Extract
Advances in orthopaedic surgery have been based on biomechanical principles for years. Engineering advances in metallurgy, polymer science, ceramic technology, and manufacturing processes lead to improvements in the devices that orthopaedic surgeons use and thus to better performance in the fields of orthopaedics, especially arthroplasty, spine surgery, trauma, and arthroscopy.
Advances in orthopaedic surgery have been based on biomechanical principles for years. Engineering advances in metallurgy, polymer science, ceramic technology, and manufacturing processes lead to improvements in the devices that orthopaedic surgeons use and thus to better performance in the fields of orthopaedics, especially arthroplasty, spine surgery, trauma, and arthroscopy.
Biomechanics and biomaterials as broad topics may seem overwhelming, but they are manageable when approached by reviewing applications directly related to clinical practice. The present review covers basic aspects of biomaterials relevant to clinical practice, including their mechanical properties and uses for implants, and it outlines the clinically relevant aspects, principles, and facts that are germane to many surgical decisions.
The surgeon's understanding of orthopaedic implant technology is enhanced by a basic knowledge of mechanics. Once one has an understanding of how mechanical properties of materials are determined, one has a better appreciation of why certain materials in orthopaedics are utilized rather than others.
The first concept to emphasize is the stress-strain diagram, a useful graphical tool for illustrating the static behavior of systems that are elastic, viscoelastic, and plastic. The full meaning of these terms is clarified in the following sections.
Stress is defined as the applied force per unit cross-sectional area of the test piece (newtons per square millimeter [N/mm2]). Strain is defined as the increase in length (in millimeters) as a fraction of the original length (in millimeters)1. A servohydraulic materials testing machine allows for either load or displacement control to test the mechanical integrity of a specimen or device. A standard materials test piece uses a small volume with a constant cross-sectional area loaded to failure according to protocols that are standardized by ASTM International and the International Organization for Standardization (ISO). With use of these testing protocols, the stress-strain curve can be plotted and multiple material properties of the specimen can be calculated. A representative stress-strain diagram is shown in Figure 1.
Several features of the stress-strain diagram in Figure 1 are important and define the mechanical properties of the material or device being tested. These features are the linear elastic region, yield point, plastic region, ultimate strength, and failure. As a convention, the properties of materials are most often illustrated with use of a longitudinal force and measurement of the material's tensile strength. A similar curve can be constructed for materials subjected to other forces such as compression or shear.
Linear Elastic Region
In the linear region of the stress-strain diagram, the test piece behaves as a simple spring. When the stress is increased, the strain increases proportionally2. If the same amount of stress is let off, the strain decreases to the previous length. No permanent deformation of the test piece occurs. The strain may be altered any number of times with the same results. The slope of the linear region equals the modulus of elasticity (or Young's modulus) of the material. On the stress-strain curve, stiffer materials have greater slope on the linear portion of the curve.
Yield Point
The yield point is the stress at which there is a change from elastic to plastic deformation. Graphically, on the stress-strain curve, the yield point occurs at the transition of a straight line with constant slope to a curved line of variable slope. On the stress-strain diagram, the yield point is not always visually apparent as it is in Figure 1. Consequently, a stress resulting in a 0.2% change in strain is conventionally chosen as the numerical definition of the yield point.
Plastic Region
In the plastic region, when the stress is increased the strain increases in a more complex way than it does in the elastic region. The stress-strain curve may decrease for a small interval or it may continue to increase but at a lower or more variable rate relative to that in the linear elastic region.
The essential feature of plastic deformation is that it is not completely reversible. If the stress is let off, the test piece will not return to its original length. The molecular mechanisms that mediate plastic deformation are complex. Plastic deformation may result in a phenomenon known as work hardening or strain hardening.
Ultimate Strength
The ultimate strength is the maximum stress that a material can withstand before impending failure. The ultimate strength is not as important for orthopaedic implants as it is in other settings. For orthopaedic implants, fatigue strength is more important and is not necessarily related to ultimate strength.
Failure
Failure occurs when the test piece or material fractures. Numerous modes of failure are possible3. Ductile materials have a process of impending failure that occurs immediately after the stress surpasses the material's ultimate strength. Brittle materials are the opposite of ductile materials. Very brittle materials, such as some ceramics, fail in the linear elastic region, or after a very small amount of plastic deformation. Brittle materials have a simple stress-strain diagram, and their yield strength, ultimate strength, and failure strength are the same.
Modulus of Elasticity
The modulus of elasticity is the stress per unit strain in the linear elastic portion of the stress-strain curve. Young's modulus is the modulus of elasticity for tensile strength (measured with use of a longitudinal force). As a rule, the stress-strain diagram is for tensile testing (demonstrating Young's modulus) unless stated otherwise. The units for Young's modulus are megapascals (MPa)2.
Strength
The yield strength is the stress in MPa at the yield point on the stress-strain curve. This is the strength at the end of linear elastic behavior and at the onset of plastic deformation4. The ultimate strength is the stress at the apex of the stress-strain curve. This is the strength at the end of the plastic deformation portion of the curve if it is higher than the yield point. This strength denotes the end of work hardening and the beginning of necking, which precedes fracture. (See Figure 1 for graphical illustrations.)
Failure strength is defined as the point of fracture, beyond the linear elastic region and after plastic deformation. In principle, one can measure the strength at the failure or fracture point. In practice, there is rarely a distinction between the ultimate strength and the strength at failure. Since necking has begun, the failure of the test piece is impending with further strain after the ultimate strength. Therefore, the ultimate strength is often reported as the final strength measurement of a material.
Fatigue Strength
Fatigue strength is defined as the maximum stress at which a material can withstand ten million cyclic loading cycles without failure. (The number ten million is arbitrary but widely used.) In the stress-strain diagram, the test piece is subjected to a static load instead of a repetitive load. Incremental increases in loading result in elastic deformation, followed by plastic deformation and eventual failure. In fatigue failure, repetitive cyclic loading below the yield strength produces failure after numerous cycles.
Hardness
The most common measure of hardness of orthopaedic implants involves indentation of the material by a small indenter made of a very hard material (such as diamond)2. The Rockwell C scale is based on a test using a small diamond indenter with nearly 1500 N of force. The scratch resistance of materials is a distinct but related concept of hardness that is important for articulating total joint components, although it is less commonly reported as a materials property.
Toughness
Toughness is defined as the amount of energy (per unit volume) that a material can absorb up to the failure strength. Intuitively, toughness is a measure of the fracture resistance of a material when it is subjected to stress. The units are joules per cubic meter (J/m3).
Roughness
Roughness is a measurement of the surface finish of a test piece. As such, it is not a property of the material alone; it is also a property of the manufacturing processes used to create the surface finish. Ra is the average deviation of the peaks and valleys on a microscopic level (in micrometers or micro-inches). Ra is the most common roughness measure used for orthopaedic implant surfaces5. Roughness may impact the ultimate and fatigue strength material properties of an implant. The greater the roughness of the material's surface, the higher the stress concentrations at the valleys of the surface. Crack initiation and propagation in a device can originate in the valleys of the surface, and roughness has an impact on the longevity of a device that is loaded for millions of cycles during its lifetime.
Stress-strain diagrams are usually an illustration of static analysis, but materials can also be studied with dynamic analysis. Static analysis reveals the properties of materials independent of time. For example, a material in the linear region of a stress-strain diagram is held at an initial stress, with some deformation (strain). If the stress is increased a small amount and then held steady, a new strain is measured without taking into account what occurs during the time period when the testing machine is adjusting the applied stress. When measurements are made during changes that are applied very slowly, or when a substantial time period elapses between changes and remeasurement, the behavior of the test piece and material appears static.
In reality, no practical situation is truly independent of time (static). The alternative to static analysis is dynamic analysis, in which the time-dependent behavior at all time points is considered. If measurements are made when large changes in stress are applied, when changes are made very quickly, or when very little time has elapsed between changes and remeasurement, the behavior of the test piece and material will appear time-dependent. A dynamic system that reaches steady-state equilibrium with respect to applied changes appears static if measurements are made only after equilibration.
Linear Elastic Behavior
Assessment of linear elasticity is the simplest type of static analysis and is useful for materials in the linear elastic region of the stress-strain diagram undergoing slow changes in stress, especially metals and alloys under static loading2. In a linear elastic system, the measured strain is directly proportional to the applied stress at the jig. The constant of proportionality, or the slope of the stress-strain diagram, is called Young's modulus for tensile testing. There are several synonymous terms for linear elasticity, such as Hooke's law and simple spring.
The input-output diagrams for a linear elastic system are illustrated in Figure 2. When the input load changes instantaneously in the testing of these kinds of materials, the output also changes instantaneously. This is an idealization: the change only appears instantaneous at an appropriately coarse time scale. If the time scale is finer, the system may appear dynamic and viscoelastic.
Linear Viscoelastic Behavior
Linear viscoelasticity is another relatively simple type of dynamic analysis. It is useful for materials in the linear elastic region of the stress-strain diagram undergoing rapid changes in stress6, especially amorphous polymers under dynamic loading. In a linear viscoelastic system, the measured strain is directly proportional to the applied stress at the jig at final equilibrium but approaches the new steady state exponentially.
Time-dynamic behavior of viscoelastic systems can be described by the terms stress relaxation and creep.Figure 2 illustrates creep, in which deformation takes some time to reach a final value after application of a new load. A complementary view of the same phenomenon is stress relaxation, in which the measured stress gradually approaches some new final value when the test piece is suddenly adjusted to a new length.
Other Behavior
An important exception to viscoelastic behavior is plastic behavior beyond the yield point, the point of transition from elastic to plastic deformation, in the stress-strain diagram. In plastic deformation, changes in length are not reversible with changes in load. The test piece is permanently deformed, and neither static nor dynamic elastic models describe the behavior beyond the yield point. Many biological tissues exhibit nonlinear viscoelastic behavior that can be extremely complex6. Nonlinear viscoelasticity is dynamic behavior that is not described by linear time-invariant differential equations. Cartilage has complex nonlinear viscoelastic behavior. When a new load is applied slowly, such as with slow walking, the modulus of elasticity is lower, the deformation is increased, and the time to final deformation is increased. When a new load is applied quickly with a sudden weight-bearing impact, the modulus of elasticity is higher, the deformation is decreased, and the time to final deformation is decreased. Although important for biological tissues and some polymers, nonlinear viscoelasticity does not need to be taken into account for most engineering purposes involving materials.
Continuum Mechanics
Continuum mechanics is used in the study of elastic, viscoelastic, and plastic behaviors of structures containing a variety of materials with different properties, often with complex geometry. Although the principles of complex systems are the same as those of simple test geometries of a single material, the practical situation is much more complex. Computers are used to solve large systems of equations in order to simulate the mechanics of complex designs (e.g., finite element analysis)2.
Some features of continuum mechanics are important clinically. For example, stress-shielding can result when a mismatch in the elastic modulus occurs between two adjacent materials, such as a relatively stiff femoral stem and the cortical/cancellous bone in the proximal part of the femur. The proximal femoral region experiences a relative decrease in stress due to the stiff implant carrying stress to the cortical bone of the femoral diaphysis. This may result in relative osteopenia of the stress-shielded bone visible on a radiograph7.
The corrosion resistance of an implant has several components. The material and surface finish combine with the geometry of the test piece interacting with the biological environment. Corrosion varies widely among orthopaedic implants, with some polymer implants, such as bioabsorbable implants, being designed to degrade completely and other implants being designed to resist corrosion as completely as possible8.
Chemical Corrosion
Chemical corrosion is a process by which there is a chemical reaction of a material with the biological environment. The chemical reaction results in new compounds at the surface of the implant, which can change the mechanical properties. Metals often undergo oxidation, which may be catalyzed by other chemicals in the biological milieu, especially halide ions such as chloride (Cl-). Polymers undergo a variety of corrosive and degradative processes, including hydrolysis of the polymer bonds into shorter-chain polymers by thermal and enzymatic processes. The process of chemical corrosion can be altered by alterations in the chemistry of the biological environment, such as the decrease in pH that may be present with bacterial infection.
Crevice and Pitting Corrosion
These two distinct but related processes occur when local chemistry established in a small feature of the test piece causes increased corrosion in a small area. Crevice corrosion occurs when a small machined feature on the surface traps the local chemical environment and results in increased corrosion. Pitting corrosion is caused by a small pit, which can exist even on a smooth surface, with an altered chemical environment due to an impurity resulting from statistical variation. This local pit sets up a catalytic process that leads to a larger pit with even more altered chemistry. Even a tiny pit can ultimately lead to catastrophic failure.
Galvanic Corrosion
Galvanic corrosion is due to differences in electrochemical potential between two distinct metals in an electrolyte solution, creating an electric current between the two metals. This battery effect results in physical migration of metal ions. There is concern that galvanic corrosion in mixed-metal implants (especially cobalt-chromium and stainless-steel combinations) can change the geometry of the implant and weaken it over time. The evidence that this phenomenon represents a clinically relevant problem with commonly used metal combinations is limited9. Nevertheless, informed decisions that include an understanding of galvanic and other corrosion modes should be made regarding the use of mixed-metal implants.
Fretting
Fretting is a chemical and mechanical process in which corrosion occurs between two mating surfaces in micromotion. It includes both the chemical corrosion of freshly exposed surfaces and wear due to mechanical friction of corrosion products. Fretting is an important issue with articulating components of total joint arthroplasties, such as tibial trays articulating with polyethylene inserts in total knee arthroplasty, which can produce polyethylene debris affecting the tibial side10.
Corrosion and Biocompatibility
Ti-6Al-4V alloy is highly resistant to corrosion in biological environments. Commercially pure titanium has a high affinity for oxygen; however, Ti-6Al-4V forms a stable, adherent oxide layer in aqueous environments in a process termed self-passivation. The resultant film stabilizes the alloy to resist further oxidation and chemical corrosion, including that due to halide ions such as chloride. Passivation is a key to the corrosion resistance of Ti-6AL-4V8. Materials that resist corrosion by passivation are subject to pitting corrosion and crevice corrosion, in which an altered local chemistry in small pits or crevices inhibits the passivation layer, leading to progressive corrosion. This can be influenced by small scratches, pits, or design features and contributes to notch sensitivity induced by intraoperative handling and metal-metal interfaces.
Ti-6Al-4V is regarded as a highly biocompatible alloy. There is virtually no nickel in this alloy, so it is useful in patients with documented nickel sensitivity11. Corrosion modes, resistance, and biocompatibility are distinct for all orthopaedic metals and biomaterials, including 316L stainless steel, Co-Cr-Mo alloy, highly cross-linked ultra-high molecular weight polyethylene, poly-l-lactic acid polymer, and other commonly used materials. However, the considerations listed here for titanium alloys are representative of the considerations and reasoning for all materials11.
ASTM International is an international standards organization responsible for standardization of materials in industrial use. Previously known as the American Society for Testing and Materials (ASTM), the organization now has a global scope and charter. The existence of standards allows for trade of materials with known engineering performance among international vendors12. The society for medical devices is designated F-04, and the breadth and number of standards applicable to the field of orthopaedic surgery are astounding. The following are a few examples of standards used for titanium alloys commonly used in orthopaedic surgery today.
ASTM F-136: This alloy is Ti-6Al-4V and is used in many orthopaedic applications in North America13.
ASTM F-67: This alloy is commercially pure (CP) titanium and is used for spinal rods, plasma-spray coatings on Ti-6Al-4V implants, and some other applications.
ASTM F-1295: This alloy contains titanium, aluminum, and niobium and has been used for applications similar to ASTM F-136 in the European and North American markets.
ASTM F-2063: This alloy is commonly called Nitinol and contains nickel and titanium in nearly equal ratios. It has the peculiar property of being a shape memory alloy (SMA), in which heating after deformation restores the material to its prior shape. It has found wide application for intraluminal stents and arthroscopic suture passers, but only scant application for orthopaedic implants.
With an understanding of how mechanical properties of materials are determined, one has a better appreciation of why certain materials in orthopaedics are utilized over others. Comparing these properties (Table I) to bone allows a better comparison of the material's properties and use in orthopaedics. The more similar a material's mechanical properties are to bone, the less stress shielding is created. For example, if the properties of a femoral component in a total hip replacement are more similar to bone, there will be less stress shielding of proximal femoral bone and a lower stress riser at the tip of the femoral stem, thereby decreasing the risk of periprosthetic fracture. This is just one of multiple examples of how a better understanding of biomechanics and material properties enhances a surgeon's ability to make an educated decision concerning the products utilized in the treatment of his or her patients.
The terminology and principles involved in mechanics, materials, and engineering for orthopaedic surgery are manageable. Having sources to review to better their understanding of applications to orthopaedic surgery is paramount to enable surgeons to better treat their patients and comprehend the rationale behind many surgical procedures. Our hope is that this review can serve as a source to aid orthopaedic surgeons in obtaining this goal.
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