Kent State University Polymer Worksheet

1.Ethylene (CH2═CH2) and propylene (CH2═CHCH3) are each separately polymerized to yield polymer A and B respectively.
Which of the polymers would have a higher glass transition temperature? Which would more likely be crystalline? Explain
your answers. (only answer the crystalline portion)
Explain the difference between branched and cross-linked polymers. Which would be expected to be more stiff? Why?
What is the difference between isotactic, syndiotacytic, and atactic forms of polymers? Which are more likely to be
crystalline in nature?
You are given two samples of the same polymer. One is amorphous and the other is mostly crystalline. Which will be
more difficult to dissolve in a solvent? Which is more likely to be opaque?
The average molecular weight of a sample of PMMA (Figure 6.8) is 75 650. What is the degree of polymerization for the
Three polymer samples with the following properties are mixed together to get 1 kg of polymer mixture. What is overall
Mn ?
Poly(ε-caprolactone) is being considered as a potential material for a vascular graft. After a particular batch
has been made, someone gives you the following fractional distribution data and asks you to calculate M n ,
M w and PI for this polymer.
Mi (kg/mol)
You are examining a copolymer for its potential as a material for a vascular graft. You are trying to
determine whether you want a high or low degree of crystallinity in the material. What type of structures for
copolymers have a higher probability for crystallization?
How do polyesters such as PLA and PGA degrade? What are some properties that affect their rate of degradation?
Explain how autocatalysis can influence their degradation.
10. What is the difference between protein-based and polysaccharide-based materials?
11. In general, what is the molecular difference between collagen and elastin?
12. List the different criteria that natural materials need to satisfy in order for them to be used in biomedical applications.
13. Receptor binding motifs such as RGD have highly selective recognition for cell adhesion. Outline a strategy for
incorporating a cell adhesive recognition domain to a biomaterial. Address the problems surrounding domain quantity,
type, and affinity in your discussion. (Outline a strategy discussed in class to attach RGD protein to the surface
of a polystyrene material and address issues such as domain quantity type, and affinity of the protein for
your chosen surface modification)
14. Concisely describe the differences between focal contacts and hemidesmosomes. What would be the implications of
excessive formation of either at a biomaterial surface taking into consideration normal cell functions such as motility and
Introduction to Biomaterials
This textbook gives students the perfect introduction to the world of biomaterials,
linking the fundamental properties of metals, polymers, ceramics, and natural
biomaterials to the unique advantages and limitations surrounding their
biomedical applications.
! Clinical concerns such as sterilization, surface modification, cell–biomaterial
interactions, drug delivery systems, and tissue engineering are discussed in
detail, giving students practical insight into the real-world challenges associated with biomaterials engineering.
! Key definitions, equations and concepts are concisely summarized alongside
the text, allowing students to quickly and easily identify the most important
! Bringing together elements from across the book, the final chapter discusses
modern commercial implants, challenging students to consider future industrial possibilities.
Concise enough to be taught in a single semester, and requiring only a basic
understanding of biology, this balanced and accessible textbook is the ideal
introduction to biomaterials for students of engineering, materials science, and
C. Mauli Agrawal is the Vice President for Research at the University of Texas at
San Antonio (UTSA), and the Peter Flawn Professor of Biomedical Engineering.
Previously, he served as the Dean of the College of Engineering at UTSA. He
specializes in orthopedic and cardiovascular biomaterials and implants and his
inventions have been licensed to various companies. He is a member of the
International College of Fellows of Biomaterials Science and Engineering, a
Fellow of the American Institute for Medical and Biological Engineering, a former
President of the Society for Biomaterials, and was awarded the 2010 Julio Palmaz
Award for Innovation in Healthcare and the Biosciences.
Joo L. Ong is Chair of the Department of Biomedical Engineering and the USAA
Foundation Distinguished Professor at the University of Texas at San Antonio.
His research focuses on modification and characterization of biomaterials surfaces
for dental and orthopedic applications, tissue engineering ceramic scaffolds,
protein-biomaterial interactions, and bone–biomaterial interactions. He is a Fellow
of the American Institute for Medical and Biological Engineering.
Mark R. Appleford is an Assistant Professor of Biomedical Engineering at the
University of Texas at San Antonio, focusing on tissue–biomaterial interactions,
cellular engineering, reconstructive tissue engineering, and biocompatibility.
Gopinath Mani is an Assistant Professor of Biomedical Engineering at the
University of South Dakota, focusing on surface modification and characterization
of biomaterials, nanomaterials and nanomedicine, biodegradable metals and drug
delivery systems. He is the Program Chair for the Surface Characterization and
Modification Special Interest Group of the Society for Biomaterials, and has
developed and taught numerous graduate-level programs in biomaterials
Series Editors
W. Mark Saltzman, Yale University
Shu Chien, University of California, San Diego
Series Advisors
Jerry Collins, Alabama A & M University
Robert Malkin, Duke University
Kathy Ferrara, University of California, Davis
Nicholas Peppas, University of Texas, Austin
Roger Kamm, Massachusetts Institute of Technology
Masaaki Sato, Tohoku University, Japan
Christine Schmidt, University of Florida, Gainesville
George Truskey, Duke University
Douglas Lauffenburger, Massachusetts Institute of Technology
Cambridge Texts in Biomedical Engineering provide a forum for high-quality
textbooks targeted at undergraduate and graduate courses in biomedical engineering.
It covers a broad range of biomedical engineering topics from introductory texts to
advanced topics, including biomechanics, physiology, biomedical instrumentation,
imaging, signals and systems, cell engineering, and bioinformatics, as well as other
relevant subjects, with a blending of theory and practice. While aiming primarily at
biomedical engineering students, this series is also suitable for courses in broader
disciplines in engineering, the life sciences and medicine.
“This is a book that is destined to be a classic in biomaterials education. Written by
leading bioengineers and scientists, it can serve not only as a textbook to support a
semester-long undergraduate course, but also as an introduction to graduate-level
classes. It is a well-written, comprehensive compendium of traditional and also modern
knowledge on all aspects of biomaterials, and I am sure that both students and instructors will embrace it and use it widely.”
Kyriacos A. Athanasiou
University of California, Davis
Introduction to
Basic Theory with Engineering Applications
C. Mauli Agrawal
University of Texas at San Antonio
Joo L. Ong
University of Texas at San Antonio
Mark R. Appleford
University of Texas at San Antonio
Gopinath Mani
University of South Dakota
University Printing House, Cambridge CB2 8BS, United Kingdom
Published in the United States of America by Cambridge University Press, New York
Cambridge University Press is part of the University of Cambridge.
It furthers the University’s mission by disseminating knowledge in the pursuit of
education, learning and research at the highest international levels of excellence.
Information on this title:
© 2014 C. M. Agrawal, J. L. Ong, M. R. Appleford and G. Mani
This publication is in copyright. Subject to statutory exception
and to the provisions of relevant collective licensing agreements,
no reproduction of any part may take place without the written
permission of Cambridge University Press.
First published 2014
Printed and bound in the United Kingdom by TJ International Ltd. Padstow Cornwall
A catalog record for this publication is available from the British Library
Library of Congress Cataloging in Publication data
Introduction to biomaterials : basic theory with engineering applications /
C. M. Agrawal . . . [et al.].
p. cm. – (Cambridge texts in biomedical engineering)
Includes bibliographical references.
ISBN 978-0-521-11690-9 (Hardback)
I. Agrawal, C. Mauli (Chandra Mauli) II. Series: Cambridge texts in biomedical engineering.
[DNLM: 1. Biocompatible Materials. 2. Biomedical Technology. QT 37]
610.280 4–dc23 2013016363
ISBN 978-0-521-11690-9 Hardback
Additional resources for this publication at
Cambridge University Press has no responsibility for the persistence or accuracy of
URLs for external or third-party internet websites referred to in this publication,
and does not guarantee that any content on such websites is, or will remain,
accurate or appropriate.
I dedicate this work to my parents who taught me to love excellence, and to my wife and
children (Sue, Ethan and Serena), who have always supported my pursuit of it.
C. Mauli Agrawal
To my family, who have put up with me all these years.
Joo L. Ong
I express my deepest appreciation for my wife Lindsey, best friend, greatest love,
Mark R. Appleford
supplier of green limes and good joss.
I dedicate this work to my wife Priya Devendran, my daughter Manushri Gopinath,
Gopinath Mani
and my parents Mani and Bagyam Mani.
page xvii
1 Introduction
1.1 Definitions
1.2 Changing focus
1.3 Types of bonds in materials
1.3.1 Ionic bonds
1.3.2 Metallic bonds
1.3.3 Covalent bonds
1.3.4 Secondary bonds
1.4 Types of materials
1.4.1 Ceramics
1.4.2 Metals
1.4.3 Polymers
1.4.4 Composites
1.5 Impact of biomaterials
1.6 Future of biomaterials
1.7 Summary
2 Basic properties of materials
2.1 Mechanical properties
2.1.1 Tensile testing
2.1.2 Compressive testing
2.1.3 Shear testing
2.1.4 Bend or flexural tests
2.1.5 Viscoelastic behavior
2.1.6 Ductile and brittle fracture
2.1.7 Stress concentration
2.1.8 Fracture toughness
2.1.9 Fatigue
2.2 Electrochemical properties
2.2.1 Corrosion
2.2.2 Types of corrosion
2.3 Surface properties
2.3.1 Contact angle
2.3.2 Hardness
2.4 Summary
Suggested reading
3 Biological systems
3.1 The biological environment
3.2 Genetic regulation and control systems
3.3 The plasma membrane
3.3.1 Membranes are phospholipid layers
3.4 Cytoskeleton and motility
3.5 Cell to cell communication pathways
3.6 Cell junctions
3.6.1 Tight junctions
3.6.2 Gap junctions
3.6.3 Adherens and desmosomes
3.7 Cell signaling pathways
3.7.1 Receptors as signaling sensors
3.7.2 Receptor classes
3.7.3 Second messengers and their activation/
3.8 Biological testing techniques
3.8.1 Probe and labeling technologies
3.8.2 Examination of gene expression
3.8.3 The plasma membrane
3.8.4 Cytoskeleton and motility
3.8.5 Communication between cells
3.8.6 Mapping intracellular signaling
3.9 Summary
Suggested reading
4 Characterization of biomaterials
4.1 Contact angle
4.2 Infrared spectroscopy
4.2.1 Attenuated total reflection (ATR)
4.2.2 Specular reflectance
4.2.3 Infrared reflection absorption spectroscopy (IRRAS)
4.2.4 Diffuse reflectance infrared Fourier transform
spectroscopy (DRIFTS)
4.3 X-ray photoelectron spectroscopy
4.4 Secondary ion mass spectrometry
4.5 Atomic force microscopy
4.6 Scanning electron microscopy
4.7 Transmission electron microscopy
4.8 X-ray diffraction (XRD)
4.9 Chromatography
4.9.1 High performance liquid chromatography (HPLC)
4.9.2 Gel permeation chromatography (GPC)
4.10 Summary
Suggested reading
5 Metals: structure and properties
5.1 Titanium and its alloys
5.1.1 Classification of Ti and its alloys based on
crystallographic forms
5.1.2 Surface properties
5.1.3 Applications
5.2 Stainless steel
5.2.1 Martensitic stainless steels
5.2.2 Ferritic stainless steels
5.2.3 Austenitic stainless steels
5.2.4 Duplex stainless steels
5.2.5 Recent developments in stainless steel alloys
5.3 Cobalt–chromium alloys
5.3.1 ASTM F75
5.3.2 ASTM F799
5.3.3 ASTM F90
5.3.4 ASTM F562
5.4 Nitinol
5.5 Tantalum
5.6 Magnesium
5.7 Summary
Suggested reading
6 Polymers
6.1 Molecular structure of polymers
6.1.1 Molecular weight
6.2 Types of polymerization
6.3 Physical states of polymers
6.3.1 Amorphous phase
6.3.2 Crystalline phase
6.4 Common polymeric biomaterials
6.4.1 Polyethylene
6.4.2 Polymethylmethacrylate (PMMA)
6.4.3 Polylactic acid (PLA) and polyglycolic acid (PGA)
6.4.4 Polycaprolactone (PCL)
6.4.5 Other biodegradable polymers
6.4.6 Polyurethanes
6.4.7 Silicones
6.5 Hydrogels
6.5.1 Synthesis of hydrogels
6.5.2 Properties of hydrogels
6.5.3 Applications
6.6 Nanopolymers
6.7 Summary
Suggested reading
7 Ceramics
7.1 General properties
7.2 Classifications
7.2.1 Classification based on form
7.2.2 Classification based on composition
7.2.3 Classification based on reactivity
7.3 Bioceramics
7.3.1 Silicate glass
7.3.2 Alumina (Al2O3)
7.3.3 Zirconia (ZrO2)
7.3.4 Carbon
7.3.5 Calcium phosphates (CaP)
7.3.6 Hydroxyapatite (HA)
7.3.7 Tricalcium phosphate (TCP)
7.3.8 Calcium sulfate (CaSO4·H2O)
7.3.9 Bioactive glass
7.4 Nanoceramics
7.5 Summary
Suggested reading
8 Natural biomaterials
8.1 Collagen
8.2 Elastin
8.3 Silk
8.4 Chitosan
8.5 Cellulose
8.6 Alginate
8.7 Hyaluronan
8.8 Chondroitin sulfate
8.9 Coral
8.10 Summary
Suggested reading
9 Surface modification
9.1 Abrasive blasting
9.2 Plasma glow discharge treatments
9.2.1 Direct current glow discharge
9.2.2 Alternating current glow discharge
9.2.3 Capacitively coupled radiofrequency glow discharge
9.2.4 Inductively coupled radiofrequency glow discharge
9.3 Thermal spraying
9.4 Physical vapor deposition (PVD)
9.4.1 Evaporative deposition
9.4.2 Pulsed laser deposition
9.4.3 Sputter deposition
9.5 Chemical vapor deposition (CVD)
9.6 Grafting
9.7 Self-assembled monolayer (SAM)
9.7.1 Patterning of self-assembled monolayers
9.8 Layer-by-layer (LbL) assembly
9.8.1 Different layer-by-layer (LbL) assembly techniques
9.9 Summary
Suggested reading
Sterilization of biomedical implants
10.1 Common terminology
10.2 Steam sterilization
10.3 Ethylene oxide sterilization
10.4 Gamma radiation sterilization
10.5 Other sterilization methods
10.5.1 Dry heat sterilization
10.5.2 Formaldehyde and glutaraldehyde treatments
10.5.3 Phenolic and hypochloride solution treatments
10.5.4 Ultraviolet (UV) radiation
10.5.5 Electron beam sterilization
10.6 Recently developed methods
10.6.1 Low temperature gas plasma treatment
10.6.2 Gaseous chlorine dioxide treatment
10.7 Summary
Suggested reading
Cell–biomaterial interactions
11.1 The extracellular environment
11.2 Extracellular matrix mimics
11.3 Cell interactions with non-cellular substrates
11.4 Biocompatibility testing and techniques
11.4.1 Immunostaining techniques for studying cell–ECM
11.4.2 Profiling a cell line for its ECM binding
11.4.3 Immunoprecipitation and Western blotting
11.5 Summary
Suggested reading
Drug delivery systems
12.1 Diffusion controlled drug delivery systems
12.1.1 Membrane controlled reservoir systems
12.1.2 Monolithic matrix systems
12.2 Water penetration controlled drug delivery systems
12.2.1 Osmotic pressure controlled drug delivery systems
12.2.2 Swelling controlled drug delivery system
12.3 Chemically controlled drug delivery systems
12.3.1 Polymer–drug dispersion systems
12.3.2 Polymer–drug conjugate systems
12.4 Responsive drug delivery systems
12.4.1 Temperature-responsive drug delivery systems
12.4.2 pH-responsive drug delivery systems
12.4.3 Solvent-responsive drug delivery systems
12.4.4 Ultrasound-responsive drug delivery systems
12.4.5 Electrically responsive drug delivery systems
12.4.6 Magnetic-sensitive drug delivery systems
12.5 Particulate systems
12.5.1 Polymeric microparticles
12.5.2 Polymeric micelles
12.5.3 Liposomes
12.6 Summary
Suggested reading
Tissue engineering
13.1 Tissue engineering approaches
13.1.1 Assessment of medical need
13.1.2 Selecting a tissue engineering strategy
13.2 Cells
13.2.1 Stem cells
13.2.2 Biopreservation of cells
13.3 Scaffold properties
13.4 Fabrication techniques for polymeric scaffolds
13.4.1 Solvent casting and particulate leaching
13.4.2 Electrospinning
13.4.3 Solid free form fabrication (SFFF)
13.5 Fabrication of natural polymer scaffolds
13.6 Fabrication techniques for ceramic scaffolds
13.6.1 Template sponge coating
13.6.2 Non-sintering techniques
13.7 Assessment of scaffold architecture
13.8 Cell seeded scaffolds
13.8.1 Cell culture bioreactors
13.8.2 Cell seeding
13.8.3 Growth factors
13.8.4 Mechanical modulation
13.9 Assessment of cell and tissue properties
13.9.1 Cellular properties
13.9.2 Tissue properties
13.10 Challenges in tissue engineering
13.11 Summary
Suggested reading
Clinical applications
14.1 Cardiovascular assist devices
14.2 Cardiovascular stents
14.3 Dental restoration
14.4 Dental implants
14.5 Neural prostheses
14.6 Opthalmology
14.7 Orthopedic implants
14.8 Renal
14.9 Skin applications
14.10 Summary
Additional reading
Biomaterials have helped millions of people achieve a better quality of life in
almost all corners of the world. Although the use of biomaterials has been
common over many millennia, it was not until the twentieth century that the field
of biomaterials finally gained recognition. With the advent of polymers, new
processing and machining processes for metals and ceramics, and general
advances in technology, there has been an exponential growth in biomaterialsrelated research and development activity over the past few decades. This activity
has led to a plethora of biomaterials-based medical devices, which are now
commercially available.
For students in the area of biomaterials, this is an especially exciting time. On
the one hand, they have the opportunity to meet and learn from some of the
stalwarts and pioneers of the field such as Sam Hulbert, one of the founders of the
Society for Biomaterials (SFB). Other greats include Allan Hoffman and Buddy
Ratner (biomaterials surfaces), Robert Langer (polymers and tissue engineering),
Nicholas Peppas (hydrogels), Jack Lemons (orthopedic/dental implants), Joseph
Salamone (contact lenses), and Julio Palmaz (intracoronary stents). Most of these
individuals are still active in research and teaching. The authors of this book have
been privileged to interact and learn from them in various forums, and students
today have the same opportunities. On the other hand, with the current availability
of sophisticated processing and characterization technologies, present day students
also have the tools to take the field to unprecedented new levels of innovation.
This book has been written as an introduction to biomaterials for college
students. It can be used either at the junior/senior levels of undergraduate education or at the graduate level for biomedical engineering students. It is best suited
for students who have already taken an introductory course in biology. We have
felt the need for a textbook that caters to all students interested in biomaterials and
does not assume that every student intends to become a biomaterials scientist. This
book is a balance between science and engineering, and presents both scientific
principles and engineering applications. It does not assume that the student has a
background in any particular field of study. Therefore, we first cover the basics of
materials in Chapters 1 and 2 followed by basic biological principles in Chapter 3.
After presenting various techniques for the characterization of biomaterials in
Chapter 4, we dedicate a chapter each to the discussion of metals, polymers,
ceramics, and natural biomaterials (Chapters 5–8). Surface modification methods
are presented in Chapter 9, followed by sterilization techniques in Chapter 10. The
success of any biomaterial depends on the biological response to it and so protein
chemistry, cell–biomaterial interactions, and the effect of biomaterials on tissue
response are addressed in Chapter 11. The last three chapters (Chapters 12–14)
cover the application of biomaterials in the clinical world; specifically drug delivery systems, tissue engineering, and clinical applications are presented and
This book has been designed to present enough material so that it can be
comfortably covered during a regular length semester-long course. It should
provide the student with a concise but comprehensive introduction to biomaterials
and lays the foundation for more advanced courses.
The authors would like to thank the following individuals for assisting in a
variety of ways in compiling this book: Jordan Kaufmann, Ethan Agrawal, Serena
Agrawal, Tim Luukkonen, Amita Shah, Steve Lin, Angee Ong, Kevin Ong, Lisa
Actis, Marcello Pilia, and Stefanie Shiels.
1 Introduction
After reading this chapter the student will understand the following.
! History of implants and biomaterials.
Various definitions for biomaterials.
Different types of chemical bonds.
Future directions for the progress in biomaterials.
Basic families of materials.
The Rig Veda, one of the four sacred Sanskrit books of ancient India that were
compiled between 3500 and 1800 BC, relates the story of a warrior queen named
Vishpla, who lost a leg in battle and was fitted with an iron leg after the wound
healed. There is also mention of lacerated limbs treated with sutures. Sushruta,
a renowned Indian physician from circa 600 BC, wrote a very comprehensive treatise
describing various ailments as well as surgical techniques. His technique for nose
reconstruction using a rotated skin flap is still used in modern times. Sutures
made of vegetable fibers, leather, tendons, and horse hair were commonly used in
his time. There are also reports of the use of linen sutures in Egypt 4000 years ago.1
These ancient records show that, since time immemorial, humans have tried
to restore the function of limbs or organs that have ceased to perform adequately
due to trauma or disease. Often, this was attempted through the use of materials
either made or shaped by humans and used external to the body. These were the
earliest form of biomaterials. Although examples of successful external prosthetic
devices can be found in history, materials placed inside the body, also known
as implants, were usually not viable due to infection. This changed in the 1860s
with the advent of aseptic surgical techniques introduced by Dr. J. Lister. The
discovery of antibiotics in the mid 1900s also reduced the incidence of infections
related to surgery. Today, implants are very successful and are used in a wide
variety of applications in the practice of medicine, improving the quality of life for
millions and saving countless lives. However, the successes of today have come
after a long history of trial and error and scientific endeavor.
Perhaps the most common implants in ancient times were dental in nature and thus,
these implants provide an interesting history of progress over the millennia. Human
remains from the first century AD recovered from a Gallo-Roman necropolis in
France show the use of an iron implant to replace a tooth. In 1923, an archeological
dig in Honduras uncovered pieces of shells that were used as a dental implant for a
young woman in AD 600. In the Middle East, ivory implants have been discovered
with skeletons from the Middle Ages. In more modern times, in the early 1800s,
gold posts were placed in sockets immediately after tooth extraction with limited
results. In the years following, other metals such as platinum were also investigated.
In the 1940s, the Strock brothers from Boston tested Vitallium dental implants. A major
breakthrough came in 1952 when Ingvar Branemark from Sweden tried titanium
implants and found them to attach well to bone. Such implants provide a post onto
which artificial teeth or crowns can be attached. These implants are in use even today.
Box 1.1
! In coronary artery disease, blood vessels supplying blood to the heart
tissue are clogged leading to heart attacks.
! In 1977, Andreas Gruentzig introduced a procedure called balloon angioplasty
where a balloon mounted on a catheter is inserted into the artery and inflated at
the site of the blockage to compress the plaque against the arterial wall thus
restoring blood flow. The procedure was very successful but a large percent
of the patients suffered a subsequent narrowing of the artery called restenosis.
! In 1978, after hearing a lecture by Gruentzig, Julio Palmaz conceived
the idea of a stent – a metal scaffold inserted into the artery and
expanded using a balloon to keep the artery propped open.
! He initially experimented with wires wrapped on pins inserted into
pencils, and wires soldered together. He finally got his inspiration from
a tool left behind in his garage by a worker.
! Working in San Antonio he partnered with a cardiologist, Richard
Schatz, and a restaurateur and investor named Phil Ramono to patent
and develop the stent.
! Johnson & Johnson licensed the technology, which was introduced into
the market in 1991.
! Today the stent is used in more than two million procedures annually
and is credited with saving numerous lives.
Table 1.1
Significant developments in the history of biomaterials
Various sutures, metal wires, pins
J. Lister
Aseptic surgical technique
C. Hansmann
Plates, screws for fracture fixation
Adolf Fick
Glass contact lens
W. D. Sherman
Vanadium steel plates and screws
A. S. Hayman, M. C.
Portable pacemaker
A. E. Strock
Vitallium for dental implants
Philip Wiles
Total hip replacement
Sir Harold Ridley
Intraocular lens
G. K. McKee, J. WatsonFarrar
Biomechanically successful total hip
Ingvar Branemark
Osteointegration of metal implants
A. B. Voorhees
Prosthetic vascular graft
Earl E. Bakken
Wearable pacemaker
Sir John Charnley
Use of polymer in total joints
L. Edwards, A. Starr
Mitral valve replacement
W. Kolff and others
Implantable artificial heart
Julio Palmaz
Balloon expandable stent
Adapted from reference 2.
Ever since the introduction of aseptic surgical techniques by Lister in the 1860s,
there has been rapid progress in the development of biomaterials and implants
for a variety of applications in the body (Table 1.1) including dental implants,
artificial total joints for hips (Figure 1.1), shoulders (Figure 1.2), and knees
(Figure 1.3), spinal implants (Figure 1.4), fracture fixation rods and plates, cardiac
pacemakers, stents to keep blood vessels open, and endovascular grafts to repair
aneurysms. Although historically metals have been extensively used as biomaterials, there has been a significant increase in the use of ceramics and polymers over
the past 40–50 years, thereby leading to a plethora of implants now available for
clinical applications.
Figure 1.1
An implant for total hip replacement. The long metal stem is inserted into the medullary cavity of the
femur. The metal hemisphere is lined with a polymer and is fixed to the acetabulum on the pelvic
side. (Courtesy of Exactech, Inc.)
Figure 1.2
An implant for total shoulder replacement. The long metal stem is inserted into the medullary cavity
of the humerus. The polymer component serves as a bearing surface. (Courtesy of Exactech, Inc.)
1.1 Definitions
Figure 1.3
An implant for total knee replacement. The polymer component is made out of ultrahigh molecular
weight polyethylene and serves as a bearing surface (Courtesy of Exactech, Inc.)
Figure 1.4
A spinal implant made of the polymer polyetheretherketone (PEEK). The fenestrated design
facilitates the introduction of bone graft. (Courtesy of Exactech, Inc.)
1.1 Definitions
Biomaterials do not necessarily have to be natural materials as the name may
suggest. In 1974, at the 6th Annual International Biomaterials Symposium held
at Clemson University, a biomaterial was defined as
. . . a systemically, pharmacologically inert substance designed for implantation within
or incorporation with a living system.3
This definition reflected the understanding of the use and function of implants at
that time and it imposed the requirement of inertness on the material. Through
the years, as science has progressed and resulted in a better understanding
of the interaction between biology and materials in the body, the definition of
biomaterials has changed as well.
In 1986, at a consensus conference of the European Society for Biomaterials, a
biomaterial was defined as
a nonviable material used in a medical device, intended to interact with biological
Perhaps an even more complete definition was provided by Williams5 as:
a material intended to interface with biological systems to evaluate, treat, augment,
or replace any tissue, organ, or function of the body.
Box 1.2
! Today, hip joint replacement surgery is a common therapy to treat hip joints
affected by acute arthritis or trauma. The implants represent a ball and
socket joint and often comprise a polymeric cup with a metal or ceramic ball
moving in it. But it took a lot of experimentation to reach this point.
! In 1925, M. N. Smith-Petersen, a Boston-based surgeon, tried to
re-surface the natural ball of the joint using a glass hemisphere but the
glass failed under stress.
! In the 1950s a procedure called hemiarthroplasty was popular. This
consisted of leaving the natural cup in place but replacing the ball
component of the joint.
! In the 1960s, John Charnley from England developed a full hip
replacement implant with a Teflon cup and a metal ball attached to a
metal stem that was inserted into the marrow cavity. When the Teflon
failed, he tried polyethylene with success. The polyethylene–metal
combination is still used to this day.
Thus, the definition of biomaterials has changed over the years as progress in
science and technology has made it possible to:
! use implants to rapidly restore organ and/or tissue function,
1.3 Types of bonds in materials
! influence the long term viability of implants by better designing the biomaterial–
biology interface, and
! drive the inevitable biological response in desired directions.
1.2 Changing focus
Not only has the definition of biomaterials been changed and refined over
the years, but so has the emphasis on biology – the bio aspect of biomaterials.
From the earliest days of biomaterials through the middle of the twentieth century,
work on biomaterials concentrated on basic material properties such as strength,
stiffness, ductility, fracture resistance, and corrosion. Over the ensuing decades,
there was growing realization that biology plays a significant role in the success of
an implant and should be taken into account in its design. In more recent years,
emphasis has turned to the manipulation and tailoring of the biological response
using principles rooted in the natural sciences. Biomaterials do not necessarily
have to be inert in the body, but rather should interact with it at the cellular and
molecular levels to ensure the success of the implant.
Although the application of biomaterials is unique, they still fundamentally belong
to the same families of materials as those used for various other industries including
construction, aerospace and sports equipment. In this respect, biomaterials are no
different from most other materials. In the next few sections of this chapter, we will
cover some of the basics types of chemical bonds that bind atoms to form materials as
well as the major types of materials. An understanding of the various bonds is
important, as bonds are responsible for a variety of material properties.
1.3 Types of bonds in materials
At a high level of classification, materials can be divided into two classes: natural
and synthetic. At a more fundamental level, materials can be separated into
different general categories based on their molecular structure and the type of
bonding between their atoms. Since the latter plays a primary role in determining
the properties of a biomaterial, it is important to first gain an understanding of
the different types of bonds.
1.3.1 Ionic bonds
Ionic bonds are based on one atom donating an electron to form a cation and
another atom accepting the electron to form an anion. The charged anion and





Figure 1.5
Ionic bonding between Na and Cl atoms.
cation are then held together by strong electrostatic attraction to create an ionic
bond. A common example of such bonding is in sodium chloride (NaCl), where
the sodium (Na) atom transfers one of its valence electrons to the chlorine (Cl)
atom (Figure 1.5). After the transfer of an electron, the Na atom becomes a net
positively charged ion and is represented as Naþ. The transfer of an electron to the
Cl atom renders it a negatively charged ion, which is represented as Cl−. The Naþ
and Cl− ions are then held together by Coulombic forces.
To reduce the anion–anion or cation–cation repulsion, each anion is surrounded by as many cations as possible and vice versa. In general, ionic bonding is
non-directional and the bond has equal strength in all directions. These properties
give rise to highly ordered structures, which result in solids that have high strength
and stiffness but are relatively brittle due to the inability of atoms to move in
response to external forces. Additionally, the electrons are closely held in place
and are not available for charge transfer, making the materials bad conductors of
electricity and heat. Bonding energies for ionic bonds are generally high and
usually range between 600 and 1500 kJ/mol, which is manifested in high melting
1.3.2 Metallic bonds
Metal atoms are good donors of electrons and metallic bonds are characterized
by tightly packed positive ions or cores surrounded by electrons. As shown in
1.3 Types of bonds in materials
Figure 1.6
Metallic bonds showing positive ions surrounded by electron cloud.
Figure 1.6, the valence electrons are not firmly attached to any one atom but
form an electron cloud to float and drift through the material.
The cores are highly organized structures, which consist of the nuclei and the
remaining non-valence electrons. Having a net positive charge, these cores would
repel each other were it not for the electron cloud surrounding them and working
as an “adhesive.” Additionally, the loose electron cloud allows for good charge
transfer, making the metals good conductors of electricity and heat. Bonding
energies can vary in metallic bonds, as shown by the examples of mercury
(0.7 eV/atom) and tungsten (8.8 eV/atom). The non-directional nature of this
bond gives planes of these ions the ability to slide on each other, thereby allowing
the materials to deform under applied forces.
1.3.3 Covalent bonds
Some atoms have the ability to share electrons in their outer shells with other
atoms. As shown in Figure 1.7, bonds formed by this sharing of electrons are
called covalent bonds. Many molecules comprising dissimilar elements use covalent bonds. Examples include organic materials such as ethylene (C2H6), which is
a gas, or its polymerized form polyethylene, which is a solid. Other non-metallic
Figure 1.7
Covalent bonds showing atoms sharing electrons in their outer orbits.
elements also form molecules (Cl2, F2) using covalent bonds. Carbon and silicon
use covalent bonds to form molecules such as silicon carbide.
Covalent bonds are strong and highly directional. However, the rotational ability
of atoms around bonds based on single-electron sharing (single bonds) generates
flexibility, and materials with such bonds usually have good deformational characteristics. As with ionic bonds, the electrons in these bonds are held in place, and
these materials are generally not good conductors of electricity and heat.
1.3.4 Secondary bonds
In addition to the three primary bonds described above, there are weaker bonds
such as van der Waals and hydrogen bonds. Such bonds are based on the
attraction between atomic or molecular dipoles. Dipoles are formed when there
is uneven or asymmetrical charge distribution, causing separation between the
1.4 Types of materials
negative and positive portions of the molecule. These dipoles result in electrostatic
attraction between adjacent atoms or molecules.
Hydrogen bonds are a special type of bonding between polar molecules and
occur when hydrogen is covalently attached to elements such as oxygen, nitrogen or
fluorine. Sharing of a single hydrogen electron results in a molecule which is positively
charged on one end and thus forms a dipole. Hydrogen bonds are the strongest of the
secondary bonds and can have bonding energies as high as 0.52 eV/atom, whereas
other secondary bonds typically have bonding energies circa 0.1 eV/atom.
! Ionic bonds are the strongest, followed by metallic, covalent, and secondary bonds in descending order. These differences are reflected in the
mechanical and physical properties of the materials.
! Ionic bonds are formed by the exchange of electrons between atoms. The
bonds are non-directional and the materials formed by such bonds have
uniform properties in all directions. The bonds are very structured and
so the materials are usually strong but brittle.
! The valence electrons in metallic bonds form a cloud surrounding the
cores of the atoms. Materials with metallic bonds are good conductors of
heat and electricity.
! Atoms forming covalent bonds share electrons in their outer shells. The
bonds are very strong and highly directional. Since the electrons are held
in place, the materials are not good conductors.
! Secondary bonds are formed due to the attraction between dipoles –
asymmetrically charged atomic or molecules. The bond strength is weak.
Secondary bonds do not involve the exchange or sharing of electrons, are less
directional, and possess less than 10% of the strength of covalent bonds. However,
they can significantly influence the properties of materials, especially polymers.
Now that we have learned about the basic types of bonds that hold atoms
together, we will explore the different types of materials that these bonds form in
the following sections.
1.4 Types of materials
1.4.1 Ceramics
The word ceramic is derived from the Greek word, “keramos,” which stands
for pottery. It has its roots in the ancient Sanskrit term “ker” or to burn. These
linguistic roots refer to the fact that subjecting these materials to high temperatures
often yields desirable properties. Ceramics are solid materials characterized
by ionic bonds or combinations of covalent and ionic bonds. Carbons are included
in this class. Usually ceramics are compounds of metallic and non-metallic
elements and often are nitrides, carbides and oxides. Ceramics may be crystalline
or amorphous (non-crystalline) glasses.
As a consequence of the nature of their bonds, ceramics are difficult to
deform and have low ductility. Thus, in general they are extremely hard and
brittle. Owing to this lack of ductility and hence devoid of the ability to blunt
crack tips effectively, ceramics are more susceptible to the presence of sharp
microcracks. These sharp cracks act as sources of stress concentration and as
stress risers. The effective stress at these crack tips can be several fold higher
than the applied stress, thereby resulting in material failure. Ceramics usually
have a low resistance to crack propagation and possess low fracture toughness.
They have high elastic modulus, high compressive strength but low toughness
and tensile strength. If processed into fibers, they can possess superior mechanical properties and are often used as the reinforcing component in composite
In general, ceramics are good insulators of both heat and electricity. They can
withstand high temperatures and challenging chemical environments better
than most other classes of materials. If properly polished they can also exhibit
good wear properties.
As a biomaterial, ceramics are used for total joint prostheses, and tissue
engineering scaffolds in the orthopedic field, as crowns for dental implants, and
in heart valves among other applications (Table 1.2).
1.4.2 Metals
Metals are inorganic materials with atoms held together by non-directional metallic bonds. These atoms are in a closely packed configuration causing metals to
have a high density and be readily visible on X-rays. The cloud of free electrons
surrounding the atoms makes metals good conductors of electricity and heat.
The strong nature of the bonds and the close packing renders metals strong and
they have high elastic modulus, strength, and melting points. In addition they are
ductile and can be formed into complex shapes using a variety of techniques such
as machining, forging, casting, and forming. Metals, however, can be susceptible
to corrosion, especially in media with chloride ions. This can be of concern when
they are used as biomaterials, and so, proper passivation of the metallic biomaterials is required.
1.4 Types of materials
Table 1.2
Examples of biomaterials and their applications
Alumina, Zirconia
Total joint prostheses, dental implants,
implant coatings
Calcium phosphates
Tissue regeneration scaffolds, cements,
drug delivery systems
Stainless steel
Fracture fixation devices, stents, wires,
dental implants
Cobalt chrome alloy
Total joint prostheses, stents
Titanium and alloys
Total joint prostheses, dental implants
Polymers (synthetic)
Total joints for hips, knees etc.
Bone cement
Vascular grafts
Polyethylene terephthalate
Vascular grafts, hernia repair meshes
Polydimethyl siloxane
Contact lenses, breast implants
Polyhydroxyethyl methacrylate
Contact lenses
Coatings for long term implants
Polylactic acid
Resorbable screws
Poly(lactic-co-glycolic) acid
Resorbable sutures
Drug delivery devices
Polymers (natural)
Tissue engineering scaffolds, cosmetic
bulking agent
Wound dressings
Of all materials, metals have the longest history as biomaterials and are
commonly used as load-bearing implants and fixation devices in the dental and
orthopedic fields (Table 1.2). They are also used in a variety of other applications
such as guide wires, stents, heart valves, and electrodes.
1.4.3 Polymers
Polymers are organic materials with long molecular chains held together by covalent
bonds along their backbone. Usually, they are based on carbon, hydrogen and
other non-metallic elements. A single sample of polymer may contain molecular
chains of a variety of lengths. These chains may be tangled and held together by
covalent bonds (cross-linked) or weaker bonds such as van der Waals and hydrogen
bonds. Polymers may be crystalline, amorphous, or a combination of the two.
Usually, they are not good conductors of heat or electricity.
Under an applied load, the molecular chains in a polymer have the ability to
slide past each other (unless cross-linked) or may have rotational capability about
their bonds. This flexibility renders the polymer highly deformable and ductile.
However, compared to metals and ceramics, polymers have low elastic modulus,
strength, and thermal transition temperatures. From a manufacturing perspective,
they can be easily and inexpensively produced and then shaped using a variety
of techniques including melt molding.
Polymers have seen an increasing role as biomaterials over the past 50 years and
are used extensively as sutures, coatings for leads, and implantable devices such as
pacemakers, bearing surfaces for total joints, resorbable meshes, fixation plates
and screws, contact lenses, and tissue engineering scaffolds (Table 1.2).
1.4.4 Composites
Composites are materials that comprise two or more different phases or
constituents with distinct interfaces. Typically, one of these phases is a “filler,”
and the other phase is a “matrix.” Although these phases can be theoretically
identical in composition, composites usually are combinations of different materials. The properties of composites are determined by the shape, size, orientation,
and distribution of the constituent materials as well as their relative proportions.
Acting as fillers, fibers of a material with high mechanical strength are often
dispersed in a matrix of a polymeric material to make a composite material with
high strength and toughness. If the reinforcing filler is in the form of fibers, the
resulting fiber-reinforced composite can be anisotropic if the fibers, with identical
properties, are aligned in a certain direction. Particulate-reinforced composites are
more likely to have isotropic properties, that is, the reinforcing particulates having
identical properties but are not aligned in a certain direction.
For composites to be successful in enhancing mechanical properties there is a
need for effective stress transfer at the interface of the filler and the matrix. Weak
interfaces result in separation of the constituents and failure.
1.5 Impact of biomaterials
Bone is a composite at the microstructural level with ceramic particles (apatite)
acting as the filler distributed in a matrix of collagen, which is a natural polymer.
This combination gives bone its strength and toughness. However, since the
collagen fibers are organized in aligned bundles, bone is anisotropic with different
properties in different directions.
1.5 Impact of biomaterials
Biomaterials have played a very significant role in improving the quality of human
life, especially for the elderly. A good example is the development of total joint
prostheses to replace natural hip and knee joints that have become non-functional
due to trauma or conditions such as arthritis. Although surgical experiments
with total hip replacements started in the 1890s, it was not until the early 1960s
that a successful total hip replacement was developed by Sir John Charnley using
stainless steel and ultrahigh molecular weight polyethylene. This was followed
by the development of total knee replacements in the period from 1968 to 1972.
Today, joint-replacement surgeries are immensely successful in restoring function
and relieving pain for thousands of patients who usually are walking within a
few days after their surgical procedure. In the USA alone, there were 193 000 total
hip arthroplasties and 381 000 total knee replacements in 2002.6 Compared to
1990, this signified a 62% increase in total hip implants and a 195% increase in
total knee implants placed in patients. In the area of orthopedics, success is not
limited to just the hip and knee implants, but replacements have also been
developed for shoulders, elbows, ankles, and finger joints. Additionally, the use
of intramedullary nails has significantly improved the outcomes of surgeries for
long bone fractures.
! The main types of bonds that hold atoms together to form materials are:
– ionic
– metallic
– covalent
– secondary.
! Common materials can be classified as:
– metals
– polymers
– ceramics
– composites.
Dental implants with good bone integration have had a major influence on
dentistry. These implants are used as posts on which porcelain crowns can be
fixed to yield artificial teeth that restore function and are also esthetically pleasing.
Another area where biomaterials have improved the quality of life is eye care.
Contact lenses have revolutionized the treatment of impaired eyesight. Soft
contact lenses have reduced the discomfort associated with the earlier hard lenses
and have led to the development of long-wear contacts.
Biomaterials have had a very significant impact on not just improving the
quality of life but also prolonging life. In the 1980s Dr. Julio Palmaz, then on
the faculty at the University of Texas Health Science Center at San Antonio,
developed the balloon expandable coronary artery stent to keep arteries open after
occlusions caused by plaque have been cleared using balloon angioplasty. This
metal (stainless steel or cobalt chrome alloy) device is now used extensively
around the world and is instrumental in reducing the incidence of heart attacks
and saving countless lives.
Other life-prolonging implants based on successful biomaterials include pacemakers, heart valves, and polymer-based aortic grafts and endovascular stent
grafts for treating life-threatening aneurysms.
Biomaterials-based implants and treatment strategies have become an integral
part of healthcare today. It is anticipated that the use of biomaterials will continue
to grow rapidly as scientists, engineers, and clinicians continue to attain a better
understanding of the interaction between materials and biology in the body.
1.6 Future of biomaterials
The field of biomaterials is changing as science yields a more detailed understanding
of human biology at the cellular and molecular levels. A large part of current
biomaterials research is devoted to the investigation of the interaction between
materials and proteins, cells and tissue. All biomaterials are immediately covered by
deposits of proteins and other biological molecules upon insertion in the body. These
deposits then play a significant role is determining the future response of the body
to the implant. Efforts are underway in laboratories around the world to develop
hybrid biomaterials where the surfaces of traditional biomaterials (such as those
described in Table 1.2 above) are modified by using biochemical moieties such
as proteins, peptides, and receptors for antibodies. These modified surfaces can
actively interact with the body’s molecules and cells to control and drive the
biological reaction to an implant in desired directions with the goal of achieving
better long-term outcomes. No longer is the goal to make implants simply “tolerable”
to the body but the aim is to develop biomaterials that interact in an intelligent fashion
with the host biology and bond with, repel, or evaluate their in vivo surroundings.
In the future the boundaries between biomaterials and naturally occurring
biological entities and systems in the body will further diminish and scientists
will be able to reproduce or grow tissue, organs, and biological molecules to
restore function.
1.7 Summary
Biomaterials have been in use since ancient times, but their progress was severely
limited by lack of aseptic surgical techniques. In recent years, there has been an
emphasis on ensuring the long-term viability of implants and how they interact with
the biology of the body. To begin to study biomaterials, it is important to
first understand the type of chemical bonds that hold atoms together. These basic
bond types include ionic, metallic, covalent, and secondary bonds. Each type of bond
imparts characteristic properties to the materials it forms and renders them hard or soft,
brittle or ductile, and good or bad conductors of electricity and heat. The basic families
of materials are ceramics, metals, polymers, and composites. Each of these material
types have distinct properties, which determine their behavior and make them more or
less appropriate for different applications as a biomaterial. In recent years, scientists
and engineers have been working to make biomaterials more responsive to the
conditions in the body so that they can interact with its complex biological milieu in
an intelligent fashion and to ensure a better outcome for the patient.
1. Duraiswami P. K., Orth, M. and Tuli, S. M. (1971). 5000 Years of Orthopaedics in
India. Clinical Orthopaedics and Related Research, 75, 269–280.
2. Bhat, Sujata, V. (2002). Biomaterials. The Netherlands, Kluwer Academic Publishers,
ISBN 0–7923–7058–9.
3. Park, J. B. (1984). Biomaterials Science and Engineering. New York, Plenum Press.
4. Williams, D. F. (1987). Definitions in Biomaterials, Proceedings of a Consensus
Conference of the European Society for Biomaterials. Chester, England, 1986,
Volume 4. New York, Elsevier.
5. Williams, D. F. (1999). The Williams Dictionary of Biomaterials. Liverpool, UK,
Liverpool University Press.
6. Kurtz, S., Mowat, F., Ong, K. et al. (2005). Prevalence of primary and revision total hip
and knee arthroplasty in the United States from 1990 through 2002. J. Bone and Joint
Surgery, 87, 1487–1497.
What is the main difference between the old definitions for biomaterials and
those from more recent years?
Which of the following would be considered to be made from a biomaterial
according to the modern definitions for a biomaterial:
(a) pacemaker for the heart,
(b) walking stick,
(c) stethoscope,
(d) vascular graft,
(e) toothbrush,
(f) suture?
Why are polymers not good conductors of electricity in general?
Which of the basic bonds are the most directional in nature? Which are the
Of the different families of materials which are generally the most hard and
brittle? Which are the least?
Of ceramics, polymers, and metals, which have low resistance to crack
propagation and fracture?
If you had to produce an artificial artery, what are the essential properties
you will incorporate in your design? Which type of material would you first
consider? Why?
You have been asked to design the stem component of a total hip joint. This
component is inserted into the medullary cavity of the femur and serves as a
major load-carrying unit. The load on a hip joint can exceed twice the
body weight during normal walking and can be even higher when running
or jumping. What family of materials would be a good candidate for this
Why are the surface properties of a biomaterial very important?
Is it preferable to make implants bioactive or inert? Why?
2 Basic properties of materials
After reading this chapter the student will understand the following.
! The mechanical properties and behavior of materials.
Material failure under ductile or brittle conditions.
The time-dependent mechanical behavior of materials.
Concepts related to the basic surface properties of materials.
Corrosion and its various forms.
In everyday life, we often define materials using relative terms such as soft or
hard, flexible or rigid, strong or weak, tough or brittle, and in a variety of other
qualitative ways. What do these terms really mean in the world of engineering?
Is such qualitative categorization sufficient for the design and manufacture of a
product? The answer is definitely a no, especially when human health or lives may
depend on the product. For example, you certainly would not want engineers who
are building bridges to pick materials based on such relative and qualitative
descriptions! Choices based on much more rigorous, scientific, and quantitative
characterization would be expected. The same is true when selecting biomaterials.
Material properties can be characterized quantitatively using standardized tests
under defined conditions. Once characterized, these properties can be used in
conjunction with engineering design techniques to predict the behavior of the
engineered product under the expected operating conditions and to ensure that it
would function safely. This is important because properties may change based on
independent variables such as temperature or rate of application of force. Often a
variety of material properties need to be considered for each product.
Some material properties fall in the category of mechanical properties, which
predict the deformation, failure behavior, and fracture of materials under the action
of forces. As an example, mechanical properties would be very important for a
Basic properties of materials
hip-replacement implant because it would be expected to withstand heavy loads
generated during walking, and such loads can be as high as several times a person’s
body weight. For some materials and operating conditions, the electrochemical
properties, which characterize corrosion behavior, may be important. This would be
true for metals used for a stent that is placed in contact with blood inside an artery.
Surface properties such as surface energy, hydrophilicity, or hydrophobicity can be
important for implants because these characteristics influence whether cells would
attach to the material or determine how proteins will interact with the surface. Since
the in vivo environment is very complex, the careful selection of appropriate
biomaterials based on material properties is essential. In this chapter, we will
discuss some of the basic material properties and how they are measured.
2.1 Mechanical properties
The response of a material to deforming forces is characterized by its mechanical
properties. These forces can be of various types such as tensile, compressive,
torsional or combinations of these forces (Figure 2.1). In general, the type of
response varies based on the kind of force applied. Other factors that also play a
role in changing the response include the rate of application of force, the temperature of the material, and the surrounding environmental conditions.
To determine the mechanical properties of a material, force versus deformation
tests are conducted. In these tests, samples of a material are loaded at a constant
rate and both the deformation and the force required to cause that deformation are
measured at various time points. To avoid complexities that may otherwise arise,
most mechanical properties reported are based on unidirectional and uniaxial
loading. However, multi-axial load tests can be conducted if needed. Also under
dynamic loading conditions, the direction of the applied force can be reversed
and varied to give sinusoidal, saw-tooth, square-wave or other loading patterns.
Materials can also be subjected to time-dependent tests. For example a material
sample may be loaded to a certain deformation and that amount of deformation is
held constant while the behavior of the force is measured over a period of time.
Conversely, a material may be subjected to a constant force and the deformation is
recorded as a function of time.
Properties measured under various conditions and testing modes have distinct
standard definitions and nomenclature. Standardized testing methods are made
available by agencies, such as the American Society for Testing and Materials
International (ASTM) and the International Organization for Standardization
(ISO), so that properties measured at different test facilities can be meaningfully
2.1 Mechanical properties
Figure 2.1
Different types of loading: (a) tensile; (b) compressive; (c) shear.
2.1.1 Tensile testing
In the classic form of tensile testing, both ends of a material specimen are clamped
between a pair of jaws. The lower jaw or holder is held fixed in place. As shown in
Figure 2.2, the upper jaw is attached to the moving crosshead of a tensile tester via
a load cell, which measures the applied force. The material specimens tested can be
bar shaped with straight sides and uniform rectangular or round cross-sections.
Alternatively, the test specimens can have varying cross-sections and be shaped into
“dog bone” (rectangular cross-section) or “dumbbell” shapes (circular crosssection), as seen in Figure 2.3. An extensiometer is used to measure the change
in the dimension of the test sample along the axis of loading. Using this test,
information on force versus deformation is collected and used to develop a stress–
strain plot. Shown in Eqs. (2.1) and (2.2) are the formulae for stress and strain, where
stress has units of force per unit area, and strain is dimensionless and has no units:
stress ðσ Þ ¼
original cross-sectional area
strain ðεÞ ¼
change in length
original length
The gage section of the test sample is measured to determine both its crosssectional area and the length (Figure 2.4). For straight-edge samples, the gage
section is the portion of the sample between the upper and lower jaws. For the
shaped dog bone and dumbbell samples, the gage section is the mid-section,
which has a uniform but reduced cross-section.
Basic properties of materials
Figure 2.2
A tensile testing machine.
! It is imperative to understand that force and stress are not one and the
! A force of the same magnitude can generate very different stress levels
based on the area over which it is applied.
! Example: a force applied on an apple through the flat side of a knife
blade (large area) may cause a small depression but when the same force
is applied through the edge of the blade (small area), it will cut the fruit
because of the very high stress generated.
In the seventeenth century, Robert Hooke showed that when a solid material is
subjected to a tensile force, its deformation is proportional to the load applied.
This is known as Hooke’s law and is represented by the linear region of the
material’s stress–strain curve. An example of a stress–strain curve is shown in
2.1 Mechanical properties
Figure 2.3
Standard types of specimens for testing the mechanical properties of a material using a tensile
testing machine: (a) dog bone and (b) dumbbell.
s= F
e = Δ!
Figure 2.4
A rod under tensile loading showing stress and strain.
Figure 2.5. In the linear region of the stress–strain curve, the material is said to be
elastic and behaves like a spring. The slope of this linear region of the curve is
known as the elastic modulus, E. Equation (2.3) defines the relationship between
elastic modulus, stress, and strain:
σ ¼ Εε:
The proportional limit is the point on the stress–strain curve corresponding
to the highest stress at which the stress is linearly proportional to the strain.
Basic properties of materials
Figure 2.5
A stress–strain curve showing the yield stress (σy), the ultimate strength (σu), and the elastic
modulus (E) given by slope of the linear portion. The shaded area under the curve represents the
toughness of the material.
The yield stress, σy, is the stress which causes the onset of permanent deformation
in the material, and this stress is sometimes defined as the yield strength or
the elastic limit of a material. Permanent deformation implies deformation that
persists even when all loading is removed from the material. The point on the
stress–strain curve corresponding to the yield stress is known as the yield point.
The proportional limit and the yield stress can be the same value, but the proportional limit often precedes the yield stress. If the stress–strain curve is not linear
or a clear yield point is not obvious, then alternate means are used to determine
the yield stress. In such circumstances, a 0.2% strain technique is generally
used to determine the yield point. In this method, a line parallel to the linear part
of the stress–strain plot is drawn to intersect the strain axis at 0.2% strain.
The point where this line intersects the stress–strain curve is designated as the
yield point.
Under elastic conditions, the material reverts to its original dimension
when unloaded. On the other hand, if the stress is increased beyond the elastic
limit, the material may either undergo failure or become permanently or plastically deformed. Plastic deformation is made possible by the movement of
large numbers of atoms or as in the case of polymers, movement of molecular
chains. Brittle materials, such as ceramics, do not exhibit plastic deformation or
yield points. Other mechanical properties can also be obtained from the stress–
strain curve: the maximum stress reached prior to fracture is defined as
2.1 Mechanical properties
the ultimate tensile strength; the area under the curve up to the failure point
is a measure of the work required to cause fracture and is a measure of the
toughness of the material.
For many ductile metals the stress levels continue to increase after the yield
stress is reached and plastic deformation is initiated. This is due to a phenomenon
known as strain hardening, which is the resistance to further deformation.
This resistance is caused by the decreased mobility of atomic planes within
the material due to the interaction of multiple dislocations (imperfections in the
lattice of atoms in a material). As a result, the stress continues to increase until it
reaches the ultimate strength, beyond which the material deforms rapidly until
it fractures. It is interesting to note that materials with very few dislocations,
such as single crystals, or materials with large number of dislocations tend to have
high strength.
! The stress value calculated for most engineering design considerations
is called the engineering stress and is defined as σ ¼ F/A, where A
represents the original cross-sectional area of the sample.
! In reality the cross-sectional area changes with loading as the sample
! The true stress in the sample can be calculated by using the instantaneous area.
In general, the elastic modulus of polymeric materials is low and they tend to
undergo large plastic deformations. The linear region of their stress–strain plot
is usually short or may be non-existent. In some polymers, the slope of the
stress–strain plot may increase after initial loading due to the alignment of
molecular chains. Most polymers have relatively low ultimate strength but
deform significantly before fracture and possess high toughness (Figure 2.6).
Metals have higher elastic modulus, a clear linear region and often a well-defined
yield point, beyond which they undergo plastic deformation. They have high
ultimate strength. Ceramic materials have very high elastic modulus values
and very little deformation. They usually have high strength but are brittle with
very low toughness.
As a material is loaded, it undergoes deformation along the axis of the load.
However, since its volume does not change, the longitudinal deformation necessitates a corresponding transverse deformation. For example, if a cylindrical
specimen is loaded in tension and undergoes an increase in length, it will also
exhibit a simultaneous decrease in diameter in order to maintain a constant
volume. This is known as the Poisson Effect. The ratio of the transverse strain
Basic properties of materials
Figure 2.6
Representative stress–strain curves for ceramics, metals, hard tissue, and polymers.
divided by the longitudinal strain is called the Poisson’s ratio, as denoted by
ν (nu). For most isotropic simple materials, the Poisson’s ratio lies in the range
0.2–0.5. Steels, when loaded within their elastic limits (before yield), usually
exhibit a ν value of approximately 0.3. Rubber has a ν value of nearly 0.5, whereas
cork has very little transverse deformation and has a ν value close to 0. A negative
value of Poisson’s ratio indicates a material which, when stretched in one direction, becomes thicker in the perpendicular direction. Such materials are called
auxetic and are usually composed of macroscopic elements with hinge-like structures. Some polymer foams behave in this manner.
2.1.2 Compressive testing
In compressive testing, block-shaped samples are used instead of the samples
described for tensile tests. The same type of test equipment may be used as in the
tensile tests but with modified holders or fixtures. Depending on the type of the
fixture, the machine crosshead can be lowered or raised to impose compressive
loading on the specimen. Often, the failure modes as well as elastic properties
differ between tensile and compressive testing modes. For example, ceramics are
usually stronger in compression than tension.
2.1 Mechanical properties
Figure 2.7
Deformation under shear loading.
2.1.3 Shear testing
In shear tests, the load is applied parallel to the surface on which it is acting
(Figure 2.7). The deformation is measured in the direction of the force applied.
The shear stress is usually denoted by τ, the shear strain by γ, and the shear modulus
by G. Hooke’s law applies in the elastic region for shear, just as in the case of
tension and compression. The relationship between shear stress, shear strain, and
shear modulus is represented in Eq. (2.4):
τ ¼ Gγ:
For an isotropic material, the mechanical properties are uniform in all directions,
irrespective of the orientation of the test specimen or direction of testing. For such
materials, only two constants, E and G, are needed to fully characterize their
stiffness. However, for an anisotropic material, the properties differ in different
directions, and thus, the results of the tests described above would vary with
orientation. Most natural tissues, such as bone, ligaments, skin, and cartilage, are
anisotropic in nature.
2.1.4 Bend or flexural tests
Samples can also be tested under a bending load to determine their mechanical
properties. In such cases, the samples have a uniform cross-section, which can be
rectangular or circular. In the case of a three-point test (Figure 2.8a), the sample is
laid on two supports that are placed at a distance, L, apart. A load is imposed at the
mid-point of the span, causing the sample to bow. In this test configuration,
the material on the inner side of the bow is in compression while that on the
outside is under tensile loading (the reader is encouraged to sketch a bent rod to
Basic properties of materials
Figure 2.8
Bend tests: (a) three-point test, where the applied load is at one location; (b) four-point test,
in which the applied load is at two places between the supports.
see why this happens). For a beam with a rectangular cross-section, Eqs. (2.5) and
(2.6) are used to determine stress and strain, respectively:
σ f ¼ 3FL=2bd2 ,
εf ¼ 6Δd=L2 ,
where σf ¼ stress in the outer material at the mid-point (MPa),
εf ¼ strain of the outer material at the mid-point,
Δ ¼ maximum deflection of the mid-point under the load (mm),
d ¼ depth or thickness of the beam (mm),
b ¼ width of the beam (mm),
L ¼ length of the span between the two supports (mm), and
F ¼ load applied at mid-point (N).
In a four-point test (Figure 2.8b), the load, F, is delivered through two points
(F/2 at each point) within the span, L, between the two support points. In this test
configuration, the stress is constant between the two loading points. While
the four-point test is preferable for determining material properties, the threepoint test is more useful for determining the overall strength of a sample or a
2.1.5 Viscoelastic behavior
The properties described above assume that the deformation in a material is
instantaneous at the application of a force. However, in some materials, there
may be an additional time-dependent deformation component due to viscous
flow within the material. These materials are known as viscoelastic materials
and can be modeled using a combination of a spring and a dashpot, as shown in
Figure 2.9.
2.1 Mechanical properties
Figure 2.9
Models for a viscoelastic material: (a) cylinder and piston or dashpot model for viscous
deformation; (b) dashpot and spring model for combined viscous and elastic deformation.
Figure 2.10
The different stages of creep deformation. Upon loading of a viscoelastic material there is initial
elastic deformation followed by an increasing rate of strain in primary creep. This is followed by
a linear creep strain–time relationship during secondary creep. Lastly, under tertiary creep, there is
rapid deformation leading to failure.
Viscoelasticity is manifested in creep and stress relaxation behavior. Creep
occurs when a viscoelastic material is loaded quickly to a certain stress, followed
by that stress being held constant over time. Under this type of loading, the
viscoelastic material will exhibit an instantaneous deformation (the equivalent of
the spring in the model), followed by an additional deformation that is proportional to time (viscous dashpot). Creep deformation usually occurs in the
following three stages (Figure 2.10):
Basic properties of materials
! primary creep, where there is a high rate of deformation which slows down rapidly,
! secondary creep, where the deformation is constant with respect to time on a log scale,
! tertiary creep, where there is a high rate of deformation followed by failure.
However, all three stages of creep deformation may not always be evident.
Usually high temperatures are necessary for tertiary creep. An increase in temperature and stress results in an increase in creep. Although the occurrence of
creep does not require a minimum stress level, creep may not occur below a
minimum temperature.
When a viscoelastic material is subjected to a load during a test, the deformation
that is measured is a combination of the instantaneous elastic deformation and the
viscous deformation that occurs over the duration of the test. This combined
deformation yields a higher strain and thus a lower value for the elastic modulus,
which is represented by the slope of the stress–strain curve. If the rate of loading is
faster, the viscous deformation component is lower due to the shorter time
available for viscous flow to occur within the material. Hence, the calculated
elastic modulus for rapid loading is higher. If the rate of loading is progressively
increased, the viscous deformation will continue to decrease until a limit is
reached where all the deformation is elastic and the viscous component is negligible. The slope of the stress–strain curve at this limit represents the true elastic
modulus for the viscoelastic material. Owing to the relationship between the
mechanical properties measured and the loading rate used for the test, it is
important to either use standardized loading rates or clearly specify the loading
rates when reporting test results for viscoelastic materials such as polymers and
Stress relaxation is observed when a viscoelastic material is rapidly deformed to
a certain level, and the deformation is held constant while the stress is measured
over time. In this case, the stress will decrease as a function of time. Viscoelasticity is exhibited in most polymers, in some metals, but rarely in ceramics.
2.1.6 Ductile and brittle fracture
Failure in a material can be defined in a variety of ways. For example, the onset
of plastic deformation, reaching a certain percent strain, or final fracture can all be
considered failure based on the functional constraints placed on a device.
Fracture itself may be defined as the separation of a body into multiple pieces
and occurs when the cohesive strength between adjoining atoms is exceeded by
the applied stress. The failure mode at fracture can be divided into two categories:
2.1 Mechanical properties
Figure 2.11
Types of fracture: (a) moderately ductile fracture showing a cup–cone failure surface; (b) brittle
fracture usually presents a macroscopically flat surface.
ductile or brittle. Whether an engineering material undergoes ductile or brittle
fracture depends on its intrinsic material properties as well as its temperature
and rate of loading.
Ductile fracture occurs after a material undergoes plastic deformation. As the
stress level increases, microdefects, such as voids or cavities, form inside the bulk
of the material or on its surface. These defects then coalesce to form a crack,
which grows by the assimilation of more defects. This crack usually travels
perpendicular to the applied load. Once the crack attains a certain length, it reaches
a critical size and starts to propagate rapidly. In the case of a test specimen with a
round cross-section, as the fracture approaches the surface, failure occurs at 45% to
the applied load along the plane of maximum shear stress. This gives the appearance of a cone shape to one failed section and a cup shape to the mating section,
leading to a “cup and cone” fracture (Figure 2.11). When examined under the high
magnification of a scanning electron microscope, the fracture surface of a ductile
failure reveals a dimpled texture. These dimples represent the microdefects that
coalesced to form the crack.
Basic properties of materials
True brittle fracture is characterized by no plastic deformation, and failure
occurs by the cleavage of atomic bonds along crystallographic planes. The brittle
crack propagates perpendicular to the applied stress and the fracture surface is
macroscopically flat. Under the microscope, v-shaped chevron marks or a fan-like
pattern radiating from the point of origin of the crack may be seen. For materials
with a very fine grain structure, the surface may be shiny and without texture.
While ceramics usually undergo brittle fracture, metals generally exhibit ductile
failure unless hardened through heat treatment. Polymers can undergo either
type of failure, depending on whether they are in an amorphous glassy or a
crystalline state. However, even ductile materials can exhibit brittle fracture under
some circumstances, a phenomenon that can be explained through the field of
fracture mechanics.
2.1.7 Stress concentration
Based on cohesive bonds between atoms, theoretical calculations show that the
failure of a material should occur at a stress level of approximately E/10, where E
is the elastic modulus of the material. However, in engineering tests, failures occur
at stress levels vastly below this number – as low as E/10 000 in some cases. In the
first half of the twentieth century, this discrepancy was brought to the forefront by
some spectacular failures of ships and airplanes. In a few cases, unexpected,
rapidly propagating cracks broke entire ships into two while under normal sailing
conditions. In other cases, jet aircraft lost their wings while in flight due to similar
cracks. These systems had been designed properly using the engineering standards
of that time and yet underwent catastrophic failures.
These failures led to renewed interest and research in the area of fracture of
materials. In the 1920s, A. A. Griffith postulated that most materials contain inherent flaws in the form of microscopic voids or cracks. These flaws serve as points of
stress concentration or as stress risers. Consequently, the applied stress is magnified
at the tips of these flaws, and the degree of stress amplification increases as the
radius of the tip of the flaw decreases. Thus, the sharper the crack, the higher will be
the stress. The flaws do not have to be microscopic in size to cause stress concentration. In fact, any discontinuity, such as a screw hole in a fracture fixation plate or
a change in cross-section in the stem of a total hip prosthesis, can cause stress
amplification. The local stress at these stress risers can exceed the theoretical
cohesive strength of the material and cause fracture, even though the applied stress
or the stress at locations away from the stress riser are below this level. As shown in
Eq. (2.7), the stress at the edge of a defect is proportional to the applied stress and the
shape and size of the defect:
2.1 Mechanical properties
σ t ¼ k : σ,
where σ ¼ applied stress,
σt ¼ stress at edge of defect, and
k ¼ factor determined by shape and size of flaw or defect; k has a value of 2 for a
hemispherical defect but can be very high for other cases depending on the sharpness
(radius) of the crack tip.
As a crack propagates during fracture, it has a sharp tip (small radius) which causes
a stress concentration and a high stress ahead of the crack. In the case of a ductile
material, this stress causes plastic deformation of the material at the tip, thus
blunting the crack and reducing the stress. However, since a brittle material
does not exhibit plastic deformation, the crack tip in it remains sharp causing high
stresses and rapid failure. Brittle materials are usually stronger under compressive
loading compared to other modes of loading because the crack tips are compressed,
decreasing the rate of crack propagation.
2.1.8 Fracture toughness
The stress distribution near a flaw can be described in terms of the stress intensity
factor, K. This factor helps to determine the stress at any point situated at a
distance, r, and at an angle, θ, from the crack tip. For a given stress σ,
K ¼ Yσ ðaÞ1=2 ,
where Y ¼ a constant based on the geometry of the specimen and the crack,
σ ¼ the stress at (r, θ), and
a ¼ the crack length.
A critical value of K defines the conditions for brittle fracture and is known
as the fracture toughness K c . Based on the applied stress and the size of the
existing crack, brittle fracture will occur when the value of K in Eq. (2.8)
exceeds Kc. Thus, fracture toughness is a material property that is a measure
of the resistance of the material to brittle fracture in the presence of an existing
flaw. During the engineering design of a medical device, the ultimate or yield
strengths need to be taken into consideration and an ample factor of safety
should be built into the design. However, it is also prudent to consider the
fracture toughness and design to prevent brittle fractures. Such analysis is either
based on the maximum-sized defect that may potentially exist within a material
without detection, or on the size of a known notch, hole, or slit in the designed
Basic properties of materials
2.1.9 Fatigue
When subjected to a single application of load, materials fail at their ultimate
strength. However, if they are subjected to multiple load cycles, they can fail at
much lower stress levels. This is known as fatigue failure. The stress cycles do not
necessarily have to completely unload the specimens but can only partially reduce
the load. The loading profile can take a variety of forms including sinusoidal, sawtooth, or otherwise uneven. Fatigue failure is a serious issue for implants because a
variety of devices, such as total joint prostheses, dental implants, and heart valves,
are subjected to cyclic loading.
Fatigue failure has several stages and the first stage of fatigue-induced failure is
crack initiation. Cracks usually initiate either at existing sites of stress concentration
or as the result of an increase in the number of dislocations within the crystal
structure of the material, which leads to imperfections and crack formation. The
second stage of fatigue failure involves crack growth or propagation which occurs
because the tip of the newly initiated crack acts as a stress riser, leading to sharply
elevated stresses and material failure at the tip. The cyclic nature of the stress causes
the crack growth to occur intermittently with propagation taking place every time the
material is loaded. The third stage is the final fracture, and this takes place once
the crack grows to a critical size and the critical stress intensity factor is reached.
! Both strength and fracture toughness are material properties.
! Stress intensity is related to fracture toughness in the same manner as
stress is related to strength.
! Failure occurs when stress exceeds strength or stress intensity exceeds
fracture toughness.
! Fracture toughness should not be confused with toughness, which is a
different material property and a measure of the amount of energy
required to cause failure.
The fracture surfaces of materials that have failed under fatigue are characteristic in
nature and exhibit striations or wave-like patterns emanating from the point of
initiation. The distance between these striations is a measure of crack growth for
each loading cycle.
As shown in Figure 2.12, the fatigue properties of a material are graphed as a
stress (S)–number of cycles to failure (N) curve. The number of cycles is plotted
on a logarithmic scale and every point on the curve denotes a 50% probability of
failure at that given stress–cycle combination.
2.2 Electrochemical properties
Figure 2.12
Examples of fatigue S–N curves. Solid lines show the behavior of a material with an endurance
limit (below a certain load the material will not undergo fatigue failure). Dotted line shows typical
behavior for a material which does not exhibit a fatigue limit.
In most materials, there is a minimum stress level below which fatigue failure
does not occur irrespective of the number of loading cycles. This stress level is
known as the endurance limit, and the S–N plot becomes horizontal at this point.
However, certain metals, such as aluminum and its alloys, do not have an endurance limit, and as a result, their S–N curve shows a steady decrease in the failure
stress with increasing loading cycles. The use of such materials should be avoided
where cyclic loading is anticipated.
2.2 Electrochemical properties
2.2.1 Corrosion
Corrosion may be defined as the chemical deterioration of a material as a result
of interaction with its environment. Although this broad definition covers
environment-related damage in all materials, including both ceramics and plastics,
this section will be limited to corrosion in metals, while the other material types will
be addressed in later chapters. Corrosion may be broadly divided into the following
two categories:
! wet corrosion, which takes place in the presence of a liquid, and
! dry corrosion, which occurs due to corrosive gases.
Of the above two different categories of corrosion, only wet corrosion is relevant
in the case of biomaterials.
Basic properties of materials
Corrosion of implants in the body is an important issue because the in vivo
environment is complex and the implants are exposed to a multitude of chemical
moieties. The severity and speed of corrosion depend primarily on the chemical
makeup of the environment as well as other variables, such as temperature and
stress levels.
For example, when iron is placed in hydrochloric acid, a strong reaction takes
place and hydrogen gas is released. This reaction may be summarized as below:
Fe þ 2HCl ! FeCl2 þ H2 :
This outcome is the combination of partial reactions:
Fe ! Feþ2 þ 2e ðanodic reaction or oxidationÞ,
2Hþ þ 2e ! H2 ðcathodic reaction or reductionÞ,
Feþ2 þ 2Cl− ! FeCl2 :
In the above equations, e represents an electron. The net result is the loss of iron
(Eq. (2.10)) due to corrosion and its conversion to ferrous chloride (Eq. (2.12)).
In the corrosion process, both oxidation and reduction processes take place
simultaneously and at equal rates.
Another example is the corrosion of iron in the presence of an aqueous medium
with dissolved oxygen. The oxidation and reduction of this corrosion phenomenon
is represented in the following equations:
2Fe ! 2Feþ2 þ 4e ðoxidationÞ,
O2 þ 2H2 O þ 4e ! 4OH− ðreductionÞ,
2Feþ2 þ 4OH− ! 2FeðOHÞ2 :
Since ferrous hydroxide [Fe(OH)2] is unstable, this by-product is further oxidized
to form ferric hydroxide [Fe(OH)3], which is commonly known as rust:
2FeðOHÞ2 þ H2 O þ O2 ! 2FeðOHÞ3 :
Equation (2.17) can be used to describe the general oxidation reaction of a
metal (M):
M ! Mþn þ ne:
The reduction reaction can take several forms:
! hydrogen evolution, as described in Eq. (2.11),
! oxygen reduction in netural or basic solutions, as described in Eq. (2.14), and
2.2 Electrochemical properties
Figure 2.13
Corrosion behavior of an active–passive material as a function of electrode potential.
! oxygen reduction in acidic solutions, as described in Eq. (2.18):
O2 þ 4Hþ þ 4e ! 2H2 O:
Metallic passivity is an interesting phenomenon that is complex in nature, and its
reasons are difficult to explain. Usually, as the oxidizing power of a solution is
increased, the rate of corrosion increases exponentially. This is known as the
active zone. However, in certain metals, such as iron, chromium, titanium, and
their alloys, an increase in the oxidizing power beyond a certain critical point
results in a decrease in corrosion rates (Figure 2.13). This decrease in corrosion
rate can be significant and metals tend to behave as inert noble metals. The active
region is followed by the metal reaching a stable passive region wherein increasing the oxidizing power of the solution has no effect on the corrosion rate.
However, further increase in the oxidizing power beyond the stable passive region
results in an increase in the corrosion rate once again. This zone is known as the
transpassive region.
2.2.2 Types of corrosion
Wet corrosion can be categorized into several different forms. Some of these
forms of corrosion are more relevant to metallic implants than others, but all forms
need to be considered when designing products.
Basic properties of materials
Uniform attack
Uniform attack is the most prevalent of all forms of corrosion, and its most
common example is rust. This type of corrosion occurs uniformly over a surface
exposed to a solution which makes oxidative and reductive chemical reactions
possible. Inherent variations in the material may lead to different sections acting as
anodes and cathodes. In such a case, electrons flow from the anode to the cathode.
Uniform attack occurs in all cases where there is not an equilibrium concentration
of metals ions in the surrounding media. The corrosion rate is usually slow, and in
the case of cobalt chrome alloys, may result in a thickness loss of 0.1 µm per year.
Galvanic corrosion
Metals have a characteristic electrical potential. Under ideal conditions, these
potentials can be measured and referenced against the hydrogen electrode.
The potential for the hydrogen electrode is defined as zero. When two metals
with different potentials or electromotive force (EMF) are connected, there is a
possibility of current flow. Immersing this pair of metals in an ionic solution leads
to the completion of the electrical loop and corrosion ensues. The electrode
potentials for various different metals are listed in Table 2.1. However, it is not
possible to find these potentials for alloys with reactive components. In addition
to the electrical potential, the environment also has an effect on the corrosion of
the metals. Thus, a practical grouping of metals has been developed using seawater and is known as the galvanic series (Table 2.2). There is not a significant
difference between the positions of metals on the EMF and the galvanic series.
At the top of the list are the noble metals or cathodic materials, while the anodic
materials are listed at the bottom. If the metals or alloys are in their passive state,
they are listed further up in the series and closer to the noble metals.
The farther apart the two metals or alloys are in the galvanic series, the greater
the difference will be in their potential. The greater the potential difference, the
higher the rate of corrosion will be when these two metals are in contact or
connected. On the other hand, materials that are very close to each other in the
galvanic series pose little danger of galvanic corrosion. The degree of corrosion
also depends on the distance from the junction of the two metals, with the
maximum corrosion occurring at the junction site and decreasing at spots farther
away. The relative surface area of the anode and cathode also influence the amount
of corrosion, with increased corrosion observed when the anode is much smaller
than the cathode.
Implants made of more than one metal, or implanted metal devices of different
metals but are in contact with each other can undergo galvanic corrosion.
2.2 Electrochemical properties
Table 2.1
Metal ion

Standard EMF series for metals
Electrode potential
(volts) vs. normal
hydrogen electrode
at 25 % C
Noble or cathodic
Active or anodic
For additional information consult Mars G. Fontana and Norbert D. Greene,
Corrosion Engineering, ISBN-13: 978-0-071-00360-5, McGraw Hill Higher
Education, 1986.
This may occur in implants such as modular hip joints, dental implants, fracture
fixation plate and screw combinations, or an intramedullary nail in contact with a
cerclage wire holding pieces of bone together. Sometimes, an inclusion or impurity introduced into a metal implant during fabrication can act as the second metal
and set up a galvanic corrosion pair.
Crevice corrosion
Crevice corrosion occurs in restricted areas where there is fluid stagnation, such
as in cracks, modular press-fit joints, beneath screw heads or underneath gaskets.
The corrosion is driven by reactions described in Eq. (2.13) and Eq. (2.14).
Basic properties of materials
Table 2.2
Galvanic series for select metals
Noble or cathodic
Active or anodic
18–8 Stainless steel (passive)
Bronzes (Cu–Sn)
Brasses (Cu–Zn)
Nickel (active)
18–8 Stainless steel (active)
With time, the oxygen in the crevice gets depleted due to limited fluid flow and the
reduction reaction slows down in that locale. The reduction reaction continues
unfettered in areas external to the crevice, and the number of metal ions going into
solution does not decrease significantly to maintain balance between the two
processes in the crevice. This leads to a buildup of positive metal ions in the crevice,
attracting negatively charged chloride ions. The presence of chloride ions in the
crevice, for reasons not fully understood, further accelerates the dissolution of metal
ions in the crevice, causing even more chloride ions to migrate, and setting up an
autocatalytic process. Crevice corrosion is very localized and causes severe localized damage while the surrounding areas of the metal remain relatively untouched.
Pitting corrosion
Pitting corrosion is similar to crevice corrosion in its mechanism and can be
initiated by scratches, inclusions or other damage. Usually, it starts on horizontal
surfaces, and its growth follows gravity. Once initiated, the corrosion quickly
forms surface pits and can grow at ever increasing rates, causing holes in the
metal. However, since the corrosion is on a surface and not protected within a
crevice, the initiation sites can be unstable and easily disrupted or stopped by any
change in convective flow which may alter ion concentration. The pits are usually
numerous in number and small in size. Often the pits change the surface texture of
2.2 Electrochemical properties
the implant giving it a matte-finish look. Improved manufacturing processes have
led to fewer inclusions in implants and a reduction in pitting corrosion.
Intergranular corrosion
Cast or forged metals and alloys contain continuous regions known as grains.
The size and structure of these grains depend on the material and also on heat
treatment. During manufacturing, the size of the grains is manipulated intentionally by varying heating/cooling rates to obtain desired properties such as increased
hardness or ductility. The areas at the edges of the grains are disordered and are
known as grain boundaries. These boundaries often contain material compositions
different from the grains and may contain more impurities. For example, the
chromium content is depleted at the grain boundaries in stainless steel. These
variations, coupled with the high surface area of the grain boundaries, render them
more reactive and hence subject to corrosion. Such localized corrosion at or near
the grain boundaries with limited corrosion in the grains is known as intergranular
corrosion and can lead to a decrease in mechanical properties. In extreme cases,
the metal can disintegrate as the grain boundaries are destroyed, and the grains fall
out. The incidence of intergranular corrosion can be decreased by using proper
heat treatment and by reducing impurities and inclusions.
Leaching is the selective dissolution of a constituent of an alloy from the grains.
This may happen for a variety of reasons: the device or implant may be immersed
in a medium that selectively attacks one component; or more than one phase may
exist within the alloy. These phases, although having the same constituents, may
have different relative compositions with different dissolution properties. Leaching is often a concern when nickel-based alloys are used as biomaterials, due to
documented adverse responses to nickel in a segment of the patient population.
Erosion corrosion
Erosion corrosion is defined as the acceleration in corrosion damage when there is
relative movement between a metal surface and the corroding media. The actual

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