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An In Vitro Model for Fluid Pressurization of Screw Holes in Metal-Backed Total Joint Components
The Journal of Arthroplasty Vol. 20 No. 7 2005
An In Vitro Model for Fluid Pressurization of Screw Holes in Metal-Backed Total Joint Components Steven M. Kurtz, PhD,*y Timothy P. Harrigan, ScD,* Michael Herr, MS,* and Michael T. Manley, PhDz
Abstract: Fluid pressure may stimulate osteolysis near screw holes in joint arthroplasty components. We developed a generalized in vitro model of a polyethylene liner and metal backing with a screw hole to investigate whether implant design factors influence local fluid pressure. We observed an order of magnitude of variation in the peak screw hole pressure (from 16.0 and 163 kPa) under clinically relevant loading conditions. Of the implant factors investigated, the surface finish of the metallic base plate had the greatest effect on peak screw hole fluid pressures; the thickness of the polyethylene liner, as well as the gap between the liner and the base plate, were also significant design variables. Our data suggest that unpolished metal base plates, thick polyethylene liners, and tight conformity between the liner and the metal base plate will all contribute to significantly reduced peak screw hole fluid pressures in joint arthroplasty. Key words: osteolysis, fluid pressure, implant design, conformity, thickness, surface finish. n 2005 Elsevier Inc. All rights reserved.
have been observed near the locations of uncovered screw holes [7,8]. The magnitude of backside wear, between the liner and shell, is orders of magnitude lower than the volume of wear generated between the femoral head and the liner [9]. If the focal pelvic lesions are indeed provoked solely by wear, the articulating surface is a more likely culprit for the source of the debris, which then migrates through the gap between the liner and the shell to attack bone through the screw holes [5]. An alternative, or complementary, hypothesis to the debris pumping scenario is that fluid pressure —either independently or in combination with the presence of wear debris —directly stimulates the production of focal lesions in the vicinity of screw holes [3,4]. In total-knee arthroplasty (TKA), tibial osteolysis has been reported in less than 5 years of implantation when screws or cementless fixation have been used [10-12]. Implant design factors, such as the configuration of porous coating on the tibial base plate, have been reported to influence the incidence of tibial osteolysis after cementless TKA [11], suggesting that fluid transport or fluid pressure influence tibial osteolysis. Furthermore, the
Periprosthetic osteolysis is recognized as a critical and limiting problem for total joint arthroplasties [1,2]. Although osteolysis is clearly linked to ultrahigh molecular weight polyethylene (hereafter, polyethylene) particulate debris, fluid pressure and fluid flow in the tissues around a total joint implant have also been shown to directly stimulate osteolysis in animal models [3,4]. Fluid flow may influence osteolysis by providing a mechanism for wear particles to reach the vulnerable periprosthetic bone [5,6]. Fluid-induced periprosthetic osteolysis is clinically relevant for both hip and knee arthroplasties. In total-hip arthroplasty (THA), focal acetabular lesions From *Exponent, Inc, Philadelphia, Pennsylvania; y Drexel University, Philadelphia, Pennsylvania; and z Ridgewood, New Jersey. Submitted April 28, 2004; accepted November 10, 2004. Benefits or funds were received in partial or total support of the research material described in this article from Stryker Orthopedics. Reprint requests: Steven M. Kurtz, PhD, Exponent, Inc, Suite 300, 3401 Market St, Philadelphia, PA 19104. n 2005 Elsevier Inc. All rights reserved. 0883-5403/05/1906-0004$30.00/0 doi:10.1016/j.arth.2004.11.003
932
In Vitro Model of Fluid Pressurization in Screw Holes ! Kurtz et al 933
intrusion of the UHMWPE tibial insert into vacant screw holes has also been identified clinically as a mechanism of backside wear [13]. Overall, the clinical observations to date would suggest that focal osteolytic lesions around screw holes is a general concern shared by metal backed tibial as well as acetabular components for joint arthroplasty. Although many previous studies have identified focal osteolysis around screw holes to be a clinical problem for both THA and TKA, comparatively limited progress has been made in developing experimental or analytical models to predict fluid pressure around joint arthroplasties. The fluid flow in the vicinity of a reconstructed joint is likely to be complex, influenced not only by the geometry and material properties of the prosthetic components, but also by the hard and soft tissues, their porosity, and their apposition to the implant interface where the metal shell and the tissue interdigitate. Given this complex fluid and structural system, this study takes a reductionist approach, where a limited number of the relevant mechanical features of the implant are studied. Based on these results, the role of other factors can be investigated by studies that are progressively more complex, giving progressively better approximations to the clinical situation. In this study, we tested the hypothesis that implant factors, such as liner thickness, screw hole geometry, and surface roughness, influence the fluid pressure within a screw hole of the supporting metal backing of a joint arthroplasty. To test this hypothesis we developed a generalized in vitro model of a polyethylene liner and a metal backing with a screw hole. This generalized in vitro model represents the first step toward better quantitative understanding
of fluid flow and pressure development around screw holes in total joint arthroplasties.
Materials and Methods Experimental Apparatus For this study we developed a generalized in vitro model of a flat polyethylene insert, centered over a screw hole in a flat metallic base plate. The insert was loaded superiorly by a flat, aligned platen (Fig. 1A and B). This in vitro model most closely replicates the scenario of a tibial insert loading a metallic tray. However, this model is also relevant to modular acetabular components, where the local geometry at the back side of the liner-shell interface has a relatively large radius of curvature. Compression molded, medical grade polyethylene (GUR 1050: Ticona, Bishop, Tex) was used to simulate inserts of different thicknesses. The polyethylene specimens were machined to the following dimensions: 31 mm length 31 mm width 5 or 10 mm thickness. The insert thicknesses of 5 or 10 mm were chosen to be consistent with commercially manufactured hip and knee components, and the length and width were chosen to produce, on average, 5-MPa contact pressure at a load of 4805 N. The target average pressure of 5 MPa for our in vitro model was chosen because the contact stress on the backside of an acetabular liner has been shown to range between 5 and 10 MPa under clinically relevant, gait loading conditions in previous finite element simulations of THA [14].
Fig. 1. In vitro screw hole pressure testing apparatus. A, Assembled apparatus, illustrating the PBS testing environment as well as the supporting fixtures. B, Partially disassembled apparatus, showing the generalized representation of the polyethylene insert and base plate with screw hole. The initial gap between the insert and base plate has been exaggerated for demonstration purposes (the actual gap distance in the experiments ranged between 0.01 and 0.2 mm).
934 The Journal of Arthroplasty Vol. 20 No. 7 October 2005 In our apparatus, an 8-mm-diameter screw hole (either cylindrical or tapered to accept a chamfered bone screw) was drilled into a flat, 4-mm-thick stainless steel base plate. The bottom of the screw hole was tapped to allow a pressure transducer to be mounted. In this manner, the sensing membrane of the transducer was at the bottom of the screw hole volume. The geometry of the resulting recess was designed to mimic the volume of a screw hole in any of several commercially available total joint arthroplasty. To precisely control the opening of an initial gap between the polyethylene insert and the base plate, the insert was bonded to the flat steel loading platen (Fig. 1B). During loading of our in vitro model, the polyethylene insert was compressed into the base plate and screw hole. The metallic base plate containing the screw hole was supported from below with a stainless steel fixture that included a recess for the transducer (Fig. 1A and B). The transducer was mounted to the base plate using Teflon sealing tape and a gasket to limit leakage. The alignment of the steel-polyethylene surfaces was considered critical to generating reproducible results. Consequently, the loading fixture incorporated a joint filled with a bismuth-based, low-melting-point casting material (Cerrobend, McMaster-Carr), and for each experimental condition (described below) the casting material was melted, the polyethylene insert and base plate were aligned in compression, and the casting material was resolidified using a liquid cooling system. During the alignment process, the temperature of the polyethylene and tibial base plate did not change noticeably.
The apparatus was mounted in an MTS miniBionix servo-hydraulic testing machine. The fluid used in this study was phosphate-buffered saline (PBS) at room temperature (228C F 18C). Experimental Design We designed a total of 16 experiments in this study to quantify the effects of the following 5 categorical implant factors on peak fluid pressure within the screw hole (Table 1): 1. Polyethylene insert thickness (2 conditions: 5 and 10 mm). These thickness values cover the range of contemporary polyethylene components for primary hip and knee arthroplasty. 2. Initial gap between the insert and metal base plate (4 conditions: 0.01, 0.05, 0.1, and 0.2 mm). The initial gap was chosen to represent the range of tolerances potentially present between an acetabular liner and shell, as well as between the tibial insert and base plate [15]. 3. Surface finish of the metal base plate (2 conditions: as-machined vs mirror polished). The surface waviness of the as-machined surface was 1.39 F 0.54 lm, whereas the surface waviness of the polished surface was 0.27 F 0.14 lm. We chose these 2 conditions to cover a range of surface finishes available commercially. 4. Surface finish of the polyethylene insert (2 conditions: normal [as-machined] vs smooth). The average surface waviness of the 2 polyethylene surfaces was 2.74 F 0.05 lm for the normal asmachined condition and 0.60 F 0.02 lm for the smooth machined surface. (Waviness is a measurement similar to roughness, but it is made
Table 1. Test Matrix for 16 Experimental Conditions Investigated in This Study Experiment 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16
Insert thickness (mm)
Gap height (mm)
Base plate surface finish
Insert surface finish
Screw hole geometry
Testing rate (mm/s)
10 10 10 10 10 10 10 10 5 5 10 10 10 10 10 10
0.2 0.2 0.1 0.1 0.05 0.05 0.01 0.01 0.2 0.2 0.2 0.2 0.2 0.2 0.2 0.2
Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Mirror Polished Mirror Polished
Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Smooth Smooth Normal Normal Normal Normal
Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Chamfered Chamfered Cylindrical Cylindrical
1 2 1 2 1 2 1 2 1 0.5 1 2 1 2 1 2
In Vitro Model of Fluid Pressurization in Screw Holes ! Kurtz et al 935
over a larger length scale [16].) We chose these conditions to represent a range of surface finishes available commercially. 5. Screw hole geometry (2 conditions: cylindrical vs chamfered to mimic a contemporary implant design). Although the chamfered screw hole is typical in implants to accommodate the screw head, we also tested the cylindrical configuration which is more idealized geometry. As part of our test protocols we also included parametric variation of the gap closing velocity (0.5, 1, or 2 mm/s, depending upon the experiment, Table 1). Although the magnitude of the initial gap between the polyethylene insert and metal base plate was intended to reflect the manufacturing tolerances, the rate of gap closure, referred to here as the gap closing velocity, was considered to depend upon the loading rate in vivo. We calculated that based on the 0.1- to 0.2-s time necessary to generate the initial peak load during gait [17], clinically relevant gap closing velocities for our apparatus should range between 0.5 and 2 mm/s. For each of the 16 experiments in our test matrix (Table 1), we ran 2 to 3 loading trials to verify the repeatability of the setup, resulting in a series of 45 individual tests or independent observations for
this study. During each of the individual tests, screw hole pressure, load, and displacement data were recorded continuously at 1 kHz throughout 40 and 45 cycles of loading. For each loading cycle in every test, the peak pressure in the screw hole was identified (Fig. 2). For each test, the mean peak pressure was calculated across all of the individual 40 to 45 loading cycles. We used multivariate analysis of variance (ANOVA) to determine whether changes in the categorical study variables had a significant effect on the peak screw hole pressure ( P b .05 for significance). Further detailed ANOVA analyses were conducted on each block of experimental data, in which only one implant factor (eg, insert thickness) was modified, holding the other 4 implant factors constant. Post hoc t tests were also used to compare paired sets of data. JMP statistical discovery software (v. 5.0; SAS Institute, Cary, NC) was used for the statistical analyses. To assess whether the pressure peaks could be treated as individual measurements, a repeatedmeasures analysis was also performed on the peak pressure data collected throughout the cyclic tests. The results indicated that the initial 5 cycles were statistically different from the subsequent data, in terms of the peak pressures generated during an individual test. However, the changes in the statistical results due to excluding the first 5 cycles of data were small. An ANOVA for peak pressure data sets which excluded the first 5 cycles did not show a significant difference from the average peak pressure data for the original data sets ( P N .05).
Results
Fig. 2. Representative pressure–time history during 1 of the 40 to 45 loading cycles of an individual tests. In the illustrated loading cycle, the peak screw hole pressure was 28 kPa. There is an initial gap between the insert and base plate at the start and finish of the loading cycle. Under load, the polyethylene deforms into the screw hole, exerting pressure on the fluid, and squeezing some of the fluid out of the partially sealed boundary between the liner and the tray. During unloading, a negative pressure (suction) develops in the screw hole from transient fluid flow as the insert and tray are pulled apart.
By modifying the experimental design and 16 test conditions (Table 1), we were able to generate an order of magnitude of variation in the peak screw hole pressure under comparable, clinically relevant loading conditions. During our 45 independent tests, the peak screw hole pressure ranged between 16.0 and 163 kPa. Across all 45 tests, the peak pressure was, on average (FSD) 45.4 F 37.5 kPa. Although the magnitude of the peak screw hole pressure varied, depending upon the test conditions, the shape of the individual pressure–time history curves was generally consistent (Fig. 2). An initial steep rise in pressure, corresponding to initial gap closure, was followed by contact between the insert and the tray. The abrupt decrease in pressure at the conclusion of each cycle was followed by a short interval of negative pressure (bsuctionQ).
936 The Journal of Arthroplasty Vol. 20 No. 7 October 2005
Fig. 3. Effect of tray surface finish and testing rate on maximum screw hole pressure. At each gap closing velocity, the difference between polished and as-machined tray finish was significant ( P b .05).
Fig. 5. Effect of initial gap and testing rate on maximum screw hole pressure. Note that the biggest difference was between the bno gapQ condition and the 0.05-mminterface gap condition.
A multivariate ANOVA model described 97% of the variation in the peak screw hole pressure data based on r 2. The tray surface finish ( P b .0001), polyethylene insert thickness ( P b .0001), initial gap size ( P = .0009), and screw hole geometry ( P = .005) were all significant effects. Of the 5 implant factors included in the model, only the surface finish of the polyethyelene insert was not found to be significant ( P = .16). Of the implant factors studied, the surface finish of the tibial tray had the greatest effect on the peak screw hole pressure during the test (Fig. 3). For the block of experiments summarized in Fig. 3, changing the surface finish of the tray from asmachined to polished condition increased the average peak pressure from 31.8 to 134 kPa (321%). The polyethylene insert thickness had the second greatest effect on peak screw hole pressure during
the test (Fig. 4). Decreasing the insert thickness from 10 to 5 mm increased the average peak pressure from 31.8 to 50.2 kPa (58%), based on ANOVA of the block of experiments shown in Fig. 4. The initial gap between the insert and the tray had the third greatest effect on peak screw hole pressure during the test (Fig. 5). However, using post hoc analysis, we were only able to detect a significant difference between the 0.01-mm initial gap and the other initial gap conditions (0.05, 0.1, and 0.2 mm) (Fig. 5). Decreasing the initial gap between the insert and the metal tray from 0.1 to 0.01 mm decreased the average peak pressure from 38.9 to 21.0 kPa (46%) based on ANOVA of the block of experiments shown in Fig. 5. Screw hole geometry had a significant effect of screw hole pressures and was ranked fourth out of the 5 implant factors studied (Fig. 6). Changing the screw hole geometry from cylindrical to chamfered
Fig. 4. Effect of UHMWPE insert thickness on peak screw hole fluid pressure. The difference between insert thicknesses was significant ( P b .05).
Fig. 6. Effect of screw hole geometry and testing rate on peak screw hole pressure. At each gap closing velocity, the difference between cylindrical and chamfered screw holes was significant ( P b .05).
In Vitro Model of Fluid Pressurization in Screw Holes ! Kurtz et al 937
decreased the average peak pressure from 31.8 to 19.6 kPa (38%), based on ANOVA of the block of experiments shown in Fig. 6. Although unrelated to implant design, the gap closing velocity, or testing rate, was also significant covariate ( P b .0001). Increasing the gap closing velocity between 0.5 and 2 mm/s consistently increased the screw hole pressures by 25% to 34%, within each individual block of experiments in the study.
Discussion The results of this study suggest that the peak screw hole fluid pressures are sensitive to implant design. Of the implant factors investigated in the present study, the surface finish of the metallic base plate had the greatest effect on peak screw hole fluid pressures; the thickness of the polyethylene liner, as well as the gap between the liner and the base plate (at small gap sizes), also were significant contributors to screw hole fluid pressure. Our data suggest that unpolished metal base plates, thick polyethylene liners, and tight conformity between the liner and the metal base plate will all contribute to significantly reduced peak screw hole fluid pressures in joint arthroplasty. From our testing, the interface with the polished metallic base plate has the highest peak pressure (Fig. 3), which is consistent with what would be expected if fluid leakage between the 2 contacting implant components (the UHMWPE liner and the metal tray) was a key mechanical parameter. The implications for total joint components that incorporate a polished metal interface for liner support should be further studied. Our data suggest that fluid leakage at the UHMWPE–metallic shell interface is an important (but heretofore unappreciated) mechanism for minimizing fluid pressure applied to the prosthesis-bone interface. The magnitude of the peak fluid pressures measured in this study, which ranged between 16 and 163 kPa, were comparable to the thresholds of 17 to 138 kPa previously associated with producing osteolysis in cell culture models [18,19]. Cyclic fluid pressure with magnitudes of up to 138 kPa has been shown to activate macrophages, which play a major role in the onset of osteolysis in the membrane at the implant-bone interface [18,19]. Recent cell culture experiments further suggest that pressure-induced osteolysis may be synergistically mediated by the presence of wear debris at cyclic pressures as low as 17 kPa [19]. Consequently, the results of this in vitro study support plausibility of
the bfluid pressure inductionQ mechanism for focal periprosthetic osteolysis. The results of this in vitro study are limited to the effects of implant factors in isolation from the biological environment. In vivo, fluid flow from pressurization of the joint capsule, as well as from pressurization of the bone underlying the implant, may also influence the local fluid pressures in the screw hole [17,20]. Depending upon the permeability of the tissue interface and their inherent resistance to flow in vivo, the peak screw hole fluid pressures may be diminished, relative to the comparatively infinite resistance to fluid flow out of the screw holes provided by our test apparatus. Within this context, the experimental data collected in this in vitro study represent a theoretical upper bound on the fluid pressure in screw holes. The large effect of metallic implant surface finish on the induced pressures in this study suggests that the fluid leakage path from the loaded hole to the outside of the experimental fixture is very thin— comparable in height to the submicrometer surface features of the nonpolished implant. This indicates that more complex fluids, containing proteins or debris, could change the induced pressure in this experiment. In this study, PBS was used at room temperature as a baseline for further experiments. Further experiments are warranted to assess the relative effects of more complex fluids, such as 30% bovine calf serum at 378C, and to assess the behavior of this experimental system with production surface finishes of the UHMWPE and metallic parts.
Acknowledgments This study was supported by a research grant from Stryker Orthopedics. The authors thank Brian Brooks, Exponent, for his contributions.
References 1. Harris WH. Wear and periprosthetic osteolysis: the problem. Clin Orthop 2001;393:66. 2. Harris WH. The problem is osteolysis. Clin Orthop 1995;311:46. 3. Skripitz R, Aspenberg P. Pressure-induced periprosthetic osteolysis: a rat model. J Orthop Res 2000; 18:481. 4. Van der Vis HM, Aspenberg P, Marti RK, et al. Fluid pressure causes bone resorption in a rabbit model of prosthetic loosening. Clin Orthop 1998;350:201. 5. Khalily C, Tanner MG, Williams VG, et al. Effect of locking mechanism on fluid and particle flow
938 The Journal of Arthroplasty Vol. 20 No. 7 October 2005
6.
7.
8.
9.
10.
11.
12.
13.
through modular acetabular components. J Arthroplasty 1998;13:254. Crawford RW, Evans M, Ling RS, et al. Fluid flow around model femoral components of differing surface finishes: in vitro investigations. Acta Orthop Scand 1999;70:589. Manley MT, D’Antonio JA, Capello WN, et al. Osteolysis: a disease of access to fixation interfaces. Clin Orthop 2002;405:129. Schmalzried TP, Guttmann D, Grecula M, et al. The relationship between the design, position, and articular wear of acetabular components inserted without cement and the development of pelvic osteolysis. J Bone Joint Surg Am 1994;76:677. Kurtz SM, Ochoa JA, Hovey CB, et al. Simulation of initial frontside and backside wear rates in a modular acetabular component with multiple screw holes. J Biomech 1999;32:967. Peters Jr PC, Engh GA, Dwyer KA, et al. Osteolysis after total knee arthroplasty without cement. J Bone Joint Surg Am 1992;74:864. Whiteside LA. Effect of porous-coating configuration on tibial osteolysis after total knee arthroplasty. Clin Orthop 1995:92. Ezzet KA, Garcia R, Barrack RL. Effect of component fixation method on osteolysis in total knee arthroplasty. Clin Orthop 1995;321:86. Wasielewski RC. The causes of insert backside wear in total knee arthroplasty. Clin Orthop 2002:232.
14. Kurtz SM, Edidin AA, Bartel DL. The role of backside polishing, cup angle, and polyethylene thickness on the contact stresses in metal-backed acetabular components. J Biomech 1997;30:639. 15. Harrigan T, Kurtz SM, Charriere E, et al. Acetabular shell-liner mismatch: fluid volume displacement and contact conditions. Presented at the annual Meeting of the Transactions of the 50th Orthopedic Research Society, San Francisco, CA, February 21-24, 2004. 16. Kurtz SM, Turner JL, Herr M, et al. Deconvolution of surface topology for quantification of initial wear in highly cross-linked acetabular components for THA. J Biomed Mater Res 2002;63:492. 17. Downey DJ, Simkin PA, Taggart R. The effect of compressive loading on intraosseous pressure in the femoral head in vitro. J Bone Joint Surg Am 1988; 70:871. 18. Ferrier GM, McEvoy A, Evans CE, et al. The effect of cyclic pressure on human monocyte-derived macrophages in vitro. J Bone Joint Surg Br 2000;82:755. 19. Mevoy A, Jeyam M, Ferrier G, et al. Synergistic effect of particles and cyclic pressure on cytokine production in human monocyte/macrophages: proposed role in periprosthetic osteolysis. Bone 2002;30:171. 20. Robertsson O, Wingstrand H, Kesteris U, et al. Intracapsular pressure and loosening of hip prostheses. Preoperative measurements in 18 hips. Acta Orthop Scand 1997;68:231.
An In Vitro Model for Fluid Pressurization of Screw Holes in Metal-Backed Total Joint Components Steven M. Kurtz, PhD,*y Timothy P. Harrigan, ScD,* Michael Herr, MS,* and Michael T. Manley, PhDz
Abstract: Fluid pressure may stimulate osteolysis near screw holes in joint arthroplasty components. We developed a generalized in vitro model of a polyethylene liner and metal backing with a screw hole to investigate whether implant design factors influence local fluid pressure. We observed an order of magnitude of variation in the peak screw hole pressure (from 16.0 and 163 kPa) under clinically relevant loading conditions. Of the implant factors investigated, the surface finish of the metallic base plate had the greatest effect on peak screw hole fluid pressures; the thickness of the polyethylene liner, as well as the gap between the liner and the base plate, were also significant design variables. Our data suggest that unpolished metal base plates, thick polyethylene liners, and tight conformity between the liner and the metal base plate will all contribute to significantly reduced peak screw hole fluid pressures in joint arthroplasty. Key words: osteolysis, fluid pressure, implant design, conformity, thickness, surface finish. n 2005 Elsevier Inc. All rights reserved.
have been observed near the locations of uncovered screw holes [7,8]. The magnitude of backside wear, between the liner and shell, is orders of magnitude lower than the volume of wear generated between the femoral head and the liner [9]. If the focal pelvic lesions are indeed provoked solely by wear, the articulating surface is a more likely culprit for the source of the debris, which then migrates through the gap between the liner and the shell to attack bone through the screw holes [5]. An alternative, or complementary, hypothesis to the debris pumping scenario is that fluid pressure —either independently or in combination with the presence of wear debris —directly stimulates the production of focal lesions in the vicinity of screw holes [3,4]. In total-knee arthroplasty (TKA), tibial osteolysis has been reported in less than 5 years of implantation when screws or cementless fixation have been used [10-12]. Implant design factors, such as the configuration of porous coating on the tibial base plate, have been reported to influence the incidence of tibial osteolysis after cementless TKA [11], suggesting that fluid transport or fluid pressure influence tibial osteolysis. Furthermore, the
Periprosthetic osteolysis is recognized as a critical and limiting problem for total joint arthroplasties [1,2]. Although osteolysis is clearly linked to ultrahigh molecular weight polyethylene (hereafter, polyethylene) particulate debris, fluid pressure and fluid flow in the tissues around a total joint implant have also been shown to directly stimulate osteolysis in animal models [3,4]. Fluid flow may influence osteolysis by providing a mechanism for wear particles to reach the vulnerable periprosthetic bone [5,6]. Fluid-induced periprosthetic osteolysis is clinically relevant for both hip and knee arthroplasties. In total-hip arthroplasty (THA), focal acetabular lesions From *Exponent, Inc, Philadelphia, Pennsylvania; y Drexel University, Philadelphia, Pennsylvania; and z Ridgewood, New Jersey. Submitted April 28, 2004; accepted November 10, 2004. Benefits or funds were received in partial or total support of the research material described in this article from Stryker Orthopedics. Reprint requests: Steven M. Kurtz, PhD, Exponent, Inc, Suite 300, 3401 Market St, Philadelphia, PA 19104. n 2005 Elsevier Inc. All rights reserved. 0883-5403/05/1906-0004$30.00/0 doi:10.1016/j.arth.2004.11.003
932
In Vitro Model of Fluid Pressurization in Screw Holes ! Kurtz et al 933
intrusion of the UHMWPE tibial insert into vacant screw holes has also been identified clinically as a mechanism of backside wear [13]. Overall, the clinical observations to date would suggest that focal osteolytic lesions around screw holes is a general concern shared by metal backed tibial as well as acetabular components for joint arthroplasty. Although many previous studies have identified focal osteolysis around screw holes to be a clinical problem for both THA and TKA, comparatively limited progress has been made in developing experimental or analytical models to predict fluid pressure around joint arthroplasties. The fluid flow in the vicinity of a reconstructed joint is likely to be complex, influenced not only by the geometry and material properties of the prosthetic components, but also by the hard and soft tissues, their porosity, and their apposition to the implant interface where the metal shell and the tissue interdigitate. Given this complex fluid and structural system, this study takes a reductionist approach, where a limited number of the relevant mechanical features of the implant are studied. Based on these results, the role of other factors can be investigated by studies that are progressively more complex, giving progressively better approximations to the clinical situation. In this study, we tested the hypothesis that implant factors, such as liner thickness, screw hole geometry, and surface roughness, influence the fluid pressure within a screw hole of the supporting metal backing of a joint arthroplasty. To test this hypothesis we developed a generalized in vitro model of a polyethylene liner and a metal backing with a screw hole. This generalized in vitro model represents the first step toward better quantitative understanding
of fluid flow and pressure development around screw holes in total joint arthroplasties.
Materials and Methods Experimental Apparatus For this study we developed a generalized in vitro model of a flat polyethylene insert, centered over a screw hole in a flat metallic base plate. The insert was loaded superiorly by a flat, aligned platen (Fig. 1A and B). This in vitro model most closely replicates the scenario of a tibial insert loading a metallic tray. However, this model is also relevant to modular acetabular components, where the local geometry at the back side of the liner-shell interface has a relatively large radius of curvature. Compression molded, medical grade polyethylene (GUR 1050: Ticona, Bishop, Tex) was used to simulate inserts of different thicknesses. The polyethylene specimens were machined to the following dimensions: 31 mm length 31 mm width 5 or 10 mm thickness. The insert thicknesses of 5 or 10 mm were chosen to be consistent with commercially manufactured hip and knee components, and the length and width were chosen to produce, on average, 5-MPa contact pressure at a load of 4805 N. The target average pressure of 5 MPa for our in vitro model was chosen because the contact stress on the backside of an acetabular liner has been shown to range between 5 and 10 MPa under clinically relevant, gait loading conditions in previous finite element simulations of THA [14].
Fig. 1. In vitro screw hole pressure testing apparatus. A, Assembled apparatus, illustrating the PBS testing environment as well as the supporting fixtures. B, Partially disassembled apparatus, showing the generalized representation of the polyethylene insert and base plate with screw hole. The initial gap between the insert and base plate has been exaggerated for demonstration purposes (the actual gap distance in the experiments ranged between 0.01 and 0.2 mm).
934 The Journal of Arthroplasty Vol. 20 No. 7 October 2005 In our apparatus, an 8-mm-diameter screw hole (either cylindrical or tapered to accept a chamfered bone screw) was drilled into a flat, 4-mm-thick stainless steel base plate. The bottom of the screw hole was tapped to allow a pressure transducer to be mounted. In this manner, the sensing membrane of the transducer was at the bottom of the screw hole volume. The geometry of the resulting recess was designed to mimic the volume of a screw hole in any of several commercially available total joint arthroplasty. To precisely control the opening of an initial gap between the polyethylene insert and the base plate, the insert was bonded to the flat steel loading platen (Fig. 1B). During loading of our in vitro model, the polyethylene insert was compressed into the base plate and screw hole. The metallic base plate containing the screw hole was supported from below with a stainless steel fixture that included a recess for the transducer (Fig. 1A and B). The transducer was mounted to the base plate using Teflon sealing tape and a gasket to limit leakage. The alignment of the steel-polyethylene surfaces was considered critical to generating reproducible results. Consequently, the loading fixture incorporated a joint filled with a bismuth-based, low-melting-point casting material (Cerrobend, McMaster-Carr), and for each experimental condition (described below) the casting material was melted, the polyethylene insert and base plate were aligned in compression, and the casting material was resolidified using a liquid cooling system. During the alignment process, the temperature of the polyethylene and tibial base plate did not change noticeably.
The apparatus was mounted in an MTS miniBionix servo-hydraulic testing machine. The fluid used in this study was phosphate-buffered saline (PBS) at room temperature (228C F 18C). Experimental Design We designed a total of 16 experiments in this study to quantify the effects of the following 5 categorical implant factors on peak fluid pressure within the screw hole (Table 1): 1. Polyethylene insert thickness (2 conditions: 5 and 10 mm). These thickness values cover the range of contemporary polyethylene components for primary hip and knee arthroplasty. 2. Initial gap between the insert and metal base plate (4 conditions: 0.01, 0.05, 0.1, and 0.2 mm). The initial gap was chosen to represent the range of tolerances potentially present between an acetabular liner and shell, as well as between the tibial insert and base plate [15]. 3. Surface finish of the metal base plate (2 conditions: as-machined vs mirror polished). The surface waviness of the as-machined surface was 1.39 F 0.54 lm, whereas the surface waviness of the polished surface was 0.27 F 0.14 lm. We chose these 2 conditions to cover a range of surface finishes available commercially. 4. Surface finish of the polyethylene insert (2 conditions: normal [as-machined] vs smooth). The average surface waviness of the 2 polyethylene surfaces was 2.74 F 0.05 lm for the normal asmachined condition and 0.60 F 0.02 lm for the smooth machined surface. (Waviness is a measurement similar to roughness, but it is made
Table 1. Test Matrix for 16 Experimental Conditions Investigated in This Study Experiment 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16
Insert thickness (mm)
Gap height (mm)
Base plate surface finish
Insert surface finish
Screw hole geometry
Testing rate (mm/s)
10 10 10 10 10 10 10 10 5 5 10 10 10 10 10 10
0.2 0.2 0.1 0.1 0.05 0.05 0.01 0.01 0.2 0.2 0.2 0.2 0.2 0.2 0.2 0.2
Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Mirror Polished Mirror Polished
Normal Normal Normal Normal Normal Normal Normal Normal Normal Normal Smooth Smooth Normal Normal Normal Normal
Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Cylindrical Chamfered Chamfered Cylindrical Cylindrical
1 2 1 2 1 2 1 2 1 0.5 1 2 1 2 1 2
In Vitro Model of Fluid Pressurization in Screw Holes ! Kurtz et al 935
over a larger length scale [16].) We chose these conditions to represent a range of surface finishes available commercially. 5. Screw hole geometry (2 conditions: cylindrical vs chamfered to mimic a contemporary implant design). Although the chamfered screw hole is typical in implants to accommodate the screw head, we also tested the cylindrical configuration which is more idealized geometry. As part of our test protocols we also included parametric variation of the gap closing velocity (0.5, 1, or 2 mm/s, depending upon the experiment, Table 1). Although the magnitude of the initial gap between the polyethylene insert and metal base plate was intended to reflect the manufacturing tolerances, the rate of gap closure, referred to here as the gap closing velocity, was considered to depend upon the loading rate in vivo. We calculated that based on the 0.1- to 0.2-s time necessary to generate the initial peak load during gait [17], clinically relevant gap closing velocities for our apparatus should range between 0.5 and 2 mm/s. For each of the 16 experiments in our test matrix (Table 1), we ran 2 to 3 loading trials to verify the repeatability of the setup, resulting in a series of 45 individual tests or independent observations for
this study. During each of the individual tests, screw hole pressure, load, and displacement data were recorded continuously at 1 kHz throughout 40 and 45 cycles of loading. For each loading cycle in every test, the peak pressure in the screw hole was identified (Fig. 2). For each test, the mean peak pressure was calculated across all of the individual 40 to 45 loading cycles. We used multivariate analysis of variance (ANOVA) to determine whether changes in the categorical study variables had a significant effect on the peak screw hole pressure ( P b .05 for significance). Further detailed ANOVA analyses were conducted on each block of experimental data, in which only one implant factor (eg, insert thickness) was modified, holding the other 4 implant factors constant. Post hoc t tests were also used to compare paired sets of data. JMP statistical discovery software (v. 5.0; SAS Institute, Cary, NC) was used for the statistical analyses. To assess whether the pressure peaks could be treated as individual measurements, a repeatedmeasures analysis was also performed on the peak pressure data collected throughout the cyclic tests. The results indicated that the initial 5 cycles were statistically different from the subsequent data, in terms of the peak pressures generated during an individual test. However, the changes in the statistical results due to excluding the first 5 cycles of data were small. An ANOVA for peak pressure data sets which excluded the first 5 cycles did not show a significant difference from the average peak pressure data for the original data sets ( P N .05).
Results
Fig. 2. Representative pressure–time history during 1 of the 40 to 45 loading cycles of an individual tests. In the illustrated loading cycle, the peak screw hole pressure was 28 kPa. There is an initial gap between the insert and base plate at the start and finish of the loading cycle. Under load, the polyethylene deforms into the screw hole, exerting pressure on the fluid, and squeezing some of the fluid out of the partially sealed boundary between the liner and the tray. During unloading, a negative pressure (suction) develops in the screw hole from transient fluid flow as the insert and tray are pulled apart.
By modifying the experimental design and 16 test conditions (Table 1), we were able to generate an order of magnitude of variation in the peak screw hole pressure under comparable, clinically relevant loading conditions. During our 45 independent tests, the peak screw hole pressure ranged between 16.0 and 163 kPa. Across all 45 tests, the peak pressure was, on average (FSD) 45.4 F 37.5 kPa. Although the magnitude of the peak screw hole pressure varied, depending upon the test conditions, the shape of the individual pressure–time history curves was generally consistent (Fig. 2). An initial steep rise in pressure, corresponding to initial gap closure, was followed by contact between the insert and the tray. The abrupt decrease in pressure at the conclusion of each cycle was followed by a short interval of negative pressure (bsuctionQ).
936 The Journal of Arthroplasty Vol. 20 No. 7 October 2005
Fig. 3. Effect of tray surface finish and testing rate on maximum screw hole pressure. At each gap closing velocity, the difference between polished and as-machined tray finish was significant ( P b .05).
Fig. 5. Effect of initial gap and testing rate on maximum screw hole pressure. Note that the biggest difference was between the bno gapQ condition and the 0.05-mminterface gap condition.
A multivariate ANOVA model described 97% of the variation in the peak screw hole pressure data based on r 2. The tray surface finish ( P b .0001), polyethylene insert thickness ( P b .0001), initial gap size ( P = .0009), and screw hole geometry ( P = .005) were all significant effects. Of the 5 implant factors included in the model, only the surface finish of the polyethyelene insert was not found to be significant ( P = .16). Of the implant factors studied, the surface finish of the tibial tray had the greatest effect on the peak screw hole pressure during the test (Fig. 3). For the block of experiments summarized in Fig. 3, changing the surface finish of the tray from asmachined to polished condition increased the average peak pressure from 31.8 to 134 kPa (321%). The polyethylene insert thickness had the second greatest effect on peak screw hole pressure during
the test (Fig. 4). Decreasing the insert thickness from 10 to 5 mm increased the average peak pressure from 31.8 to 50.2 kPa (58%), based on ANOVA of the block of experiments shown in Fig. 4. The initial gap between the insert and the tray had the third greatest effect on peak screw hole pressure during the test (Fig. 5). However, using post hoc analysis, we were only able to detect a significant difference between the 0.01-mm initial gap and the other initial gap conditions (0.05, 0.1, and 0.2 mm) (Fig. 5). Decreasing the initial gap between the insert and the metal tray from 0.1 to 0.01 mm decreased the average peak pressure from 38.9 to 21.0 kPa (46%) based on ANOVA of the block of experiments shown in Fig. 5. Screw hole geometry had a significant effect of screw hole pressures and was ranked fourth out of the 5 implant factors studied (Fig. 6). Changing the screw hole geometry from cylindrical to chamfered
Fig. 4. Effect of UHMWPE insert thickness on peak screw hole fluid pressure. The difference between insert thicknesses was significant ( P b .05).
Fig. 6. Effect of screw hole geometry and testing rate on peak screw hole pressure. At each gap closing velocity, the difference between cylindrical and chamfered screw holes was significant ( P b .05).
In Vitro Model of Fluid Pressurization in Screw Holes ! Kurtz et al 937
decreased the average peak pressure from 31.8 to 19.6 kPa (38%), based on ANOVA of the block of experiments shown in Fig. 6. Although unrelated to implant design, the gap closing velocity, or testing rate, was also significant covariate ( P b .0001). Increasing the gap closing velocity between 0.5 and 2 mm/s consistently increased the screw hole pressures by 25% to 34%, within each individual block of experiments in the study.
Discussion The results of this study suggest that the peak screw hole fluid pressures are sensitive to implant design. Of the implant factors investigated in the present study, the surface finish of the metallic base plate had the greatest effect on peak screw hole fluid pressures; the thickness of the polyethylene liner, as well as the gap between the liner and the base plate (at small gap sizes), also were significant contributors to screw hole fluid pressure. Our data suggest that unpolished metal base plates, thick polyethylene liners, and tight conformity between the liner and the metal base plate will all contribute to significantly reduced peak screw hole fluid pressures in joint arthroplasty. From our testing, the interface with the polished metallic base plate has the highest peak pressure (Fig. 3), which is consistent with what would be expected if fluid leakage between the 2 contacting implant components (the UHMWPE liner and the metal tray) was a key mechanical parameter. The implications for total joint components that incorporate a polished metal interface for liner support should be further studied. Our data suggest that fluid leakage at the UHMWPE–metallic shell interface is an important (but heretofore unappreciated) mechanism for minimizing fluid pressure applied to the prosthesis-bone interface. The magnitude of the peak fluid pressures measured in this study, which ranged between 16 and 163 kPa, were comparable to the thresholds of 17 to 138 kPa previously associated with producing osteolysis in cell culture models [18,19]. Cyclic fluid pressure with magnitudes of up to 138 kPa has been shown to activate macrophages, which play a major role in the onset of osteolysis in the membrane at the implant-bone interface [18,19]. Recent cell culture experiments further suggest that pressure-induced osteolysis may be synergistically mediated by the presence of wear debris at cyclic pressures as low as 17 kPa [19]. Consequently, the results of this in vitro study support plausibility of
the bfluid pressure inductionQ mechanism for focal periprosthetic osteolysis. The results of this in vitro study are limited to the effects of implant factors in isolation from the biological environment. In vivo, fluid flow from pressurization of the joint capsule, as well as from pressurization of the bone underlying the implant, may also influence the local fluid pressures in the screw hole [17,20]. Depending upon the permeability of the tissue interface and their inherent resistance to flow in vivo, the peak screw hole fluid pressures may be diminished, relative to the comparatively infinite resistance to fluid flow out of the screw holes provided by our test apparatus. Within this context, the experimental data collected in this in vitro study represent a theoretical upper bound on the fluid pressure in screw holes. The large effect of metallic implant surface finish on the induced pressures in this study suggests that the fluid leakage path from the loaded hole to the outside of the experimental fixture is very thin— comparable in height to the submicrometer surface features of the nonpolished implant. This indicates that more complex fluids, containing proteins or debris, could change the induced pressure in this experiment. In this study, PBS was used at room temperature as a baseline for further experiments. Further experiments are warranted to assess the relative effects of more complex fluids, such as 30% bovine calf serum at 378C, and to assess the behavior of this experimental system with production surface finishes of the UHMWPE and metallic parts.
Acknowledgments This study was supported by a research grant from Stryker Orthopedics. The authors thank Brian Brooks, Exponent, for his contributions.
References 1. Harris WH. Wear and periprosthetic osteolysis: the problem. Clin Orthop 2001;393:66. 2. Harris WH. The problem is osteolysis. Clin Orthop 1995;311:46. 3. Skripitz R, Aspenberg P. Pressure-induced periprosthetic osteolysis: a rat model. J Orthop Res 2000; 18:481. 4. Van der Vis HM, Aspenberg P, Marti RK, et al. Fluid pressure causes bone resorption in a rabbit model of prosthetic loosening. Clin Orthop 1998;350:201. 5. Khalily C, Tanner MG, Williams VG, et al. Effect of locking mechanism on fluid and particle flow
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6.
7.
8.
9.
10.
11.
12.
13.
through modular acetabular components. J Arthroplasty 1998;13:254. Crawford RW, Evans M, Ling RS, et al. Fluid flow around model femoral components of differing surface finishes: in vitro investigations. Acta Orthop Scand 1999;70:589. Manley MT, D’Antonio JA, Capello WN, et al. Osteolysis: a disease of access to fixation interfaces. Clin Orthop 2002;405:129. Schmalzried TP, Guttmann D, Grecula M, et al. The relationship between the design, position, and articular wear of acetabular components inserted without cement and the development of pelvic osteolysis. J Bone Joint Surg Am 1994;76:677. Kurtz SM, Ochoa JA, Hovey CB, et al. Simulation of initial frontside and backside wear rates in a modular acetabular component with multiple screw holes. J Biomech 1999;32:967. Peters Jr PC, Engh GA, Dwyer KA, et al. Osteolysis after total knee arthroplasty without cement. J Bone Joint Surg Am 1992;74:864. Whiteside LA. Effect of porous-coating configuration on tibial osteolysis after total knee arthroplasty. Clin Orthop 1995:92. Ezzet KA, Garcia R, Barrack RL. Effect of component fixation method on osteolysis in total knee arthroplasty. Clin Orthop 1995;321:86. Wasielewski RC. The causes of insert backside wear in total knee arthroplasty. Clin Orthop 2002:232.
14. Kurtz SM, Edidin AA, Bartel DL. The role of backside polishing, cup angle, and polyethylene thickness on the contact stresses in metal-backed acetabular components. J Biomech 1997;30:639. 15. Harrigan T, Kurtz SM, Charriere E, et al. Acetabular shell-liner mismatch: fluid volume displacement and contact conditions. Presented at the annual Meeting of the Transactions of the 50th Orthopedic Research Society, San Francisco, CA, February 21-24, 2004. 16. Kurtz SM, Turner JL, Herr M, et al. Deconvolution of surface topology for quantification of initial wear in highly cross-linked acetabular components for THA. J Biomed Mater Res 2002;63:492. 17. Downey DJ, Simkin PA, Taggart R. The effect of compressive loading on intraosseous pressure in the femoral head in vitro. J Bone Joint Surg Am 1988; 70:871. 18. Ferrier GM, McEvoy A, Evans CE, et al. The effect of cyclic pressure on human monocyte-derived macrophages in vitro. J Bone Joint Surg Br 2000;82:755. 19. Mevoy A, Jeyam M, Ferrier G, et al. Synergistic effect of particles and cyclic pressure on cytokine production in human monocyte/macrophages: proposed role in periprosthetic osteolysis. Bone 2002;30:171. 20. Robertsson O, Wingstrand H, Kesteris U, et al. Intracapsular pressure and loosening of hip prostheses. Preoperative measurements in 18 hips. Acta Orthop Scand 1997;68:231.