**2.4 Material properties**

168 Recent Advances in Arthroplasty

This type of wear occurs by the combination of mechanical wear and chemical reaction.

Fatigue acts during repeated sliding or rolling over the same wear track. The repeated loading and unloading can induce the initiation and propagation of microcracks parallel and orthogonal to the surfaces for mechanical or material-related reasons. As a result,

The wear mode defines the general mechanical conditions under which the bearing is functioning when wear occurs. Wear modes are defined by two sets of criteria: first by the macroscopic structure of the tribosystem and the kinematic interaction of its elements and second by the combination of acting wear mechanisms. It should be noted that the wear mode is not a steady-state condition and can change from one form to another. For example, particulate debris generated by two-body abrasion may function as an interfacial medium and turn the problem into a particle-related (third-body) phenomenon. Depending on the circumstances, this may reduce or increase the wear rate. Mode 1 wear results from the motion of two primary bearing surfaces one against each other, as intended. Mode 2 refers to the condition of a primary bearing surface moving against a secondary surface, which is not intended. Usually, this mode of wear occurs after excessive wear in mode 1. Mode 3 refers to the condition of the primary surfaces moving against each other, but with thirdbody particles interposed. In mode 3, the contaminant particles directly abrade one or both of the primary bearing surfaces. This is known as three-body abrasion or three-body wear. The primary bearing surfaces may be transiently or permanently roughened by this interaction, leading to a higher mode 1 wear rate. Mode 4 wear refers to two secondary (nonprimary) surfaces rubbing together. Examples of mode 4 wear include wear due to metal-cement or bone-cement interface motion or from relative motion of a porous coating, or other metallic surface, against bone; relative motion of the superior surface of a modular

Fig. 3. Examples of abrasive wear

**2.3.1.3 Corrosive wear** 

**2.3.1.4 Surface fatigue wear** 

shallow pits and filaments (delaminations) are generated.

(Howcroft 2008)

**2.3.2 Wear mode** 

Materials used in the manufacturing of femoral heads for metal-on-polyethylene (MOP) total hip replacement (THR) include the metal alloys, stainless steel, cobalt-chromium, and titanium alloy as well as ceramic materials, aluminum oxide, and zirconium oxide. Properties to consider when evaluating materials for bearings in THR include corrosion resistance, strength, ductility, hardness, and frictional characteristics. Frictional characteristics are a result of material properties such as wettability (related to surface energy), manufacturing variables such as surface finish, and operating conditions such as lubrication. The degree of resistance is proportional to the load. Because both chemical and mechanical interactions may occur, frictional forces depend on both the material composition and the roughness of the opposed surfaces. Lubricating conditions can change the nature of the interface between the moving surfaces and decrease friction. As described earlier, the coefficients of friction depend upon the nature and amount of lubricant present, as well as the speed of relative motion and the applied load.

### **2.5 Surface roughness**

A thorough surface roughness evaluation should include a visual comparison of the actual tracings and representative photomicrographs (scanning electron microscope) of the surface. The surface roughness of currently available femoral heads ranges from an average height (Ra) of less than 0.03 µm to about 0.10 µm and maximum height ranging from less than 0.10 to about 0.40 µm. The surface roughness of a femoral bearing can change over time in vivo. In the presence of hard third bodies, as can occur in vivo, surface abrasions (scratches) result

The Bearing Surfaces in Total Hip Arthroplasty – Options, Material Characteristics and Selection 171

cemented acetabular component and lower than that reported for surface replacement components. Contrary to theoretical considerations, frictional torque has not been demonstrated to be important in the initiation of aseptic loosening of either femoral or acetabular components. Accumulating evidence indicates that PE wear particles have a greater effect on the durability of implant fixation than frictional torque. From this perspective, the success of the Charnley low-friction arthroplasty is primarily a function of the low volumetric wear of the 22-mm bearing, not low frictional torque. Higher bearing surface friction and frictional torques can be tolerated if the release of wear particles to periprosthetic tissues is sufficiently low. Within the range of frictional torques generated by implants used to date, wear is a more important factor in survivorship than frictional torque. Large-diameter bearings can be successful if the wear rate is low. This is an

When contact occurs between metal and polyethylene components, both surfaces deform, but the deformation of the metal component is negligible and the metal component behaves like a rigid indenter. Thus, when an artificial joint is loaded, the polyethylene is squeezed between the rigid metal component and the supporting material (bone, cement, or metal backing), and in the region of contact, the articulating surface of the polyethylene is forced to conform to the shape of the metal surface. The resulting deformation causes compressive, tensile, and shear stresses in the polyethylene. The magnitude of the stress depends on the magnitude of the joint load. These loads are large (3-5 times body weight) and in combination with the relative motion of the articulating surfaces, cause damage that increases with time of implantation (number of cycles of loading) and patient weight (magnitude). This provides strong evidence that surface damage in total joint replacements is the result of fatigue processes. The stresses associated with damage to the articulating surfaces occur both at the surface and within the polyethylene component. Two types of stresses can be applied to the surface: normal compressive stresses (contact stresses), and tangential shear stresses due to friction. Metal and polyethylene were originally chosen for bearing surface materials to produce low-friction total joint replacements. The stresses acting on the surface produce normal and shear stresses within the polyethylene. At the surface, the largest compressive stresses are the contact stresses that act perpendicular to the surface**.** They decrease nonlinearly with depth through the thickness of the polyethylene. Joint contact also produces compressive and tensile stresses within the polyethylene component that act tangent to the articulating surface. Tangential compressive stresses occur because the polyethylene under the center of the contact area expands radially as the component is compressed. This expansion is resisted by the surrounding material, and tangential compressive stresses are produced. Tangential tensile stresses near the articulating surface occur because the surface must stretch as the polyethylene conforms to the shape of the metal component when the joint is loaded. The stretching occurs near the edge of the contact area. The resulting tensile stresses are largest at the surface of the component. Surface damage is most likely due to combinations of stress components. The maximum shear stress occurs at the surface for conforming joints like hip joints. Although artificial joints produce a range of particle sizes, conforming joints such as hip replacements produce a number of smaller particles greater than the number of larger particles produced

important consideration as alternatives to PE bearings are being investigated.

**2.7 Stresses caused by contact** 

by nonconforming total knee replacements.

in an increased surface roughness, and the wear rate of polyethylene (PE) can increase. Conversely, in "clean" operating conditions with little or no hard third bodies, motion against PE may result in polishing of the metal surface and a lower surface roughness. This suggests that differences in the initial surface roughness of a femoral head may not be as important in the long run as the in vivo operating conditions. The susceptibility to scratching is a function of the hardness of the material. The decreased hardness of titanium alloy results in decreased abrasion resistance. Although the initial surface roughness of a titanium alloy femoral head may be equivalent to that of other bearing materials, there is greater potential for surface roughness to increase in vivo. In an environment with few or no hard third bodies, the wear performance of titanium alloy against PE can be comparable to the other metals, but the performance of titanium alloy against PE is affected to a greater degree by the presence of hard third bodies. The abrasion resistance of cobalt-chromium alloy is considered superior to that of titanium alloy and stainless steel.(Fisher 2004)

The abrasive particles are mainly released from modular interfaces, metal backing, and/or porous coating. Increased femoral head surface roughness may dramatically accelerate twobody abrasive wear of PE. Experimental studies indicate that a threefold increase in femoral roughness can cause at least a tenfold increase in the wear rate of PE. Specific increases in wear rate are dependent on the nature of the damage to the femoral head. Ceramics are harder and therefore more resistant to damage by third-body particles than metal counterfaces. For this reason, the increased hardness of ceramic materials is considered advantageous. The pattern of damage by a hard third body in metals and ceramics differs.(Raimondi 2008) (Fig.5)

Fig. 5. Scratch profiles of metal and ceramic femoral heads

#### **2.6 Frictional torque**

The frictional torque of the Charnley prosthesis with a load of 890 N has been reported between 0.4 to 1.2 Nm. Under similar conditions, the frictional torque of a 28-mm prosthesis, a 43-mm prosthesis, and a 51-mm prosthesis averaged 1.3, 2.7, and 3.2 Nm, respectively. These values are 20 to 100 times smaller than the reported static torques to failure for cemented acetabular components. The coefficient of friction for the MOM bearing of the McKee-Farrar hip is roughly two to three times greater than that for the Charnley. The larger diameter of the McKee-Farrar (about 40 mm) amplifies this difference, and the result is a frictional torque that is up to 10 times greater than that in the Charnley. This value is still an order of magnitude less than the static torque-to-failure of an acutely implanted cemented acetabular component and lower than that reported for surface replacement components. Contrary to theoretical considerations, frictional torque has not been demonstrated to be important in the initiation of aseptic loosening of either femoral or acetabular components. Accumulating evidence indicates that PE wear particles have a greater effect on the durability of implant fixation than frictional torque. From this perspective, the success of the Charnley low-friction arthroplasty is primarily a function of the low volumetric wear of the 22-mm bearing, not low frictional torque. Higher bearing surface friction and frictional torques can be tolerated if the release of wear particles to periprosthetic tissues is sufficiently low. Within the range of frictional torques generated by implants used to date, wear is a more important factor in survivorship than frictional torque. Large-diameter bearings can be successful if the wear rate is low. This is an important consideration as alternatives to PE bearings are being investigated.

### **2.7 Stresses caused by contact**

170 Recent Advances in Arthroplasty

in an increased surface roughness, and the wear rate of polyethylene (PE) can increase. Conversely, in "clean" operating conditions with little or no hard third bodies, motion against PE may result in polishing of the metal surface and a lower surface roughness. This suggests that differences in the initial surface roughness of a femoral head may not be as important in the long run as the in vivo operating conditions. The susceptibility to scratching is a function of the hardness of the material. The decreased hardness of titanium alloy results in decreased abrasion resistance. Although the initial surface roughness of a titanium alloy femoral head may be equivalent to that of other bearing materials, there is greater potential for surface roughness to increase in vivo. In an environment with few or no hard third bodies, the wear performance of titanium alloy against PE can be comparable to the other metals, but the performance of titanium alloy against PE is affected to a greater degree by the presence of hard third bodies. The abrasion resistance of cobalt-chromium

alloy is considered superior to that of titanium alloy and stainless steel.(Fisher 2004)

differs.(Raimondi 2008) (Fig.5)

**2.6 Frictional torque** 

Fig. 5. Scratch profiles of metal and ceramic femoral heads

The abrasive particles are mainly released from modular interfaces, metal backing, and/or porous coating. Increased femoral head surface roughness may dramatically accelerate twobody abrasive wear of PE. Experimental studies indicate that a threefold increase in femoral roughness can cause at least a tenfold increase in the wear rate of PE. Specific increases in wear rate are dependent on the nature of the damage to the femoral head. Ceramics are harder and therefore more resistant to damage by third-body particles than metal counterfaces. For this reason, the increased hardness of ceramic materials is considered advantageous. The pattern of damage by a hard third body in metals and ceramics

The frictional torque of the Charnley prosthesis with a load of 890 N has been reported between 0.4 to 1.2 Nm. Under similar conditions, the frictional torque of a 28-mm prosthesis, a 43-mm prosthesis, and a 51-mm prosthesis averaged 1.3, 2.7, and 3.2 Nm, respectively. These values are 20 to 100 times smaller than the reported static torques to failure for cemented acetabular components. The coefficient of friction for the MOM bearing of the McKee-Farrar hip is roughly two to three times greater than that for the Charnley. The larger diameter of the McKee-Farrar (about 40 mm) amplifies this difference, and the result is a frictional torque that is up to 10 times greater than that in the Charnley. This value is still an order of magnitude less than the static torque-to-failure of an acutely implanted When contact occurs between metal and polyethylene components, both surfaces deform, but the deformation of the metal component is negligible and the metal component behaves like a rigid indenter. Thus, when an artificial joint is loaded, the polyethylene is squeezed between the rigid metal component and the supporting material (bone, cement, or metal backing), and in the region of contact, the articulating surface of the polyethylene is forced to conform to the shape of the metal surface. The resulting deformation causes compressive, tensile, and shear stresses in the polyethylene. The magnitude of the stress depends on the magnitude of the joint load. These loads are large (3-5 times body weight) and in combination with the relative motion of the articulating surfaces, cause damage that increases with time of implantation (number of cycles of loading) and patient weight (magnitude). This provides strong evidence that surface damage in total joint replacements is the result of fatigue processes. The stresses associated with damage to the articulating surfaces occur both at the surface and within the polyethylene component. Two types of stresses can be applied to the surface: normal compressive stresses (contact stresses), and tangential shear stresses due to friction. Metal and polyethylene were originally chosen for bearing surface materials to produce low-friction total joint replacements. The stresses acting on the surface produce normal and shear stresses within the polyethylene. At the surface, the largest compressive stresses are the contact stresses that act perpendicular to the surface**.** They decrease nonlinearly with depth through the thickness of the polyethylene. Joint contact also produces compressive and tensile stresses within the polyethylene component that act tangent to the articulating surface. Tangential compressive stresses occur because the polyethylene under the center of the contact area expands radially as the component is compressed. This expansion is resisted by the surrounding material, and tangential compressive stresses are produced. Tangential tensile stresses near the articulating surface occur because the surface must stretch as the polyethylene conforms to the shape of the metal component when the joint is loaded. The stretching occurs near the edge of the contact area. The resulting tensile stresses are largest at the surface of the component. Surface damage is most likely due to combinations of stress components. The maximum shear stress occurs at the surface for conforming joints like hip joints. Although artificial joints produce a range of particle sizes, conforming joints such as hip replacements produce a number of smaller particles greater than the number of larger particles produced by nonconforming total knee replacements.

The Bearing Surfaces in Total Hip Arthroplasty – Options, Material Characteristics and Selection 173

where Δ is the displacement of the rod, E is the elastic modulus of the rod material, A is the cross-sectional area of the rod, and L is the length of the rod, which corresponds to the thickness of the polyethylene component. It can be seen from this equation that the structural stiffness of the rod increases as the length of the rod decreases. In a similar way, the structural stiffness of polyethylene components increases with decreasing thickness. When the stiffness of the component increases, the indenter does not displace as much, the contact area decreases, and the contact stresses increase. The rod analogy further shows that structural stiffness also increases when the elastic modulus of the rod material increases. Similarly, the structural stiffness of a polyethylene component increases when the elastic modulus of the polyethylene increases, leading to a decrease in contact area and an increase in contact stress. The contact stresses for acetabular components are within the elastic limit of the polyethylene. For perfectly conforming spherical contact, it can be shown that the contact stresses are independent of cup thickness and modulus. This is consistent with observations of in vivo head penetration rates, which are constant with time after an initial wearing-in period. On the basis of contact stress alone, larger head sizes would be preferred. But it has been shown that wear at a point is proportional to pressure times sliding distance. Consequently, there is a trade-off between the relative sliding of the contacting surfaces and the contact stress, both of which are a function of diameter. The maximum shear stress in polyethylene acetabular components occurs very close to the articulating surface and this fact has not yet been directly linked to the wear seen in these components.(Teoh 2002)

The generation of particulate debris is a central focus of attention in the arthroplasty literature. The biologic response to wear debris is currently heralded as the single most important factor limiting the long-term durability of contemporary total hip and total knee

Current data suggest that tissues adjacent to a failed joint prosthesis contain billions of particles per gram of tissue. A wide variety of particle types have been retrieved from periprosthetic tissues at the time of autopsy or revision, as well as from joint simulators. In general, these particles types may be classified as metallic, polymeric, and ceramic. The majority of reports on periprosthetic metal debris pertain to Co-Cr and titanium. Metallic particles are characteristically gray to black, and although they may appear weakly birefringent under polarized microscopy, the appearance of birefringence is an optical artifact because the particles are actually opaque. The metal particles are generally smaller

Submicrometer metallic particles have been described as globular or irregularly shaped as well as elongated with sharp corners. Larger Co-Cr particles, in the 1- to 5-µm size range, have been described more often as needle, rod, or splinter shaped. Even larger Co-Cr particles, from 5 µm to over 1 mm, have been described as irregularly shaped or globular. These particles tend to be extracellular and may actually represent aggregates or clusters of smaller particles. The majority of titanium particles likewise range from under a micrometer to less than 5 µm in size and have been described as blackish-gray material and fine powder. Occasional titanium particles from 5 µm up to 1 mm in size have been

**2.9 Particle debris** 

replacement arthroplasty.

**2.9.1 Types of particles** 

reported. (Malviya 2010)

than polymer particles but larger than ceramic particles.

### **2.8 Reducing surface damage**

By minimizing the stresses associated with the types of damage that can occur or by increasing the strength of the polyethylene, the risk of surface damage can be decreased. We know that abrasive wear which occurs in hip joints can be minimized by a reduction in contact stress. Interestingly, both the range of the maximum principal stress and the maximum shear stress generally decrease when contact stress is decreased. Therefore, one can reduce the risk of abrasive wear and pitting and delamination by reducing the contact stress. The overall design goal is to choose the geometry of the articulating surfaces and material properties of the polyethylene that minimize contact stress.

### **2.8.1 Minimizing contact stress**

Contact stresses in acetabular components are affected by changes in loading, conformity of the articulating surfaces, thickness of the polyethylene, and stiffness of the material. Contact stresses increase with increasing load. If the same prosthesis (same conformity, thickness, and material) is used in patients with different weights, the stresses will be higher in the heavier patients. The stresses are not directly proportional to the load. As the load between the contacting surfaces increases, the contact area also gets larger. The contact stresses, of course, are not uniform over the contact area. The metal component of a total joint replacement is a rigid indenter. The contact stress will be greatest where the surface displacement of the polyethylene is greatest. Consequently, the displacement of the polyethylene surface in a direction normal to the surface will be determined by the shapes of the two contacting surfaces. For example, if the indenter and the polyethylene are both spherical, as they are in an ideal total hip replacement, then the maximum displacement of the polyethylene will occur at the center of contact. Therefore, the maximum contact stress will occur at the center of the contact area, and the minimum contact stress (zero) will occur at the edge of contact. Furthermore, the shape of the contact area in this ideal case will be circular. If the surfaces are not spherical because of either design or manufacturing variations, then the maximum contact stress may not be at the center of the contact area. The acetabular surface has ripples in it. As a result, when the femoral head is pressed into the polyethylene, the largest displacement normal to the surface of the polyethylene component will be at the apexes of the ripples. Therefore, the greatest contact stresses will also occur at these points. The stress in the valleys in the contact area will be small, because the deformation (difference between the dashed and solid lines) will be small at these points. Surface waviness can be caused by normal variations in manufacturing processes. Changes in conformity, thickness, and material properties cause changes in the contact area. In general, changes that decrease the contact area will increase the stresses, because the same load must be distributed over a smaller region. The contact area decreases when the conformity between the articulating surfaces decreases, when the thickness of the material decreases, and when the stiffness of the material increases. The effects of changes in thickness and elastic modulus on contact stresses may be understood as follows. Conceptually, the polyethylene may be considered to be supporting the metal indenter by a collection of parallel rods that are aligned along the direction of loading. Each rod supports a portion, δP, of the total load P. The stiffness of a rod under axial load is given by

$$\mathbf{K}\_{\text{rad}} = \text{SP}/\Delta = \text{EA}/\text{L} \tag{3}$$

where Δ is the displacement of the rod, E is the elastic modulus of the rod material, A is the cross-sectional area of the rod, and L is the length of the rod, which corresponds to the thickness of the polyethylene component. It can be seen from this equation that the structural stiffness of the rod increases as the length of the rod decreases. In a similar way, the structural stiffness of polyethylene components increases with decreasing thickness. When the stiffness of the component increases, the indenter does not displace as much, the contact area decreases, and the contact stresses increase. The rod analogy further shows that structural stiffness also increases when the elastic modulus of the rod material increases. Similarly, the structural stiffness of a polyethylene component increases when the elastic modulus of the polyethylene increases, leading to a decrease in contact area and an increase in contact stress. The contact stresses for acetabular components are within the elastic limit of the polyethylene. For perfectly conforming spherical contact, it can be shown that the contact stresses are independent of cup thickness and modulus. This is consistent with observations of in vivo head penetration rates, which are constant with time after an initial wearing-in period. On the basis of contact stress alone, larger head sizes would be preferred. But it has been shown that wear at a point is proportional to pressure times sliding distance. Consequently, there is a trade-off between the relative sliding of the contacting surfaces and the contact stress, both of which are a function of diameter. The maximum shear stress in polyethylene acetabular components occurs very close to the articulating surface and this fact has not yet been directly linked to the wear seen in these components.(Teoh 2002)
