential exists for increasing corrosion. In addition, if any difference in OCP exists, it is typically that Ti is more active and CoCrMo is more noble, thus it would be the titanium that would be accelerated in its corrosion. The passive oxide films also serve as a kinetic barrier to corrosion and limit galvanic effects (Jacobs et al., 1998). When modular tapers of Ti–6Al–4V and CoCrMo are examined, both the titanium and the CoCrMo are corroded and there is no evidence that either protects the other. Finally, MACC is observed in all possible alloy combinations and thus is not dependent on any galvanic effects. Therefore galvanic corrosion, while a mechanism that may play out in some circumstances (e.g., corrosion near carbides vs. bulk alloy, for example), does not contribute significantly to the MACC mechanism. Examples of the type of damage resulting from MACC are shown in Figs. 2.4.3.6–2.4.3.8, where fretting damage, intergranular corrosion, direct oxide conversion, and pitting of CoCrMo and/or Ti–6Al–4V are all observed. Fig. 2.4.3.6 demonstrates the basic fretting damage that results in small parallel abrasion patterns due to asperity–asperity contact and cyclic micromotion. Examples of more severe fretting corrosion damage in Fig. 2.4.3.6 show generation of oxide debris that accumulates within the taper region. Fig. 2.4.3.6A is an undamaged region of a CoCrMo head taper after retrieval. Fig. 2.4.3.6B–D are examples demonstrating

minor fretting damage resulting primarily from contact and motion. Fig. 2.4.3.7 shows an example of the more severe corrosion that can arise in CoCrMo alloy modular tapers resulting from MACC where fretting corrosion within the crevice led to intergranular corrosion that is affiliated with the imprinting of the male trunnion. Intergranular corrosion at the grain boundaries of this high-carbon CoCrMo alloy surface is evident. Examples of corrosion damage in Ti–6Al–4V alloy modular taper junctions, Fig. 2.4.3.8, show the severe damage possible in  vivo for these alloy systems. Fig. 2.4.3.8A is a scanning electron micrograph of a retrieved Ti–6Al–4V acetabular shell taper region (CoCrMo alloy was on the other side). Note the pitting and dissolution of the surface. Fig. 2.4.3.8B–D are scanning electron micrography (SEM) micrographs of sectioned regions through the taper junction of Ti–6Al–4V alloy modular body tapers (where Ti–6Al–4V was on both sides of the taper). These micrographs show the corrosion reaction leading to several hundred micron-thick oxide corrosion products (Fig. 2.4.3.8B), a sectioned region through a large (200 μm) pit in Ti–6Al–4V filled with oxide, and a section through a Ti–6Al–4V/Ti–6Al–4V alloy junction still intact (Fig. 2.4.3.8D), where about 10 μm of oxide formation has occurred at the taper junction. These images show that Ti alloys are susceptible to severe corrosion damage under MACC circumstances. The fretting corrosion aspects of MACC have been extensively

CHAPTER 2.4.3   Metallic Degradation and the Biological Environment

(A)

(B)

(C)

(D)

947



Figure 2.4.3.6 Evidence of fretting crevice corrosion in retrieved modular head/neck tapers of total hip replacements. (A) Undamaged machining ridges in a retrieved modular CoCrMo head taper. (B)–(D) Examples of minor fretting damage seen in retrieved CoCrMo head components.

• Figure 2.4.3.7  Retrieved CoCrMo head taper with extensive intergranular corrosion damage associated with mechanically assisted crevice corrosion. Corrosion damage around the grain boundaries and between the alloy and the second-phase particles is evident and associated with the fretting-based imprinting from the countersurface.

studied and reproduced in device and basic pin-on-disk tests. However, the severe corrosion damage of both CoCrMo and Ti–6Al–4V alloys seen in these retrievals (e.g., intergranular corrosion of CoCrMo, thick oxide formation, and pitting in Ti–6Al–4V) has not been reproduced in any in vitro laboratory tests to date that represent the in  vivo conditions, and the detailed conditions and mechanisms of generation of this type of damage are still under investigation. 

Mechanically Assisted Corrosion Basics for CoCrMo and Ti–6Al–4V Alloys MACC is a conjoint degradation process (Fig. 2.4.3.9) that combines mechanics, materials, surfaces, chemistry, and biology. A hard asperity makes contact with the metal

oxide passive film-covered surface where the contact stresses reach the yield point of the substrate. This results in disruption (removal) of the oxide film covering the surface and generates oxide-based debris that can be released from the surface. Metal ions may be released into the solution from the regions where the oxide film has been removed; however, these regions rapidly repassivate (within milliseconds) and return the surface to a passive state. Because the contact stresses reach or exceed yielding, the substrate alloy undergoes plastic deformation (dislocation motion) of the near-surface region, which can work harden the outer alloy layer and develop very high dislocation densities. The corrosion reactions present are associated with both the ionic release of metal and the reaction of metal with oxygen and water for form metal oxides (repassivation). These reactions, because they involve the liberation of electrons, result in currents that can be detected and used to monitor the

948 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

(A)

(B)

(C)

(D)

• Figure 2.4.3.8  Examples of corrosion damage in retrieved Ti–6Al–4V alloy components. (A) Ti–6Al–4V

shell taper from a retrieved acetabular component. (B)–(D) examples of retrieved, sectioned Ti–6Al–4V alloy-corroded taper regions that show thick oxide formation (B), pitting (C), and direct conversion of the alloy (bright regions) to oxide (darker gray region in center). All examples were from retrieved components demonstrating in vivo corrosion damage. (Fig. 2.4.3.8B,C are used with permission from Gilbert, J.L., Mali, S.A., Urban, R.M., Silverton, C.D., Jacobs, J.J., 2012. In-vivo oxide-induced stress corrosion cracking of Ti-6Al-4V in a neck-stem modular taper: emergent behavior in a new mechanism of in-vivo corrosion. J. Biomed. Mater. Res. B Appl. Biomat. 100B(2), 584–594.)



Figure 2.4.3.9 Schematic of mechanically assisted crevice corrosion. (Reprinted with permission. Adapted from Gilbert, J.L., Sivan, S., Mali, S., 2015. Corrosion of modular tapers in total joint replacements: a critical assessment of design, materials, surfaces structure, mechanics, electrochemistry and biology. In: Greenwald, K., Lemons, M. (Eds.), ASTM Special Technical Publication on Implant Modularity, STP 1591. ASTM Int, 192–223 and Gilbert, J.L., Mali, S., 2012. Medical implant corrosion: electrochemistry at metallic biomaterial surfaces, Degradation of Implant Materials. In: Eliaz, N. (Ed.), Springer Press, New York, NY.)

corrosion processes arising from mechanical disruption of the surface. Reduction reactions are also present, but can be remote from the site of abrasion. Much has been learned about MACC from retrievals, but also from fundamental studies of oxide film disruption and repassivation tests that include high-speed single

asperity scratch tests (Gilbert et al., 1996; Goldberg et al., 1997; Goldberg and Gilbert, 2004), to pin-on-disk fretting corrosion tests (Swaminathan and Gilbert, 2012, 2013), and in vitro implant tests (Goldberg and Gilbert, 2003; Gilbert et al., 2009 Jan; Mali and Gilbert, 2015). These tests have, for example, shown that during oxide

CHAPTER 2.4.3   Metallic Degradation and the Biological Environment

949

• Figure 2.4.3.10  Schematic of the area abraded analysis based on plastic deformation. Scanning electron

micrography micrograph is of a 16 μm radium diamond stylus scratching a high-carbon CoCrMo alloy surface under a controlled load (0.4 N). The width of the permanent scratch is determined by the hardness (4 GPa) and the normal load applied.

disruption, the corrosion reactions release free electrons into the metal (due to oxidation), which accumulate within the metal electrode and cause a negative excursion in potential. These electrons are eventually consumed in reduction reactions; however, it is well known that the open circuit potential of alloys engaged in MAC can drop significantly. The potential drop is limited by the potential where metal oxide films will begin to form (passivation potential). For titanium this potential is −1 V versus Ag/ AgCl, while it is about −0.5 to −0.8 V for CoCrMo. The amount the potential drops in these cases depends on the rate of oxide abraded (and repassivated), the amount of available area for the reduction reactions to take place, and the kinetics of the reduction reaction. These effects are also reflective in the changes in impedance of the electrode where larger area electrodes have lower resistances and higher capacitances that will affect the extent of voltage change (Gilbert et al., 2016). In addition, the thickness of passive oxide film that repassivates depends primarily on the electrode potential above the passivation potential. This thickness is linearly dependent on the electrode potential difference (when within the passive range of the oxide). Thus the potential of the metal changes with MACC and the changing potential affects the repassivation processes, and these effects may have significant impacts on the nature of the corrosion damage observed (Figs. 2.4.3.6–2.4.3.8). There are a number of models under development to describe the tribocorrosion behavior of passive film-covered alloys. These rely on the processes of oxide disruption and current arising from film growth (repassivation) and ionic dissolution (Gilbert et al., 1996; Goldberg et al., 1997). The mechanics of abrasion relate to asperity contact and sliding mechanics (Swaminathan and Gilbert, 2012). Here, a hard

(rigid) spherical single asperity of some radius, R, indenting into a surface will penetrate to some load-dependent depth and will result in a contact area (Fig. 2.4.3.10) of 2

Acontact = πa =

( ) L

(

L

)1/2

(2.4.3.5) or acontact = H πH where a is the contact radius, L is the normal load, and H is the hardness of the metal. When this asperity slides some distance, δ, then the area abraded is twice the contact radius times the sliding distance: ( )1/2 L (2.4.3.6) Aabraded = 2aδ = 2δ πH and the volume of oxide abraded (and repassivated) is ( )1/2 L Vabraded = Aabraded h = 2δ h (2.4.3.7) πH Since the repassivation rate is typically much faster than the rates of abrasion of the oxide, the sliding abrasion is the rate limiting step in the generation of currents resulting from MACC. The current due to film formation can be found by determining the volume of oxide film disrupted per unit time (which corresponds to the volume reformed) times the charge per unit volume of oxide created by the oxidation reactions. The volume of oxide abraded (disrupted) depends on the true contact area diameter times the sliding distance per unit time times the oxide film thickness: I ( t) =

ρnF dV Mw dt

(2.4.3.8)

950 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

TABLE   Physical Properties for TiO2, CoCrMo, 2.4.3.1  and 316L SS Oxides (Based on Alloy

Composition) TiO2

CoCrMo Oxide

316L SS

Units



4.45

6.06

5.81

g/cm3

n

4

2.37

2.9

Mw

79.9

76.91

130.09

g/mol

m

1.8

2.0

2.0

nm/V

Eo

−1,000

−500

−100

mV

⍴nF/Mw

21,498

18,020

12,498

C/cm3

where: dV dt

=h

dA dt



= 2h

L dδ πH dt

(2.4.3.9)

Here, δ is the sliding distance, L is the normal load, H is the hardness of the alloy, and h is the thickness of the oxide film. The film thickness, h, varies linearly with potential above the passivation potential by about 2 nm/V (known as m, the anodization rate, see Table 2.4.3.1) and with these, one can obtain the current that arises from oxide film abrasion. Table 2.4.3.1 summarizes the various physical properties of the oxides of Ti, CoCrMo, and 316L SS used in the estimation of currents arising from tribocorrosion processes. For example, if 1 mm2 of oxide (2 nm thick) is abraded and reformed on titanium in 1 s, the associated currents would be about 43 μA of current. This analysis ignores the currents arising from ionic dissolution reactions as these currents are typically much smaller than the film currents (Gilbert et al., 1996; Goldberg et al., 1997). As a comparison, the currents typically measured in fretting corrosion tests of modular tapers are on the order of 1–10 μA, or about 0.1 mm2 abraded area per second. This shows that currents that result from MAC can be highly sensitive to the rate of the area abraded, that the true contact area undergoing oxide disruption is only a small fraction of the nominal area, and that monitoring of these currents is an excellent means for assessing the damage process in experimental setups. The thickness of oxide film reformed depends on the potential. It must be above the passivating potential to form oxide and for CoCrMo and 316L SS must be below the breakdown potential for the alloy (about +500 mV vs. Ag/ AgCl). If the potential exceeds this latter potential, then the passive film is no longer passive and this analysis does not hold. The variation of single asperity scratch current with potential can be seen in Fig. 2.4.3.11A,B, which are the results of single diamond asperity high-speed scratch tests (Gilbert et al., 1996; Goldberg et al., 1997; Goldberg and Gilbert, 2004) on both CoCrMo and Ti–6Al–4V alloy surfaces. Using controlled loads and scratch distances, the current response as a function of applied potential shows that

• Figure 2.4.3.11  Single asperity high-speed scratch test peak currents

plotted against applied potential. These plots show the passivation potential for the two alloys and how the peak scratch current varies linearly with potential over the passive range of the alloy. (Adapted with permission from Goldberg, J.R., Gilbert, J.L., 2004. The electrochemical and mechanical behavior of passivated and TiN/AlN-coated CoCrMo and Ti6Al4V alloys Biomaterials 25 (5), 851–864.)

there is a lower scratch current with more negative potentials that approach zero at unique potentials (−1 V for Ti–6Al– 4V and −0.5 V for CoCrMo), which are the onset potentials or passivating potentials for the respective oxides (primarily TiO2, Cr2O3). The linear increase of current above the passivation potential relates to the linear dependence of the oxide film thickness with potential. Because MAC processes include tribological factors and electrochemical factors, the details of these interactions are important to study. Test methods, including electrochemical fretting pin-on-disk tests, have been developed to explore the range of interactions present (Swaminathan and Gilbert, 2012, 2013). These studies have shown, for example, that there are effects of potential on the fretting coefficient of friction (Swaminathan and Gilbert, 2012, 2013; Liu and Gilbert, 2017) and that the currents generated during fretting corrosion scale directly with the work of fretting (i.e., the energy dissipated per cycle of fretting motion). Fretting corrosion currents generated by oxide film disruption and repassivation can be modeled using a heredity integral approach based on the idea (Fig. 2.4.3.12) that an asperity abrasion process can be thought of as consisting of infinitesimal increments of oxide disruption, dV, and repassivation, and that each increment in oxide disruption

CHAPTER 2.4.3   Metallic Degradation and the Biological Environment

951

• Figure 2.4.3.13  Comparison of predicted currents and experimental cur-

• Figure 2.4.3.12  Schematic of an asperity removing oxide film and the

film regrowing where each infinitesimal volume of oxide removal and regrowth results in a current spike and decay transient. The overall current response is the sum (integral) of each of these infinitesimal processes over time and is related to the sliding speed. (Used with permission from Gilbert, J.L., Zhu, D., October, 2019. A tribocorrosion model linking fretting mechanics, currents and potentials: model development and experimental comparison. J. Biomed. Mat. Res. B Appl. Biomat. (In Review).)

and repassivation results in current transient responses that can be superimposed and summed over time to obtain the full current response. This is similar to the approach of Boltzmann superposition used in viscoelastic theory to predict stresses in a viscoelastic material in response to an arbitrary strain path over time. Here, the currents can be determined from the integral over time of the current transient response of each infinitesimal oxide disruption event. The equation that describes this behavior, known as a Dumahel integral, is dV I (t) = AT (t) V (0) + ∫ tλ = 0 AT (t − λ) dλ dλ

(2.4.3.10)

where V is the oxide volume abraded (and repassivated), λ is a dummy variable in time, and AT(t – λ) is the tribocorrosion admittance (or transfer function) relating currents to volume abraded. This function is ( ) η ρnF io βa − t−λ I T o τ A (t − λ) = e = + e V Mw τ h t − λ t − λ ρnF − τ − e τ ≈ e Mw τ

(2.4.3.11)

where ⍴ is the oxide density, n is the valence of the cation, F is Faraday’s constant (96,485 C/mol), Mw is the molecular weight of the oxide, h is the oxide thickness, io is the exchange current density for ionic release through bare metal, βa is the anodic Tafel slope for ionic release, τ is the time constant for the current transient, and η is the overpotential for ionic release. This equation shows that the

rents resulting from a pin-on-disk fretting corrosion experiment of CoCrMo alloy in phosphate-buffered saline. Tribocorrosion currents can be modeled theoretically with this approach when the potential remains fixed.

current transient that results from an instantaneous removal and repassivation of oxide film will result in an exponential decay of the current and has been developed previously from single asperity high-speed scratch tests of titanium and CoCrMo alloy surfaces in physiologically representative solutions (Gilbert et al., 1996; Goldberg et al., 1997; Goldberg and Gilbert, 2004). Eq. (2.4.3A.10) describes how the current will vary over time, depending on the rate of volume abraded over time. This equation predicts the tribocorrosion currents for any arbitrary sliding distance time response and has been verified experimentally (Fig. 2.4.3.13) where step-like displacements of pin-on-disk coupled to CoCrMo/CoCrMo were tested and comparison of the experimental currents with the theoretical currents shows high similarity. These studies demonstrated that when mechanical disruption of the passive oxide film on titanium and CoCrMo alloys occurs, the resulting corrosion reactions are generated by the repassivation reaction and ionic release and these can be measured and calculated such that the degradation response can be modeled and assessed systematically.

Tribocorrosion Layer and Surface Damage on Metallic Biomaterials Surfaces Studies of the resultant damage to metal surfaces experiencing tribocorrosion processes have identified a mixed tribocorrosion layer that may form within the top few nanometers of the surface wherein the oxide debris mixes with the alloy to result in a layer of mixed metal, metal oxide, and adsorbed protein (Wimmer et al., 2003, 2010). This tribocorrosion layer may alter the properties of the surface (hardness, corrosion response) and affect subsequent tribocorrosion processes. Other studies have investigated the plastic deformation damage observed in the near-surface alloy grains resulting from fretting contacts. Here, highly deformed surface grains may be induced to undergo strain-induced transformation and/or develop

952 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

highly deformed dislocation substructures that appear to be similar to nanocrystalline regions. Such fine crystalline zones may well egress from the surface and be a source of metal particulates released into the surrounding tissues (Oladokun et al., 2019).

 Biology and Corrosion: Additional Insights Starting in the 1970s and 1980s, Hans Willert et al. 1990, 2005) and Jasty and Harris (Jasty et al., 1994) recognized that particulates from a range of biomaterials can induce significant adverse biological reactions. These reactions may be chronic inflammatory reactions, dominated by macrophages, lymphocytes, and foreign body giant cells within granulomatous tissues (Davies et al., 2005), or may result in osteolysis (bone destruction) or soft tissue inflammation. For metallic biomaterials, the effects of tribocorrosion on local tissue responses have only more recently been identified as a result of the introduction of metal-onmetal hip implants. Here, metallic or metal oxide debris generated from tribocorrosion processes may be released into the joint fluid and surrounding joint tissues and be associated with significant adverse local tissue reactions. While the specific causal relationships and mechanisms of interaction are still not well understood, there appears to be an effect of CoCrMo degradation products on the local tissue reaction that is the result of an immunological response to the alloy degradation products (Davies et al., 2005). These effects are described in more detail in other sections of this text (Chapter 2.4.3) and will not be focused on here. However, there are certain interactions between the biological system and metallic biomaterials that merit further attention. In particular, the biological system can affect the corrosion behavior of these alloys and the electrochemical behavior of the alloys can affect the local biology in ways that have not been fully appreciated. This feedback interaction may be a significant factor in why some people are more reactive to metallic degradation debris than others and why some implants show much more severe corrosion damage than others. 

Reduction Reactions Affect Cells The first new insight has to do with how cells respond to metal surfaces that are held to different electrochemical potentials. Work by Ehrensberger and Gilbert (Ehrensberger et al., 2010), Kalbacova et al. (2007), and Haeri et al. (2012) have shown that preosteoblast-like cells (e.g., MC3T3-E1 preosteoblasts) cultured on titanium and CoCrMo surfaces are highly sensitive to the electrode potential of the metal and, in particular, that cathodic potentials (wherein reduction reactions are taking place) can induce cell death (by an apoptotic process) within a few hours. That is, reduction reactions involving the reduction of oxygen and water induce the killing of cells on or near to the metal surface. It is thought that this killing effect may be the result of

consumption of oxygen, shifting of the near-surface solution to higher pH, and/or the generation of reactive oxygen intermediates (ROS) from the reduction process. As shown in Fig. 2.4.3.2, reduction of oxygen and water can create small fractions of ROS and these chemistries can have significant effects on cells cultured on the surface (Wiegand et al., 2019). This is true regardless of the alloy. The killing effect occurs for both mammalian cells and bacteria (Ehrensberger et al., 2015). This approach is now being adopted for the possible control of cancer cells (Kim and Gilbert, 2018) or bacterial biofilms on metallic implant surfaces (Ehrensberger et  al., 2015). Reduction reactions can be induced on metallic medical devices, and such reactions can kill bacteria on the surface. These observations represent a new and potentially significant means for treating infection. In addition, tribocorrosion processes can induce large drops in potential and the increased associated reduction reactions have also been shown to induce cell killing on the surface (Kubacki et  al., 2017). Thus redox reactions not specifically associated with metal ion release or oxide film formation can result in significant cellular effects that are as yet unrecognized and/or understudied. Recent work (Wiegand et al., 2018) has sought to identify an “electrochemical zone of viability,” which is defined as the electrode potential range wherein cells can remain viable and outside of which cells die. For titanium, cells can remain viable on the surface up to at least +1 V versus Ag/AgCl, but that by −0.4 V and lower, cells will rapidly die. For CoCrMo, the zone of viability is limited to +0.3 V where cells on CoCrMo above this potential rapidly die by necrosis and metal ion toxicity, while below −0.4 V cells die by reduction-induced apoptosis. These zones of viability may also be dependent of cell type (Wiegand et al., 2018) where it has been shown that macrophage-like cells can survive for longer periods to more negative potentials than preosteoblast-like cells. This is likely due to their increased ability to control their redox homeostasis in the presence of ROS generated by the electrode surface. Such electrochemical effects on cells may induce a wider range of responses than have been previously understood, and possible sensing and control of systems by electrochemical means remain a potentially important area of study. 

Reactive Oxygen Species May Enhance Corrosion Reactions In addition to the fact that reduction reactions affect biological systems, the effect of immune cell responses on the corrosion of metallic biomaterials is also a new area of focus. In these circumstances, it is hypothesized that the respiratory burst processes associated with a range of phagocytic cells can release a range of ROS and reactive nitrogen species molecules into the near implant space. These oxidizing species, if they come into contact with the alloy surface, have been shown to have significantly deleterious effects on the corrosion resistance, particularly

CHAPTER 2.4.3   Metallic Degradation and the Biological Environment

in CoCrMo alloys (Liu and Gilbert, 2018; Kubacki and Gilbert, 2018). That is, chemistry that includes hydrogen peroxide or hypochlorous acid, or other oxidizing species, lowers the impedance of the oxide films on CoCrMo and titanium alloys and can shift the electrode potential of the metal such that this combined effect can raise the rate of corrosion significantly and alter tribocorrosion processes (Liu and Gilbert, 2017). Such effects may play a significant role in the corrosion behavior of these alloys in vivo and additional work is needed to better understand this corrosion–biology interplay. 

Summary This chapter has outlined the principal means by which metallic biomaterial surfaces undergo degradation in the body. The central role of passive oxide films on these alloys and how mechanical disruption in conjunction with corrosion and biological environments can lead to severe degradation mechanisms. Degradation by corrosion can increase reduction reactions as well and has been shown to have killing effects on nearby cells. Inflammatory species (e.g., ROS) can increase the corrosion response of titanium, CoCrMo, and 316L SS alloys. In all, metal degradation in the biological environment is a complex interplay of conjoint mechanisms and can result in a wide range of damage modes and potential biological outcomes. Our understanding of the tribocorrosion processes present in many metallic implant systems and the physics, chemistry, and mechanics at play will assist in designing metal-based implants for the future.

Acknowledgments There were many students and colleagues that participated in the development of this work and shared in the efforts to document and characterize the degradation of metals in the body. These include: Joshua J. Jacobs, Robert Urban, Jay Goldberg, Christine Buckley, Mark Ehrensberger, Viswanathan Swaminathan, Shiril Sivan, Jua Kim, Sachin Mali, Greg Kubacki, Yangping Liu, Morteza Haeri, and Michael Wiegand. Their efforts and contributions are greatly appreciated by the author.

References Cao, S., Mischler, S., 2018. “Modeling tribocorrosion of passive metals – a review”. Curr. Opin. Solid State Mater. Sci. 22, 127–141. Davies, A.P., Willer, H.G., Capbell, P.A., Learmonth, I.D., Case, C.P., 2005. An unusual lymphocytic perivascular infiltration in tissues around contemporary metal-on-metal joint replacements. J. Bone Jt. Surg. 87A (1), 18–27. Ehrensberger, M., Sivan, S., Gilbert, J.L., 2010. “Titanium is NOT “the most biocompatible metal” under cathodic potentials: the relationship between voltage and MC3T3 pre-osteoblast behavior on electrically polarized cpTi surfaces”. J. Biomed. Mat. Res. A 93A (4), 1500–1509.

953

Ehrensberger, M.T., Tobias, M.E., Nodzo, S.R., Hansen, L.A., LukeMarshall, N.R., Cole, R.F., Wild, L.M., Campagnari, A.A., 2015. Cathodic voltage-controlled electrical stimulation of titanium implants as treatment for methicillin-resistant Staphylococcus Aureus periprosthetic infections. Bioimaterials 41, 97–105. Biological Responses to Metal Implants, US FDA. , September 2019. Center for Devices and Radiological Health. https://www.fda.gov/ media/131150/download. Gilbert, J.L., 2017. Corrosion in the human body: metallic implants in the complex body environment. Corrosion 73 (12), 1478– 1495. Gilbert, J.L., Kubacki, G.W., 2015. “Oxidative stress, inflammation and the corrosion of metallic biomaterials: corrosion causes biology and biology causes corrosion”. In: Dziubla, T.D., Butterfield, D.A. (Eds.), Oxidative Stress and Biomaterials. Elsevier Press (Chapter 3). Gilbert, J.L., Mali, S., 2012. “Medical implant corrosion: electrochemistry at metallic biomaterial surfaces”. In: Eliaz, N. (Ed.), Degradation of Implant Materials. Springer Press, New York, NY. Gilbert, J.L., Zhu, D., October, 2019. “A tribocorrosion model linking fretting mechanics, currents and potentials: model development and experimental comparison”. J. Biomed. Mat. Res. B Appl. Biomat. In Review. Gilbert, J.L., Buckley, C.A., Jacobs, J.J., 1993. In-vivo corrosion of modular hip prosthesis components in mixed and similar metal combinations: the effect of crevice, stress, motion and alloy coupling. J. Biomed. Mater. Res. 27 (12), 1533–1544. Gilbert, J.L., Buckley, C.A., Lautenschlager, E.P., 1996. “Titanium Oxide Film Fracture and Repassivation: The Effect of Potential, pH and Aeration”, Medical Applications of Titanium and its Alloys the Materials and Biological Issues. ASTM Special Technical Publication 1272, American Society for Testing and Materials, Philadelphia, PA, pp. 199–215. Gilbert, J.L., Mehta, M., Pinder, B., January 2009. In-vitro fretting crevice corrosion of stainless steel-cobalt chrome modular hip stems: effect of material, assembly and offset. J. Biomed. Mater. Res. B 88B (1), 162–173. Gilbert, J.L., Mali, S.A., Urban, R.M., Silverton, C.D., Jacobs, J.J., 2012. “In-vivo oxide-induced stress corrosion cracking of Ti-6Al4V in a neck-stem modular taper: emergent behavior in a new mechanism of in-vivo corrosion”. J. Biomed. Mater. Res. B Appl. Biomat. 100B (2), 584–594. Gilbert, J.L., Sivan, S., Mali, S., 2015. “Corrosion of modular tapers in total joint replacements: a critical assessment of design, materials, surfaces structure, mechanics, electrochemistry and biology”. In: Greenwald, K., Lemons, M. (Eds.), ASTM Special Technical Publication on Implant Modularity, STP 1591. ASTM Int, pp. 192–223. Gilbert, J.L., Mali, S.A., Liu, Y., 2016. Area-dependent impedance based voltage shifts during tribocorrosion of Ti-6Al-4V biomaterials: theory and experiment. In: IOP Surface Topography: Metrology and Properties, Special Topic Issue: Surfaces and Interfaces in Bioengineering Systems” 034002; 1–18. Goldberg, J.R., Gilbert, J.L., 2003. In-vitro corrosion testing of modular hip tapers. Appl. Biomater. 64B (2), 78–93. Goldberg, J.R., Gilbert, J.L., 2004. The electrochemical and mechanical behavior of passivated and TiN/AlN-coated CoCrMo and Ti6Al4V alloys. Biomaterials 25 (5), 851–864. Goldberg, J.R., Lautenschlager, E.P., Gilbert, J.L., 1997. Electrochemical response of CoCrMo to high speed fracture of its metal oxide using an electrochemical scratch test method. J. Biomed. Mater. Res. 37 (2), 421–433.

954 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Goldberg, J., Gilbert, J.L., Jacobs, J.J., 2002. “A multicenter retrieval analysis of taper fretting-crevice corrosion of modular femoral total hip prostheses”. Clin. Orthop. Relat. Res. 401, 149–161. Haeri, M., Wollert, T., Langford, G.M., Gilbert, J.L., 2012. Electrochemical control of cell death by reduction-induced intrinsic apoptosis and oxidation-induced necrosis on CoCrMo alloy invitro. Biomaterials 33, 6295–6304. Jacobs, J.J., Gilbert, J.L., Urban, R.M., 1998. Current concepts review, “corrosion of metal orthopaedic implants”. J. Bone Jt. Surg. 80-A (2), 268–282. Jasty, M., Bragdon, C., Jiranek, W., Chandler, H., Mahoney, W., Harris, W.H., 1994. Etiology of osteolysis around porous-coated cementless total hip arthroplasties. Clin. Orthop. Relat. Res. 308, 111–126. Kalbacova, M., Roessler, S., Hempel, U., Tsaryk, R., Peters, K., Scharnweber, D., Kirkpatrick, J., Dieter, P., 2007. The effect of electrochemically simulated titanium cathodic corrosion products on ROS production and metabolic activity of osteoblasts and monocytes/macrophages. Biomaterials 28, 3263–3272. Kim, J., Gilbert, J.L., 2018. In-vitro cytotoxicity of galvanically coupled magnesium-titanium particles on human osteosarcoma SAOS2 cells: a potential cancer therapy. J. Biomed. Mater. Res. B 107B, 178–189. Kubacki, G.W., Gilbert, J.L., 2018. The effect of the inflammatory species hypochlorous acid on the corrosion and surface damage of Ti-6Al-4V and CoCrMo alloys. J. Biomed. Mater. Res. A 106A, 3185–3194. Kubacki, G.M., Hui, T., Gilbert, J.L., 2017. Voltage and wear debris from Ti-6Al-4V interact to affect cell viability during in-vitro fretting corrosion. J. Biomed. Mater. Res. A 106A, 160–167. Liu, Y., Gilbert, J.L., 2017. The effect of simulated inflammatory conditions and pH on fretting corrosion of CoCrMo alloy surfaces. Wear 390–391, 302–311. Liu, Y., Gilbert, J.L., 2018. The effect of simulated inflammatory solutions and fenton chemistry on the electrochemistry of CoCrMo alloy. J. Biomed. Mater. Res. B Appl. Biomat. 106B, 209–220. Mali, S., Gilbert, J.L., 2015. “Correlating fretting corrosion and micromotions in modular tapers: test method development and assessment”. In: Greenwald, K., Lemons, M. (Eds.), ASTM Special

Technical Publication 1591 on Implant Modularity. ASTM International, W., Conshohocken, PA, pp. 259–282. Mischler, S., Debaud, S., 1998. Landolt, “wear-accelerated corrosion of passive metals in tribocorrosion systems”. J. Electrochem. Soc. 145 (3), 750–758. Oladokun, A., Hall, R.M., Neviille, A., Bryant, M.G., 2019. The evolution of subsurface microstructure and tribochemical processes in CoCrMo-Ti-6Al04V fretting corrosion contacts: what lies at and below the surface? Wear 440–441 203095. Popov, V.L., 2010. Contact Mechanics and Friction. Springer Press, New York, NY. Swaminathan, V., Gilbert, J.L., 2012. Fretting corrosion of CoCrMo and Ti6Al4V interfaces. Biomaterials 33, 5487–5503. Swaminathan, V., Gilbert, J.L., 2013. Potential and frequency effects on fretting corrosion of Ti6Al4V and CoCrMo surfaces. J. Biomed. Mater. Res. A 101A (9), 2602–2612. Wiegand, M.J., Kubacki, G.W., Gilbert, J.L., September, 2018. Electrochemical potential zone of viability on CoCrMo surfaces is affected by cell type: macrophages under cathodic bias are more resistant to killing. J. Biomed. Mater. Res. A 107A, 526–534 2019. Wiegand, M.J., Benton, T., Gilbert, J.L., 2019. A fluorescent approach for detection and measuring reduction reaction byproducts near a cathodically-biased metallic surface: reactive oxygen species production and quantification. J. Bioelectrochem. 129, 235–241. Willert, H.G., Bertram, H., Buchhorn, G.H., 1990. Osteolysis in Alloarthroplasty of the Hip: the role of ultrahigh molecular weight polyethylene wear particles. Clin. Orthop. Relat. Res. 258, 95– 107. Willert, H.G., Buchhorn, G.H., Fayyazi, A., Flurry, R., Windler, M., Koster, G., Lohmann, C.H., 2005. Metal-on-Metal bearings and hypersensitivity in patients with artificial hip joints: a clinical and histomorphological study. J. Bone Jt. Surg. 87-A (1), 28–36. Wimmer, M.A., Sprecherb, C., Hauert, R., Tager, G., Fischer, A., 2003. Tribochemical reaction on metal-on-metal hip joint bearings: a comparison between in-vitro and in-vivo. Wear 225, 1007–1014. Wimmer, M.A., Fischer, A., Buscher, R., Porsal, R., Sprecher, C., Hauert, R., Jacobs, J.J., 2010. Wear Mechanisms in Metal-onMetal Bearings: the important of tribochemical reaction layers. J. Orthop. Res. 28, 436–443.

2.4.4

Degradative Effects of the Biological Environment on Ceramic Biomaterials MARIA VALL ET REGI, PEDRO ESBRIT, ANTONIO J. SALINAS Chemical Department of Pharmaceutical Sciences, Faculty of Pharmacy, Universidad Complutense of Madrid, Spain

Introduction Ceramic materials are important sources of biomaterials for clinical applications. In this field, they are usually denoted as bioceramics and continue to be used to regenerate and repair ill or damaged tissues mainly in orthopedics and dentistry. Ceramics are inorganic materials exhibiting a combination of ionic and covalent bonding that show high melting temperatures, low conduction of electricity and heat, and high hardness. In addition, ceramics show high wetting degrees and surface tensions, which favor the adhesion of proteins and cells. Moreover, ceramics show great compression but low tensile strengths, stiffness and little plastic deformation. Bioceramics as implants are often manufactured with interconnected hierarchical porosity, which decreases the mechanical properties but allows ingrowth of cells and the necessary blood vessels. The biological environment is both very aggressive and susceptible to material-provoked adverse reactions (Williams, 2008). Thus implanted bioceramics might suffer surface modifications and degradation during their required lifespan. Plasma is an aqueous medium containing anions (bicarbonate, chloride, phosphate, sulfate), cations (calcium, magnesium, potassium, sodium), as well as more than 20 different types of proteins: albumin, immunoglobulins, and fibrinogen being the most abundant. It also contains dissolved gases, mainly oxygen, carbon dioxide, and nitrogen. Finally, blood also includes red blood cells, white blood cells (neutrophils, eosinophils, basophils, lymphocytes, and monocytes), and platelets. This set of blood components is normally buffered at pH in the range 7.35–7.45, at a temperature of 37°C. Other biological fluids, such as cerebrospinal fluid, synovial fluid, saliva, tears, and lymph present some variations with respect to the components of blood but they are as complex and aggressive as blood itself on any material in contact with them.

Many of the blood components are extremely corrosive such as chlorine ions—present at a concentration that is one-third of that of sea water—or oxygen concentration in venous blood, a quarter of that in air. In addition, the presence of proteins is known to have a significant influence on the corrosive nature of the body fluids (Black and Hastings, 1998). Also of note, in inflammatory or infectious conditions, the biological medium pH can be acidified for short periods, reaching values of 4 or 5. 

Reactivity of Bioceramics When bioceramics are in contact with the biological fluids, three types of behavior can be expected. Whereas some ceramics, named bioinert, remain almost unchanged, others end up being reabsorbed after a certain period of time, and so are named resorbable; a third group of ceramics, called bioactive, undergo a set of reactions on their surface promoting binding to bone and, occasionally, soft tissues. Based on this different behavior, bioceramics are normally classified according to three subdivisions, which are described in depth in other chapters of this book: 1. Inert (or almost inert) ceramics. 2. Resorbable ceramics 3. Ceramics with surface reactivity, i.e., bioactive ceramics. Bioinert ceramics, namely Al2O3, ZrO2, TiO2, or TiN, which exclusively contain covalent bonds that require a high energy to break, are very stable in the biological environmental, as observed in clinical practice. However, these bioceramics are also denoted as “almost inert” (Hench, 1991), since in contact with biological fluids they are usually covered with water, proteins, and other biological components. The former concept of “aging” of inert ceramics, which decreases their mechanical properties, appears to be more associated with an incorrect way of their preparation (Chevalier, 2006). As an example, zirconia, ZrO2, like 955

956 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

most of this type of ceramics, is brittle but when used to design implant components, it is prepared in a polymorphic phase with great toughness. The stability of this polymorph was achieved with addition of yttria (Y2O3) to reach partially stabilized zirconia (PSZ). However, incorrect thermal treatment of this material could cause the deterioration of mechanical properties of zirconia. Resorbable bioceramics are ceramic materials containing ionic bonds that are stable in the air but dissolve in aqueous media. From their chemical structure, it is possible to identify their potential to produce bioceramics with a controlled degradation rate when implanted. Since any material that degrades in the body will release its constituents to the surrounding tissues, it is necessary to select as constituents of bioceramics cations and anions that are harmlessly incorporated in the metabolic processes. Ionic compounds containing sodium or calcium, including calcium phosphates and calcium carbonates, are widely used in this respect. The degradation rate of these bioceramics depends on their chemical composition and microstructure (Bohner, 2000). Hence, tricalcium phosphate [TCP, Ca3(PO4)2] degrades quite rapidly, while hydroxyapatite [HA, Ca10(PO4)6(OH)2] is relatively stable. However, these compounds present variations in this regard: the α polymorphic form of [Ca3(PO4)2] is more soluble than the β form. For this reason, α-TCP is frequently used as an ingredient of calcium phosphate cements, whereas β-TCP is used in bone grafts, pure, or as a component of biphasic calcium phosphates, containing HA and β-TCP. HA-based ceramics exhibit different solubility degrees depending on various factors. In general, HA can only be considered a resorbable material when obtained with a particle size smaller than 10 μm. Moreover, porosity highly affects ceramic solubility. Therefore dense ceramic materials will degrade slowly, while a porous material will be susceptible to a more rapid degradation. It is considered that dissolution rates in  vivo can be predicted from their behavior in simple aqueous solutions mimicking the biological fluid, although there are some differences within the body in this regard, especially in the rate of degradation observed at the different implantation sites. It is possible that cell activity, related to phagocytosis or the release of free radicals, could be responsible for such variations. Between the two extremes in terms of stability, represented by the (almost) inert and resorbable ceramics, there is a relatively small group of ceramics that exhibits a limited reactivity. This is particularly observed with a number of silica-based glasses and glass-ceramics mainly containing Si, Ca, Na, and P. In these ceramics, a selective dissolution of the surface takes place, involving the release of Ca and P, until a SiO2 layer is formed. This is of considerable interest due to the ability of such surface layer to attract calcium and phosphate ions from the body fluids forming a calcium phosphate layer that makes possible the binding to bone, as discussed in other parts of this book. This bone-bonding ability was also reported for HA. In fact, in spite of some controversy regarding HA surface reactivity when soaked in simulated biological solutions, this ceramic is the best-known example



Figure 2.4.4.1 Factors governing the reactivity of ceramics and a classification of bioceramics according to their in vivo reactivity.

of bioactive ceramics. These implanted bioceramics are characterized by the absence of a fibrous tissue layer formation isolating them from the surrounding tissues, as occurs with inert ceramics as implants. A common method to evaluate the bioactivity of a ceramic under in vitro conditions is based on the use of a simulated body fluid (SBF), proposed by Kokubo (Kokubo et al., 1990). SBF is an aqueous acellular solution with a composition of inorganic ions almost equal to that of human plasma. In this assay, the formation of a layer of hydroxycarbonate apatite (HCA) nanocrystals on the material surface indicates an expected positive bioactive response in vivo (Vallet-Regí, 2001). Fig. 2.4.4.1 shows representative examples of these three categories of bioceramics, including the family of carbon-based materials, which exhibit many ceramic-like properties. 

Factors Influencing the Degradation of Bioceramics The stability of ceramics in different environments displays a variable behavior depending on their chemical reactivity. In this regard, it is critical to bear in mind that selection of a particular bioceramic would be governed by a combination of its mechanical and physicochemical properties, depending on the intended application (e.g., as long-term or shortterm implants). For a ceramic to be used as an implant for bone arthroplasty, inert ceramics are often used due to their very low friction coefficient that produces a minimum amount of wear particles. However, for those implant parts in contact with bone, a bioactive or biodegradable ceramic is mandatory to contribute to the process of bone regeneration, since inert ceramics cannot do that task. Besides composition, other factors are known to affect bioceramic reactivity affecting their dissolution behavior in the biological medium (Fig. 2.4.4.1). Among them, the size of the crystalline domains (crystallinity), the particle size, or the presence of crystalline defects are critical. Thus decreasing the crystallinity and grain size augments the presence of

CHAPTER 2.4.4   Degradative Effects of the Biological Environment on Ceramic Biomaterials

957

• Figure 2.4.4.2  Comparison between the structure of a silica-based glass (disordered) and hydroxyapatite (ordered). The corresponding X-ray diffraction patterns are shown at the bottom.

network defects, which increase ceramic solubility. In addition, an increase in the surface area and porosity also promotes the latter. Fig. 2.4.4.2 shows in a comparative manner the different crystallinities exhibited by bioceramics, exemplified by the characteristic disordered structure of a silicabased glass containing a small proportion of phosphorus, calcium, and sodium as network modifiers, and the ordered crystalline structure of HA. These different structures produce very different X-ray diffraction patterns. Increasing the crystallinity of a ceramic such as HA also tends to diminish its reactivity and, consequently, its bioactivity. However, since crystallinity is not the only factor responsible for ceramic reactivity, HA and other crystalline ceramics might exhibit a surface reactivity that makes them capable of binding to bone, i.e., to be bioactive. On the other hand, most of the available compositions of glasses behave as inert compounds when implanted. In fact, in traditional glasses obtained by quenching of a melt, the SiO2 content should not exceed 60 mol%, and must contain over 10%–15% of monovalent oxides, such as Na2O or K2O, and divalent oxides, such as CaO or MgO, to be bioactive. The method of glass synthesis producing materials with radically different textural properties will also have a marked influence on their reactivity. Thus glasses obtained by wet methods, such as the sol–gel route, have high specific surface and porosity, which allow them to display bioactivity even with a SiO2 content of 90 mol%. Likewise, glasses with compositions that make them totally degradable in the biological environment can also be obtained. Concerning the influence of particle size on ceramic bioactivity, glasses with a composition within the bioactivity window and a

particle size of less than 90 μm can be completely degraded in  vivo. Equally, by sintering at high temperatures, HA may reach a large particle size, which promotes formation of a thick fibrous nonadherent layer around the implanted material, thus preventing its bioactivity. In contrast, HA can be obtained by wet methods with very small particle size. This fact together with the presence of crystalline defects or ionic substitutions favors the release of reactive substances in the biological medium and its degradation. The complex behavior of a ceramic as determined by different structural variables justifies that HA was included in the three ceramic categories depicted in Fig. 2.4.4.1. Considering the foregoing, reactivity has turned out to be a more suitable criterion than chemical composition or crystallinity for classifying bioceramics. Therefore slight differences in glass composition give rise to a bioinert, bioactive, or resorbable behavior (Hench and Anderson, 1993). In addition, it is possible to find glasses with identical composition behaving as bioinert or bioactive ceramics when synthesized by either melting or a sol–gel method (ValletRegi et al., 2003a; Salinas and Vallet-Regi, 2017). Moreover, some glass compositions considered bioactive can be completely resorbed when used as particles under a certain size limit, namely 90 μm for Bioglass 45S5 (Hench and Polak, 2002; Salinas et al., 2018). Similarly, in the case of crystalline ceramics the in vivo reactivity of HA can range from almost bioinert, when it is sintered at high temperatures as dense monoliths, to resorbable (Kumta, 2006), when having a particle size under 10 μm, and in between we have the bioactive character generally attributed to HA (Vallet-Regi and Gonzalez-Calbet, 2004).

958 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Biological apatites Ca8.3

0.7

(PO4)4.3(HPO4,CO3)1.7(2OH,CO3)0.15

1.7

dentin 26

30

34



bone 26

30

34



Trabecular bone

enamel

26

30

34



100µm

• Figure 2.4.4.3  Biological apatites present in bone and in tooth enamel and dentin.

An implanted bioceramic may induce extracellular and/ or intracellular responses in  vivo. Implant biocompatibility ensures that these responses are not only harmless but indeed promote the function of target host tissues. In addition, degradability of bioceramics, which in part depends on their physicochemical structure, is also an important aspect to define their biocompatibility in vivo. In general, their predominant ionic bonds and chemical stability make bioceramics highly biocompatible materials for a range of clinical applications. 

Reactivity and Degradation of Natural Apatites A good example for understanding the complexity of factors that influence the reactivity and degradation of bioceramics can be provided by the three types of natural apatites present in tooth enamel, dentin, and bone (Fig. 2.4.4.3). Bones of vertebrates are made of organic–inorganic composite materials whose structure includes an inorganic component made of carbonated and calcium-deficient nonstoichiometric HA. These biological apatite crystals, in the range of 25–50 nm (Vallet-Regi and ArcosNavarrete, 2008; LeGeros, 1991), grow at the mineralization sites of the collagen molecules, which are grouped together forming collagen fibers. Furthermore, a certain hierarchical bone porosity is needed to accomplish bone physiological functions (Frieb and Warner, 2002). The order of magnitude of biological apatites and bone pores is shown in Fig. 2.4.4.4. Bone remodeling, involving a coordinated process of bone degradation and new bone formation, is required for both bone repair and mineral homeostasis. For this reason,



Figure 2.4.4.4  Layout of different orders of magnitude in biological apatite and bone pores.

bone mineral contains nanometric pores and many crystallographic vacancies in the structure (i.e., empty positions in which there should be ions), which produces a calciumdeficient nonstoichiometric apatite. Likewise, this apatite structure can accommodate numerous types of extra ions in the three sublattices: Ca2+, (PO4)3–, and OH−; specifically, carbonate groups that are flat in shape and replace tetrahedral phosphate groups. Their presence in a proportion higher than 4% in bone apatite is very important since it produces structure stresses and increases apatite solubility. Therefore bone apatite is based on nanometric calciumdeficient carbonate HA. All these characteristics give them a high reactivity and ability to readily dissolve during bone remodeling. However, in dentin and especially in enamel, mechanical behavior is most important; thus enamel contains much larger apatite crystals but lower defects, ion substitutions, and nonstoichiometry, compared to bone (Fig. 2.4.4.3). 

CHAPTER 2.4.4   Degradative Effects of the Biological Environment on Ceramic Biomaterials

Evolution in the Use of Bioceramics for Bone Repair By the 1950s, inert ceramics like alumina, Al2O3, began to be used for structural applications in bone replacement due to their in  vivo biocompatibility and excellent mechanical features. In the 1980s, ceramics based on calcium phosphates or calcium sulfates, bioactive glasses, and glass-ceramics started to be used as bone grafts or metallic implant coatings due to their degradation pattern and bioactive behavior. Current developments focus on new porous ceramics designed as scaffolds for hosting cells and bioactive molecules to help drive bone tissue regeneration. New advanced bioceramics, namely mesoporous silica-based materials, ordered glasses, and organic–inorganic hybrids, as well as mesoporous nanoparticles, are under investigation. The field of bioceramics is an excellent example of translational biomedicine, since the synthesized material by chemists and engineers must be functionally validated as an implant into the host. The interaction of bioceramics with living tissues controls the biological behavior and degradation of the implanted materials, which requires the crossover of various clinical specialties and experimental sciences. The development of a bioceramic implant for a specific clinical need in this field involves various stages, including in  vitro and in  vivo biocompatibility tests and other preclinical testing discussed elsewhere in this text. Regarding their mechanical behavior, ceramics exhibit great compression and low tensile strengths; they are stiff and brittle materials wherein failure takes place without plastic deformation. In addition, ceramics show high wetting degrees in a variety of forms: powder, porous pieces, dense pieces, injectable mixtures and cements, or as coatings. Ceramic surface tension favors the adhesion of proteins, cells, and other biological moieties, and the surface might exhibit antibacterial properties after appropriate treatment. However, recently developed ceramics with interconnected porosity show a significant decrease of their mechanical features. Until recently, the most popular solutions in bone repair involved the use of natural materials, using autologous bone or bone allographs from a donor bank, or animal bone. This approach presents some drawbacks such as the need for two surgical interventions (autologous implant) and the risk of infection. Alternative inert, resorbable, or bioactive materials (ceramics) have found promising applications in this respect (Fig. 2.4.4.5). The outcome of different ceramics as implants for bone repair and bone regeneration can be summarized as follows: inert ceramics induce the formation of an acellular collagen capsule, which isolates the material from the host tissue, and thus the artificial nature of the material prevails; bioactive ceramics react with the biological environment giving rise to bony apatite that, in concert with osteogenic cells, promotes new bone formation. These ceramics have excellent biocompatibility and bioactivity, but their mechanical properties are poor, rendering them unsuitable for the

Natural

Autologous bone: Self donor

959

Artificial Bioceramics

Allogenic bone: Tissue bank

Heterologous bone: animal source



Figure 2.4.4.5 Natural bone and artificial bioceramics for human bone repair.

repair of large osseous defects. However, they are excellent for the filling of small defects. Glass-ceramics exhibiting improved mechanical properties with respect to the parent glass can be obtained by thermal treatments. Furthermore, organic–inorganic hybrid materials containing bioactive glasses as the inorganic component can also be manufactured. These hybrid materials have mechanical features like natural bone and are bioactive. The present technology allows design of bioactive ceramics with mechanical properties matching those of bone tissue in bone repair applications. As mentioned earlier, ceramics with a hierarchical porous structure are now being investigated for putative use as scaffolds in bone tissue engineering. Pores, in the range of 2–50 nm (mesopores), enable the loading and release of biological molecules. Moreover, they must exhibit interconnected porosity with macropores and channels in the 2–1000 μm range, which allow cell interaction and vascularization when implanted (Fig. 2.4.4.4). This is an important point because the reactivity of ceramics begins at their surface in contact with a wet medium and in the presence of cells and proteins. This porosity though implies a certain sacrifice of their mechanical properties. The pathway to reach third-generation bioceramics based on porous ceramic scaffolds as implants is depicted in Fig. 2.4.4.6. All of the ceramics mentioned earlier could be conformed into pieces with interconnected and hierarchical porosity within the micron range. Some sort of “smart” behavior is also required, so that these materials can modify their behavior in response to external or internal stimuli or to variations in the environment. The possible ceramic degradation in vivo is an important issue to consider when designing a bioceramic as an implant. This applies to the so-called “traditional”

960 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

• Figure 2.4.4.6  Requirements to be met by scaffolds for bone tissue engineering.

bioceramics of well-proven use in the clinic, and also to the novel bioceramics, whose development is at the forefront of knowledge and currently in progress. This aspect will be dealt with in more detail later on in this chapter. As an example of these new bioceramics, silica-based mesoporous nanoparticles are currently at a preliminary research stage for putative cancer therapy. For this application their small size allows their incorporation into the bloodstream to carry antitumor drugs where necessary and deliver this cargo only where needed after receiving a stimulus that can be internal or external. 

Bioceramic Interactions With the Biological Environment When implanted, bioceramics trigger various cell and tissue responses, including acute and/or chronic inflammation, formation of granulation tissue, and fibrous encapsulation, which depend on the ceramic chemical composition and its structural features. In Table 2.4.4.1, the most important bioceramics used clinically are shown.

Inert Ceramics: First-Generation Bioceramics The most widely used near bioinert ceramics are alumina (Al2O3), zirconia (ZrO2), TiO2, and diverse forms of carbon, such as the low-temperature isotropic (LTI) form of pyrolytic carbon (PyC), glassy (vitreous) carbon, the ultralow temperature isotropic (ULTI) form of PyC and carbon fibers, carbon nanotubes, and graphene. Applications of alumina and zirconia as load-bearing prosthetics in orthopedics and dentistry are similar, based

on their excellent mechanical features such as wear resistance, high strength, and smooth surface; zirconia being superior to alumina for its higher toughness (Chevalier, 2006). Moreover, titanium and its alloys, first used as dental implants, have been subsequently and continuously used as long-lasting implants for their excellent biocompatibility and resistance to corrosion. Carbons are mainly used as coatings for improving the physicochemical and biological (e.g., antithrombogenic and antibacterial) properties of implants (Hu et  al., 2010; Fan et  al., 2010; Allen et  al., 1994; Zhu et al., 2007; Bloomfield et al., 1991) or as fibers in reinforced composites. Despite an almost bioinert character, first-generation ceramics can elicit a foreign body reaction, involving the recruitment of macrophages and the secretion of inflammatory cytokines (Rodrigo et al., 2006), and ultimately the formation of an acellular collagenous capsule isolating the implant from the host tissue (Ratner and Castner, 2002). This provokes interfacial micromovements that might eventually lead to implant failure. However, for most medical applications of these nearly bioinert ceramics, this foreign body reaction is rarely an issue. The first applications of alumina as bone substitute (Smith, 1963) and dental implant were reported in the 1960s (Sandhaus, 1967). In addition, the first bioceramic couple (alumina–alumina), with enhanced wear resistance, was implanted in 1970 (Boutin, 1971). Alumina has been extensively used as an implant material for hip and knee arthoplastia (Oonishi et al., 1981) and as single crystal alumina bone screws and in dental applications since the 1980s (Kawahara, 1981). The mechanical properties (toughness and wear resistance) of alumina were improved by decreasing the grain size of alumina from 10 to 2 μm, increasing purity, lowering porosity, and using hot isostatic pressure as a processing method. Current research studies are focused

CHAPTER 2.4.4   Degradative Effects of the Biological Environment on Ceramic Biomaterials

961

TABLE 2.4.4.1    First, Second, and Third-Generation Bioceramics

Type of Bioceramic

In Vivo Reactivity

Examples

First Generation:

Isolated by a nonadherent fibrous capsule

Alumina: Al2O3 Zirconia: ZrO2

Bioinert nonabsorbable

Second Generation:

Carbons, mainly pyrolytic and as fibers in composites Dissolved after a specific time

Calcium phosphates

Biodegradable resorbable

Calcium sulfate Calcium phosphates and sulfates + ZnO, Al2O3, Fe2O3 Coralline: CaCO3

Bioactive surface reactive

Tightly bonded to the living tissues through a surface reaction

Hydroxyapatite, pure and substituted Hydroxycarbonate apatite Glasses: by melting and sol–gel Glass ceramics: apatite/wollastonite glass-ceramic and Ceravital

Third Generation:

Stimulating living tissue regeneration

Scaffolds of biologically active molecules

on reducing the alumina ball size to make it more appropriate for patients shorter in height; ZrO2-based prostheses being more efficient in this respect. Postoperative failure of alumina, and other bioinert ceramic implants, is mainly a consequence of crack propagation (due to fabrication flaws or corrosion) under loading conditions (Morrell et  al., 2012). Alumina has been used in dentistry for root analogs, endosteal screws, blades, and pin-type dental implants, and even ceramic crowns (Li and Hastings, 1998). However, mechanical tests showed functional limitations in some of these long-term devices. In this regard, single crystalline alumina showed mechanical strength superior to polycrystalline alumina as high-load devices. Alumina ceramics have also been used for total knee prosthesis, middle ear implants, and in ophthalmology (Schulte, 1990; Polack and Heimke, 1980). Partially stabilized (with MgO or Y2O3) zirconia has been gaining market in orthopedics and dentistry as an option to alumina because of its enhanced mechanical properties, mainly toughness (Christel et  al., 1988; Afzal, 2014; Akagawa et al., 1993). Moreover, in dentistry, compared to conventional titanium implants, the white color of zirconia makes it more similar to natural teeth. Large volume variations of around 5% take place when zirconia transforms from tetragonal to monoclinic polyform, which leads to an increase in the risk of cracking. Thus zirconia is usually stabilized by adding small amounts of oxides such as calcia (CaO), magnesia (MgO), or yttria (Y2O3), leading to PSZ or tetragonal zirconia polycrystals, which are partially or totally in tetragonal phase, respectively. PSZ zirconia

Bioglass: in particulate form Porous bioactive and biodegradable ceramics Advanced bioceramics: Mesoporous materials, organic–inorganic hybrids

represents the main form of modern medical-grade zirconia. This material also presents advantages over alumina in terms of lower friction and wear resistance. However, the use of zirconia as a bioceramic elicits adverse outcomes produced by surface degradation releasing particles and also radioactive impurities. Degradation is a major drawback in this respect, due to the phase transformation accelerated with material aging in the aqueous body environment, particularly under dynamic loads. Although surface degradation of zirconia balls due to phase transformation does not seem to be significant, and particulate debris is chemically stable and biocompatible, femoral head zirconia prostheses have a relatively short history and further investigation is required. Putative improvements of these bioinert ceramics might come from the preparation of alumina/zirconia composites (Kurtz et al., 2014). Titanium and its alloys have been used as orthopedic and dental implants for more than 40 years (Elias et al., 2008). Upon air exposure, the formed outer TiO2 layer prevents corrosion and greatly improves the biocompatibility of titanium by adsorbing proteins of the host environment (Liu et al., 2004). Attempts to further improve the biocompatibility of Ti-based biomaterials by HA coating and adsorption of osteogenic peptides have been carried out (Khorasani et al., 2015; Van der Stok et al., 2015). Several carbon allotropes, including nanocrystalline glassy carbon, graphene, diamond-like carbon, and PyC, are widely used as implant coatings, since they show optimal adherence to both metallic and polymeric substrates and are not degraded in  vivo. The most common form of carbon

962 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

TABLE 2.4.4.2    Clinical Applications of First- and Second-Generation Bioceramics

Ceramic

Clinical Application

Examples

Bioinert

Coatings for tissue ingrowth: orthopedic, dental, maxillofacial

Al2O3

Orthopedic load-bearing applications

Al2O3, ZrO2, PE–HA composite

Artificial heart valves

Pyrolytic carbon coatings

Artificial tendon and ligament

PLA–carbon fiber composites

Coatings for prosthesis fixation: orthopedic, dental, maxillofacial

HA, bioactive glasses, and glass-ceramics

Percutaneous devices

Bioactive glasses, bioactive composites

Periodontal pocket elimination

HA, HA–PLA composite, sodium phosphate, calcium phosphates, bioactive glasses

Spinal surgery

Bioactive glass-ceramics, HA

Dental implants

Al2O3, HA, bioactive glasses

Alveolar ridge augmentations

Al2O3, HA, HA–autogenous bone, bioactive glasses

Maxillofacial reconstruction

Al2O3, HA, HA–PLA composite, bioactive glasses

Otolaryngological applications

Al2O3, HA, glasses, glass-ceramics

Orthopedic fixation devices

PLA–carbon fibers, PLAca/p glass fibers

Temporary bone spacers, fillers

Calcium phosphates (α- and β-TCP, etc.), calcium sulfate

Bioactive

Bioinert and bioactive

Resorbable

HA, Hydroxyapatite; PE, polyethylene; PLA, poly(lactic acid); TCP, tricalcium phosphate.

used in this regard is PyC, presented in two forms: LTI and ULTI (Dauskardt and Ritchie, 1993). The former material presents excellent biocompatibility with the cell and protein components of blood and soft tissues, and it has received special attention related to its nonthrombogenic behavior which has led to its wide use in prosthetic heart valves. Moreover, LTI carbon shows no degradation and excellent strength and resistance to wear and fatigue in vivo. At present, over 90% of mechanical heart valve prostheses use LTI–PyC as a coating on a polycrystalline substrate or as a monolithic material. PyC can also be used in small orthopedic joints such as fingers and as spinal inserts. However, current efforts are being addressed to eliminate the traces of silicon carbide in PyC, due to its thrombogenicity, and to improve PyC mechanical properties. Future improvements of almost bioinert ceramics could be derived from manufacturing alumina/zirconia composites or, alternatively, from research on nonoxide bioinert ceramics, namely nitrides and carbides like Si3N4 or SiC. Table 2.4.4.2 shows important clinical applications of bioinert ceramics and second-generation ceramics that will be described in the next section. 

Resorbable and Bioactive Ceramics: SecondGeneration Bioceramics Bioactive ceramics denote those ceramics whose surface is able to form a strong bond with soft and hard tissues (Hench

et al., 1971), making them suitable for many clinical applications. As mentioned in another section of this chapter, typical examples of bioactive ceramics are HA and some compositions of glasses and glass-ceramics (Vallet-Regi et al., 2006; Hench and Wilson, 1984; Yuan et  al., 1999; Vallet-Regi et al., 1999; Livingston et al., 2002; Kokubo et al., 2003). In contact with biological fluids, dissolution products of these ceramics cause the formation of an HA-like layer on their surface, whereby they can promote cell attachment and cell differentiation (Kokubo, 1998; Sanders and Hench, 1973; Xynos et  al., 2001). The apatite particles on the surface of bioactive ceramics are easily recognizable by scanning electron microscopy because they have a needle-like shape of around 500 nm in length, agglomerated in pseudospherical particles of 10 μm in diameter. These bioceramics are osteoinductive materials, so that they activate bone tissue regeneration (Hench and Polak, 2002). They have excellent features in terms of biocompatibility and bioactivity, which are strongly associated with ion release from their surface into the surrounding tissue (Hoppe et  al., 2011). Furthermore, these implanted bioceramics induce no deleterious effects related to tissue inflammation. However, prepared as dense monoliths, they are brittle, which renders them unsuitable for repairing large osseous defects. These bioceramics are available in different formats: powder, porous pieces, dense pieces, injectable mixtures, and coatings, dependent on the intended application. Thus a bioactive glass shaped as a porous piece exhibits a

CHAPTER 2.4.4   Degradative Effects of the Biological Environment on Ceramic Biomaterials

fast bioactive response. In this form, the biological fluid can reach the inner structure of the glass more effectively, thus improving its reactivity. These bioceramics are excellent for the filling of small bone defects, where the rate of bone regeneration is the main concern, but the mechanical properties are a secondary aspect. In addition, glasses can also be used for the production of glass-ceramics (see later) by thermal treatment (Roman et al., 2003), which exhibit a somewhat lower bioactivity than glasses but mechanical properties closer to natural bone. In addition, both bioactive glasses and glass-ceramics can be used to design magnetic materials for promoting bone regeneration and, simultaneously, killing bone cancer cells by hyperthermia (Del Real et al., 2002). Biodegradable ceramics are designed to fulfill a specific function for a certain period of time to help tissue regeneration, and they are resorbed thereafter. In this manner, possible deleterious effects associated with a prolonged permanence of a synthetic biomaterial in the host can be avoided. Calcium phosphates and calcium sulfates are typically considered major members of degradable ceramics (Elliott, 1994; De Groot, 1991; Dorozhkin and Epple, 2002; Pietrzak and Ronk, 2000). In vivo ceramic degradation usually occurs by passive dissolution of ions and/or by cell-mediated resorption in the biological environment (i.e., by enzymatic activities and pH changes). Host monocytes/macrophages and even osteoclasts from the bone marrow rapidly invade the bioceramic implant surface and play a key role in material degradation (Rae, 1986). Thus phagocytosis by monocytes/macrophages or acidification by osteoclasts causes the resorption of the bioceramic implant (Heymann et  al., 1999). In addition, osteoblasts and other mesenchymal cells present at the implantation site have been shown to be able to phagocytose calcium phosphate (Heymann et al., 1999; Gregoire et al., 1990). As mentioned earlier in this chapter, decreasing the particle size or increasing the hydrophilic character can promote bioceramic degradation (Vallet-Regi et  al., 2001; Colilla et  al., 2006). Moreover, the loss of mechanical integrity of these ceramics during degradation might contribute to material disintegration (Tamimi et al., 2012; LeGeros, 1993). The latter generates particles that can adversely affect extracellular matrix production by osteoblasts and may result in periimplant osteolysis (Grover et al., 2003). Hence, a critical aspect in designing resorbable ceramics will be to adjust the ceramic degradation rate, which is usually faster than the process of tissue healing, to a specific biomedical application. Present research is focused on the development of porous and new advanced bioceramics as scaffolds for hosting cells and bioactive molecules (growth factors, small peptides) to be released into the biological environment in a controlled manner for tissue-engineering applications (Langer and Vacanti, 1993; Salinas et  al., 2013). This is discussed in detail in other sections of this chapter. Calcium phosphates. Most of the second-generation ceramics used for bone regeneration are based on calcium

963

phosphates Zhang et  al. (2007); Dorozhkin (2010); Lew et al. (2012); Daculsi et al. (2014); Wang et al. (2014); Habraken et al. (2016). These include: synthetic apatites, pure and substituted biphasic mixtures of calcium phosphates, and calcium cements, containing calcium phosphates and calcium sulfate. Synthetic apatites. HA and type B HCA are the most investigated calcium phosphate bioceramics due to similarities with the inorganic component of bone (Elliott, 1994). Submicrometric particles of HA can be obtained by aerosol pyrolysis (Vallet-Regi et al., 1994), precipitation (Naasaraju and Phebe, 1996; Vallet-Regi et  al., 1997), or the liquid mix technique (Pechini, 1967; Peña and Vallet-Regi, 2003). The biological type B HCA, where CO32− are located in the PO43− positions, is synthesized using moderate temperatures (Doi et al., 1998). At higher temperatures a high CO32− content is included in the structure but located in the OH– sublattice; in this way a nonbiological type A HCA is obtained (LeGeros, 1991). HA structure can accept many compositional variations in the Ca2+, PO43−, or OH– sublattices. Thus apatite crystals usually included ions such as Na+, K+, Mg2+, Sr2+, Cl–, F–, or HPO42− (Jha et al., 1997). The migration of these extra ions toward the surrounding tissues produces an increase in the crystal size, decreasing the solubility of the bioceramic. This effect has positive physiological consequences in the rapidly growing young—and less crystalline—bone. Ion substitutions in the HA structure turn bone into an important system to quench toxic heavy metals in the biological milieu. In addition, the ability to host ions allows the design of synthetic calcium phosphates with improved properties for specific biomedical applications. The ionic substitutions can modify the surface structure and electric charge of HA, with potential influence on the biological environments. In addition, while carbonate and strontium ions facilitate apatite dissolution, the release of certain chemical elements, namely strontium, zinc, or silicate, during resorption facilitates bone formation. Silicates in the network increase the mechanical strength of HA, and accelerate its bioactive response (Vallet-Regi and Arcos, 2005; Gibson et  al., 1999). Thus the current trend is to obtain synthetic calcium phosphates partially substituted with different ions for use as implants. Biphasic mixtures of calcium phosphates. The most popular materials of this type are based on HA and β-TCP, β-Ca3(PO4)2, mixtures (Daculsi, 1998; Tancred et  al., 1998; Kivrak and Tas, 1998; Sanchez-Salcedo et al., 2008a), which can evolve to HCA in the biological environment. This process involves the gradual dissolution of the mixture, releasing Ca2+ and PO43− to the local environment, thus providing a source of newly formed bone minerals. This material can be injected and used as a coating or as bulk in bone replacement. Moreover, other biphasic mixtures have been investigated, including calcium phosphates together with other second-generation ceramics, namely bioactive glasses and calcium sulfates (Ramila et al., 2002; Cabañas et al., 2002).

964 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Bone cements. Cements based on calcium phosphates, calcium carbonates, or calcium sulfates have attracted much attention due to their excellent biocompatibility. Most of the injectable calcium phosphate cements evolved to either a calcium apatite or octacalcium phosphate during the setting reaction. Physical–chemical properties of these materials, such as the setting time, porosity, and mechanical behavior, depend on the cement formulation and the presence of additives (Ginebra et al., 2004; Otsuka et al., 1995; Nilsson et al., 2002). These cements cured in field are biocompatible and they slowly resorb, while newly formed bone grows and replaces the cement in vivo. Research is under way on shortening the curing time in contact with blood and improving the mechanical toughness. Calcium phosphate coatings. HA and other calcium phosphates are also used as coatings to protect the surface of metals (i.e., titanium alloys, Ti6Al4V, and commercially pure Ti) from the biological environment. Commercial implants with such coatings are produced by plasma spray at high temperature (Koch et al., 1990; García-Sanz et al., 1997; Sun et al., 2001), which is advantageous in terms of a rapid deposition rate and a relatively low cost. However, the formation of resorbable amorphous calcium phosphate (ACP) in the coating by this method compromises implant integrity. Other unresolved issues related to these coatings are the adherence to substrate and instability of HA at high temperatures or the phase transition of titanium at 1156 K. Alternative techniques to obtain the coatings at lower temperature, including physical vapor deposition (PVD) (Park et  al., 2005), chemical vapor deposition (CVD) (Cabañas et  al., 2003), magnetron sputtering (Wolke et  al., 1998), electrophoretic deposition (Kannan et  al., 2002), pulsed laser deposition (PLD) (Arias et  al., 2003), and sol–gelbased dip coating (Hijon et al., 2006), allow a higher control of the coating thickness and crystallinity of phases. Glasses. Specific compositions of glasses react with body fluids forming an HCA layer, which favors the generation of a bone matrix and bone growth. These glasses are used for filling osseous cavities, substitution of ear ossicles, maxillofacial reconstruction, and dental applications (Vallet-Regi et al., 2003a,b). Those glasses exhibiting a high bioactivity are eligible as bioactivity accelerators of mineral apatites or as bioactivity inductors of magnetic materials for hyperthermia treatment of osseous tumors. Melt glasses. The first bioactive glasses, prepared by quenching of a melt, contained SiO2 and P2O5 as network formers, and CaO and Na2O as network modifiers (Hench et  al., 1971). By using the SBF test, the silica-rich layer formed on the glass surface by the ionic exchange of calcium and sodium ions in the glass and protons in the solution were initially found to attract calcium, phosphate, and carbonate ions in the surrounding fluid to form ACP. This then crystallizes into a biologically active HCA layer that promotes new bone formation (Vallet-Regi, 2001). Those glasses with the highest bioactivity, as tested by this method, were able to bond to both hard and soft tissues. To explain these differences in reactivity, in 1994 Hench defined two

classes of bioactivity: class A (osteoproductive) and class B (osteoconductive) (Hench, 1994, 2016). The high bioactivity of sol–gel glasses (SGGs) makes them excellent candidates to be used in coatings, mixed materials, or as porous scaffolds in third-generation biomaterials, which will be described later on in this chapter. Sol–gel glasses. The sol–gel synthesis of bioactive glasses produces materials (fibers and highly porous monoliths, or coatings) with high textural properties and excellent biocompatibility in both cell cultures (Olmo et al., 2003) and animal models (Gil-Albarova et al., 2005). The sol–gel route is more time consuming than the traditional quenching of a melt, but it requires noticeably lower temperatures. The CaO–P2O5–SiO2 system has been the most studied type of SGGs (Li et  al., 1991; Pereira et  al., 1994; Peltola et  al., 1999; Salinas et  al., 2001), containing additions such as MgO, ZnO, or removing P2O5, to modify the mechanical properties, the resorption rate, or the interaction with osteoblasts (Perez-Pariente et  al., 1999; Du and Chang, 2004; Oki et al., 2004). CaO plays an essential role in the texture and bioactive response of SGGs (Martinez et  al., 2000); Ca2+ ions released to the medium increase exposure of the silanol (Si–OH) groups on the glass surface, which favors the formation of the HCA layer. Of note, formation of the latter depends on the presence of P2O5 in the glass (Salinas et  al., 2002). CaO–SiO2 glasses exhibit higher reactivity. In contrast, in CaO–P2O5–SiO2 glasses, ACP formation is slower but evolves more rapidly to HCA nanocrystals than CaO–SiO2 glasses. Mesoporous bioactive glasses. (MBGs), first synthesized by Zhao et al. in 2004 (Yan et al., 2004), have more controlled and reproducible textural properties than SSGs. They also exhibit twofold higher surface areas and pore volumes than those of SGG analogs with identical composition (Izquierdo-Barba et al., 2013), consistent with the observed faster bioactivity kinetics than the latter glasses. Also, in contrast to SGGs, MBGs exhibit ordered mesopores with a pore diameter between 2 and 10 nm in a very narrow pore size distribution. MBGs also have abundant silanol groups on their surface that allow their functionalization as required and facilitate an excellent performance in SBF (Vallet-Regi and Salinas, 2017). The large surface area of MBGs is thought to be responsible for the high reactivity in SBF medium reaching a pH of 6.7 on the surface during the first stages of the assay, instead of a pH of 7.4 commonly observed in these assays for bioactive melt and SSGs in this assay (Vallet-Regi and Salinas, 2018). This lower pH made possible the formation of octacalcium phosphate as an intermediate phase between ACP and calcium-deficient HA (CDHA), as reported to occur during bone maturation in vivo. All these characteristics make MBGs excellent candidates for bone tissue-engineering applications. As for the SGGs, most of the investigated MBGs belong to the SiO2–CaO–P2O5 system synthesized by the evaporation-induced self-assembly method (Brinker et al., 1999). The CaO content and the temperature used for the aging of gel during the process of synthesis influence the

CHAPTER 2.4.4   Degradative Effects of the Biological Environment on Ceramic Biomaterials

order of mesopores obtained (García et al., 2009). Analysis of the mesostructure of MBGs and SGGs by highresolution transmission electron microscopy allowed a better understanding of the role of the P2O5 content in the behavior of these glasses in SBF medium (Vallet-Regi et  al., 2005b; Vallet-Regi and Salinas, 2018). In both types of glasses, P2O5 binds to calcium to form calcium phosphate nuclei, removing calcium ions from the glass network and, as a consequence, retarding ACP formation in the in  vitro assay. On the other hand, the formation of CDHA identical to that of biological apatite is accelerated. In  vitro studies using SBF assay showed the different bioactive responses of two SGGs and two MBGs, with or without P in their composition, respectively. The presence of P was quite important for SGGs and minimal for MBGs. There is now a growing tendency to include small amounts of other oxides with biological activity, such as CeO2, Ga2O3, ZnO, CuO, SrO, CoO, or Ag2O, for improving the biological performance of MBGs (Salinas et al., 2011; Kargozar et al., 2018; Sanchez-Salcedo et al., 2018). Surfactants are a necessary addition during MBG syntheses for self-assembly in aqueous solution to produce mesophases acting as structure-directing agents of the glass network. The most common mesoporous arrangements obtained for MBGs are 2D hexagonal, spatial group (SG) P6mm, and 3D cubic bicontinuous SG Ia-3d. Bioactive glass coatings. As previously mentioned for calcium phosphate, many techniques can be used to obtain ceramic coatings on metallic substrates, including PLD, PVD, CVD, plasma spray, and dip coating. The latter method controls the porosity, roughness, and composition of the bioactive film formed on the metal surface. As an example of the stability of such coatings, a bioactive glass composed of 20%CaO–80%SiO2 (mol%) coating the Ti6Al4V surface completely dissolved in 14 days in the SBF assay (Izquierdo-Barba et al., 2003). However, in the presence of osteoblasts, this coating not only remained stable in the culture medium, but it promoted cell adhesion, proliferation, and differentiation (Izquierdo-Barba et al., 2006). The influence of the substrate on the behavior of SGG coatings is well illustrated by using 316L stainless steel as substrate. In this case, silicon in the glass bonds to chromium in the substrate during the dip-coating synthesis, subsequently increasing the corrosion of this metallic alloy (Vallet-Regi et al., 2003b). Mixed materials containing bioactive SGGs. Mixed materials, including bioactive SGGs, have been investigated, including: SGG–polymer–drug, SGG–magnetic component, and SGG–HA. SGG–polymer–drug materials were designed to obtain systems with controlled release of bioactive drugs. In this case, bone integration is improved, and the drug release is favored by the ionic exchange between the glass and the biological medium. SGG–magnetic glass–ceramics have been designed for treating bone tumors, based on the higher sensitivity of cancer cells to hyperthermia (Arcos et al., 2003). The in vitro

965

biocompatibility of these mixed ceramics has been established (Serrano et al., 2008). SGG–HA materials have also been designed to ameliorate the bioactivity of HA. Thus a 30%CaO–70%SiO2 glass/HA mixture was found to be covered with an HCA layer after 12 h in SBF, whereas HA alone was not modified in this regard after 45 days. The highest reactivity in this in vitro setting was achieved with biphasic HA and dried-gel glass mixtures with an average particle size of 30 μm. Glass-ceramics. This family of compounds is obtained by the thermal treatment of glasses that produced the formation of crystalline phases inside the glass matrix. The synthesis of a glass-ceramic requires the stages of nucleation and growth. The former is favored by the addition of nucleating agents. Further annealing causes uniform crystal growth. Thus glass-ceramics behave as composites with the crystalline phases strengthening the glassy matrix. Consequently, they exhibit superior mechanical properties compared to the parent glass, and also to sintered crystalline ceramics, but without matching those of cortical bone (Vallet-Regi et al., 2005a,b). Glass-ceramics with rapid bioactive response were synthesized by thermal treatment of SGGs (Padilla et  al., 2005; Boccardi et  al., 2017). However, glass-ceramics present a poorer bioactive response than glasses due to a lower abundance of Si–OH groups on their surface and a decreased release of calcium ions from the crystalline phases. The bioactive response, expressed as the bioactivity index IB = 100/t0,5bb (day−1) (t0,5bb days required for 50% bioceramic surface bonding to bone) varied for different ceramics between 0 for alumina and other bioinert ceramics; 2.3 for HA; 3.2 for apatite/wollastonite glass-ceramic; and 12.5 for Bioglass 45S5 (Hench, 1998). Table 2.4.4.3 shows important bioactive and resorbable bioceramics and their clinical applications. 

Third-Generation Ceramics Third-generation bioceramics are based on resorbable or bioactive materials obtained in a porous form that allows them to act as scaffolds for cells and inducting molecules to drive tissue regeneration. This category includes bioceramics based on porous second-generation bioceramics, like nanometric apatites, shaped as pieces with interconnected porosity, and new advanced bioceramics like silica mesoporous materials, mesoporous ordered glasses, or organic–inorganic hybrids. Manufacturing of these ceramic scaffolds for tissue engineering requires the use of conformation methods that yield pieces with interconnected porosity and pores in the 2–400 μm range (Sanchez-Salcedo et al., 2008b). Several of these methods proceed at room temperature (Roman et al., 2008), which makes it possible to include molecules with biological activity. Silica-based mesoporous materials. These materials were initially proposed for use as drug delivery systems in biomedical applications (Vallet-Regi et al., 2001), and thereafter as bioactive ceramics (Vallet-Regi et al., 2006).

966 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

TABLE 2.4.4.3   Bioactive and Resorbable Bioceramics and Their Uses as Implants

Ceramic

Form

Application

Function

Calcium phosphates

Bulk

Bone graft substitutes, cell scaffolds

Replace the bone loss

Coatings

Surface coatings on total joint prosthesis

Provide bioactive bonding to bone

Glasses

Bulk

Endosseous alveolar ridge maintenance

Space filling and tissue bonding

Middle ear prosthesis

Replacement of part of the ossicular chain

Orbital floor prosthesis

Repair damaged bone supporting eye

Fixation or revision arthoplasty

Restore bone after prostheses loss

Filler in periodontal defects

Periodontal disease treatment

Bone graft substitutes and cranial repair

Augmentation after diverse illness or traumas

Vertebral prostheses

Replace vertebrae removed by surgery

Iliac crest prostheses

Substitute bone removed for autogenous graft

Coatings

Fixation of hip prosthesis

Provide bioactive bonding

Bulk and powder

Bone graft substitutes

Repairing osseous tissues

Powder

Glassceramics

Calcium sulfate

Bulk

They are synthesized with surfactants acting as templates to obtain materials with ordered porosity—in the nanometer range—and amorphous silica pore walls. Previously described MBGs obtained in this manner as powder can be processed to form scaffolds enabling decoration with osteogenic substances (Perez et  al., 2018). Moreover, bioactive glass microspheres with accelerated deposition of HCA have shown to exhibit hemostatic efficacy (Ostomel et al., 2006) and magnetic silica microspheres for drug targeting have been reported (Ruiz-Hernandez et al., 2008). Organic–inorganic hybrid materials. These materials are synthesized at room temperature to preserve the organic component as bulk, coatings, and fibers. Organic– inorganic hybrids containing SiO2–CaO glasses as the inorganic component and several biocompatible polymers as the organic component have been investigated with regard to their mechanical features comparable to bone (Vallet-Regi and Arcos, 2006). In some hybrid systems, only weak physical interactions exist between inorganic and organic domains with several containing poly(vinyl alcohol) (PVA) (Pereira et  al., 2000; Martin et al., 2005) or poly(hydroxyethyl methacrylate) (Schiraldi et al., 2004). In other cases, chemical links between both components are formed, as occurs with hybrids containing poly(methyl methacrylate) (Rhee and Choi, 2002), poly(ε-caprolactone) (Rhee, 2004), or gelatin (Ren et al., 2002; Coradin et  al., 2004). To improve the bioactive response of these hybrids, inorganic components such as TiO2 were added to hybrids with poly(dimethylsiloxane) (PDMS) (Chen et  al., 2000; Manzano et  al., 2006a) or poly(tetramethylene oxide) (Miyata et al., 2004). Examples

of hybrid systems investigated for bone repair include: (1) CaO–SiO2–PDMS, obtained as bioactive coatings (Hijon et  al., 2005) or as bioactive monoliths with or without P additions to improve bioactivity (Salinas et  al., 2007; Manzano et al., 2006b); (2) CaO–SiO2–P2O5–PVA, synthesized as transparent films (Pereira et  al., 2000) or as bioactive and degradable monoliths (Martin et al., 2005); and (3) CaO–SiO2-based hybrids with organic polymers containing methacriloxy and amino groups (Colilla et al., 2006), designed with tailored bioactivity and degradability proposed for time release of bioencapsulates (Gonzalez et al., 2008). Components of bioactive organic–inorganic hybrids are based on the so-called star gels, a family of materials that exhibit remarkable mechanical properties (Michalczyk and Sharp, 1995). These materials have a unique structure made of an organic core surrounded by flexible arms that yield materials with a mechanical behavior between that of glasses and rubbers (Sharp, 1998). Addition of Ca2+ ions conferred bioactivity in SBF to the star gels (Manzano et al., 2006a,b). Notably, these hybrids exhibit fracture toughness comparable to that of the human tibia. Organic–inorganic hybrid materials were also synthesized as porous scaffolds (Tsuru et al., 2008) with interconnected pore structure. Thus SiO2–PDMS with 90% porosity and pore size between 200 and 500 μm, obtained using sieved sucrose particles as a template (Yabuta et  al., 2003), were evaluated in a bioreactor and implanted in vivo into brain defects. Gelatin-siloxane (3-glicidopropyltrimethoxisilane, GPTMS) hybrid scaffolds with different orders of porosity (5–10, 30–50, and 300–500 μm), obtained by freeze drying wet hybrid gels (Ren et  al., 2002), were also assessed after

CHAPTER 2.4.4   Degradative Effects of the Biological Environment on Ceramic Biomaterials

967

CASE STUDY: BIPHASIC CALCIUM PHOSPHATES What Problem Was Addressed?

What Partially Degradable Bioceramic Was Used?

The treatment of bone defects resulting from surgery, trauma, or disease requires the use of biomaterials to achieve faster and better healing outcomes. Indeed, the use of biomaterials is an essential requirement to treat critical bone defects where their size is too large to heal on their own. Autografts and allografts are the gold standard to be used in this respect. However, these approaches have important drawbacks, including the limited supply of donor bone grafts, secondary trauma in harvesting autografts, and the immune response and/or the possibility of disease transmission for allografts. These facts stimulated interest in the development of new synthetic materials with similar properties to native bone for bone replacement and regeneration. 

BCPs are a combination between the more stable bioactive phase of hydroxyapatite (Ca10(PO4)6(OH)2, HA) and the more soluble β-tricalcium phosphate (Ca3(PO4)2, β-TCP), thus allowing control of the material dissolution rate and the tailoring of mechanical properties. BCPs were first elaborated in 1986 by LeGeros, although some of the materials previously implanted as HA were actually a mixture of HA and β-TCP. The most common way to prepare BCPs is the thermal treatment of calcium-deficient apatites obtained by precipitation. When the calcium deficiency is small, that is, the Ca/P molar ratio is close to 1.67, the amount of β-TCP formed after calcination is very small. However, when the deficiency in calcium increases to Ca/P values close to 1.50, the amount of β-TCP in the BCP becomes increasingly high. Therefore by varying the calcium deficiency of the precursor hydroxyapatite, BCPs with the desired proportion of HA and β-TCP can be obtained. Currently, there are over 30 commercially available BCPs as bone substitutes for various orthopedic and maxillofacial applications. At present, there is no general agreement on the optimum proportions of HA and β-TCP in BCP. The majority of these commercial products contain a 60:40 ratio of HA:βTCP, but there are also several products with other ratio values: 20:80, 65:35, or even 96:4. Variations of pH, temperature, and duration of the sintering process produce unique materials with different chemical and physical properties that lead to different tissue responses. Furthermore, BCP has been used as either a carrier or delivery system for therapeutic drugs (antibiotics, hormones, and growth factors) in bone tissue engineering. Despite the many BCPs in use, much research must be done concerning the optimum features for specific applications, including composition and grain size, to improve the desirable osteoinductive behavior of these materials.

What Properties Were Required of the Material? In the context of bone regeneration, the design of new biocompatible, osteoconductive, and osteoinductive materials is an important challenge. With this aim, calcium phosphate bioceramics are processed as porous materials, which makes tissue ingrowth possible. Moreover, they exhibit a bioactive surface that facilitates cell attachment, proliferation, and differentiation. With this application, a family of compounds designated biphasic calcium phosphates (BCPs) is a promising alternative for bone reconstruction. BCPs are formed by two phases: one is more stable, which contributes to its bioactive and osteoconductive surface; the other degrades with time and releases phosphate and calcium ions to the medium enabling new bone formation and causing an increase in material porosity, which facilitates tissue ingrowth. The partial degradation of BCP allows the maintenance of sufficient mechanical properties of the implanted biomaterial until new bone is formed. 

  

implantation in brain defects where no inflammation was observed (Deguchi et al., 2006). Moreover, in vitro biocompatibility of chitosan–GPTMS hybrids, with or without calcium addition, with 90% porosity and 100 μm pores, prepared by freeze-drying methods, were proven in cell cultures (Shirosaki et al., 2008). Taken together, these data point to organic–inorganic hybrids as promising candidates to be used as scaffolds in tissue engineering.

 Summary and Future Perspectives Ceramics are widely used for bone repair and bone regeneration in orthopedics and dentistry. This chapter described the key role played by the interactions of these materials with the biological environment in bioceramic efficiency when implanted. Specifically, ceramic degradation behavior allows classification of bioceramics into three categories. Inert (or almost inert) ceramics have no tendency to degrade in vivo. Resorbable ceramics, on the contrary, are degradable to various degrees. The degradation products of these ceramics are directly phagocytosed by inflammatory cells, osteoclasts, or even mesenchymal cells of the host, or chemically disintegrated (i.e., by local cell-mediated pH

changes) after implantation. A third category of bioceramics, usually referred to as bioactive ceramics, shows a surface affinity for bone binding. Inert and second-generation (bioactive and resorbable) ceramics are currently used as implants and other medical devices for bone repair. The degradation rate of these implanted bioceramics should match or at least approximate the osteogenic rate. In addition, resorbable and bioactive bioceramics, obtained as porous forms, are used for manufacturing scaffolds susceptible to hosting osteogenic cells and molecules for bone tissue-engineering applications. Incorporation of small peptides (osteostatin) or inorganic compounds (i.e., Sr, Cu, Zn) influencing bone growth and function into these ceramics appears to be an attractive approach in this context. It is obvious that despite extensive research carried out in the field, the perfect bioceramic implant is still not available. In this regard, third-generation bioceramics, which refer to those actively driving tissue regeneration and displaying optimal controlled degradability, are currently under investigation. Likewise, bioceramic nanoparticles might be envisioned as easily functionalized materials for intravenous injection or integrated into polymeric scaffolds to specifically target damaged tissues.

968 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

References Afzal, A., 2014. Implantable zirconia bioceramics for bone repair and replacement: a chronological review. Materials Express 4, 1–12. Akagawa, Y., Ichikawa, Y., Nikai, H., Tsuru, H., 1993. Interface histology of unloaded and early loaded partially-stabilized zirconia endosseous implant in initial bone healing. J. Prosthet. Dent. 69, 599–604. Allen, M., Law, F., Rushton, N., 1994. The effects of diamond-like carbon coatings on macrophages, fibroblasts and osteoblast-like cells in vitro. Clin. Mater. 17, 1–10. Arcos, D., del Real, R.P., Vallet-Regí, M., 2003. Biphasic materials for bone grafting and hyperthermia treatment of cancer. J. Biomed. Mater. Res. 65A, 71–78. Arias, J.L., Mayor, M.B., Pou, J., Leng, Y., León, B., Pérez-Amor, M., 2003. Micro- and nano-testing of calcium phosphate coatings produced by pulsed laser deposition. Biomaterials 24, 3403–3408. Black, J., Hastings, G. (Eds.), 1998. Handbook of Biomaterial Properties. Chapman & Hall, New York, NY, pp. 114–124. Bloomfield, P., Wheatley, D.J., Prescott, R.J., Miller, H.C., 1991. Twelve-year comparison of a Bjork–Shiley mechanical heart valve with porcine bioprostheses. N. Engl. J. Med. 324, 573–579. Boccardi, E., Ciraldo, F.E., Boccaccini, A.R., 2017. Bioactive glassceramic scaffolds: processing and properties. MRS Bull. 42, 226– 232. Bohner, M., 2000. Calcium orthophosphates in medicine: from ceramics to calcium phosphate cements. Injury 31 (S-4), 37–47. Boutin, P., 1971. Alumina and its use in surgery of the hip. Presse Med. 79, 639–640. Brinker, C.J., Lu, Y., Sellinger, A., Fan, H., 1999. Evaporation‐ induced self‐assembly: nanostructures made easy. Adv. Mater. 11, 579–601. Cabañas, M.V., Vallet-Regí, M., 2003. Calcium phosphate coatings deposited by aerosol chemical vapour deposition. J. Mater. Chem. 13, 1104–1107. Cabañas, M.V., Rodríguez-Lorenzo, L.M., Vallet-Regi, M., 2002. Setting behavior and in  vitro bioactivity of hydroxyapatite/calcium sulphate cements. Chem. Mater. 14, 3550–3555. Chen, Q., Miyaji, F., Kokubo, T., Nakamura, T., 2000. Bioactivity and mechanical properties of PDMS-modified CaO-SiO2-TiO2 hybrids prepared by sol-gel process. J. Biomed. Mater. Res. 51, 605–611. Chevalier, J., 2006. What future for zirconia as a biomaterial? Biomaterials 27 (4), 535–543. Christel, P., Meunier, A., Dorlot, J.M., Crolet, J.M., Witvoet, J., Sedel, L., Boutin, P., 1988. Biomechanical compatibility and design of ceramic implants for orthopaedic surgery. Bioceramics: material characteristics versus in vivo behaviour. Ann. NY Acad. Sci. 523, 234–256. Colilla, M., Salinas, A.J., Vallet-Regi, M., 2006. Amino-polysiloxane hybrid materials for bone reconstruction. Chem. Mater. 18, 5676–5683. Coradin, T., Bah, S., Livage, J., 2004. Gelatine/silicate interactions: from nanoparticles to composite gels. Colloids Surf. B 35, 53–58. Daculsi, G., 1998. Biphasic calcium phosphate concept applied to artificial bone, implant coating and injectable bone substitute. Biomaterials 19, 1473–1478. Daculsi, G., Fellah, B.H., Miramond, T., 2014. The essential role of calcium phosphate bioceramics in bone regeneration. In: BenNissan, B. (Ed.). Springer Series in Biomaterials Science and Engineering, vol. 2. Springer, Berlin, Heidelberg.

Dauskardt, R.H., Ritchie, R.O., 1993. Pyrolytic carbon coatings. In: Hench, L.L., Wilson, J. (Eds.), An Introduction to Bioceramics. World Scientific, Singapore, pp. 261–280. De Groot, K., 1991. Bioceramics consisting of calcium phosphate salts. Biomaterials 1, 47–50. Deguchi, K., Tsuru, K., Hayashi, T., Takaishi, M., Nagahara, M., Nagotani, S., Sehara, Y., Jin, G., Zhang, H., Hayakawa, S., Shoji, M., Miyazaki, M., Osaka, A., Huh, N.-H., Abe, K., 2006. Implantation of a new porous gelatin-siloxane hybrid into a brain lesion as a potential scaffold for tissue regeneration. J. Cereb. Blood Flow Metab. 26, 1263–1273. Del Real, R.P., Arcos, D., Vallet-Regi, M., 2002. Implantable magnetic glass-ceramic based on (Fe, Ca) SiO3 solid solutions. Chem. Mater. 14, 64–70. Doi, Y., Shibutani, T., Moriwaki, Y., Kajimoto, T., Iwayama, Y.J., 1998. Sintered carbonate apatites as bioresorbable bone substitutes. J. Biomed. Mater. Res. 39, 603–610. Dorozhkin, S.V., 2010. Bioceramics of calcium orthophosphates. Biomaterials 31, 1465–1485. Dorozhkin, S.V., Epple, M., 2002. Biological and medical significance of calcium phosphates. Angew. Chem. Int. Ed. 41, 3130–3146. Du, R.L., Chang, J., 2004. Preparation and characterization of Zn and Mg doped bioactive glasses. J. Inorg. Mater. 19, 1353–1358. Elias, C.N., Lima, J.H.C., Valiev, R., Meyers, M.A., 2008. Biomedical applications of titanium and its alloys. JOM J. Miner. Metals Mater. Soc. 60 (3), 46–49. Elliott, J.C., 1994. Structure and Chemistry of the Apatites and Other Calcium Orthophosphates. Elsevier, London. Fan, H., Wang, L., Zhao, K., Li, N., Shi, Z., Ge, Z., Jin, Z., 2010. Fabrication, mechanical properties, and biocompatibility of graphene-reinforced chitosan composites. Biomacromolecules 11, 2345–2351. Frieb, W., Warner, J., 2002. Biomedical applications. In: Schuth, F., Sing, K.S.W., Weitkamp, J. (Eds.), Handbook of Porous Solids. Wiley-VCH, Weinheim, pp. 2923–2970. Garcia, A., Cicuendez, M., Izquierdo-Barba, I., Arcos, D., ValletRegi, M., 2009. Essential role of calcium phosphate heterogeneities in 2D-hexagonal and 3D-cubic SiO2-CaO-P2O5 mesoporous bioactive glasses. Chem. Mater. 21, 5474–5484. Garcia-Sanz, F.J., Mayor, M.B., Arias, J.L., Pou, J., Leon, B., PerezAmor, M., 1997. Hydroxyapatite coatings: a comparative study between plasma-spray and pulsed laser deposition techniques. J. Mater. Sci. Mater. Med. 8, 861–865. Gibson, I.R., Best, S.M., Bonfield, W., 1999. Chemical characterization of silicon substituted hydroxyapatite. J. Biomed. Mater. Res. 44, 422–428. Gil-Albarova, J., Salinas, A.J., Bueno-Lozano, A.L., Román, J., Aldini-Nicolo, N., García-Barea, A., Giavaresi, G., Fini, M., Giardini, R., Vallet-Regí, M., 2005. The in vivo behaviour of a sol-gel glass and a glass-ceramic during critical diaphyseal bone defects healing. Biomaterials 26, 4374–4382. Ginebra, M.P., Driessens, F.C.M., Planell, J.A., 2004. Effect of the particle size on the micro and nanostructural features of a calcium phosphate cement: a kinetic analysis. Biomaterials 25, 3453–3462. González, B., Colilla, M., Vallet-Regí, M., 2008. Time-delayed release of bioencapsulates: a novel controlled delivery concept for bone implant technology. Chem. Mater. 20, 4826–4834. Gregoire, M., Orly, I., Menanteau, J., 1990. The influence of calcium phosphate biomaterials on human bone cell activities. An in vitro approach. J. Biomed. Mater. Res. 24, 165–177.

CHAPTER 2.4.4   Degradative Effects of the Biological Environment on Ceramic Biomaterials

Grover, L.M., Knowles, J.C., Fleming, G.J.P., Barralet, J.E., 2003. In vitro ageing of brushite calcium phosphate cement. Biomaterials 24, 4133–4141. Habraken, W., Habibovic, P., Epple, M., Bohner, M., 2016. Calcium phosphates in biomedical applications: materials for the future? Mater. Today 19, 69–87. Hench, L.,L., 1991. Bioceramics: from concept to clinic. J. Am. Ceram. Soc. 74, 1487–1510. Hench., L.L., 1994. Bioactive ceramics: theory and clinical applications. In: Anderson, Ö.H., Happonen, R.-P., Yli-Urpo, A. (Eds.), Bioceramics 7. Butterworth-Heinemann Ltd, Oxford, pp. 3–14. Hench, L.L., 1998. Bioceramics. J. Am. Ceram. Soc. 81, 1705–1728. Hench, L.L., 2016. Bioglass: 10 milestones from concept to commerce. J. Non-cryst. Solids 432, 2–8. Hench, L.L., Anderson, O., 1993. Bioactive glasses. In: Hench, L.L., Wilson, J. (Eds.), An Introduction to Bioceramics. World Scientific, Singapore, pp. 41–62. Hench, L.L., Polak, J.M., 2002. Third-generation biomedical materials. Science 295, 1014–1017. Hench, L.L., Wilson, J., 1984. Surface-active biomaterials. Science 226, 630–636. Hench, L.,L., Splinter, R.J., Allen, W.C., Greenlee, T.K., 1971. Bonding mechanisms at the interface of ceramic prosthetic materials. J. Biomed. Mater. Res. 2 117–4. Heymann, D., Pradal, G., Benahmed, M., 1999. Cellular mechanisms of calcium phosphate ceramic degradation. Histol. Histopathol. 14, 871–877. Hijon, N., Manzano, M., Salinas, A.J., Vallet-Regí, M., 2005. Bioactive CaO-SiO2-PDMS coatings onto Ti6Al4V substrates. Chem. Mater. 17, 1591–1596. Hijon, N., Cabañas, M.V., Izquierdo-Barba, I., Garcia, M.A., ValletRegi, M., 2006. Nanocrystalline bioactive apatite coatings. Solid State Sci. 8, 685–689. Hoppe, A., Güldal, N.S., Boccaccini, A.R., 2011. A review of the biological response to ionic dissolution products from bioactive glasses and glass-ceramics. Biomaterials 32, 2757–2774. Hu, W., Peng, C., Luo, W., Lv, M., Li, X., Li, D., Huang, Q., Fan, C., 2010. Graphene-based antibacterial paper. ACS Nano 4, 4317–4323. Izquierdo-Barba, I., Asenjo, A., Esquivias, L., Vallet-Regí, M., 2003. SiO2-CaO vitreous films deposited onto Ti6Al4V Substrates. Eur. J. Inorg. Chem. 1608–1613. Izquierdo-Barba, I., Conde, F., Olmo, N., Lizarbe, M.A., García, M.A., Vallet-Regí, M., 2006. Vitreous SiO2–CaO coatings onto Ti6Al4V alloys: reactivity in a cellular solution vs osteoblast cell culture. Acta Biomater. 2, 445–455. Izquierdo-Barba, I., Salinas, A.J., Vallet-Regi, M., 2013. Bioactive glasses: from macro to nano. Int. J. Appl. Glass Sci. 4 149–16. Jha, L.J., Best, S.M., Knowles, J.C., Rehman, I., Santos, J.D., Bonfield, W., 1997. Preparation and characterization of fluoride-substituted apatites. J. Mater. Sci. Mater. Med. 8, 185–191. Kannan, S., Balamurugan, A., Rajeswari, S., 2002. Development of calcium phosphate coatings on type 316L SS and their in-vitro response. Trends Biomater. Artif. Organs 16, 8–11. Kargozar, S., Baino, F., Hamzehlou, S., Hill, R.G., Mozafari, M., 2018. Bioactive glasses: sprouting angiogenesis in tissue engineering. Trends Biotechnol. 36, 430–444. Kawahara, H., 1981. Implant biomaterial and ceramics. Orthopaedic Ceramic Implants 1, 1–10. Khorasani, A.M., Goldberg, M., Doeven, E.H., Littlefair, G., 2015. Titanium in biomedical applications—properties and fabrication: a review. J. Biomater. Tissue Eng. 5 (8), 593–619.

969

Kivrak, N., Tas, A.C., 1998. Synthesis of calcium hydroxyapatitetricalcium phosphate (HA-TCP) composite bioceramic powders and their sintering behavior. J. Am. Ceram. Soc. 82, 2245–2252. Koch, B., Wolke, J.G., de Groot, K., 1990. X-ray diffraction studies on plasma-sprayed calcium phosphate-coated implants’. J. Biomed. Mater. Res. 24, 655–667. Kokubo, T., 1998. Apatite formation on surfaces of ceramics, metals and polymers in body environment. Acta Mater. 46, 2519–2527. Kokubo, T., Kushitani, H., Sakka, S., Kitsugi, T., Yamamuro, T., 1990. Solutions able to reproduce in  vivo surface-structure changes in bioactive glass-ceramic A-W. J. Biomed. Mater. Res. 24, 721–734. Kokubo, T., Kim, H.M., Kawashita, M., 2003. Novel bioactive materials with different mechanical properties. Biomaterials 24, 2161–2175. Kumta, P.N., 2006. In: Guelcher, S.A., Hollinger, J. (Eds.), Ceramic Biomaterials in: An Introduction to Biomaterials. CRC Taylor & Francis, Boca Raton, pp. 311–340. Kurtz, S.M., Kocagöz, S., Arnholt, C., Huet, R., Ueno, M., Walter, W.L., 2014. Advances in zirconia toughened alumina biomaterials for total joint replacement. J. Mech. Behav. Biomed. Mater. 31, 107–116. Langer, R., Vacanti, J.P., 1993. Tissue engineering. Science 260, 920– 926. LeGeros, R.Z., 1991. Calcium phosphates in oral biology and medicine. In: Myers, H.M. (Ed.). Monographs in Oral Science, vol. 15. S. Karger, Basel. LeGeros, R.Z., 1993. Biodegradation and bioresorption of calcium phosphate ceramics. Clin. Mater. 14, 65–88. Lew, K.S., Othman, R., Ishikawa, K., Yeoh, F.Y., 2012. Macroporous bioceramics: a remarkable material for bone regeneration. J. Biomater. Appl. 27, 345–358. Li, J., Hastings, G.W., 1998. Oxide bioceramics: inert ceramic materials in medicine and dentistry. In: Black, J., Hastings, G. (Eds.), Handbook of Biomaterial Properties. Chapman Hall, Oxford, pp. 340–354. Li, R., Clark, A.E., Hench, L.L., 1991. An investigation of bioactive glass powders by sol-gel processing. J. Appl. Biomater. 2, 231–239. Liu, X., Chu, P.K., Ding, C., 2004. Surface modification of titanium, titanium alloys, and related materials for biomedical applications. Mater. Sci. Eng. R Rep. 47 (3), 49–121. Livingston, T., Ducheyne, P., Garino, J., 2002. In vivo evaluation of a bioactive scaffold for bone tissue engineering. J. Biomed. Mater. Res. 62, 1–13. Manzano, M., Arcos, D., Delgado, M.R., Ruiz, E., Gil, F.J., ValletRegi, M., 2006a. Bioactive star gels. Chem. Mater. 18, 5696– 5703. Manzano, M., Salinas, A.J., Vallet-Regí, M., 2006b. P-containing ormosils for bone reconstruction. Prog. Solid State Chem. 34, 267–277. Martin, A.I., Salinas, A.J., Vallet-Regí, M., 2005. Bioactive and degradable organic-inorganic hybrids. J. Eur. Ceram. Soc. 25, 3533–3538. Martínez, A., Izquierdo-Barba, I., Vallet-Regí, M., 2000. ´Bioactivity of a CaO-SiO2 binary glasses system’. Chem. Mater. 12, 3080– 3088. Michalczyk, M.J., Sharp, K.G., 1995. US Patent 5378790. Miyata, N., Fuke, K., Chen, Q., Kawashita, M., Kokubo, T., Nakamura, T., 2004. Apatite-forming ability and mechanical properties of PTMO-modified CaO–SiO2–TiO2 hybrids derived from sol– gel processing. Biomaterials 25, 1–7. Morrell, R., Danzer, R., Milosev, I., Trebse, R., 2012. An assessment of in vivo failures of alumina ceramic total hip joint replacements. J. Eur. Ceram. Soc. 32, 3073–3084.

970 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Narasaraju, T.S.B., Phebe, D.E., 1996. Some physico-chemical aspects of hydroxyapatite. J. Mater. Sci. 31, 1–21. Nilsson, M., Fernández, E., Sarda, S., Lidgren, L., Planell, J.A., 2002. Characterization of a novel calcium phosphate/sulphate bone cement. J. Biomed. Mater. Res. 61, 600–607. Oki, A., Parveen, B., Hossain, S., Adeniji, A., Donahue, H., 2004. Preparation and in  vitro bioactivity of zinc containing sol–gelderived bioglass materials. J. Biomed. Mater. Res. 69A, 216–221. Olmo, N., Martín, A.I., Salinas, A.J., Turnay, J., Vallet-Regí, M., Lizarbe, M.A., 2003. Bioactive sol-gel glasses with and without a hydroxycarbonate apatite layer as substrates for osteoblast cell adhesion and proliferation. Biomaterials 24, 3383–3393. Oonishi, H., Okabe, N., Hamaguchi, T., Nabeshima, T., 1981. Cementless alumina ceramic total knee prosthesis. Orthopaedic Ceramic Implants 1, 11–18. Ostomel, T.A., Shi, Q.H., Tsung, C.K., Liang, H.J., Stucky, G.D., 2006. Spherical bioactive glass with enhanced rates of hydroxyapatite deposition and hemostatic activity. Small 2, 1261–1265. Otsuka, M., Matsuda, Y., Suwa, Y., Fox, J.L., Higuchi, W., 1995. Effect of particle size of metastable calcium phosphates on mechanical strength of a novel self-setting bioactive calcium phosphate cement. J. Biomed. Mater. Res. 29, 25–32. Padilla, S., Roman, J., Carenas, A., Vallet-Regí, M., 2005. Influence of the phosphorus content on the bioactivity in sol-gel glass ceramics. Biomaterials 26, 475–483. Park, Y.S., Yi, K.Y., Lee, I.S., Han, C.H., Jung, Y.C., 2005. The effects of ion beam-assisted deposition of hydroxyapatite on the gritblasted surface of endosseous implants in rabbit tibiae’. Int. J. Oral Maxillofac. Implant. 20, 31–38. Pechini, M.P., July 11, 1967. U. S. Patent 3 330,697. Peltola, T., Jokinen, M., Rahiala, H., Levänen, E., Rosenhold, J.B., Kangasniemi, I., Yli-Urpo, A., 1999. Calcium phosphate formation on porous sol-gel-derived SiO2 and CaO-P2O5-SiO2 substrates in vitro. J. Biomed. Mater. Res. 44, 12–21. Peña, J., Vallet-Regi, M., 2003. Hydroxyapatite, tricalcium phosphate and biphasic materials prepared by a liquid mix technique. J. Eur. Ceram. Soc. 23, 1687–1696. Pereira, M.M., Clark, A.E., Hench, L.L., 1994. Calcium-phosphate formation on sol-gel-derived bioactive glasses in-vitro. J. Biomed. Mater. Res. 28, 693–698. Pereira, A.P.V., Vasconcelos, W.L., Oréfice, R.K., 2000. Novel multicomponent silicate-poly(vinyl alcohol) hybrids with controlled reactivity. J. Non-cryst. Solids 273, 180–185. Pérez-Pariente, J., Balas, F., Román, J., Salinas, A.J., Vallet-Regi, 1999. Influence of composition and surface characteristics on the in  vitro bioactivity of SiO2-CaO-P2O5-MgO sol-gel glasses. J. Biomed. Mater. Res. 47, 170–175. Perez, P., Sanchez-Salcedo, S., Lozano, D., Heras, C., Esbrit, P., ValletRegí, M., Salinas, A.J., 2018. Osteogenic effect of ZnO-mesoporous glasses loaded with osteostatin. Nanomaterials 8 (8). Pietrzak, W.S., Ronk, R., 2000. Calcium sulphate bone void filler: a review and a look ahead. J. Craniofac. Surg. 11, 327–334. Polack, F.M., Heimke, G., 1980. Ceramic kerato-prostheses. Ophtalmology 87, 693–698. Rae, T., 1986. The macrophage response to implant materials-with special reference to those used in orthopedics. CRC Crit. Rev. Biocompat. 2, 97–126. Ramila, A., Padilla, S., Muñoz, B., Vallet-Regí, M., 2002. A new hydroxyapatite/glass biphasic material: in vitro bioactivity’. Chem. Mater. 14, 2439–2943. Ratner, B.D., Castner, D.G., 2002. Biomedical surface science: foundations to frontiers. Surf. Sci. 500, 28–60.

Ren, L., Tsuru, K., Hayakawa, S., Osaka, A., 2002. Novel approach to fabricate porous gelatin-siloxane hybrids for bone tissue engineering. Biomaterials 23, 4765–4773. Rhee, S.H., 2004. Bone-like apatite-forming ability and mechanical properties of poly(ε -caprolactone)/silica hybrid as a function of poly(ε-caprolactone) content. Biomaterials 25, 1167–1175. Rhee, S., Choi, J., 2002. Preparation of a bioactive poly(methyl methacrylate)/silica nanocomposite. J. Am. Ceram. Soc. 85, 1318–1320. Rodrigo, A., Valles, G., Saldaña, L., Rodrıguez, M., Martinez, M.E., Munuera, L., Vilaboa, N., 2006. Alumina particles influence the interactions of cocultured osteoblasts and macrophages. J. Orthop. Res. 24, 46–54. Roman, J., Padilla, S., Vallet-Regí, M., 2003. Sol-gel glasses as precursors of bioactive glass-ceramics. Chem. Mater. 15, 798–806. Román, J., Cabañas, M.V., Peña, J., Doadrio, J.C., Vallet-Regi, M., 2008. An optimised ß-tricalcium phosphate and agarose scaffold fabrication technique. J. Biomed. Mater. Res. A 84A, 99–107. Ruiz-Hernández, E., López-Noriega, A., Arcos, D., Vallet-Regí, M., 2008. Mesoporous magnetic microspheres for drug targeting. Solid State Sci. 10, 421–426. Salinas, A.J., Vallet-Regí, M., 2017. Sol-gel silica-based biomaterials and bone tissue regeneration. In: Klein, L., Aparicio, M., Jitianu, A. (Eds.), Handbook of Sol-Gel Science and Technology, second ed. The Springer, pp. 3597–3618. Salinas, A.J., Vallet-Regí, M., Izquierdo-Barba, I., 2001. Biomimetic apatite deposition on calcium silicate gel glasses. J. Sol. Gel Sci. Technol. 21, 13–25. Salinas, A.J., Martín, A.I., Vallet-Regi, M., 2002. Bioactivity of three CaO-P2O5-SiO2 sol-gel glasses. J. Biomed. Mater. Res. 61, 524–532. Salinas, A.J., Merino, J.M., Gil, F.J., Babonneau, F., Vallet-Regí, M., 2007. Microstructure and macroscopic properties of CaO-SiO2-PDMS hybrids for use in implants. J. Biomed. Mater. Res. B 81B, 274–282. Salinas, A.J., Shruti, S., Malavasi, G., Menabue, L., Vallet-Regi, M., 2011. Substitutions of cerium, gallium and zinc in ordered mesoporous bioactive glasses. Acta Biomater. 7, 3452–3458. Salinas, A.J., Esbrit, P., Vallet-Regí, M., 2013. A tissue engineering approach based on the use of bioceramics for bone repair. Biomater. Sci. 1, 40–51. Salinas, A.J., Vallet-Regi, M., Heikkilä, J., 2018. Use of bioactive glasses as bone substitutes in orthopedics and traumatology. In: Bioactive Glasses, Materials, Properties and Applications. Woodhead Publishing Series in Biomaterials, pp. 337–364. Sánchez-Salcedo, S., Nieto, A., Vallet-Regí, M., 2008a. Hydroxyapatite/ß-tricalcium phosphate/agarose macroporous scaffolds for bone tissue engineering. Chem. Eng. J. 137, 62–71. Sanchez-Salcedo, S., Werner, J., Vallet-Regi, M., 2008b. Hierarchical pore structure of calcium phosphate scaffolds by combination of the gel casting and multiple tape casting methods. Acta Biomater. 4, 913–922. Sanchez-Salcedo, S., Malavasi, G., Salinas, A.J., Lusvardi, G., Rigamonti, L., Menabue, L., Vallet-Regi, M., 2018. Highly-bioreactive silica-based mesoporous bioactive glasses enriched with gallium(III). Materials 11, 367. Sanders, D., Hench, L., 1973. Mechanisms of glass corrosion. J. Am. Ceram. Soc. 56, 373–377. Sandhaus, S., 1967. Bone Implants and Drills and Taps for Bone Surgery. British Patent 1083769. Schiraldi, C., D’Agostino, A., Oliva, A., Flamma, F., De la Rosa, A., Apicella, A., Aversa, R., De la Rosa, M., 2004. Development of hybrid materials based on hydroxyethylmethacrylate as supports for improving cell adhesion and proliferation. Biomaterials 25, 3645–3653.

CHAPTER 2.4.4   Degradative Effects of the Biological Environment on Ceramic Biomaterials

Schulte, W., 1990. The Frialit Tuebingen implant system. In: Heimke, G. (Ed.), Osseointegrated Implants Volume II. Implants in Oral and ENT Surgery. CRC Press, Boca Raton, pp. 1–34. Serrano, M.C., Portolés, M.T., Pagani, R., Sáez de Guinoa, J., RuízHernández, E., Arcos, D., Vallet-Regí, M., 2008. In vitro positive biocompatibility evaluation of glass-glass ceramic thermoseeds for hyperthermic treatment of bone tumors. Tissue Eng. A 14, 617–627. Sharp, K.G., 1998. Inorganic/organic hybrid materials. Adv. Mater. 10, 1243–1248. Shirosaki, Y., Tsuru, K., Hayakawa, S., Osaka, A., 2008. Biodegradable chitosan-silicate porous hybrids for drug delivery. Key Eng. Mater. 361–363, 1219–1222 (Switzerland, Trans Tech Publications). Smith, L., 1963. Ceramic plastic material as a bone substitute. Arch. Surg. 87, 653–661. Sun, L.M., Berndt, C.C., Gross, K.A., Kucuk, A., 2001. Material fundamentals and clinical performance of plasma-sprayed hydroxyapatite coatings: a review. J. Biomed. Mater. Res. 58, 570–592. Tamimi, F., Sheikh, Z., Barralet, J., 2012. Dicalcium phosphate cements: brushite and monetite. Acta Biomater. 8, 474–487. Tancred, D.C., McCormack, B.A.O., Carr, A.,J., 1998. A synthetic bone implant macroscopically identical to cancellous bone. Biomaterials 19, 2303–2311. Tsuru, K., Hayakawa, S., Osaka, A., 2008. Cell proliferation and tissue compatibility of organic-inorganic hybrid materials. Key Eng. Mater. 377, 167–180 Progress in Bioceramics, Vallet-Regi, M. (ed), Switzerland, Trans Tech Publications. Vallet-Regí, M., 2001. Ceramics for medical applications. J. Chem. Soc., Dalton Trans. 97–108. Vallet-Regí, M., 2006. Revisiting ceramics for medical applications. Dalton Trans. 5211–5220. Vallet-Regi, M., Arcos, D., 2005. Silicon substituted hydroxyapatites. A method to upgrade calcium phosphate based implants. J. Mater. Chem. 15, 1509–1516. Vallet-Regí, M., Arcos, D., 2006. Nanostructured hybrid materials for bone tissue regeneration. Curr. Nanosci. 2, 179–189. Vallet-Regí, M., Arcos-Navarrete, D., 2008. Biomimetic Nanoceramics in Clinical Use: From Materials to Applications. Royal Society of Chemistry, Cambridge. Vallet-Regí, M., González-Calbet, J., 2004. Calcium phosphates in the substitution of bone tissue. Prog. Solid State Chem. 32, 1–31. Vallet-Regí, M., Salinas, A.J., 2017. Mesoporous bioactive glasses in tissue engineering and drug delivery. In: Boccaccini, A.R., Brauer, D.S., Hupa, L. (Eds.), Bioactive Glasses: Fundamentals, Technology and Applications. The Royal Society of Chemistry, London, UK, pp. 393–419. Vallet-Regí, M., Salinas, A.J., 2018. Role of the short distance order in glass reactivity. Materials 11, 415. Vallet-Regi, M., Gutiérrez-Ríos, M.T., Alonso, M.P., de Frutos, M.I., Nicolopoulos, S., 1994. Hydroxyapatite particles synthesized. by pyrolysis of an aerosol. J. Solid State Chem. 112, 58–64. Vallet-Regi, M., Rodríguez-Lorenzo, L.M., Salinas, A.J., 1997. Synthesis and characterisation of calcium deficient apatite. Solid State Ion. 101–103, 1279–1285.

971

Vallet-Regí, M., Romero, A.M., Ragel, C.V., LeGeros, R.Z., 1999. XRD, SEM-EDS and FTIR studies of in  vitro growth of an apatite-like layer on sol-gel glasses. J. Biomed. Mater. Res. 44, 416–421. Vallet-Regi, M., Rámila, A., del Real, R.,P., Pérez-Pariente, J., 2001. A new property of MCM-41: drug delivery system. Chem. Mater. 13, 308–311. Vallet-Regí, M., Ragel, C.V., Salinas, A.J., 2003a. Glasses with medical applications. Eur. J. Inorg. Chem. (6) 1029–1042. Vallet-Regi, M., Izquierdo-Barba, I., Gil, F.J., 2003b. Localized corrosion of 316L Stainless Steel with SiO2–CaO films obtained by means of sol-gel treatment. J. Biomed. Mater. Res. 67A, 674–678. Vallet-Regí, M., Román, J., Padilla, S., Doadrio, J.C., Gil, F.J., 2005a. Bioactivity and mechanical properties of SiO2-CaO-P2O5 glassceramics. J. Mater. Chem. 15, 1353–1359. Vallet-Regí, M., Salinas, A.J., Ramírez-Castellanos, J., González-Calbet, J.M., 2005b. Nanostructure of bioactive sol-gel glasses and organic inorganic hybrids. Chem. Mater. 17, 1874–1879. Vallet-Regí, M., Ruiz-González, M.L., Izquierdo-Barba, I., GonzálezCalbet, J.M., 2006. Revisiting silica based ordered mesoporous materials: medical applications. J. Mater. Chem. 16, 26–31. Van der Stok, J., Lozano, D., Chai, Y.C., Yavari, S.A., Bastidas Coral, A.P., Verhaar, J.A.N., Gomez-Barrena, E., Schrooten, J., Jahr, H., Zadpoor, A.A., Esbrit, P., Weinans, H., 2015. Tissue Eng. A 21, 1495–1506. Wang, G.C., Lu, Z.F., Zreiqat, H., 2014. Bioceramics for skeletal bone regeneration. In: Bone Substitute Biomaterials. Woodhead Publishing Series in Biomaterials, pp. 180–186. 187e-190e, 187216. Williams, D.F., 2008. On the mechanisms of biocompatibility. Biomaterials 29, 2041–2953. Wolke, J.G.C., de Groot, K., Jansen, J.A., 1998. Subperiosteal implantation of various RF magnetron sputtered Ca-P coatings in goats. J. Biomed. Mater. Res. 43, 270–276. Xynos, I.D., et al., 2001. Gene-expression profiling of human osteoblasts following treatment with the ionic products of Bioglass® 45S5 dissolution. J. Biomed. Mater. Res. 55 (2), 151–157. Yabuta, T., Bescher, E.P., Mackenzie, J.D., Tsuru, K., Hayakawa, S., Osaka, A., 2003. Synthesis of PDMS-based porous materials for biomedical applications. J. Sol. Gel Sci. Technol. 26, 1219–1222. Yan, X., Yu, C., Zhou, H., Tang, J., Zhao, D., 2004. Highly ordered mesoporous bioactive glasses with superior in vitro bone‐forming bioactivities. Angew. Chem. Int. Ed. 43, 5980–5984. Yuan, H.P., Kurashina, K., de Bruijn, J.D., Li, Y.B., de Groot, K., Zhang, X.D., 1999. A preliminary study on osteoinduction of two kinds of calcium phosphate ceramics. Biomaterials 20, 1799– 1806. Zhang, F., Chang, J., Lu, J., Lin, K., Ning, C., 2007. Bioinspired structure of bioceramics for bone regeneration in load-bearing sites. Acta Biomater. 3, 896–904. Zhu, L., Chang, D.W., Dai, L., Hong, Y., 2007. DNA damage induced by multiwalled carbon nanotubes in mouse embryonic stem cells. Nano Lett. 7, 3592–3597.

2.4.5

Pathological Calcification of Biomaterials FREDERICK J. SCHOEN 1 , ROBERT J. LEVY 2 , HOBEY TAM 3 , NAREN VYAVAHARE 3 1Department

of Pathology, Brigham and Women’s Hospital, and Harvard Medical School, Boston, MA, United States 2Department

of Pediatrics, The Childrens’ Hospital of Philadelphia, The Perelman School of Medicine at the University of Pennsylvania, Philadelphia, PA, United States 3Department

of Bioengineering, Rhodes Engineering Research Center, Clemson University, Clemson, SC, United States

C

alcification of biomaterial implants is an important pathologic process affecting a variety of tissuederived biomaterials, as well as synthetic polymers in various functional configurations. The pathophysiology has been partially characterized by some useful animal models; common features are the involvement of devitalized cells and cellular debris and often extracellular matrix proteins, and oxidative damage to the extracellular matrix and/or other reactive sites within the implant. Calcification of biomaterials often contributes to deleterious outcomes. Clinically useful preventive approaches based on modifying biomaterials appear to be promising in some contexts. The formation of nodular deposits of calcium phosphate or other calcium-containing compounds may occur on biomaterials and prosthetic devices used in the circulatory system and at other sites. This process, known as calcification or mineralization, has been encountered in association with both synthetic and biologically derived biomaterials in diverse experimental and clinical settings, including chemically treated tissue (bioprosthetic), homografts, and polymeric cardiac valve substitutes (Mitchell et al., 1998; Schoen and Levy, 1999, 2005; Claiborne et al., 2012), blood pumps used as cardiac assist devices (Schoen and Edwards, 2001), breast implants (Peters et al., 1998; Legrand et al., 2005), intrauterine contraceptive devices (Patai et al., 1998), urological stents (Vanderbrink et al., 2008), intraocular lenses (Neuhann et  al., 2008; Nakanome et  al., 2008; Rimmer et al., 2010; Rahimi et al., 2018), and scleral buckling materials (Brockhurst et al., 1993; Yu et al., 2005). Vascular access grafts for hemodialysis and synthetic vascular replacements composed of Dacron or expanded polytetrafluoroethylene

(e-PTFE) also calcify in some patients (Tomizawa et  al., 1998; Schlieper et  al., 2008). Importantly, calcification can lead to important clinical complications of implanted biomaterials and medical devices, examples of which are stiffening and/or tearing of tissue heart valve substitutes, encrustation with blockage of a urinary stent or clouding of intraocular lenses (Table 2.4.5.1). Deposition of mineral salts of calcium (as calcium phosphates, especially calcium hydroxyapatite) occurs normally in bones and teeth and is a critical determinant of their mechanical and biological properties (called physiologic mineralization). Mineralization of skeletal tissues is both controlled and restricted to specific anatomic sites. The mature mineral phase of bone is a poorly crystalline calcium phosphate known as calcium hydroxyapatite, which has the chemical formula Ca10(PO4)6(HO)2. Mineralization of some implant biomaterials is desirable for proper function, e.g., osteoinductive materials used for orthopedic or dental applications (Habibovic and de Groot, 2007), and materials used to engineer skeletal and dental tissues (Kretlow and Mikos, 2007; Bueno and Glowacki, 2009). However, severe consequences can occur if mineralization occurs in regions that do not normally calcify (pathologic or ectopic mineralization) (Kirsch, 2008). Because the biomaterials used in medical devices outside of the musculoskeletal and dental systems are not intended to calcify, calcification of these biomaterials is pathologic. The mature mineral phase of biomaterial-related and other forms of pathologic calcification is a poorly crystalline calcium phosphate that closely resembles the calcium hydroxyapatite present in bones and teeth. Indeed, as we 973

974 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

TABLE 2.4.5.1    Representative Prostheses and Devices With Clinical Consequences due to Biomaterials Calcification

Configuration

Biomaterial

Clinical Consequence

Cardiac valve prostheses

Glutaraldehyde-pretreated porcine aortic valve or bovine pericardium, and allograft aortic/ pulmonary valves

Valve obstruction or incompetency

Cardiac ventricular assist bladders

Polyurethane

Dysfunction by stiffening or cracking

Vascular grafts

Dacron® grafts and aortic allografts

Graft obstruction or stiffening

Soft contact lens

Hydrogels

Opacification

Intrauterine contraceptive devices

Silicone rubber, polyurethane or copper

Birth control failure by dysfunction or expulsion

Urinary prostheses

Silicone rubber or polyurethane

Incontinence and/or infection

Breast implants

Silicone

Implant contracture

will see later, biomaterials-related calcification shares many features with other conditions of pathologic calcification and physiologic mineralization. Pathologic calcification of natural structures is also a common feature of important disease processes; for example, in native arteries and heart valves, calcification occurs as an important feature of disease processes such as atherosclerosis and calcific aortic stenosis (Mitchell and Schoen, 2010; Schoen and Mitchell, 2010; Andrews et al., 2018; Pawade et al., 2015). Pathologic calcification is further classified as either dystrophic or metastatic, depending on its setting (Kumar et al., 2015). In dystrophic calcification the deposition of calcium salts (usually calcium phosphates) occurs in damaged or diseased tissues or in or related to biomaterials; moreover, dystrophic calcification usually occurs in the setting of normal systemic calcium metabolism (generally defined by a product of the serum levels of calcium and phosphorus within the physiologic range). In contrast, metastatic calcification comprises the deposition of calcium salts in otherwise normal tissues in individuals who have deranged mineral metabolism, usually with markedly elevated blood calcium or phosphorus levels. Conditions favoring dystrophic and metastatic calcification can act synergistically. Thus, the rate and extent of dystrophic mineralization within abnormal tissues are accelerated when calcium and/or phosphorus serum levels are high, for example, in kidney failure or calcium supplementation (Umana et al., 2003; Peacock, 2010), and in osteoporosis (Hofbauer et al., 2007; Hjortnaes et al., 2010). Moreover, the ability to form physiologic mineral (e.g., in bone) is regulated through adjustment of enhancing and inhibiting substances, many of which circulate in the blood (Weissen-Planz et al., 2008). In young individuals (especially into early adulthood) the balance appropriately favors bone formation; moreover, the biochemical environment that favors physiologic bone formation in children also promotes calcification of biomaterials (Bass and Chan, 2006; Peacock, 2010). Additionally, calcification of an implanted biomaterial can occur either within the material (intrinsic calcification)

or at the surface, usually associated with attached cells and proteins or in the implant fibrous capsule (extrinsic calcification). A notable instance of extrinsic calcification is that associated with prosthetic valve infection (also called prosthetic valve endocarditis) or thrombi.

The Spectrum of Pathologic Biomaterial and Medical Device Calcification Bioprosthetic Heart Valves Calcific degeneration of bioprosthetic heart valves composed of glutaraldehyde-pretreated heterograft materials, typically either porcine aortic valves or bovine pericardium (Fig. 2.4.5.1) is the most prevalent clinically significant and well-characterized dysfunction of a medical device due to biomaterials calcification (Schoen and Levy, 1999, 2005). The predominant pathologic process is intrinsic calcification of the valve cusps, largely initiated in the deep interstitial cells of the tissue from which the valve was fabricated and often involves collagen and elastin. Calcification leads to valve failure most commonly by causing tissue tears or tissue stiffening, and rarely by inducing distant emboli. Overall, approximately half of contemporary bioprostheses fail within 15–20 years. Calcification is more rapid and aggressive in younger recipients; to exemplify, the rate of failure of bioprostheses is approximately 10% in 10 years in elderly recipients but is nearly uniform in less than 4 years in most adolescent and preadolescent children (Chan et  al., 2006). In some young individuals with congenital cardiac defects or acquired aortic valve disease, human allograft/homograft aortic or pulmonary valves surrounded by a sleeve of the aorta or pulmonary artery, respectively, are used. Allograft valves are valves removed from a person who has died and transplanted to another individual; the tissue is usually cryopreserved (and thawed immediately prior to implantation), but not chemically cross-linked. Allograft vascular segments (without a valve) can be used to replace a large blood vessel.

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

(A)

(B)

975

(C)



Figure 2.4.5.1 Calcification of a pericardial bioprosthetic heart valve, implanted in a person for many years: (A) gross photograph, demonstrating marked thickening of cusps by nodular calcific deposits; (B) radiograph of another long-term valve, demonstrating predominant deposits at commissures; (C) photomicrograph demonstrating calcific nodule deeply embedded within cuspal tissue; von Kossa stain (calcium phosphates black) 100×.

Allograft vascular tissue, whether containing an aortic valve or not, can undergo severe calcification in the wall; calcification can lead to allograft valve dysfunction or deterioration (Mitchell et al., 1998).

2012). Most of the TAVRs currently used are treated with proprietary anticalcification treatments.

Transcatheter (or Percutaneous) Cardiac Valve Replacements

Calcification has also complicated the use of heart valves and artificial mitral valve chordae tendineae composed of polymers (e.g., polyurethane) (Fukunaga et al., 2010; Hilbert et  al., 1987; Schoen et  al., 1992a; Schoen and Levy, 1999; Wang et al., 2010) and the flexing bladder surfaces of blood pumps used as ventricular assist devices or total artificial hearts (Schoen and Edwards, 2001) (Fig. 2.4.5.2). Massive deposition of mineral leading to failure of cardiac assist devices has been noted in experimental animals, but only a lesser degree of calcification has been encountered following extended human implantation. Mineral deposits can result in deterioration of pump or valve performance through loss of bladder pliability or the initiation of tears. Blood pump calcification, regardless of the type of polyurethane used, generally predominates along the flexing margins of the diaphragm, emphasizing the important potentiating role of mechanical factors in this system (Coleman et al., 1981; Harasaki et al., 1987; Kantrowitz et al., 1995). Calcific deposits associated with a polymeric heart valve or blood pump components can occur either within the adherent layer of deposited thrombus, proteins, and cells on the blood-contacting surface (a form of extrinsic mineralization) or below the surface (a form of intrinsic calcification) (Joshi et al., 1996). In some cases, calcific deposits are associated with microscopic surface defects, originating either during bladder fabrication or resulting from cracking during the function. 

Transcatheter aortic valve replacement (TAVR, also known as transcatheter aortic valve implantation [TAVI]) has emerged over the past decade as a widely used minimally invasive therapeutic option for patients with symptomatic severe aortic stenosis who are ineligible for or at excessive risk for conventional surgical aortic valve replacement (Baumgartner et al., 2017). In this procedure, the diseased valve is not removed; instead, it is simply pushed out of the flow stream by the expanded replacement valve. Transcatheter valves are usually fabricated from bovine pericardium and mounted on a self-expanding or balloon-expandable stent crimped into and deployed via a catheter inserted through the femoral artery and threaded retrograde through the aorta to the heart, or across the apex of the ventricle into the heart. Although this procedure was initially preferred for people who have severe valve disease but are not suitable for openheart surgery, recent demonstration of favorable outcomes (relatively to open surgery) in intermediate- and low-risk patients suggests that the use of transcatheter valve replacement may be expanding. These valve replacements are also deployed after a failure of a bioprosthetic heart valve implant in a valve-in-valve procedure. The old replacement valve is left inside the patient, and the TAVR is deployed inside of the old valve, thereby pushing open the old valve and opening up a smaller but unblocked/functional valve with effective orifice area. The overall long-term durability of transcatheter replacement valves is yet uncertain; nevertheless, calcification has been observed in some cases (Zareian et al., 2019; Kataruka and Otto, 2018). Furthermore, studies have shown that crimping is associated with higher level of passive tissue calcification (Ramin et al., 2019; Munnelly et al.,

Polymeric Heart Valves and Blood Pump Bladders

Breast Implants Calcification of silicone-gel breast implant capsules occurs as discrete calcified plaques at the interface of the inner fibrous capsule with the implant surface (Peters et al.,1998; Gumus, 2009). Capsular calcification has also been encountered with breast implants in patients with silicone envelopes filled with saline. Calcification could interfere with effective

976 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

(A)

(C)

(D)

(B)



Figure 2.4.5.2 Calcification of an experimental polymeric (polyurethane) heart valve: (A) gross photograph of valve; (B) photomicrograph of calcified material at surface of polymer (polymer at bottom of photo); and calcification of the flexing bladder of a ventricular assist pump removed from a person after 257 days. (C) Gross photograph. Calcific masses are noted by arrows; (D) photomicrograph. (B) and (D) von Kossa stain (calcium phosphates black) 100×.

tumor detection and diagnosis, which could potentially delay treatment, particularly during monitoring for recurrence in patients who have breast implants following reconstructive surgery for breast cancer. In a study of breast implants removed predominantly for capsular contraction, 16% overall demonstrated calcific deposits, including 26% of implants inserted for 11–20 years, and all those inserted for more than 23 years (Peters and Smith, 1995). Another study demonstrated calcification associated with virtually all implants examined after more than 20 years (Legrand et al., 2005). Ivalon (polyvinyl alcohol) sponge prostheses, used quite extensively during the 1950s, were also frequently associated with calcification. In Japan, where augmentation mammoplasty was frequently performed using an injection of foreign material (liquid paraffin from approximately 1950 until 1964, and primarily liquid silicone injections thereafter), the incidence of calcification has been much higher. One study showed calcification in 45% of breast augmentations which were done by injection (Koide and Katayama, 1979). 

Intrauterine Contraceptive Devices Intrauterine contraceptive devices (IUDs) are composed of plastic or metal and placed in a woman’s uterus chronically to prevent implantation of a fertilized egg. Device dysfunction due to calcific deposits can be manifested as a contraceptive failure or device expulsion. For example, accumulation of calcific plaque could prevent the release of the active contraception-preventing agent—either ionic copper from

copper-containing IUDs or an active agent from hormonereleasing polymer IUD systems. Studies of explanted IUDs using transmission and electron microscopy coupled with X-ray microprobe analysis have shown that surface calcium deposition is virtually ubiquitous, but highly variable among patients (Khan and Wilkinson, 1985; Patai et al., 1998). 

Urinary Stents and Prostheses Mineral crusts form on the surfaces of urinary stents and nephrostomy tubes, which are used extensively in urology to alleviate a urinary obstruction or incontinence (Goldfarb et  al., 1989; Vanderbrink et  al., 2008). Observed in both male and female urethral implants and artificial ureters, this calcification can lead to obstruction and device failure. The mineral crust typically consists of either calcium oxalate or calcium phosphate mineral such as hydroxyapatite or struvite, ammonium- and magnesium-containing phosphate mineral derived from urine. There is some evidence that encrustation may both result from and predispose to bacterial infection. 

Intraocular and Soft Contact Lenses and Scleral Buckles Calcium phosphate deposits can opacify intraocular and soft contact lenses, typically composed of poly-2-hydroxyethyl-methacrylate polymers (pHEMA) (Bucher et  al., 1995; Nakanome et al., 2008; Neuhann et al., 2008; Rimmer et al., 2010; MacLean et al., 2015; Phogat et al., 2017).

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

977

TABLE 2.4.5.2   Methods for Assessing Calcification

Technique

Sample Preparation

Analytical Results

Gross examination

Gross specimen

Overall morphology

Radiographs

Gross specimen

Calcific deposit distribution

Microcomputed tomography

Gross specimen

Three-dimensional reconstruction of calcific deposit morphology, localization, and quantification

Light microscopy—von kossa or alizarin red stains

Formalin or glutaraldehyde fixed

Microscopic phosphate or calcium distribution, respectively

Transmission electron microscopy

Glutaraldehyde fixed

Mineral ultrastructure

Scanning electron microscopy with electron microprobe

Glutaraldehyde fixed

Elemental localization and quantitation

Electron energy loss spectroscopy

Glutaraldehyde fixed or rapidly frozen

Elemental localization and quantitation (high sensitivity)

Atomic absorption

Ash or acid hydrolysate

Bulk calcium

Colorimetric phosphate analysis

Ash or acid hydrolysate

Bulk phosphorus

X-ray diffraction

Powder

Nature of crystal phase

Infrared spectroscopy

Powder

Carbonate mineral phase

Morphologic Procedures

Chemical Procedures

Tear fluid provides the source of the calcium in the deposits found on HEMA contact lens, and calcification may be potentiated in patients with systemic and ocular conditions associated with elevated tear calcium levels (Klintworth et  al., 1977). Encircling scleral bands, used in surgery for retinal detachment, and composed of silicone or hydrogel materials, also can calcify (Lane et al., 2001). 

Assessment of Biomaterial Calcification Calcific deposits are investigated using morphologic and chemical techniques (Table 2.4.5.2). Morphologic techniques facilitate detection and characterization of the microscopic and ultrastructural sites and distribution of calcific deposits, and their relationship to tissue or biomaterials structural details. Such analyses directly yield important qualitative (but not quantitative) information. In contrast, chemical techniques, which require destruction of the tissue specimen, permit identification and quantitation of both bulk elemental composition and determination of crystalline mineral phases. However, such techniques generally cannot relate the location of the mineral to the details of the underlying tissue structure. The most comprehensive studies use several analytical modalities to simultaneously characterize both morphologic and chemical aspects of calcification. Moreover, newer techniques are available for nondestructive and potentially noninvasive characterizing of calcification, both in specimens (microcomputed tomography [micro-CT]) (Ford-Hutchinson et  al., 2003; Rousselle et al, 2019; Neues and Epple, 2008), and in  vivo using molecular imaging (Aikawa et  al.,

2007). Micro-CT has been used extensively in studies of bone-regenerative biomaterials (Jiang et  al., 2009). Molecular imaging, which probes biomarkers of particular targets or pathways of the cellular and molecular mechanisms of calcification, is particularly exciting in this context to enable the visualization of the ongoing and dynamic process of calcification, potentially quantitatively and repetitively in living organisms (New and Aikawa, 2011).

Morphologic Evaluation Morphologic assessment of calcification is done using several readily available and well-established techniques that range from macroscopic (gross) examination and radiographs (X-rays) of explanted prostheses to more sophisticated analytical techniques such as electron energy loss spectroscopy. Each method has advantages and limitations; several techniques are often used in combination to obtain an understanding of the structure, composition, and mechanism of each type of calcification. Careful visual examination of the specimen, often under a dissecting (low-power) microscope and radiography assess the distribution of mineral. Specimen radiography typically involves placing the explanted prosthesis on an X-ray film plate and exposing to an X-ray beam in a special device used for small samples (e.g., Faxitron, Hewlett–Packard, McMinnville, CA, with an energy level of 35 keV for 1 min for valves). Deposits of mineral appear as bright densities which have locally blocked the beam from exposing the film (see Fig. 2.4.5.1B).

978 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

(A)

(A)

(B)

(B)



Figure 2.4.5.3 Light microscopic appearance of calcification of experimental porcine aortic heart valve tissue. Note cell-based orientation of initial deposits (arrows). (A) Implanted subcutaneously in 3-week-old rats for 72 h; (B) implanted in growing sheep for 5 months, demonstrating predominant site of growing edge of calcification in cells of the residual porcine valve matrix (arrows). ((B) Reproduced by permission from Schoen, F. J. (2001). Pathology of heart valve substitution with mechanical and tissue prostheses, In: Cardiovascular Pathology, third ed., Silver, M. D., Gotlieb, A. I. and Schoen, F. J. (Eds.). W. B. Saunders: Philadelphia, PA, pp. 629–677.)

Light microscopy of calcified tissues is widely used. Identification of calcium-phosphate mineral is facilitated through the use of either calcium- or phosphorus-specific stains, such as alizarin red (which stains calcium red) or von Kossa (which stains phosphates brown-black) (Figs. 2.4.5.2B and 2.4.5.3). The von Kossa stain generally permits the most discrete localization of the mineral deposits. Staining for alkaline phosphatase, an enzyme involved in the pathogenesis of some forms of calcification, can also be useful in some contexts (Maranto and Schoen, 1988; Levy et al., 1991). These histologic stains are readily available, can be easily applied to tissue sections embedded in either paraffin or plastic, and are most useful for confirming and characterizing suspected calcified areas which have been noted by routine hematoxylin and eosin staining techniques. Sectioning of calcified tissue which has been embedded in paraffin often leads to considerable artifacts due to fragmentation; embedding of tissue with calcific deposits in a harder medium such as glycolmethacrylate polymer can yield superior section quality. On light microscopic sections, the extent of calcification can be semiquantitatively graded, and its locations described relative to key landmarks of the implant (Bennink et al., 2018). Transmission electron microscopy may facilitate determination of early sites of calcific deposits. In TEM, an electron beam traverses an ultra-thin (0.05 μm) section of tissue (Fig. 2.4.5.4); observation of the ultrastructure (i.e., submicron tissue features) of calcification by TEM may contribute to an understanding of the mechanisms by which calcific



Figure 2.4.5.4 Transmission electron microscopy of calcification of experimental porcine aortic heart valve implanted subcutaneously in 3-week-old rats. (A) 48-hour implant demonstrating focal calcific deposits in nucleus of one cell (closed arrows) and cytoplasm of two cells (open arrows), n, nucleus; c, cytoplasm. (B) 21-day implant demonstrating collagen calcification. Bar  =  2  μm. Ultrathin sections stained with uranyl acetate and lead citrate. ((A) Reproduced with permission from Schoen, F.J., Levy, R.J., Nelson, A.C., Bernhard, W.F., Nashef, A., et al. (1985). Onset and progression of experimental bioprosthetic heart valve calcification. Lab. Investig., 52, 523–532.)

crystals form. Scanning electron microscopy (SEM) images the specimen surface and can be coupled with elemental localization by energy-dispersive X-ray analysis (EDXA), allowing a semiquantitative evaluation of the local progression of calcium and phosphate deposition in a site-specific manner. Electron energy loss spectroscopy (EELS) couples transmission electron microscopy with highly sensitive elemental analyses to provide a most powerful localization of incipient nucleation sites and early mineralization (Webb et al., 1991). In general, the more highly sensitive and sophisticated morphologic techniques require more demanding and expensive preparation of specimens to avoid unwanted artifacts, which may complicate interpretation of high-resolution morphologic information. Since choice

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

979

(A)

(B) Ca2+

Ca2+ - ATPase

[Ca2+] approx 10−3 M

EXTRACELLULAR

[Ca2+] approx 10−7 M ATP

calcium binding molecule

ADP+P

Ca2+

INTRACELLULAR

Ca2+

mitochondrion

• Figure 2.4.5.5  Extended hypothetical model for the calcification of bioprosthetic tissue. (A) Overall model,

which considers host factors, implant factors, and mechanical damage, and relates initial sites of mineral nucleation to increased intracellular calcium in residual cells and cell fragments in bioprosthetic tissue. The ultimate result of calcification is valve failure, with tearing or stenosis. The key contributory role of existing phosphorus in membrane phospholipids and nucleic acids in determining the initial sites of crystal nucleation is emphasized, and a possible role for the independent mineralization of collagen is acknowledged. Mechanical deformation probably contributes to both nucleation and growth of calcific crystals. (B) Events at the cell membrane and other calcium-binding structures. There is a substantial physiologic (normal) gradient of free calcium across the cell membrane (10−3 M outside, 10−7 M inside) which is maintained as an energy-dependent process. With cell death or membrane dysfunction, calcium phosphate formation can be initiated at the membranous cellular structures. ((A) Reproduced by permission from Schoen, F. J. and Levy, R. J. (2005). Calcification of tissue heart valve substitutes: Progress toward understanding and prevention. Ann. Thorac. Surg., 79, 1072–1080. (B) Reproduced by permission from Schoen, F. J. et al. (1988). Biomaterials-associated calcification: Pathology, mechanisms, and strategies for prevention. J. Appl. Biomater., 22, 11–36.)

of analytical methodology involves trade-offs, forethought about, and careful planning, specimen handling optimizes the yield provided by the array of available techniques and allows multiple approaches to be used on specimens from a particular experiment. 

Chemical Assessment Quantitation of calcium and phosphorus in biomaterial calcifications permits characterization of the progression of deposition, comparison of severity of deposition among specimens, and determination of the effectiveness of preventive measures (Levy et al., 1983a, 1985a; Schoen et al., 1985, 1986, 1987; Schoen and Levy, 1999). However, such techniques destroy the specimen (and hence obliterate any spatial information) during preparation. A widely used technique for quantitation of calcium is atomic absorption spectroscopy of acid-hydrolyzed or ashed samples. Phosphorus is usually quantitated as phosphate, using a molybdate complexation technique with spectrophotometric detection. The degree of calcification associated with a gross explant specimen can also be semiquantitatively graded by specimen

radiographs (Schoen et  al., 1987). The specific crystalline form of the mineral phase can be determined by X-ray diffraction. Carbonate-containing mineral phases may also be analyzed by infrared spectroscopy. 

Mechanisms of Biomaterial Calcification Regulation of Pathologic Calcification Three factors regulate biomaterial mineralization: (1) biological factors (local environment of function and recipients systemic metabolic state and circulating regulatory factors); (2) biomaterial factors (the structure and chemistry of the substrate biomaterial); and (3) biomechanical factors (degree and locations of stress and strain). (see Fig. 2.4.5.5). 

Role of Biological Factors The most important host metabolic factors relate to young age, with more rapid calcification taking place in younger patients or experimental animals (Levy et  al., 1983a), and pathologic alterations of mineral concentrations in the

980 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

circulating blood. Although the relationship with age is wellestablished and incontrovertible, the mechanisms accounting for this “physiologic” effect of age are uncertain but may be related to a more active mineral metabolism that exists at a young age to create and strengthen the skeleton. Other factors that can change blood chemistry such as comorbidities (e.g., chronic kidney impairment or diabetes) may aggravate the mineral deposition process. Local tissue factors such as oxidative potential and pH alteration due to inflammation can also stimulate material degradation and calcification. Dystrophic pathologic calcification has typically been considered a passive, unregulated, and degenerative process. However, the observations of matrix vesicles, hydroxyapatite mineral, and bone-related morphogenetic and noncollagenous proteins in situations of pathological calcification have suggested that the mechanisms responsible for pathologic calcification may be at least partially regulated, similarly to physiologic mineralization of bone and other hard tissues. Moreover, some of the mechanisms regulating biomaterial-associated calcification may be shared with other pathological process that have calcification as a prominent feature, such as atherosclerosis and calcific aortic valve disease (Johnson et al., 2006; Demer and Tintut, 2008; Persy and D’Haese, 2009). In healthy blood vessels and valves, inhibitory mechanisms outweigh procalcification-inductive mechanisms; in contrast, in bone and pathologic tissues, inductive mechanisms dominate. Naturally occurring inhibitors to crystal nucleation and growth (of which many have been identified) may also play a role in biomaterial and other cardiovascular calcification (Weissen-Plenz et al., 2008). Specific inhibitors in this context include osteopontin (Steitz et al., 2002) and phosphocitrate (Tew et al., 1980). Naturally occurring mineralization cofactors, such as inorganic phosphate (Jono et al., 2000), bone morphogenic protein (a member of the transforming growth factor [TGF] beta family) (Bostrom et al., 1993), proinflammatory lipids (Demer, 2002), and other substances (e.g., cytokines) may also play a role in pathologic calcification. The noncollagenous proteins osteopontin, TGF-beta1, and tenascin-C involved in bone matrix formation and tissue remodeling have been demonstrated in clinical calcified bioprosthetic heart valves, natural valves, and atherosclerosis, suggesting that they play a regulatory role in these forms of pathologic calcification in humans (Srivasta et al., 1997; Bini et al., 1999; Jian et al., 2001, 2003), but a direct mechanistic relationship has not yet been demonstrated. 

Role of Biomaterial Factors Biomaterial structural and chemical properties play a significant role in the propensity of calcification. An important implant factor for calcification of bioprosthetic tissue is pretreatment with glutaraldehyde, done to preserve the tissue (Golomb et al., 1987; Grabenwoger et al., 1996). It has been hypothesized that the cross-linking agent glutaraldehyde stabilizes and perhaps modifies phosphorous-rich calcifiable structures in the bioprosthetic tissue. These sites seem to be capable of mineralization upon implantation when exposed to the comparatively high calcium levels of extracellular fluid. Calcification of the two principal types of tissue used

in bioprostheses—glutaraldehyde-pretreated porcine aortic valve or glutaraldehyde-pretreated bovine pericardium—is similar in extent, morphology, and mechanisms (Schoen and Levy, 2005). Glutaraldehyde also has been shown to be an inefficient cross-linker for other tissue-based materials. Nevertheless, the elastin and glycosaminoglycans present in the tissue are not stabilized, and degradation and leaching of those components can create voids and a site for additional calcification (Lovekamp et al., 2006). Thus stabilizing those components in the biological tissue has been shown to improve calcification resistance (Tripi and Vyavahare, 2014). Degraded elastin fibers can also be calcificationprone sites (Bailey et al., 2003). Calcification mechanisms of synthetic polymers are different from those of tissue-based materials and mostly occur on the surface of the materials, a feature generally considered to be related to protein and platelet attachment (Bernacca et al., 1997). Surface properties of the materials play a significant role in material calcification. Segmented polyether urethanes are prone to oxidation, and that can cause surface cracking that would provide sites for mineralization (Stachelek et  al., 2006). Increasing hydrophilicity of degradable polyurethanes has been shown to increase the extent of calcification (Gorna and Gogolewski, 2002). Surface modification of synthetic polymers has been extensively investigated to prevent calcification (Bernacca and Wheatley, 1998). 

Role of Biomechanical Factors Mechanical factors also regulate calcification. Both intrinsic and extrinsic mineralization of a biomaterial are generally enhanced at the sites of intense mechanical deformations generated by motion, such as the points of flexion in heart valves (Thubrikar et  al., 1983; Schoen and Levy, 1999). The mechanisms underlying mechanical potentiation of calcification associated with biomaterials are incompletely understood, but the effect mimics the well-known Wolf ’s law in bone, in which formation and adaptation occurs in response to the mechanical forces that it experiences (Chen et al., 2009). Moreover, enhancement of mineral deposition is seen in systems where static but not dynamic mechanical strain is applied (Levy, 1983a), and in systems where live stem cells are subjected to a spectrum of mechanical environments (Yip et al., 2009). Furthermore, long-term cyclic wear on materials has been demonstrated to cause permanent geometric deformations that are due to ECM degradation in tissue-based valves, specifically the collagen triple helix unraveling (Sun et  al., 2004). Recently, it has been suggested that preimplantation crimping of the pericardium in transcatheter aortic valves can lead to material damage that promotes calcification (Zareian et al., 2019). 

Experimental Models for Biomaterial Calcification Animal models have been developed for the investigation of the calcification of bioprosthetic heart valves, aortic homografts, cardiac assist devices, and trileaflet polymeric valves

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

981

TABLE 2.4.5.3   Experimental Models of Calcification and Typical Durations

Type

System

Typical Duration

Calcification of bioprosthetic or other tissue heart valve

In vitro incubation of tissue fragment or flexing valves

Days to weeks

Rat subdermal implant of tissue fragment

3 weeks

Calf or sheep orthotopic valve replacement

3–5 months

Rat or sheep descending aorta

1–5 months

Rat subdermal implant of material sample after induction of calciphylaxis

1–3 months

Calf or sheep artificial heart implant

5 months

Trileaflet polymeric valve implant in calf or sheep

5 months

Calcification of hydrogel

Rat subdermal implant of material sample

3 weeks

Calcification of collagen or elastin

Rat subdermal implant of material sample

3 weeks

Urinary encrustation

In vitro incubation

Hours to days

In vivo bladder implants (rats and rabbits)

10 weeks

Calcification of polyurethane

(Table 2.4.5.3). The most widely used experimental models are used as a preclinical screen in the development of new or modified materials and design configurations, and to investigate the pathophysiology of bioprosthetic tissue calcification. They include orthotopic valve replacements (in the tricuspid, pulmonary, mitral, or aortic positions) or conduit-mounted valves in sheep or calves, and isolated tissue (i.e., not in a valve) samples implanted subcutaneously in mice, rabbits, or rats (Levy et al., 1983a; Schoen et al., 1985, 1986). In both circulatory and noncirculatory models, bioprosthetic tissue calcifies progressively with a morphology similar to that observed in clinical specimens, but with markedly accelerated kinetics. In vitro models of biomaterial calcification have been investigated, but have not yielded useful information in studying mechanisms or preventive strategies (Schoen et al., 1992a; Mako and Vesely, 1997). Compared with the several years normally required for calcification of clinical bioprostheses, valve replacements in sheep or calves calcify extensively in 3–6 months (Schoen et al., 1985, 1986, 1994). However, these models are expensive, limiting the number of subjects, and they are technically complex, demanding a high level of surgical expertise, postoperative care, and stringent housing and management procedures. Also, the available large animal models grow rapidly, have differences in anatomy from that of humans, and are without the pathologic substrate into which human implants are placed, and often yield results that vary among subjects treated similarly. These limitations stimulated the development of subdermal (synonym, subcutaneous—under the skin) implant models. Subdermal bioprosthetic implants in rats, rabbits, and mice provide the following useful features: (1) a markedly accelerated rate of calcification in a morphology comparable to that seen in circulatory explants; (2) sufficient economies that permit many specimens to be studied with a given set of experimental conditions, thereby

allowing quantitative characterization and statistical comparisons; and (3) quick retrieval of specimens from the experimental animals, which facilitates the careful manipulation and rapid processing required for detailed analyses (Levy et al., 1983a,b; Schoen et al., 1985, 1986). Moreover, controls are easily generated so that specific factors that may place a role in calcification can be isolated, and hypothetical mechanisms studied in detail. Moreover, these models provide a data set to which inhibitory approaches can be screened. Polymer calcification has also been studied with subdermal implants in rats (Joshi et al., 1996). Calcification of a synthetic polymer such as polyurethane mostly occurs on the surface due to the hydrophobic nature of the materials, and it is therefore much more difficult to assess. Changes in surface or bulk properties or topology of the polymer due to these degradation processes or adherent biological material may promote extrinsic calcification. Oxidation of polymer chains may promote polymer degradation. For example, polyether urethanes calcify more than polycarbonate urethanes due to susceptibility of ether bonds to oxidation (Bernacca et  al., 1997; Joshi et  al., 1996). These degradation and subsequent calcification processes often mechanically weaken the polymer (surface cracking and tearing) and enhance vulnerability to failure in a dynamic setting of cardiovascular applications. Aggravated calcification animal models such as calciphylaxis are often needed to accelerate and augment calcification of polymers (Joshi et al., 1996). Thus, the subcutaneous model provides a technically convenient and economically advantageous vehicle for investigating host and implant determinants and mechanisms of mineralization, as well as for an initial screening of potential strategies for its inhibition (anticalcification). Promising approaches may be investigated further in a large animal valve implant model. Large animal implants

982 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

as valve replacements are also used: (1) to elucidate further the processes accounting for clinical failures; (2) to evaluate the performance of design and biomaterial modifications in valve development studies; (3) to assess the importance of blood–surface interactions; and (4) to provide data required for approval by regulatory agencies (Schoen et al., 1992b, 1994). Indeed, regulatory guidelines require such large animal implants and prescribe many of the parameters of these animal studies and postexplant analysis (Gallegos et  al., 2005; Zhang et al., 2019).  Data from valve explants from patients and subdermal and circulatory experiments in animal models using bioprosthetic heart valve tissue have elucidated the pathophysiology of this significant clinical problem and enhanced our understanding of pathologic biomaterial calcification more generally (Fig. 2.4.5.4). The similarities of calcification in the different experimental models and clinical bioprostheses suggest common mechanisms, independent of the implant site (across vascular and extravascular locations). Indeed, calcification of bioprosthetic tissue appears to depend on exposure of a susceptible substrate to calcium-containing extracellular fluid; other local implant-related or circulating substances may play a regulatory role. Residual phosphorus in cell-related membranes localizes the Ca–P mineral nucleation sites. Consistent with a dystrophic mechanism, and encountered in other situations (Cardoso and Weinbaum, 2018), the initial calcification sites in bioprosthetic tissue are predominantly dead cells and cell membrane fragments (Levy et al., 1983a; Schoen et al., 1985, 1986; Schoen and Levy, 1999) (see Fig. 2.4.5.4).

Role of Cells Normally, extracellular calcium concentration in tissues is approximately 1 mg/mL (10−3 M). In contrast, the intracellular concentration of calcium (in the cytoplasm) is 1000– 10,000 times lower (approximately 10−7 M). This steep calcium gradient is maintained in healthy cells because mechanisms in their intact membranes pump calcium to the exterior using calcium ATP-ases and other vehicles for extrusion. In cells which have been rendered nonviable by glutaraldehyde fixation, the physiologic cellular handling of calcium ions is disabled. The cell membranes and other intercellular structures are high in phosphorus (as phospholipids, especially phosphatidyl serine, which can bind calcium); such sites can serve as nucleators. Glutaraldehyde pretreatment also stabilizes these phosphorous stores. Other susceptible sites can include collagen and elastic fibers of the extracellular matrix, denatured proteins, phosphoproteins, fatty acids, blood platelets and, in the case of infection, bacteria. Moreover, since the morphology and extent of calcification in subcutaneous implants is analogous to that observed in clinical and experimental circulatory implants, despite the lack of dynamic mechanical activity characteristic of the circulatory environment, it is clear that dynamic stress promotes, but is not prerequisite for, calcification of

bioprosthetic tissue. In the subcutaneous model, calcification is enhanced in areas of tissue folds, bends, and areas of shear, suggesting that static mechanical deformation also potentiates mineralization (Levy et  al., 1983a). Also, long-term bioprosthetic valve function is accompanied by collagen degradation in the most highly stressed regions (Sun et al., 2004; Vyavahare et al., 1997; Sacks and Schoen, 2002); highly stressed regions typically manifest the earliest and most substantial calcification. Nevertheless, although these data suggest that local tissue disruption mediates the mechanical effect, the precise mechanisms by which mechanical factors influence calcification are uncertain. 

Role of Collagen and Elastin Calcification of the extracellular matrix structural proteins collagen and elastin has been observed in clinical and experimental implants of bioprosthetic and homograft valvular and vascular tissue and has been studied using a rat subdermal model. Collagen-containing implants are widely used in various surgical applications, such as tendon prostheses and surgical absorptive sponges, but their usefulness is compromised owing to calcium phosphate deposits and the resultant stiffening. Cross-linking by either glutaraldehyde or formaldehyde promotes the calcification of collagen sponge implants made of purified collagen, but the extent of calcification does not correlate with the degree of cross-linking (Levy et al., 1986). Elastin calcification has also been studied (Vyavahare et al., 1999; Lee et al., 2006). As mentioned previously, elastin is not stabilized by glutaraldehyde cross-linking in bioprosthetic valves and its degradation, in particular in an aortic portion of the valve, can lead to elastin-specific calcification (Isenburg et al., 2005). 

Role of Glutaraldehyde The role of glutaraldehyde (GA) in promoting calcification is well established and strong (Golomb et al., 1987). Two key processes, not yet fully understood, yet both directly related to the chemical effects of GA, are believed to dominate: (1) the reaction between the residual phosphorus and associated phospholipids with calcium in the surrounding fluid to yield calcium-phosphate mineral associated with the cells of the tissue devitalized by GA (as described above), and (2) chemical reactions related to the presence of active residual-free aldehyde functional groups induced by GA fixation (Simionescu, 2004). Collagen and elastic fibers can also serve as calcium-phosphate nucleation sites, independent of cellular components, perhaps through reactive aldehyde-related mechanisms. It is unknown whether these several mechanisms are independent or interrelated. In addition to cellular devitalization and stabilizing phosphorus in the tissue, glutaraldehyde treatment produces functional group residuals (i.e., aldehydes, various carbonyls, Schiff bases, etc.) by which not only cells, but potentially

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

matrix glycosaminoglycans and structural proteins, such as collagen and elastin, may also bind calcium. Tissue levels of reactive functional group residuals can be directly correlated to bioprosthetic tissue calcification (Manji et al., 2006; Zilla et al., 1997a,b; Cunanan et al., 2001; Tod and Dove, 2016). Indeed, tissue levels of calcification correlate with aldehyde content. The close association between aldehyde content and the calcification of bovine pericardial tissue suggests that processing of bioprosthetic valves to reduce the content or reactivity of reactive residual aldehydes may offer a significant advantage in terms of reducing the potential for long-term bioprosthetic valve calcification (Webb et al., 1988; Shang et al., 2017; Christian et al., 2015; Everaerts et  al., 2006). Oxidative mechanisms may also play a role (Christian et al., 2014; Lee et al., 2017). 

Role of Immunologic Factors A potential role for inflammatory and immune processes has been postulated by some investigators (Love, 1993; Human and Zilla, 2001a,b), since (1) experimental animals can be sensitized to both fresh and cross-linked bioprosthetic valve tissues; (2) antibodies to valve components can be detected in some patients following valve dysfunction; and (3) some failed tissue valves have brisk mononuclear inflammation. However, no definite evidence for a role of immunologic mechanisms in calcification has been demonstrated, and many lines of evidence suggest that neither nonspecific inflammation nor specific immunologic responses appear to favor bioprosthetic tissue calcification. Specific evidence against a contributory role of immunologic process is as follows: (1) calcification in either circulatory or subcutaneous locations is not usually associated with inflammation, and (2) in experiments in which valve cusps were enclosed in filter chambers that prevent host cell contact with tissue, but allow free diffusion of extracellular fluid and implantation of valve tissue in congenitally athymic (“nude”) mice who have essentially no T-cell function, calcification morphology, and extent were unchanged (Levy et al., 1983a,b). Clinical and experimental data detecting antibodies to valve tissue after failure might reflect a secondary response to valve damage, rather than a cause of failure. Calcification of the adjacent aortic wall portion of glutaraldehyde-pretreated porcine aortic valves, and valvular allografts and vascular segments, is also observed clinically and experimentally. Mineral deposition occurs throughout the vascular cross-section but is accentuated in the dense bands at the inner and outer media, and cells and elastin (itself not a prominent site of mineralization in cusps) are the major sites. In nonstented porcine aortic valves which have higher portions of the aortic wall exposed to blood than in currently used stented valves, calcification of the aortic wall could stiffen the root, altering hemodynamic efficiency, causing nodular calcific obstruction, potentiating wall rupture or providing a nidus for emboli. Some anticalcification agents, including 2-amino-oleic acid (AOA) and ethanol, prevent experimental cuspal but not aortic wall calcification (Chen et al., 1994a,b). 

983

Prevention of Calcification The major strategy for preventing calcification of biomaterial implants has been biomaterial modifications, whether by removal of a calcifiable component, the addition of an exogenous agent, or chemical alteration. The principles are well-illustrated by the work done with bioprosthetic heart valves, in which the subcutaneous model has been used to screen potential strategies for calcification inhibition (anticalcification or antimineralization). Promising approaches have been investigated further in a large animal valve implant model. However, some strategies that appeared efficacious in subcutaneous implants have not proven favorable when used on valves implanted into the circulation. Analogous to any new or modified drug or device, a potential antimineralization treatment must be effective and safe. The treatment should not impede valve performance such as hemodynamics or durability. Investigations of an anticalcification strategy must demonstrate not only the effectiveness of the therapy but also the absence of adverse effects (Schoen et  al., 1992b). Adverse effects in this setting could include systemic or local toxicity, a tendency toward thrombosis or infection, induction of immunological effects, or accelerated structural degradation, with either immediate loss of mechanical properties or premature deterioration and failure. Indeed, there are several examples whereby an antimineralization treatment contributed to an unacceptable degradation of the tissue (Jones et al., 1989; Gott et al., 1992; Schoen, 1998). As summarized in more detail in Table 2.4.5.4, a rational approach for preventing bioprosthetic calcification must integrate safety and efficacy considerations with the scientific basis for inhibition of calcium-phosphate crystal formation. This will of necessity involve the steps summarized in Table 2.4.5.5, before appropriate clinical trials can be undertaken (Schoen et al., 1992b; Vyavahare et al., 1997a). Experimental studies have demonstrated that adequate doses of systemic agents used to treat clinical metabolic bone TABLE   Criteria for Efficacy and Safety of 2.4.5.4  Antimineralization Treatments

Efficacy • Effective and sustained calcification inhibition

Safety • Adequate valve performance (i.e., unimpaired hemodynamics and durability) • Does not cause adverse blood–surface interactions (e.g., hemolysis, platelet adhesion, coagulation protein activation, complement activation, inflammatory cell activation, binding of vital serum factors) • Does not enhance local or systemic inflammation (e.g., foreign body reaction, immunologic reactivity, hypersensitivity) • Does not cause local or systemic toxicity • Does not potentiate infection Modified from Schoen F.J. et al. (1992). Antimineralization treatments for bioprosthetic heart valves. Assessment of efficacy and safety. J. Thorac. Cardiovasc. Surg., 104, 1285–1288.

984 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

TABLE   Preclinical Efficacy and Safety Testing of 2.4.5.5  Antimineralization Treatments

Type of Study

Information Derived

Subcutaneous implantation in rats

• Initial efficacy screen • Mechanisms • Dose–response • Toxicity

Biomechanical evaluation

TABLE   Prototypical Agents for Mechanism-Based 2.4.5.6  Prevention of Calcification in Bioprosthetic

Heart Valves

Mechanisms

Strategy/Agent

Inhibition of hydroxyapatite formation

Ethane hydroxybisphosphonate (EHBP)

• Hemodynamics • Accelerated wear

Inhibition of calcium uptake

2-Alpha-amino-oleic acid (AOA™)

Morphologic studies of unimplanted valves

• Structural degradation assessed by light and transmission electron microscopy • Scanning electron microscopy

Ferric/aluminum chloride exposure

Circulatory implants in large animals

• Device configuration, surgical technique, in vivo hemodynamics, explant valve pathology • Durability, thrombi, thromboembolism, hemolysis, cardiac and systemic pathology

Inhibition of Ca–P crystal growth; inhibition of alkaline phosphate; chemical modification of elastin Phospholipid extraction

Sodium dodecyl sulfate (SDS)

Phospholipid extraction and collagen conformation modification

Ethanol exposure

Binding of agents that prevent calcification

Osteopontin, EDTA, heparin

Eliminate glutaraldehyde potentiation of calcification: • Amino acid neutralization of glutaraldehyde residues • Polyepoxide (polyglycidal ether), acyl azide, carbodiimide, cyanimide and glycerol cross-linking • Dye-mediated photooxidation • Natural crosslinkergenipin, tannic acid • Polymerization of aldehydes by thermal treatment

Modification of (alternatives to) glutaraldehyde fixation

Modified from Schoen F.J. et al. (1992). Antimineralization treatments for bioprosthetic heart valves. Assessment of efficacy and safety. J. Thorac. Cardiovasc. Surg., 104, 1285–1288.

disease, including calcium chelators (e.g., diphosphonates such as ethane hydroxybisphosphonate [EHBP]), can dramatically lower the calcification potential of bioprosthetic tissue implanted subcutaneously in rats (Levy et al., 1987). However, systemic therapy is unlikely to be safe, because such chemicals generally also interfere with physiologic calcification (i.e., bone growth), causing growth retardation. To avoid this difficulty, coimplants of a drug-delivery system adjacent to the prosthesis, in which the effective drug concentration would be confined to the site where it is needed (i.e., the implant), were investigated, in the hope that systemic sideeffects would be prevented (Levy et al., 1985b). A localized anticalcification effect would be particularly attractive in young people. Although studies incorporating EHBP in nondegradable polymers, such as ethylene-vinyl acetate (EVA), polydimethylsiloxane (silicone), silastic, and polyurethanes showed effectiveness of this strategy in animal models, this approach has been challenging to implement clinically. Thus, modification of the substrate, either by removing or altering a calcifiable component or binding an inhibitor has been an area of focus. Forefront strategies should also consider: (1) a possible synergism provided by multiple anticalcification agents and approaches used simultaneously; (2) new materials; and (3) the possibility of tissue-engineered heart valve replacements. The agents most widely studied, for efficacy, mechanisms, lack of adverse effects, and potential clinical utility are summarized below and in Table 2.4.5.6. Combination therapies using multiple agents may provide a synergy of beneficial effects (Levy et  al., 2003). The paragraphs below describe some of the approaches that have been investigated in preclinical studies; several have achieved regulatory approval and proven successful in clinical bioprosthetic valves.

Inhibitors of Hydroxyapatite Formation Bisphosphonates Ethane hydroxybisphosphonate (EHBP) has been approved by the FDA for human use to inhibit pathologic calcification, and to treat hypercalcemia of malignancy. Compounds of this type probably inhibit calcification by poisoning the growth of calcific crystal by chelation. As discussed above, either cuspal pretreatment or systemic or local therapy of the host with diphosphonate compounds inhibits experimental bioprosthetic valve calcification (Levy et al., 1985a,b, 1987; Johnston et  al., 1993). Recently, controlled clinical trials which have orally administered bisphosphonates have demonstrated the ability to stabilize osteoporosis. 

Trivalent Metal Ions Pretreatment of bioprosthetic tissue with iron and aluminum (e.g., FeCl3 and AlCl3) inhibited calcification of subdermal implants with glutaraldehyde-pretreated porcine cusps or

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

985

Alcohol Treatments



Figure 2.4.5.6 Reduction of calcification of bioprosthetic tissue by preincubation in 1% SDS demonstrated in a rat subcutaneous model of glutaraldehyde cross-linked porcine aortic valve. These results support the concept that phospholipid extraction is an important but perhaps not the only mechanism of SDS efficacy. (Reproduced by permission from Schoen, F. J., Levy, R. J. and Piehler, H. R. (1992). Pathological considerations in replacement cardiac valves. Cardiovasc. Pathol., 1, 29–52.)

Ethanol preincubation of glutaraldehyde cross-linked porcine aortic valve bioprostheses prevents calcification of the valve cusps in both rat subdermal implants, and sheep mitral valve replacements (Vyavahare et al., 1997b, 1998). Eighty percent ethanol pretreatment: (1) extracts almost all phospholipids and cholesterol from glutaraldehyde cross-linked cusps; (2) causes a permanent alteration in collagen conformation as assessed by attenuated total reflectance-Fourier transform infrared spectroscopy (ATR-FTIR); (3) affects cuspal interactions with water and lipids; and (4) enhances cuspal resistance to collagenase. Ethanol is in clinical use as a porcine valve cuspal pretreatment in both Europe and the United States, and use in combination with aluminum treatment of the aortic wall of a stentless valve is under consideration. Other alcohols that have shown promise to remove phospholipids include methanol, isopropanol, and 1,2 octanediol (Vyavahare et al., 2000; Pettenazzo et al., 2008). 

Glutaraldehyde Neutralization pericardium (Webb et  al., 1991; Carpentier et  al., 1997). Such compounds are hypothesized to act through complexation of the cation (Fe or Al) with phosphate, thereby preventing calcium phosphate formation. Both ferric ion and the trivalent aluminum ion inhibit alkaline phosphatase, an important enzyme in bone formation. This may be related to their mechanism for preventing the initiation of calcification. Furthermore, research from our laboratories has demonstrated that aluminum chloride prevents elastin calcification through a permanent structural alteration of the elastin molecule (Vyavahare et al., 1999). These compounds are also active when released from polymeric controlled-release implants. 

Calcium Diffusion Inhibitor 2-Alpha-amino-oleic acid (AOA, Biomedical Designs, Inc., Atlanta, GA) bonds covalently to bioprosthetic tissue through an amino linkage to residual aldehyde functions, and inhibits calcium flux through bioprosthetic cusps (Chen et al., 1994a,b). AOA is effective in mitigating cusp, but not aortic wall, calcification in rat subdermal and cardiovascular implants. This compound is used in FDA-approved porcine aortic valves (Fyfe and Schoen, 1999; Celiento et al., 2012; El-Hamamsy et al., 2010). 

Removal/Modification of Calcifiable Material Surfactants Incubation of bioprosthetic tissue with sodium dodecyl sulfate (SDS) and other detergents extracts the majority of acidic phospholipids (Hirsch et al., 1993); this is associated with reduced mineralization, probably resulting from suppression of the initial cell-membrane-oriented calcification (Fig. 2.4.5.6). This compound is used in an FDA-approved porcine aortic valve bioprosthesis (David et al., 1998; Bottio et al., 2003). 

As glutaraldehyde cross-linking itself has been shown to aggravate tissue calcification (Levy, 1994; Gong et al., 1991), several approaches have been suggested to neutralize remnant aldehyde residues in the tissue after glutaraldehyde cross-linking. Several amino acids, amines, and natural compounds such as taurine have been used to bind to unreacted glutaraldehyde residues (Jorge-Herrero et  al., 1996; Meuris et  al., 2018). These treatments have shown promise in preventing tissue calcification in rodent subcutaneous models and sheep heart valve replacement models. Storage of tissues in glutaraldehyde after the anticalcification treatment has shown a partial return of calcification for the tissues. Thus, other storage conditions without glutaraldehyde have been investigated. One product where the bioprosthetic valve is stored in dry conditions after glycerol treatment is now clinically available (Inspiris Resilia by Edwards Lifesciences, Irvine, California). New methods where the combination of multiple treatments such as removing phospholipids and glutaraldehyde neutralization followed by storage in glutaraldehyde free solutions have shown promise in animal studies to prevent calcification (Meuris et al., 2018). 

Decellularization Since the initial mineralization sites are devitalized connective cells of bioprosthetic tissue, these cells may be removed from the tissue, with the intent of making the bioprosthetic matrix less prone to calcification (Courtman et  al., 1994; Wilson et al., 1995). Decellularized tissues have been used as scaffolds for tissue-engineered heart valves that would be populated with host cells and would grow with patients and hopefully not calcify. In vivo animal studies have seen some success for patency of a functional valve (Dohmen et  al., 2006; Baraki et  al., 2009). However, clinical experiments have shown valve failure due to inflammation and calcification (Simon et  al., 2003). It is clear from these studies that extracellular matrix in the tissue (collagen and elastin) can degrade and create sites for calcification. Physical stabilization of collagen and elastin by pentagalloyl glucose

986 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

TABLE 2.4.5.7   Currently Used Anticalcification Treatments for Bioprosthetic Heart Valves

Heart Valve Producer

Anticalcification Trade Name

Anticalcification Method

Edwards Lifesciences

Thermafix™

Proprietary annealing process developed internally to increase the durability of the material and rid excess aldehyde residues

Edwards Lifesciences

Inspiris Resilia

Dehydration of the tissue and reduction in lipid content, dry storage

Medtronic

AOA™

Alpha-oleic-acid treatment to decrease high-affinity sites of calcification, neutralizes residual glutaraldehyde

Abbott (formerly St. Jude Medical)

Linx™

Ethanol treatment to decrease high-affinity sites of calcification combined with a proprietary method to rid excess aldehyde residues

has been shown to reduce valvular calcification (Deborde et  al., 2016). CorMatrix, a decellularized porcine intestinal submucosa, is also being investigated for cardiovascular implants (Zaidi et al., 2014). 

Modification of Glutaraldehyde Fixation and Other Tissue Fixatives Since previous studies have demonstrated that conventional glutaraldehyde fixation is conducive to calcification of bioprosthetic tissue, several studies have investigated modifications of, and alternatives to, conventional glutaraldehyde pretreatment. For example, and paradoxically, fixation of bioprosthetic tissue by extraordinarily high concentrations of glutaraldehyde (5–10× those normally used) appear to inhibit calcification (Zilla et al., 1997a,b, 2000). Moreover, residual glutaraldehyde residues in bioprosthetic tissue can be neutralized (“detoxified”) by treatment with lysine or diamine, which inhibits calcification of subdermal implants (Grabenwoger et al., 1992; Zilla et  al., 2000, 2005; Trantina-Yates et  al., 2003). Removing phospholipids or neutralizing aldehyde residues by end capping with amines, reduction with borohydride, or using hightemperature fixation strategies have been now clinically used for preventing tissue calcification (Table 2.4.5.7) Nonglutaraldehyde cross-linking of bioprosthetic tissue with epoxides, carbodiimides, acyl azides, and other compounds reduces their calcification in rat subdermal implant studies (Xi et al., 1992; Myers et al., 1995), and with triglycidylamine (TGA), an epoxy cross-linker (Connolly et al., 2005). Additionally, alternative cross-linking chemistries (i.e., carbodiimide, triglycidylamine, or epoxies; Connolly et al., 2011), and photooxidative preservatives in sheep implants (Moore and Phillips, 1997) with tissue still embedded with devitalized cells results in a marked reduction of calcification (Tam et al., 2017). This further supports the notion that glutaraldehyde products themselves, in conjunction with devitalized cells, may serve as high-affinity sites for calcification. It is also known that glutaraldehyde does not stabilize many components in the tissue such as glycosaminoglycans and elastin. The degradation and depletion of those components can create voids and

high-affinity calcium-binding sites that can aggravate calcification (Isenburg et al., 2005; Raghavan et al., 2010). Novel methods of material fabrication can also include combining methods in more irreversible cross-linking chemistries to produce a more chemically resilient material to biological and chemical insult. Moreover, the discussion above emphasizes that mechanisms of calcification can be complex. Indeed, several material factors can work in concert to cause calcification, and inhibition of one factor alone is not likely to prevent long-term calcification. For instance, carbodiimide can be used with a network extender to provide a more stable, irreversible cross-linking mesh in a tissue-based material than carbodiimide or glutaraldehyde can do alone (Leong et al., 2013). Other current approaches include combination of decellularization/delipidation, endcapping aldehydes, storage in nonglutaraldehyde-based conditions so that calcification of the valve materials could be completely prevented (Meuris et  al., 2018). Additives like pentagalloyl glucose that have been shown to physically stabilize elastin and collagen as well as prevent calcification can be imparted into the implants and stacked together in these fixation techniques (Tedder et  al., 2009; Chuang et al., 2009). These different additives that modularly protect different parts of the base tissue material ECM are being combined to produce superior biomaterials that resist calcification and cyclic wear while maintaining native-like biomechanics (Tam et al., 2017). 

Alternative Materials Synthetic polymers have been investigated as possible candidates for heart valve leaflets. Polyurethane trileaflet valves have been fabricated and studied as a reasonable alternative to bioprostheses or mechanical valve prostheses (Ghanbari et al., 2009). Despite versatile properties, such as superior abrasion resistance, hydrolytic stability, high flexural endurance, excellent physical strength, and acceptable blood compatibility, the use of polymers has been hampered by calcification, thrombosis, tearing, and biodegradation. Although the exact mechanism of polyurethane calcification is as yet unclear, it is

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

987

TABLE 2.4.5.8   Design Considerations for Preventing Calcification of Biomaterials

Material Considerations • Principal chemical bond stability • Chemical bond oxidation potential • High-affinity binding sites for calcium phosphates • Toxicity of residuals and leachables • Surface texture and topography

Material Characterization Techniques • In vitro aqueous hydrolytic degradation • Water and lipid swelling study • In vitro oxidative stress testing • Incubation in supersaturated Ca:P solution • SEM for surface characterization • Uniaxial/biaxial testing • SFDI imaging analysis aThese

Biomechanical Environment Considerations

Biological Environment Considerations

Structural degradation leading to creation of voids, new nucleation sites, or fragmentation of material during cyclic wear and shear flow

Patient factors • Age • Metabolism • Comorbidities Local factors • pH • Oxidative potential

Biomechanical Testing Conditions

Biological Testing Conditions

• Prolonged static or dynamic strain/ stress to study creep/stress relaxation; determines material’s ability to resist permanent geometric deformationsa • Cyclic fatigue for >500M cycles to determine long-term material durabilityb

• In vitro enzymatic challenges that can degrade the material • Cytotoxicity testing with leachables • In vitro/in vivo hemocompatibility testing • Implantation in rodents and small animals as a subcutaneous implant in either healthy animals or under harsher conditions like calciphylaxis or diabetes • Large animal model device implantation

tests can be done in biological testing conditions for more robust model. calcification studies can be performed on fatigued material.

bIn vitro/in vivo

believed that several physical, chemical, and biologic factors (directly or indirectly) play an important role in initiating this pathologic disease process (Schoen et al., 1992c; Joshi et al., 1996; Hyde et al., 1999). A new generation of polyurethanes and other materials are showing promise as an alternative for fabrication of calcification-resistant bioprosthetic heart valves (Bezuidenhout et al., 2015). 

Design Considerations and Selection of Materials to Avoid Calcification Preclinical studies comprise in  vitro (i.e., engineering and materials characterization) and in vivo components (i.e., animal models). For example, in the United States, animal model testing of heart valve substitutes is guided by document ISO5840 from the International Standards Organization, supplemented by appropriate FDA guidance documents, which identifies the goals of preclinical in vivo assessment in evaluation of performance characteristics not amenable to be assessed by in vitro testing (Zhang et al., 2019). Selecting a material candidate appropriate for heart valve replacements remains to this day a difficult task because of the inherent mechanical requirements (hundreds of millions of cycles with high flexural pressure) as well as the short- and long-term failure modes it must resist—namely calcification and/or noncalcific structural deterioration, manifest in vivo as leaflet stiffening and/ or tearing. These two failure modes are not mutually exclusive and have covariate factors that are illustrated in Table 2.4.5.8. This table illustrates the design considerations from material

• Figure 2.4.5.7  Biomaterials calcification triad. The pathophysiology of calcification involves three interrelated factors:

(column 1), biomechanical (column 2), and biological (column 3) perspectives as well as the associated tests used to assess material characteristics. Finding the ideal material candidate for heart valve replacements involves designing around a multitude of criteria that could very quickly narrow down the list of suitable materials that merit further development. The animal model for in vivo studies must be carefully considered, for example, it is well known that juvenile animals of some species (e.g., sheep) have accelerated calcification of tissue heart valve and other biomaterials in device configurations. The design considerations discussed above must be applied to any candidate material that, using appropriate tests, shows it is not likely to calcify. This should be done before investing substantial amounts of resources for further verification or validation testing of new materials in novel configurations (Fig. 2.4.5.7). 

988 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Conclusions Calcification of biomaterial implants is an important pathologic process affecting a variety of tissue-derived biomaterials, as well as synthetic polymers in various functional configurations. The pathophysiology has been partially characterized through the judicious use of animal models. Common features of biomaterial calcification in several clinical applications includes the involvement of devitalized cells and cellular debris, extracellular matrix degradation, and oxidative degradation of materials. Preventive approaches based on modifying biomaterials to make them resistant to degradation due either to biological insult or biomechanical factors appear to be promising. Several anticalcification strategies are already used clinically, particularly for extending the functional lifetime of bioprosthetic heart valves. The long-term data suggest that these new technologies may be efficacious and safe. Further materials development based on the principles described in this chapter may be beneficial and translatable to the clinic.

References Aikawa, E., Nahrendorf, M., Figueiredo, J.L., Swirski, F.K., Shtatland, T., et al., 2007. Osteogenesis associates with inflammation in early-stage atherosclerosis evaluated by molecular imaging in vivo. Circulation 116, 2841–2850. Andrews, J., Psaltis, P.J., Bartolo, B.A.D., Nicholls, S.J., Puri, R., 2018. Coronary arterial calcification: a review of mechanisms, promoters and imaging. Trends Cardiovasc. Med. 28, 491–501. Bailey, M.T., Pillarisetti, S., Xiao, H., Vyavahare, N.R., 2003. Role of elastin in pathologic calcification of xenograft heart valves. J. Biomed. Mater. Res. A 66, 93–102. Baraki, H., Tudorache, I., Braun, M., Höffler, K., Görler, A., Lichtenberg, A., et  al., 2009. Orthotopic replacement of the aortic valve with decellularized allograft in a sheep model. Biomaterials 30 6240-6. Bass, J.K., Chan, G.M., 2006. Calcium nutrition and metabolism during infancy. Nutrition 22, 1057–1066. Baumgartner, H., Falk, V., Bax, J.J., et al., 2017. ESC/EACTS guidelines for the management of valvular heart disease. Eur. Heart J. 38, 2739–2791. Bennink, G., Torii, S., Brugmans, M., Cox, M., Svanidze, O., Ladich, E., Carrel, T., Virmani, R., 2018. A novel restorative pulmonary valved conduit in a chronic sheep model: mid-term hemodynamic function and histologic assessment. J. Thorac. Cardiovasc. Surg. 155 (6), 2591–2601. Bernacca, G.M., Mackay, T.G., Wilkinson, R., Wheatley, D.J., 1997. Polyurethane heart valves: fatigue failure, calcification, and polyurethane structure. J. Biomed. Mater. Res. 34, 371–379. Bernacca, G.M., Wheatley, D.J., 1998. Surface modification of polyurethane heart valves: effects on fatigue life and calcification. Int. J. Artif. Organs 21, 814–819. Bini, A., Mann, K.G., Kudryk, B.J., Schoen, F.J., 1999. Non-collagenous bone proteins, calcification and thrombosis in carcinoid artery atherosclerosis. Arterio Sci. Thromb. Vasc. Biol. 19, 1852– 1861. Bezuidenhout, D., Williams, D.F., Zilla, P., 2015. Polymeric heart valves for surgical implantation, catheter-based technologies and heart assist devices. Biomaterials 36, 6–25.

Bostrom, K., Watson, K.E., Horn, S., Wortham, C., Herman, I.M., et  al., 1993. Bone morphogenetic protein expression in human atherosclerotic lesions. J. Clin. Investig. 91, 1800–1809. Bottio, T., Thiene, G., Pettenazzo, E., Ius, P., Bortolotti, U., et  al., 2003. Hancock II bioprostheses: a glance at the microscope in mid-long-term explants. J. Thorac. Cardiovasc. Surg. 126, 99– 105. Brockhurst, R.J., Ward, P.C., Lou, P., Ormerod, D., Albert, D., 1993. Dystrophic calcification of silicone scleral buckling implant materials. Am. J. Ophthalmol. 115, 524–529. Bucher, P.J., Buchi, E.R., Daicker, B.C., 1995. Dystrophic calcification of an implanted hydroxyethylmethacrylate intraocular lens. Arch. Ophthalmol. 113, 1431–1435. Bueno, E.M., Glowacki, J., 2009. Cell-free and cell-based approaches for bone regeneration. Nat. Rev. Rheumatol. 5, 685–697. Cardoso, L., Weinbaum, S., 2018. Microcalcifications, their genesis, growth, and biomechanical stability in fibrous cap rupture. Adv. Exp. Med. Biol. 1097, 129–155. https://doi.org/10.1007/978-3319-96445-4_7. Review. Carpentier, S.M., Carpentier, A.F., Chen, L., Shen, M., Quintero, L.J., et al., 1997. Calcium mitigation in bioprosthetic tissues by iron pretreatment: the challenge of iron leaching. Ann. Thorac. Surg. 63, 1514–1515. Celiento, M., Ravenni, G., Milano, A.D., Pratali, S., Scioti, G., Nardi, C., et al., 2012. Aortic valve replacement with Medtronic Mosaic bioprosthesis: a 13-year follow-up. Ann. Thorac. Surg. 93, 510–515. Chan, V., Jamieson, W.R., Germann, E., Chan, F., Miyagishima, R.T., et al., 2006. Performance of bioprostheses and mechanical prostheses assessed by composites of valve-related complications to 15 years after aortic valve replacement. J. Thorac. Cardiovasc. Surg. 131, 1267–1273. Chen, W., Kim, J.D., Schoen, F.J., Levy, R.J., 1994a. Effect of 2-amino oleic acid exposure conditions on the inhibition of calcification of glutaraldehyde crosslinked porcine aortic valves. J. Biomed. Mater. Res. 28, 1485–1495. Chen, W., Schoen, F.J., Levy, R.J., 1994b. Mechanism of efficacy of 2-amino oleic acid for inhibition of calcification of glutaraldehyde-pretreated porcine bioprosthetic valves. Circulation 90, 323–329. Chen, J.-H., Liu, C., You, L., Simmons, C.A., 2009. Boning up on Wolff’s Law: mechanical regulation of the cells that make and maintain bone. J. Biomech. 43, 108–118. Christian, A.J., Lin, H., Alferiev, I.S., Connolly, J.M., Ferrari, G., Hazen, S.L., Ischiropoulos, H., Levy, R.J., 2014. The susceptibility of bioprosthetic heart valve leaflets to oxidation. Biomaterials 35, 2097–2102. Christian, A.J., Alferiev, I.S., Connolly, J.M., et al., 2015. The effects of the covalent attachment of 3-(4-hydroxy-3,5-di-tert-butylphenyl)propyl amine to glutaraldehyde pre-treated bovine pericardium on structural degeneration, oxidative modification and calcification of rat subdermal implants. J. Biomed. Mater. Res. 103, 2441–2448. Chuang, T.H., Stabler, C., Simionescu, A., Simionescu, D.T., 2009. Polyphenol-stabilized tubular elastin scaffolds for tissue engineered vascular grafts. Tissue Eng. A 15, 2837–2851. Claiborne, T.E., Slepian, M.J., Hossainy, S., Bluestein, D., 2012. Polymeric trileaflet prosthetic heart valves: evolution and path to clinical reality. Expert Rev. Med. Devices 9, 577–594. Coleman, D., Lim, D., Kessler, T., Andrade, J.D., 1981. Calcification of nontextured implantable blood pumps. Trans. Am. Soc. Artif. Intern. Organs 27, 97–103.

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

Connolly, J.M., Alferiev, I., Clark-Gruel, J.N., Eidelman, N., Sacks, M., Palmatory, E., Kronsteiner, A., Defelice, S., Xu, J., Ohri, R., Narula, N., Vyavahare, N., Levy, R.J., 2005. Triglycidylamine crosslinking of porcine aortic valve cusps or bovine pericardium results in improved biocompatibility, biomechanics, and calcification resistance: chemical and biological mechanisms. Am. J. Pathol. 166 (1), 1–13. Connolly, J.M., Bakay, M.A., Alferiev, I.S., Gorman, R.C., Gorman, J.H., Kruth, H.S., Ashworth, P.E., Kutty, J.K., Schoen, F.J., Bianco, R.W., Levy, R.J., 2011. Triglycidylamine cross-linking combined with ethanol inhibits bioprosthetic heart valve calcification. Ann. Thorac. Surg. 92, 858–865. Courtman, D.W., Pereira, C.A., Kashef, V., McComb, D., Lee, J.M., et  al., 1994. Development of a pericardial acellular matrix biomaterial: biochemical and mechanical effects of cell extraction. J. Biomed. Mater. Res. 28, 655–666. Cunanan, C.M., Cabiling, C.M., Dinh, T.T., et  al., 2001. Tissue characterization and calcification potential of commercial bioprosthetic heart valves. Ann. Thorac. Surg. 71, S417–S421. David, T.E., Armstrong, S., Sun, Z., 1998. The Hancock II bioprosthesis at 12 years. Ann. Thorac. Surg. 66, S95–S98. Deborde, C., Simionescu, D.T., Wright, C., Liao, J., Sierad, L.N., Simionescu, A., 2016. Stabilized collagen and elastin-based scaffolds for mitral valve tissue engineering. Tissue Eng. A 22, 1241–1251. Demer, L.L., 2002. Vascular calcification and osteoporosis: inflammatory responses to oxidized lipids. Int. J. Epidemiol. 31, 737–741. Demer, L.L., Tintut, Y., 2008. Vascular calcification: pathobiology of a multifaceted disease. Circulation 117, 2938–2948. Dohmen, P.M., da Costa, F., Holinski, S., Lopes, S.V., Yoshi, S., et al., 2006. Is there a possibility for a glutaraldehyde-free porcine heart valve to grow? Eur. Surg. Res. 38, 54–61. El-Hamamsy, I., Clark, L., Stevens, L.M., Sarang, Z., Melina, G., Takkenberg, J.J., et  al., 2010. Late outcomes following freestyle versus homograft aortic root replacement: results from a prospective randomized trial. J. Am. Coll. Cardiol. 55, 368–376. Everaerts, F., Gillissen, M., Torrianni, M., et al., 2006. Reduction of calcification of carbodiimide-processed heart valve tissue by prior blocking of amine groups with monoaldehydes. J. Heart Valve Dis. 15, 269–277. Ford-Hutchinson, A.F., Cooper, D.M., Hallgrimsson, B., Jirik, F.R., 2003. Imaging skeletal pathology in mutant mice by microcomputed tomography. J. Rheumatol. 30, 2659–2665. Fukunaga, S., Tomoeda, H., Ueda, T., Mori, R., Aovagi, S., Kato, S., 2010. Recurrent mitral regurgitation due to calcified synthetic chordae. Ann. Thorac. Surg. 89, 955–957. Fyfe, B., Schoen, F.J., 1999. Pathologic analysis of removed nonstented Medtronic Freestyle™ aortic root bioprostheses treated with amino oleic acid (AOA). Semin. Thorac. Cardiovasc. Surg. 11 (4), 151–156. Gallegos, R.P., Nockel, P.J., Rivard, A.L., Bianco, R.W., 2005. The current state of in-vivo pre-clinical models for heart valve evaluation. J. Heart Valve Dis. 14, 423–432. Ghanbari, H., Viatge, H., Kidane, A.G., Burriesci, G., Tavakoli, M., et al., 2009. Polymeric heart valves: new materials, emerging hopes. Trends Biotechnol. 27, 359–367. Goldfarb, R.A., Neerhut, G.J., Lederer, E., 1989. Management of acute hydronephrosis of pregnancy by urethral stenting: risk of stone formation. J. Urol. 141, 921–922. Golomb, G., Schoen, F.J., Smith, M.S., Linden, J., Dixon, M., et al., 1987. The role of glutaraldehyde-induced crosslinks in calcification of bovine pericardium used in cardiac valve bioprostheses. Am. J. Pathol. 127, 122–130.

989

Gong, G., Ling, Z., Seifter, E., Factor, S.M., Frater, R.W., 1991. Aldehyde tanning: the villain in bioprosthetic calcification. Eur. J. Cardiothorac. Surg. 5, 288–299. Gott, J.P., Chih, P., Dorsey, L.M., Jay, J.L., Jett, G.K., et al., 1992. Calcification of porcine valves: a successful new method of antimineralization. Ann. Thorac. Surg. 53, 207–216. Gorna, K., Gogolewski, S., 2002. Biodegradable polyurethanes for implants. II. In  vitro degradation and calcification of materials from poly(ε‐caprolactone)–poly(ethylene oxide) diols and various chain extenders. J. Biomed. Mater. Res. 60, 592–606. Grabenwoger, M., Grimm, M., Ebyl, E., Leukauf, C., Müller, M.M., et al., 1992. Decreased tissue reaction to bioprosthetic heart valve material after L-glutamic acid treatment. A morphological study. J. Biomed. Mater. Res. 26, 1231–1240. Grabenwoger, M., Sider, J., Fitzal, F., Zelenka, C., Windberger, U., et  al., 1996. Impact of glutaraldehyde on calcification of pericardial bioprosthetic heart valve material. Ann. Thorac. Surg. 62 772–62,777. Gümüş, N., 2009. Capsular calcification may be an important factor for the failure of breast implant. Plast. Reconstr. Aesthet. Surg. 62, e606–e608. Habibovic, P., de Groot, K., 2007. Osteoinductive biomaterials – properties and relevance in bone repair. J. Tissue Eng. Regenerat. Med. 1, 25–32. Harasaki, H., Moritz, A., Uchida, N., Chen, J.F., McMahon, J.T., et al., 1987. Initiation and growth of calcification in a polyurethane coated blood pump. Trans. Am. Soc. Artif. Intern. Organs 33, 643–649. Hilbert, S.L., Ferrans, V.J., Tomita, Y., Eidbo, E.E., Jones, M., 1987. Evaluation of explanted polyurethane trileaflet cardiac valve prostheses. J. Thorac. Cardiovasc. Surg. 94, 419–429. Hirsch, D., Drader, J., Thomas, T.J., Schoen, F.J., Levy, J.T., et al., 1993. Inhibition of calcification of glutaraldehyde pretreated porcine aortic valve cusps with sodium dodecyl sulfate: preincubation and controlled release studies. J. Biomed. Mater. Res. 27, 1477–1484. Hjortnaes, J., Butcher, J., Figueiredo, J.-H., Riccio, M., Kohler, R.H., et al., 2010. Arterial and aortic valve calcification inversely correlates with osteoporotic bone remodeling: a role for inflammation. Eur. Heart J. 31, 1975–1984. Hofbauer, L.C., Brueck, C.C., Shanahan, C.M., Schoppet, M., Dobnig, H., 2007. Vascular calcification and osteoporosis – clinical observation towards molecular understanding. Osteoporos. Int. 18, 251–259. Human, P., Zilla, P., 2001a. Inflammatory and immune processes: the neglected villain of bioprosthetic degeneration? J. Long Term Eff. Med. Implant. 11, 199–220. Human, P., Zilla, P., 2001b. The possible role of immune responses in bioprosthetic heart valve failure. J. Heart Valve Dis. 10, 460–466. Hyde, J.A., Chinn, J.A., Phillips Jr., R.E., 1999. Polymer heart valves. J. Heart Valve Dis. 8, 331–339. Isenburg, J.C., Simionescu, D.T., Vyavahare, N.R., 2005. Tannic acid treatment enhances biostability and reduces calcification of glutaraldehyde fixed aortic wall. Biomaterials 26, 1237–1245. Jian, B., Jones, P.L., Li, Q., Mohler 3rd, E.R., Schoen, F.J., et  al., 2001. Matrix metalloproteinase-2 is associated with tenascin-C in calcific aortic stenosis. Am. J. Pathol. 159, 321–327. Jian, B., Narula, N., Li, Q.Y., Mohler 3rd, E.R., Levy, R.J., 2003. Progression of aortic valve stenosis: TGF-beta 1 is present in calcified aortic valve cusps and promotes aortic valve interstitial cell calcification via apoptosis. Ann. Thorac. Surg. 75, 457–465. Jiang, X., Zhao, J., Wang, S., Sun, X., Zhang, X., et al., 2009. Mandibular repair in rats with premineralized silk scaffolds and BMP-2 modified bMSCs. Biomaterials 30, 4522–4532.

990 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Johnson, R.C., Leopold, J.A., Loscalzo, J., 2006. Vascular calcification. Pathobiological mechanisms and clinical implications. Circ. Res. 99, 1044–1059. Johnston, T.P., Webb, C.L., Schoen, F.J., Levy, R.J., 1993. Sitespecific delivery of ethanehydroxy diphosphonate from refillable polyurethane reservoirs to inhibit bioprosthetic tissue calcification. J. Control. Release 25, 227–240. Jones, M., Eidbo, E.E., Hilbert, S.L., Ferrans, V.J., Clark, R.E., 1989. Anticalcification treatments of bioprosthetic heart valves: in vivo studies in sheep. J. Card. Surg. 4, 69–73. Jono, S., McKee, M.D., Murry, C.E., Shiroi, A., Nishizawa, Y., et al., 2000. Phosphate regulation of vascular smooth muscle cell calcification. Circ. Res. 87, e10–e17. Jorge-Herrero, E., Fernández, P., Escudero, C., García-Páez, J.M., Castilo-Olivares, J.L., 1996. Calcification of pericardial tissue pretreated with different amino acids. Biomaterials 17, 571–575. Joshi, R.R., Underwood, T., Frautschi, J.R., Phillips Jr., R.E., Schoen, F.J., et al., 1996. Calcification of polyurethanes implanted subdermally in rats is enhanced by calciphylaxis. J. Biomed. Mater. Res. 31, 201–231 207. Kantrowitz, A., Freed, P.S., Zhou, Y., Mandell, G., DeDecker, P., Riddle, J., et  al., 1995. A mechanical auxiliary ventricle. Histologic responses to long-term, intermittent pumping in calves. AS70 J 41, M340–M345. Kataruka, A., Otto, C.M., 2018. Valve durability after transcatheter aortic valve implantation. J. Thorac. Dis. 10 (Suppl. 30), S3629– S3636. Khan, S.R., Wilkinson, E.J., 1985. Scanning electron microscopy, X-ray diffraction, and electron microprobe analysis of calcific deposits on intrauterine contraceptive devices. Hum. Pathol. 16, 732–738. Kirsch, T., 2008. Determinants of pathologic mineralization. Crit. Rev. Eukaryot. Gene Expr. 18, 1–9. Klintworth, G.K., Reed, J.W., Hawkins, H.K., Ingram, P., 1977. Calcification of soft contact lenses in patient with dry eye and elevated calcium concentration in tears. Investig. Ophthalmol. Vis. Sci. 16, 158–161. Koide, T., Katayama, H., 1979. Calcification in augmentation mammoplasty. Radiology 130, 337–338. Kretlow, J.D., Mikos, A.G., 2007. Mineralization of synthetic polymer scaffolds for bone tissue engineering. Tissue Eng. 13, 927– 938. Kumar, V., Abbas, A., Aster, J.C., 2015. Robbins/Cotran Pathologic Basis of Disease, ninth ed. W. B. Saunders, Philadelphia, PA. Lane, J.I., Randall, J.G., Campeau, N.G., Overland, P.K., McCannell, C.A., Matsko, T.A., et  al., 2001. Imaging of hydrogel episcleral buckle fragmentation as a late complication after retinal reattachment surgery. Am. J. Neuroradiol. 22, 1199–1202. Lee, J.S., Basalyga, D.M., Simionescu, A., Isenburg, J.C., Simionescu, D.T., et al., 2006. Elastin calcification in the rat subdermal model is accompanied by up-regulation of degradative and osteogenic cellular responses. Am. J. Pathol. 168, 490–498. Lee, S., Levy, R.J., Christian, A.J., Hazen, S.L., Frick, N.E., Lai, E.K., Grau, J.B., Bavaria, J.E., Ferrari, G., 2017. Calcification and oxidative modifications are associated with progressive bioprosthetic heart valve dysfunction. J. Am. Heart Assoc. 8, 6. Legrand, A.P., Marinov, G., Pavlov, S., Guidoin, M.F., Famery, R., et al., 2005. Degenerative mineralization in the fibrous capsule of silicone breast implants. J. Mater. Sci. Mater. Med. 16, 477–485. Leong, J., Munnelly, A., Liberio, B., Cochrane, L., Vyavahare, N., 2013. Neomycin and carbodiimide crosslinking as an alternative to glutaraldehyde for enhanced durability of bioprosthetic heart valves. J. Biomater. Appl. 27, 948–960.

Levy, R.J., Schoen, R.J., Levy, J.T., Nelson, A.C., Howard, S.L., et al., 1983a. Biologic determinants of dystrophic calcification and osteocalcin deposition in glutaraldehyde-reserved porcine aortic valve leaflets implanted subcutaneously in rats. Am. J. Pathol. 113, 142–155. Levy, R.J., Schoen, F.J., Howard, S.L., 1983b. Mechanism of calcification of porcine aortic valve cusps: role of T-lymphocytes. Am. J. Cardiol. 52, 629–631. Levy, R.J., Hawley, M.A., Schoen, F.J., Lund, S.A., Liu, P.Y., 1985a. Inhibition by diphosphonate compounds of calcification of porcine bioprosthetic heart valve cusps implanted subcutaneously in rats. Circulation 71, 349–356. Levy, R.J., Wolfrum, J., Schoen, F.J., Hawley, M.A., Lund, S.A., et al., 1985b. Inhibition of calcification of bioprosthetic heart valves by local controlled-released diphosphonate. Science 229, 190–192. Levy, R.J., Schoen, F.J., Sherman, F.S., Nichols, J., Hawley, M.A., et al., 1986. Calcification of subcutaneously implanted type I collagen sponges: effects of glutaraldehyde and formaldehyde pretreatments. Am. J. Pathol. 122, 71–82. Levy, R.J., Schoen, F.J., Lund, S.A., Smith, M.S., 1987. Prevention of leaflet calcification of bioprosthetic heart valves with diphosphonate injection therapy. Experimental studies of optimal dosages and therapeutic durations. J. Thorac. Cardiovasc. Surg. 94, 551–557. Levy, R.J., 1994. Glutaraldehyde and the calcification mechanism of bioprosthetic heart valves. J. Heart Valve Dis. 3, 101–104. Levy, R.J., Schoen, F.J., Flowers, W.B., Staelin, S.T., 1991. Initiation of mineralization in bioprosthetic heart valves: studies of alkaline phosphatase activity and its inhibition by AlCl3 or FeCl3 preincubations. J. Biomed. Mater. Res. 25, 905–935. Levy, R.J., Vyavahare, N., Ogle, M., Ashworth, P., Bianco, R., et al., 2003. Inhibition of cusp and aortic wall calcification in ethanol- and aluminum-treated bioprosthetic heart valves in sheep: background, mechanisms, and synergism. J. Heart Valve Dis. 12, 209–216. Love, J.W., 1993. Autologous Tissue Heart Valves. R. G. Landes, Austin, TX. Lovekamp, J.J., Simionescu, D.T., Mercuri, J.J., Zubiate, B., Sacks, M.S., Vyavahare, N.R., 2006. Stability and function of glycosaminoglycans in porcine bioprosthetic heart valves. Biomaterials 27, 1507–1518. MacLean, K.D., Apel, A., Wilson, J., Werner, L., 2015. Calcification of hydrophilic acrylic intraocular lenses associated with intracameral air injection following DMEK. J. Chem. Res., Synop. 41, 1310–1314. Mako, W.J., Vesely, I., 1997. In vivo and in vitro models of calcification in porcine aortic valve cusps. J. Heart Valve Dis. 6, 316–323. Manji, R.A., Zhu, L.F., Nijjar, N.K., et  al., 2006. Glutaraldehydefixed bioprosthetic heart valve conduits calcify and fail from xenograft rejection. Circulation 114, 318–327. Maranto, A.R., Schoen, F.J., 1988. Alkaline phosphatase activity of glutaraldehyde-treated bovine pericardium used in bioprosthetic heart valves. Circ. Res. 63, 844–848. Meuris, B., De Praetere, H., Strasly, M., Trabucco, P., Lai, J.C., et al., 2018. A novel tissue treatment to reduce mineralization of bovine pericardial heart valves. J. Thorac. Cardiovasc. Surg. 156, 197–206. Mitchell, R.N., Jonas, R.A., Schoen, F.J., 1998. Pathology of explanted cryopreserved allograft heart valves: comparison with aortic valves from orthotopic heart transplants. J. Thorac. Cardiovasc. Surg. 115, 118–127. Mitchell, R.N., Schoen, F.J., 2010. Blood vessels. In: Kumar, V., Fausto, N., Aster, J.C., Abbas, A. (Eds.), Robbins/Cotran Pathologic Basis of Disease, eighth ed. W.B. Saunders, Philadelphia, pp. 487–528. Moore, Phillips, R.E., 1997. Biocompatibility and immunologic properties of pericardial tissue stabilized by dye-mediated photooxidation. J. Heart Valve Dis. 6, 307–315.

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

Munnelly, A.E., Cochrane, L., Leong, J., Vyavahare, N.R., 2012. Porcine vena cava as an alternative to bovine pericardium in bioprosthetic percutaneous heart valves. Biomaterials 33, 1–8. Myers, D.J., Nakaya, G., Girardot, G.M., Christie, G.W., 1995. A comparison between glutaraldehyde and diepoxide-fixed stentless porcine aortic valves: biochemical and mechanical characterization and resistance to mineralization. J. Heart Valve Dis. 4, S98–S101. Nakanome, S., Watanabe, H., Tanaka, K., Tochikubo, T., 2008. Calcification of Hydroview H60M intraocular lenses: aqueous humor analysis and comparisons with other intraocular lens materials. J. Cataract Refract. Surg. 34, 80–86. Neues, F., Epple, M., 2008. X-ray microcomputer tomography for the study of biomineralized endo- and exoskeletons of animals. Chem. Rev. 108, 4734–4741. Neuhann, M., Kleinmann, G., Apple, D.J., 2008. A new classification of calcification of intraocular lenses. Ophthalmology 115, 73–79. New, S.E., Aikawa, E., 2011. Molecular imaging insights into early inflammatory stages of arterial and aortic valve calcification. Circ. Res. 108, 1381–1391. Patai, K., Berényi, M., Sipos, M., Noszál, B., 1998. Characterization of calcified deposits on contraceptive intrauterine devices. Contraception 58, 305–308. Pawade, T.A., Newby, D.E., Dweck, M.R., 2015. Calcification in aortic stenosis: the skeleton key. J. Am. Coll. Cardiol. 66, 561–577. Peacock, M., 2010. Calcium metabolism in health and disease. Clin. J. Am. Soc. Nephrol. 5, S23–S30. Persy, V., D’Haese, P., 2009. Vascular calcification and bone disease: the calcification paradox. Trends Mol. Med. 15, 405–416. Peters, W., Pritzker, K., Smith, D., Fornasier, V., Holmyard, D., et al., 1998. Capsular calcification associated with silicone breast implants: incidence, determinants, and characterization. Ann. Plast. Surg. 41, 348–360. Peters, W., Smith, D., 1995. Calcification of breast implant capsules: incidence, diagnosis, and contributing factors. Ann. Plast. Surg. 34, 8–11. Pettenazzo, E., Valente, M., Thiene, G., 2008. Octanediol treatment of glutaraldehyde fixed bovine pericardium: evidence of anticalcification efficacy in the subcutaneous rat model. Eur. J. Cardiothorac. Surg. 34, 418–422. Phogat, J., Rathi, M., Verma, R., Marwah, N., Sachdeva, S., et al., 2017. Calcification of hydrophilic intraocular lenses: laboratory analysis and surgical technique for intraocular lens exchange. MSJCRS Online Case Reports 5, 64–66. Raghavan, D., Shah, S.R., Vyavahare, N.R., 2010. Neomycin fixation followed by ethanol pretreatment leads to reduced buckling and inhibition of calcification in bioprosthetic valves. J. Biomed. Mater. Res. Part B, Appl. Biomater. 92, 168–177. Rahimi, M., Azimi, A., Hosseinzadeh, M., 2018. Intraocular lens calcification: clinico-pathological report of two cases and literature review. J. Ophthalmic Vis. Res. 13, 195–199. Ramin, Z., Jen-Chieh, T., Robert, F., Jeffrey, M., et al., 2019. Effect of stent crimping on calcification of transcatheter aortic valves. Interact. Cardiovasc. Thorac. Surg. 29(1), 64–73. Rimmer, T., Hawkesworth, N., Kirkpatrick, N., Price, N., Manners, R., et  al., 2010. Calcification of Hydroview lenses implanted in the United Kingdom during 2000 and 2001. Eye 24, 199–200. Rousselle, SD., Wicks, JR., Tabb, BC., Price, N., Tellez, A.O’Brien, M, 2019. Histology strategies for medical implants and interventional device studies. Toxicol Pathol 47, 235–249. Sacks, M.S., Schoen, F.J., 2002. Collagen fiber disruption occurs independent of calcification in clinically explanted bioprosthetic heart valves. J. Biomed. Mater. Res. 62, 359–371.

991

Schlieper, G., Krűger, T., Djuric, Z., Damjanovic, T., Markovic, N., et  al., 2008. Vascular access calcification predicts mortality in hemodialysis patients. Kidney Int. 74, 1582–1587. Schoen, F.J., 1998. Pathologic findings in explanted clinical bioprosthetic valves fabricated from photooxidized bovine pericardium. J. Heart Valve Dis. 7, 174–179. Schoen, Edwards, W.D., 2001. Pathology of cardiovascular interventions, including endovascular therapies, revascularization, vascular replacement, cardiac assist/replacement, arrhythmia control and repaired congenital heart disease. In: Silver, M.D., Gotlieb, A.I., Schoen, F.J. (Eds.), Cardiovascular Pathology, third ed. W. B. Saunders, Philadelphia, PA, pp. 678–723. Schoen, F.J., Kujovich, J., Webb, C.L., Levy, R.J., 1987. Chemically determined mineral content of explanted porcine aortic valve bioprostheses: Correlation with radiographic assessment of calcification and clinical data. Circulation 76, 1061–1066. Schoen, F.J., Levy, R.J., 1999. Tissue heart valves: current challenges and future research perspectives. J. Biomed. Mater. Res. 47, 439– 465. Schoen, F.J., Levy, R.J., 2005. Calcification of tissue heart valve substitutes: progress toward understanding and prevention. Ann. Thorac. Surg. 79, 1072–1080. Schoen, F.J., Levy, R.J., Nelson, A.C., Bernhard, W.F., Nashef, A., et al., 1985. Onset and progression of experimental bioprosthetic heart valve calcification. Lab. Investig. 52, 523–532. Schoen, F.J., Tsao, J.W., Levy, R.J., 1986. Calcification of bovine pericardium used in cardiac valve bioprostheses. Implications for mechanisms of bioprosthetic tissue mineralization. Am. J. Pathol. 123, 143–154. Schoen, F.J., Golomb, G., Levy, R.J., 1992a. Calcification of bioprosthetic heart valves: a perspective on models. J. Heart Valve Dis. 1, 110–114. Schoen, F.J., Levy, R.J., Hilbert, S.L., Bianco, R.W., 1992b. Antimineralization treatments for bioprosthetic heart valves. Assessment of efficacy and safety. J. Thorac. Cardiovasc. Surg. 104, 1285–1288. Schoen, F.J., Levy, R.J., Piehler, H.R., 1992c. Pathological considerations in replacement cardiac valves. Cardiovasc. Pathol. 1, 29–52. Schoen, F.J., Hirsch, D., Bianco, R.W., Levy, R.J., 1994. Onset and progression of calcification in porcine aortic bioprosthetic valves implanted as orthotopic mitral valve replacements in juvenile sheep. J. Thorac. Cardiovasc. Surg. 108, 880–887. Shang, H., Claessens, S.M., Tian, B., et al., 2017. Aldehyde reduction in a novel pericardial tissue reduces calcification using rabbit intramuscular model. J. Mater. Sci. Mater. Sci. 28, 16. Simon, P., Kasimir, M.T., Seebacher, G., Weigel, G., Ullrich, R., et  al., 2003. Early failure of the tissue engineered porcine heart valve SYNERGRAFT in pediatric patients. Eur. J. Cardiothorac. Surg. 23, 1002–1006. Simionescu, D., 2004. Prevention of calcification in bioprosthetic heart valves: challenges and perspectives. Biol. Ther. 4, 1971– 1985. Srivasta, S.S., Maercklein, P.B., Veinot, J., Edwards, W.D., Johnson, C.M., et  al., 1997. Increased cellular expression of matrix proteins that regulate mineralization is associated with calcification of native human and porcine xenograft bioprosthetic heart valves. J. Clin. Investig. 5, 996–1009. Stachelek, S.J., Alferiev, I., Choi, H., Chan, C.W., Zubiate, B., et al., 2006. Prevention of oxidative degradation of polyurethane by covalent attachment of di-tert-butylphenol residues. J. Biomed. Mater. Res. A 78, 653–661. Steitz, S.A., Speer, M.Y., McKee, M.D., Liaw, L., Almeida, M., et al., 2002. Osteopontin inhibits mineral deposition and promotes regression of ectopic calcification. Am. J. Pathol. 161, 2035–2046.

992 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Sun, W., Sacks, M., Fulchiero, G., Lovekamp, J., Vyavahare, N., et al., 2004. Response of heterograft heart valve biomaterials to moderate cyclic loading. J. Biomed. Mater. Res. A 69, 658–669. Tam, H., Zhang, W., Infante, D., Parchment, N., Sacks, M., Vyavahare, N., 2017. Fixation of bovine pericardium-based tissue biomaterial with irreversible chemistry improves biochemical and biomechanical properties. J. Cardiovasc. Transl. Res. 10, 194–205. Tedder, M.E., Liao, J., Weed, B., Stabler, C., Zhang, H., Simionescu, A., et al., 2009. Stabilized collagen scaffolds for heart valve tissue engineering. Tissue Eng. A 15, 1257–1268. Tew, W.P., Mahle, C., Benavides, J., Howard, J.E., Lehninger, A.L., 1980. Synthesis and characterization of phosphocitric acid, a potent inhibitor of hydroxylapatite crystal growth. Biochemistry 19, 1983–1988. Thubrikar, M.J., Deck, J.D., Aouad, J., Nolan, S.P., 1983. Role of mechanical stress in calcification of aortic bioprosthetic valves. J. Thorac. Cardiovasc. Surg. 86, 115–125. Tod, T.J., Dove, J.S., 2016. The association of bound aldehyde content with bioprosthetic tissue calcification. J. Mater. Sci. Mater. Med. 27, 8. Tomizawa, Y., Takanashi, Y., Noishiki, Y., Nishida, H., Endo, M., et  al., 1998. Evaluation of small caliber vascular prostheses implanted in small children: activated angiogenesis and accelerated calcification. Am. Soc. Artif. Intern. Organs J. 44, M496–M500. Trantina-Yates, A.E., Human, P., Zilla, P., 2003. Detoxification on top of enhanced, diamine-extended glutaraldehyde fixation significantly reduces bioprosthetic root calcification in the sheep model. J. Heart Valve Dis. 12, 93–100. Tripi, D.R., Vyavahare, N.R., 2014. Neomycin and pentagalloyl glucose enhanced cross-linking for elastin and glycosaminoglycans preservation in bioprosthetic heart valves. J. Biomater. Appl. 28, 757–766. Umana, E., Ahmed, W., Alpert, M.A., 2003. Valvular and perivalvular abnormalities in end-stage renal disease. Am. J. Med. Sci. 325, 237–242. Vanderbrink, B.A., Rastinehad, A.R., Ost, M.C., Smith, A.D., 2008. Encrusted urinary stents: evaluation and endourologist management. J. Endocrinol. 22, 905–912. Vyavahare, N.R., Chen, W., Joshi, R., Lee, C.-H., Hirsch, D., et al., 1997a. Current progress in anticalcification for bioprosthetic and polymeric heart valves. Cardiovasc. Pathol. 6, 219–229. Vyavahare, N., Hirsch, D., Lerner, E., Baskin, J.Z., Schoen, F.J., et al., 1997b. Prevention of bioprosthetic heart valve calcification by ethanol preincubation. Efficacy and mechanism. Circulation 95, 479–488. Vyavahare, N.R., Hirsch, D., Lerner, E., Baskin, J.Z., Zand, R., et al., 1998. Prevention of calcification of glutaraldehyde-crosslinked porcine aortic cusps by ethanol preincubation: mechanistic studies of protein structure and water–biomaterial relationships. J. Biomed. Mater. Res. 40, 577–585. Vyavahare, N., Ogle, M., Schoen, F.J., Levy, R.J., 1999. Elastin calcification and its prevention with aluminum chloride pretreatment. Am. J. Pathol. 155, 973–982. Vyavahare, N.R., Jones, P.L., Hirsch, D., Schoen, F.J., Levy, R.J., 2000. Prevention of glutaraldehyde-fixed bioprosthetic heart valve calcification by alcohol pretreatment: further mechanistic studies. J. Heart Valve Dis. 9, 561–566. Wang, Q., McGoron, A.J., Bianco, R., Kato, Y., Pinchuk, L., Schoephoerster, R.T., 2010. In-vivo assessment of a novel polymer (SIBS) trileaflet heart valve. J. Heart Valve Dis. 19, 499–505. Weissen-Plenz, G., Nitschke, Y., Rutsch, F., 2008. Mechanisms of arterial calcification: spotlight on the inhibitors. Adv. Clin. Chem. 46, 263–293.

Webb, C.L., Benedict, J.J., Schoen, F.J., Linden, J.A., Levy, R.J., 1988. Inhibition of bioprosthetic heart valve calcification with aminodiphosphonate covalently bound to residual aldehyde groups. Ann. Thorac. Surg. 46, 309–316. Webb, C.L., Schoen, F.J., Flowers, W.E., Alfrey, A.C., Horton, C., et  al., 1991. Inhibition of mineralization of glutaraldehyde-pretreated bovine pericardium by AlCl3. Mechanisms and comparisons with FeCl3 LaCl3 and Ga(NO3)3 in rat subdermal model studies. Am. J. Pathol. 138, 971–981. Wilson, G.J., Courtman, D.W., Klement, P., Lee, J.M., Yeger, H., 1995. Acellular matrix: a biomaterials approach for coronary artery and heart valve replacement. Ann. Thorac. Surg. 60, S353–S358. Xi, T., Ma, J., Tian, W., Lei, X., Long, S., et  al., 1992. Prevention of tissue calcification on bioprosthetic heart valve by using epoxy compounds: a study of calcification tests in vitro and in vivo. J. Biomed. Mater. Res. 26, 1241–1251. Yip, C.Y.Y., Chen, J.-H., Zhao, R., Simmons, C.A., 2009. Calcification by valve interstitial cells is regulated by the stiffness of the extracellular matrix. Arterioscler. Thromb. Vasc. Biol. 29, 936–942. Yu, S.-Y., Viola, F., Christoforidis, J.B., D’Amico, D.J., 2005. Dystrophic calcification of the fibrous capsule around a hydrogel explant 13 years after scleral buckling surgery: capsular calcification of a hydrogel explant. Retina 25, 1104–1107. Zaidi, A.H., Nathan, M., Emani, S., Baird, C., del Nido, P.J., Gauvreau, K., Harris, M., Sanders, S.P., Padera, R.F., 2014. Preliminary experience with porcine intestinal submucosa (CorMatrix) for valve reconstruction in congenital heart disease: histologic evaluation of explanted valves. J. Thorac. Cardiovasc. Surg. 148 2216-4. Zareian, R., Tseng, J.C., Fraser, R., Meganck, J., Kilduff, M., Sarraf, M., Dvir, D., Kheradvar, A., 2019. Effect of stent crimping on calcification of transcatheter aortic valves. Interact. Cardiovasc. Thorac. Surg. 1–10. Zhang, B., Bianco, R.W., Schoen, F.J., 2019. Preclinical assessment of cardiac valve substitutes. Current Status and considerations for engineered tissue heart valves. Front. Cardiovasc. Med. (in press). Zilla, P., Weissenstein, C., Bracher, M., Zhang, Y., Koen, W., et al., 1997a. High glutaraldehyde concentrations reduce rather than increase the calcification of aortic wall tissue. J. Heart Valve Dis. 6, 490–491. Zilla, P., Fullard, L., Trescony, P., et al., 1997b. Glutaraldehyde detoxification of aortic wall tissue: a promising perspective for emerging bioprosthetic valve concepts. J. Heart Valve Dis. 6, 510–520. Zilla, P., Weissenstein, C., Human, P., Dower, T., von Oppell, U.O., 2000. High glutaraldehyde concentrations mitigate bioprosthetic root calcification in the sheep model. Ann. Thorac. Surg. 70, 2091–2095. Zilla, P., Bezuidenhout, D., Torrianni, M., Hendriks, M., Human, P., 2005. Diamine-extended glutaraldehyde- and carbodiimide crosslinks act synergistically in mitigating bioprosthetic aortic wall calcification. J. Heart Valve Dis. 14, 538–545.

Further Reading Anderson, H.C., 1988. Mechanisms of pathologic calcification. Rheum. Dis. Clin. N. Am. 14, 303–319. Anderson, H.C., 1989. Mechanism of mineral formation in bone. Lab. Investig. 60, 320–330. Arora, S., Ramm, C.J., Misenheimer, J.A., Vavalle, J.P., 2017. Early transcatheter valve prosthesis degeneration and future ramifications. Cardiovasc. Diagn. Ther. 7, 1–3. Barrere, F., van Blitterswijk, C.A., de Groot, K., 2006. Bone regeneration: molecular and cellular interactions with calcium phosphate ceramics. Int. J. Nanomed. 1, 317–332.

CHAPTER 2.4.5   Pathological Calcification of Biomaterials

Bechtel, J.F., Muller-Steinhardt, M., Schmidtke, C., Bruswik, A., Stierle, U., et  al., 2003. Evaluation of the decellularized pulmonary valve homograft (SynerGraft). J. Heart Valve Dis. 12, 734–739. Bonucci, E., 1987. Is there a calcification factor common to all calcifying matrices? Scanning Microsc. 1, 1089–1102. Cheng, P.T., 1988. Pathologic calcium phosphate deposition in model systems. Rheum. Dis. Clin. N. Am. 14, 341–351. Courtman, D.W., Pereira, C.A., Omar, S., Langdon, S.E., Lee, J.M., et  al., 1995. Biomechanical and ultrastructural comparison of cryopreservation and a novel cellular extraction of porcine aortic valve leaflets. J. Biomed. Mater. Res. 29, 1507–1516. Deutsch, M.-A., Mayr, N.P., Assmann, G., Will, A., Krane, M., Piazza, N., et  al., 2015. Structural valve deterioration 4 years after transcatheter aortic valve replacement. Circulation 131, 682. Discher, D.E., Mooney, D.J., Zandstra, P.W., 2009. Growth factors, matrices, and forces combine and control stem cells. Science 324, 1673–1677. Everaerts, F., Torrianni, M., van Luyn, M.J., Van Wachem, P.B., Feijen, J., et  al., 2004. Reduced calcification of bioprostheses, cross-linked via an improved carbodiimide based method. Biomaterials 25, 5523–5530. Flameng, W., Ozaki, S., Meuris, B., Herijgers, P., Yperman, J., et al., 2001. Antimineralization treatments in stentless porcine bioprostheses. An experimental study. J. Heart Valve Dis. 10, 489–494. Fradet, G., Bleese, N., Busse, E., Jamieson, E., Raudkivi, P., et  al., 2004. The mosaic valve clinical performance at seven years: results from a multicenter prospective clinical trial. J. Heart Valve Dis. 13, 239–247. Gertz, S.D., Kurgan, A., Eisenberg, D., 1988. Aneurysm of the rabbit common carotid artery induced by periarterial application of calcium chloride in vivo. J. Clin. Investig. 81, 649–656. Giachelli, C.M., 1999. Ectopic calcification: gathering hard facts about soft tissue mineralization. Am. J. Pathol. 154, 671–675. Gott, J.P., Girardot, M.N., Girardot, J.M., Hall, J.D., Whitlark, J.D., et al., 1997. Refinement of the alpha aminooleic acid bioprosthetic valve anticalcification technique. Ann. Thorac. Surg. 64, 50–58. Grunkemeier, G.L., Jamieson, W.R., Miller, D.C., Starr, A., 1994. Actuarial versus actual risk of porcine structural valve deterioration. J. Thorac. Cardiovasc. Surg. 108, 709–718. Harbaoui, B., Courand, P.-Y., Schmitt, Z., Farhat, F., Dauphin, R., Lantelme, P., 2016. Early Edwards SAPIEN valve degeneration after transcatheter aortic valve replacement. JACC Cardiovasc. Interv. 9, 198. Hendriks, M., Eveaerts, F., Verhoeven, M., 2001. Alternative fixation of bioprostheses. J. Long Term Eff. Med. Implant. 11, 163–183. Human, P., Bezuidenhout, D., Torrianni, M., Hendriks, M., Zilla, P., 2002. Optimization of diamine bridges in glutaraldehyde treated bioprosthetic aortic wall tissue. Biomaterials 23, 2099–2103. Johnston, T.P., Webb, C.L., Schoen, F.J., Levy, R.J., 1992. Assessment of the in vitro transport parameters for ethanehydroxy diphosphonate through a polyurethane membrane. A potential refillable reservoir drug delivery device. Am. Soc. Artif. Intern. Organs Trans. 38, M611–M616. Langer, R., Vacanti, J.P., 1993. Tissue engineering. Science 260, 920–926. Lentz, D.L., Pollock, E.M., Olsen, D.B., Andrews, E.J., 1982. Prevention of intrinsic calcification in porcine and bovine xenograft materials. Trans. Am. Soc. Artif. Intern. Organs 28, 494–497.

993

Luo, G., Ducy, P., McKee, M.D., Pinero, G.J., Loyer, E., et al., 1997. Spontaneous calcification of arteries and cartilage in mice lacking matrix GLA protein. Nature 386, 78–81. Mayer Jr., J.E., Shin’oka, T., Shum-Tim, D., 1997. Tissue engineering of cardiovascular structures. Curr. Opin. Cardiol. 12, 528–532. McGonagle-Wolff, K., Schoen, F.J., 1992. Morphologic findings in explanted Mitroflow pericardial bioprosthetic valves. Am. J. Cardiol. 70, 263–264. Meuris, B., Phillips, R., Moore, M.A., Flameng, W., 2003. Porcine stentless bioprostheses: prevention of aortic wall calcification by dye-mediated photooxidation. Artif. Organs 27, 537–543. Ogle, M.F., Kelly, S.J., Bianco, R.W., Levy, R.J., 2003. Calcification resistance with aluminum-ethanol treated porcine aortic valve bioprostheses in juvenile sheep. Ann. Thorac. Surg. 75, 1267–1273. Ong, S.H., Mueller, R., Iversen, S., 2011. Early calcific degeneration of a CoreValve transcatheter aortic bioprosthesis. Eur. Heart J. 33, 586. Parhami, F., Basseri, B., Hwang, J., Tintut, Y., Demer, L.L., 2002. High-density lipoprotein regulates calcification of vascular cells. Circ. Res. 91, 570–576. Pascual, I., Avanzas, P., Morı´s, C., 2017. Degenerative pattern of a percutaneous aortic valve. Rev. Esp. Cardiol. 70, 772. Piatti, F., Sturla, F., Marom, G., Sheriff, J., Claiborne, T.E., et  al., 2015. Hemodynamic and thrombogenic analysis of a trileaflet polymeric valve using a fluid-structure interaction approach. J. Biomech. 48, 3641–3649. Richardt, D., Hanke, T., Sievers, H.H., 2015. Two cases of heart failure after implantation of a CoreValve prosthesis. N. Engl. J. Med. 372, 1079–1081. Schoen, F.J., 2001. Pathology of heart valve substitution with mechanical and tissue prothesis. In: Silver, M.D., Gotlieb, A.I., Schoen, F.J. (Eds.), Cardiovascular Pathology, third ed. W. B. Saunders, Philadelphia, PA, pp. 629–677. Schoen, F.J., Mitchell, R.N., 2015. The heart. In: Kumar, V., Abbas, A., Aster, J.C. (Eds.), Robbins/Cotran Pathologic Basis of Disease, ninth ed. W.B. Saunders, Philadelphia, pp. 523–578. Schoen, F.J., 1988. Biomaterials-associated calcification: pathology, mechanisms, and strategies for prevenion. J. Appl. Biomater. 22, 11–36. Shinoka, T., Ma, P.X., Shum-Tim, D., Breuer, C.K., Cusick, R.A., et  al., 1996. Tissue-engineered heart valves. Autologous valve leaflet replacement study in a lamb model. Circulation 94 II-164–II-168. Speer, M.Y., Giachelli, C.M., 2004. Regulation of vascular calcification. Cardiovasc. Pathol. 13, 63–70. Speer, M.Y., McKee, M.D., Guldberg, R.E., Liaw, L., Yang, H.-Y., et  al., 2002. Inactivation of the osteopontin gene enhances vascular calcification of matrix Gla protein-deficient mice: evidence for osteopontin as an inducible inhibitor of vascular calcification in vivo. J. Exp. Med. 196, 1047–1055. Stock, U.A., Nagashima, M., Khalil, P.N., Nollert, G.D., Herden, T., et al., 1999. Tissue engineered valved conduits in the pulmonary circulation. J. Thorac. Cardiovasc. Surg. 119, 732–740. Thoma, R.J., Phillips, R.E., 1995. The role of material surface chemistry in implant device calcification: a hypothesis. J. Heart Valve Dis. 4, 214–221. van Steenberghe, M., de Vasconcelos, C.-Y., Delay, D., Niclauss, L., Kirsch, M., 2016. Early transcatheter aortic valve degeneration in the young. Int. J. Cardiol. 222, 786–787. Van Wachem, P.B., Brouwer, L.A., Zeeman, R., Dijkstra, P.J., Feijen, J., et al., 2000. In vivo behavior of epoxy-crosslinked porcine heart valve cusps and walls. J. Biomed. Mater. Res. 53, 18–27.

994 SEC T I O N 2. 4     Degradation of Materials in the Biological Environment

Van Wachem, P.B., Brouwer, L.A., Zeeman, R., Dijkstra, P.J., Feijen, J., et  al., 2001. Tissue reactions to epoxy-crosslinked porcine heart valves post-treated with detergents or a dicarboxylic acid. J. Biomed. Mater. Res. 55, 415–423. Wada, T., McKee, M.D., Steitz, S., Giachelli, C.M., 1999. Calcification of vascular smooth muscle cell cultures. Inhibition by osteopontin. Circ. Res. 84, 166–178.

Webb, J.G., Dvir, D., 2015. Is transcatheter aortic valve replacement a durable therapeutic strategy? JACC Cardiovasc. Interv. 8, 1092. Zilla, P., Weissenstein, C., Bracher, M., Human, P., 2001. The anticalcific effect of glutaraldehyde detoxification on bioprosthetic aortic wall tissue in the sheep model. J. Card. Surg. 16, 467–472.

S E C T I ON 2 . 5

Applications of Biomaterials

2.5.1

Introduction to Applications of Biomaterials MICHAEL J. YASZEMSKI 1 , FREDERICK J. SCHOEN 2 , JACK E. LEMONS 3 1Orthopaedic

Surgery and Biomedical Engineering, Mayo Clinic College of Medicine, Rochester, MN,

United States 2Department

of Pathology, Brigham and Women’s Hospital and Harvard Medical School, Boston, MA, United States 3Schools

of Dentistry, Medicine and Engineering, University of Alabama at Birmingham, Birmingham, AL, United States

M

ost students of biomaterials have a strong interest in medical, surgical, or dental applications. Biomaterials are used for the construction of components in an extensive array of devices across a wide range of biomedical disciplines. When considering the applications of biomaterials as a section in Biomaterials Science, Fourth Edition: An Introduction to Materials in Medicine, the primary consideration is on outcomes of treatments. Outcomes are evaluated in terms of the medical discipline involved and the specific biomaterial properties needed for a targeted application to improve the outcomes of patients with specific clinical problems. For example, a total joint replacement has very different considerations than a tooth root replacement, although both anchor in bone for function. Similarly, requirements for a heart valve are very different compared to a vessel replacement or an endovascular stent, although all have extensive blood contact and some biomaterial properties are in common. A central theme is the generation and use of design criteria based on desired functionality, potentially deleterious biomaterials–tissue interaction mechanisms, pathologies of the underlying conditions for which the implant is needed, and the basic properties of the various biomaterials available or needing to be developed. It should not be surprising that considerable research and development has led to clinically used devices with active mechanical, electrical, biologic, or mass exchange functions.

Most implants serve their recipients well for extended periods by alleviating the conditions for which they were implanted. Considerable effort is expended in understanding biomaterials–tissue interactions and eliminating patient–device complications (the clinically important manifestations of biomaterials–tissue interactions). Moreover, many patients receive substantial and extended benefit, despite complications. For example, heart valve disease is a serious medical problem affecting over 100,000 people per year in the United States. Patients with aortic stenosis (the most common form of heart valve disease) have a 50% chance of dying within approximately 3 years without surgery. Surgical replacement of a diseased valve leads to an expected survival of 70% at 10 years, a substantial improvement over the natural course. However, of these patients whose longevity and quality of life have clearly been enhanced, approximately 60% will suffer a serious valve-related complication within 10 years after the operation. Thus long-term failure of biomaterials leading to a clinically significant event does not preclude clinical success for a significant duration. The estimated annual expenditure for implanted medical devices in the United States is 100 billion dollars, and the worldwide expenditure is 300 billion dollars. The following chapters that constitute Section 2.5 present a broad spectrum of biomaterials applications and the key properties needed for specific physiological environments. 995

996 SEC T I O N 2. 5     Applications of Biomaterials

Section 2.5 begins with Chapter 2.5.2A, “Cardiovascular Medical Devices: Heart Valves, Pacemakers and Defibrillators, Mechanical Circulatory Support and Other Intracardiac devices,” by Robert F. Padera and Frederick J. Schoen and Chapter 2.5.2B “Cardiovascular medical devices: stents, grafts, stent-grafts” by Michael A. Seidman, Robert F. Padera and Frederick J. Schoen. Multifunction endovascular implants, for example, an aortic stent that includes an aortic valve, can significantly lessen the procedural risk in treating aortic valvular disease when compared to open valve replacement in patients who often have multiple comorbidities. Chapter 2.5.3, “Extracorporeal Artificial Organs and Therapeutic Devices,” by Rei Ukita, Alastair Campbell Richie, Angela Lai and Keith E. Cook, discusses the role of functional biomaterials that provide bodily functions that no longer exist because of disease or trauma. Chapter 2.5.4, “Orthopedic Applications,” by Joshua J. Jacobs and Nadim James Hallab, presents the spectrum of biomaterials that replace musculoskeletal functions, which include joints, cartilaginous surfaces, tendons, ligaments, and bones. Chapter 2.5.5, “Dental Applications,” by David H. Kohn and Jack E. Lemons, discusses the hard and soft tissues whose function is replaced by biomaterials when those tissues’ functions are absent because of normal wear, disease, or trauma. Chapter 2.5.6, “Ophthalmologic Applications,” by Heather Sheardown, Emily Anne Hicks, Ben Muirhead, and Aftab Taiyab, gives an introductory overview on biomaterials used to treat diseases of the eye. Chapter 2.5.7, “Bioelectronic and Neural Implants,” by Jeffrey R. Capadona and Andrew J. Shoffstall, presents the fundamentals of these devices, including the materials used, their selection criteria, the neuroinflammatory response to them, and the brain–computer interface. Chapter 2.5.8, “Burn Dressings and Skin Substitutes,” by Steven T. Boyce, Petra M. Warner, and Philip Hyungjin Chang, discusses the sequence of treatment for burn patients, beginning with debridement/escharotomy, selection of wound dressings, including subatmospheric (“negative”) pressure dressings, management of microbial contamination, fluid balance, and temporary and/or permanent skin substitutes (some of which are polymeric biomaterials), all leading to wound closure. As an example of the advances made in this area by Dr. Boyce’s team, Dr. Yannas’ team and other colleagues working in this field, my medical/surgical colleagues and I treated a wounded soldier with an 81% body surface area burn in a middle east theater hospital. After escharotomy and temporary wound coverage with a degradable polymeric dressing, he was flown directly, nonstop, from the combat zone to the US Army Institute of Surgical Research Burn Center at the San Antonio Military Medical Center for his definitive treatment. He survived; an engineered skin substitute with which he was treated at the Burn Center, and the temporary biodegradable wound

dressings applied at the field hospital, played significant roles in his survival. Chapter 2.5.9, “Description and Definition of Adhesives, and Related Terminology,” by Bryan K. Lawson and Darshan Shah, discusses three distinct and yet related biomaterial categories. Adhesives effect a bond between tissues or between a tissue and a biomaterial that is the same or different from the adhesive biomaterial. In addition, the adhesive biomaterials may link hard tissues together, soft tissues together, or a hard tissue to a soft tissue. Hard tissue adhesives find many applications in both dental and bone applications. Sealants find uses in minimizing blood loss during surgery, after trauma, and/or after surgery. Sutures keep wound edges apposed as the wound, whether it be surgical or traumatic, heals to a point that it can resist failure without the presence of the sutures. Lawson and Shah discuss natural and synthetic sutures, degradable and nondegradable sutures, polymeric and metallic sutures, and the appropriate uses for these myriad biomaterials. Chapter 2.5.10, “Biomaterials for Immunoengineering,” by Susan Thomas, Margaret P. Manspeaker, and Paul A. Archer, describes interactions between the immune system and biomaterials with the goal of inducing the immune system to behave in a desired manner to treat disease. The sections of this chapter include biomaterial use in vaccine development, vaccine efficacy, and increased vaccine performance via biomaterial adjuvants. The chapter also covers biomaterial use for B cell activation and to effect an increased humoral response. Finally, the authors discuss biomaterials use in the targeting and modulation of T cell therapies. Chapter 2.5.11, “Biomaterials-based model systems to study tumor-microenvironment interactions,” by Claudia Fischbach-Teschl, Brittany Elizabeth Schutrum, and Matthew A. Whitman, presents the tumor microenvironment and its importance in understanding cancer behavior. The authors then discuss the development of biomaterial systems to model the function of the tumor microenvironment to learn how best to alter the cancer’s natural behavior in a way that is optimal for the patient. Chapter 2.5.12, “Drug Delivery Systems,” by Danielle Benoit, Marian Adriana Ackun-Farmmer, Kenneth R. Sims Jr., and Clyde Thomas Overby III, discusses the history of controlled drug delivery, and presents the design factors that are important for the development of drug delivery systems. These include pharmacokinetics, targeting, solubility behavior, delivery routes, biomaterial properties, and biomaterial stability. Chapter 2.5.13, “Responsive Polymers in the Fabrication of Enzyme-Based Biosensors,” by John R. Aggas and Anthony Guiseppi-Elie, presents the role of polymers in the design of enzyme-based biosensors: bioconjugation, bioimmobilization, biohosting, biocompatibility, and active transduction. Examples of polymers that fulfill these roles include inherently conductive polymers, responsive hydrogels, polymeric redox mediators, and ferroelectric, piezoelectric, and pyroelectric polymers.

CHAPTER 2.5.1   Introduction to Applications of Biomaterials

The range of tolerable risks of adverse effects varies directly with the medical benefit obtained by the therapy. Benefit and risk go hand in hand, and clinical decisions are made to maximize the ratio of benefit to risk. The tolerable benefit–risk ratio may depend on the type of implant and the medical problem it addresses. Thus more risk can be tolerated with a heart assist device (a life-sustaining implant) than with a prosthetic hip joint (an implant that relieves pain and disability and enhances function), or with a breast implant (an implant with predominantly cosmetic benefit). As an example, total hip arthroplasties (THAs) with metal-on-metal (MoM), or more correctly cobalt alloy-on-cobalt alloy articulating surfaces (the metal stem trunion–metal femoral head junction, and the metal femoral head–metal acetabular component junction) have been used clinically since the 1950s. The history of implant innovation is replete with examples of implants that performed well during the in vitro and in vivo phases of their development, and performed well during the Food and Drug Administration (FDA) clinical studies to determine safety and effectiveness, only to have something unexpected occur after FDA marketing approval and subsequent widespread clinical use. MoM THAs were introduced to general clinical use in the 2000–02 timeframe. After being in use for about 5–7 years after FDA marketing approval, concerns arose regarding both foreign body

997

reactions to metallic wear debris that produced benign “pseudotumors” and neurologic illnesses from cobalt toxicity. Both the United Kingdom and the United States performed product recalls, and warned patients about these previously unknown responses to the cobalt metal debris. Further information regarding this specific topic can be found on the following FDA Website: http://www.fda.gov /MedicalDevices/ProductsandMedicalProcedures/Implan tsandProsthetics/MetalonMetalHipImplants/default.htm. The message for us is that we must constantly be on the lookout for potential complications that may arise both before and after FDA approval to market a new biomaterial implant. We must expect that something unexpected may occur even after a thorough evaluation of an implant leading to marketing approval. MoM THAs, once thought to herald the arrival of the ultimate THA implant, have disappeared (correctly so) from the armamentarium of the total joint reconstructive surgeon. In summary, Section 2.5 explores the most widely used applications of materials in medicine, biology, and artificial organs. The progress made in many of these areas has been substantial. In most cases, the individual chapters describe a device category from the perspective of the clinical need, the spectrum of devices available to the practitioner, the results and complications, and the challenges to the field that must be addressed and solved to optimize success.

2.5.2A

Cardiovascular Medical Devices: Heart Valves, Pacemakers and Defibrillators, Mechanical Circulatory Support, and Other Intracardiac Devices ROBERT F. PADERA, FREDERICK J. SCHOEN Department of Pathology, Brigham and Women’s Hospital and Harvard Medical School, Boston, MA, United States

Introduction Cardiovascular disease continues to be the leading cause of mortality and morbidity in the Western world, resulting in over 800,000 (nearly one-third of ) deaths in the United States each year. Moreover, as the leading cause of death globally, cardiovascular disease kills more than 17 million individuals per year, a number that is expected to grow to more than 23 million by 2030. The most important subtype is coronary heart disease accounting for more than 1 in 7 deaths, killing over 360,000 people a year in the United States. Additionally, over 25,000 persons per year in the United States succumb to valvular heart disease, of which 17,000 have aortic valve disease, a number which is expected to double by 2040 and triple by 2060 (Benjamin et al., 2019). The good news is that the past several decades have witnessed a virtual explosion in the number and scope of innovative surgical and interventional diagnostic and therapeutic procedures for patients with cardiovascular diseases. Data from the National Center for Health Statistics and the American Heart Association indicate that approximately 8 million major cardiac and vascular operations are performed annually in the United States. Concurrent with and integral to the broad application of these surgical and interventional procedures is the use of various prostheses and medical devices composed of highly advanced biomaterials. Data from 2014 show 475,000 percutaneous

coronary interventions (almost all using endovascular bare-metal or drug-eluting stents), 371,000 coronary artery bypass graft procedures, 156,000 cardiac valve procedures, pacemakers, leads, and cardioverter-defibrillators (420,000), and use of many cardiac assist devices, vascular grafts, umbrellas, patches, and other devices (D’Agostino et al., 2018; Benjamin et al., 2019). Thus cardiovascular prostheses and medical devices, and their constituent biomaterials, are of critical importance to the practices of interventional cardiologists and cardiac and vascular surgeons. The number and complexity of devices permit choices among surgical or catheter-based interventional options that optimize short- and long-term patient management. The recognition and understanding of complications of these devices, many of them related to the biomaterials that comprise them, has led to iterative efforts to improve their performance and safety through biomaterials and device research and development. The result has been highly significant improvements, which have been translated into better patient care. The nature, frequency, and pathologic anatomy of their complications, as well as the responsible blood–tissue–biomaterials interaction mechanisms, have been published for widely used devices used for many years, but are less well appreciated for recently introduced or modified devices and those currently in development (Schoen and Gotlieb, 2016; Buja and Schoen, 2016). This chapter and the one following summarize key considerations in cardiovascular medical devices, including the 999

1000 SE C T I O N 2. 5     Applications of Biomaterials

underlying pathology of the conditions they are designed for and used to treat, relevant biomaterials information, and the most important complications that need to be circumvented. The first chapter emphasizes biomaterials and engineering design issues relevant to cardiac valve prostheses, which have been used extensively for approximately six decades and are clinically important; the outcomes and pathological descriptions of complications encountered with many different types of valve prostheses are well known. The chapter also discusses pacemakers and implantable cardioverter-defibrillators (ICDs), implantable cardiac assist devices and artificial hearts, and miscellaneous intracardiac devices, including percutaneous catheter-based techniques to treat cardiovascular disease in a minimally invasive manner, such as septal defect closure devices, and left atrial occlusion devices. The second chapter discusses devices used for vascular repair and replacement (including vascular grafts and endovascular stents and stent grafts), filters to prevent pulmonary embolism, and other catheters and cardiovascular devices that reside outside the heart. 

Heart Valve Function and Valvular Heart Disease The four intracardiac valves play a critical role in assuring unidirectional forward blood flow through the heart. The tricuspid valve allows one-way flow from the right atrium to the right ventricle, and correspondingly the pulmonary valve from the right ventricle to the pulmonary artery, the mitral valve from the left atrium to the left ventricle, and the aortic valve from the left ventricle to the aorta. The heart valves open and close with each cardiac cycle, i.e., approximately once per second, which equates to approximately 40 million times per year and 3 billion times in a 75-year lifetime. Disorders of heart valves can cause stenosis (i.e., obstruction to flow) or regurgitation (i.e., reverse flow across the valve) (Schoen and Mitchell, 2015; Schoen and Butany, 2016). Sometimes, both stenosis and regurgitation are present in the same valve. Some disease processes such as infective endocarditis (infection of a heart valve) can cause rapid (in days) destruction of the affected valve and lead to abrupt heart failure and death, while others such as calcific aortic stenosis can take many decades to develop clinical manifestations. Progress has been made in recent years toward elucidating a conceptual framework that integrates the dynamic functional structure of heart valves from macro- to micro- to ultrastructure, the biomechanical properties, and the pathobiological behavior of the cardiac valves (Ayoub et al., 2017). There are several major forms of valvular heart disease; most involve the aortic and/or the mitral valve. The most common type of valve disease and most frequent indication for valve replacement overall is calcific aortic stenosis—obstruction at the aortic valve secondary to age-related calcification of the cusps of a valve that was previously anatomically normal

(Fig. 2.5.2A.1A) (Carabello and Paulus, 2009). Elucidating the precise mechanisms of calcification of the aortic valve has been elusive, but recent insights have been gained through the study of the effects of continuous and cyclical mechanical forces on the behavior of the valvular interstitial cells and the associated inflammatory cytokine milieu within the valve tissue (Merryman and Schoen, 2013). Calcific nodules form in the valve cusps, which do not allow the valve to fully open, resulting in pressure overload of the left ventricle, which induces hypertrophy (enlargement of the mass) of the walls of this chamber. This condition takes decades to develop and typically produces symptoms in approximately the seventh and eighth decades of life. Although the normal aortic valve has three cusps, 1%–2% of all individuals are born with a bicuspid aortic valve (i.e., with only two cusps), a condition called congenital bicuspid aortic valve (Mathieu et al., 2015). A congenitally bicuspid valve generally functions well initially and into adulthood, but persons who have this condition develop valve dysfunction and thereby symptoms at relatively younger ages— approximately 10 years earlier than in a patient having a valve with three cusps. Aortic regurgitation (also known as insufficiency) is a less frequent but nevertheless important problem, most often caused by dilation of the aortic root. This prevents complete and effective closure of the cusps, allowing backflow across the valve and leading to volume overload of the left ventricle (Goldbarg and Halperin, 2008). Mitral stenosis (Fig. 2.5.2A.1B) has a single predominant cause—chronic rheumatic heart disease—which leads to scarring and stiffening of the mitral leaflets. This condition usually becomes clinically manifest many years or even decades following an episode of acute rheumatic fever secondary to streptococcal pharyngitis (a common form of childhood throat infection) (Chandrashekhar et  al., 2009). Mitral regurgitation, on the other hand, results from many different conditions; the most frequent is myxomatous valve degeneration (also known as floppy mitral valve), in which the strength of the mitral valve tissue is deficient, thereby causing the valve leaflets to deform excessively (Carabello, 2008) (Fig. 2.5.2A.1C). Conditions in which the left ventricle is abnormally dilated and/or scarred and consequently the valve is not supported properly, and infective endocarditis (i.e., infection of the valve), are among the other major causes of mitral regurgitation. Diseases of the right-sided valves (tricuspid and pulmonic) are much less common than those of the left-sided valves. However, in children with congenital heart disease, there is a need for valves in the pulmonary position (in addition to left-sided valves); replacement of valves for congenital anomalies account for approximately 5% of valve replacements (Barnett and Ad, 2009). The major clinical complication of valvular heart disease is cardiac failure secondary to changes in the myocardium induced by pressure or volume overload of the chambers either upstream or downstream of the diseased valve.

CHAPTER 2.5.2A   Cardiovascular Medical Devices

(A)

1001

(C)

(B)

• Figure 2.5.2A.1  Types of heart valve disease. (A) Severe degenerative calcification of a previously ana-

tomically normal tricuspid aortic valve, the predominant cause of aortic stenosis, and the leading form of valvular heart disease. (B) Chronic rheumatic heart disease, manifest as mitral stenosis, viewed from the left atrium. (C) Myxomatous degeneration of the mitral valve, demonstrating hooding with prolapse of the posterior mitral leaflet into the left atrium (arrow). ((A, B) Reproduced by permission from Schoen, F.J., Edwards, W.D., 2001. Valvular heart disease: General principles and stenosis. In: Silver, M.D., Gotlieb, A.I., Schoen, F.J. (Eds.), Cardiovascular Pathology, third ed. Churchill Livingstone, New York. (C) Reproduced by permission from Schoen, F.J., Mitchell, R.N., 2015. The heart. In: Kumar, V., et  al., (Eds.), Robbins Pathologic Basis of Disease, ninth ed. W.B. Saunders, Philadephia.)

Heart valve disease is common and serious, and individuals with its various forms have significant mortality and morbidity. For example, the mortality of nonsurgically treated critical aortic stenosis, the most deleterious functional abnormality, is approximately 50% at 2–3 years following the onset of symptoms (Fig. 2.5.2A.2A); thus this natural history is more severe than that of many cancers. Valve replacement is a highly beneficial therapy for such patients; survival following valve replacement is 50%–70% and serious complication-free survival is approximately 30%–50% at 10–15 years (Rahimtoola, 2003). Operative mortality for aortic and mitral valve replacement is 3% and 6%, respectively. While valve replacement thus provides a substantial improvement over the natural history of disease, patients with artificial valves can suffer complications related to the device (Fig. 2.5.2A.2A). The surgical treatments available for valvular heart disease include replacement of the valve with a prosthesis and repair of the existing abnormal valve tissue to make it functional (Fedak et  al., 2008). According to estimates derived from data collected by the Healthcare Cost and Utilization Project, Agency for Healthcare Research and Quality, in 2005, 36,678 individuals had aortic or mitral valve replacement in the United States and 8669 had valve repairs (Barnett and Ad, 2009). Reconstructive/repair procedures to eliminate

mitral insufficiency and to minimize the severity of rheumatic mitral stenosis are now highly effective and commonplace. A recent survey of practice in the United States showed that 69% of mitral valve operations for mitral regurgitation currently use repair rather than replacement (Gammie et al., 2009). Whenever possible, repair of a valve is preferable over replacement; advantages of repair relate to the elimination of both the risk of prosthesis-related complications and the need for chronic anticoagulation that is required in many patients with substitute valves, and mandatory in patients with mechanical valves. Surgical valve repair is often accompanied by stabilization of the annulus with or without implantation of a prosthetic annuloplasty ring. Unfortunately, repair is usually not possible for most forms of aortic valve disease. When repair is not possible, severe symptomatic valvular heart disease is treated by surgical valve replacement, which comprises excision of part or all of the diseased valve and replacement by a functional substitute. From a design standpoint, the ideal replacement valve would be nonthrombogenic, nonhemolytic, infection resistant, chemically inert, durable, and easily inserted. It would open fully and close quickly and completely, heal appropriately in place, and not be noticed by the patient (for example, it would be noise free) (Harken et al., 1962; Sapirstein and Smith, 2001).

1002 SE C T I O N 2. 5     Applications of Biomaterials

(A) CUMULATIVE SURVIVAL RATE (%)

100 COMPARISON POPULATION

80

AORTIC VALVE (95) REPLACEMENT

60

Due to (approx. 50-50): • Cardiovascular disease • Prosthesis-associated complications

40

NATURAL HISTORY OF AORTIC VALVE DISEASE

20 IMo 1

2

3

4 5 6 YEARS

7

8

9

ALL VALVE-RELATED COMPLICATIONS (%)

(B)

100

AVR

80

p=0.26

60 40 20

Bioprosthesis Mechanical Prosthesis

0 100 80

0 1 2 3 4 5 6 7 8 9 10111213141516 MVR p=0.26

60 40 20 0

Bioprosthesis Mechanical Prosthesis 0 1 2 3 4 5 6 7 8 9 10111213141516 YEARS AFTER VALVE REPLACEMENT



Figure 2.5.2A.2 Outcome following cardiac valve replacement. (A) Survival curves for patients with untreated aortic valve stenosis (natural history of valve disease) and aortic valve stenosis corrected by valve replacement, as compared with an age-matched control population without a history of aortic valve stenosis. The numbers presented in this figure for survival following valve replacement nearly four decades ago remain accurate today. This reflects the fact that improvements in valve substitutes and patient management have been balanced by a progressive trend toward operations on older and sicker patients with associated medical illnesses. (B) Frequency of valve-related complications for mechanical and tissue valves following mitral valve replacement (MVR) and aortic valve replacement (AVR). ((A) Reproduced by permission from Roberts, L., et al., 1976. Long-term survival following aortic valve replacement. Am. Heart J. 91, 311–317. (B) Reproduced by permission from Hammermeister, K., et al., 2000. Outcomes 15 years after valve replacement with a mechanical versus a bioprosthetic valve: Final report of the Veterans Affairs Randomized Trial. J. Am. Coll. Cardiol. 36, 1152–1158.)

Surgical Bioprosthetic and Mechanical Heart Valves The evolution of prosthetic heart valves and related cardiovascular surgical technology was enabled during the first half of the 20th century by multiple key developments, including cardiac catheterization, innovative surgical techniques, cardiopulmonary bypass machines, and the anticoagulant heparin (Chaikoff, 2007). In the late 1950s, stimulated by collaborations established between surgeons and biomedical engineers, innovative procedures and device technology matured in the surgical research laboratory were translated to clinical practice. These developments fostered new opportunities to replace dysfunctional cardiovascular components with biologic or synthetic prostheses. A key step in modern valve replacement technology was the Hufnagel ball valve, designed to be implanted rapidly into the descending thoracic aorta with the use of proximal and distal fixation rings in patients with aortic regurgitation (Butany et al., 2002). However, with this valve, regurgitant flow from the lower body was prevented, but cardiac work was only partially relieved and coronary flow was not improved. Subsequently, cardiac surgeon Dr. Albert Starr and his colleagues, along with a mechanical engineer, Lowell Edwards, fabricated a valve consisting of a stainless-steel cage, a heat-cured Silastic

ball, and a base surrounded by a Teflon fabric sewing cuff, the latter component permitting the surgeon to suture the valve in place orthotopically (i.e., in the anatomically appropriate location within the heart). The three generic components just described, moving part (either synthetic or biologic), superstructure to guide the motion of the moving occluder, and sewing cuff (anchored at the anastomotic site), comprise the key parts of all previous and present surgical heart valve prostheses. Now used extensively for more than a half century, cardiac valve prostheses are a clinically important achievement of biomaterials science and biomedical engineering. Indeed, the prestigious 2007 Lasker Award for Clinical Medical Research was granted to Drs. Albert Starr and Alain Carpentier to recognize the importance of cardiac valve replacement as a major clinical success (Chaikoff, 2007; Lifton, 2007). Starr performed the first successful valve replacement in the heart in 1960 by implanting a caged-ball mechanical valve prosthesis in the mitral position (Starr, 2007). Carpentier fabricated a “bioprosthesis,” combining chemically treated biologic tissue and a mechanical structure to create a tissue-based (though nonliving) heart valve replacement (Carpentier, 2007). Relevant outcome data and pathological descriptions of complications of many different types of valve prostheses are well known (Vongpatanasin et  al., 1996; Bonow et al., 2006; Schoen and Butany, 2016).

CHAPTER 2.5.2A   Cardiovascular Medical Devices

1003

• Figure 2.5.2A.3  Mechanical prosthetic heart valves. (A) Starr-Edwards caged-ball valve. (B) Bjork-Shiley

tilting disk valve. (C) St. Jude Medical bileaflet tilting disk heart valve. (Reproduced by permission from Schoen, F.J., 2001. Pathology of heart valve substitution with mechanical and tissue prostheses. In: Silver, M.D., Gotlieb, A.I., Schoen, F.J. (Eds.), Cardiovascular Pathology, third ed. Churchill Livingstone, New York.)



Figure 2.5.2A.4 Tissue heart valve replacement devices. (A) Hancock porcine valve. (B) CarpentierEdwards bovine pericardial valve. (Reproduced by permission from Schoen, F.J., 2001. Pathology of heart valve substitution with mechanical and tissue prostheses. In: Silver, M.D., Gotlieb, A.I., Schoen, F.J. (Eds.), Cardiovascular Pathology, third ed. Churchill Livingstone, New York.)

The achievements of Starr and Carpentier provided the foundation on which the clinical success of heart valve replacement is built. Today, more than 80,000 valves are replaced each year in the United States and more than 275,000 worldwide. Moreover, devices and techniques for minimally invasive and percutaneous (catheter-based) valve replacement and repair and other interventional techniques are undergoing rapid innovation and development, and there has been exciting progress toward the creation of a living tissue-engineered heart valve replacement. Hundreds of designs of substitute heart valve replacement devices have been explored experimentally and in patients; most have been abandoned because of design and materials deficiencies that manifest in complications that became apparent only in clinical use (Dewall et al., 2000; Edmunds, 2001; Schoen and Butany, 2016). The opening and closing of a prosthetic valve are purely passive, with the moving parts (occluder or disk[s]) responding to changes in pressure and blood flow within the chambers of the heart and great vessels. Today’s cardiac valvular substitutes are of two generic types: mechanical valves and biological tissue valves. The choice of which valve to use in a particular patient is often difficult (Rahimtulla, 2003; Head et al., 2017; Hirji et al., 2018), even though the overall complication rates for mechanical versus bioprosthetic valves is similar over time in both the aortic and mitral positions (Fig. 2.5.2A.2B).

Mechanical prosthetic heart valves (Fig. 2.5.2A.3) are composed of nonphysiologic biomaterials that employ rigid, mobile occluders in a metallic cage (cobalt-chrome or titanium alloy) as in the Bjork-Shiley, Hall-Medtronic, and OmniScience valves, or two carbon hemidisks in a carbon housing as in the St. Jude Medical (the most widely used), CarboMedics CPHV, Medical Carbon Research Institute, or On-X prostheses. Visually, mechanical valves do not resemble the natural heart valves. Today, all mechanical valve occluders are fabricated from pyrolytic carbon. Pyrolytic carbon has high strength, fatigue and wear resistance, and exceptional biocompatibility, including relative thromboresistence. Patients receiving mechanical valves must be treated with lifelong anticoagulation to reduce the risk of thrombosis and thromboembolic events. Having a trileaflet configuration with a central orifice, tissue valves (Fig. 2.5.2A.4) resemble natural valves. The term “bioprosthesis” describes a special type of tissue valve composed of three cusps of tissue derived from animals— most frequently either a porcine (pig) aortic valve or bovine (cow) pericardium—each treated with glutaraldehyde. Glutaraldehyde fixation preserves the tissue, decreases its (already relatively low) immunological reactivity, and kills the cells within the valve tissue. No immunosuppression is required for these xenografts as is required for whole organ transplants (e.g., kidney, liver, or heart). However, since

1004 SE C T I O N 2. 5     Applications of Biomaterials

these valves no longer contain viable cells, the cusps themselves cannot remodel or respond to injury as does normal tissue. Fabricated tissue valve cusps are usually mounted on a metal or plastic stent with three posts (or struts) to simulate the geometry of a native semilunar valve. As with mechanical valves, the base ring is covered by a Dacron- or Teflon-covered sewing cuff to facilitate surgical implantation and healing. The most widely used valve type is the Carpentier-Edwards pericardial valve. Also used occasionally are tissue valves derived from human cadaveric aortic or pulmonary valves with or without the associated vascular conduit (called allografts, or homografts). These valves have good hemodynamic profiles, a low incidence of thromboembolic complications without chronic anticoagulation, and a low reinfection rate following valve replacement for endocarditis (O’Brien et  al., 2001). Several decades ago, when the use of valve allografts began, they were sterilized and/or preserved with chemicals or irradiation; such valves suffered a high rate of leaflet calcification and rupture. Subsequent technical developments have led to the current practice—allografts that are cryopreserved rather than chemically preserved. Freezing is performed with protection from ice crystal formation using dimethyl-sulfoxide. The valves are subsequently stored until use at −196°C in liquid nitrogen. Contemporary allograft valves yield freedom from degeneration and tissue failure equal to or better than those of conventional porcine bioprosthetic valves, but their use is limited by availability, difficulty in obtaining the proper size, and a more complex surgical procedure for implantation. The reliability of a valve prosthesis and its interactions with the host and local tissues play a major role in patient outcome. Four categories of valve-related complications have been most important in limiting success (Fig. 2.5.2A.5): thrombosis and thromboembolism, infection, structural dysfunction (i.e., failure or degeneration of the biomaterials comprising a prosthesis), and nonstructural dysfunction (i.e., miscellaneous complications and modes of failure not encompassed in the previous groups). The major advantages of tissue valves compared to mechanical prostheses are their pseudoanatomic central flow and relative nonthrombogenicity (see later); consequently, patients with tissue valves usually do not require anticoagulant therapy unless they have atrial fibrillation (AF) or another specific propensity to thrombose the valve. Thromboembolic complications are the major cause of mortality and morbidity after cardiac valve replacement with mechanical valves. No synthetic or modified biological surface is as resistant to thrombosis (thromboresistant) as normal unperturbed endothelium. As in the cardiovascular system in general, Virchow’s triad (surface thrombogenicity, hypercoagulability, and locally static blood flow) largely predicts the relative propensity of a device to thrombus formation and location of thrombotic deposits with cardiovascular prostheses (Bennett et al., 2009). Exposure of blood to an artificial surface can induce thrombosis, embolization, and consumption of platelets and plasma coagulation factors, as well as

the systemic effects of activated coagulation, complement products, and platelets. Thus patients who have received mechanical substitute heart valves require lifetime therapeutic anticoagulation with warfarin derivatives, which induces a risk of hemorrhage, is potentially serious, and in some cases is fatal (Vahanian, 2008). Thrombotic deposits forming on valve prostheses can immobilize the occluder or shed emboli to downstream arterial beds (Fig. 2.5.2A.5A–C). Prosthetic valve infection (endocarditis) occurs in 3%–6% of recipients of substitute valves (Fig. 2.5.2A.5D). When endocarditis was the reason for the original valve replacement, the risk is markedly increased. Rates of infection of bioprostheses and mechanical valves are similar. However, since mechanical valve biomaterials cannot themselves become infected, endocarditis on mechanical valves is localized to the prosthesis/tissue junction at the sewing ring, with accompanying tissue destruction in this area (Piper et  al., 2001). While bioprosthetic valve endocarditis can also be localized to the host tissue/prosthesis junction, biological tissue (despite being chemically fixed) can support growth of bacteria and other microorganisms, and thus the cusps are involved in some cases. The most frequent portals of entry include the mouth via dental procedures, urologic infections and interventions, and indwelling catheters; all comprise breaches of a natural mucosal or cutaneous barrier that may release organisms into the blood. Prosthetic valve endocarditis can occur either early (less than 60 days postoperatively) or late (can be years). The microbial etiology of early prosthetic valve endocarditis is dominated by the staphylococcal species Staphylococcus epidermidis and Staphylococcus. aureus, even though prophylactic antibiotic regimens used routinely at the time of implantation are targeted against these microorganisms. The clinical course of early prosthetic valve endocarditis tends to be fulminant. The most common organisms in late prosthetic valve endocarditis are S. epidermidis, S. aureus, Streptococcus viridans, and Enterococci. Prosthetic valve endocarditis is very difficult to eradicate by antibiotics alone, and thus usually necessitates surgical reintervention. Prosthetic valve dysfunction because of materials degradation can necessitate reoperation or cause prosthesisassociated death. Many valve models have been withdrawn from clinical use because of poor durability. Durability considerations vary widely for mechanical valves and bioprostheses, for specific types of each, for different models of a particular prosthesis utilizing different materials or having different design features, and even for the same model prosthesis placed in the aortic rather than the mitral position. Fractures of metallic or carbon components of mechanical valve prostheses are usually catastrophic but are fortunately rare (Fig. 2.5.2A.5E). Contemporary single-leaflet or bileaflet tilting disk valves with pyrolytic carbon occluders and either metallic struts or carbon housing have generally favorable durability. Fractures related to past design defects are noteworthy in two valve cohorts. In one instance, the Bjork-Shiley single-leaflet tilting disk valve was redesigned with the intention of enhancing disk opening and

CHAPTER 2.5.2A   Cardiovascular Medical Devices

1005



Figure 2.5.2A.5 Complications of prosthetic heart valve. (A) Thrombosis on a Bjork-Shiley tilting disk aortic valve prosthesis, localized to outflow strut near minor orifice, a point of flow stasis. (B) Thrombosis of Hancock porcine bioprosthetic valve. (C) Thromboembolic infarct of the spleen (light area on left) secondary to embolus from valve prosthesis. (D) Prosthetic valve endocarditis with large ring abscess (arrow), viewed from the ventricular aspect of an aortic Bjork-Shiley tilting disk aortic valve. (E) Strut fracture of Bjork-Shiley valve, showing valve housing with single remaining strut and adjacent disk. Sites of prior attachment of missing fractured strut designated by arrows. (F) Structural valve dysfunction (manifest as calcific degeneration with cuspal tear) of porcine valve. ((D) Reproduced by permission from Schoen, F.J., 1987. Cardiac valve prostheses: pathological and bioengineering considerations. J. Card. Surg. 2 (65). (A and E) Reproduced by permission from Schoen, F.J., Levy, R.J., Piehler, H.R., 1992. Pathological considerations in replacement cardiac valves. Cardiovasc. Pathol. 1 (29).)

relieving obstruction and thromboembolic complications that occurred with the original and widely used model. The resultant Bjork-Shiley 60- and 70-degree convexoconcave tilting disk valves suffered fractures of the welded metallic outlet strut and separation from the valve, leading

to frequently fatal disk escape (Blot et  al., 2005; Harrison et  al., 2013). Over 80,000 valves of this model were implanted and at least 600 fractured in this manner. The underlying problem was due to the unanticipated consequence of disk closure at a higher velocity and force, causing

1006 SE C T I O N 2. 5     Applications of Biomaterials

100 80 60

AVR Bioprosthesis Mechanical Prosthesis

40 PRIMARY VALVE FAILURE (%)

the overrotation and an excessively hard contact with the metallic outlet strut. When the outlet strut stress exceeded its endurance limit, fatigue fracture occurred, most frequently in the region of the welds anchoring this strut to the housing. In another instance, fractures of carbon valve components (hemidisk or housing) occurred in implanted Edwards (previously Hemex)-Duromedics bileaflet tilting disk valves. At least 46 valves of this type failed in this manner (Mastroroberto et al., 2000). Studies of these explants suggest that valve fracture with leaflet escape resulted from variable combinations of five factors: (1) microporosity in the pyrolytic carbon coating in the leaflets, (2) cavitation bubbles impacting on the carbon surfaces during function, (3) unusual combinations of dimensional tolerances, (4) poor shock-absorbing qualities of the annular tissues in some patients (perhaps due to calcification-induced rigidity), and (5) structural defects in the valve prosthesis induced by fabrication or surgical mishandling. Fractures of carbon components have been encountered only rarely with other carbon bileaflet tilting disk valves, such as the St. Jude Medical valve. In contrast, structural valve failure with dysfunction is frequent and is the major cause of failure of the most widely used bioprostheses (Fig. 2.5.2A.5F). Bioprosthetic valve structural tissue failure usually results in progressive symptomatic deterioration, which requires reoperation (Schoen and Levy, 2005). Within 15 years following implantation, 30%–50% of porcine aortic valves implanted as either mitral or aortic valve replacements require replacement because of primary tissue failure (Fig. 2.5.2A.6). Cuspal calcification (see Chapter 2.4.5) is the major responsible pathologic mechanism, with regurgitation through tears the most frequent failure mode in porcine valves. The more frequently used contemporary bovine pericardial valves also suffer design-related tearing and/or calcification. Calcification is markedly accelerated in younger patients, with children and adolescents having an especially accelerated course (Saleeb et al., 2014). Within the group of complications causing nonstructural failure are those that relate to healing of the valve in the site of implantation, either too little or too much. Inadequate healing can cause paravalvular leaks, which permit reverse flow usually through a relatively small defect at the junction of prosthesis and host tissue when the valve is closed. Paravalvular leaks may be clinically inconsequential, may cause hemolysis (i.e., destruction of red blood cells through mechanical destruction of their membranes by the high shear stresses engendered by blood being forced at high velocity through small spaces), or may cause heart failure through regurgitation. In contrast, overexuberant healing, called tissue overgrowth (or pannus), can block occluder motion or lead to secondary thrombus. Various incremental improvements to valve prostheses are being investigated in preclinical studies and clinical research and implementation. For example, methods are being actively studied and some are being used clinically to prevent calcification of bioprosthetic valves. The confidence engendered by early data that these methods may have extended the durable

p=0.0001

20 0 0 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 100 80 60

MVR Bioprosthesis Mechanical Prosthesis

p=0.0002

40 20 0 0 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 YEARS AFTER VALVE REPLACEMENT



Figure 2.5.2A.6 Frequency of primary valve failure (nonthrombotic valve obstruction or central valvular regurgitation) for mechanical and tissue valves following mitral valve replacement (MVR) and aortic valve replacement (AVR). Cuspal mineralization is the major responsible pathologic mechanism with regurgitation through tears the most frequent failure mode. (Reproduced by permission from Hammermeister, K., et  al., 2000. Outcomes 15 years after valve replacement with a mechanical versus a bioprosthetic valve: Final report of the Veterans Affairs Randomized Trial. J. Am. Coll. Cardiol. 36, 1152–1158.)

lifetime of bioprosthetic valves has led to a dramatic resurgence of their use. Thus as reflected in overall heart valve replacement industry data, innovations in tissue valve technologies and design have stimulated this segment of the market to grow disproportionately in the last decade by expanding indications for tissue valve use and potentially enhanced durability. Thus there has been a trend toward an increasing fraction of tissue valves implanted relative to mechanical valves. Tissue valve use continues to expand, and studies estimate that tissue valves now represent as much as 80% of all substitute heart valves used in some countries (Mohr, 2014; Isaacs et al., 2015; Goldstone et al., 2017). The trend toward increasing use of bioprostheses (relative to mechanical valves) is especially high in older recipients, who generally have diminished rates of calcific failure and in whom anticoagulation carries increased risk of serious hemorrhage. The advent of so-called sutureless valves for aortic valve disease has allowed for quicker implantations using a more minimally invasive surgical approach (Zannis et al., 2012). The surgery involves removing the diseased aortic cusps and deploying the sutureless valve device. While the valves are implanted surgically, they incorporate several design features discussed later for percutaneous transcatheter valves. The 3f Enable valve (Medtronic, Inc., Minneapolis, MN) consists of three equal pieces of equine pericardium mounted

CHAPTER 2.5.2A   Cardiovascular Medical Devices

on a self-expanding Nitinol frame, which contains a polyester flange on the inflow aspect to prevent migration and minimize paravalvular leak. The Perceval S valve (LivaNova, London, UK) contains bovine pericardial cusps mounted within a Nitinol frame. The frame consists of proximal and distal segments with connecting elements to support the valve cusps and allow anchoring to the aortic root. Both of these valves have shown promise clinically for selected patient populations (Chauvette et al., 2018; Fuzellier et al., 2016). Other approaches to provide improved valves include modifications of bioprosthetic valve stent design and tissue-mounting techniques to reduce cuspal stresses, tissue treatment modifying or an alternative to conventional glutaraldehyde pretreatment to enhance durability and postimplantation biocompatibility, and minimally cross-linked autologous pericardial valves. Near-anatomic configuration, central flow trileaflet prostheses using three flexible synthetic polymeric leaflets in an anatomy that resemble the natural aortic valve, may be facilitated by major developments in the technology of polymeric materials, particularly in the thermoplastic polyurethanes (Bezuidenhout et  al., 2015). Although a polymer valve has been used clinically in a cardiac assist device (Leat and Fisher, 1993), durability limitations have been the major concern in an orthotopic (in the natural site) flexible polymeric valve replacement, with preclinical valve failures being marked by thrombosis, tearing, and/or calcification of the cusps (Fishbein et  al., 1975; Claiborne et al., 2012). Sutureless valves provide the opportunity in minimally invasive and conventional aortic valve surgery to minimize aortic cross-clamp time and maximize effective valve area. Performance and safety have been demonstrated for up to 5 years (Meuris et al., 2015; Di Eusanio and Phan, 2015). Scientific and technological progress has stimulated the goal of generating a living valve replacement that would obviate the complications of conventional valve replacement, adapt to changing environmental conditions in the recipient, and potentially grow with a growing patient (Schoen, 2011; Rippel et  al., 2012; Bouten et  al., 2012; Emmert et  al., 2014). The long-term success of a tissueengineered (living) valve replacement will depend on the ability of its viable cellular components (particularly valvular interstitial cells) to assume normal function with the capacity to repair structural injury, remodel the extracellular matrix, and potentially grow in a growing patient. Translational challenges are substantial (Rabkin et al., 2002; Stassen et al., 2017; Zhang et al. 2019). Tissue-engineered heart valves grown as valved conduits from autologous cells (either vascular wall cells or bone marrow-derived mesenchymal stem cells) seeded on biodegradable synthetic polymers (e.g., polyglycolic acid mesh coated in poly-4-hydroxybutarate) grown in vitro have functioned in the pulmonary circulation of growing lambs for up to 5 months (Hoerstrup et  al., 2000; Rabkin et  al., 2002). In some studies, these grafts evolved in  vivo to a specialized layered structure that resembled that of a native semilunar valve. Pulmonary vascular walls fabricated from vascular

1007

wall cells and biodegradable polymer and implanted into very young lambs enlarged proportionally to overall animal growth over a 2-year period (Hoerstrup et al., 2006). Innovative heart valve tissue approaches may be enabled by emerging biomaterials technologies, including 3D bioprinting, multilayer biomaterials, metal mesh scaffolds, and decellularized valves (Cheung et  al., 2015; Lueders et  al., 2014; Zhang et  al., 2015; Alavi and Kheradvar, 2015; Masoumi et al., 2014). To eliminate the need for in vitro cell seeding and culture steps, an alternative tissue-engineering strategy has used a scaffold of either decellularized naturally derived biomaterial (such as animal xenograft or human allograft valve, decellularized sheep intestinal submucosa) or a porous polymer matrix implanted without prior seeding but with the intent of harnessing intrinsic circulating cells to populate and potentially remodel the scaffold (Matheny et al., 2000; Ghazanfari et al., 2015; Boroumand et al., 2018). Tissuederived scaffolds must possess desirable three-dimensional architecture, mechanical properties, and potential adhesion/ migration sites for cell attachment and ingrowth. Nevertheless, decellularized porcine valves implanted in humans had a strong inflammatory response and suffered structural failure (Simon et al., 2003). Work is also being done on cell-seeded, engineered tissue valves for transcatheter implantation (Driessen-Mol et al., 2014). 

Percutaneous Transcatheter Valves and Other Devices Surgical implantation of bioprosthetic and mechanical valves (discussed previously) has a long and proven track record of success, with symptom and quality of life improvements, and enhanced survival. However, a substantial fraction of patients with aortic stenosis, estimated to be 30%–40% overall, is deemed unsuitable for surgical aortic valve replacement because of advanced age, frailty, and often multiple comorbidities (Goldbarg et  al., 2007). However, a highly significant development over the past several decades is a minimally invasive alternative to conventional aortic valve replacement, called transcatheter aortic valve implantation (TAVI), which was initially used clinically in 2002, and has extended the opportunity for effective mechanical correction of valve disease to a potentially large population of otherwise untreatable individuals (Rodes-Cabau, 2012; Genereux et al., 2012; Hamm et al., 2016). Thus TAVI, with its associated novel valve replacement devices, has rapidly become the new standard of care for many (especially elderly) patients who would otherwise be deemed inoperable. In contrast to classical open surgical treatment of heart valve disease, catheter-based valve implantation uses peripheral arterial access into the femoral artery via a catheter passed retrograde up the aorta to the aortic valve where a novel type of valve device is deployed (i.e., expanded); thus this procedure avoids opening the chest. Alternatively, the device can be deployed in the heart via an antegrade and minimally invasive surgical approach that exposes the apex

1008 SE C T I O N 2. 5     Applications of Biomaterials

(A)

(D)

(B)

(E)

(C)

(F)

• Figure 2.5.2A.7  Percutaneous valve replacement technology. (A) The Edwards-Sapien balloon-expand-

able aortic valve replacement designed for percutaneous implantation, constructed from bovine pericardium attached to a stainless-steel stent. A fabric sealing cuff covers the ventricular aspect to prevent leaks between the prosthesis and surrounding tissues. The valve is mechanically crimped onto a valvuloplasty balloon catheter and expanded within the aortic annulus to displace and exclude the stenotic native valve from the circulation. (B) The Sapien valve in situ, which has been deployed within the aortic valve to treat aortic stenosis, viewed from the ascending aorta. (C) Corevalve aortic bioprosthesis, constructed of bovine pericardium attached to a self-expanding nickel-titanium alloy (Nitinol) stent. The ventricular portion has a high radial force to compress the native valve. The midportion is tapered to avoid interference with the coronary arteries. The aortic portion is flared to provide additional fixation against the wall of the ascending aorta. Nitinol can be made soft at cold temperatures allowing the stent to be tightly compressed within a delivery sheath. Once positioned within the native valve the sheath is withdrawn allowing the stent to assume its predetermined shape. There is adequate radial force to compress the native valve. (D) The Corevalve in situ, which has been deployed within the aortic valve to treat aortic stenosis viewed from the ascending aorta. A coronary artery stent that was placed in the left main coronary artery can be seen peeking through the Nitinol stent of the Corevalve at the left of the image. (E) The Melody pulmonary valve is constructed from a bovine jugular venous valve attached with sutures to a platinum-iridium alloy stent. The relatively delicate venous valve functions well in the pulmonary circulation but is too fragile for use in the systemic circulation. Although often referred to as a pulmonary valve, its maximum expanded diameter of 22 mm largely limits its use for surgically constructed right ventricular to pulmonary artery conduits in the pediatric population. (F) A Sapien valve has been deployed within a failing surgically placed bioprosthesis in the mitral position in a “valve-in-valve” application, viewed from the left ventricle. ((A) and (C) reproduced by permission from Schoen, F.J., Webb, J.G., 2008. Prosthetics and the Heart. In: McManus, B.M., Braunwald, E. (Eds.), Atlas of Cardiovascular Pathology for the Clinician, Current Medicine, Philadelphia, 241–256.)

of the left ventricle (called transapical implantation). This approach is favored in patients in whom manipulation of catheters through severely atherosclerotic sites in the femoral artery and aorta might dislodge debris leading to emboli. The delivery strategy involves collapsing the device and placing it within a catheter-based sheath; for balloon expandable devices, they must be collapsed over a balloon. In the case of aortic stenosis, the valve device is deployed between the cusps of the calcified aortic valve, pushing the diseased cusps out of the flow stream (i.e., in contrast to

open surgery, the diseased valve tissue is not removed during TAVI—Fig. 2.5.2A.7B). During TAVI, delivery, positioning, and permanent fixation in the optimal location are critical to procedural success and usually guided by a combination of external imaging modalities. Clinical experience with TAVI is growing rapidly; since the first clinical implantations in 2002, an estimated 300,000+ TAVI procedures have been performed worldwide. Randomized and observational clinical trials from different countries, including observational studies and

CHAPTER 2.5.2A   Cardiovascular Medical Devices

randomized trials, have compared TAVI to classical aortic valve replacement (Smith et al., 2011; Kodali et al., 2012; Kim et  al., 2014). The consensus of these studies is that TAVI is at least as good as classical aortic valve replacement in terms of procedure “success” regarding morbidity and mortality in high-risk patients to at least 2 years (Tice et  al., 2014; Reardon et  al., 2015). TAVI and surgical valve replacement are comparable hemodynamically. Recent data suggest that TAVI may also be an excellent alternative to open surgery in patients with intermediate or lower surgical risk (Reardon et al., 2017; Mack et al., 2019; Popma et al., 2019). The devices used in transcatheter valve implantation have an outer stent-like structure that contains leaflets. The stent holds open a valve annulus and resists the tendency of a valve annulus or diseased native leaflets to recoil following balloon dilation, supports the valve leaflets, and provides the means for seating the prosthesis in the annulus. Tissues used for the valve component include bovine, equine, or porcine pericardium and bovine jugular venous valves. The stents are made from self-expandable stainless steel, platinum-iridium, or other alloys, or shape-memory materials such as nickel-titanium alloys (e.g., Nitinol). Several catheter-based devices are currently in various stages of development and clinical use in the aortic and pulmonary position. The two transcatheter aortic valves with the largest clinical experience are the Edwards-Sapien family of devices (Edwards LifeSciences) and the CoreValve ReValving system (Medtronic). The Sapien device is composed of a balloon-expandable stainless-steel stent that houses a bovine pericardial trileaflet valve (Fig. 2.5.2A.7A and B). There is a polymer skirt circumferentially attached to the stent to reduce paravalvular leaks. The CoreValve device is composed of a self-expandable Nitinol stent that houses a porcine pericardial trileaflet valve (Fig. 2.5.2A.7C and D). For children with failed right-ventricular to pulmonary artery devices used to correct certain types of congenital heart defects, the Medtronic Melody transcatheter pulmonary valve, composed of a balloon-expandable platinumiridium alloy stent that houses a segment of bovine jugular vein with its native venous valve, is threaded from the femoral vein to the inferior vena cava through the right side of the heart (Fig. 2.5.2A.7E) (Lurz et al., 2009). Catheterbased valves may also play a role in the treatment of surgically implanted bioprosthetic valves that are failing due to stenosis or regurgitation in a so-called “valve-in-valve” application in which a new prosthesis is inserted directly into a prior one (Fig. 2.5.2A.7F). Transcatheter valve implantation presents novel challenges (Fishbein et al., 2014). Valved stents are significantly larger than most existing percutaneous cardiac catheters and devices, and thus vascular access is difficult, potential damage along the course of the catheter passage is possible, and dislodging debris that can become emboli is a significant risk. In the aortic position, there is the potential to impede coronary flow, or interfere with anterior mitral leaflet mobility, the conduction system, or the native diseased

1009

leaflets. Stent architecture may also preclude future catheter access to the coronaries for possible interventions. Secure seating within the aortic annulus or a pulmonary conduit and long-term durability of both the stent and the valve tissue are also major challenges. Surgical complications include most frequently paravalvular leak, vascular injury with hemorrhage, and embolic stroke (Fassa et  al., 2013; Van Mieghem et al., 2015). Transcatheter heart valves are also likely susceptible to prosthesis-associated failure modes typical of surgical bioprostheses and unique to their specific design; prosthetic valve endocarditis (Mylotte et al., 2015; Neraqi-Miandoab et al., 2015; Amat-Santos et al., 2015) and structural valve failure due to leaflet calcification and thrombosis are the most frequent complications. A critical unknown in the expanding use of transcatheter techniques is the durability of the prostheses. Many patients with mitral valve disease also are in need of less invasive transcatheter therapies. The first transcatheter mitral valve replacement (TMVR) in a native valve was performed in 2012. However, the complexities and variability of the mitral valve anatomy and its relationship to neighboring structures have resulted in slower progress with this new therapy compared to the rapid uptake that has occurred with transcatheter aortic valve implantation (Patel and Bapat, 2017; Wyler von Ballmoos et  al., 2018). TMVR can be applied to degenerated prosthetic valves and annuloplasty rings or to a wide variety of native mitral valve diseases. In cases of degenerated bioprosthetic valves, annuloplasty ring, and native valve mitral annular calcification, transcatheter heart valves designed for the aortic position can be implanted with high procedural safety and success rates. In the case of native valve mitral regurgitation, the complexities have led to the development of several TMVR systems for native valve disease with different anchoring mechanisms and geometry; all are currently investigational and none are Food and Drug Administration (FDA) approved at this time. Percutaneous mitral valve replacement has been investigated using a device similar to that used for TAVI (Webb et al., 2019). Other percutaneous devices for the treatment of mitral regurgitation are under development or in clinical use. The MitraClip device (Abbott Laboratories, Abbott Park, IL) is a transcatheter-delivered device that reduces mitral regurgitation by fastening the anterior and posterior leaflets together in an edge-to-edge fashion (Panaich and Eleid, 2018). Approved in 2013 by the FDA, the device is composed of a polyester-covered implant that consists of two cobaltchromium metallic arms that can be opened and closed to capture the edges of the anterior and posterior leaflets of the mitral valve (Fig. 2.5.2A.8). When the arms are closed, the leaflets are held together and the valve orifice approximates a “Fig. 2.5.2A.8" with two openings, rather than the single opening of the native valve. The polyester is macroporous, allowing for tissue ingrowth with the goal both of anchoring the device and preventing thrombosis on the foreign material. Clinical studies have been done; a recent clinical trial

1010 SE C T I O N 2. 5     Applications of Biomaterials

of MitraClip versus medical therapy demonstrated safety and efficacy of the device in patients with heart failure due to moderate-to-severe or severe mitral regurgitation (Stone et al., 2018). 

heart’s natural pacemaker, which is located in the right atrium near the junction with the superior vena cava. The impulse spreads through the muscle of both left and right atrial walls, causing depolarization of the cardiac myocytes that results in atrial contraction. The impulse arrives at the atrioventricular (AV) node, which is located in the posterior right atrium enclosed by the ostium of the coronary sinus, the septal leaflet of the tricuspid valve, and the membranous portion of the interatrial septum (called the

Cardiac Arrhythmias The normal cardiac electrical cycle (Fig. 2.5.2A.9A) begins with an impulse initiated by the sinoatrial (SA) node, the (A)

(C)

(B)

(D)

• Figure 2.5.2A.8  The MitraClip device for mitral regurgitation. (A) MitraClip device on delivery apparatus.

(B) Edge-to-edge approximation of the anterior and posterior leaflets of the mitral valve is achieved by deployment of the MitraClip device that is analogous to an Alfieri stitch, thereby creating a double orifice with improved leaflet coaptation. (C) Two MitraClip devices are seen attached to a portion of the mitral valve. The portion of valve and devices were removed surgically during a valve replacement necessitated by worsening mitral regurgitation. The cloth covering has facilitated tissue ingrowth into the device to help passivate the blood-contacting surface and minimize thrombosis. (D) A specimen radiograph shows the structure of the two cobalt-chromium metallic arms in the closed position. ((A) and (B) reproduced with permission from Schoen, F.J., Butany, J., 2016. Cardiac valve replacement and related interventions. In: Buja, L.M., Butany, J. (Eds.), Cardiovascular Pathology, fourth ed. Elsevier, 529–576.)

CHAPTER 2.5.2A   Cardiovascular Medical Devices

triangle of Koch). After a short delay within the AV node, the impulse passes to the bundle of His and into the left and right bundle branches, located in the intraventricular septum. The impulse spreads through the right and left ventricular myocardium causing a wave of myocyte depolarization and thereby coordinated ventricular contraction. The SA and AV nodes and the bundles of His and its right and (A)

1011

left bundle branches are composed of cardiac muscle cells specialized for conduction. Cardiac arrhythmias (Huikuri et  al., 2001) reflect disturbances of either impulse initiation or impulse conduction. Foci of impulse-generating (automatic) cells outside the SA node, called ectopic foci, may initiate cardiac impulses that generate suboptimal ventricular (B)

Pacemaker or implantable cardioverter-defibrillator

Left Atrium Sinoatrial (SA) Node

Pacing Leads Bundle of His Left Atrium

Right Atrium Left Bundle Branch

Atrioventricular (AV)node

Left Ventricle

1

Right Atrium

Left Ventricle

Right Ventricle

2

Right Ventricle Right Bundle Branch

1 Bradyarrhythmia/heart block (disruption of conduction) 2 Tachyarrhythmia (ectopic, irritable focus) (C)

• Figure 2.5.2A.9  Cardiac arrhythmias and device therapy. (A) The normal cardiac electrical cycle show-

ing schematically sites of both conduction blocks and ectopic foci of impulse generation. (B) Schematic demonstrating implantable cardioverter-defibrillator (ICD) lead placement in the right ventricle. (C) Guidant Prizm II DR ICD, introduced to the US market in 2000, and withdrawn in 2005. (D) Transvenous pacing lead placed in right ventricle demonstrating fibrosis of the distal portion of the lead (arrow). (E) Fibrous capsule surrounding pacemaker electrode in the right ventricle. Low-power photomicrograph demonstrating the space previously occupied by the electrode (e), fibrous tissue separating electrode from blood in the right ventricular chamber (between arrows) and extending around the electrode to separate it from myocardium (m), potentially creating a barrier to conduction of the pacing impulse. (F) Micra Transcatheter Pacing System capsule, which is meant to be implanted in the right ventricle as a leadless pacemaker. (G) Photograph depicting failure of the polymer insulation surrounding the wires (blue) of an ICD, allowing the wires to directly contact the myocardium in areas away from the lead tip. ((E) Reproduced by permission from Schoen, F.J., Webb, J.G., 2008. Prosthetics and the Heart. In: McManus, B.M., Braunwald, E. (Eds.), Atlas of Cardiovascular Pathology for the Clinician, Current Medicine, Philadelphia, 241–256.) (Continued)

1012 SE C T I O N 2. 5     Applications of Biomaterials

(D)

(E)

(F)

(G)

Figure 2.5.2A.9 cont’d

contractions. These arrhythmias are usually fast, i.e., tachyarrhythmias, and can result in ventricular fibrillation, which can be fatal. Intrinsic SA node dysfunction also can account for disturbances of impulse initiation. In contrast, disturbances of impulse conduction mainly consist of conduction blocks or reentry. Conduction blocks constitute a failure of propagation of the usual impulse through the specialized muscle as a result of a disease process (such as ischemia or inflammation) or certain drugs. Blocks can be complete (no impulse propagation) or incomplete (impulse propagates more slowly than normal), and can be permanent or transient. Reentry is said to occur when a cardiac impulse traverses a loop of cardiac fibers and reexcites previously excited tissue without a second impulse from the SA node. For patients in whom these cardiac arrhythmias cannot be controlled pharmacologically by antiarrhythmic drugs, two therapeutic options are available: (1) electrical therapy to control the cardiac rhythm, such as direct current cardioversion or implantable devices such as pacemakers and ICDs, and (2) interventional/surgical therapy to remove the affected tissue or interrupt the abnormal pathway such as endocardial resection, cryoablation, or radiofrequency ablation (Halbfass et al., 2018).

Cardiac Pacemakers Cardiac pacemakers are medical devices that provide impulses to the conduction system to initiate contraction. The first cardiac pacemaker was implanted (Atlee and Bernstein, 2001) in 1958 and since then cardiac pacing has become a wellestablished therapeutic tool. The first pacemakers were large (40–200 cm3) by today’s standards (9–45 cm3) and contained few of the features that are standard in current devices. Over a million patients in the United States currently have pacemakers (Fig. 2.5.2A.9B and C) and over 250,000 new permanent pacemakers are implanted each year; pacemaker placement, revision or removal is a commonly performed procedure. Most cardiac pacemakers are implanted in patients over 60 years old but they are also used in children, including infants, when necessary. The most common indications for permanent cardiac pacing are various types of conduction block. Some conduction blocks lead to bradycardia (abnormally low heart rate), while others, predominantly in the left or right bundles, will result in ventricular dyssynchrony and inefficient ventricular contraction in the setting of a normal heart rate. These conduction blocks can result in decreased cardiac output and the signs and symptoms of congestive heart failure, but can be well treated by cardiac pacing.

CHAPTER 2.5.2A   Cardiovascular Medical Devices

Modern cardiac pacing (Kusumoto and Goldschlager, 2002), either temporary or permanent, is achieved by a system of interconnected components consisting of (1) a pulse generator, which includes a power source and circuitry to initiate the electric stimulus and to sense cardiac electrical activity; (2) one or more electrically insulated conductors leading from the pulse generator to the heart, with a bipolar electrode at the distal end of each; and (3) a tissue, or blood and tissue, interface between electrode and adjacent myocardial cells. The pacemaker delivers a small current (2–4 mA) to the myocardium via the electrodes, resulting in depolarization and contraction of the heart. Temporary pacing is most frequently used for patients with acute myocardial infarction that is complicated by cardiac conduction system disturbances that could progress to complete heart block. Leads for temporary cardiac pacing are generally directed transvenously into the apex of the right ventricle and the pulse generator is located outside the body. In the context of cardiac surgery when the epicardium is already exposed, temporary pacing is achieved by placing insulated wires with bare ends to the epicardial surfaces of the atria or ventricles with the leads emerging transthoracically from the anterior chest to permit easy withdrawal. Ultimately, the temporary pacemaker is either replaced by a permanent device or discontinued. Permanent cardiac pacing involves long-term implantation of both pulse generator and electrode leads. The generator, usually made of a titanium alloy, is placed in a tissue pocket beneath the skin on the left anterior chest with the leads advanced transvenously through the left subclavian vein to terminate at the endocardial surface of the heart. The conducting elements are typically made of MP-35N (an alloy of nickel, cobalt, chromium, and molybdenum with excellent strength and corrosion resistance) in a composite with higher electrical conductance materials such as silver or stainless steel; these are typically insulated with an outer coating of silicone and/or urethane. The tips of the electrodes are typically placed within the right atrium and/or right ventricle depending on the pacing modality. A single chamber pacemaker delivers a stimulus based on a programmed timing interval. The pacemaker also senses intrinsic cardiac activity and can be inhibited from providing unnecessary or inappropriate stimuli. This “demand” pacing is valuable in a patient whose problem is intermittent. A dual-chamber pacemaker with electrodes in both the atrium and ventricle delivers the sequential atrial and ventricular signals to approximate the timing of the normal heartbeat. This device also senses intrinsic atrial and ventricular depolarizations and delivers stimuli at the appropriate time to maintain proper synchrony of the chambers. Patients with ventricular conduction delays such as left bundle branch block may suffer from heart failure due to dyssynchrony of ventricular contraction, where the right and left ventricles do not contract simultaneously. Cardiac resynchronization therapy via biventricular pacing is an intervention in which pacing electrodes are placed in the right atrium, right ventricle, and coronary sinus. The

1013

coronary sinus electrode stimulates the lateral wall of the left ventricle to allow for simultaneous excitation of the right and left ventricles, and for more uniform contraction of the entire left ventricle. Cardiac resynchronization therapy has been shown to significantly improve cardiac function in these patients (McAlister et al., 2007). Permanent implantable pacemakers are powered by lithium-iodide batteries with a finite lifespan of 5–8 years, requiring removal and reimplantation of a new device when the battery is exhausted. In fact, the first patient to receive an implantable pacemaker in 1958 required 22 different pulse generators until his death in 2001 at the age of 86. Improving battery technology to allow for longer lifespan would minimize the number of reimplantations that a patient would require and the complications that arise from these procedures. The interface between the electrode and depolarizable myocardial tissue is of critical importance in the proper functioning of the pacemaker (Fig. 2.5.2A.9D). Typically, a layer of nonexcitable fibrous tissue induced by the electrode forms around the tip of the electrode, which is undesirable as it increases the strength of the threshold pacing stimulus required to initiate myocyte depolarization (Fig. 2.5.2A.9E). Strategies to reduce this fibrosis include improved lead designs, and the use of slow, local release of corticosteroids to minimize the thickness of fibrous tissue formed after lead implantation (Mond and Grenz, 2004). The practical point is that, if pulse generator output is not set sufficiently high in the early postimplantation phase, loss of pacing with potentially fatal consequences can result. By contrast, maintaining output at such high levels once thresholds have stabilized greatly shortens battery life. Thus pacemakers with adjustable variations in output have been developed. An ideal endocardial pacing lead should provide stable fixation immediately from the time of implantation, achieve and maintain a minimal threshold for stimulation, maximize sensing, and function reliably for many years. Electrode fixation to the endocardium may be active or passive. In active fixation, the electrode is designed to grasp the endocardial surface to achieve immediate fixation at implantation. A very effective aid to passive fixation is the addition of projecting “tines,” or fins, in the region of the electrode tip. A different approach to improving fixation has been the development of electrodes with porous metal surfaces to foster tissue ingrowth. An endocardial pacemaker lead may require a special design if it is implanted at a particular site. One example is the J-shaped atrial lead, which is curved to facilitate placing the electrode tip in the right atrial appendage, inherently the most stable site for fixation. “Leadless” pacemaker therapy (Della Rocca et al., 2018) is a new technology that has been introduced into clinical practice in the form of two novel devices: the Nanostim Leadless Pacing System (St. Jude Medical) and the Micra Transcatheter Pacing System (Medtronic, Inc.). These devices (Fig. 2.5.2A.9F) consist of a single capsule-like module that contains all of the functions of a traditional single-chamber pacemaker, but is implanted entirely within

1014 SE C T I O N 2. 5     Applications of Biomaterials

the apical aspect of the right ventricle and attached to the interventricular septum via a primary fixation method involving a helical coil or tines. Implantation is performed percutaneously via a transvenous route under fluoroscopic guidance; the devices have a design feature at their proximal end such that they can be recaptured for repositioning during implantation or removal at a later date. The aim of these devices is to provide the same pacing function as a traditional transvenous pacemaker, but without the complications discussed later associated with a large subcutaneous generator pocket and associated long leads traversing the cardiac chambers and valves. 

Implantable Cardioverter-Defibrillators The first implantable cardioverter defibrillator (ICD) was placed in 1980; currently, more than 100,000 ICDs are implanted annually in the United States. The goal of ICDs is to prevent sudden death in patients with certain lifethreatening arrhythmias by resetting the heart’s electrical activity and stimulating a normal cardiac rhythm. ICDs have been shown to revert sustained ventricular tachycardia (abnormally high ventricular rate) and ventricular fibrillation (uncoordinated electrical/myocardial activity) in multiple prospective clinical trials. Benefit in overall mortality has been well documented (The Antiarrhythmics versus Implantable Defibrillators (AVID) Investigators, 1997). A transvenous ICD consists of similar components to a pacemaker, namely a pulse generator and leads for tachydysrhythmia detection and therapy. The pulse generator is a self-powered, self-contained computer with one or two 3.2 V lithium-silver vanadium oxide batteries used to power all components of the system, including aluminum electrolytic storage capacitors. The devices have a service life of 3–5 years, at which point they require removal and implantation of a new device. The lead is generally placed in the right ventricle through a transvenous approach. ICD leads typically contained several coaxial conducting elements that can sense as well as deliver an appropriate shock, each conducting element may be coated in an insulator such as polytetrafluoroethylene, with all of the elements encased within silicone and coated with an outer layer of urethane insulation. The ICD constantly monitors the ventricular rate, and when the rate exceeds a certain value, provides therapy. Current devices will initially provide a short burst of rapid ventricular pacing that terminates some types of ventricular tachyarrhythmias without providing a large shock. This approach can terminate up to 96% of episodes of ventricular tachycardia without the need for a shock. If this pacing fails to break the arrhythmia, the ICD delivers a shock of 10–30 J between the electrode in the right ventricle and the surface of the pulse generator to terminate the dysrhythmic episode. These devices also keep a running record of arrhythmias and treatment results. ICDs are indicated in patients at high risk for ventricular arrhythmias (primary prevention) and in patients who have already had an episode of aborted sudden cardiac death (secondary prevention).

In contrast to the transvenous ICD, a type of ICD termed a subcutaneous ICD (S-ICD) has been developed and is in clinical use for the detection and termination of malignant arrhythmias (Willcox et  al., 2016). The S-ICD uses an extravascular lead that is implanted in the subcutaneous tissue parallel to the sternum just to the left of midline in most cases. The lead is connected to the generator, which is implanted in the midaxillary line on the left side of the thorax. The lead senses a “far field” signal from the electrical activity of the heart in a manner similar to a surface electrocardiogram, and the shock is delivered by the generator in the subcutaneous tissue overlying the heart in a way similar to an external defibrillator. The S-ICD is therefore not in contact with the heart and blood and is in a more stable mechanical environment, with the hope that some of the complications noted later will be reduced. 

Complications of Pacemakers and ICDs These devices share many of the same complications, many of them requiring device removal and replacement. Like many cardiovascular devices, these are life-sustaining technologies and the implications of device failure can be fatal due to a lack of appropriate cardiac pacing (for pacemakers) or inability to sense or deliver appropriate therapy for a lethal arrhythmia (for ICDs). While normal device endof-service from a depleted battery may not be technically considered a device malfunction, it certainly requires device replacement, and may happen prematurely due to increased fibrosis at the lead–tissue interface requiring a higher stimulus threshold. Failures of the hardware, including the battery/capacitor and charge circuit, connectors, and leads, are the most common device malfunctions, with software problems being less prevalent. Some mechanical failures include electrode dislodgment, lead fractures, electrode corrosion, and insulation failure (Fig. 2.5.2A.9G) (Zeitler et al., 2015). Complications of leads may be related to the body of the lead, as distinct from the lead–device pack interface or the electrodes. Several devices and components have been recalled in recent years for these modes of failure (Amin and Ellenbogen, 2010). Lead improvements over the years have included helical coil and multifilament designs to decrease electrical resistance and enhance flexibility and durability (Haqqani and Mond, 2009). In the past, many reports appeared on interference with pacemaker function by devices ranging from electric razors, toothbrushes, and microwave ovens at home to electrosurgical and diathermy apparatus in hospitals. Fortunately, recent generations of cardiac pacemakers have been greatly improved with regard to their resistance to electromagnetic interference. Many complications relate to the interaction of the device biomaterials with the host tissue. These include infection, thrombosis and thromboembolism, myocardial penetration or perforation, pressure necrosis of the skin overlying the pulse generator, and migration or rotation of the pulse generator. Infection is a dreaded complication of implantable devices in general, and this is certainly true for

CHAPTER 2.5.2A   Cardiovascular Medical Devices

pacemakers and ICDs (Joy et al., 2017). The infection may originate in the subcutaneous pocket and track along the lead, which acts as a contaminated foreign body. Alternatively, it may occur by implantation of bacteria on traumatized endocardium or thrombus contiguous with the lead. The most common organisms responsible for these infections are coagulase-negative Staphylococcus species such as S. epidermidis. Septicemia may develop and septic pulmonary emboli may occur. The fundamental therapeutic principle in device-related endocarditis is treatment of the infection with antibiotics followed by removal of at least the lead and, when the pacemaker pocket is involved, the entire pacing system (Baddour et al., 2010). ICDs contain more extensive hardware than pacemakers, and this may contribute an increased relative frequency of complications. Several additional considerations are specific to ICDs. The consequences of repeated defibrillations can cause the following effects: (1) direct effect of repeated discharges on the myocardium and vascular structures, and (2) possible thrombogenic potential of the indwelling intravascular electrodes. Another major complication of ICDs from the standpoint of the patient, other than the inability to sense or terminate an arrhythmia leading to sudden death, is an inappropriate shock. In addition to being startling and quite painful at the time of the shock, patients receiving multiple inappropriate shocks have been known to develop posttraumatic stress disorder symptoms. As mentioned earlier, the leads are designed to optimize their interactions with the adjacent myocardium; this can be problematic when a complication arises in which the leads must be removed. Some leads can be removed by prolonged gentle traction, but many require additional tools and techniques to free them from the venous wall through which the body of the lead travels and from the myocardium to which they are often tenaciously adherent (Krainski et al., 2018). Recourse to cardiotomy with cardiopulmonary bypass may be needed if the lead is densely incarcerated in fibrous tissue. 

Congestive Heart Failure Congestive heart failure (Jessup and Brozena, 2003) is a deficiency of the pumping function of the heart and is an extremely common condition, affecting approximately 6.2 (A)

1015

million Americans. Each year in the United States, congestive heart failure is the principal cause of death in 60,000 individuals, a contributing factor in over 280,000 deaths, and the primary discharge diagnosis in over 1.1 million hospitalizations, all increases over previous years. Cardiac transplantation is a potential solution for some of these patients (Mancini and Lietz, 2010). However, the increasing discrepancy between the number of acceptable donor hearts (only 2500 per year) and the number of patients who might benefit from cardiac transplantation (estimated at greater than 100,000 per year) has prompted efforts toward the development of mechanical devices to augment or replace cardiac function (Baughman and Jarcho, 2007; Boilson et al., 2010; Krishnamani et al., 2010). Congestive heart failure is the final common pathway of many specific cardiac conditions, including valvular heart disease, coronary artery atherosclerosis with resultant ischemic heart disease, and diseases that affect the cardiac muscle directly (termed cardiomyopathies). Heart failure can occur precipitously, as in myocardial infarction or viral myocarditis, or it can be a slow, progressive worsening of exercise tolerance and shortness of breath over many months or years because of ongoing deterioration of the heart muscle. It can manifest itself in the postoperative period after both cardiac surgery (e.g., valve replacement, cardiac transplantation) and noncardiac surgery (e.g., abdominal aortic aneurysm repair). When the left ventricle is failing of whatever etiology, blood backs up into the pulmonary circulation raising the pulmonary vascular resistance and increasing the pulmonary arterial pressure necessary to overcome that resistance. Hence, the pulmonary arterial pressure can be a surrogate marker of the degree of left heart failure in patients with chronic congestive heart failure. The CardioMEMS HF System (Micro-Electrico-Mechanical HF System, Abbott Medical, Inc., Abbott Park, IL) provides hemodynamic information that can be used for the monitoring and management of heart failure (Ayyadurai et al., 2019). The device consists of a wireless sensor that is implanted in the distal pulmonary artery (Fig. 2.5.2A.10) generally via a catheterbased procedure. The sensor consists of a three-dimensional coil and pressure-sensitive capacitor encased between two wafers of fused silica, which is further encased in silicone. (B)

• Figure 2.5.2A.10  CardioMEMS device. (A) Unimplanted CardioMEMS pressure sensor. (B) In situ view of a CardioMEMS pressure sensor implanted in a branch of the pulmonary artery within the lung.

1016 SE C T I O N 2. 5     Applications of Biomaterials

The coil electromagnetically couples the pressure-sensitive capacitor to the electronics system, allowing the measurement of the resonant frequency of the circuit without the need for an implanted battery. The resonant frequency is continuously converted to a pressure measurement, and the pressure waveform (including systolic, diastolic, and mean arterial pressures) and heart rate are transmitted to a receiver either in the hospital or the patient’s home if they are ambulatory. The treating physician can access the data remotely and in real time to evaluate the patient and make any changes to the medical regimen. This strategy can be beneficial in certain patient groups or when used in structured programs (Dickinson et al., 2018). As one might expect therefore, the natural history of heart failure depends on the cause and progression of the underlying disease process. For example, patients with heart failure after cardiac surgery (called postcardiotomy shock) often recover the vast majority of their cardiac function after a short period of time if they are otherwise sustained by mechanical circulatory support. In contrast, patients with dilated cardiomyopathy, one of the most common indications for cardiac transplantation, often need long-term mechanical support; studies have shown that at least 50% of such individuals would die in 3–5 years from their disease without it. One must take these clinical considerations into account when designing mechanical support systems, as different devices may best serve patients with different problems (DiGiorgi et al., 2003).

Cardiopulmonary Bypass First used in 1953 by Dr. John H. Gibbon, cardiopulmonary bypass devices pump blood external to the body and thereby permit complex cardiac surgical procedures to be done safely and effectively. Bypass machines are useful in extracorporeal membrane oxygenation (ECMO) to assist in the transport of oxygen and carbon dioxide for patients (especially neonates and infants) with pulmonary diseases such as the respiratory distress syndrome (Alpard and Zwischenberger, 2002). The basic operating principles of the current heart–lung machines are quite straightforward and have changed little in the past half century. Deoxygenated blood returning from the systemic circulation into the right atrium is withdrawn by gravity siphon into a cardiotomy reservoir and is then pumped into an oxygenator. The most common type of oxygenator is a membrane oxygenator, where oxygen is passed through the tube side of a shell-and-tube-type device while the blood passes through the shell side. Oxygen and carbon dioxide are exchanged via diffusion through synthetic membranes (usually polypropylene or silicone) with high permeability to these respiratory gases. The oxygenated blood is then passed through a heat exchanger to adjust the temperature of the blood and the blood is returned to the systemic circulation via the aorta. At the beginning of the procedure, the patient is anticoagulated with heparin to reduce the risk of thrombosis within the device; as the patient is weaned from bypass, the anticoagulation can be quickly reversed by the use of a drug called protamine.

During an operation, the heat exchanger lowers the temperature of the blood and therefore the core body temperature, decreasing the metabolic requirements of the body and protecting the organs (including the heart) against ischemic damage. At the end of the operation, the blood can be warmed to normal physiologic temperature as the patient is weaned from the bypass machine. A specially trained perfusionist controls the operation of the heart–lung machine, allowing the surgeon and anesthesiologist to concentrate on their respective tasks. This device therefore provides the function of both the heart (maintaining systemic blood flow and pressure) and the lungs (oxygenating blood and removing carbon dioxide), allowing the heart to be effectively stopped for delicate surgical procedures that would be more difficult or impossible to perform on a beating, moving heart. Many improvements to the original design of cardiopulmonary bypass machines have been made since their inception. One of the problems with the original heart–lung machines was the trauma that they would cause to the blood cells. Hemolysis of red blood cells would lead to functional anemia and loss of oxygen-carrying capacity of the blood; damage to platelets would lead to thrombocytopenia (low numbers of or dysfunctional platelets), resulting in bleeding problems. The problem of blood cell damage has been largely overcome with advanced pump designs and the use of the membrane oxygenators. Roller pumps and centrifugal pumps are commonly used because they cause a lesser degree of hemolysis and shear forces; it is important in the design of these pumps to determine the optimum balance between pumping function and hemolysis/shear stress to the formed blood elements. Bubble oxygenators, which directly pass bubbles of oxygen gas through the blood, cause more hemolysis, protein denaturation, and platelet dysfunction than membrane oxygenators and are currently less frequently used. In addition, newer devices allow blood that has escaped the circulation within the sterile operating field around the heart to be processed and returned to the patient, reducing the need for blood transfusion during the procedure. Cardiopulmonary bypass can result in many pathophysiologic changes, including complement activation from the prolonged interaction of blood with synthetic surfaces, platelet and neutrophil activation and aggregation, changes in systemic vascular resistance, and expression of other proinflammatory mediators (Levy and Tanaka, 2003). When these changes are severe, the use of the heart–lung machine can result in complications, including confusion, renal insufficiency, pulmonary dysfunction, low-grade hepatic dysfunction, and increased susceptibility to infection. Together, these manifestations are termed the postperfusion syndrome. The last decade has seen development of mini extracorporeal circuit (MECC) cardiopulmonary bypass systems (Curtis et al., 2010) with a goal of providing cardiopulmonary bypass with a reduction in this harmful systemic inflammatory response. The MECC has a greatly reduced tubing length, smaller priming volumes, reduction in the blood–air interface, and fewer components than the standard bypass systems, and utilizes heparin-coated components and centrifugal blood

CHAPTER 2.5.2A   Cardiovascular Medical Devices

pumps. These systems show a reduction in postoperative cytokine levels, organ damage, postoperative complications, and the need for blood transfusions compared to standard circuits (Vohra et al., 2009).

 Percutaneous Mechanical Circulatory Support Devices Percutaneous ventricular assist devices (VADs) (Mandawat and Rao, 2017; Miller et al., 2017) are used primarily for potentially reversible acute heart failure or cardiogenic shock, in which cardiac function is likely to recover with cardiac rest (e.g., postcardiotomy shock), or in cases of acute heart failure in which the patient needs to be stabilized quickly so that the possibilities for additional lifesaving therapies (e.g., coronary artery bypass surgery, valve surgery, durable VAD placement, heart transplantation) can be assessed (“bridge-to decision”). The typical patient is critically ill with acute cardiogenic shock, mechanical complications of myocardial infarction such as ventricular septal or papillary muscle rupture, unrelenting ventricular arrhythmias, or advanced heart failure. Also, patients undergoing high-risk cardiac surgical procedures or percutaneous revascularization may benefit from the use of devices in the periprocedure period to reduce myocardial oxygen demand. Percutaneous devices (Fig. 2.5.2A.11A) used in these situations include the intraaortic balloon pump (IABP), TandemHeart (Cardiac Assist, Inc., Pittsburgh, PA), Impella (ABIOMED, Inc., Danvers, MA), and ECMO. Since the original use of the IABP in 1968 by Kantrowitz, the basic design and function of the current device has remained relatively similar during the ensuing decades. IABPs (Fig. 2.5.2A.11B) are catheter-based polyethylene or polyurethane balloons with volumes of 25–50 mL, although smaller devices are used in the pediatric population. Helium is most often used as the inflating gas; its low viscosity allows for rapid inflation and deflation and it is rapidly dissolved in the bloodstream in the event of inadvertent balloon rupture. IABPs (Baskett et al., 2002; De Sousa et al., 2010) are generally positioned under fluoroscopic guidance in the descending thoracic aorta after percutaneous insertion via the femoral artery. They are timed to inflate during diastole (ventricular filling) and deflate during systole (ventricular contraction) using the patient’s electrocardiogram or arterial pressure curve for synchronization; the devices discussed later do not have this requirement. This is termed counterpulsation (Trost and Hillis, 2006), which is out of phase with the patient’s heartbeat, and causes volume displacement of blood proximally and distally within the aorta. Several beneficial effects serve to improve cardiac function. Coronary blood flow (the majority of which occurs in diastole) is increased by the rise in diastolic pressure, delivering more oxygenated blood to the myocardium. In addition, left ventricular afterload (the pressure the myocardium must attain to pump blood into the aorta) is decreased, reducing the workload and therefore the oxygen requirement of the myocardium. The combination of these two hemodynamic

1017

factors therefore improves the balance between myocardial oxygen supply and demand, and results in improved cardiac performance. The device also directly improves systemic circulation to a modest degree (approximately 10%). IABP therapy permits the heart to rest and recover enough function to support adequate circulation after the device has been removed, usually after only a few days. The major contraindications for IABP use include severe peripheral vascular disease, including aneurysms, aortic valve regurgitation, and aortic dissection (because of the need to thread the balloon through peripheral arteries and the aorta). Complications, which occurred in approximately 7% of patients with IABPs in a registry study, include limb ischemia from insertion site problems, bleeding, thrombosis with embolization, aortic dissection, balloon rupture, and sepsis. The TandemHeart (Fig. 2.5.2A.11C) is a percutaneous VAD that supports the systemic circulation by withdrawing blood from the left atrium and reinjecting it into the abdominal aorta or iliac artery. The device has been commercially available since 2004. The pump, which sits in an extracorporeal location, is a continuous flow centrifugal pump that can provide up to 4–5 L/min of blood flow. The inflow cannula (usually 21-Fr in size) is typically inserted into a femoral vein, advanced into the right atrium and then pushed across the interatrial septum via transseptal puncture into the left atrium; when the device is withdrawn at the end of use, a small atrial septal defect remains. The outflow cannula (usually 15- to 17-Fr in size) is typically inserted into a femoral artery and advanced into the iliac artery or abdominal aorta. TandemHeart has been successfully used in a wide variety of clinical scenarios, including during high-risk percutaneous coronary interventions, bridge-to-recovery, bridge-to-decision, and bridge-to-transplant for patients with acute and advanced heart failure (Tempelhof et al., 2011; Kar et al., 2006). A modification of the TandemHeart allows it also to be used as a right VAD; a dual lumen cannula sits such that the inflow cannula of the device resides in the right atrium while the outflow cannula sits in the pulmonary artery distal to the pulmonic valve. One complication of the leftsided TandemHeart is the migration of the inflow cannula (meant to be in the left atrium) back into the right atrium. When this occurs, deoxygenated blood from the systemic venous return is drawn into the pump and returned into the systemic circulation, resulting in a right-to-left shunt. In addition, this steals blood from the right heart resulting in decreased pulmonary arterial flow leading to hypoxemic respiratory failure. The most common complication with the TandemHeart is bleeding at the cannula insertion site(s), exacerbated by the need for anticoagulation to prevent pump thrombosis. The Impella device (Fig. 2.5.2A.11D) is a continuous, nonpulsatile axial flow pump that employs an Archimedes-screw impeller that draws blood from the left ventricular cavity and expels blood into the ascending aorta distal to the aortic valve. There are currently three versions of the device for left-sided support. The Impella 2.5 provides 2.5 L/min of flow, the Impella CP provides 3.7 L/min, and the Impella 5.0 provides

1018 SE C T I O N 2. 5     Applications of Biomaterials

(A)

(B)

A

IABP

B Impella

C

Tandem Heart

D

ECMO

(C)

• Figure 2.5.2A.11  Percutaneous mechanical circulatory support devices. (A) The

intraaortic balloon pump (IABP) is inserted through the femoral artery and resides within the aorta. The Impella is inserted through the femoral artery and passed retrograde across the aortic valve, with the pump inflow within the left ventricle and the outflow in the ascending aorta. The TandemHeart pump is extracorporeal, with the inflow cannula tip in the left atrium and the outflow in the iliac artery or abdominal aorta. Extracorporeal membrane oxygenation (ECMO) consists of both a pump and oxygenator, drawing deoxygenated blood from the systemic venous system and returning newly oxygenated blood to the systemic arterial circulation, similar in principle to cardiopulmonary bypass. (B) Percutaneous IABP. Left, balloon deflated for insertion. Right, balloon inflated. (C) Tandem Heart pump, with the central inflow port and peripheral outflow port. (D) Impella device with the inflow (yellow arrow) and outflow (red arrow) areas that would sit in the left ventricle and ascending aorta, respectively. The motor is housed in the gray area adjacent to the outflow area. ((B) Courtesy S. Volvek, Datascope Corp., Oakland, NJ.)

CHAPTER 2.5.2A   Cardiovascular Medical Devices

1019

(D)

Figure 2.5.2A.11 cont’d

5.0 L/min. The device is typically inserted into the femoral or axillary artery and advanced retrograde across the aortic valve so the device inlet rests in the left ventricular cavity and the outlet sits distal to the aortic valve. While the first two devices can generally be inserted purely in a percutaneous fashion, the Impella 5.0 is large enough that it often requires a surgical cut down at the arterial insertion site. By virtue of its position in the left ventricle, the Impella unloads the left ventricular pressure as well as unloading volume, which is advantageous for reducing ventricular oxygen consumption and demand. In contrast, the left atrial inlet cannula position of the TandemHeart only unloads volume without significantly changing left ventricular pressure directly. Impella devices have been extensively used for patients undergoing high-risk percutaneous coronary interventions and for patients in cardiogenic shock secondary to acute myocardial infarction. They have also been investigated as a bridge-to-decision and even as a bridge-to-transplant in appropriate patients (Cheng et al., 2018). They are contraindicated in patients who have mechanical aortic valves or left ventricular thrombi. The more common complications of the device include device migration, thrombosis/thromboembolism, bleeding, hemolysis, and damage to left ventricular cardiac structures such as the aortic and mitral valves. The Impella RP is a similar device designed for right heart failure, which is inserted in the systemic venous circulation and sits such that the device inflow is in the right heart and the outflow is in the pulmonary artery distal to the pulmonic valve. ECMO, in contrast to the other percutaneous devices, provides the functions of both the heart (blood pressure and flow) and lungs (gas exchange) and is increasingly being used for patients with both acute cardiovascular and pulmonary diseases. The ECMO circuit is analogous to cardiopulmonary bypass in the operating room, but in a system that can be utilized at the bedside. The usual configuration is venoarterial (VA-ECMO) where systemic venous blood is removed from the patient, and passed through a continuous flow centrifugal pump, a heat exchanger, and a membrane oxygenator to provide full biventricular support and gas exchange. Cannulation can occur peripherally and percutaneously using the femoral vein and artery, or can use a surgical approach with cannulation of the right atrium and aorta in an open procedure. Just as with cardiopulmonary bypass, it is critical to have an experienced multidisciplinary team ensuring adequate functioning of the device and monitoring of the patient.

 Durable Ventricular Assist Devices and Total Artificial Hearts Durable VADs, first successfully employed by DeBakey in 1963, can replace ventricular function for extended periods, in contrast to the short-term duration of cardiopulmonary bypass, IABP, and percutaneous VADs. Durable VADs are currently used primarily in three settings: (1) for end-stage cardiac failure not likely to recover and where mechanical support will provide a “bridge-to-transplantation”; (2) for longterm cardiac support for patients with end-stage congestive heart failure that are not transplant candidates (“destination therapy”) (Christiansen et al., 2008); and (3) for chronic congestive heart failure where the unloading of the left ventricle may induce myocardial changes that might lead to normalization of cardiac function and eventually allow device removal (“bridge-to-recovery”). Research in this latter area focuses on the mechanisms of cardiac recovery, identification of patients who could achieve recovery, and specifics such as the timing and duration of therapy (Maybaum et al., 2008; Birks, 2010). The first generation of durable VADs used as bridge-totransplant or destination therapy (Hunt and Frazier, 1998) were large, pulsatile systems, with the inflow cannula of the device generally connected to the left ventricular apex and the outflow cannula connected to the ascending aorta. The pump itself would either be implanted in the peritoneal cavity with a driveline traversing the skin to provide power and controller functions, or would remain extracorporeal with the inflow and outflow cannulae each traversing the skin. Examples of first-generation pulsatile left ventricular assist devices (LVADs) included Thoratec HeartMate XVE (Fig. 2.5.2A.12) (Rose et  al., 2001) and Novacor Ventricular Assist System (Dagenais et al., 2001); the Thoratec PVAD (Farrar et al., 1990) is an example of a paracorporeal VAD. These devices generally consisted of a flexible polymer pumping bladder or diaphragm actuated by a pusher plate to allow filling and emptying of the pumping chamber. Valves on the inflow and outflow aspect of the pump ensured unidirectional flow of blood. These devices were very large and posed many challenges; nevertheless, these devices were critical in bridging patients to transplant from the 1980s into the mid-2000s. While the Thoratec PVAD is still occasionally used, the other devices are obsolete. The second generation of durable VADs consists of implantable continuous axial flow devices, where the long

1020 SE C T I O N 2. 5     Applications of Biomaterials

• Figure 2.5.2A.12  (A) Diagram of the HeartMate XVE pulsatile ventricular assist device. (B) Photograph of

HeartMate XVE and heart after removal from a patient at autopsy. ((A) Reproduced with permission from Slaughter, M.S., Rogers, J.G., Milano, C.A., Russell, S.D., Conte, J.V., Feldman, D., Sun, B., Tatooles, A.J., Delgado, R.M., Long, J.W., Wozniak, T.C., Ghumman, W., Farrar, D.J., Frazier, O.H., 2009. Advanced heart failure treated with continuous-flow left ventricular assist device. N. Engl. J. Med. 361, 2241–2251.)



Figure 2.5.2A.13 (A) Diagram of the HeartMate II continuous axial flow ventricular assist device. (B) Photograph of HeartMate II and heart after removal from patient at autopsy. ((A) Reproduced with permission from Slaughter, M.S., Rogers, J.G., Milano, C.A., Russell, S.D., Conte, J.V., Feldman, D., Sun, B., Tatooles, A.J., Delgado, R.M., Long, J.W., Wozniak, T.C., Ghumman, W., Farrar, D.J., Frazier, O.H., 2009. Advanced heart failure treated with continuous-flow left ventricular assist device. N. Engl. J. Med. 361, 2241–2251.)

axis of the impeller is parallel to the direction of blood flow. Examples of such devices include Thoratec HeartMate II (Fig. 2.5.2A.13) (Slaughter et  al., 2009), BerlinHeart INCOR (Schmid et  al., 2005) (pumps reside within the peritoneal cavity), and Jarvik 2000 FlowMaker (Sorensen et al., 2012) (the pump is intraventricular). These secondgeneration devices are connected in much the same way as the implantable first-generation devices but are much smaller, making implantation easier and allowing smaller patients to receive them; they are also more durable than the first-generation devices. Since these are all continuous flow devices, the pumps themselves do not impart pulsatility to the blood resulting in reduced pulse pressure for the patient;

this reduction of pulse pressure does not seem to have significant clinical effects. Data demonstrate lower complication rates and improved outcomes over this time, especially with newer continuous flow LVADs as compared to the first-generation pulsatile devices (Kirklin et al., 2017). The third generation of VADs consists of implantable continuous centrifugal flow devices, where the impeller creates a centrifugal force to add kinetic energy to the flowing blood. Examples of such devices include HeartWare HVAD (Aaronson et al., 2012), Thoratec HeartMate 3 (Schmitto et al., 2015), and Evaheart LVAS (Saito et al., 2014). The use of these devices is accelerating, with a respective decline in the use of the second-generation VADs.

CHAPTER 2.5.2A   Cardiovascular Medical Devices

1021

(A)

(B)

• Figure 2.5.2A.14  (A) Diagram of the HeartWare Ventricular Assist Device (HVAD), a continuous flow cen-

trifugal pump. (B) Photograph of HVAD and heart after surgical explantation from a patient who was bridged to cardiac transplantation with this device. ((A) Reproduced with permission from Rogers, J.G., Pagani, F.D., Tatooles, A.J., Bhat, G., Slaughter, M.S., Birks, E.J., Boyce, S.W., Najjar, S.S., Jeevanandam, V., Anderson, A.S., Gregoric, I.D., Mallidi, H., Leadley, K., Aaronson, K.D., Frazier, O.H., Milano, C.A., 2017. Intrapericardial left ventricular assist device for advanced heart failure. N. Engl. J. Med. 376, 451–460.)

The HeartWare Ventricular Assist Device (HVAD, Medtronic, Inc., Minneapolis, MN) was the first implantable third-generation centrifugal flow device available in the United States, with important design differences from the commercially available axial flow devices (Larose et al., 2010); it received CE Mark approval in 2008 followed by FDA approval in 2012. The pump is smaller and resides directly on the epicardial surface of the left ventricle (Fig. 2.5.2A.14), with the inflow cannula residing within the left ventricular cavity. The rotor produces continuous, centrifugal flow with a magnetically levitated impeller rather than

the axial flow system of the Thoratec HeartMate II, for example, which requires inflow and outflow bearings to support and align the impeller. The smaller pump size and driveline, intrathoracic positioning, and flow characteristics of the HVAD are thought to be advantageous in reducing common device complications such as infection, thrombosis, and bleeding. The HVAD ADVANCE trial demonstrated noninferiority (91% patient survival at 6 months) to other commercially available devices in a bridge-to-transplant setting, and there are longer follow-up data in a postmarket registry (Streuber et al., 2014). The original HVAD inflow

1022 SE C T I O N 2. 5     Applications of Biomaterials

(A)

(B)

(C)

• Figure 2.5.2A.15  (A) Diagram of the HeartMate 3 Ventricular Assist Device, a continuous flow centrifugal pump.

(B and C) Photograph of HeartMate 3 and heart after surgical explantation from a patient who was bridged to cardiac transplantation with this device. ((A) Reproduced with permission from Mehra, M.R., Naka, Y., Uriel, N., Goldstein, D.J., Cleveland, J.C., Colombo, P.C., Walsh, M.N., Milano, C.A., Patel, C.B., Jorde, U.P., Pagani, F.D., Aaronson, K.D., Dean, D.A., McCants, K., Itoh, A., Ewald, G.A., Horstmanshof, D., Long, J.W., Salerno, C., 2017. A fully magnetically levitated circulatory pump for advanced heart failure. N. Engl. J. Med. 376, 440–450.)

cannula had a smooth, polished titanium outer surface, which raised concerns about the increased risk of cerebrovascular accidents secondary to device-related thromboemboli (Najjar et al., 2014). A change in the HVAD inflow cannula design was therefore implemented in an attempt to promote nonthrombotic passivating tissue overgrowth by replacing the smooth, polished titanium surface with one incorporating a collar of sintered titanium microspheres (Soltani S, 2015). However, this created a discontinuity at the smoothsintered interface on the outer aspect of the inflow cannula that appears to serve as a nidus for thrombus formation (Glass et al., 2019).The HeartMate 3 (Fig. 2.5.2A.15) is the latest third-generation centrifugal flow pump with a fully magnetically levitated motor along with active magnetic

mounting (Chatterjee et  al., 2018). It was first implanted in humans in 2014 by a group in Germany (Schmitto et al., 2015). The motor incorporates a contactless bearing technology and consists of the rotor with passive magnets for drive and bearing, the stator with electromagnetic coils for drive and levitation, along with distance sensors and a microcontroller. This pump is approximately one-third the size of the HeartMate I and is implanted in the pericardial space rather than in the abdominal cavity. The inflow cannula is fully sintered and resides within the left ventricular chamber, with the outflow graft anastomosed to the ascending aorta. It can deliver up to 10 L/min of flow, and can also generate an “artificial pulse” by periodically increasing and decreasing the pump speed mimicking a pulse rate of

CHAPTER 2.5.2A   Cardiovascular Medical Devices

1023

• Figure 2.5.2A.16  (A) Diagram of the SynCardia Total Artificial Heart. (B) Photograph of an unimplanted SynCardia Total Artificial Heart. ((A) Reproduced with permission from Copeland, J.G., Smith, R.G., Arabia, F.A., Nolan, P.E., Sethi, G.K., Tsau, P.H., McClellan, D., Slepian, M.J., 2004. Cardiac replacement with a total artificial heart as a bridge to transplantation. N. Engl. J. Med. 351, 859–867.)

30 beats per minute. The MOMENTUM 3 clinical trial was designed to compare the centrifugal flow HeartMate 3 with the axial flow HeartMate II in patients with advanced heart failure. The HeartMate 3 showed lower rates of pump thrombosis, stroke, and reoperation to remove or replace a malfunctioning pump compared to the HeartMate II at 2-year follow-up (Mehra et al., 2018). In contrast to VADs where the native heart remains in place, a total artificial heart is composed of two pumping chambers that together replace the entire heart and provide both right and left ventricular function, analogous to heart transplantation (Fig. 2.5.2A.16). The SynCardia Total Artificial Heart (SynCardia Systems, Inc.; Tucson, AZ) has been implanted worldwide in more than 1000 patients with biventricular heart disease (Copeland et  al., 2003; Copeland et al., 2012). The device was approved by the FDA in 2004 and has proven to be an effective and reliable device for successful bridge-to-transplant. The clinical indications for implantation have included biventricular failure, left ventricular failure with prior mechanical heart valves, left ventricular failure with severe anatomical damage (ventricular septal defect, AV disruption), intractable malignant arrhythmias, massive ventricular thrombus, cardiac allograft failure, hypertrophic or restrictive cardiomyopathy, and complex congenital heart disease. The device is an implantable, pneumatically driven pulsatile pair of pumps consisting of polyurethane ventricles, whose inflows are anastomosed to the left and right atria and whose outflows are anastomosed to the ascending aorta and pulmonary artery after complete removal of the native cardiac ventricles and all four valves. Medtronic-Hall mechanical valves on the inflow and outflow aspects of the pump ensure unidirectional flow.

Systemic infection and thromboembolic or hemorrhagic events have been reported as the most common complications that prevent successful bridge-to-transplant. The major complications of cardiac assist devices are hemorrhage, thrombosis/thromboembolism, infection, interactions with host tissue, and device component failure, including the pump and peripheral electrical systems (Fig. 2.5.2A.17). Hemorrhage continues to be a problem in device recipients, although the risk of major hemorrhage has been decreasing with improved devices, therapies, patient selection, and surgical methods. Many factors predispose to perioperative hemorrhage, including (1) anticoagulation therapy and its management along with coagulopathy secondary to liver dysfunction and poor nutritional status, (2) contact of the blood with the device resulting in intrinsic platelet dysfunction and acquired von Willebrand disease, and (3) the extensive nature of the required surgery. Nonthrombogenic blood-contacting surfaces are essential for a clinically useful cardiac assist device or artificial heart. Indeed, thromboembolism occurred in most patients having long-term implantation of the Jarvik-7 artificial heart and is a major design consideration for current devices. In the absence of adequate anticoagulation and despite the development of minimally thrombogenic blood-contacting surfaces and appropriate blood flow characteristics, thrombi can form in areas of disturbed blood flow such as connections of conduits and other components to each other and to the natural heart. The current generation of continuous flow LVADs is carefully designed to minimize thrombosis, but oral anticoagulation is still required. Pump thrombosis in the continuous axial flow HeartMate II LVAD at the inflow bearing was one of the major sources of morbidity and mortality for patients with this device. Thrombi also may form outside the

1024 SE C T I O N 2. 5     Applications of Biomaterials

(A)

(B)

(C)

(D)

• Figure 2.5.2A.17  Complications of cardiac assist devices. (A) Hemorrhage into the brain in a patient with

a left ventricular assist device (LVAD). (B) Thrombus on the inflow flow straightener and inflow bearing of the impeller of the HeartMate II LVAD. (C) Thrombus on the inflow valve of a Thoratec PVAD. (D) Fungal infection in LVAD outflow graft. ((A) and (C) Reproduced with permission from Padera, R.F., 2008. Pathology of ventricular assist devices. In: McManus, B.M., Braunwald, E. (Eds.), Atlas of Cardiovascular Pathology for the Clinician, second ed. Current Medicine Group LLC, Philadelphia. (D) Reproduced by permission from Schoen, F.J., Edwards, W.D., 2001. Pathology of cardiovascular interventions. In: Silver, M.D., Gotlieb, A.I., Schoen, F.J. (Eds.), Cardiovascular Pathology, third ed. Churchill Livingstone, New York.)

LVAD, often in association with crevices and voids as in the HeartWare HVAD discussed earlier, and in areas of disturbed blood flow such as near connections of conduits and other components to the native heart. These thrombi can detach and lead to catastrophic embolic events such as ischemic stroke. Accounting for significant morbidity and mortality following the prolonged use of cardiac assist devices, infection can occur either within the device or associated with percutaneous drive lines (Padera, 2006). Susceptibility to infection is potentiated not only by the usual prosthesis-associated factors, but also by the multisystem organ damage from the underlying disease, the periprosthetic culture medium provided by postoperative hemorrhage, and by prolonged hospitalization with the associated risk of nosocomial infections. Assist device-associated infections are often resistant to antibiotic therapy and host defenses,

but are generally considered not an absolute contraindication to subsequent cardiac transplantation. Novel device designs, including alternative sites for driveline placement and the elimination of the driveline altogether with transcutaneous energy transmission technology, may play a role in further decreasing infection. 

Atrial Septal Defects and Other Intracardiac Defects In prenatal life, circulation is different than it is in postnatal life (Schoen and Mitchell, 2015). The lungs of the fetus are not providing gas exchange, so oxygenation of fetal blood is provided via the placenta and maternal circulation. This

CHAPTER 2.5.2A   Cardiovascular Medical Devices

requires two important shunts that need to close immediately after birth to separate the circulation into the pulmonary and the systemic arms. The foramen ovale, a hole in the fetal intraatrial septum, allows oxygenated blood returning to the right atrium from the placenta to preferentially pass into the left atrium. This blood passes through the mitral valve into the left ventricle and is pumped out through the aorta into the systemic circulation. The ductus arteriosus, present between the pulmonary artery and aorta, allows deoxygenated blood pumped from the right ventricle to bypass the lungs and directly reenter the systemic circulation, as the prenatal pulmonary circulation has a high vascular resistance (because of the nonexpanded lungs). After birth, these functional shunts should close to completely separate the right and left circulations; failure to do so results in a patent foramen ovale (PFO) or patent ductus arteriosus (PDA) that can allow inappropriate shunting of blood in the postnatal circulation. In addition, atrial septal defects (ASDs) or ventricular septal defects (VSDs) can also result from abnormal formation of the atrial septum or ventricular septum. While these defects can be closed via an open surgical procedure (sutures and/or fabric patches for PFO, ASD, or VSD, ligation for PDA), efforts have been made to allow closure of these defects using a minimally invasive approach. The decision to close a PFO, ASD, VSD, or PDA depends on the size of the shunt and the symptoms of the patient; the choice of technique (surgical vs. percutaneous) depends on the anatomy and structure of the defect.

Closure Devices The first catheter-based closure of a PDA was performed in 1967 by Porstmann using an Ivalon plug to occlude flow through the ductus arteriosus. Of the many PDA closure devices that have been developed over the years, most are metal-based devices that work by causing thrombosis of the PDA with subsequent organization and fibrosis, permanently preventing flow through the residual ductus arteriosus. Three of the more commonly used devices are the Gianturco coil (stainless-steel coil containing polyester fibers to promote thrombosis), the Amplatzer Duct Occluder (ADO, a conical device consisting of Nitinol wires and a polyester fiber patch to promote thrombosis and tissue integration), and the next-generation ADO II (Baruteau et al., 2014). Mills and King reported the first transcatheter closure of an ASD in 1976 using a double umbrella device that covered the opening from both the right and left atrial sides. Their occlusion device consisted of a skeleton of expanded polytetrafluoroethylene (ePTFE)-coated wire supporting an occluder of Dacron fabric delivered through a catheter. Improvements over the years include better device fixation methods and smaller caliber introducers. Several designs (Fig. 2.5.2A.18) of PFO/ASD closure devices are currently in use (Jung and Choi, 2018). The Amplatzer device (St. Jude Medical, St. Paul, MN) is a self-centering device that consists of double Nitinol disks filled with polyester patches

1025

connected by a small waist; the waist sits within the defect to connect the disks, which sit on either side of and are selected to be larger than the defect. The Gore Cardioform Septal Occluder (W.L. Gore and Associates, Flagstaff, AZ) is a nonself-centering device composed of a platinum-filled Nitinol wire framework, which is covered by an ePTFE patch designed to promote endothelialization. Advantages of nonsurgical closure devices such as these include shorter hospital stay, more rapid recovery, and no residual thoracotomy scar. With the experience gained in the transcatheter closure of PFOs and ASDs, this interventional technology is being extended to the closure of some VSDs, particularly in patients thought to be poor operative risks. Several types of complications have been reported for closure devices. The most straightforward is the failure to fully close the defect resulting in residual shunting. Erosion of the device through the interatrial septum occurs in some patients with perforation and device embolization; this is thought to be secondary to the specific anatomy of the defect, but the stiffness of the device may also play a role. Inadequate fixation of the device within the defect or a device-defect size mismatch can also result in device embolization. Closure devices are effective in closing defects at least in part via thrombosis; if the thrombosis extends beyond the defect on the device, thromboemboli may result. Fractures of various device components, air embolism at the time of device deployment, infection, and development of new arrhythmias have also been reported. A number of devices are in development, with the trends being defect-specific design and minimization of the amount of foreign material left in the patient, including the use of biodegradable components (O’Byrne and Levi, 2019). 

Atrial Fibrillation AF affects more than 3 million individuals in the United States, making it the most common cardiac arrhythmia. Instead of orderly atrial contraction initiated by the SA node, rapid disorganized electrical activity in the atria causes these upper chambers of the heart to quiver or fibrillate resulting in poor contractile function and irregular flow within the chamber. As would be predicted by Virchow’s triad, thrombosis may occur due to these flow abnormalities within the atria, especially within the atrial appendage. Atrial appendage thrombi are an important source of thromboemboli, explaining why patients with AF are at a fivefold greater risk for embolic stroke than individuals in sinus rhythm. Anticoagulation is effective in reducing the risk of atrial thrombosis and stroke, but this therapy has many drawbacks, including a narrow therapeutic window, variability in metabolism of the drug, interactions with other drugs and metabolites, need for frequent monitoring by blood drawing, poor patient compliance, and, most importantly, the risk of life-threatening bleeding. These side effects are reduced in a class of drugs called nonvitamin K antagonist oral anticoagulants when compared to the established vitamin K antagonists such as warfarin (Granger et al., 2011), but there is still a

1026 SE C T I O N 2. 5     Applications of Biomaterials

(A)

(B)

(C)

(D)

(E)

(F)

(G)

(H)

(I)

(J)



Figure 2.5.2A.18 Closure devices for intraatrial septal defects. (A) Amplatzer Septal Occluder; (B) Occlutech Figulla Flex II device; (C) Gore Cardioform Septal Occluder; (D) Cocoon Septal Occluder; (E) CeraFlex ASD device; (F) Nit Occlud ASD-R device; (G) Cardi-O-Fix Septal Occluder; (H) Ultracept II ASD Occluder; (I) Carag Bioresorbable Septal Occluder; (J) Amplatzer Septal Occluder after 2 years with fibrous tissue overgrowth. ((A–I) Reproduced with permission from Jung, S.Y., Choi, J.Y., 2018. Transcatheter closure of atrial septal defect: principles and available devices. J. Thorac. Dis. 10, S2909–S2922.)

CHAPTER 2.5.2A   Cardiovascular Medical Devices

(A)

1027

(B)

Aorta Watchman device placed in left atrial appendage Left atrium Right atrium

(C) Transseptal catheterization

Flexible catheter inserted through right femoral vein and up the inferior vena cava

(D)

• Figure 2.5.2A.19  The Watchman left atrial appendage occluder device. (A) Diagrammatic. (B) Photograph

of Watchman device showing coverage by host tissue. (C) Specimen radiograph of Watchman device showing metallic framework. (D) Photograph of heart at autopsy with Watchman device successfully occluding the left atrial appendage. ((A) Reproduced with permission from Maisel, W.H., 2009. Left atrial appendage occlusion – closure or just the beginning? N. Engl. J. Med. 360, 2601–2603.)

bleeding risk with these newer medications. An approach for reducing the risk of thromboembolic stroke in patients with AF is to remove or ligate the left atrial appendage, first proposed in the 1930s and first performed in 1949 via a surgical approach. This approach is typically used when a patient is undergoing a concomitant cardiac surgical procedure such as valve replacement or bypass surgery.

Left Atrial Appendage Occlusion Devices Nonsurgical device-based approaches to close the left atrial appendage have been developed, including several devices that can be deployed percutaneously to occlude the opening to the appendage and isolate it from the blood in the left atrium (Pacha et al., 2019). The Watchman (Fig. 2.5.2A.19)

1028 SE C T I O N 2. 5     Applications of Biomaterials

left atrial appendage system (Boston Scientific, Marlborough, MA) is an FDA-approved, percutaneously deployed parachute-shaped device consisting of a Nitinol cage with a polyethylene terephthalate membrane on its surface and fixation barbs along the perimeter, which allows it to anchor in the atrial appendage. This device has been shown to be noninferior to standard warfarin therapy for prevention of embolic stroke and systemic embolization, and superior to warfarin for prevention of cardiovascular death and hemorrhagic stroke in recent clinical trials (Holmes et al., 2009; Reddy et al., 2017). The Amplatzer Amulet (St. Jude Medical, Minneapolis, MN) device is another current generation percutaneously deployed left atrial occlusion device that is currently in clinical trials (Kleinecke et al., 2017) prior to FDA approval. The Atriclip Device System (Atricure, Inc., West Chester, OH) is an occlusion device applied on the epicardial surface at the base of the left atrial appendage to close the entrance to the left atrial appendage by external compression (Ailawadi et al., 2011). The device consists of two parallel rigid titanium tubes with elastic Nitinol springs that hold the device closed.

References Aaronson, K.D., Slaughter, M.S., Miller, L.W., McGee, E.C., Cotts, W.G., Acker, M.A., et al., 2012. Use of an intrapericardial, continuous-flow, centrifugal pump in patients awaiting heart transplantation. Circulation 125, 3191–3200. Ailawadi, G., Gerdisch, M.W., Harvey, R.L., Hooker, R.L., Damiano, R.J., Salamon, T., Mack, M.J., 2011. Exclusion of the left atrial appendage with a novel device: early results of a multicenter study. J. Thorac. Cardiovasc. Surg. 142, 1002–1009. Alavi, S.H., Kheradvar, A., 2015. A hybrid tissue-engineered heart valve. Ann. Thorac. Surg. 99, 2183–2187. Alpard, S.K., Zwischenberger, J.B., 2002. Extracorporeal membrane oxygenation for severe respiratory failure. Chest Surg. Clin, N. Am. 12, 355–378. Amat-Santos, I.J., Ribeiro, H.B., Urena, M., et al., 2015. Prosthetic valve endocarditis after transcatheter valve replacement: a systemic review. JACC Cardiovasc. Interv. 8, 334–336. Amin, M.S., Ellenbogen, K.A., 2010. The effect of device advisories on implantable cardioverter-defibrillator therapy. Curr. Cardiol. Rep. 12, 361–366. Atlee, J.L., Bernstein, A.D., 2001. Cardiac rhythm management devices (part I): indications, device selection, and function. Anesthesiology 95, 1265–1280. Ayoub, S., Ferrari, G., Gorman, R.C., Gorman III, J.H., Schoen, F.J., Sacks, M.S., 2017. Heart valve biomechanics and underlying mechanobiology. Compr. Physiol. 6, 1743–1780. Ayyadurai, P., Alkhawam, H., Saad, M., Al-Sadawi, M.A., Shah, N.N., Kosmas, C.E., Vittorio, T.J., 2019. An update on the CardioMEMS pulmonary artery pressure sensor. Ther. Adv. Cardiovasc. Dis 13, 1–11. Baddour, L.M., Epstein, A.E., Erickson, C.C., Knight, B.P., Levison, M.E., Lockhart, P.B., Masoudi, F.A., Okum, E.J., Wilson, W.R., Beerman, L.B., Bolger, A.F., Estes 3rd, N.A., Gewitz, M., Newburger, J.W., Schron, E.B., Taubert, K.A., 2010. Update on cardiovascular implantable electronic device infections and their management: a scientific statement from the American Heart Association. Circulation 121, 458–477.

Barnett, S.C., Ad, N., 2009. Surgery for aortic and mitral valve disease in the United States: a trend of change in surgical practice between 1998 and 2005. J. Thorac. Cardiovasc. Surg. 137, 1422–1429. Baruteau, A.-E., Hascoet, S., Baruteau, J., Boudjemline, Y., Lambert, V., Angel, C.-Y., Belli, E., Petit, J., Pass, R., 2014. Transcatheter closure of patent ductus arteriosus: past, present and future. Arch. Cardiovasc. Dis. 107, 122–132. Baskett, R.J.F., Ghali, W.A., Maitland, A., Hirsch, G.M., 2002. The intraaortic balloon pump in cardiac surgery. Ann. Thorac. Surg. 74, 1276–1287. Baughman, K.L., Jarcho, J.A., 2007. Bridge to life – cardiac mechanical support. N. Engl. J. Med. 357, 846–849. Benjamin, E.J., Munter, P., Alonso, A., 2019. Heart disease and stroke statistics-2019 update. Circulation 139 (10), e56–e528. Bennett, P.C., Silverman, S.H., Gill, P.S., Lip, G.Y., 2009. Peripheral arterial disease and Virchow’s triad. Thromb. Haemost. 101, 1032–1040. Bezuidenhout, D., Williams, D.F., Zilla, P., 2015. Polymeric heart valves for surgical implantation, catheter-based technologies and heart assist devices. Biomaterials 36 , 6–25. Birks, E.J., 2010. Myocardial recovery in patients with chronic heart failure: is it real? J. Card. Surg. 25, 472–477. Blot, W.J., Ibrahim, M.A., Ivey, T.D., Acheson, D.E., Brookmeyer, R., Weyman, A., Defauw, J., Smith, J.K., Harrision, D., 2005. Twentyfive year experience with the Bjork-Shiley convexoconcave heart valve: a continuing clinical concern. Circulation 111, 2850–2857. Boilson, B.A., Raichlin, E., Park, S.J., Kushwaha, S.S., 2010. Device therapy and cardiac transplantation for end-stage heart failure. Curr. Probl. Cardiol. 35, 8–64. Bonow, R.O., Carabello, B.A., Kanu, C., et al., 2006. ACC/AHA 2006 guidelines for the management of patients with valvular heart disease: a report of the American College of cardiology/ American heart association task force on practice guidelines (writing committee to revise the 1998 guidelines for the management of patients with valvular heart disease): developed in collaboration with the Society of cardiovascular anesthesiologists: endorsed by the Society for cardiovascular angiography and interventions and the Society of thoracic surgeons. Circulation 114, e84–e231. Boroumand, S., Asadpour, S., Akbarzadeh, A., Faridi-Majidi, R., Ghanbari, H., 2018. Heart valve tissue engineering: an overview of heart valve decellularization processes. Regen. Med. 13, 41–54. Bouten, C.V., Driessen-Mol, A., Baaijens, F.P., 2012. In situ heart valve tissue engineering: simple devices, smart materials, complex knowledge. Expert Rev. Med. Devices 9, 453–457. Buja, L.M., Schoen, F.J., 2016. The pathology of cardiovascular interventions and devices for coronary artery disease, vascular disease, heart failure, and arrhythmias. In: Buja, L.M., Butany, J. (Eds.), Cardiovascular Pathology, fourth edition. Elsevier, pp. 577–610. Butany, J., Ahluwalia, M.S., Payet, C., et al., 2002. Hufnagel valve: the first prosthetic mechanical valve. Cardiovasc. Pathol. 11, 351–353. Carabello, B.A., 2008. The current therapy for mitral regurgitation. J. Am. Coll. Cardiol. 52, 319–326. Carabello, B.A., Paulus, W.J., 2009. Aortic stenosis. Lancet 373, 956–966. Carpentier, A., 2007. Lasker Clinical Research Award. The surprising rise of nonthrombogenic valvular surgery. Nat. Med. 13, 1165–1168. Chaikoff, E.L., 2007. The development of prosthetic heart valves – lessons in form and function. N. Engl. J. Med. 357, 1368–1371. Chandrashekhar, Y., Westaby, S., Narula, J., 2009. Mitral stenosis. Lancet 374, 1271–1283.

CHAPTER 2.5.2A   Cardiovascular Medical Devices

Chatterjee, A., Feldman, C., Hanke, J.S., Ricklefs, M., Shrestha, M., Dogan, G., Haverich, A., Schmitto, J.D., 2018. The momentum of HeartMate 3: a novel active magnetically levitated centrifugal left ventricular assist device. J. Thorac. Dis. 10, S1790–S1793. Chauvette, V., Mazine, A., Bouchard, D., 2018. Ten-year experience with the Perceval S sutureless prosthesis: lessons learned and future perspectives. J. Vis. Surg. May 3; 4, 87. https://doi.org/10.21037/ jovs.2018.03.10. Cheng, R., Tank, R., Ramzy, D., Azarbai, B., Chung, J., Esmailian, F., Kobashigawa, J.A., Moriguchi, J., Arabia, F.A., 2018. Clinical outcomes of Impella microaxial devices used to salvage cardiogenic shock as a bridge to durable circulatory support or cardiac transplantation. ASAIO J. 65, 642–648. Cheung, D.Y., Duan, B., Butcher, J.T., 2015. Current progress in tissue engineering of heart valves: multiscale problems, multiscale solutions. Expert Opin. Biol. Ther. 15, 1155. Christiansen, S., Kolcke, A., Autschback, R., 2008. Past, present and future of long-term mechanical circulatory support in adults. J. Card. Surg. 23, 664–676. Claibome, T.E., Slepian, M.J., Hossainy, S., Bluestein, D., 2012. Polymeric trileaflet prosthetic heart valves: evolution and path to clinical reality. Expert Rev. Med. Devices 9, 577–594. Copeland, J.G., Arabia, F.A., Tsau, P.H., Nolan, P.E., McClellan, D., Smith, R.G., Slepian, M.J., 2003. Total artificial hearts: bridge to transplantation. Cardiol. Clin. 21, 101–113. Copeland, J.G., Copeland, H., Gustafson, M., Mineburg, N., Covington, D., Smith, R.G., Friedman, M., 2012. Experience with more than 100 total artificial heart implants. J. Thorac. Cardiovasc. Surg. 143, 727–734. Curtis, N., Vohra, H.A., Ohri, S.K., 2010. Mini extracorporeal circuit cardiopulmonary bypass system: a review. Perfusion 25, 115–124. Dagenais, F., Portner, P.M., Robbins, R.C., Oyer, P.E., 2001. The Novacor left ventricular assist system: clinical experience from the Novacor registry. J. Card. Surg. 16, 267–271. De Sousa, C.F., Brito, F.D., de Lima, V.C., Carvalho, A.C., 2010. Percutaneous mechanical assistance for the failing heart. J. Interv. Cardiol. 23, 195–202. Della Rocca, D.G., Gianni, C., Di Base, L., Natale, A., Al-Ahmad, A., 2018. Leadless pacemakers: state of the art and future perspectives. Card. Electrophysiol. Clin. 10, 17–29. DeWall, R.A., Qasim, N., Carr, L., 2000. Evolution of mechanical heart valves. Ann. Thorac. Surg. 69, 1612–1621. Di Eusanio, M., Phan, K., 2015. Sutureless aortic valve replacement. Am. Cardiothorac. Surg. 4, 123–130. Dickinson, M.G., Allen, L.A., Albert, N.A., Disalvo, T., Ewald, G.A., Vest, A.R., Whellan, D.J., Zile, M.R., Givertz, M.M., 2018. Remote monitoring of patients with heart failure: a white paper from the heart failure society of America scientific statements commmittee. J. Card. Fail. 24, 682–694. DiGiorgi, P.L., Rao, V., Naka, Y., Oz, M.C., 2003. Which patient, which pump? J. Heart Lung Transplant. 22, 221–235. Driessen-Mol, A., Emmert, M.Y., Dijkman, P.E., et  al., 2014. Transcatheter implantation of homologous “off-the-shelf ” tissue-engineered heart valves with self-repair capacity: long-term functionality and rapid in vivo remodeling in sheep. J. Am. Coll. Cardiol. 63, 1320–1329 . D’Agostino, R.S., Jacobs, J.P., Bodhwar, V., Fernandez, F.G., Paone, G., Wormuth, D.W., Shahian, D.M., 2018. The Society of thoracic surgeons adult cardiac surgery database: 2018 update on outcomes and quality. Ann. Thorac. Surg. 105, 15–23. Edmunds Jr., L.H., 2001. Evolution of prosthetic heart valves. Am. Heart J. 141, 849–855.

1029

Emmert, M.Y., Weber, B., Falk, V., Hoerstrup, S.P., 2014. Transcatheter tissue engineered heart valves. Expert Rev. Med. Devices 11, 15–21. Farrar, D.J., Lawson, J.H., Litwak, P., Cederwall, G., 1990. Thoratec VAD system as a bridge to heart transplantation. J. Heart Lung Transplant. 9, 415–422. Fassa, A.A., Himbert, D., Vahanian, A., 2013. Mechanisms and management of TAVR-related complications. Nat. Rev. Cardiol. 10, 685–695. Fedak, P.W., McCarthy, P.M., Bonow, R.O., 2008. Evolving concepts and technologies in mitral valve repair. Circulation 117, 963–974. Fishbein, M.C., Roberts, W.C., Golden, A., Hufnagel, C.A., 1975. Cardiac pathology after aortic valve replacement using Hufnagel trileaflet prostheses: a study of 20 necropsy patients. Am. Heart J. 89, 443–448. Fishbein, G.A., Schoen, F.J., Fishbein, M.C., 2014. Transcatheter aortic valve implantation: status and challenges. Cardiovasc. Pathol. 23, 65. Fuzellier, J.F., Campisi, S., Gerbay, A., Haber, B., Ruggieri, V.G., Vola, M., 2016. Two hundred consecutive implantation of the sutureless 3f Enable aortic valve: what we have learned. Ann. Thorac. Surg. 101, 1716–1723. Gammie, J.S., Sheng, S., Griffith, B.P., et al., 2009. Trends in mitral valve surgery in the United States: results from the Society of thoracic surgeons adult cardiac surgery database. Ann. Thorac. Surg. 87, 1431–1437. Généreux, P., Head, S.J., Wood, D.A., et  al., 2012. Transcatheter aortic valve implantation 10-year anniversary: review of current evidence and clinical implications. Eur. Heart J. 33, 2399–2402. Ghazanfari, S., Driessen-Mol, A., Sanders, B., et  al., 2015. In  vivo collagen remodeling in the vascular wall of decellularized stented tissue-engineered heart valves. Tissue Eng. A 21, 2206–2215. Glass, C.H., Christakis, A., Fishbein, G.A., Watkins, J.C., Strickland, K.C., Mitchell, R.N., Padera, R.F., 2019. Thrombus on the inflow cannula of the HeartWare HVAD: an update. Cardiovasc. Pathol. 38, 14–20. Goldbarg, S.H., Halperin, J.L., 2008. Aortic regurgitation: disease progression and management. Nat. Clin. Pract. Cardiovasc. Med. 5, 269–279. Goldbarg, S.H., Elmariah, S., Miller, M.A., Fuster, V., 2007. Insights into degenerative aortic valve disease. J. Am. Coll. Cardiol. 50, 1205. Goldstone, A.B., Chiu, P., Baiocchi, M., Lingala, B., Patrick, W.L., Fischbein, M.P., Woo, Y.J., 2017. Mechanical or biologic prosthesis for aortic valve and mitral valve replacement. N. Engl. J. Med. 377, 1847–1857. Granger, C.B., Alexander, J.H., McMurray, J.J., Lopes, R.D., Hylek, E.M., Hanna, M., Al‐Khalidi, H.R., Ansell, J., Atar, D., Avezum, A., Bahit, M.C., Diaz, R., Easton, J.D., Ezekowitz, J.A., Flaker, G., Garcia, D., Geraldes, M., Gersh, B.J., Golitsyn, S., Goto, S., Hermosillo, A.G., Hohnloser, S.H., Horowitz, J., Mohan, P., Jansky, P., Lewis, B.S., Lopez‐Sendon, J.L., Pais, P., Parkhomenko, A., Verheugt, F.W., Zhu, J., Wallentin, L., 2011. Apixaban versus warfarin in patients with atrial fibrillation. N. Engl. J. Med. 365, 981–992. Halbfass, P., Sonne, K., Nentwich, K., Ene, E., Deneke, T., 2018. Current developments in cardiac rhythm management devices. Clin. Res. Cardiol. 107, S100–S104. Hamm, C.W., Arsalan, M., Mack, M.J., 2016. The future of transcatheter aortic valve implantation. Eur. Heart J. 37, 203–210. Haqqani, H.M., Mond, H.G., 2009. The implantable cardioverterdefibrillator lead: principles, progress and promises. Pacing Clin. Electrophysiol. 32, 1336–1353.

1030 SE C T I O N 2. 5     Applications of Biomaterials

Harken, D.F., Taylor, W.J., LeFemine, A.A., et al., 1962. Aortic valve replacement with a caged ball valve. Am. J. Cardiol. 9, 292–299. Harrison, D.C., Ibrahim, M.A., Weyman, A.E., Kuller, L.H., Blot, W.J., Miller, D.E., 2013. The Bjork-Shiley convexo-concave heart valve experience from the perspective of the supervisory panel. Am. J. Cardiol. 112, 1921–1931. Head, S.J., Celik, M., Kappetein, A.P., 2017. Mechanical versus bioprosthetic aortic valve replacement. Eur. Heart J. 38, 2183–2191. Hirji, S.A., Kolkailah, A.A., Ramirez-Del Val, R., Lee, J., McGurk, S., Pelletier, M., Singh, S., Mallidi, H.R., Aranki, S., Shekar, P., Kaneko, T., 2018. Mechanical versus bioprosthetic aortic valve replacement in patients aged 50 years and younger. Ann. Thorac. Surg. 106, 1113–1120. Hoerstrup, S.P., Sodian, R., Daebritz, S., et  al., 2000. Functional living trileaflet heart valves grown in-vitro. Circulation 102, III 44–III 49. Hoerstrup, S.P., Cummings, I., Lachat, M., et al., 2006. Functional growth in tissue engineered living vascular grafts: follow up at 100 weeks in a large animal model. Circulation 114, I159–I166. Holmes, D.R., Reddy, V.Y., Turi, Z.G., Doshi, S.K., Sievert, H., Buchbinder, M., Mullin, C.M., Sick, P., 2009. Percutaneous closure of the left atrial appendage versus warfarin therapy for prevention of stroke in patients with atrial fibrillation: a randomized non-inferiority trial. Lancet 374, 534–542. Huikuri, H.V., Castellanos, A., Myerburg, R.J., 2001. Sudden death due to cardiac arrhythmias. N. Engl. J. Med. 345, 1473–1482. Hunt, S.A., Frazier, O.H., 1998. Mechanical circulatory support and cardiac transplantation. Circulation 97, 2079–2090. Isaacs, A.J., Shuhaiber, J., Salemi, A., Isom, O.W., Sekrakyan, A., 2015. National trends in utilization and in-hospital outcomes of mechanical versus bioprosthetic aortic valve replacements. J. Thorac. Cardiovasc. Surg. 149, 1262–1269. Jessup, M., Brozena, S., 2003. Heart failure. N. Engl. J. Med. 348, 2007–2018. Joy, P.S., Kumar, G., Poole, J.E., London, B., Olshansky, B., 2017. Cardiac implantable electronic device infections: who is at greatest risk? Heart Rhythm 14, 839–845. Jung, S.Y., Choi, J.Y., 2018. Transcatheter closure of atrial septal defect: principles and available devices. J. Thorac. Dis. 10, S2909–S2922. Kar, B., Adkins, L.E., Civitello, A.B., Loyalka, P., Palanichamy, N., Bemmato, C.J., MyersTJ, Gregoric, I.D., Delgado 3rd, R.M., 2006. Clinical experience with the TandemHeart percutaneous ventricular assist device. Tex. Heart Inst. J. 33, 111–115. Kim, S.J., Samad, Z., Bloomfield, G.S., Douglas, P.S., 2014. A critical review of hemodynamic changes and left ventricular remodeling after surgical aortic valve replacement and percutaneous aortic valve replacement. Am. Heart J. 168, 150–159. Kirklin, J.K., Pagani, F.D., Kormos, R.L., Stevenson, L.W., Blume, E.D., Myers, S.L., et  al., 2017. Eighth annual INTERMACS report: special focus on framing the impact of adverse events. J. Heart Lung Transplant. 36, 1080–1086. Kleinecke, C., Park, J.W., Godde, M., Zintl, K., Schnupp, S., Brachmann, J., 2017. Twelve-month follow-up of left atrial appendage occlusion with Amplatzer Amulet. Cardiol. J. 24, 131–138. Kodali, S.K., Williams, M.R., Smith, C.R., Svensson, L.G., Webb, J.G., Makkar, R.R., et al., 2012. Two-year outcomes after transcatheter or surgical aortic-valve replacement. N. Engl. J. Med. 366, 1686–1695. Krainski, F., Pretorius, V., Birgersdotter-Green, U., 2018. A practical approach to lead removal: transvenous tools and techniques. Card. Electrophysiol. Clin. 10, 637–650.

Krishnamani, R., DeNofrio, D., Konstam, M.A., 2010. Emerging ventricular assist devices for long-term cardiac support. Nat. Rev. Cardiol. 7, 71–76. Kusumoto, F.M., Goldschlager, N., 2002. Device therapy for cardiac arrhythmias. JAMA 287, 1848–1852. Larose, J.A., Tamez, D., Ashenuga, M., Reyes, C., 2010. Design concepts and principle of operation of the HeartWare ventricular assist system. ASAIO J. 56, 285–289. Leat, M.E., Fisher, J., 1993. Comparative study of the function of the Abiomed polyurethane heart valve for use in left ventricular assist devices. J. Biomed. Eng. 15, 516–520. Levy, J.H., Tanaka, K.A., 2003. Inflammatory response to cardiopulmonary bypass. Ann. Thorac. Surg. 75, S715–S720. Lifton, R.P., 2007. Lasker Award to heart valve pioneers. Cell 130, 971–974. Lueders, C., Jastram, B., Hetzer, R., Schwandt, H., 2014. Rapid manufacturing techniques for the tissue engineering of human heart valves. Eur. J. Cardiothorac. Surg. 46, 593–601. Lurz, P., Bonhoeffer, P., Taylor, A.M., 2009. Percutaneous pulmonary valve implantation: an update. Expert Rev. Cardiovasc. Ther. 7, 823–833. Mack, M.J., Leon, M.B., Thourani, V.H., Makkar, R., Kodali, S.K., Russo, M., Kapadia, S.R., Malaisrie, S.C., Cohen, D.J., Pibarot, P., Leipsic, J., Hahn, R.T., Blanke, P., Williams, M.R., McCabe, J.M., Brown, D.L., Babaliaros, V., Goldman, S., Szeto, W.Y., Genereux, P., Pershad, A., Pocock, S.J., Alu, M.C., Webb, J.G., Smith, C.R., PARTNER 3 Investigators., 2019. Transcatheter aortic valve replacement with a ballon-expandable valve in low-risk patients. N. Engl. J. Med. 380, 1695–1705. Mancini, D., Lietz, K., 2010. Selection of cardiac transplantation candidates in 2010. Circulation 122, 173–183. Mandawat, A., Rao, S.V., 2017. Percutaneous mechanical circulatory support devices in cardiogenic shock. Circ. Cardiovasc. Interv. 10 (5), e004337. Masoumi, N., Annabi, N., Assmann, A., et al., 2014. Tri-layered elastomeric scaffolds for engineering heart valve leaflets. Biomaterials 35, 7774–7785. Mastroroberto, P., Chello, M., Bevacqua, E., Cirillo, F., Covino, E., 2000. Duromedics original prosthesis: what do we really know about diagnosis and mechanism of leaflet escape? Can. J. Cardiol. 16, 1269–1272. Matheny, R.G., Hutchison, M.L., Dryden, P.E., Hiles, M.D., Shaar, C.J., 2000. Porcine small intestine submucosa as a pulmonary valve leaflet substitute. J. Heart Valve Dis. 9, 769. Mathieu, P., Bosse, Y., Huggins, G.S., Della Corte, A., Pibarot, P., Michelena, H.I., Limongelli, G., Boulanger, M.-C., Evangelista, A., Bedard, E., Citro, R., Body, S.C., Nemer, M., Schoen, F.J., 2015. The pathology and pathobiology of bicuspid aortic valve: state of the art and novel research perspectives. J. Pathol.: Clin. Res. 1, 195–206. Maybaum, S., Kamalakannan, G., Murthy, S., 2008. Cardiac recovery during mechanical assist device support. Semin. Thorac. Cardiovasc. Surg. 20, 234–246. McAlister, F.A., Ezekowitz, J., Hooton, N., Vandermeer, B., Spooner, C., Dryden, D.M., Page, R.L., Hlatky, M.A., Rowe, B.H., 2007. Cardiac resynchronization therapy for patients with left ventricular systolic dysfunction: a systematic review. JAMA 297, 2502–2514. Mehra, M.R., Goldstein, D.J., Uriel, N., Cleveland, J.C., Yuzefpolskaya, M., Salerno, C., Walsh, M.N., Milano, C.A., Patel, C.B., Ewald, G.A., Itoh, A., Dean, D., Krishnamoorthy, A., Cotts, W.G., Tatooles, A.J., Jorde, U.P., Bruckner, B.A., Estep, J.D.,

CHAPTER 2.5.2A   Cardiovascular Medical Devices

Jeevanandam, V., Sayer, G., Horstmanshof, D., Long, J.W., Gulati, S., Skipper, E.R., O’Connell, J.B., Heatley, G., Sood, P., Naka, Y., For the MOMENTUM 3 Investigators, 2018. Two-year outcomes with a magnetically levitated cardiac pump in heart failure. N. Engl. J. Med. 378, 1386–1395. Merryman, W.D., Schoen, F.J., 2013. Mechanisms of calcification in aortic valve disease: role of mechanokinetics and mechanodynamics. Curr. Cardiol. Rep. 15 (5) 355. Meuris, B., Flameng, W.J., Laborde, F., et al., 2015. Five-year results of the pilot trial of a stentless valve. J. Thorac. Cardiovasc. Surg. 150, 84. Miller, P.E., Solomon, M.A., McAreavey, D., 2017. Advanced percutaneous mechanical circulatory support devices for cardiogenic shock. Crit. Care Med. 45, 1922–1929. Mohr, F.W., 2014. Decade in review – valvular disease: current perspectives on treatment of valvular heart disease. Nat. Rev. Cardiol. 11, 637–638. Mond, H.G., Grenz, D., 2004. Implantable transvenous pacing leads: the shape of things to come. Pacing Clin. Electrophysiol. 27, 887–893. Mylotte, D., Andalib, A., Theriault-Lauzier, P., et  al., 2015. Transcatheter heart valve failure: a systematic review. Eur. Heart J. 36, 1306–1327. Najjar, S.S., Slaughter, M.S., Pagani, F.D., Starling, R.C., McGee, E.C., Eckman, P., et  al., 2014. An analysis of pump thrombus events in patients in the HeartWare ADVANCE bridge to transplant and continued access protocol trial. J. Heart Lung Transplant. 33, 23–34. Neraqi-Miandoab, S., Ewstbrook, B., Flynn, J., Blakely, J., Baribeau, Y., 2015. Prosthetic valve endocarditis five months following transcatheter aortic valve implantation and review of literature. Heart Surg. Forum 18, E20–E22. O’Byrne, M.L., Levi, D.S., 2019. State-of-the-art atrial septal defect closure devices for congenital heart. Interv. Cardiol. Clin. 8, 11–21. O’Brien, M.F., Harrocks, S., Stafford, E.G., et al., 2001. The homograft aortic valve: a 29-year, 99.3% follow up of 1022 valve replacements. J. Heart Valve Dis. 10, 334–344. Pacha, H.M., Al-khadra, Y., Soud, M., Darmoch, F., Pacha, A.M., Alraies, M.C., 2019. Percutaneous devices for left atrial appendage occlusion: a contemporary review. World J. Cardiol. 11, 57–70. Padera, R.F., 2006. Infection in ventricular assist devices: the role of biofilm. Cardiovasc. Pathol. 15, 264–270. Panaich, S.S., Eleid, M.F., 2018. Current status of MitraClip for patients with mitral and tricuspid regurgitation. Trends Cardiovasc. Med. 28, 200–209. Patel, A., Bapat, V., 2017. Transcatheter mitral valve replacement: device landscape and early results. Eurointervention 13, AA31–AA39. Piper, C., Korfer, R., Horstkotte, D., 2001. Prosthetic valve endocarditis. Heart 85, 590–593. Popma, J.J., Deeb, G.M., Yakubov, S.J., Mumtaz, M., Gada, H., O’Hair, D., Bajwa, T., Heiser, J.C., Merhi, W., Kleiman, N.S., Askew, J., Sorajja, P., Rovin, J., Chetcuti, S.J., Adams, D.H., Teirstein, P.S., Zorn 3rd, G.L., Forrest, J.K., Tchétché, D., Resar, J., Walton, A., Piazza, N., Ramlawi, B., Robinson, N., Petrossian, G., Gleason, T.G., Oh, J.K., Boulware, M.J., Qiao, H., Mugglin, A.S., Reardon, M.J., Evolut Low Risk Trial Investigators, 2019. Transcatheter aortic valve replacement with a self-expanding valve in low-risk patients. N. Engl. J. Med. 380, 1706–1715. Rabkin, E., Hoerstrup, S.P., Aikawa, M., et  al., 2002. Evolution of cell phenotype and extracellular matrix in tissue-engineered heart valves during in-vitro maturation and in-vivo remodeling. J. Heart Valve Dis. 11, 308–314.

1031

Rahimtulla, S.H., 2003. Choice of prosthetic heart valve for adult patients. J. Am. Coll. Cardiol. 41, 893–904. Reardon, M.J., Adams, D.H., Kleiman, N.S., et al., 2015. 2-year outcomes in patients undergoing surgical or self-expanding transcatheter aortic valve replacement. J. Am. Coll. Cardiol. 66, 113–121 ([Epub]). Reardon, M.J., Van Mieghem, N.M., Popma, J.J., Kleiman, N.S., Søndergaard, L., Mumtaz, M., Adams, D.H., Deeb, G.M., Maini, B., Gada, H., Chetcuti, S., Gleason, T., Heiser, J., Lange, R., Merhi, W., Oh, J.K., Olsen, P.S., Piazza, N., Williams, M., Windecker, S., Yakubov, S.J., Grube, E., Makkar, R., Lee, J.S., Conte, J., Vang, E., Nguyen, H., Chang, Y., Mugglin, A.S., Serruys, P.W., Kappetein, A.P., SURTAVI Investigators, 2017. Surgical or transcatheter aortic valve replacement in intermediate-risk patients. N. Engl. J. Med. 376, 1321–1331. Reddy, V.Y., Doshi, S.K., Kar, S., Gibson, D.N., Price, M.J., Huber, K., Horton, R.P., Buchfinder, M., Neuzil, P., Gordon, N.T., Holmes, D.R., PREVAIL and PROTECT AF Investigators, 2017. 5-year outcomes after left atrial appendage closure: from the PREVAIL and PROTECT AF trials. J. Am. Coll. Cardiol. 70, 2964–2975. Rippel, R.A., Ghanbari, H., Seifalian, A.M., 2012. Tissue-engineered heart valve; future of cardiac surgery. World J. Surg. 36, 1581–1591. Rodés-Cabau, J., 2012. Transcatheter aortic valve implantation: current and future approaches. Nat. Rev. Cardiol. 9, 15–29. Rose, E.A., Gelijns, A.C., Moskowitz, A.J., Heitjan, D.F., Stevenson, L.W., Dembitsky, W., et  al., 2001. Long-term use of a left ventricular assist device for end-stage heart failure. N. Engl. J. Med. 345, 1435–1443. Saito, S., Yamazaki, K., Nishinaka, T., Ichihara, Y., Ono, M., Kyo, S., et al., 2014. Post-approval study of a highly pulsed, low-shearrate, continuous-flow, left ventricular assist device, EVAHEART: a Japanese multicenter study using J-MACS. J. Heart Lung Transplant. 33, 599–608. Saleeb, S.F., Newburger, J.W., Geva, T., Baird, C.W., Gauvreau, K., Padera, R.F., Del Nido, P.J., Borisuk, M.J., Sanders, S.P., Mayer, J.E., 2014. Accelerated degeneration of a bovine pericardial bioprosthetic aortic valve in children and young adults. Circulation 130, 51–60. Sapirstein, J.S., Smith, P.K., 2001. The “ideal” replacement heart valve. Am. Heart J. 141, 856–860. Schmid, C., Tjan, T.D., Etz, C., Schmidt, C., Wenzelburger, F., Wilhelm, M., et al., 2005. First clinical experience with the Incor left ventricular assist device. J. Heart Lung Transplant. 24, 1188–1194. Schmitto, J.D., Hanke, J.S., Rosa, S.V., Asvar, M., Haverich, A., 2015. First implantation in man of a new magnetically levitated left ventricular assist device (HeartMate 3). J. Heart Lung Transplant. 34, 858–860. Schoen, F.J., 2011. Heart valve tissue engineering: quo vadis? Curr. Opin. Biotechnol. 22, 698–705 ([Epub ahead of print]). Schoen, F.J., Butany, J., 2016. Cardiac valve replacement and related interventions. In: Buja, L.M., Butany, J. (Eds.), Cardiovascular Pathology, fourth ed. Elsevier, pp. 529–576. Schoen, F.J., Gotlieb, A.I., 2016. Heart valve health, disease, replacement and repair: a 25-year cardiovascular pathology perspective. Cardiovasc. Pathol. 25, 341–352. Schoen, F.J., Levy, R.J., 2005. Calcification of tissue heart valve substitutes: progress toward understanding and prevention. Ann. Thorac. Surg. 79, 1072–1080. Schoen, F.J., Mitchell, R.N., 2015. The heart. In: Kumar, V., Abbas, A.K., Aster, J.C. (Eds.), Robbins and Cotran Pathologic Basis of Disease, ninth ed. W.B. Saunders, Philadelphia, PA, pp. 523–578.

1032 SE C T I O N 2. 5     Applications of Biomaterials

Schoen, F.J., Webb, J.G., 2008. Prosthetics and the heart. In: McManus, B.M., Braunwald, E. (Eds.), Atlas of Cardiovascular Pathology for the Clinician, pp. 241–256 Current Medicine, Philadelphia. Simon, P., Kasimir, M.T., Seebacher, G., et al., 2003. Early failure of the tissue engineered porcine heart valve SYNERGRAFT in pediatric patients. Eur. J. Cardiothorac. Surg. 23, 1002–1006. Slaughter, M.S., Rogers, J.G., Milano, C.A., Russell, S.D., Conte, J.V., Feldman, D., et al., 2009. Advanced heart failure treated with continuous-flow left ventricular assist device. N. Engl. J. Med. 361, 2241–2251. Smith, C., Leon, M., Mack, M., et al., 2011. Transcatheter versus surgical aortic-valve replacement in high-risk patients. N. Engl. J. Med. 364, 2187–2198. Soltani, S., Kaufmann, F., Vierecke, J., Kretzschmar, A., Hennig, E., Stein, J., et al., 2015. Design changes in continuous-flow left ventricular assist devices and life-threatening pump malfunction. Eur. J. Cardiothorac. Surg. 47, 984–989. Sorensen, E.N., Pierson 3rd, R.N., Feller, E.D., Griffith, B.P., 2012. University of Maryland surgical experience with the Jarvik 2000 axial flow ventricular assist device. Ann. Thorac. Surg. 93, 133–140. Starr, A., 2007. Lasker clinical medical research award. The artificial heart valve. Nat. Med. 10, 1160–1164. Stassen, O.M.J.A., Muylaert, D.E.P., Bouten, C.V.C., Hjortnaes, J., 2017. Current challenges in translating tissue-engineered heart valves. Curr. Treat. Options Cardiovasc. Med. 19 (9), 71. Stone, G.W., Lindenfeld, J., Abraham, W.T., Kar, S., Lim, D.S., Mishell, J.M., Whisenant, B., Grayburn, P.A., Rinaldi, M., Kapadia, S.R., Rajagopal, V., Sarembock, I.J., Brieke, A., Marx, S.O., Cohen, D.J., Weissman, N.J., Mack, M.J., COAPT Investigators, 2018. Transcatheter mitral valve repair in patients with heart failure. N. Engl. J. Med. 379, 2307–2318. Strueber, M., Larbalestier, R., Jansz, P., Zimpfer, D., Fiane, A.E., Tsui, S., Simon, A., Schmitto, J.D., Khaghani, A., Wieselthaler, G.M., Najarian, K., Schueler, S., 2014. Results of the postmarket registry to evaluate the HeartWare left ventricular assist system (ReVOLVE). J. Heart Lung Transplant. 33, 486–491. Tempelhof, M.W., Klein, L., Cotts, W.G., Benzuly, K.H., Davidson, C.J., Meyers, S.N., McCarthy, P.M., Malaisrie, C.S., McGee, E.C., Beohar, N., 2011. Clinical experience and patient outcomes associated with the TandemHeart percutaneous transseptal assist device among a heterogeneous patient population. Am. Soc. Artif. Intern. Organs J. 57, 254–261. The Antiarrhythmics versus Implantable Defibrillators (AVID) Inevstigators, 1997. A comparison of antiarrhythmic-drug therapy

with implantable defibrillators in patients resuscitated from nearfatal ventricular arrhythmias. N. Engl. J. Med. 337, 1576–1583. Tice, J.A., Sellke, F.W., Schaff, H.V., 2014. Transcatheter aortic valve replacement in patients with severe aortic stenosis who are at high risk for surgical complications: Summary assessment of the California Technology Assessment Forum. J. Thorac. Cardiovasc. Surg. 148, 482–491. Trost, J.C., Hillis, L.D., 2006. Intra-aortic balloon counterpulsation. Am. J. Cardiol. 97, 1391–1398. Vahanian, A., 2008. Antithrombotic therapy for patients with valvular heart disease. Herz 33, 44–51. Van Mieghem, N.M., El Faquir, N., Rahhab, Z., et al., 2015. Incidence and predictors of debris embolizing to the brain during transcatheter aortic valve implantation. JACC Cardiovasc. Interv. 8, 718. Vohra, H.A., Whistance, R., Modi, A., Ohri, S.K., 2009. The inflammatory response to miniaturised extracorporeal circulation: a review of the literature. Mediat. Inflamm. 2009, 2009, 707042. Vongpatanasin, W., Hillis, L.D., Lange, R.A., 1996. Prosthetic heart valves. N. Engl. J. Med. 335, 407–416. Webb, J.G., Murdoch, D.J., Boone, R.H., Moss, R., Attinger-Toller, A., Blanke, P., Cheung, A., Hensey, M., Leipsic, J., Ong, K., Sathananthan, J., Wood, D.A., Ye, J., Tartara, P., 2019. Percutaneous transcatheter mitral valve replacement: first-in-human experience with a new transseptal system. J. Am. Coll. Cardiol. 73, 1239–1246. Willcox, M.E., Prutkin, J.M., Bardy, G.H., 2016. Recent developments in the subcutaneous ICD. Trends Cardiovasc. Med. 26, 526–535. Wyler von Ballmoos, M.C., Kalra, A., Reardon, M.J., 2018. Complexities of transcatheter mitral valve replacement (TMVR) and why it is not transcatheter aortic valve replacement (TAVR). Ann. Cardiothorac. Surg 7, 724–730. Zannis, K., Folliguet, T., Laborde, F., 2012. New sutureless aortic valve prosthesis: another tool in less invasive aortic valve replacement. Curr. Opin. Cardiol. 27, 125–129. Zeitler, E.P., Pokorney, S.D., Zhou, K., Lewis, R.K., Greenfield, R.A., Daubert, J.P., Matchar, D.B., Piccini, J.P., 2015. Cable externalization and electrical failure of the Riata family of implantable cardioverter-defibrillator leads: a systematic review and meta-analysis. Heart Rhythm 12, 1233–1240. Zhang, X., Xu, B., Puperi, D.S., et  al., 2015. Integrating valveinspired design features into poly(ethylene glycol) hydrogel scaffolds for heart valve tissue engineering. Acta Biomater. 14, 11–21. Zhang, B.L., Bianco, R.W., Schoen, F.J., 2019. Preclinical assessment of cardiac valve substitutes: current status and considerations for engineered tissue heart valves. Front. Cardiovasc. Med. 6, 72.

2.5.2B

Cardiovascular Medical Devices: Stents, Grafts, Stent-Grafts and Other Endovascular Devices MICHAEL A. SEIDMAN a , ROBERT F. PADERA, FREDERICK J. SCHOEN Department of Pathology, Brigham and Women's Hospital and Harvard Medical School, Boston, MA, United States

V

ascular disorders and their downstream sequelae are responsible for more morbidity and mortality than any other category of human disease. The top two causes of death worldwide, ischemic heart disease and stroke, collectively comprising approximately 13 million persons annually, more than one-quarter of all deaths in both men and women (Lozano et  al., 2012), are both manifestations of vascular pathology. Additionally, aortic aneurysms and related pathologies are associated with over 175,000 deaths a year globally (GBD 2013 Mortality and Causes of Death Collaborators, 2015). Thus, the vasculature has been a longstanding and important target for both open surgical procedures and endovascular interventions using biomaterials and associated devices. The devices that have facilitated these procedures are highly specialized to the anatomic site and functional role of the segment of vascular bed targeted, and to the specific pathological processes the procedures and devices are intended to manage. This chapter focuses on the biomaterials and devices used to manage vascular disease. However, before discussing specific device categories, we provide a brief overview of the vascular system and the key pathologic processes that affect it.

Key Concepts in Vascular Structure and Function Architecture of the Circulation The circulation (a.k.a. the cardiovascular system) transports and distributes essential substances throughout the body and aDr.

Seidman is presently affiliated with the University Health Network and the University of Toronto, Toronto, Canada.

helps eliminate waste products. The cardiovascular system is composed of a pump, a series of distributing and collecting tubes that carry blood away from the heart (arteries), vessels that carry blood toward the heart (veins), and an extensive network of thin-walled vessels (capillaries) that provide nutrient and gas exchange to the tissues and remove byproducts of metabolism (i.e., oxygen and nutrients in; carbon dioxide and other waste products out). To accomplish these functions, the cardiovascular system must regulate blood pressure and volume at all levels of the vasculature, buffer pulsatile flow in order to ensure steady and slow velocity flow in the capillaries, maintain circulatory continuity while permitting free and bidirectional exchange of substances between blood and the extravascular compartments, and control hemostasis, i.e., prevent clots (thrombi) when the circulation is intact, and prevent hemorrhage when the vessel is injured. Other functions of the circulation include regulation of body temperature, and systemic distribution of various regulating substances (e.g., hormones, inflammatory mediators, growth factors, and therapeutic substances). Moreover, the circulatory system distributes immune and inflammatory cells, and the lining cells of the circulatory system (endothelial cells) facilitate egress of the immune cells from the blood to the tissues in which they are needed. The human circulation is composed of two parallel circuits with this generalized configuration (Fig. 2.5.2B.1). The systemic circulation originates at the aortic valve of the heart; the single lead artery, the aorta, branches into successively smaller arteries until ultimately reaching the capillaries, the level of the circulatory system at which metabolic exchange with the tissues occurs. The capillaries then “coalesce” into small and progressively larger veins, ultimately returning to the right side of the heart via the largest veins, the inferior and superior venae cavae. Similarly, and 1033

1034 SEC T I O N 2. 5     Applications of Biomaterials

• Figure 2.5.2B.1  The cardiovascular system. Large and medium-sized

vessels carrying oxygenated (red) and deoxygenated (blue) blood are shown with respect to the heart (orange). The ascending aorta (red) and pulmonary artery (blue) carry blood away from the heart, the superior and inferior vena cavae (blue) carry blood back to the heart (the pulmonary veins, in red behind the pulmonary artery, also return blood to the heart). The descending aorta passes behind the heart but does not anastomose with it. (Reproduced from https://commons.wikimedi a.org/w/index.php?curid=10005587. (Accessed 3 September 2019).)

in parallel, the pulmonary circulation begins at the pulmonary valve, from which arises the main pulmonary artery (the only large artery in the body which carries blood which has not yet been oxygenated). In the lung, the pulmonary artery branches successively into smaller arteries, then capillaries, and collects oxygenated blood into successively larger veins, returning blood to the left atrium via the pulmonary veins. In most organs and tissues, one or a few larger arteries bring blood in, then branch into capillaries in a manner

that creates vascular beds within each organ. The branching structure dramatically increases the surface area available for gas and nutrient exchange at the capillary level. The entire cardiovascular system follows a consistent architectural paradigm with regional differences (Figs. 2.5.2B.2 and 2.5.2B.3). Three basic structural constituents comprise the walls of blood vessels: endothelium, muscular tissue, and extracellular matrix (ECM), including collagen and elastic elements. These tissues are arranged in concentric layers (called tunica, meaning coats): (1) the intima—consisting of endothelial cells (ECs), subendothelial connective tissue and the internal elastic lamina; (2) the media—consisting of smooth muscle cells (SMCs) and elastic laminas, including the external elastic lamina; and (3) adventitia— collagenous connective tissue. The histologic composition and organization, as well as the thickness of these three layers, vary characteristically with the physiologic functions performed by specific segments of the vascular system (see Fig. 2.5.2B.3). The layered structure of the blood vessels progressively undergoes segmental differentiation and adaptation to local factors, both mechanical (especially blood pressure, which determines the amount and arrangement of muscular tissue), and metabolic (reflecting the local nutritional needs of the tissues). As emphasized above, the structure of vessels varies with their functional role. Arteries are usually divided into three major categories. In sequential order from the heart they are: elastic arteries (large or conducting arteries), muscular arteries (medium or distributing arteries), and arterioles (smallest arteries, 15

Ti oxide

Titanium–aluminum– vanadium

90Ti–6Al–4V

120

860–930

>8

Ti oxide

Cobalt–chromium– molybdenum (casting)

66Co–28Cr–6Mo

235

520–1170

>8

Cr oxide

Stainless steel (316L)

70Fe–18Cr–12Ni

190

480–1000

>30

Cr oxide

Zirconium

+

99 Zr

100

450–550

>15

Zr oxide

Tantalum

+



170–520

30

Au

+

165

130

40

Pt

Gold Platinum

99 Ta

99 Pt

w/o, weight percent. aMinimum values from the American Society for Testing and Materials Committee F4 documents are provided. Selected products provide a range of properties.

TABLE a 2.5.5.2    Engineering Properties of Inert Ceramics Used as Biomaterials

Material

Modulus of Elasticity GPa

Ultimate Bending Strength MPa

Surface

Poly-crystalline

370

300–550

Al2O3

Single crystal (sapphire)

390

640

Al2O3

Zirconium oxide (partially stabilized zirconia)

150–240

500–650

ZrO2

Titanium oxide (titania)

280

70–100

TiO2

Carbon

25–40 (4–6)

150–250 (22–36)

C

Carbon-silicon (LTI)

25–40 (4–6)

200–700 (29–101)

CSi

Aluminum oxide

aThese

high ceramics have 0% permanent elongation at fracture.

TABLE a 2.5.5.3    Engineering Properties of Bioactive and Biodegradable Ceramics

Material

Modulus of Elasticity GPa

Ultimate Bending Strength MPa

Surface

Hydroxyapatite

10–120

15–300

Ca10(PO4)6(OH)2

Tricalcium phosphate

30–120

15–120

Ca3(PO4)2

Bioglass or Ceravital

40–140

20–350

CaPO4

Apatite–Wollastonite ceramic

125

215

CaPO4 + F

aThese

ceramics have 0% permanent elongation at fracture.

CHAPTER 2.5.5  Dental Applications

Metals Metallic implants used in clinical dentistry today are almost exclusively manufactured from titanium and its alloys. The strength of titanium is due to its hexagonal close-packed crystal lattice and crystallographic orientation, whereas its biocompatibility (corrosion resistance) is attributed to its stable, passive oxide layer. Titanium implants are in a passive state (their oxide is chemically stable) in a normal physiological microenvironment. Both CPTi and Ti-6Al4V possess excellent corrosion resistance for a full range of oxide states and pH levels. The coherent oxide layer and ability of titanium to repassivate results in corrosion resistance. Titanium ions are released from transient chemical dissolution of the oxide. However, the low dissolution rate and relative nonreactivity of titanium dissolution products allow bone to thrive and osseointegrate with titanium. The mechanical properties of titanium-based materials are well established. Microstructures with small (4–6 nm) to pass through the glomeruli of the kidneys (Wilhelm et al., 2016). In cases where drug elimination is similar to or faster than absorption or pharmacodynamic effects, a DDS can be critical for the clinical application of a drug candidate. 

Dosage and Distribution Control DDSs allow better control of dosage to match the TI. Control can come via feedback (i.e., electrical control via sensors or from physical interactions with the body) or simple ratelimited release to improve convenience or better match the therapeutic concentration of a drug with a narrow TI. As an example, insulin is a hormone with a narrow and rapidly changing therapeutic concentration that is dependent upon food intake and composition, activity level, and blood glucose concentration. A fully integrated glucose sensor coupled with insulin release could act as an artificial pancreas and remove the burden of constant monitoring and calculation of insulin dosages, providing improved quality of life and reduced risk of hypo- or hyperglycemia. Current insulin delivery pumps, such as wearable modules that deliver insulin through a persistent subcutaneous injection site, offer excellent control over blood glucose levels and improve patient compliance, but still require user input despite attempts at automation (Bergenstal et al., 2010). One biomaterial DDS solution under development is the use of phenylboronic acid, a glucose-responsive moiety, as a sensor for insulin release from systemically delivered mesoporous silica nanoparticles (Zhao et  al., 2009). DDSs designed for convenience include extended-release oral capsules, transdermal delivery patches, and implanted depots, which are covered in the section “DDSs to Overcome Biological Barriers.” 

Controlling Drug Release Kinetics Modifying drug delivery rate can be sufficient to ensure maintenance of the TW. Drug release rate from a DDS can be controlled via several mechanisms, including diffusion, dissolution, affinity, swelling, and ion exchange (Fig. 2.5.12.7). Table 2.5.12.2 lists examples of FDA-approved DDSs that exploit these delivery mechanisms. Diffusion from DDSs

is common and driven by concentration gradients. The general categories for these DDSs are matrix or reservoirtype systems. Matrix-based systems, which are also known as monoliths, contain drug uniformly dispersed within the material and are commonly accompanied by an initial “burst” release of drug upon placement due to rapid diffusion of surface-localized drug. If the material is hydrophilic, swelling mediates diffusion-controlled release. If the material is hydrophobic, drug releases after water penetration enables a diffusive path. In addition to “burst” release, diffusive release from matrix devices inherently follows first-order kinetics and requires coupling with degradable DDS materials to modify the release profile. Reservoir systems contain drug within an inner core surrounded by a permeable membrane layer that controls release and can achieve zero-order release with appropriately designed membrane–drug combinations. Dissolution is liberation of matrix-entrapped drug as a function of matrix degradation or dissolution rate (Fig. 2.5.12.7). The kinetics of drug delivery by dissolution is controlled by DDS properties, including pore size, degradable bonds, and hydrophobicity, rather than drug solubility and mobility, as in diffusion. However, drug release is often dependent upon both drug diffusion and dissolution, as dissolution or degradation alters the DDS pore size, thus liberating greater amounts of drug over time. Affinity-based systems exploit noncovalent interactions, including electrostatic, van der Waals, hydrophobic, and hydrogen-bonding interactions between drug and DDS to control drug release rate (Fig. 2.5.12.7). Affinity-based DDS release rates are tunable and based on the association constant of drug–ligand interactions, which allow for release of multiple drugs with various kinetics. However, a priori identification of affinity ligands is necessary for affinity-based release. Several ligands have been identified; these include cyclodextrin, heparin, albumin, and various cationic DDSs to release a variety of drugs, including small molecule antibiotics, proteins, and nucleic acids (Bader et al., 2014; Fu et al., 2011; Oss-Ronen and Seliktar, 2011; Rivera-Delgado et al., 2016; Vulic and Shoichet, 2014; Wang and von Recum, 2011). Charged drug molecules can be loaded into an ion exchange resin to provide control over release rate as a function of the ionic environment. Ion exchange is particularly suitable for enteral delivery routes, as ion exchange resins are typically inert, functionalized poly(styrene) derivatives formed into micron- or millimeter-scale beads that pass through the digestive tract safely (Guo et al., 2009). These resins have found commercial success in many over-thecounter extended release formulations (Table 2.5.12.2). Swelling of osmotic pumps can control mechanical dispensing systems to achieve variable drug release (Table 2.5.12.2). An osmotic pump is a compartment surrounded by a semipermeable membrane typically composed of cellulose acetate. The membrane controls diffusion of water into the osmogen: a material with high osmotic pressure, such as sugars or salts, embedded into a carrier such as poly(ethylene oxide) or poly(hydroxypropyl methylcellulose). The influx

CHAPTER 2.5.12   Drug Delivery Systems

1245

• Figure 2.5.12.7 Mechanisms of drug release from drug delivery systems (DDSs). Release of drugs from

DDSs can be controlled by a number of mechanisms. (A) Drug is encapsulated within a DDS with mesh/pore size to allow for diffusive release of the encapsulated drug with optional diffusional membrane barrier. (B) Drug is tethered to a DDS that degrades hydrolytically, oxidatively, photolytically, or proteolytically to control release. (C) Drug is tethered to the DDS by a degradable tether, and released upon linker cleavage via hydrolysis, oxidation, photolysis, or proteolysis. (D) Diffusive release of encapsulated drug is controlled by affinity interactions between the DDS and the drug. (E) Diffusive release of encapsulated drug is prolonged by delayed release of the drug from a matrix or reservoir. (F) Drug is encapsulated within a degradable DDS and released by dissolution as the material degrades. Not to scale. (Adapted from Van Hove, A.H., Benoit, D.S., 2015. Depot-based delivery systems for pro-angiogenic peptides: a review. Front. Bioeng. Biotechnol. 3, 102.)

of water to the osmogen increases the pressure inside the container, forcing the drug or drug carrier through microdrilled pores within the membrane. There are many variations of osmotic pumps, from the single-component elementary osmotic pump (Theeuwes, 1975) to multistage, multichamber systems, all of which can range from ingestible pills to implantable devices. Release kinetics can be tuned from zero-order to complex profiles simulating multiple separate doses (Malaterre et al., 2009). More detailed reviews of osmotic pumps can be found within Malaterre et al. (2009) and Herrlich et al. (2012). 

DDSs to Improve Drug Solubility Drug solubility is vital for successful delivery. Dose variations, poor and unknown absorption profiles, low bioavailability, and subpar therapeutic efficacy are limitations associated with systemically delivered, poorly soluble, small molecule drugs. The intrinsic link between solubility and drug efficacy is described by Lipinski’s rule of 5, which predicts that compounds with molecular weight 5 will not penetrate cell membranes and is unlikely to be drug-like. Compound hydrophobicity determined based on meeting Lipinski’s rule of 5: (A) MW < 500 g/mol, H-donor = 1, H-acceptor = 4, log P < 8 (B) MW = 542.53 g/mol, H-donor = 6, H-acceptor = 12, log P = 1.3 (C) MW < 500 g/mol, H-donor = 3, H-acceptor = 5, log P > 3—passively diffuses through stratum corneum per Table 2.5.12.4 (D) MW > 500 g/mol, H-donor = 0, H-acceptor = 12, log P>5 The drugs that are able to permeate the epidermis will need to be small ( 1). Drug C meets these criteria and can be expected to penetrate lipids well. Drug A is not currently intended for transdermal delivery; however, it has been shown to penetrate the stratum corneum through handling by nurses. Drug B is too hydrophilic for these criteria, while drug D is too large and hydrophobic. Per Table 2.5.12.4, drug C has been formulated to be used for transdermal delivery as well, likely due to MW < 500 g/mol and log P > 3.

1266.e3

Approaches that can be used to enhance transdermal drug delivery of the remaining drugs could be iontophoresis, electrophoresis, microneedles, transdermal microjet, etc. (Table 2.5.12.4). 6.  Several key scientific findings in the 19th and early 20th centuries (e.g., Ehrlich’s magic bullet and Banting, Best, and Macleod’s discovery that insulin treated diabetes) laid the groundwork for DDS development in the late 20th century. What recent scientific findings have occurred since the start of the 21st century and how might they impact the future of DDS design and development? What challenges will need to be overcome to take full advantage of these scientific findings in DDS design and development? Answers could include discoveries such as mapping of the human genome, siRNA, and CRISPR (as well as others). Certainly, these technologies will be useful for future design of targeted DDSs. For example, the human genome map will likely highlight genes or proteins that enable targeted cancer therapies to only destroy cancerous and potentially cancerous cells while leaving healthy cells unaffected. Technologies such as siRNA or CRISPR will likely provide specific molecular tools that facilitate this targeting process, but suitable DDS designs using these technologies will need to overcome multiple challenges, including biocompatibility, preliminary drug degradation, biological barriers to enable targeted drug delivery, etc.

2.5.13

Responsive Polymers in the Fabrication of Enzyme-Based Biosensors JOHN R. AGGAS 1, 2 , ANTHONY GUISEPPI-ELIE 1 ,2,3 1Center

for Bioelectronics, Biosensors and Biochips (C3B), Department of Biomedical Engineering, Texas A&M University, College Station, TX, United States 2Department 3ABTECH

of Biomedical Engineering, Texas A&M University, College Station, TX, United States

Scientific, Inc., Biotechnology Research Park, Richmond, VA, United States

Introduction The detection and measurement of glucose (static or continual; in vitro, in situ, or in vivo) has been studied in great depth since the description of the first biosensor by Clark and Lyons in 1962 and the subsequent patent of the original amperometric glucose enzyme electrode in 1970 (Clark and Lyons, 1962; Clark et  al., 1970). In addition to the well-studied benefits of glucose sensing for diabetic patients, continuous measurement of glucose can serve as a gauge to identify shock states for patients who have undergone hemorrhage-associated trauma (Guiseppi-Elie, 2011) and may be experiencing insulin resistance. A biosensor is a fully integrated system capable of receiving an input/stimulus from its environment through a biotransducer and outputting a signal that can be processed into readable, actionable information through its instrumentation. The development of biosensors over the last several decades has been driven not only by need in the medical industry, but by needs in the food, environmental, and national defense industries (Idros et al., 2015; Neethirajan et al., 2018; Justino et al., 2017). Biosensors can be split into four main groups, based on the type of physicochemical transduction employed: electrochemical, optical, thermal, and piezoelectric. From 2000 to the present, over 5000 peer-reviewed articles have been recorded in the ISI Web of Knowledge relevant to these four types of glucose biosensors. From the data shown in Fig. 2.5.13.1A, the number of articles written for each type of glucose biosensor has increased sixfold over the past

2 decades. The most widely studied glucose biosensor systems are electrochemical (48%), followed by optical (32%), thermal (15%), and piezoelectric (5%). Electrochemical glucose biosensors have been the most widely researched type of biosensor since their inception. Electrochemical glucose biosensors, specifically amperometric biosensors, offer many advantages over the other types of transduction mechanisms. Specifically, the low cost of instrumentation, the small footprint that is possible, and the low power requirements coupled with high sensitivity and fast response times (dictated by enzyme kinetics/diffusion) have all been exploited to realize miniaturized, portable, pointof-care glucose-biosensing systems. Moreover, amperometric glucose biosensors are highly studied because they offer the highest sensitivity and are more easily produced on a large scale than other enzymatic electrochemical biosensors (Gaudin, 2017). Optical glucose biosensors, which continue to be highly researched, also offer several advantages. Optical biosensors offer label-free detection of many types of biological or chemical analytes at high specificity and sensitivity. In addition, optical biosensors show great promise for noninvasive biosensing (Bandodkar and Wang, 2014). However, the equipment required for optical measurements is often costly and requires higher power consumption than most electrochemical glucose biosensors. Enzymatic electrochemical glucose biosensors, which utilize glucose oxidase (GOx) (an oxido-reductase: EC number 1.1.3.4) as the bioreceptor, can be classified into four groups based on the type of electrochemical transduction 1267

1268 SEC T I O N 2. 5     Applications of Biomaterials

(A)

(B)

450 400 350 300

Amperometric Electrochemical+Glucose+Biosensor Opcal+Glucose+Biosensor

Potenometric

Thermal+Glucose+Biosensor

Conductometric

Piezoelectric+Glucose+Biosensor

250

Impedimetric

200 150 100 50 0

• Figure 2.5.13.1  (A) The number of electrochemical, optical, thermal, and piezoelectric glucose sensor-

related articles since the year 2000, by year. (B) The total number of electrochemical glucose biosensors since the year 2000, split into the transduction type: amperometric, conductometric, impedimetric, and potentiometric. (Data collected from ISI Web of Knowledge.)

employed: amperometric (apply a constant voltage and measure current), conductometric (apply a current and measure a voltage drop), impedimetric (apply a sinusoidally varying voltage and measure the ensuing phase-shifted current as the transfer function), and potentiometric (measure the time-varying voltage relative to a reference voltage). Of the articles recorded on the ISI Web of Knowledge since 2000, over 90% are related to amperometric sensing, followed by potentiometric, impedimetric, and conductometric, as shown in Fig. 2.5.13.1B. Recently, responsive polymers have been incorporated into enzymatic electrochemical biosensor systems. In part, stimuli-responsive polymers (and addition of nanoparticles) have been incorporated into the transduction element and bioreceptor element in passive and active roles. Many important reviews have been written on glucose biosensors, but a deep understanding and topical review of the incorporation of stimuli-responsive polymers into the various parts of the biosensing system has not been a major area for review. This review discusses the basic principles and types of enzymatic biosensors, with a focus on electrochemical glucose biosensors, the roles and advantages of incorporating stimuli-responsive polymers in electrochemical glucose biosensors, and the current and future outlooks of systems integration of biosensor fabrication.

Classic Biosensor System The classic biosensor system consists of a bioreceptor established within a recognition layer, a physicochemical transducer, instrumentation or hardware, and signal/computer processing with accompanying signal processing, conditioning, and the production of actionable information as output. The latter is often overlooked but is in fact the central reason for adopting a biosensor format for any analyte measurement.

Bioreceptor (Recognition Layer) The bioreceptor layer of a biosensor system can be any organic or biomimetic component which has high selectivity toward the target analyte to be measured. The role of the bioreceptor is to receive a stimulus from a target analyte and to present that stimulus to a transducer. Commonly used bioreceptors include antibodies/antigens, nucleic acids/DNA, whole cells, synthetic biomimetics (molecularly imprinted polymers), and enzymes (Tavakoli and Tang, 2017). Antibodies are large, Y-shaped glycoproteins used by the immune system of the body to identify and neutralize bacteria and viruses by binding to a specific epitope of the pathogen called the antigen. The high specificity of the antibody–antigen binding (kaff > 10−4 M) has led to its widespread use as a bioreceptor in immuno-biosensors. Sandwich assays utilize a capture antibody and a detection antibody labeled with an enzyme to “sandwich” the analyte of interest, followed by a measurement of the catalytic activity between the enzyme attached to the detection antibody and an enzymatic substrate (Cho et al., 2018). In general, antibodies are not available for small molecules (haptens) such as glucose. However, the potential does exist for the development of peptide mimetics with high binding affinities to build biosensors (Thyparambil et al., 2017). Nucleic acids and DNA used as bioreceptors undergo a recognition process called complementary base-pairing (i.e., adenine-thymine and cytosine-guanine in DNA). Spontaneous hydrogen bonding between the bioreceptor and target complementary strand can be measured by labeling with an electroactive indicator or measurement of a signal produced by the DNA (Asal et al., 2018). The use of RNA or DNA as a bioreceptor is easily prepared via polymerase chain reaction (PCR). Aptamers, which are short, single-stranded oligonucleotides that are able to bind to target molecules have been used as bioreceptors, but the similarity between glucose and other monosaccharides such as fructose and galactose

CHAPTER 2.5.13   Responsive Polymers in the Fabrication of Enzyme-Based Biosensors

(similar structures and hydroxyl groups) creates difficulty in selecting an aptamer with high specificity (Ruscito and DeRosa, 2016). Hence, high-specificity aptamers are generally not available for glucose detection and monitoring. Immobilized cells, organelles, and tissues have been used as transducers in the detection of many biological agents which affect cells. The response of the cell, and signals that form their homeostasis, can be transduced into measurable signals. These cellular biosensors utilize cells which rely on several different biological signals. For example, pancreatic β-cells are able to sense multiple nutrient levels (glucose, lipids, amino acids) and signal for insulin via hormones (i.e., GLP-1, GIP), which were measured through electrode arrays and amplified to extract glucose concentration (Nguyen et  al., 2013). Kim et  al. utilized red blood cell membrane (RBCM) to fabricate a permiselective glucose diffusion barrier based on transport of glucose by glucose transporter-1 in tandem with glucose dehydrogenase to create an amperometric glucose biosensor with high selectivity against interfering molecules (uric acid, ascorbic acid, cysteine, and galactose) (Kim et  al., 2018). These cellular bioreceptor systems show future promise for biocompatible, closed-loop feedback systems for sensing and regulation of insulin within the human body. Molecularly imprinted polymers (MIPs) are polymers which utilize the analyte of interest, acting as a template, to create a cavity in the polymer matrix with affinity toward a target molecule. Monomers are polymerized around the target analyte (via gelation, dispersion, bulk polymerization, etc.), which are then removed (usually by cleavage) and extracted via a solvent. The cavities left behind are expected to have high affinity for the template molecule. Several materials have been reported for synthesis of glucose sensitive MIPs, such as methacrylic acid (MAA) and ethylene glycol dimethyl acylate (EGDMA) (Widayani, 2017), poly(Nisopropylacrylamide-acrylamide-vinylphenylboronic acid) [poly(NIPAM-AAm-VPBA)] (Wang et  al., 2015a,b), O-phenylenediamine (o-PD) (Cheng et  al., 2001), and metal–organic copolymers (Qian et al., 2016). This nonenzymatic detection mechanism shows great promise due to the direct electron transfer (DET) functionality, freedom of molecular design and functional groups (to respond to different stimuli such as pH, electric field strength, etc.), high stability, high sensitivity, and low production cost (Alexander et al., 2017). Recently, other enzyme-free systems have been used for detection of glucose, including synthetic boronic acids as molecular receptors (Wang et al., 2013). These systems which do not rely on proteins are less susceptible to enzyme denaturation and short shelf-life. Boronic acid forms reversible covalent complexes of boronate esters with the cis-diols of sugars, which alter changes in optical properties or conductivity (Wang and Lee, 2015). Hence, boronic acid-containing sensors are usually optically or electrochemically based (Cambre and Sumerlin, 2011). Boronic acid derivatives have been directly immobilized on electrodes using various surface chemistries as well as electropolymerized with various

1269

conductive polymers to create a conductive network capable of glucose sensing (Wang et al., 2017; Çiftçi et al., 2013). Enzymes, which are macromolecular biological catalysts, are the most widely used bioreceptors in biosensing. Specifically, the most commonly used enzyme is glucose oxidase for the sensing of glucose. Biosensors utilizing enzymes as the molecular recognition layer offer several distinct advantages over other types of bioreceptors, including high sensitivity, high selectivity, and short response time (Mulchandani, 1998). However, variables such as bioreceptor film thickness, temperature, pH, and amount of active enzyme contribute to instability of electrode devices (Hossain and Park, 2017). Methods such as electropolymerization have been developed to control the thickness of conductive and nonconductive polymers using various electrochemical techniques (cyclic voltammetry, amperometry, etc.) (Mousa et  al., 2019). The known temperature and pH dependence of enzymatic electrochemical biosensors can be accounted for and have been shown to be accurate even outside of their optimum conditions (Rocchitta et al., 2016).

 Physicochemical Transducers

The transducer in a biosensor system converts physicochemical changes (temperature, light intensity, pH, conductivity, etc.) which take place at the bioreceptor into another type of energy signal through a process called “signalization” (Rocchitta et al., 2016). The signal measured should be proportional to the concentration (chemical potential) of the target analyte. The type of transduction used in the biosensor dictates the type of biosensor (i.e., electrochemical, optical, colorimetric, gravimetric). Examples of transducers are shown in Fig. 2.5.13.2. 

Computer Processing The computer processing and electronics of a biosensor system are comprised of various complex electronics used to digitize, amplify, filter, and multiplex electronic signals and present the processed signal in a display. Recently, many advances in miniaturization of electronics, low power consumption, and wireless communication (Bluetooth, ZigBee, RFID, NFC, etc.) have been utilized to either bring computer processing as well as power delivery “on-board” for implantable biosensors or wireless transmission of signal data for “off-board” processing (Kotanen C and Guiseppi-Elie, 2014; Jung et al., 2017; Zhao et al., 2017; Kassal et al., 2018). 

Types of Enzymatic Glucose Biosensors Electrochemical Biosensors A major percentage of developed enzymatic glucose biosensors used are electrochemical, because they offer sensitivity and reproducibility at a low cost that cannot be offered by many of the other sensing modalities. In fact, many of the commercially available point-of-care glucose biosensors measure glucose electrochemically (Yoo and Lee, 2010). Electrochemical biosensors can be grouped into four main transduction types: amperometric, conductimetric, impedimetric, and potentiometric.

1270 SEC T I O N 2. 5     Applications of Biomaterials

Bioreceptor

Analyte

Anbodies

∆ pH

Electron tunneling ∆ Light intensity

Cells

Amplificaon

Opcal

A/D Conversion

+

Output

Piezoelectric

∆ Microviscocity

Filtering

Piezoelectric crystals Surface acousc device Wave guide device

∆ Temperature

Enzymes

Electrochemical

Photon Counter Photodiode Fiber-opcs

∆ Mass

MIPs

Signal Processing

Electrodes Nanowire FETs Nanoparcles

∆ Conducvity

Nucleic Acids

Transducer

H(f)

Thermal Thermistor Thermopile

Heat Release Recognion event Physicochemical changes

f

Signalizaon

• Figure 2.5.13.2  General mechanism of biosensor operation. Analytes of interest are recognized by vari-

ous types of bioreceptors, where some form of energy transduction can be displayed as an output through computation.

Gluconolactone Glucose

FAD (Ox.)

FADH2 (Red.)

O2

Gluconolactone Glucose

FAD (Ox.)

FADH2 (Red.)

2Mox

H2 O 2

2e-

2Mred

Gluconolactone Glucose

FAD (Ox.)

FADH2 (Red.)

2e-

2e-

Electrode

Electrode

Electrode

Gen-1

Gen-2

Gen-3

• Figure 2.5.13.3  The three generations of enzymatic glucose sensors. (Adapted from Karunwi et al. (2013)).

Amperometric Biosensors Amperometric glucose enzyme biosensors measure the movement of electrons (current) produced in a redox reaction beginning with the oxidation of glucose to gluconolactone via GOx (the redox cofactor, FAD, serves as the electron acceptor) (Eq. 2.5.13.1). Gluconolactone is converted to gluconate− and H+ (Eq. 2.5.13.2) and H2O2 is converted to O2, 2H+ and electrons in the form of current (Eq. 2.5.13.3) (Weibel and Bright, 1971; Chou et al., 2018). Glucose + O2 + H2 O GOx → gluconolactone + H2 O2

(2.5.13.1) Gluconolactone→gluconate− +H+ (2.5.13.2)

H2 O2 → O2 + 2H + + 2e −

(2.5.13.3)

In an electrochemical cell, measurements can be made to monitor the consumption of oxygen or the production of hydrogen peroxide in the reaction, the current of which is directly proportional to the glucose concentration. Amperometric glucose biosensors have been divided into three generations, based on the mechanism of electron transfer used to measure the concentration of target analyte. The reader is directed to several reviews for a comprehensive discussion of the three generations (Rocchitta et al., 2016; Wang, 2008; Chen et al., 2013). Briefly, the three generations of amperometric glucose biosensor are summarized below and visualized in Fig. 2.5.13.3: • Generation 1: Generation 1 glucose biosensors monitor the consumption of oxygen or the production of hydrogen peroxide (Eq. 2.5.13.1) at the cathode or anode, respectively.

CHAPTER 2.5.13   Responsive Polymers in the Fabrication of Enzyme-Based Biosensors

Instrumentaon

(A)

1271

(B)

∆S

L W

Chemo-sensive enzyme/ polymer film

Conducvity (S)

(C)

[Glucose]

• Figure 2.5.13.4  (A) Typical interdigitated conductometric biosensor. A chemi-sensitive film is deposited

on an interdigitated electrode system. (B) Protonation and deprotonation of polyaniline as a result of hydrogen ion formation. (C) Typical conductometric glucose biosensor response.

• Generation 2: The second generation of glucose biosensors utilized a redox mediator, such as ferrocenes, viologens, quinones, etc., to shuttle electrons from the redox center of the GOx molecule to the electrode, effectively eliminating the need for oxygen. • Generation 3: The third generation of glucose biosensors involves a direct transport of the electron (DET) from the active site of the GOx to the electrode itself, without the need of a mediator. 

Conductometric Biosensors Conductometric enzyme biosensors measure the changes in conductivity of a membrane or medium, usually between two interdigitated electrodes spaced very close to each other (tens of microns) and sometimes a pair of interdigitated electrodes, one serving as measurement and one serving as reference (Nouira et al., 2013). Conductometric devices are sometimes presented as a subset of impedimetric devices but are discussed separately here. As glucose oxidase reacts with glucose, the local concentration of H+ ions increases (Eq. 2.5.13.3), resulting in an increase in acidity and conductivity near the electrode (Kucherenko et al., 2016). However, since most assays are done in buffered solutions, these ions are neutralized quickly and don’t take part in measurable net charge transport (Bănică, 2012). In fact, solution parameters such as ion strength, buffer capacity, and pH have a significant impact on solution conductivity and in turn the biosensor response (higher ion strength or buffer concentration tend to reduce sensitivity) (Pyeshkova et al., 2009). An elegant solution involves preparation of conductometric transducers utilizing chemoresistive materials, such as conductive polymers. The pH of responses of conductive polymers such as polyaniline and polypyrrole have been used as stimuli-responsive films in conductometric glucose biosensors (Miwa et  al., 1994). Polyaniline, which can be protonated to its conductive emeraldine salt or deprotonated to its insulating emeraldine base, has been used as films on interdigitated microelectrodes (IDEs) with GOx

(Fig. 2.5.13.4). In particular, the conductivity of the conductive polymer polyaniline is determined by the degree of protonation, and therefore unless in contact with a very strong acid, the conductivity can vary several orders of magnitude over several pHs. The resistance of such films is measured using small-amplitude AC signals (see “Impedimetric biosensor” section) or mV level DC signals, according to the type of instrumentation used. While conductometric biosensors offer advantages such as nonsensitivity to light, no need for a reference electrode, and low cost, many efforts have afforded poor signal-tonoise ratios in comparison with other types of enzymatic biosensors. 

Impedimetric Biosensors Based on the fact that changes in conductivity of a membrane or medium are a factor of resistance and reactance, impedimetric biosensors emerged as a means to deconvolute and monitor the capacitive and resistive changes occurring in a biosensor system (Katz and Willner, 2003). Electrochemical impedance spectroscopy (EIS) is a commonly used nondestructive measurement tool that has been utilized in applications including electrode characterization (Aggas et al., 2018), corrosion (Ribeiro and Abrantes, 2016), and biosensors (Kotanen et al., 2018). Impedance is measured by application of a small sinusoidal voltage (3% strain, soft bioelectronics are able to withstand strains that the human body experiences (∼30%) (Sekitani, 2016). Recently, Fuketa et al. developed a 1 μm thick ultraflexible surface EMG electrode using organic transistors capable of measurement of the myoelectric signal (Fuketa et al., 2014). The organic transistors microfabricated on polyethylene naphthalate (PEN) films boasted high flexibility, high electrode density, and high signal intensity. Several soft enzymatic biosensors utilizing similar fabrication methods have been created to measure analytes such as glucose/lactate in bodily fluids (sweat, saliva, tears). Of particular interest is a newly developed

Three-Dimensional (3-D) Bioprinting More recently, microfabrication techniques have utilized hydrogels (rather than semiconductive metals) to create 3-D tissue arrays, organ-on-chips, and biosensors (Verhulsel et  al., 2014). Micromolding, photolithography, stereolithography and 3-D bioprinting have led to construction of precise, 3-D hydrogel microstructures. Here, the focus of discussion is on the growing use of 3-D bioprinting. Detailed reviews on other fabrication methods can be found elsewhere (Yanagawa et al., 2016; Li et al., 2015; Vaezi et al., 2013). 3-D bioprinting utilizing an array of biocompatible materials has become of great interest recently, due to commercialization of 3-D bioprinting technologies. Supported by the growing catalog of “bio-inks,” 3-D bioprinting technology has the ability to produce biosensors without the need for masks required in photolithography (Sharafeldin et al., 2018). The inclusion of stimuli-responsive polymers within bioinks to create 3-D-printed biosensors has led to the so-called 4-D printed constructs: 3-D printed structures capable of biomolecular recognition by response to stimuli (Mandon et  al., 2017). This research area utilizes all concepts presented in this review and is indeed at the forefront of ongoing research. Li et  al. 3-D printed stretchable interdigitated PDMS electrodes with CNTs for use in electrochemical sensing (Kai et  al., 2018). Stretchable bioelectronic elements with

(B) 10.0

ECC MDEA 5037

8.0

Omnetics Quick Connect Connector Sense Region 1 (Ch1, Glucose) Sense Region 2 (Ch2, Lactate)

MDEA 5037 gold electrode biosensor

Amperometric Current (µA)

Wireless DualPotentiostat

16.0 Tissue Lactate (uA)

9.0

Blood Lactate (mM)

14.0

MAP (mmHg)

12.0

7.0

BLOOD

6.0 5.0

8.0

INTRAMUSCULAR

4.0

10.0

6.0

3.0

4.0

2.0

2.0

1.0

0.0

0.0 0

10

20 30 Bleeding Time (min)

• Figure 2.5.13.10  Microdisc electrode array (MDEA) architecture used for glucose and lactate biosensors capable of transmitting data wirelessly via Bluetooth (A). Amperometric readings of analyte measured from blood and intramuscularly (B) (Kotanen et al., 2012).

40

Mean Arterial Pressure (x10 mm Hg) Blood Lactate (mM)

(A)

flexible glove biosensor, termed “lab-on-a-glove” or “forensic finger” (Mishra et al., 2017). The system incorporates a microfabricated three-electrode enzymatic system produced directly on low-cost disposable polymer gloves capable of rapid detection of organophosphate and subsequent wireless transmission, as shown in Fig. 2.5.13.11. Common microfabrication steps have been combined with micromachining techniques to create 3-D microstructures for biosensing, termed microelectromechanical systems (MEMS), which have been utilized for the base electrodes in electrochemical, optical, and thermal glucose biosensors (Huang et al., 2009). 

CHAPTER 2.5.13   Responsive Polymers in the Fabrication of Enzyme-Based Biosensors

1281

• Figure 2.5.13.11  Flexible glove biosensor: fabrication, design, and performance. (A) Image of the serpentine stencil design employed for printing the glove-based stretchable device. (B) Schematic of (left) the biosensing scan finger (index finger) containing smiling face shape carbon-based counter (CE), working (WE) electrodes and Ag/AgCl-based reference electrode (RE), and (right) collecting thumb finger with its printed carbon pad; scale bar 10 mm. (C) Photographs of the biosensing index finger under 0% (left) and 50% (right) linear stretch; scale bar, 10 mm. (D) On-glove swiping protocol for sampling chemical threat residues from tomato and stainless steel surfaces. (E) On-glove sensing procedure by joining the index (scan) and thumb (collector) fingers to complete the electrochemical cell. (F, G) Photographs of the wearable glove biosensor, consisting of a sensing finger, containing the immobilized OPH enzyme layer, and the collector/sampling finger. The electrodes are connected via an adjustable ring bandage to the portable potentiostat (attached to the back of the hand) for on-site detection with wireless communication to a smartphone for rapid presentation of the voltammetric results. (Inset) Schematic of the interface between potentiostat and glove sensor. The connections consist of a (iii) velcro fabric containing (ii) the aluminum-tape based pins that are adjusted as a ring with the glove sensing connectors and (i) the wiring with the potentiostat (Mishra et al., 2017).

modulus closer to that of human tissue present a clear advantage over traditional solid electronics for both wearable and implantable biosensors. Gowers et al. 3-D printed a microfluidic system using FDA-approved microdialysis probes for continuous monitoring of glucose and lactate (Gowers et  al., 2015). Integration with existing technologies, such as cellphones for point-of-care diagnostics, has also been of great interest. Roda et al. introduced low-cost 3-D printed optics for a microfluidic chemiluminescence glucose sensor as an autonomous lab-on-a-chip (Roda et al., 2014).

 Future Outlook Since its inception, electrochemical detection of glucose has undergone radical development toward the goal of sensitive, selective, miniaturized sensors for point of care, bed-side, or implantable devices. The use of stimuli-responsive polymers

in biosensors has been studied for decades, but only recently have they been used to their full extent either integrated with enzymes or as transduction elements. The integration of conductive stimuli-responsive polymers in biosensors and a push toward autonomous sensing and self-sufficiency seems to be an area requiring further study. As new materials are created or discovered for biosensing, such as the little understood carbon dot, considerable work is required to functionalize and utilize these molecules in biosensing. With new developments of systems integration, including the boom in 3-D bioprinting within the last decade, there are new opportunities for low-cost, rapid prototyping of 4-D responsive biosensor architectures which were not achievable before. The pairing of these new types and architectures of biosensors, nanomaterials, and wireless communication and handheld devices, such as cell phones, remains a promising endeavor in telemedicine, remote health care, and biosensing.

1282 SEC T I O N 2. 5     Applications of Biomaterials

References Aggas, J.R., Harrell, W., Lutkenhaus, J., Guiseppi-Elie, A., 2018. Metal–polymer interface influences apparent electrical properties of nano-structured polyaniline films. Nanoscale 10 (2), 672–682. Aini, B.N., Siddiquee, S., Ampon, K., Rodrigues, K.F., Suryani, S., 2015. Development of glucose biosensor based on ZnO nanoparticles film and glucose oxidase-immobilized eggshell membrane. Sens. Bio Sens. Res. 4, 46–56. Alexander, S., Baraneedharan, P., Balasubrahmanyan, S., Ramaprabhu, S., 2017. Highly sensitive and selective non enzymatic electrochemical glucose sensors based on Graphene Oxide-Molecular Imprinted Polymer. Mater. Sci. Eng. C 78, 124–129. Alexeev, V.L., Sharma, A.C., Goponenko, A.V., Das, S., Lednev, I.K., Wilcox, C.S., Finegold, D.N., Asher, S.A., 2003. Anal. Chem. 75, 2316–2323. Anthony, G.-E., Chenghong, L., Ray, H.B., 2002. Direct electron transfer of glucose oxidase on carbon nanotubes. Nanotechnology 13 (5), 559. Ayranci, R., Kirbay, F.O., Demirkol, D.O., Ak, M., Timur, S., 2018. Methods Appl. Fluoresc. 6, 035012. Asal, M., Özen, Ö., Şahinler, M., 2018. Polatoğlu İ. Recent developments in enzyme, DNA and immuno-based biosensors. Sensors 18 (6), 1924. Bai, Y.-F., Xu, T.-B., Luong, J.H.T., Cui, H.-F., 2014. Direct electron transfer of glucose oxidase-boron doped diamond interface: a new solution for a classical problem. Anal. Chem. 86 (10), 4910–4918. Bandodkar, A.J., Wang, J., 2014. Non-invasive wearable electrochemical sensors: a review. Trends Biotechnol. 32 (7), 363–371. Barone, P.W., Yoon, H., Ortiz-García, R., Zhang, J., Ahn, J.-H., Kim, J.-H., et al., 2009. Modulation of single-walled carbon nanotube photoluminescence by hydrogel swelling. ACS Nano 3 (12), 3869–3877. Barsan, M.M., David, M., Florescu, M., Ţugulea, L., Brett, C.M.A., 2014. A new self-assembled layer-by-layer glucose biosensor based on chitosan biopolymer entrapped enzyme with nitrogen doped graphene. Bioelectrochemistry 99, 46–52. Basabe-Desmonts, L., Reinhoudt, D.N., Crego-Calama, M., 2007. Design of fluorescent materials for chemical sensing. Chem. Soc. Rev. 36 (6), 993–1017. Ben-Moshe, M., Alexeev, V.L., Asher, S.A., 2006. Fast responsive crystalline colloidal array photonic crystal glucose sensors. Anal. Chem. 78 (14), 5149–5157. Besic, S., Minteer, S.D., 2017. Micellar polymer encapsulation of enzymes. In: Minteer, S.D. (Ed.), Enzyme Stabilization and Immobilization: Methods and Protocols. Springer New York, New York, NY, pp. 93–108. Beyene, N.W., Moderegger, H., Kalcher, K., 2004. Simple and effective procedure for immobilization of oxidases onto MnO2-bulkmodified, screen-printed carbon electrodes. S. Afr. J. Chem. 57. Boztas, A.O., Guiseppi-Elie, A., 2009. Immobilization and release of the redox mediator ferrocene monocarboxylic acid from within cross-linked p (HEMA-co-PEGMA-co-HMMA) hydrogels. Biomacromolecules 10 (8), 2135–2143. Brahim, S., Narinesingh, D., Guiseppi-Elie, A., 2002. Polypyrrolehydrogel composites for the construction of clinically important biosensors. Biosens. Bioelectron. 17 (1–2), 53–59. Bănică, F.G., 2012. Electrical-Impedance-Based Sensors. Chemical Sensors and Biosensors. Bünsow, J., Enzenberg, A., Pohl, K., Schuhmann, W., Johannsmann, D., 2010. Electroanalysis 22, 978–984.

Cambre, J.N., Sumerlin, B.S., 2011. Biomedical applications of boronic acid polymers. Polymer 52 (21), 4631–4643. Chen, T., Chang, D.P., Liu, T., Desikan, R., Datar, R., Thundat, T., Berger, R., Zauscher, S., 2010. J. Mater. Chem. 20, 3391–3395. Chen, T., Chang, D.P., Liu, T., Desikan, R., Datar, R., Thundat, T., et al., 2010. Glucose-responsive polymer brushes for microcantilever sensing. J. Mater. Chem. 20 (17), 3391–3395. Chen, C., Xie, Q., Yang, D., Xiao, H., Fu, Y., Tan, Y., et al., 2013. Recent advances in electrochemical glucose biosensors: a review. RSC Adv. 3 (14), 4473–4491. Chen, L., Hwang, E., Zhang, J., 2018. Fluorescent nanobiosensors for sensing glucose. Sensors 18 (5), 1440. Cheng, Z., Wang, E., Yang, X., 2001. Capacitive detection of glucose using molecularly imprinted polymers. Biosens. Bioelectron. 16 (3), 179–185. Cho, I.-H., Lee, J., Kim, J., Kang, M-s, Paik, J., Ku, S., et al., 2018. Current technologies of electrochemical immunosensors: perspective on signal amplification. Sensors 18 (1), 207. Chou, F.-F., Chang, H.-W., Li, T.-L., Shih, J.-S., 2008. Piezoelectric crystal/surface acoustic wave biosensors based on fullerene C60 and enzymes/antibodies/proteins. J. Iran. Chem. Soc. 5 (1), 1–15. Chou, J., Yan, S., Liao, Y., Lai, C., Wu, Y., Wu, C., 2018. Remote detection for glucose and lactate based on flexible sensor array. IEEE Sens. J. 18 (8), 3467–3474. Chuang, C.-W., Shih, J.-S., 2001. Preparation and application of immobilized C60-glucose oxidase enzyme in fullerene C60-coated piezoelectric quartz crystal glucose sensor. Sens. Actuators, B Chem. 81 (1), 1–8. Çiftçi, H., Tamer, U., Teker, M.Ş., Pekmez, N.Ö., 2013. An enzyme free potentiometric detection of glucose based on a conducting polymer poly (3-aminophenyl boronic acid-co-3-octylthiophene). Electrochim. Acta 90, 358–365. Clark, L.C., Lyons, C., 1962. Electrode systems for continuous monitoring in cardiovascular surgery. Ann. N. Y. Acad. Sci. 102 (1), 29–45. Clark Jr, LC., inventor., Leland Jr, CC., assignee., 1970. Membrane Polarographic Electrode System and Method with Electrochemical Compensation. United States. Çolak, Ö., Yaşar, A., Çete, S., Arslan, F., 2012. Glucose biosensor based on the immobilization of glucose oxidase on electrochemically synthesized polypyrrole-poly(vinyl sulphonate) composite film by cross-linking with glutaraldehyde. Artif. Cells Blood Substit. Immobil. Biotechnol. 40 (5), 354–361. Collyer, S.D., Davis, F., Higson, S.P.J., 2010. Sonochemically fabricated microelectrode arrays for use as sensing platforms. Sensors 10 (5), 5090–5132. Costantini, F., Tiggelaar, R., Sennato, S., Mura, F., Schlautmann, S., Bordi, F., Gardeniers, H., Manetti, C., 2013. Glucose level determination with a multi-enzymatic cascade reaction in a functionalized glass chip. Analyst 138(17), 5019–5024. https://doi. org/10.1039/C3AN00806A. Dastider, S.G., Barizuddin, S., Wu, Y., Dweik, M., Almasri, M., January 2013. Impedance biosensor based on interdigitated electrode arrays for detection of low levels of E.coli O157:H7. In: 2013 IEEE 26th International Conference on Micro Electro Mechanical Systems (MEMS), vol. 2013, pp. 20–24. Degani, Y., Heller, A., 1987. Direct electrical communication between chemically modified enzymes and metal electrodes. I. Electron transfer from glucose oxidase to metal electrodes via electron relays, bound covalently to the enzyme. J. Phys. Chem. 91 (6), 1285–1289.

CHAPTER 2.5.13   Responsive Polymers in the Fabrication of Enzyme-Based Biosensors

Dhand, C., Das, M., Datta, M., Malhotra, B.D., 2011. Recent advances in polyaniline based biosensors. Biosens. Bioelectron. 26 (6), 2811–2821. Dübner, M., Cadarso, V.J., Gevrek, T.N., Sanyal, A., Spencer, N.D., Padeste, C., 2017. Reversible light-switching of enzymatic activity on orthogonally functionalized polymer brushes. ACS Appl. Mater. Interfaces 9 (11), 9245–9249. Ferraz, N., Strømme, M., Fellström, B., Pradhan, S., Nyholm, L., Mihranyan, A., 2012. In vitro and in vivo toxicity of rinsed and aged nanocellulose–polypyrrole composites. J. Biomed. Mater. Res. A 100A (8), 2128–2138. Flexer, V., Mano, N., 2014. Wired pyrroloquinoline quinone soluble glucose dehydrogenase enzyme electrodes operating at unprecedented low redox potential. Anal. Chem. 86 (5), 2465–2473. Fuketa, H., Yoshioka, K., Shinozuka, Y., Ishida, K., Yokota, T., Matsuhisa, N., et  al., 2014. 1μm-thickness ultra-flexible and high electrode-density surface electromyogram measurement sheet with 2 V organic transistors for prosthetic hand control. IEEE Trans. Biomed. Circuits Syst. 8 (6), 824–833. Gaudin, V., 2017. Advances in biosensor development for the screening of antibiotic residues in food products of animal origin – a comprehensive review. Biosens. Bioelectron. 90, 363–377. Gerard, M., Chaubey, A., Malhotra, B., 2002. Application of conducting polymers to biosensors. Biosens. Bioelectron. 17 (5), 345–359. Gowers, S.A.N., Curto, V.F., Seneci, C.A., Wang, C., Anastasova, S., Vadgama, P., et  al., 2015. 3D printed microfluidic device with integrated biosensors for online analysis of subcutaneous human microdialysate. Anal. Chem. 87 (15), 7763–7770. Gracia, R., Mecerreyes, D., 2013. Polymers with redox properties: materials for batteries, biosensors and more. Polym. Chem. 4 (7), 2206–2214. Grieshaber, D., MacKenzie, R., Vörös, J., Reimhult, E., 2008. Electrochemical biosensors – sensor principles and architectures. Sensors 8 (3), 1400–1458. Guiseppi-Elie, A., 2010. Electroconductive hydrogels: synthesis, characterization and biomedical applications. Biomaterials 31 (10), 2701–2716. Guiseppi-Elie, A., 2011. An implantable biochip to influence patient outcomes following trauma-induced hemorrhage. Anal. Bioanal. Chem. 399 (1), 403–419. Haarindraprasad, R., Hashim, U., Gopinath, S.C.B., Perumal, V., Liu, W.-W., Balakrishnan, S.R., 2016. Fabrication of interdigitated high-performance zinc oxide nanowire modified electrodes for glucose sensing. Anal. Chim. Acta 925, 70–81. Harsányi, G., 2000. Polymer films in sensor applications: a review of present uses and future possibilities. Sens. Rev. 20 (2), 98– 105. Hierlemann, A., Brand, O., Hagleitner, C., Baltes, H., 2003. Microfabrication techniques for chemical/biosensors. Proc. IEEE 91 (6), 839–863. Hoang Hiep, N., Moonil, K., 2017. An overview of techniques in enzyme immobilization. Appl. Sci. Converg. Technol. 26 (6), 157–163. Hossain, M.F., Park, J.Y., 2017. Fabrication of sensitive enzymatic biosensor based on multi-layered reduced graphene oxide added PtAu nanoparticles-modified hybrid electrode. PLoS One (3), 12 e0173553-e. Hu, J., Liu, S., 2010. Responsive polymers for detection and sensing applications: current status and future developments. Macromolecules 43 (20), 8315–8330.

1283

Huang, X., Li, S., Schultz, J.S., Wang, Q., Lin, Q., 2009. A MEMS affinity glucose sensor using a biocompatible glucose-responsive polymer. Sens. Actuators B Chem. 140 (2), 603–609. Idros, N., Ho, M., Pivnenko, M., Qasim, M., Xu, H., Gu, Z., et al., 2015. Colorimetric-based detection of TNT explosives using functionalized silica nanoparticles. Sensors 15 (6), 12891. Hahm, J-i, 2011. Functional polymers in protein detection platforms: optical, electrochemical, electrical, mass-sensitive, and magnetic biosensors. Sensors 11 (3), 3327–3355. James, T.D., 2006. Boronic Acid-Based Receptors and Sensors for Saccharides. Boronic Acids. Jenekhe, S.A., Kiserow, D.J., 2004. Chromogenic effects in polymers: an overview of the diverse ways of tuning optical properties in real time. In: Chromogenic Phenomena in Polymers. ACS Symposium Series, vol. 888. American Chemical Society, pp. 2–15. Jin, P., Yamaguchi, A., Oi, F.A., Matsuo, S., Tan, J., Misawa, H., 2001. Glucose sensing based on interdigitated array microelectrode. Anal. Sci. 17 (7), 841–846. Jung, J., Lee, J., Shin, S., Kim, Y.T., 2017. Development of a telemetric, miniaturized electrochemical amperometric analyzer. Sensors 17 (10), 2416. Justin, G., Finley, S., Abdur Rahman, A.R., Guiseppi-Elie, A., 2009. Biomimetic hydrogels for biosensor implant biocompatibility: electrochemical characterization using micro-disc electrode arrays (MDEAs). Biomed. Microdevices 11 (1), 103–115. Justino, C.I.L., Duarte, A.C., Rocha-Santos, T.A.P., 2017. Recent progress in biosensors for environmental monitoring: a review. Sensors 17 (12), 2918. Kai, L., Hong, W., Wenguang, L., Hong, M., Peixin, Z., Chaoyi, Y., 2018. 3D printed stretchable capacitive sensors for highly sensitive tactile and electrochemical sensing. Nanotechnology 29 (18), 185501. Karunwi, O., Guiseppi-Elie, A., 2013. Supramolecular glucose oxidase-swnt conjugates formed by ultrasonication: effect of tube length, functionalization and processing time. J. Nanobiotechnology 11(1), 6. https://doi.org/10.1186/1477-3155-11-6. Kassal, P., Steinberg, M.D., Steinberg, I.M., 2018. Wireless chemical sensors and biosensors: a review. Sens. Actuators B Chem. 266, 228–245. Katz, E., Willner, I., 2003. Probing biomolecular interactions at conductive and semiconductive surfaces by impedance spectroscopy: routes to impedimetric immunosensors, DNA-sensors, and enzyme biosensors. Electroanalysis 15 (11), 913–947. Khan, M.R.R., Khalilian, A., Kang, S.-W., 2016. Fast, highly-sensitive, and wide-dynamic-range interdigitated capacitor glucose biosensor using solvatochromic dye-containing sensing membrane. Sensors 16 (2) 265-. Kim, J., Kim, J.H., Ariga, K., 2017. Redox-Active Polymers for Energy Storage Nanoarchitectonics. Kim, I., Kwon, D., Lee, D., Lee, T.H., Lee, J.H., Lee, G., et al., 2018. A highly permselective electrochemical glucose sensor using red blood cell membrane. Biosens. Bioelectron. 102, 617–623. Klonoff, D.C., 2012. Overview of fluorescence glucose sensing: a technology with a bright future. J. Diabetes Sci. Technol. 6 (6), 1242–1250. Klotzbach, T.L., Watt, M., Ansari, Y., Minteer, S.D., 2008. Improving the microenvironment for enzyme immobilization at electrodes by hydrophobically modifying chitosan and Nafion® polymers. J. Membr. Sci. 311 (1), 81–88. Kotanen, C., Guiseppi-Elie, A., 2010. Development of an implantable biosensor system for physiological status monitoring during long duration space exploration. Gravitational Space Biol. 23 (2), 55–64.

1284 SEC T I O N 2. 5     Applications of Biomaterials

Kotanen, C., Guiseppi-Elie, A., 2014. Characterization of a Wireless Potentiostat for Integration with a Novel Implantable Biotransducer. 768–76 pp. Kotanen, C.N., Guiseppi-Elie, A., 2012. Bioactive electroconductive hydrogels yield novel biotransducers for glucose. Macromol. Symp. 317–318 (1), 187–197. Kotanen, C., Karunwi, O., Guiseppi-Elie, A., 2014. Biofabrication using pyrrole electropolymerization for the immobilization of glucose oxidase and lactate oxidase on implanted microfabricated biotransducers. Bioengineering 1 (1), 85. Kotanen, C.N., Karunwi, O., Alam, F., Uyehara, C.F.T., GuiseppiElie, A., 2018. Fabrication and in  vitro performance of a dual responsive lactate and glucose biosensor. Electrochim. Acta 267, 71–79. Kotanen, C.N., Moussy, F.G., Carrara, S., Guiseppi-Elie, A., 2012. Implantable enzyme amperometric biosensors. Biosens. Bioelectron. 35(1), 14–26. https://doi.org/10.1016/j.bios.2012.03.016. Kucherenko, I.S., Kucherenko, D.Y., Soldatkin, O.O., Lagarde, F., Dzyadevych, S.V., Soldatkin, A.P., 2016. A novel conductometric biosensor based on hexokinase for determination of adenosine triphosphate. Talanta 150, 469–475. Lee, K.M., Kim, K.H., Yoon, H., Kim, H., 2018. Chemical design of functional polymer structures for biosensors: from nanoscale to macroscale. Polymers 10 (5), 551. Li, C.-C., Kharaziha, M., Min, C., Maas, R., Nikkhah, M., 2015. Microfabrication of cell-laden hydrogels for engineering mineralized and load bearing tissues. In: Bertassoni, L.E., Coelho, P.G. (Eds.), Engineering Mineralized and Load Bearing Tissues. Springer International Publishing, Cham, pp. 15–31. Li, H., Zhao, F., Yue, L., Li, S., Xiao, F., 2016. Nonenzymatic electrochemical biosensor based on novel hydrophilic ferrocene-terminated hyperbranched polymer and its application in glucose detection. Electroanalysis 28 (5), 1003–1011. Li, H., Yan, X., Qiao, S., Lu, G., Su, X., 2018. Yellow-emissive carbon dot-based optical sensing platforms: cell imaging and analytical applications for biocatalytic reactions. ACS Appl. Mater. Interfaces 10 (9), 7737–7744. Lin, H., Li, M., Ding, L., Huang, J., 2019. Appl. Biochem. Biotechnol. 187, 1569–1580. Luong, J.H.T., Glennon, J.D., Gedanken, A., Vashist, S.K., 2017. Achievement and assessment of direct electron transfer of glucose oxidase in electrochemical biosensing using carbon nanotubes, graphene, and their nanocomposites. Microchimica Acta 184 (2), 369–388. Ma, Y., Promthaveepong, K., Li, N., 2016. Anal. Chem. 88, 8289– 8293. Maji, S., Cesur, B., Zhang, Z., De Geest, B.G., Hoogenboom, R., 2016. Polym. Chem. 7, 1705–1710. Mandon, C.A., Blum, L.J., Marquette, C.A., 2017. Adding biomolecular recognition capability to 3D printed objects: 4D printing. Procedia Technol. 27, 1–2. Mano, N., Yoo, J.E., Tarver, J., Loo, Y.-L., Heller, A., 2007. An electron-conducting cross-linked polyaniline-based redox hydrogel, formed in one step at pH 7.2, wires glucose oxidase. J. Am. Chem. Soc. 129 (22), 7006–7007. Mariani, A.M., Natoli, M.E., Kofinas, P., 2013. Enzymatic activity preservation and protection through entrapment within degradable hydrogels. Biotechnol. Bioeng. 110 (11), 2994–3002. Microfabricated biosensors and microsystems. In: Hesketh, P.J., Zivanovic, S., Ming, Y., Park, S., Svojanovsky, S., Cunneen, J., et al. (Eds.), Sept. 1997. 1997 21st International Conference on Microelectronics Proceedings, Sept. 1997. vol. 1997, pp. 14–17.

Mishra, R.K., Hubble, L.J., Martín, A., Kumar, R., Barfidokht, A., Kim, J., et al., 2017. Wearable flexible and stretchable glove biosensor for on-site detection of organophosphorus chemical threats. ACS Sens. 2 (4), 553–561. Miwa, Y., Miyake, T., Matsue, T., Uchida, I., 1994. A Conductometric Glucose Sensor Based on a Twin-Microband Electrode Coated with a Polyaniline Thin Film. 2864–2866 pp. Mousa, H.M., Aggas, J.R., Guiseppi-Elie, A., 2019. Electropolymerization of aniline and (N-phenyl-o-phenylenediamine) for glucose biosensor application. Mater. Lett. 238, 267–270. Mugo, S.M., Berg, D., Bharath, G., 2019. Anal. Lett. 52, 825–838. Mulchandani, A., 1998. Principles of enzyme biosensors. In: Mulchandani, A., Rogers, K.R. (Eds.), Enzyme and Microbial Biosensors: Techniques and Protocols. Humana Press, Totowa, NJ, pp. 3–14. Naderi Asrami, P., Mozaffari, S.A., Saber Tehrani, M., Aberoomand Azar, P., 2018. A novel impedimetric glucose biosensor based on immobilized glucose oxidase on a CuO-Chitosan nanobiocomposite modified FTO electrode. Int. J. Biol. Macromol. 118, 649–660. Nagel, B., Warsinke, A., Katterle, M., 2007. Langmuir 23, 6807– 6811. Neethirajan, S., Weng, X., Tah, A., Cordero, J.O., Ragavan, K.V., 2018. Nano-biosensor platforms for detecting food allergens – new trends. Sens. Bio Sens. Res. 18, 13–30. Nguyen, Q.V., Caro, A., Raoux, M., Quotb, A., Floderer, J., Bornat, Y., et al. (Eds.), July 2013. A Novel Bioelectronic Glucose Sensor to Process Distinct Electrical Activities of Pancreatic Beta-Cells. 2013 35th Annual International Conference of the IEEE Engineering in Medicine and Biology Society (EMBC), vol. 2013, pp. 3–7. Nouira, W., Maaref, A., Elaissari, H., Vocanson, F., Siadat, M., Jaffrezic-Renault, N., 2013. Comparative study of conductometric glucose biosensor based on gold and on magnetic nanoparticles. Mater. Sci. Eng. C 33 (1), 298–303. Pan, L., Yu, G., Zhai, D., Lee, H.R., Zhao, W., Liu, N., et al., 2012. Hierarchical nanostructured conducting polymer hydrogel with high electrochemical activity. Proc. Natl. Acad. Sci. U.S.A. 109 (24), 9287–9292. Pan, H., Gonuguntla, S., Li, S., Trau, D., 2017. Conjugated Polymers for Biosensor Devices II. Pänke, O., Balkenhohl, T., Kafka, J., Schäfer, D., Lisdat, F., 2007. Impedance spectroscopy and biosensing. In: Biosensing for the 21st Century: Springer, pp. 195–237. Park, H.S., Cho, M.Y., Noh, Y.-W., Hong, K.S., Lim, Y.T., 2017. Dyes Pigments 136, 583–589. Parrilla, M., Cánovas, R., Andrade, F.J., 2017. Paper-based enzymatic electrode with enhanced potentiometric response for monitoring glucose in biological fluids. Biosens. Bioelectron. 90, 110–116. Pham, X.-H., Bui, M.-P.N., Li, C.A., Han, K.N., Kim, J.H., Won, H., Seong, G.H., 2010. Anal. Chim. Acta 671, 36–40. Pietschnig, R., 2016. Polymers with pendant ferrocenes. Chem. Soc. Rev. 45 (19), 5216–5231. Pisoschi, A., 2016. Potentiometric biosensors: concept and analytical applications—an editorial. Biochem. Anal. Biochem. 5 e164. Pohanka, M., 2018. Overview of piezoelectric biosensors, immunosensors and DNA sensors and their applications. Materials 11 (3), 448. Putzbach, W., Ronkainen, N.J., 2013. Immobilization techniques in the fabrication of nanomaterial-based electrochemical biosensors: a review. Sensors 13 (4), 4811–4840. Pyeshkova, V., Saiapina, O., Soldatkin, O., Dzyadevych, S., 2009. Enzyme Conductometric Biosensor for Maltose Determination. 272–8 pp.

CHAPTER 2.5.13   Responsive Polymers in the Fabrication of Enzyme-Based Biosensors

Qian, K., Deng, Q., Fang, G., Wang, J., Pan, M., Wang, S., et al., 2016. Metal–organic frameworks supported surface–imprinted nanoparticles for the sensitive detection of metolcarb. Biosens. Bioelectron. 79, 359–363. Ribeiro, D.V., Abrantes, J.C.C., 2016. Application of electrochemical impedance spectroscopy (EIS) to monitor the corrosion of reinforced concrete: a new approach. Constr. Build. Mater. 111, 98–104. Rocchitta, G., Spanu, A., Babudieri, S., Latte, G., Madeddu, G., Galleri, G., et al., 2016. Enzyme biosensors for biomedical applications: strategies for safeguarding analytical performances in biological fluids. Sensors 16 (6), 780. Roda, A., Guardigli, M., Calabria, D., Calabretta, M.M., Cevenini, L., Michelini, E.A., 2014. 3D-printed device for a smartphonebased chemiluminescence biosensor for lactate in oral fluid and sweat. Analyst 139 (24), 6494–6501. Rukiye, A., Fatma Ozturk, K., Dilek Odaci, D., Metin, A., Suna, T., 2018. Copolymer based multifunctional conducting polymer film for fluorescence sensing of glucose. Methods Appl. Fluoresc. 6 (3), 035012. Ruscito, A., DeRosa, M.C., 2016. Small-molecule binding aptamers: selection strategies, characterization, and applications. Front. Chem. 4, 14. Russell, T.P., 2002. Surface-responsive materials. Science 297 (5583), 964–967. Saifuddin, N., Raziah, A.Z., Junizah, A.R., 2013. Carbon nanotubes: a review on structure and their interaction with proteins. J. Chem. 2013, 18. Sau, A., Bera, K., Pal, U., Maity, A., Mondal, P., Basak, S., et  al., 2018. Design and synthesis of fluorescent carbon-dot polymer and deciphering its electronic structure. J. Phys. Chem. C 122 (41), 23799–23807. Saxena, A.P., Deepa, M., Joshi, A.G., Bhandari, S., Srivastava, A.K., 2011. Poly(3,4-ethylenedioxythiophene) – ionic liquid functionalized graphene/reduced graphene oxide nanostructures: improved conduction and electrochromism. ACS Appl. Mater. Interfaces 3 (4), 1115–1126. Schmidt, U., Guenther, M., Gerlach, G., 2016. Curr. Dir. Biomed. Eng. 117. Schuhmann, W., Ohara, T.J., Schmidt, H.L., Heller, A., 1991. Electron transfer between glucose oxidase and electrodes via redox mediators bound with flexible chains to the enzyme surface. J. Am. Chem. Soc. 113 (4), 1394–1397. Sekitani, T., 2016. Soft biosensor systems using flexible and stretchable electronics technology. In: Rogers, J.A., Ghaffari, R., Kim, D.-H. (Eds.), Stretchable Bioelectronics for Medical Devices and Systems. Springer International Publishing, Cham, pp. 133–149. Shakya, A.K., Nandakumar, K.S., 2018. An update on smart biocatalysts for industrial and biomedical applications. Journal of the Royal Society. Interface 15 (139), 20180062. Shan, X., Chai, L., Ma, J., Qian, Z., Chen, J., Feng, H., 2014. B-doped Carbon Quantum Dots as a Sensitive Fluorescence Probe for Hydrogen Peroxide and Glucose Detection. Sharafeldin, M., Jones, A., Rusling, J., 2018. 3D-Printed biosensor arrays for medical diagnostics. Micromachines 9 (8), 394. Singh, P., Shukla, S.K., 2018. Opto-chemical glucose sensing over NiO/polyaniline hybrid matrix using optical fiber approach. Optik 165, 94–101. Sirca, D., Vardeu, A., Pinna, M., Diana, M., Enrico, P., 2014. A robust, state-of-the-art amperometric microbiosensor for glutamate detection. Biosens. Bioelectron. 61, 526–531.

1285

Sirkar, K., Revzin, A., Pishko, M.V., 2000. Glucose and lactate biosensors based on redox polymer/oxidoreductase nanocomposite thin films. Anal. Chem. 72 (13), 2930–2936. Sodzel, D., Khranovskyy, V., Beni, V., Turner, A.P., Viter, R., Eriksson, M.O., et al., 2015. Continuous sensing of hydrogen peroxide and glucose via quenching of the UV and visible luminescence of ZnO nanoparticles. Microchimica Acta 182 (9–10), 1819–1826. Swamy, N.K., Sandeep, S., Santhosh, A., 2017. Conductive polymers and their nanohybrid transducers for electrochemical biosensors applications: a brief review. Indian J. Adv. Chem. Sci. S2 6, 9. Syshchyk, O., Skryshevsky, V.A., Soldatkin, O.O., Soldatkin, A.P., 2015. Enzyme biosensor systems based on porous silicon photoluminescence for detection of glucose, urea and heavy metals. Biosens. Bioelectron. 66, 89–94. Takasu, K., Kushiro, K., Hayashi, K., Iwasaki, Y., Inoue, S., Tamechika, E., et  al., 2015. Polymer brush biointerfaces for highly sensitive biosensors that preserve the structure and function of immobilized proteins. Sens. Actuators B Chem. 216, 428–433. Tavakoli, J., Tang, Y., 2017. Hydrogel based sensors for biomedical applications: an updated review. Polymers 9 (8), 364. Tereshchenko, A., Bechelany, M., Viter, R., Khranovskyy, V., Smyntyna, V., Starodub, N., et al., 2016. Optical biosensors based on ZnO nanostructures: advantages and perspectives. A review. Sens. Actuators B Chem. 229, 664–677. Thyparambil, A.A., Bazin, I., Guiseppi-Elie, A., 2017. Molecular modeling and simulation tools in the development of peptidebased biosensors for mycotoxin detection: example of ochratoxin. Toxins 9 (12), 395. Tierney, S., Volden, S., Stokke, B.T., 2009. Glucose sensors based on a responsive gel incorporated as a Fabry-Perot cavity on a fiber-optic readout platform. Biosens. Bioelectron. 24 (7), 2034–2039. Tuller, H., 2017. Ionic conduction and applications. In: Kasap, S., Capper, P. (Eds.), Springer Handbook of Electronic and Photonic Materials. Springer International Publishing, Cham. 1-. Usman Ali, S.M., Nur, O., Willander, M., Danielsson, B., 2010. A fast and sensitive potentiometric glucose microsensor based on glucose oxidase coated ZnO nanowires grown on a thin silver wire. Sens. Actuators, B Chem. 145 (2), 869–874. Vaezi, M., Seitz, H., Yang, S., 2013. A review on 3D micro-additive manufacturing technologies. Int. J. Adv. Manuf. Technol. 67 (5), 1721–1754. Vargas-Bernal, R., Rodríguez-Miranda, E., Herrera-Pérez, G., 2012. Evolution and Expectations of Enzymatic Biosensors for Pesticides, pp. 329–356. Vasylieva, N., Maucler, C., Meiller, A., Viscogliosi, H., Lieutaud, T., Barbier, D., et al., 2013. Immobilization method to preserve enzyme specificity in biosensors: consequences for brain glutamate detection. Anal. Chem. 85 (4), 2507–2515. Verhulsel, M., Vignes, M., Descroix, S., Malaquin, L., Vignjevic, D.M., Viovy, J.-L., 2014. A review of microfabrication and hydrogel engineering for micro-organs on chips. Biomaterials 35 (6), 1816–1832. Wahab, H.A., Salama, A.A., El Saeid, A.A., Willander, M., Nur, O., Battisha, I.K., 2018. Zinc oxide nano-rods based glucose biosensor devices fabrication. Res. Phys. 9, 809–814. Wan, D., Yuan, S., Li, G.L., Neoh, K.G., Kang, E.T., 2010. Glucose biosensor from covalent immobilization of chitosan-coupled carbon nanotubes on polyaniline-modified gold electrode. ACS Appl. Mater. Interfaces 2 (11), 3083–3091. Wang, J., 1999. Sol–gel materials for electrochemical biosensors. Anal. Chim. Acta 399 (1), 21–27.

1286 SEC T I O N 2. 5     Applications of Biomaterials

Wang, J., 2008. Electrochemical glucose biosensors. Chem. Rev. 108 (2), 814–825. Wang, H.-C., Zhou, H., Chen, B., Mendes, P.M., Fossey, J.S., James, T.D., Long, Y.-T., 2013. Analyst 138, 7146–7151. Wang, H.-C., Lee, A.-R., 2015. Recent developments in blood glucose sensors. J. Food Drug Anal. 23 (2), 191–200. Wang, X., Uchiyama, S., 2013. Polymers for biosensors construction. In: State of the Art in Biosensors-General Aspects InTech, Chapter 3. 67–84. Wang, J., Myung, N.V., Yun, M., Monbouquette, H.G., 2005. Glucose oxidase entrapped in polypyrrole on high-surface-area Pt electrodes: a model platform for sensitive electroenzymatic biosensors. J. Electroanal. Chem. 575 (1), 139–146. Wang, H.-C., Zhou, H., Chen, B., Mendes, P.M., Fossey, J.S., James, T.D., et al., 2013. A bis-boronic acid modified electrode for the sensitive and selective determination of glucose concentrations. Analyst 138 (23), 7146–7151. Wang, H., Yi, J., Velado, D., Yu, Y., Zhou, S., 2015a. Immobilization of carbon dots in molecularly imprinted microgels for optical sensing of glucose at physiological pH. ACS Appl. Mater. Interfaces 7 (29), 15735–15745. Wang, H., Ohnuki, H., Endo, H., Izumi, M., 2015b. Impedimetric and amperometric bifunctional glucose biosensor based on hybrid organic–inorganic thin films. Bioelectrochemistry 101, 1–7. Wang, W., Kong, L., Zhu, J., Tan, L., 2017. One-pot preparation of conductive composite containing boronic acid derivative for nonenzymatic glucose detection. J. Colloid Interface Sci. 498, 1–8. Weibel, M.K., Bright, H.J., 1971. The glucose oxidase mechanism: interpretation of the pH dependence. J. Biol. Chem. 246 (9), 2734–2744. Welch, M.E., Doublet, T., Bernard, C., Malliaras, G.G., Ober, C.K., 2015. A glucose sensor via stable immobilization of the GOx enzyme on an organic transistor using a polymer brush. J. Polym. Sci. Part A Polym. Chem. 53 (2), 372–377. Widayani, Y., Wungu, T.D.K., Suprijadi, 2017. Preliminary study of molecularly imprinted polymer-based potentiometric sensor for glucose. Procedia Eng. 170, 84–87. Wilson, A.M., Justin, G., Guiseppi-Elie, A., 2010. Electroconductive Hydrogels. Biomedical Applications of Hydrogels Handbook: Springer, pp. 319–337. Xiao, Y., Patolsky, F., Katz, E., Hainfeld, J.F., Willner, I., 2003. “Plugging into enzymes”: nanowiring of redox enzymes by a gold nanoparticle. Science 299 (5614), 1877–1881. Xie, W., Bülow, L., Xie, B., 2018. Pyrroloquinoline quinone glucose dehydrogenase adopted in thermometric analysis for enhancement of glucose determination. J. Therm. Anal. Calorim. 134 (3), 1913–1919. Xiong, M., Gu, B., Zhang, J.-D., Xu, J.-J., Chen, H.-Y., Zhong, H., 2013. Glucose microfluidic biosensors based on reversible enzyme

immobilization on photopatterned stimuli-responsive polymer. Biosens. Bioelectron. 50, 229–234. Xu, Z.J., 2017. From two-phase to three-phase: the new electrochemical interface by oxide electrocatalysts. Nano-Micro Lett. 10 (1), 8. Yakovleva, M., Bhand, S., Danielsson, B., 2013. The enzyme thermistor—a realistic biosensor concept. A critical review. Anal. Chim. Acta 766, 1–12. Yanagawa, F., Sugiura, S., Kanamori, T., 2016. Hydrogel microfabrication technology toward three dimensional tissue engineering. Regen. Ther. 3, 45–57. Yang, L., Guiseppi-Wilson, A., Guiseppi-Elie, A., 2010. Design Considerations in the Use of Interdigitated Microsensor Electrode Arrays (IMEs) for Impedimetric Characterization of Biomimetic Hydrogels. 279–89 pp. Yi-Hua, Z., Tse-Chao, H., Fei, X. (Eds.), Jan. 2006. A Thermal Biosensor Based on Enzyme Reaction. 2005 IEEE Engineering in Medicine and Biology 27th Annual Conference, vol. 2005, pp. 17–18. Yoo, E.-H., Lee, S.-Y., 2010. Glucose biosensors: an overview of use in clinical practice. Sensors 10 (5), 4558–4576. Yu, M., Zhao, K., Zhu, X., Tang, S., Nie, Z., Huang, Y., Zhao, P., Yao, S., 2017. Biosens. Bioelectron. 95, 41–47. Qu, Z-b, Zhou, X., Gu, L., Lan, R., Sun, D., Yu, D., et al., 2013. Boronic Acid Functionalized Graphene Quantum Dots as a Fluorescent Probe for Selective and Sensitive Glucose Determination in Microdialysate. Zayats, M., Katz, E., Willner, I., 2002. Electrical contacting of glucose oxidase by surface-reconstitution of the apo-protein on a relay-boronic acid-FAD cofactor monolayer. J. Am. Chem. Soc. 124 (10), 2120–2121. Zhang, M., Liao, C., Mak, C.H., You, P., Mak, C.L., Yan, F., 2015. Highly sensitive glucose sensors based on enzyme-modified wholegraphene solution-gated transistors. Sci. Rep. 5, 8311. Zhao, X., Sadhu, V., Le, T., Pompili, D., Javanmard, M. (Eds.), May 2017. Towards Low-Power Wearable Wireless Sensors for Molecular Biomarker and Physiological Signal Monitoring. 2017 IEEE International Symposium on Circuits and Systems (ISCAS), vol. 2017, pp. 28–31. Zhou, M., Dong, S., 2011. Bioelectrochemical interface engineering: toward the fabrication of electrochemical biosensors, biofuel cells, and self-powered logic biosensors. Acc. Chem. Res. 44 (11), 1232–1243. Zhou, Y., Chiu, C.-W., Liang, H., 2012. Interfacial structures and properties of organic materials for biosensors: an overview. Sensors 12 (11), 15036. Zohourtalab, A., Razmi, H., 2018. Selective determination of glucose in blood plasma by using an amperometric glucose biosensor based on glucose oxidase and a chitosan/Nafion/IL/ferrocene composite film. Iran. J. Anal. Chem. 5 (1), 9–16.

S E C TI ON 2 . 6

Applications of Biomaterials in Functional Tissue Engineering

2.6.1

Rebuilding Humans Using Biology and Biomaterials SHELLY E. SAKIYAMA-ELBERT Department of Biomedical Engineering, The University of Texas at Austin, Austin, TX, United States

O

ver the past four decades, biomaterials have been increasingly used for applications in tissue engineering to rebuild tissues lost due to disease or injury, as well as to build in vitro models of complex tissue systems. For applications in tissue engineering, our growing understanding of how biological systems function at a mechanistic level and our ability to modulate them at both genetic and epigenetic levels have allowed the development of more elegant approaches to tissue engineering. However, underlying these approaches is still the critical need to understand how the biomaterials used for these applications interface with the physiology environment. This section focuses on the fundamental principles and approaches used for designing and fabricating biomaterials for tissueengineering applications, stressing common themes and challenges. This edition of the textbook has refined the focus more on the materials, so the chapters in this section stress examples and a model selection of tissues rather than seeking to cover all topics. For additional discussion of specifics for given tissues, there are a variety of references on the topic of tissue engineering, which we do not seek to reproduce in this section. The first half of this section focuses on tissue engineering broadly, the materials used for tissue-engineering applications, and the transmission of mechanical cues between materials and tissues. The first chapter focuses on a general overview of tissue engineering, including a brief history and common components of tissue-engineering approaches. It highlights common applications, including new areas such as organoids, and newer approaches, such as genome editing. The second chapter focuses specifically on the use of biomaterial scaffolds for tissue-engineering applications. It highlights the key scaffold design criteria, applications,

and materials. It also highlights options for fabrication techniques (which are covered extensively in Section 1.4) and characterization methods. The third chapter focuses on micromechanical criteria of biomaterials scaffolds for tissue engineering. It covers mechanotransduction from scaffolds and extracellular matrices to cells and tissues. It also discusses the effects of scaffold topography and roughness on transduction, as well as the effects of mechanical stimulation on tissue formation. The remaining three chapters in this section provide examples of three tissues, and the thought processes behind selecting materials and designing biomaterial scaffolds for tissue-engineering applications in each area: musculoskeletal (tendon), cardiovascular, and soft tissues. The fourth chapter focuses on the design of scaffolds for the tissue engineering of tendons, as an example of a tissue where mechanical properties of the resulting tissue are critical for the return of load-bearing function after injury. Examples of using material fiber alignment to guide cell infiltration and extracellular matrix deposition are given that apply broadly to many tissues. The fifth chapter focuses on biomaterials used for cardiovascular tissue engineering. It highlights different types of materials used for different applications along with the advantages and disadvantages of those materials for each type of application. The final chapter in this section focuses on the use of biomaterials in soft tissue engineering, such as skin, adipose, and gastrointestinal tissue. The pros and cons of different materials are highlighted as well as design criteria for each type of tissue with regard to scaffold design. Overall this is a large research and commercial space for biomaterials-derived products and those interested in this field can find extensive primary literature, reviews, and texts on the topics, which are further described in Appendix E. 1287

2.6.2

Overview of Tissue Engineering Concepts and Applications SARAH MIHO VAN BELLEGHEM 1 ,2 *, BHUSHAN MAHADIK 1,2 *, KIRSTIE LANE SNODDERLY 1,2 *, JOHN P. FISHER 1 ,2 1University

of Maryland College Park

2NIH/NIBIB

Center for Engineering Complex Tissues

General Introduction History of Tissue Engineering The term “tissue engineering” as recognized today was likely first introduced at a panel meeting of the National Science Foundation in 1987, which led to the first tissue-engineering meeting in early 1988 (Vacanti, 2006). Despite these early approaches for replacement, repair, and regeneration of failing organs, the true emergence of tissue engineering as a medical field began in the early 1990s when tissue engineering was defined as an interdisciplinary field that applies the principles of engineering and life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function (Langer and Vacanti, 1993). The field has since rapidly progressed worldwide with an increasing market size, with nearly $4 billion invested in the field in the 1990s resulting in over 70 companies and several products on the market by the end of the millennium (Lysaght and Reyes, 2001). From a commercial perspective, tissue engineering carries a broad definition and can include companies focused on several products such as cells/stem cells, biomaterials, preclinical testing, clinical trials, commercially approved products, and specialized services (e.g., cord blood banking, tissue grafts). By 2011, 106 companies selling tissue-engineering-based products or services and employing almost 14,000 people were documented worldwide, with a total estimated sales of $3.5 billion, a roughly threefold increase from 2007. As of mid-2018, in the United States alone there were 49 public tissue-engineering companies with nearly 146,000 employees and a cumulative of $9 billion in sales. Well over a million patients have been treated with tissue-engineered products. Additionally, 66 interventional clinical trials in various phases involving * all authors contributed equally.

tissue-engineering products were active between 2011 and 2018 and both public and private sector companies continue R&D investment for new products and technologies (Kim et al., 2019). While early success has been challenging, technological advances alongside biological discoveries continue to propel the field of tissue engineering into exciting new frontiers (Vacanti, 2006).

 Goals of Tissue Engineering and Classification Goals of Tissue Engineering Tissue engineering aims to restore tissue and organ function by employing biological and engineering strategies to clinical problems. The functional failure of tissues and organs is a severe and costly healthcare problem as their replacements are limited by the availability of compatible donors. Artificial prostheses and mechanical devices help millions of patients; however, they are not ideal due to poor long-term performance. Furthermore, mechanical devices rarely integrate with host tissues and can trigger a host immune response that may damage healthy tissue around the implant. In addition, surgical reconstruction of organs and tissues involves the replacement of damaged tissue from the patient’s own tissue, e.g., saphenous vein as bypass graft, patella tendon for anterior cruciate ligament repair, and autologous skin grafts. However, often this strategy fails to replace all the functions of the original tissue and can cause issues such as development of malignant tumors, surgical complications, and morbidity at the donor sites. Thus tissue engineering has emerged as an alternative for tissue or organ transplantation with the primary goal of providing a clinically relevant substitute by integrating engineering, biology, materials science, chemistry, and medicine. By recapitulating the normal tissue development process, tissue engineering represents a 1289

1290 SEC T I O N 2. 6    Applications of Biomaterials in Functional Tissue Engineering

• Figure 2.6.2.1  Tissue engineering approaches for regeneration can consist of various methods. In acellular grafts donor-derived or synthetic biomaterial platforms devoid of any cellular components are implanted into the patient body to promote and aid native regeneration. Cellular grafts consist of patient-derived or donor cells that are used to cellularize and mature a scaffold prior to implantation. Cell therapy consists only of the desired cell and biological populations devoid of any scaffolds, that is administered to the patient.

strategy to restore, maintain, and improve tissue function, with the ultimate aim of complete tissue or organ regeneration. 

Classification of Tissue-Engineering Approaches Strategies for the clinical implementation of tissue-engineering methods can be broadly classified according to the implantation or application of the materials or biologics. In addition to traditional tissue engineering, cell therapies and closed looped systems used as implantable or extracorporeal devices are also examples of regenerative medicine approaches. While classically these methods are not regarded as tissue engineering, they have contributed significantly to tissue regeneration and are therefore briefly introduced here. Traditional tissue engineering: Traditional tissue engineering includes two main approaches: (1) transplantation of a tissue grown in  vitro consisting of an artificial matrix with cells and growth factors, and (2) in situ regeneration of tissue utilizing a combination of an artificial matrix and growth factors as a guiding template to induce host cell regeneration of the tissue in vivo (Fig. 2.6.2.1). These two strategies and their components are explained in detail throughout this chapter. Cell therapy: The main strategy in cell therapy is the harvesting of desired cell populations and their expansion in

large numbers for in vivo transplantation. This involves the delivery of cells through systemic injection into the bloodstream or through direct transplantation into a local tissue. A few examples of cell therapy include cell transplants from bone marrow, peripheral blood, or umbilical cord, which have been used to treat several blood-related diseases, including leukemia, multiple myeloma, and immune deficiencies (Rao et al., 2012). The main challenges for cell transplantation are growing large numbers of cells without bacterial contamination; preservation of cell phenotype; and preventing accumulation of genetic mutations during culture expansion. These challenges must be combated for cell therapy to become clinically successful and federally approved. Closed-loop methods: This involves the use of extracorporeal or implantable devices that house transplanted cell-laden polymers within a semipermeable membrane, which allows diffusion of nutrients and excreted products, but prevents the movement of antibodies, pathogens, or immunocompetent cells (Murua et  al., 2008). Microcarrier or microcapsule-based cell encapsulation can function as an effective drug delivery system and requires the use of biomaterials engineered for the desired cell–matrix interactions, therapeutic drug release, and transport properties to influence the targeted cell population, defect, or wound site (Hernández et  al., 2010). Closed-loop extracorporeal

CHAPTER 2.6.2   Overview of Tissue Engineering Concepts and Applications

devices have been used for the treatment of liver, pancreas, and kidney pathologies (Patzer, 2006; Hueso et al., 2019). Natural as well as synthetic biomaterials such as alginate, collagen, and polyethylene glycol (PEG)-based systems have been explored as carriers where factors such as material degradation, immunogenicity, encapsulation, and release kinetics, among others, have been studied to achieve the desired therapeutic effect (Hernández et al., 2010).

 Components of Tissue Engineering Tissue-engineering strategies typically involve multiple components, including cells, a physical template (scaffold), and a combination of biological cues that promote regeneration and integration of the construct into a functional and organized tissue.

The Cell Cells are the building blocks of tissues and play a critical role in promoting tissue healing and regeneration. Within tissue engineering, cells may be a component of the in vitro construct or may be recruited in vivo with the aid of immobilized or soluble signals. Cell types utilized for tissue engineering are selected from a variety of sources, which include autologous cells from the patient, allogeneic cells from another human, and xenogeneic cells from a different species. However, allogeneic and xenogeneic cells often suffer from immune rejection. Cell type: Common cell types utilized for tissue-engineering purposes include stem cells, differentiated mature cells, or a mixture of differentiated cells. Stem cells are capable of self-renewal and differentiation into multiple lineages and may include embryonic stem cells (ESCs), adult stem cells such as mesenchymal stem cells (MSCs) and hematopoietic stem cells, or induced pluripotent stem cells: • Embryonic stem cells: ESCs are capable of self-renewal without differentiation, can be culture expanded, and most importantly can differentiate into any cell type. Since ESCs are isolated from an embryonic stage, they can develop into any of the three germ layers: endoderm (interior stomach lining, gastrointestinal tract, and lungs); mesoderm (muscle, bone, blood, and urogenital); or ectoderm (epidermal tissues and nervous system) (Song et  al., 2018). However, because of several ethical issues and regulatory limitations concerning human ESCs, induced pluripotent stem cells (iPSCs) are now an attractive alternative in tissue engineering for clinical approaches. • Adult stem cells: Adult stem cells include, for example, hematopoietic, mesenchymal, neural, and hepatic stem cells. In particular, hematopoietic stem cells have been used in clinics for a few decades for treating blood diseases (i.e., bone marrow transplantation) (Mosaad, 2014). MSCs, which can be transplanted as an allogenic cell source to another patient without immunosuppressive drugs, are capable of differentiation into multiple lineages

1291

that may produce tissues, including bone, cartilage, and muscle, and have been approved for use in multiple systems to treat bone defects (Su et  al., 2018). MSCs can also modulate the host immune response through paracrine or endocrine mechanisms, and are currently being applied in clinical trials for the treatment of immune diseases (Leyendecker Jr. et  al., 2018). Nevertheless, adult stem cells are rare, challenging to isolate and expand without altering cell phenotype, and limited in their differentiation potential. • Induced pluripotent stem cells: Successful reprogramming of differentiated human somatic cells into pluripotent cells has led to the creation of iPSCs (Takahashi et  al., 2006). These cells are functionally similar to ESCs, but do not require the destruction of an embryo and can be created from a patient’s own cells, eliminating the risk of host rejection (Yamanaka, 2008). Since their groundbreaking discovery in 2006, there have been significant advances in our ability to control the induction of pluripotency from multiple tissue sources such as skin fibroblasts, peripheral blood, hair-derived keratinocytes, and urine-derived epithelial cells. Studies have also focused on their targeted differentiation, as well as applications in tissue engineering and therapeutics (Mora et al., 2017). iPSCs hold significant promise as a universal, patient-derived cell source for the generation of multiple cell types in a single application. However, iPSC differentiation and subsequent expansion of targeted cell populations is technically challenging and often limited to low passage numbers, making it difficult to obtain clinically relevant cell numbers. In the case of uncontrolled downstream differentiation, it is further necessary to isolate cell populations for desired purity. Thus the application of patient-derived iPSCs for on-site therapeutics requires facilities, equipment, and expertise that may not be readily available. Additionally, challenges such as maintaining genomic stability, role of epigenetic factors, downstream mutations, potential immunogenicity, and regulatory hurdles need to be overcome for iPSCs to be applied clinically (Menon et al., 2016). • Mature cells: These cells, either parenchymal or stromal, represent the large population of fully differentiated, functionally, as well as tissue-specific cell types in the body such as fibroblasts, smooth muscle cells, epithelial cells, endothelial cells, hepatocytes, etc. They are routinely used in research, either as immortalized cell lines or as primary cells isolated from native tissue to help improve our understanding of cell biology as well as investigate drug screening and toxicity (Jackson and Lu, 2016). Within a clinical setting, these cells can be classified as autologous (patient derived), allogeneic (donor derived), or xenogeneic (nonhuman derived). Examples include the use of chondrocytes for cartilage repair, keratinocytes and fibroblasts for skin repair and cardiac cells for heart patches. Although autologous cells are less likely to cause an immune response, their

1292 SEC T I O N 2. 6    Applications of Biomaterials in Functional Tissue Engineering

availability depends on the tissue type, injury, or patient in question. In contrast, allogeneic and xenogeneic cells are more widely available but often suffer from immune rejection. Another potential complication is the possibility of the undesired dedifferentiation of cells during in  vitro expansion, which should be avoided prior to any in  vivo or clinical application. Although advanced reprogramming strategies and adult stem cells are a more efficient cell source, mature cell types are easier to use for translation because of their ease of availability and application. • Direct reprogrammed cells: The process of direct reprogramming of cells, also known as transdifferentiation, circumvents the iPSC induction of the seed (original) cells and differentiates them directly into the target cell type. Reprogramming can occur via many ways such as transfection via small molecules, small RNAs, or proteins. Successful reprogramming of adult cells such as fibroblasts into neuronal cells cardiac progenitors, hepatocytes, and more has already been demonstrated (Qin et al., 2017). It can be argued that based on the method of reprogramming, this process is safer than that for iPSCs in terms of mitigating downstream epigenetic or mutagenic factors. Particularly during the miRNA-mediated direct conversion of human fibroblasts into neurons, nonneuronal cells were induced to be postmitotic, thus limiting the proliferation of cells that have not undergone full reprogramming (Lu and Yoo, 2018). In tissue engineering, direct reprogramming can be utilized as an alternative site-specific or patient-specific cell source with potential regenerative and clinical applications such as cardiac, renal, and neural tissue engineering. Extrinsic factors such as the 3D microenvironment, the substrate biomaterial, and scaffold fabrication method have also been shown to play a role in guiding cellular reprogramming (Lee et al., 2016). Although still in their early stages, these approaches are extremely promising for the availability and expansion of an ideal cell source for tissue-engineering-based clinical applications. Further investigation is needed to identify the best source of cells for each tissue-engineering application. Multiple variables should be addressed in the selection of the best stem cell population, including: (1) stem cell accessibility (e.g., the isolation of autologous or allogenic neural stem cells is invasive and relatively difficult compared to other stem cells); (2) number of cells needed, where undesired differentiation during expansion has to be prevented; (3) proliferation capacity, where extra resources and expertise are needed for cells with limited proliferation; (4) differentiation profile; (5) cell population purity, depending on whether the cell source is a specific tissue or differentiated from pluripotent/stem cells, population homogeneity, and purity; (6) absence of random mutations that can potentially cause uncontrolled proliferation and tumors; and (7) ethical issues. In addition to primary cells, the intrinsic biological potential and performance of a cell can be modified by transient or permanent alteration of specific genes, often

accomplished with vectors created by modifying naturally occurring viruses such as retrovirus, lentivirus, adenovirus, or adeno-associated virus (Hannallah et al., 2003). There are several concerns for these approaches, such as transformation efficiency, safety of viral transfection, vector stability, and optimal function of the inserted genes. Nonviral transfection techniques have been developed to circumvent some of these issues; however, the long-term fate of these genetically modified cells still presents a potential risk. The cell phenotype can also be regulated through manipulation of isolation and culture conditions. Although this may expand the available tools to manipulate cell characteristics, it can create more hurdles for consistently manufacturing high-quality cells. Rigorous characterization must be completed to ensure maintenance of cell phenotype, purity, and differentiation state. In addition, it is important to consider that cell culture may “activate” cells, altering their phenotype from those found in situ. Factors such as the ideal number of cells to be transplanted, the maximum number of times cells can be passaged, the maximum length of time cells should be maintained in culture, the ideal differentiated state of the cell (determined via phenotypic or genotypic expression) to produce a therapeutic effect, and the ideal cell storage conditions can significantly alter the outcome of a regenerative approach, and often need to be optimized on a case-by-case basis for specific animal models. 

Materials Biomaterials are used to develop scaffolds that provide a template for cells to organize and restore structure and function to damaged or dysfunctional tissues. The bioactive matrix provided by these materials presents welldefined biochemical (e.g., growth factors and surface chemistry) and biophysical (e.g., fibrous structure, hydrophilicity, and stiffness) cues to effectively regulate cellular behaviors such as attachment, migration, proliferation, and differentiation for restoring the functionality of damaged tissues. Biomaterials can, at the same time, be used to supply nutrients, drugs, and bioactive factors that direct specific tissue growth. Accordingly, the material should be nontoxic and be fully biocompatible upon degradation. Ideally, the material should also possess physicomechanical and engineering properties suitable for the intended application and be compatible for further functionalization with bioactive molecules. Biomaterial scaffolds for tissue engineering are also discussed in more detail in Chapter 2.6.3. To accommodate these material requirements, the field of tissue engineering has witnessed tremendous development of new biomaterials over the past few decades. These materials are derived from both natural and synthetic sources and possess a broad spectrum of structural and functional properties that make them suitable for many clinical applications. Natural materials: A wide range of natural-origin polymers, generally including proteins and polysaccharides, are

CHAPTER 2.6.2   Overview of Tissue Engineering Concepts and Applications

used as carriers for cells and bioactive molecules (Seyednejad et  al., 2011). Natural materials are advantageous due to their inherent biological recognition through receptor– ligand interactions, cell-mediated proteolysis and remodeling, and low toxicity. Protein-based natural polymers include collagen, gelatin, silk fibroin, fibrin, and elastin. Collagen is the most prevalent protein in the human body and offers both physical support and essential signals for cell anchorage, migration, proliferation, differentiation, and survival. As a result, collagen has been studied for engineering artificial skin (collagen IV), bone (collagen I), and cartilage (collagen II), resulting in several tissue-engineering products. For example, bilayered collagen gels seeded with human fibroblast and keratinocytes are used as a bioengineered artificial skin by Organogenesis, Inc. under the name of Apligraf. Collagraft (Angiotech Pharmaceuticals), a composite of fibrillar collagen, hydroxyapatite, and tricalcium phosphate, has also been approved by the Food and Drug Administration (FDA) as a biodegradable synthetic bone graft substitute. Gelatin is a natural polymer derived from collagen that has significantly lower antigenicity in contrast to collagen and has been used for engineering bone, cartilage, and skin. Gelatin is also well suited to serve as a carrier for bioactive molecules, including growth factors, cells, and drugs, and displays a useful gelation mechanism at low temperatures that is commonly utilized in extrusion-based 3D printing practices. Silk fibroin has received significant attention as a versatile natural polymer due to its high strength-to-weight ratio and slow degradation. Its semicrystalline structure and self-healing modifications (via cross-linking) make it an attractive material to include as part of a composite scaffold for tissue engineering. Another commonly used natural material is fibrin, the structural component of blood clots that provides a transitory matrix for cell migration during wound healing. Fibrin has been used as a matrix for studying the regeneration of numerous tissues, such as adipose, bone, cardiac, cartilage, muscle, nervous, ocular, respiratory, skin, tendons and ligaments, and vascular tissue, as well as a carrier vehicle for bioactive molecules. Elastin is an insoluble, high-elastic polymer that is a major component in vascular and lung tissue. It is a promising material for synthetic vascular grafts: however, its insolubility and its ability to elicit an immune response have limited its use. Another class of natural polymers is polysaccharides that contain monosaccharide units joined together by glycosidic linkages (Celikkin et  al., 2017). Several polysaccharides, including chitosan, hyaluronic acid (HA), and alginate, have been studied for tissue-engineering applications. Chitosan, the fully or partially deacetylated form of chitin found particularly in the shell of crustaceans, has attracted significant attention, due to its biocompatible properties, for regeneration of ligaments, muscle, skin, bone, cartilage, nerves, and for vascular grafts. Alginate, a polysaccharide derived from the cell wall of brown seaweed, is a widely utilized biomaterial due to its biocompatibility, mild and physical gelation process, chemical and physical cross-linking

1293

abilities, nonthrombogenic nature, and the resemblance of its hydrogel matrix texture to that of the native extracellular matrix (ECM). Alginate happens to be easily modified into any form, such as microspheres, sponges, foams, elastomers, fibers, and hydrogels, thereby broadening the scope of application of this biomaterial. HA is a linear anionic polysaccharide, a member of the glycosaminoglycan family, and plays an important structural role in articular cartilage and skin tissues. Moreover, HA has been shown to promote both epithelial and mesenchymal cell migration and differentiation, regulate injury-induced inflammation as a free radical scavenger, and stimulate angiogenesis making it vital for tissue repair. Other commonly employed polysaccharides include starch, chrondoitin sulfate, cellulose, and dextran. Unfortunately, there are several limitations of natural materials, which include purification, cost, immunogenic responses, and lack of control over mechanical properties. There also exists the potential for a natural polymer to carry microbes or viruses. Although some of these disadvantages have been avoided through recombinant protein expression (Werkmeister and Ramshaw, 2012), synthetic biomaterials present a paradigm shift that overcomes many of these challenges while enabling custom-designed biomaterials. Natural materials are discussed in further detail in Chapter 1.3.6. Decellularized extracellular matrix: Decellularizing tissues, a process that eliminates all cellular and nuclear materials with various detergents, and using the remaining matrix as building blocks for therapeutic purposes, has provided a facilitated approach that contains critical physical and chemical properties for site-specific tissue regeneration. The use of decellularized ECM from donor tissue has been utilized in the repair of skin bladder, heart valve, and small intestinal submucosa. In addition, several commercialized decellularized scaffolds have received FDA approval for use in humans, including dermis tissue (Alloderm; LifeCell), porcine heart valves (Synergraft; Cryolife), porcine urinary bladder (urinary bladder matrix; ACell), and processed nerve allografts (Axogen; Acell). Decellularization of organs such as kidney, heart, lung, and liver has been documented and further seeding of functional parenchymal or stem cell populations onto these acellular three-dimensional (3D) biologic scaffolds provides the opportunity for direct connection to the patient’s vasculature. Decellularized tissues have also been used as lyophilized powders that are incorporated into natural or synthetic hydrogels (Kuo et al., 2018a). This allows chemical cues to be present to influence the behavior of coencapsulated cells, and gives customizable shapes for the application of the scaffold. The process of decellularization is preferred in the use of donor grafts for regenerative purposes since allogeneic and xenogeneic antigens are usually recognized as foreign by the host and can either cause destructive inflammatory response or overt immune-mediated rejection. Though the type of tissue governs its decellularization process, a general protocol contains physical (mechanical agitation, freeze/thaw, sonication), enzymatic (trypsin, endonucleases, exonucleases), and chemical (alkaline/acid, hypotonic and hypertonic solutions,

1294 SEC T I O N 2. 6    Applications of Biomaterials in Functional Tissue Engineering

EDTA/EGAT, Triton X-100, sodium dodecyl sulfate, deoxycholate, tri(n-butyl) phosphate) degradation steps (Brown et al., 2017). Since it is exceedingly difficult to rid of all cells in dense tissues, nearly all decellularization protocols result in residual DNA and other cytoplasmic and nuclear material. Additional applications of decellularized ECM are discussed in further detail in Chapter 1.3.6A. Synthetic materials: Many synthetic polymers have been designed and fabricated for tissue-engineering purposes. Biodegradable synthetic polymers offer several advantages compared to natural materials such as controlled mechanical properties and degradation kinetics, easy processability into custom shapes and structures, and easy modification of the material for specific applications. Saturated aliphatic polyesters, such as polylactide (PLA), polyglycolide (PGA), and poly(caprolactone) (PCL), are the most commonly used biodegradable synthetic polymers for 3D scaffolds in tissue engineering (Seyednejad et al., 2011). The chemical properties of these polymers allow hydrolytic degradation, and the rate and extent of degradation depends on the polymer’s molecular weight, structure, and composition. Polyesters have been widely used for scaffold construction because of their merits such as ease in synthesis, controllable degradable properties, and minimal inflammatory response. They have long been used in the clinic in the form of degradable sutures, but have more recently been used to mimic the aligned structures of native fibrous ECM in tissues (e.g., nerve, heart, skin, and tendon) with fibers ranging from several nanometers to several micrometers in diameter via electrospinning and 3D printing fabrication processes. A copolymer formulation of these two materials—poly(l-lactide-co-caprolactone)—has also been documented to improve the quality of regenerated cartilage tissue during microfracture when 3D printed into a scaffold that directs chondrocyte attachment and alignment (Guo et  al., 2018a). Poly(lactide-co-glycolide) (PLGA) is another widely investigated copolymer formulation of PLA and PGA, and its properties can easily be tailored to its application. Specifically, with 25%–75% lactide composition, PLGA forms amorphous polymers, which are very hydrolytically unstable when compared to PCL. This allows for various degradation times depending on the ratio of lactidyl:glycolidyl used (50:50 can be degraded as quickly as 1–2 months, while 85:15 can sustain 5–6 months in vivo) (Ulery et al., 2011). Poly(ortho esters) are hydrophobic, surface-eroding polymers, and their capacity to be used as tissue-engineering scaffolds is limited due to their weak mechanical properties and their capacity to induce a mild to moderate inflammatory response (Tschan et  al., 2017). Alternatively, poly(propylene fumarate) (PPF) is a high-strength, elastic polymeric biomaterial that can be cross-linked through unsaturated bonds in its backbone. The polymer degradation is therefore dependent not only on its molecular weight and cross-linker, but also on the material’s cross-linking density. For osteogenic tissue engineering, PPF can be mixed with ceramics such as hydroxyapatite or natural biomacromolecules like fibrin to create stronger, more bioactive scaffolds (Trachtenberg et al., 2017). Polyhydroxyalkanoates are semicrystalline isotactic polymers that

also undergo surface erosion and are excellent candidates for use in long-term tissue-engineering applications (Ulery et al., 2011). Elastomeric polyesters such as poly(glycerol sebacate) and poly(diol citrate) have shown significant promise as tissueengineering biomaterials that mimic the mechanical properties of soft tissues. Poly(anhydrides) are synthetic polymers developed from the condensation of diacids or a mixture of diacids. These polymers are biocompatible and have well-defined degradation characteristics. They were initially designed for drug-delivery applications due to their hydrophobicity and surface erosion that allowed a constant release profile for certain drugs. Several poly(anhydride) homopolymers and copolymers have been synthesized to control their degradation rate and mechanical properties. Polyurethanes are also favorable in tissue engineering because of their modifiable mechanical properties and biodegradability due to their segmented block structural character. These elastomers have also been developed as synthetic biomaterials for the regeneration of soft tissues, vasculature, and cartilage. As the foregoing materials are used for the creation of dry, fibrous scaffolds, synthetic polymers can also be dissolved and chemically cross-linked in solution to form 3D hydrogel networks. Photopolymerizable groups are needed for the creation of these gels so that the matrix remains intact once submerged in in vitro culture. PEG, poly(N-isopropylacrylamide), poly(vinyl alcohol), and poly(acrylates) are examples of these materials, and upon exposure to ultraviolet radiation or cross-linking agents, these polymers can cross-link to form polymeric networks. 

Biological Factors Biomolecules are the next essential component of a tissueengineered construct and an important part of the strategy to guide and regulate cell response, both in vitro and in vivo. They are the mediators of molecular signaling mechanisms and crosstalk between cells and their immediate microenvironment, with one biomolecule often serving multiple functionalities. Type of biological factors: Biological factors satisfy a broad category, including hormones, cytokines, growth factors, ECM molecules, cell surface molecules, and nucleic acids. The temporal and spatial coordination of cellular processes is orchestrated by these signals from the extracellular environment (Lee, 2000). A large number of biomolecules have been explored to induce tissue regeneration and can be broadly categorized as follows (1) small molecules (e.g., corticosteroids, hormones) are used for intercellular and intracellular signaling by binding to specific protein receptors; (2) proteins and peptides act on the cells as mitogens, morphogens, growth factors, and cytokines where they bind to a target cell receptor, triggering an intracellular signal transduction and a biological response; and (3) oligonucleotides (DNA or RNA) can affect either gene transcription and/or translation or can be incorporated into the cell’s genome. Cell-secreted biomolecules regulate cell–cell communication through

CHAPTER 2.6.2   Overview of Tissue Engineering Concepts and Applications

different mechanisms (endocrine, autocrine, paracrine, juxtacrine, intracrine) and have significant implications on the cell’s response to its external environment. Delivery and presentation of biological factors: The challenges and strategies to deliver these factors vary due to differences in their chemistry, biological response, as well as the scaffold biomaterial in question (Magin et al., 2016). Factors can be presented as soluble cues in vitro, where diffusive properties of the scaffold may be leveraged to further regulate cell response through autocrine–paracrine mechanisms (Mahadik et  al., 2017), although controlling their release kinetics can be challenging. Several strategies exist for the encapsulation and release of biological factors, often via degradation of the scaffold or diffusive release of the agent. For example, researchers have demonstrated the delivery of vascular endothelial growth factor and platelet-derived growth factor for improved angiogenesis, and of dextrin-conjugated growth factors for neural stem cell culture. Additionally, controlling the spatiotemporal presentation of biomolecules via advanced fabrication methods (such as 3D printing) is particularly advantageous for complex tissues, e.g., multiple drug release from a single scaffold as an advanced drug delivery mechanism (Liu et al., 2016). Biomolecules can also remain active when membrane bound or tethered to surfaces via immobilization techniques. Noncovalent association of the matrix components, e.g., glycosaminoglycans, can slowly release and potentiate binding to the cell membrane receptors (Sakiyama-Elbert, 2014). Alternatively, covalently bound growth factors can also influence cell behavior, such as the immobilized stem cell factor on hematopoietic stem cells (Mahadik et  al., 2015), or the osteochondral differentiation of MSCs (Di Luca et al., 2017). A class of conjugation method known as “click” chemistry is widely used in tissue engineering for the functionalization of biomolecules (McKay and Finn, 2014). A subset of photocross-linkable thiol-ene chemistry is used extensively due to the high degree of spatial control and multiple conjugation offered by this technique. This is regularly used for biomolecule patterning within hydrogels, as well as for photo-controlled degradation for biomolecule release. In addition, cell adhesion on the scaffold can be controlled by the presentation of specific peptides and carbohydrates. In particular, a prototypical three amino acid sequence arginine–glycine–aspartic acid (RGD) is frequently found in many adhesion proteins and binds to many integrin receptors on cells. RGD peptide sequences have therefore been covalently immobilized on a synthetic material surface at a defined density and orientation to guide cell adhesion. Other sequences have also been successfully explored in the literature (Gao et al., 2017). Finally, the covalent immobilization of protease-sensitive cleavable linkages (i.e., matrix metalloproteinase [MMP]-sensitive links) on synthetic matrices mimics the native ECM, which can be used, in addition to soluble growth factors and other chemotactic cues, to guide migration of cells (Stevens et al., 2015). Mechanochemical factors in tissue growth: The physical and/or chemical nature of the scaffold is an equally

1295

important factor in regulating cell fate. Integrin-mediated cell adhesion to the ECM as well as the corresponding mechanotransduction signals facilitate cell response to their immediate microenvironment (Sun et  al., 2016). Modifying the material surface chemistry can control many aspects of cell response such as adhesion, migration, and differentiation. Moreover, nanoscale geometry, surface topography, matrix size, and fiber alignment can also influence cell adhesion, proliferation, and migration. Altering surface topography leads to cytoskeletal reorganization, which directly influences molecular and biomechanical signals (Wu et al., 2017). Furthermore, mechanical forces exerted by the scaffold matrix and the elasticity of the material also influence cell fate. For instance, matrix stiffness in concert with surface chemistry regulates the differentiation of stem cells (Engler et al., 2006; Choi and Harley, 2012). This response is further enhanced via mechanical loading and cyclic or constant tension exposed to the scaffold, and consequently the cells, and is a strategy commonly used for engineering scaffolds for native tissues that are under constant stress (e.g., bone, cartilage, muscle, tendon etc.). In conclusion, the physical environment, consisting of geometry, time-varying stress, strain, fluid flow, pressure, and potentially other biophysical parameters (e.g., osmotic pressure and electrical field), can regulate cell phenotype and tissue structure in a 3D environment. 

Scaffold Design Scaffolds act as the synthetic analog of the natural ECM. The role of scaffolds is to recapitulate the normal tissue development process by allowing cells to formulate their own microenvironment. The scaffold provides the necessary support for cells to attach, proliferate, and maintain their differentiated function and subsequent regeneration of new tissues. Ideally, a scaffold should have the following characteristics: (1) 3D highly porous structure with an interconnected pore network to facilitate cell/tissue growth and diffusion of nutrients, metabolic waste, and paracrine factors; (2) biodegradable or bioresorbable features with controllable degradation and resorption rates to match cell/ tissue growth in vitro and in vivo; (3) suitable surface chemistry for cell attachment, proliferation, and differentiation; (4) mechanical properties to match those of the tissues at the site of implantation; and (5) easy processability to form a variety of shapes and sizes (O’Brien, 2011). Conventionally fabricated scaffolds: The formation of a porous structure is the main goal of scaffold fabrication. Most methods for fabricating porous scaffolds, including particulate leaching, freeze drying, gas infusion, and phase separation, rely on casting the scaffold within molds with specific designs to create isotopically distributed interconnected pores. Porous structures can be developed by introducing particles or bubbles when the scaffold is solidified, which are later removed leaving behind an interconnected network of pores. Although these techniques are relatively simple for developing a 3D structure, they are limited by

1296 SEC T I O N 2. 6    Applications of Biomaterials in Functional Tissue Engineering

uncontrolled pore size and connectivity, poor mechanical strength, and residual solvent/porogens. Hydrocarbon templating is a process where a polymer is dissolved in a solvent, mixed with a hydrocarbon porogen, compacted, immersed in a solvent to precipitate the porogen polymer solvent, and vacuum dried to create a porous foam (Shastri et  al., 2000). This process has several advantages, including control over scaffold thickness and pore structure, although use of organic solvents is typically a disadvantage for biological applications as it is difficult to remove completely. Hydrogel scaffolds: Hydrogels are commonly used as tissueengineering scaffolds because of their ability to be formed into specific shapes based on their applications. They are typically cross-linked networks swollen in water that have the capability to encapsulate cells in a 3D network. The specific hydrogel material determines the particular crosslinking capabilities. Some hydrogels are cross-linked using light, while others may be cross-linked through heat, ionic bonding, or covalent bonding. Photocross-linkable hydrogels are polymerized into 3D networks in the presence of cross-linkers under ultraviolet or visible light radiation and are frequently used for cell encapsulation (Ferreira et  al., 2007). By controlling the structure with defined crosslinking density, mechanical properties, mass transport, and degradation characteristics, the gels can be tuned for a range of applications. Hydrogels are an appealing 3D scaffold because they are structurally similar to the ECM of many tissues, can often be processed under relatively mild conditions, and may be delivered in a minimally invasive manner (Drury and Mooney, 2003). They can either be durable or biodegradable and are advantageous due to their high water content, facile transport properties, and controlled degradation kinetics. Furthermore, hydrogels can be chemically modified to improve the adhesion and proliferation of the cells on the gel matrices through inclusion of adhesion peptides. However, major problems with using hydrogels include poor mechanical properties that lead to their use in only soft and nonload-bearing tissues and difficulty in engineering complex tissues with multiple cell types due to unique ECM requirements. Fiber-based scaffolds: Fibrous scaffold structures may be developed by electrospinning of polymers to generate continuous micro- or nanoscale diameter fibers. Electrospinning has the ability to mimic the ECM due to its ability to create nanofibers on the scale of naturally occurring collagen fibrils such as those in tendons and ligaments (Sensini and Cristofolini, 2018). Additionally, the orientation of fibers can be controlled during electrospinning to develop random or aligned fibers. Electrospun scaffolds are advantageous due to their high surface-to-volume ratio and structural similarity to natural ECM. Although it is possible to control the orientation of the fibers, it can be difficult to control the distance between fibers, an important factor that influences the migration of cells. Self-assembly of biopolymers, e.g., peptides and nucleic acids, using noncovalent interactions, including H-bonding, hydrophobic, electrostatic interactions, and van der Waals forces, has also

been used for scaffold development. The main advantage of such biopolymers is that their self-assembly relies on specific biorecognitions (e.g., DNA hybridization), which therefore make the formation of scaffold highly predictable and programmable. This type of fabrication has been used to create biocompatible fibers on the micro- or nanoscale and has the ability to enhance cell attachment (DeFrates et al., 2018). However, due to its self-assembly capabilities, the composition and size of these fibers can be difficult to regulate. Additive manufacturing: 3D printing is a modern fabrication method where digital data of a 3D structure is converted into a physical object. In contrast to traditional fabrication methods (involving molds or solvents), 3D printing can create complex structures in a layer-by-layer fashion. Generally, a structure is first visualized digitally with the help of computer-aided software (such as SolidWorks and CreoParametric) and later sliced into individual layers by the specific printer software. Three major approaches for 3D bioprinting currently exist: inkjet printing, extrusion printing, and stereolithography. The chosen fabrication method will vary in terms of materials used, time required, and printing parameters used to develop the desired part. The following is a brief description of each major printing method, as well as its positives and pitfalls: • Inkjet printing: Much like how a traditional printer utilizes ink cartridges to deliver droplets to paper to create documents, inkjet bioprinters can deliver biomaterials and/ or cells in controlled small volumes. These printers have an extremely high resolution in bioink deposition (up to 50 μm) and the highest printing speeds (15–25 kHz) (Gudapati et  al., 2016). However, droplets are synthesized through either thermal or mechanical means, which can perturb cells during the printing process. Instant heat exposure and shear stress can also induce cell damage. In addition, ink for this printer must be fluid so that droplets can form, which limits cell density of the ink (Skardal and Atala, 2015). As such, inkjet printing is best for twodimensional (2D) tissues with relatively low cell density and complexity. • Extrusion printing: This printing process extrudes continuous filaments of shear-thinning, thixotropic material (of which either biomacromolecules or cells are seeded) through a nozzle under pneumatic pressure. Though extrusion printing is an efficient and low-cost method for printing cell-encapsulated constructs, the printing resolution is limited by the print head’s diameter, the viscosity of the biomaterial, and size of the incorporated biologic (Ozbolat and Hospodiuk, 2016). Shear-thinning biomaterials, in regard to strain rate or temperature-dependent properties, are limited as this fabrication process requires the material to maintain strand shape. Examples of inks compatible with this deposition method include gelatin, fibrin, alginate, HA mixtures, and silicate nanoplatelet mixtures (Panwar and Tan, 2016). The combination of multiple materials via different printing cartridges allows for facile engineering of chemically and structurally complex tissues.

CHAPTER 2.6.2   Overview of Tissue Engineering Concepts and Applications

• Stereolithography: A photopolymerizable material is printed layer by layer in this printing process. The light source, located directly below the printing platform, is instructed to focus at specific points in space to project a 2D mask with a specific light intensity that locally polymerizes the biomaterial. As the print progresses, the build platform moves up, and new masks are projected and cross-linked onto the existing shape. The overall print is complete when all mask layers have been projected and the material has been cross-linked into the desired 3D shape. Digital micromirror device-based printing is a similar platform in which an array of digitally controlled mirrors reflects light to create a 2D projection on a surface. Advantages of this process include the high precision able to be obtained by the light source, and that a 3D object of any desired shape can be made by changing the 2D projection over the height of the printing stage (Wang et al., 2017). One major disadvantage of this methodology is that only one material can be printed at a time. 

Integration of Multiple Factors The environment in which a particular cell resides in vivo, also known as its “niche” microenvironment, plays a significant role in determining cell functionality. This environment is composed of other cell types, the ECM proteins, and a host of biomolecules, with a high degree of crosstalk between all these components. In vitro, successful integration of the appropriate cells, scaffolds, soluble cues, and mechanochemical factors is key in regulating cell fate and ultimately regenerating a functional tissue. Hence, combinatorial approaches that recapitulate key components of this niche microenvironment are extremely important in tissue engineering, as well as to advance our basic understanding of cell biology. A number of studies have utilized this approach for tissue niche engineering, such as bone marrow (Torisawa et al., 2014), lung (Petersen et al., 2010), cardiac (Jang et al., 2017), and more (Donnelly et al., 2018; Kobolak et  al., 2016), with promising results. However, static in vitro tissue-engineering systems often fail to consider the multiple factors that contribute to the healing and regeneration process. Thus it is necessary to understand the system as a whole through dynamic in vitro systems, such as bioreactors and relevant in vivo models. 

Models for Tissue Engineering Bioreactors The traditional tissue-engineering approach of growing tissues on 2D surfaces (i.e., Petri dishes) or in 3D scaffolds is limited by mass transfer, since the diffusion of metabolites, oxygen, and carbon dioxide under static conditions can only support a certain thickness of tissue. Depending on the biomaterial, this limits the maximum thickness of a tissue-engineered construct to