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Dr. Brown is Richard and Janice Johnston Chair of Orthopaedic Biomechanics, and Professor, Department of Orthopaedics and Rehabilitation, University of Iowa, Iowa City, IA. Dr. Bartel is Willis H. Carrier Professor in Engineering Emeritus, Cornell University, Ithaca, NY, and Senior Scientist, Hospital for Special Surgery, New York, NY.
*The Implant Wear Symposium 2007 Engineering Work Group included Donald L. Bartel, PhD, Thomas D. Brown, PhD, Ian C. Clarke, PhD, Roy D. Crowninshield, PhD, Darryl D'Lima, MD, PhD, A. Seth Greenwald, DPhil(Oxon), Steven M. Kurtz, PhD, Jack Lemons, PhD, Michael T. Manley, PhD, Harry A. McKellop, PhD, Orhun K. Muratoglu, PhD, Ebru Oral, PhD, Lisa Pruitt, PhD, Clare Rimnac, PhD, Peter S. Walker, PhD, and Timothy Wright, PhD.
Dr. Brown or a member of his immediate family has received research or institutional support from DePuy, Arthrosurface, and the National Institutes of Health and is a consultant to or an employee of Smith & Nephew. Dr. Bartel or a member of his immediate family has received research or institutional support and royalties from the Hospital for Special Surgery, holds stock in Exactech, and is a consultant to or employee of Stryker Orthopaedics.
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Aprincipal strategy for minimizing the osteolysis problem is to reduce its root cause—wear debris. Implant mechanical design is a key consideration in that regard. Mechanical design issues are closely coupled with bearing surface material combinations, the options for which have become ever more diverse. Metal-on-polyethylene bearing couples have been largely supplanted by metal-on–highly cross-linked polyethylene, ceramic-on-polyethylene, ceramic-on-ceramic, and newer-generation metal-on-metal bearings. These newer bearing surface combinations unquestionably far outperform metal-on–conventional polyethylene bearings in terms of volumetric wear rate,1 although various ancillary issues (eg, absolute particle numbers, relative particle osteolytic potency, metal ion release) currently confound interpretation of clinical efficacy.
The design of successful total joint arthroplasty involves both functional and structural goals. The joint arthroplasty must provide the normal range of motion while transmitting forces across the joint that are generally several times body weight. Forces and motions in natural and prosthetic joints are coupled—changing one will affect the other. Structurally, the bone-implant system is a composite. Therefore, designers must be concerned with the strength of the individual prosthetic components, damage to the articulating surfaces, the strength of the interfaces between the implant and the bone, the interfaces between components of the prosthesis (eg, a polyethylene insert to a metallic backing or the taper connection between a femoral head and stem), and the potential bone adaptation to altered loading. Thus, the reduction of wear and damage to bearing surfaces is only one aspect of bone-implant design and must be considered within the context of the composite system. Solutions that minimize wear and damage may limit function or place interfaces at risk.
A wealth of information exists regarding design factor influences on wear of metal-on–conventional polyethylene bearings, which make up the vast majority of implants currently in service. Much of this information is applicable to the newer bearing surfaces. For the abrasive and adhesive wear mechanisms widely held to engender the micron- or submicron-sized polyethylene particles responsible for osteolysis, particularly in total hip arthroplasty (THA), debris removal from the bearing surface depends on the combined interaction of the tribologic properties of the bearing surface couple (eg, materials, surface finish, friction/lubrication), the contact stress at the bearing surfaces, and joint kinematics. In combination, laboratory simulators and computational models have provided compelling evidence that volumetric wear increases in approximately linear proportion to head diameter, about 4% to 6% per millimeter of head size increase.2
Other studies have demonstrated that bearing surface wear is far greater than backside wear (at least for reasonably well-fixed liners),3 that thin polyethylene in combination with nonuniform support from the metallic backing leads to increased wear, and that the type of fixation resulting in well-fixed cups (cemented versus cementless) has little direct kinetic influence on bearing surface wear behavior.4 For articulations against polyethylene, initial curvature radius mismatches on the order of a few tenths of a millimeter (the current industrial tolerance range in THA) have minimal consequence in the long term, although they do increase wear rates for the first few hundred thousand cycles as the polyethylene surface is reshaped to match the head’s radius of curvature.5
The basic premise is that bearing surfaces are damaged or exhibit wear because stresses have exceeded the strength of the articulating surfaces. Thus, the basic question posed in Figure 1 is, Do the stresses exceed the strength of the material? The performance of the system is affected by implant, patient, and surgical factors that interact in complex ways. These factors may be further sorted into design variables that are under the control of the designers and environmental variables over which they have little or no control. The variables associated with the implant (eg, material, geometry, manufacturing processes) are design variables. The patient variables (eg, bone structure, weight, activity level) are clearly environmental variables, which are known at best only stochastically. Surgical factors include both design and environmental variables. For example, the desired component position and orientation with respect to the bone are design variables; the variations about a desired position or orientation due to the limited precision of a surgical procedure are environmental variables.
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In THAs, the contacting surfaces are nearly conforming; diametral clearances are on the order of 0.1 mm for contemporary designs. Elasticity theory and finite-element models show that if perfect conformity is achieved, the maximum contact stresses are independent of the elastic modulus of the polyethylene. The effect of thickness on maximum contact stresses depends on whether the head diameter or the inner diameter of the shell is being changed. When the head size is held fixed and the inner diameter of the shell is increased, the stresses are independent of thickness. If the shell is fixed and the head size is increased, the stresses are reduced because the applied load is distributed over a greater projected area (Figure 2). As noted earlier, increasing the head size also reduces the linear wear rate. However, volumetric wear increases2 (Figure 3).
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In TKAs with fixed bearings, the maximum contact stresses are a function of the conformity between the femoral and tibial components. Most designs fall somewhere between toroidal and flat contact in the lateral-medial direction. In designs in which the contacting surfaces are toroidal (or may be approximated by toroidal surfaces), the maximum contact stress is most sensitive to the lateral-medial radius and relatively insensitive to the radius in the anterior-posterior plane. The maximum contact stresses are least when the contact is perfectly conforming in the lateral-medial direction. In this case, however, torsional laxity is greatly reduced, and torques about the tibial axis are transmitted primarily through the prosthesis instead of being shared with soft-tissue structures about the knee, a situation that may increase the risk of loosening. The maximum contact stresses increase with decreasing thickness and become increasingly sensitive to changes in thickness as the thickness is decreased.
In designs in which the contacting surfaces are flat or nearly so in the lateral-medial direction, the maximum contact stress is dependent on the radius at the edge of contact.9 If the radius is small, the condyle acts like a punch and produces high stresses in the polyethylene near the lateral-medial edges of contact. As the edge radius increases, these stresses can be reduced to acceptable levels, provided that both condyles remain in contact. If liftoff occurs and only one condyle is in contact, stresses may increase, and the center of contact will move closer to the edge of the component.
Substantial improvements have been made to polyethylene through improved sterilization and cross-linking processes. In addition, the strength and wear resistance of specific versions of polyethylene components can be quantified using relatively simple tests and material models.2,10 Consequently, the "strength" side of the question in Figure 1 (Does the stress exceed strength?) can now be quantified. As a result, design analyses can now include constraints on stress instead of making stress minimization the objective of the design process. For example, improved function in TKAs may result in higher contact stresses. The design optimization problem can then be formulated as follows: maximize function subject to upper bounds on the stresses associated with wear or fracture.
As suggested above, the general effects of design factors on wear and damage are well known from basic elasticity theory. Better estimates of stresses for specific designs and materials require detailed finite element analyses. These analyses are computationally expensive, which renders optimal design using typical iterative algorithms difficult if not impossible to accomplish. Statistically based algorithms for optimal design have been developed that make finite-element analysis–based, optimal design economically feasible.11-13 Because the underlying methodology is based on the mathematics of statistics, a natural platform is provided for including environmental variables that may be known only stochastically. This provides the opportunity to include patient and surgical environmental variables (Figure 1) in optimal design studies, to determine the relative influence of design and environmental variables, and to calibrate complex computational models with experimental studies in which uncertainty exists both in model parameters and experimental data.
By contrast, appropriate surface clearance is an absolutely crucial consideration for modern hard-on-hard bearings (metal-on-metal and ceramic-on-ceramic), for which successful function relies on fluid-film lubrication. Appropriate clearances for that purpose can be determined from formal engineering lubrication analysis and are on the order of 100 µm.14 For purposes of achieving fluid-film lubrication (thereby avoiding direct contact between microasperities on the respective bearing surfaces), a rule-of-thumb metric is the lambda parameter, defined as the ratio of the minimum film thickness (determined from lubrication analysis) to the characteristic surface roughness of the respective bearing surfaces. For typical conditions existing in THAs, maintaining suitably high lambda values (>3) requires highly polished surfaces (with roughnesses on the order of a few hundredths of a micron) and very small deviations (micron range) in spherical concentricity of the respective surfaces. Under such conditions, wear rates for hard-on-hard bearings actually decrease as nominal head diameter increases, the opposite of the relationship for abrasive/adhesive wear of metal-on-polyethylene bearings. Unfortunately, idealized tribologic behavior for hard-on-hard bearings can be severely compromised under conditions of component malalignment causing edge loading and/or if even a small amount (a few hundredths of a millimeter) of component geometric distortion exists— for example, from interference-fit deformation of very thin acetabular shells impacted into an underreamed bone bed.
Although most design-related analyses of implant wear address as-manufactured components operating under ideal laboratory conditions, increasing evidence exists to suggest that damage in vivo from third-body mechanisms is a "wild card" responsible for many of the outlier cases most predisposed to osteolysis. Between cohorts of otherwise similar constructs performed by the same surgeon, a statistically significant association exists between elevated third-body burden and decreased survivorship.15 Moreover, within specific cohorts, the distribution of wear rate among patients is distinctly nongaussian due to the presence of high-wear outliers.16
The list of design variables potentially contributing to elevated third-body particulate burden is extensive. Even the ideally pristine situation represented in a cemented monoblock design has several third-body sources—cement and cement radio-opacifier particles, bone particles, metal particles from fretting of trochanteric reattachment wires, and metal particles from burnishing of loose stems. Cementless fixation and modularity add yet more sources to the list: the femoral cement/metal interface, especially for matte and precoated stems; fretting debris from trunnion-tapered head-stem connections or from modular stem junctions; broken-off porous-coating fragments (either late or during impaction at insertion); hydroxyapatite particles; metal fragments from the liner-backing interface of metal-backed acetabular components; and fretting debris from screw attachment of cementless metal-backed cups. Moreover, there is a statistically significant association between third-body embedment in retrieved acetabular cups and impingement-associated rim damage.17 This association, corroborated by laboratory physical tests and by computational simulations, suggests that impingement/subluxation may facilitate access of third-body debris to wear-critical regions of otherwise closely conforming bearing surfaces.18
Thus, design factors contributing to impingement appear to be deleterious from a wear perspective, even in clinically "stable" cups. Furthermore, when frank dislocations occur, there is a strong possibility of head macrodamage from scraping contact with the shell, potentially exacerbated by attempts at closed reduction.19
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