Excluding the thick-walled MOM shell, which showed consistently low maximum deformation throughout the testing period, an overall reduction in deformation (p < 0.001) was observed from initial implantation to 120 hours after implantation in both the pinch and relief regions of the model. An average of 18% reduction in compression over all shell designs was observed 48 hours after implantation, while an average of 7% additional pinch reduction was observed between 48 and 120 hours after implantation. Similarly, we observed a mean 13% reduction in diametral expansion in the relief regions 48 hours after implantation, with an additional 6% mean reduction through 120 hours after implantation.
Acetabular cup deformation, an inevitable mechanical consequence in uncemented press-fit components, can lead to a breakdown in lubrication mechanisms resulting in suboptimal conditions for long-term success for MOM and modular cup designs. Our study represents an effort to quantify deformation over a segment of the wide variety of design and material variations available in the current orthopaedic marketplace. Specifically, we investigated whether (1) a change in component geometry and design altered maximum shell deformation of MOM and metal-backed modular cups in a biomechanical model and (2) any time-dependent deformational relaxation processes were exhibited in the model.
We recognize limitations in our methodology and clinical extrapolation of our findings. First, we used a static, polyurethane model to represent a dynamic in vivo environment. Clinical projection of true component deformation in the native acetabulum may be limited. This model is validated for recreating and measuring initial cup deformation . This model has not been validated for measuring cup deformation beyond initial implantation. Second, a single material density and a single interference fit were tested though clinically they vary widely. Interference fit is subjective, fluctuating from case to case as the strength in the bone and surgeon preference dictates. The amount of variance that typically occurs clinically with 1 mm of underreaming is not only unknown but rather broad. Variation in reaming derives from variation in surgeon-applied force, instrument manufacturing, and bone quality; therefore, control of these variables within a representative density and underream setting provides a baseline for comparison in this study. Third, we used only a single-sized thick-walled MOM component; thus, conclusions based on altered component size from these data can only be confidently applied to thin-walled MOM designs. Fourth, this study incorporated a single loading condition while varied deformational response could be expected in dynamic loading or in conditions of rim loading and incomplete shell seating; however, this loading condition provides a relative baseline of comparison among varied cup designs.
Recent studies have begun to examine the effect of polyethylene liner deformation in regard to frictional torque, liner fracture, and volumetric wear in foam block and finite element analyses [8, 11, 14]. Schmidig et al.  observed higher magnitudes of deformation in the polyethylene liner than in the shell in which they were inserted. Because maximal deformation occurs at the periphery of the cup, this change could potentially lead to adverse consequences with respect to the peripheral locking mechanisms in modular cups. This finding may be of particular concern with respect to the reduction of fracture toughness (increased stress and crack initiation) of the newer polyethylene. However, with stress relaxation and diminished deformation of the polyethylene liner over time, reduced effect on fracture and wear would be expected. Likewise, the negative clinical impact of this polyethylene deformation may be ameliorated through modern polyethylene processing techniques . It is important to note these potential adverse consequences of liner deformation have yet to be demonstrated clinically.
The viscoelasticity of bone lends to the hypothesis of diminished cup deformation over time in both MOM and metal-backed polyethylene acetabular cups. However, to our knowledge, no extensive long-term in vivo cup deformation data are available in the current scientific literature. In this study, a reduction in initial deformation magnitude was observed across all but the stiffest MOM designs in this nonphysiologic model. Our data would seem to support the hypothesis of long-term reduction in initial cup deformation in the more transient in vivo environment. More extensive evaluation in vivo should be performed to fully characterize the time dependence of cup deformation in press-fit implantation.
In summary, our findings agree with those of previously published experimental and finite element studies maintaining decreased wall thickness and increased cup diameter lead to higher initial cup deformations in a worst-case-scenario pinch relief cancellous foam model. Thick-walled MOM acetabular shells exhibited small overall shell deformations, while thin-walled components were susceptible to larger deformations. Furthermore, conventional metal-backed shells for modular polyethylene liners exhibited higher deformations than the majority of the MOM designs and sizes tested. This study indicates component material and design as factors in initial acetabular cup deformation in these specific devices. Further study is needed to assess the effects of pelvic viscoelasticity, bony remodeling, and polyethylene liner insertion on long-term component deformation and its impact on implant survivability.
The authors thank Michael Volitich BS, Ron Hofmann BS, and Gary Burgess BS, for their assistance in experimental setup. We additionally thank Kenneth Davis MS, and Matthew Brunsman MS, for their assistance with statistical analysis.
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