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The biomechanical effect of anteversion and modular neck offset on stress shielding for short-stem versus conventional long-stem hip implants
Medical Engineering and Physics 38 (2016) 232–240
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The biomechanical effect of anteversion and modular neck offset on stress shielding for short-stem versus conventional long-stem hip implants Peter Goshulak a,c, Saeid Samiezadeh b, Mina S.R. Aziz c, Habiba Bougherara b, Radovan Zdero b,c,d,∗, Emil H. Schemitsch a,c a
Institute of Biomaterials and Biomedical Engineering, University of Toronto, Toronto M5S 3G9, ON, Canada Department of Mechanical and Industrial Engineering, Ryerson University, Toronto M5B 2K3, ON, Canada c Martin Orthopaedic Biomechanics Laboratory, St. Michael’s Hospital, Toronto M5B 1W8, ON, Canada d Department of Mechanical and Industrial Engineering, University of Toronto, Toronto M5S 3G8, ON, Canada b
a r t i c l e
i n f o
Article history: Received 2 May 2015 Revised 22 October 2015 Accepted 6 December 2015
Keywords: Biomechanics Stress shielding Short-stem Hip arthroplasty Finite element analysis
a b s t r a c t Short-stem hip implants are increasingly common since they preserve host bone stock and presumably reduce stress shielding by improving load distribution in the proximal femur. Stress shielding may lead to decreased bone density, implant loosening, and fracture. However, few biomechanical studies have examined short-stem hip implants. The purpose of this study was to compare short-stem vs. standard length stemmed implants for stress shielding effects due to anteversion–retroversion, anterior–posterior position, and modular neck offset. Twelve artificial femurs were implanted with either a short-stem modular-neck implant or a conventional length monolithic implant in 0° or 15° of anteversion. Three modular neck options were tested in the short-stem implants. Three control femurs remained intact. Femurs were mounted in adduction and subjected to axial loading. Strain gauge values were collected to validate a Finite Element (FE) model, which was used to simulate the full range of physiologically possible anteversion and anterior–posterior combinations (n = 25 combinations per implant). Calcar stress was compared between implants and across each implant’s range of anteversion using one and two-way ANOVA. Stress shielding was defined as the overall change in stress compared to an intact femur. The FE model compared well with experimental strains (intact: slope = 0.898, R = 0.943; short-stem: slope = 0.731, R = 0.948; standard-stem: slope = 0.743, R = 0.859); correction factors were used to adjust slopes to unity. No implant anteversion showed significant reduction in stress shielding (α = 0.05, p > 0.05). Stress shielding was significantly higher in the standard-stem implant (63% change from intact femur, p < 0.001) than in short-stem implants (29–39% change, p < 0.001). Short-stem implants reduce stress shielding compared to standard length stemmed implants, while implant anteversion and anterior–posterior position had no effect. Therefore, short-stem implants have a greater likelihood of maintaining calcar bone strength in the long term. © 2015 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction Hip replacement surgery is becoming more prevalent in younger populations [1,2]. In order to facilitate possible future revision surgeries in young patients, maximum bone stock must be conserved [3]. Maintaining bone strength has always been a concern for load-bearing prostheses, but an increasingly active patient population [4] increases the demand for implant designs which maximize the load-bearing capabilities of the bone. As bone re∗
Corresponding author. Tel.: +1 416 953 5328. E-mail address: [email protected] (R. Zdero).
http://dx.doi.org/10.1016/j.medengphy.2015.12.005 1350-4533/© 2015 IPEM. Published by Elsevier Ltd. All rights reserved.
models itself in response to stress, regions of poor stress distribution become resorbed [5]. This “stress shielding” in the femur occurs primarily in the calcar region [6,7]. This decreased bone density may lead to aseptic loosening and stem migration, as well as periprosthetic fractures [6,8,9]. Implants should aim to replicate the physiological stress distribution, or that of an intact femur [10]. To this end, short-stem implants were relatively recently developed to conserve more bone stock in order to reduce stress shielding. These implants have good short-term clinical results such as reduced fracture and revision rate compared to conventional implants [9,11–13]. Specifically, for example, recent clinical studies on short-stem hip implants have shown certain advantages: vertical
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center of rotation, horizontal center of rotation, and leg length of the implanted hip vs. non-operated contralateral hip were within 5 mm, respectively, for 89%, 80%, and 96% of patients [14]; shortstem hips implants vs. hip resurfacing implants demonstrated 40% shorter surgery time and similar pre- and post-surgery outcomes for HOOS (hip disability and osteoarthritis outcome score) [15]; and revision rates over a 10-year follow-up period due to aseptic loosening, undersized stems, periprosthetic fracture, or via falsa are required in less than 2% of cases [16]. Similarly, recent biomechanical studies have shown several benefits compared to traditional hip stems: short-stem hip implants are more resistant to axial migration and retroversion measured at 100,000 cycles [17], whilst short-stem implants with predominantly metaphyseal fixation can provide cortical strain patterns in the proximal femur (including the calcar) more similar to an intact femur thereby potentially reducing “stress shielding” and improving implant stability [18]. However, to the authors’ knowledge, the biomechanical effects on stress shielding of a short-stem hip implant’s position in the host bone and modular neck selection have not been investigated. The goal of this experimental and computational study was to compare stress shielding between a short-stem implant versus a conventional length stemmed implant by analyzing all theoretically possible combinations of implant orientations plus clinicallyfeasible modular neck options. 2. Methods 2.1. General strategy Three study phases were used to quantify the stress shielding caused by surgical anteversion, anterior–posterior translation, and modular neck selection. Phase 1 involved mechanical experimentation on synthetic femurs to establish strain profiles for all implant orientations and modular neck combinations over a range of physiological load conditions. Phase 2 involved constructing a Finite Element (FE) model of the experimental setup and using the strain data to validate this model. Phase 3 used “Design of Experiments” on the validated FE model to analyze stress shielding for all possible implant orientations and modular neck options. This allowed clinical recommendations to be made about optimal surgical use of a typical short-stem hip implant.
Fig. 1. Superior view of the 4 implant and neck combinations inserted into a left femur: a) SMF stem, neutral neck; b) SMF stem, anteverted neck; c) SMF stem, retroverted neck; d) Synergy stem, no modular neck options. Note that the stems are inserted in neutral alignment (no anteversion), and the apparent anteversion/retroversion in b) and c) are only due to the modular neck itself.
2.2. Phase 1: Mechanical experimentation 2.2.1. Implant insertion Fifteen large left 4th generation synthetic femurs (Model #3406, Sawbones, Vashon, WA, USA) were potted in cement blocks in 7° of adduction to simulate the single-leg stance of walking [17,19]. The femurs had an intramedullary canal with a 16 mm diameter, a solid cancellous foam with a 0.27 g/cm3 density; and a cortical bone with a 1.64 g/cm3 density. The working length of the femurs was 330 mm from top of the cement block to the apex of the greater trochanter. Three femurs were left unimplanted to serve as control specimens. Six femurs were implanted with a size 5 Short Modular Femoral (SMF) hip stem (Smith & Nephew, Cordova, TN, USA): 3 in neutral alignment and 3 in maximum anteversion. Six femurs were implanted with a size 17 Synergy hip stem (Smith & Nephew): 3 in neutral alignment and 3 in maximum anteversion. Implant sizes were selected by preoperative templating by an orthopaedic resident and visually confirmed by a second observer. Measurements and photographs of the implants were confirmed as acceptable by an experienced orthopaedic surgeon. 2.2.2. Implant configuration Five implant configurations were initially considered for testing: 3 modular necks for the SMF system, 1 monolithic Synergy stem,
and 1 intact femur. For the SMF system, the “Standard” modular neck produced no additional anteversion (i.e. “SMF-Neutral”) in the intact left femur, the “Left” modular neck produced additional anterior offset in the intact left femur (i.e. “SMF-Anteverted”), and the “Right” modular neck produced posterior offset in the intact left femur (i.e. “SMF-Retroverted”) (Fig. 1). Consequently, only stem and neck combinations which produced an overall neutral or anteverted head position were used in testing. Thus, 7 configurations were tested in total: 1) intact femur; 2) SMF-Neutral, neutral stem; 3) SMF-Anteverted, neutral stem; 4) SMF-Neutral, anteverted stem; 5) SMF-Retroverted, anteverted stem; 6) Synergy, neutral stem; 7) Synergy, anteverted stem. Each configuration had 3 specimens tested 3 times each. 2.2.3. Strain gauges Each femur was instrumented with 5 linear strain gauges and 1 rosette strain gauge (Models #125UW and #062UR, Vishay Precision Group, Malvern, PA, USA). For the SMF-implanted specimens, 3 linear gauges were mounted along the lateral cortex at 1.5, 3.5, and 6 inches below the widest point of the greater trochanter, thus corresponding to the middle of the stem, the tip of the stem, and just distal to tip of the stem. Two linear gauges were mounted on the medial cortex to be located at the same height as the 2
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to an 8-channel Cronos-PL data acquisition unit (IMC MeβSysteme GmbH, Berlin, Germany). This unit was connected to a laptop running IMC Device Control Software V2.6, which collected and stored the data for later analysis using FAMOS V5.0 software (IMC MeβSysteme GmbH). Strain data was collected at 10 Hz, and stored in 30-s segments. 2.2.4. Loading regime Femurs were secured distally to an angle vice mounted in a mechanical tester (Model 8874, Instron, Norwood, MA, USA). Three load angles were tested which represent a full gait cycle of singleleg-stance loading: flexion (15°), neutral (0°), and extension (−15°). Axial compression (preload = 50 N, rate = 25, 50, 75, or 100 N/s, max = 250, 500, 750, and 1000 N) was applied to the femoral head and maintained for 90 s to collect strain gauge data. These subclinical loads prevented specimen damage. Load was applied via a flat plate in order to transmit vertical load only by allowing lateral sliding of the femoral head under the plate; this produced pure bending and torsion in the femoral shaft. An acetabular cup-shaped load applicator was avoided, since it might have constrained the femoral head to produce buckling (rather than bending) of the femoral shaft and also a moment in the load cell. While cup loading has been used [19,20], others acknowledge its drawbacks and use flat plates instead [17,18,21]. 2.3. Phase 2: Finite element modeling
Fig. 2. Strain gauge placement pattern for Phase 1 experiments.
distal lateral gauges. The rosette gauge was mounted in the calcar region, directly anterior to the lesser trochanter (Fig. 2). Strain gauges were mounted in identical locations for the unimplanted control femurs. For the 6 Synergy-implanted femurs, strain gauges were located in the same pattern as above, but at different locations: 3.5, 7.5, and 9.5 inches below the greater trochanter. Wire leads were soldered to the strain gauge contacts and connected
2.3.1. CAD models A previously validated FE model of synthetic femur geometry was used (Table 1) [20,22,23]. FE models of SMF and Synergy stems were created using a HandySCAN 3D laser scanner (Creaform Inc., Levis, Quebec, Canada) (Table 1). The scanned files were imported into Geomagic Studio 12 (Geomagic, Morrisville, NC, USA) where the model geometry was virtually repaired. SMF implant head and neck geometries were created in Solidworks 2013 (Dassault Systemes, Waltham, MA, USA) using caliper measurements and dimensional drawings available from the implant manufacturer [24]. Head and neck geometries were merged with stem geometry in Geomagic, creating a single part for each implant configuration. 2.3.2. Material properties Material properties were assigned to the various solid bodies as follows (Table 1) [25]: Femurs were modeled as having cortical bone (Young’s Modulus, 16.7 GPa, Poisson’s ratio, 0.26) and cancellous bone (Young’s Modulus, 155 MPa, Poisson’s ratio,
Table 1 FE model geometric dimensions and material properties. Feature
Sawbone femur
SMF implant
Synergy implant
Max length (mm) Max diameter (mm)
485 52 (head) 37 (neck) 32 (shaft) Glass fibers in epoxy (cortical) Polyurethane (cancellous)
135 28 (head) 13.5 (neck) 38 (stem) CoCrMo (head) CoCrMo (neck) Ti alloy (stem) 240 (head) 240 (neck) 114 (stem) 0.30 (head) 0.30 (neck) 0.30 (stem) 8589
223 28 (head) 12.9 (neck) 37 (stem) CoCrMo (head) CoCrMo (neck) Ti alloy (stem) 240 (head) 240 (neck) 114 (stem) 0.30 (head) 0.30 (neck) 0.30 (stem) 15,345
13,387
23,490
Material
Young’s Modulus (GPa)
16.7 (cortical) 0.155 (cancellous)
Poisson’s Ratio (–)
0.26 (cortical) 0.30 (cancellous)
FE Elements
55,939 (intact) 52,344 (with SMF implant) 48,137 (with Synergy implant) 89,650 (intact) 84,531 (with SMF implant) 77,143 (with Synergy implant)
FE Nodes
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0.3). Implant stems were modeled as being made from titanium alloy (ASTM F1472) (Young’s Modulus, 114 GPa, Poisson’s ratio, 0.3). Implant heads and necks were modeled as being made from cobalt-chromium-molybdenum alloy (ASTM F799) (Young’s Modulus, 240 GPa, Poisson’s ratio, 0.3). All specimen components were modeled to behave as linear, elastic, homogenous, and isotropic materials [20,23,26]. 2.3.3. Elements and nodes All contact between bodies was “bonded” to simulate full bony ongrowth around implants. All bodies were meshed with patchconforming tetrahedral 10-node elements (SOLID187) with a maximum size of 4 mm. A mesh relevance of 80% was determined by iterative simulations to produce convergence in cortical bone Von Mises peak strain of less than 1% in successive simulations. The total number of elements and nodes, respectively, depended on the particular item being modeled (Table 1). 2.3.4. Simulating experiments Femurs and implants were modeled using 14 adjustable design parameters to account for variations in implant orientation. Parameters were measured for each experimental specimen using calipers or photographs using MB-Ruler (Markus Bader, www. markus-bader.de) and verified to ±0.5° or ±0.5 mm. Femur models were imported into Ansys Workbench 15.0 (Ansys Inc., Canonsburg, PA, USA) and virtually positioned. Distally, femur condyles were removed and constrained to replicate the cement potting blocks using 2 design parameters (i.e. axial working length and adduction angle). Proximally, femurs were osteotomized to remove proximal bone, followed by implant positioning using body transformations in 3 translational (superior– inferior, anterior–posterior, and medial–lateral) and 2 rotational (anteversion–retroversion and varus–valgus) design parameters. Boolean subtraction was used to remove cortical and cancellous bone from regions occupied by the implant. All strain gauges were modeled using 1-mm square surfaces projected onto the cortical bone surface to match locations of the experimental strain gauges. Specifically, linear strain gauges along the medial and lateral femur were given an anterior–posterior design parameter. This helped match patch location to the strain gauge center by measuring the strain gauge’s position relative to the anterior–posterior femoral shaft surface on medial or lateral photographs using MB-Ruler. These measurements varied by <3 mm across specimens and were entered directly into FE software to locate the strain gauge patch. Superior–inferior positioning had <0.5 mm variation across experimental specimens, measured by Vernier caliper from the greater trochanter and confirmed with medial and lateral photographs measured with MBRuler. Similarly, rosette gauges were given superior–inferior and medial–lateral design parameters. Anterior photographs were measured using MB-Ruler, landmarked from the greater trochanter level (superior–inferior measurement) and medial and lateral shaft surface (medial–lateral measurement). Measurements varied by <4 mm across specimens. Finally, a vertical vector load was applied directly to the implant femoral head, or to a patch on the superior surface of the intact femoral head for 250, 500, 750, and 1000 N. The load vector’s direction could be varied to match the flexion, neutral, or extension stances used experimentally. Strain data were then gathered from the 6 strain gauge locations on the femurs to compare to the experimental strain gauge data. 2.4. Phase 3: Stress shielding analysis and optimization 2.4.1. Finite element modeling For 5 implant configurations (intact femur, SMF-Neutral, SMFAnteverted, SMF-Retroverted, and Synergy), 3 important regions
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were defined. The calcar region was the primary region of interest, since prior studies have shown the highest stress shielding to occur here [6–8]. The medial femur was also chosen, since preliminary simulations revealed intact femurs to have a stress concentration in the calcar region, while implanted femurs showed a distal shift of this stress concentration, which could be monitored more easily in the medial femur. The “overall femur” was also considered for the intact femur and all implant configurations, since this allowed stress to be compared in the same volume of bone. Load was applied as in Phase 2, except at a clinical-level of 3000 N, i.e. 4 × body weight of a 75 kg patient, as done previously [19,27]. 2.4.2. Design of experiments (DOE) strategy Two design parameters were selected for DOE analysis across the 3 SMF configurations (i.e. anteversion–retroversion angle and anterior–posterior position). The other 5 design parameters from Phase 2 were held constant as they were irrelevant to the implant (i.e. femoral working length and adduction), or indicated by the manufacturer to be held constant (implant superior–inferior, medial–lateral, and varus–valgus) [24]. The two parameters were assigned levels which produce physiologically-possible implant orientations, where the implant does not excessively impinge on the cortical bone or protrude through the outer cortical surface. The range for anteversion was 0°–15°, relative to the femur’s existing angle of torsion. In the synthetic femur geometry, this angle is 12°, making the total anteversion 12°–27° with respect to the knee’s transcondylar axis. The remaining anterior–posterior ranges were found by trial-and-error, and varied based on anteversion. Once the allowable range was established for each parameter, 5 DOE levels were evenly distributed through this range. Since 2 inputs were tested at 5 levels each, the total number of experiments for each SMF stem was 25. For the SMF-Anteverted and SMF-Retroverted implants (± 6° anteversion caused by the modular neck), stem anteversion ranges were chosen such that both the implant stem and the resulting head position were within the 0°–15° overall anteversion envelope. Due to the large size of the Synergy stem, the range of anterior–posterior positions was small enough to eliminate this parameter, leaving only anteversion to be measured at 5 levels. The intact femur required no DOE. 2.4.3. DOE analysis of stress shielding The degree of stress shielding caused by an implant can be measured as a change in stress magnitude, with respect to an intact femur [10,18,28]. Four outcome measures were collected from each DOE run: 1) Mean Calcar Stress, which is an average of equivalent Von Mises stress in the calcar region nodes, indicating overall load transmitted to this region; 2) Standard Deviation of Calcar Stress, which is the standard deviation of equivalent Von Mises stress across calcar region nodes, approximately indicating the range of stress magnitudes and the presence or absence of stress concentrations; 3) Peak Stress Location, which is the distal shift in load distribution along the medial shaft; and 4) Peak Overall Stress, which is the maximum stress across the entire femoral cortical bone, indicating stress concentrations and potential failure sites. Mean and Standard Deviation of stress were calculated as percent changes in stress from the intact femur. Each outcome measure was collected for loading in 3 phases of gait (flexion, neutral, and extension). These results were combined, creating 4 aggregate stress results for each implant orientation and neck configuration. These aggregate results were used as the primary analysis of stress shielding, both in analyzing various orientations and comparing entire implants. Analysis was performed in Minitab 16 (Minitab Inc., State College, PA, USA), using ANOVA at a 95% confidence interval. Pairwise comparisons were calculated using a General Linear Model, and p-values were calculated using the Tukey test; statistical significance was set at p = 0.05.
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2.4.4. DOE optimization of implant orientation DOE results were used as an input for Minitab’s “Result Optimizer” function. For each implant configuration, the optimal level of anteversion and anterior–posterior offset was calculated, which reduced the change in calcar stress (Mean and Standard Deviation outputs; target = 0%, lower = −100%, upper = 100%, weight = 1) while keeping stress below the fracture limit (Peak Overall Stress output; target = 106 MPa [29], upper = 500 MPa, weight = 10). Peak Stress Location was not used, since variation was only present between implants. These 3 inputs were used for each of the 3 stances (flexion, neutral, and extension) to find a single orientation for each implant that optimized results across all stances. 3. Results 3.1. FE model validation (Phases 1 and 2) Correlation plots from the FE vs. experimental strain data are shown (Fig. 3). Correlation was strong for all models (Intact femur: slope = 0.898, R = 0.943; SMF implants: slope = 0.731, R = 0.948; Synergy implants: slope = 0.743, R = 0.859). Correction factors were calculated (as 1/slope) to adjust FE model data to unity versus experimental data. 3.2. Calcar stress shielding (Phase 3) Boxplots were first generated using one-way ANOVA in Minitab, P-values for pairwise comparisons were calculated with the Tukey test at 95% confidence interval, and post-hoc two-tailed power analysis of each pairwise comparison yielded 100% statistical power for all significantly different pairs. Primary DOE analysis of mean stress show greater stress shielding in the Synergy implant than in all SMF implants with all pairwise comparisons significantly different (p > 0.001), except for SMF-Anteverted and SMFRetroverted (p > 0.05) (Figs. 4 and 5). Similar results were obtained for standard deviation of stress. Calcar stress concentration was shifted distally for both implants (SMF: 38-47 mm, Synergy: 141 mm). Overall peak stresses were below the 106 MPa ultimate tensile strength of cortical bone in neutral stance, but above this limit when loaded in flexion and extension. 3.3. Optimal implant orientation (Phase 3) The “Response Optimizer” function yielded stem angles (SMFNeutral: 0.0°; SMF-Anteverted: 0.0°; SMF-Retroverted: 15.0°, Synergy: 10.3°) to minimize change in stress and maintain safe peak stresses. Anteverted and retroverted necks produced an additional ±6° of anteversion, when measured from the lateral aspect of the stem. Anterior–posterior positioning was unlikely to be considered clinically relevant owing to varying patient anatomy. To determine the magnitude of this anteversion optimization effect, 5 levels of anteversion for each implant were compared using the Tukey test for pairwise comparisons (Fig. 4). No significant differences among any anteversion level were present, except for SMF-Neutral at 15.0° compared to 0°–7.5° (p < 0.05). 4. Discussion 4.1. Overall stress maps Several aspects of the stress magnitude and distribution for the cases examined are worth highlighting. The current intact femur showed peak compressive stress of 50–60 MPa on the inferior side of the neck, indicating that this is an area of high risk for fracture
Fig. 3. FE versus experimental strain validation plots of a) intact femur, b) SMFimplanted femurs, and c) Synergy-implanted femurs.
during over-loading (Fig. 5). This is consistent with a prior thermographic imaging study of synthetic intact femurs during 5 Hz cyclic vertical loading at 1500, 1800, or 2100 N with the femur in 15° adduction, which indicated peak stresses were on the inferior neck and could reach 91, 96, and 104 MPa [19]. Similarly, a previous investigation of quasi-static vertical load-to-failure tests on the same model of the intact synthetic femur showed failure always occurred through the neck [30]. However, at present, once a hip implant was inserted, regardless of type or orientation, peak stresses shifted distally toward the level of the hip stem’s tip because of the impingement of the metal tip against the surrounding bone, thus creating a stress riser (Fig. 5). This is consistent with a prior biomechanical examination of short-stem hip stems, which illustrated that experimental strains increased at the level of the hip stem tip compared to more proximal regions of bone [18].
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Fig. 4. Average change in mean calcar stress (%) compared to the intact femur. A greater decrease indicates greater stress shielding. Error bars indicate ±1 standard deviation. Individual levels of anteversion are given (0° to 15°) as well as an aggregate average. A preliminary statistical analysis showed no significant difference in calcar stress (with exception of “Short Stem, Neutral Neck” at 15° versus 0°– 7.50°) caused by anteversion angle; this justified combining all anteversion points into aggregate averages. Each (non-aggregate) point represents n = 5 simulations for Short Stem and n = 1 simulation for Standard Stem configurations. Asterisks above Short Stem Anteverted and Retroverted aggregates indicate no significant difference between these configurations (p > 0.05); all other pairwise comparisons yielded significant differences (p < 0.001).
Fig. 5. Calcar Von Mises stress (MPa) for anterior (top) anterior and medial (bottom) views in a) Intact, b) SMF-Neutral, c) SMF-Anteverted, d) SMF-Retroverted, and e) Synergy implanted femurs.
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Finally, the current stress peak FE stresses were predicted to be 60 MPa (Fig. 5), which are far below the ultimate tensile stress of synthetic cortical bone (i.e. 106 MPa) [31], indicating that intact and implanted femurs would never be in danger of fracture at the applied loads used. 4.2. Calcar stress shielding Calcar stress was significantly reduced in the short-stem SMF-implanted femurs and further reduced in the conventional Synergy-implanted femur (Figs. 4 and 5). Direct evidence for stress shielding caused by reduced calcar stress has been shown in previous biomechanical studies [18,28,32–34]. Indirect evidence has been suggested for stress shielding by the 8%–24% reductions in bone mineral density in standard length stemmed hip implant patients at 2 years post-surgery [8,21]; however, long-term results for short-stem implants have not yet been published. Moreover, SMFAnteverted and SMF-Retroverted necks created significantly greater resistance to stress shielding than the SMF-Neutral implant, due to the additional 4 mm anterior–posterior and 6 mm medial offset in head position. This confirms a previous study in modular neck selection [35], which found higher strains in longer or anteverted necks than in shorter or neutral necks. 4.3. Implant stem anteversion The secondary goal of this study was to determine the effect of implant anteversion on stress shielding in the calcar region (Fig. 4). The results of the optimization procedure were not statistically important, since pairwise comparisons revealed very few statistically significant differences in stress shielding caused by anteversion for an implant type. As well, the effect of anteversion is not as large as that of implant selection, since there is a large degree of overlap in data groups of the same implant. These conclusions match a previous study on implant anteversion in short-stem implants [34], which concluded that small changes in implant anteversion had no significant effect on cortical strains. 4.4. Clinical relevance Several findings of clinical relevance are worth highlighting. First, the short-stem SMF implant with 3 of its various modular necks produced stress shielding at about half the magnitude of the conventional long-stem Synergy implant. Second, modular necks that increase the implant head’s distance from the stem (e.g. the SMF-Anteverted and SMF-Retroverted necks) present no disadvantage in terms of stress shielding. This may be particularly important for patients with unique anatomies who may require anteverted, retroverted, or otherwise offset modular necks. Third, current predictions of long-term stress shielding may be helpful to orthopaedic surgeons and biomechanical engineers because there are currently no long-term clinical analyses evaluating the success of short-stem implants. Fourth, current stress results could potentially be correlated with BMD measurements, as previously investigated [21,36,37]. If successful, this correlation could yield more accurate predictors of long-term bone integrity. It must be cautioned that too high a stress level, especially in the calcar, may result in short-term destabilization and micromotion of the implant. Fifth, the degree of anteversion had no significant effect on stress shielding, regardless of the implant selected. However, this was only applicable to a 0°–15° anteversion range (in addition to the femur’s natural 12° anteversion with respect to the knee) and assumed the implant is inserted according to manufacturer instructions.
Finally, one possible disadvantage to reducing stress shielding is the loss of a clinical indicator of osseointegration. When an implant is fully osseointegrated and stable, some stress shielding will occur over time, which can be detected via radiographs. If an implant is not osseointegrated, no stress shielding will occur, though this may ultimately lead to aseptic loosening. Therefore, some degree of stress shielding is beneficial as an indicator of proper fixation. 4.5. Limitations Synthetic femurs in present experiments cannot fully predict the in-vivo implantation outcomes or bony ongrowth over time in human femurs. However, synthetic bones have been validated previously against human cadaveric bone and have enjoyed increasing use in biomechanical studies by reducing inter-specimen variation [20]. Experimental tests were carried out at subclinical loads of 250, 500, 750, and 1000 N, which is acceptable for the following reasons. Subclinical loads were used only to experimentally validate the stresses (and strains) generated by the FE model, rather than to provide absolute stresses (and strains) for higher clinical-level loads. But, the final FE stresses (and strains) used to evaluate the various implants, in fact, were generated using the realistic clinical-level load of 3000 N, i.e. 4 x body weight of a 75 kg patient. Moreover, since the FE model is linearly elastic, therefore, the slope and linearity results from the validation graphs up to 1000 N (Fig. 3) apply equally to the final FE analysis at 3000 N. Furthermore, subclinical load experiments prevented gross damage to the physical specimens, since specimens needed to be tested multiple times to generate enough strain data for FE validation purposes. Finally, the above strategy is commonly used in similar biomechanics studies [20,23,38–40]. The change in hip contact force orientation and magnitude were not examined in experiments or FE models due to change in femoral head center caused by stem anteversion and medialized offset. Rather, hip contact force was applied only vertically at a fixed magnitude to simulate the stance phase of walking, as often done to simplify experiments and FE computations in other biomechanical hip implant studies [38–42]. In reality, with respect to an intact femur’s head center during walking, a medialized hip implant offset can produce 3D hip contact force components that are increased by 3% (medial), while simultaneously reduced by 15% (anterior) and 2% (superior), causing more cortical strain on the anterior, posterior, medial, and lateral surfaces in the proximal femur down to the stem tip [34]. But, stem anteversion can produce 3D hip contact force components that are increased by 11% (medial), 23% (anterior), and 6% (superior), causing increased medial and lateral cortical strain, but not medial and lateral cortical strain [34]. However, even with such considerations, the current overall comparative trends between configurations would have remained similar. Muscle forces were not replicated for experiments or FE models which would reduce superior neck stresses, but this is a common and acceptable simplification in biomechanical studies that are primarily focused on comparative analysis of a series of implant/bone constructs. Varus–valgus orientation was currently not investigated for experiments or FE models, since it has already been well-established biomechanically that some valgus angulation of bone conserving hip implants is preferable with respect to the native femoral neck angle [41–45]. Thus, it was felt that varus–valgus analysis would not have yielded any novel results at present. Fully bonded contact was chosen to reduce the time required for FE solution convergence, thus the current model replicated full bony ongrowth around the hip implant, rather than the immediate
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post-operative situation. This has also been done in some previous studies [5,28]. The FE model did not produce a 1:1 correspondence to experimental strain gauge data. However, the reasonable slopes (= 0.731– 0.898) and strong correlations R (= 0.859–948) of the data allow minor slope correction factors to be used. More importantly, the relative trends in stress shielding would still be maintained even for perfect 1:1 correlation. 5. Conclusions This study presents the first biomechanical analysis of stress shielding, stem anteversion, anterior–posterior translation, and modular neck options for the SMF short-stem hip system versus a conventional long-stem implant. Mechanical tests on synthetic femurs validated an FE model used to analyze a range of implant stem anteversion angles and neck options. The intact femur case showed peak stress occurring at the inferior neck, suggesting that any over-loading would begin to cause injury in this region first. However, once femurs were implanted with a hip stem, regardless of type or configuration, the peak stress shifted distally towards the level of the hip stem tip, indicating that this region was now more vulnerable to failure than the inferior neck. Moreover, while stress shielding was not eliminated entirely, short-stem implants significantly reduced stress shielding in the calcar region compared to conventional long-stem implants. Changes in implant stem anteversion and anterior–posterior translation had no significant effect on stress shielding. These results have important implications for implant selection in the young patient undergoing stemmed total hip arthroplasty. Author contributions statement Peter Goshulak – Study concept, experimentation, FE modeling and analysis, design of experiments, data interpretation, and article preparation. Saeid Samiezadeh – Design, analysis, and interpretation of FE model. Mina S.R. Aziz – Specimen preparation, and design and execution of experimental analysis. Habiba Bougherara – Design, analysis, and interpretation of FE model. Radovan Zdero – Study concept, FE model and design of experiments, data interpretation, and article preparation. Emil H. Schemitsch – Supervised first author, study concept, FE model and design of experiments, data interpretation, and article preparation. Funding Mr. and Mrs. Frank and Barbara Milligan Graduate Fellowship, the University of Toronto, and St. Michael’s Hospital. Ethical approval Not required. Competing interests None declared. References [1] Polkowski G, Gallaghan J, Mont M, Clohisy J. Total hip arthroplasty in the very young patient. J Am Acad Orthop Surg 2012;20(8):487–97.
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[2] Launay F, Jouve J, Guillaume J, Viehweger E, Jacquemier M, Bollini G. Total hip arthroplasty without cement in children and adolescents: 17 cases. Revue de Chirurgie Orthopedique et Reparatrice de l’Appareil Moteur 2002;88(5): 460–466. [3] Smith & Nephew, “SMF Short Modular Femoral Hip,” Smith & Nephew Corporate, 2011. [Online]. Available: http://www.smith-nephew.com/key-products/ orthopaedic-reconstruction/smf-short-modular-femoral-hip/ [4] Mulholland S, Wyss U. Activities of daily living in non-Western cultures: range of motion requirements for hip and knee joint implants. Int J Rehabil Res 2001;24(3):191–8. [5] Huiskes R, Weinans H, Rietbergen BV. The relationship between stress shielding and bone resorption around total hip stems and the effects of flexible materials. Clin Orthop Relat Res 1992;274:124–34. [6] Kroger H, Venesmaa P, Jurvelin J, Miettinen H, Suomalainen O, Alhava E. Bone density at the proximal femur after total hip arthroplasty. Clin Orthop Relat Res 1998;352:66–74. [7] Engh C, McGovern T, Bobyn J, Harris W. A quantitative evaluation of periprosthetic bone-remodeling after cementless total hip arthroplasty. J Bone Joint Surg (Am) 1992;74(7):1009–20. [8] Karachalios T, Tsatsaronis C, Efraimis G, Papadelis P, Lyritis G, Diakoumopoulos G. The long-term clinical relevance of calcar atrophy caused by stress shielding in total hip arthroplasty. J Arthroplast 2004;19(4):469–75. [9] Lombardi AVJ, Berend KR, Ng VY. Stubby stems: good things come in small packages. Orthopedics 2011;34(9):464–6. [10] Pettersen SH, Wik TS TS, Skallerud B. Subject specific finite element analysis of stress shielding around a cementless femoral stem. Clin Biomech 2009;24(2):196–202. [11] Wang YS, Liu M, Li JW, Hao YJ, Li JF, Yang J, et al. Application of CFP short-stem prosthesis in the treatment of osteonecrosis of the femoral head. Zhonghua Yi Xue Za Zhi 2011;91(47):3320–3. [12] Braun A, Sabah A. Two-year results of a modular short hip stem prosthesis – a prospective study. Z für Orthopädie und Unfallchirurgie 2009;147(6):700–6. [13] Soderman P, Malchau H. Is the Harris hip score useful to study the outcome of total hip replacement? Clin Orthop Relat Res 2001;384:189–97. [14] Amenabar T, Marimuthu K, Hawdon G, Gildone A, McMahon S. Total hip arthroplasty using a short-stem prosthesis: restoration of hip anatomy. J Orthop Surg 2015;23(1):90–4. [15] Dettmer M, Pourmoghaddam A, Kreuzer SW. Comparison of patient-reported outcome from neck-preserving, short-stem arthroplasty and resurfacing arthroplasty in younger osteoarthritis patients. Adv Orthop 2015:1–7 ID# 817689. [16] Von Lewinski G, Floerkemeier T. 10-year experience with short stem total hip arthroplasty. Orthopedics 2015;38(3):S51–6. [17] Bieger R, Ignatius A, Reichel H, Durselen L. Biomechanics of a short stem: in vitro primary stability and stress shielding of a conservative cementless hip stem. J Orthop Res 2013;31(8):1180–6. [18] Gronewold J, Berner S, Olender G, Hurschler C, Windhagen H, von Lewinski G, et al. Changes in strain patterns after implantation of a short stem with metaphyseal anchorage compared to a standard stem: and experimental study in synthetic bone. Orthop Rev 2014;6(5211):25–32. [19] Shah S, Bougherara H, Schemitsch E, Zdero R. Biomechanical stress maps of an artificial femur obtained using a new infrared thermography technique validated by strain gauges. Med Eng Phys 2012;34(10):1496–502. [20] Papini M, Zdero R, Schemitsch EH, Zalzal P. The biomechanics of human femurs in axial and torsional loading: comparison of finite element analysis, human cadaveric femurs, and synthetic femurs. J Biomech Eng 2007;129(1):12– 19. [21] Ike H, Inaba Y, Kobayashi N, Hirata Y, Yukizawa Y, Aoki C, et al. Comparison between mechanical stress and bone mineral density in the femur after total hip arthroplasty by using subject-specific finite element analyses. Comput Methods Biomech Biomed Eng 2015;8(10):1056–65. [22] M. Papini, and P. Zalzal, 2013. [Online]. Available: http://www.tecno.ior.it/ VRLAB/researchers/repository/BEL_repository.html. [23] Cheung G, Zalzal P, Bhandari M, Spelt J, Papini M. Finite element analysis of a femoral retrograde intramedullary nail subject to gait loading. Med Eng Phys 2004;26(2):93–108. [24] Smith & Nephew Inc., “SMF surgical technique,” 2011. [Online]. Available: www.smith-nephew.com/global/assets/pdf/products/surgical/71381476_ rev0_smf_stem_st_0610_lores(1).pdf. [25] Matweb, 2014. [Online]. Available: http://www.matweb.com/search/DataSheet. aspx?MatGUID=42cd25dc7f414bfcb4432a8bcb969889. [26] Cheal E, Spector M, Hayes W. Role of loads and prosthesis material properties on the mechanics of the proximal femur after total hip arthroplasty. J Orthop Res 1992;10(3):405–22. [27] Bergmann G, Deuretzbacher G, Heller M, Graichen F, Rohlmann A, Strauss J. Hip contact forces and gait patterns from routine activities. J Biomech 2001;34(7):859–71. [28] Baharuddin M, Salleh S, Zulkifly A, Lee M, Noor A, Harris A, et al. Design process of cementless femoral stem using a nonlinear three dimensional finite element analysis. BMC Musculoskelet Disord 2014;15:30. [29] MatWeb, 2014. [Online]. Available: http://www.matweb.com/search/DataSheet. aspx?MatGUID=14f152aec09e43e8a7861742be6c6cd3. [30] Nicayenzi B, Shah S, Schemitsch EH, Bougherara H, Zdero R. The biomechanical effect of changes in cancellous bone density on synthetic femur behavior. Proc Inst Mech Eng (Part H): J Eng Med 2011;225(11):1050–60. [31] Sawbone Worldwide, http://www.sawbones.com
240
P. Goshulak et al. / Medical Engineering and Physics 38 (2016) 232–240
[32] Pettersen S, Wik T, Skallerud B. Subject specific finite element analysis of stress shielding around a cementless femoral stem. Clin Biomech 2009;24(2):196–202. [33] Yamako Y, Chosa E, Zhao X, Totoribe K, Watanabe S, Sakamoto T, et al. Loadtransfer analysis after insertion of cementless anatomical femoral stem using pre- and post-operative CT images based patient-specific finite element analysis. Med Eng Phys 2014;36(6):694–700. [34] Speirs A, Heller M, Taylor W, Duda G, Perka C. Influence of changes in stem positioning on femoral loading after THR using a short-stemmed hip implant. Clin Biomech 2007;22(4):431–9. [35] Politis A, Siogkas G, Gelalis I, Xenakis T. Patterns of stress distribution at the proximal femur after implantation of a modular neck prosthesis. A biomechanical study. Clin Biomech 2013;28(4):415–22. [36] Herrera A, Panisello J, Ibarz E, Cegonino J, Puertolas J, Gracia L. Comparison between DEXA and finite element studies in the long-term bone remodeling of an anatomical femoral stem. J Biomech Eng 2009;131(4):041013-1-10. [37] Kwon J, Naito H, Matsumoto T, Tanaka M. Estimation of change of bone structures after total hip replacement using bone remodeling simulation. Clin Biomech 2013;28(5):514–18. [38] Ebrahimi H, Rabinovich M, Vuleta V, Zalcman D, Shah S, Dubov A, et al. Biomechanical properties of an intact, injured, repaired, and healed femur: an experimental and computational study. J Mech Behav Biomed Mater 2012;16:121– 35.
[39] Shah S, Kim SYR, Dubov A, Schemitsch EH, Bougherara H, Zdero R. The biomechanics of plate fixation of periprosthetic femoral fractures near the tip of a total hip implant: cables, screws, or both? Proc Inst Mech Eng (Part H): J Eng Med 2011;225(9):845–56. [40] Bougherara H, Zdero R, Dubov A, Shah S, Khurshid S, Schemitsch EH. A preliminary biomechanical study of a novel carbon-fibre hip implant versus standard metallic hip implants. Med Eng Phys 2011;33(1):121–8. [41] Anglin C, Masri B, Tonetti J, Hodgson A, Greidanus N. Hip resurfacing femoral neck fracture influenced by valgus placement. Clin Orthop Relat Res 2007;465:71–9. [42] Davis E, Olsen M, Zdero R, Waddell J, Schemitsch E. Femoral neck fracture following hip resurfacing: the effect of alignment of the femoral component. J Bone Jt Surg (Br) 2008;90(11):1522–7. [43] Nabavi A, Yeoh K, Shidiac L, Appleyard R, Gillies R, Turnbull A. Effects of positioning and notching of resurfaced femurs on femoral neck strength: a biomechanical test. J Orthop Surg (Hong Kong) 2009;17(1):47–50. [44] Olsen M, Lewis P, Waddell J, Schemitsch E. A biomechanical investigation of implant alignment and femoral neck notching with the Birmingham Mid-Head Resection. J Arthroplast 2010;25(6 (Suppl)):112–17. [45] Radcliffe I, Taylor M. Investigation into the effect of varus–valgus orientation on load transfer in the resurfaced femoral head: a multi-femur finite element analysis. Clin Biomech 2007;22(7):780–6.
Contents lists available at ScienceDirect
Medical Engineering and Physics journal homepage: www.elsevier.com/locate/medengphy
The biomechanical effect of anteversion and modular neck offset on stress shielding for short-stem versus conventional long-stem hip implants Peter Goshulak a,c, Saeid Samiezadeh b, Mina S.R. Aziz c, Habiba Bougherara b, Radovan Zdero b,c,d,∗, Emil H. Schemitsch a,c a
Institute of Biomaterials and Biomedical Engineering, University of Toronto, Toronto M5S 3G9, ON, Canada Department of Mechanical and Industrial Engineering, Ryerson University, Toronto M5B 2K3, ON, Canada c Martin Orthopaedic Biomechanics Laboratory, St. Michael’s Hospital, Toronto M5B 1W8, ON, Canada d Department of Mechanical and Industrial Engineering, University of Toronto, Toronto M5S 3G8, ON, Canada b
a r t i c l e
i n f o
Article history: Received 2 May 2015 Revised 22 October 2015 Accepted 6 December 2015
Keywords: Biomechanics Stress shielding Short-stem Hip arthroplasty Finite element analysis
a b s t r a c t Short-stem hip implants are increasingly common since they preserve host bone stock and presumably reduce stress shielding by improving load distribution in the proximal femur. Stress shielding may lead to decreased bone density, implant loosening, and fracture. However, few biomechanical studies have examined short-stem hip implants. The purpose of this study was to compare short-stem vs. standard length stemmed implants for stress shielding effects due to anteversion–retroversion, anterior–posterior position, and modular neck offset. Twelve artificial femurs were implanted with either a short-stem modular-neck implant or a conventional length monolithic implant in 0° or 15° of anteversion. Three modular neck options were tested in the short-stem implants. Three control femurs remained intact. Femurs were mounted in adduction and subjected to axial loading. Strain gauge values were collected to validate a Finite Element (FE) model, which was used to simulate the full range of physiologically possible anteversion and anterior–posterior combinations (n = 25 combinations per implant). Calcar stress was compared between implants and across each implant’s range of anteversion using one and two-way ANOVA. Stress shielding was defined as the overall change in stress compared to an intact femur. The FE model compared well with experimental strains (intact: slope = 0.898, R = 0.943; short-stem: slope = 0.731, R = 0.948; standard-stem: slope = 0.743, R = 0.859); correction factors were used to adjust slopes to unity. No implant anteversion showed significant reduction in stress shielding (α = 0.05, p > 0.05). Stress shielding was significantly higher in the standard-stem implant (63% change from intact femur, p < 0.001) than in short-stem implants (29–39% change, p < 0.001). Short-stem implants reduce stress shielding compared to standard length stemmed implants, while implant anteversion and anterior–posterior position had no effect. Therefore, short-stem implants have a greater likelihood of maintaining calcar bone strength in the long term. © 2015 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction Hip replacement surgery is becoming more prevalent in younger populations [1,2]. In order to facilitate possible future revision surgeries in young patients, maximum bone stock must be conserved [3]. Maintaining bone strength has always been a concern for load-bearing prostheses, but an increasingly active patient population [4] increases the demand for implant designs which maximize the load-bearing capabilities of the bone. As bone re∗
Corresponding author. Tel.: +1 416 953 5328. E-mail address: [email protected] (R. Zdero).
http://dx.doi.org/10.1016/j.medengphy.2015.12.005 1350-4533/© 2015 IPEM. Published by Elsevier Ltd. All rights reserved.
models itself in response to stress, regions of poor stress distribution become resorbed [5]. This “stress shielding” in the femur occurs primarily in the calcar region [6,7]. This decreased bone density may lead to aseptic loosening and stem migration, as well as periprosthetic fractures [6,8,9]. Implants should aim to replicate the physiological stress distribution, or that of an intact femur [10]. To this end, short-stem implants were relatively recently developed to conserve more bone stock in order to reduce stress shielding. These implants have good short-term clinical results such as reduced fracture and revision rate compared to conventional implants [9,11–13]. Specifically, for example, recent clinical studies on short-stem hip implants have shown certain advantages: vertical
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center of rotation, horizontal center of rotation, and leg length of the implanted hip vs. non-operated contralateral hip were within 5 mm, respectively, for 89%, 80%, and 96% of patients [14]; shortstem hips implants vs. hip resurfacing implants demonstrated 40% shorter surgery time and similar pre- and post-surgery outcomes for HOOS (hip disability and osteoarthritis outcome score) [15]; and revision rates over a 10-year follow-up period due to aseptic loosening, undersized stems, periprosthetic fracture, or via falsa are required in less than 2% of cases [16]. Similarly, recent biomechanical studies have shown several benefits compared to traditional hip stems: short-stem hip implants are more resistant to axial migration and retroversion measured at 100,000 cycles [17], whilst short-stem implants with predominantly metaphyseal fixation can provide cortical strain patterns in the proximal femur (including the calcar) more similar to an intact femur thereby potentially reducing “stress shielding” and improving implant stability [18]. However, to the authors’ knowledge, the biomechanical effects on stress shielding of a short-stem hip implant’s position in the host bone and modular neck selection have not been investigated. The goal of this experimental and computational study was to compare stress shielding between a short-stem implant versus a conventional length stemmed implant by analyzing all theoretically possible combinations of implant orientations plus clinicallyfeasible modular neck options. 2. Methods 2.1. General strategy Three study phases were used to quantify the stress shielding caused by surgical anteversion, anterior–posterior translation, and modular neck selection. Phase 1 involved mechanical experimentation on synthetic femurs to establish strain profiles for all implant orientations and modular neck combinations over a range of physiological load conditions. Phase 2 involved constructing a Finite Element (FE) model of the experimental setup and using the strain data to validate this model. Phase 3 used “Design of Experiments” on the validated FE model to analyze stress shielding for all possible implant orientations and modular neck options. This allowed clinical recommendations to be made about optimal surgical use of a typical short-stem hip implant.
Fig. 1. Superior view of the 4 implant and neck combinations inserted into a left femur: a) SMF stem, neutral neck; b) SMF stem, anteverted neck; c) SMF stem, retroverted neck; d) Synergy stem, no modular neck options. Note that the stems are inserted in neutral alignment (no anteversion), and the apparent anteversion/retroversion in b) and c) are only due to the modular neck itself.
2.2. Phase 1: Mechanical experimentation 2.2.1. Implant insertion Fifteen large left 4th generation synthetic femurs (Model #3406, Sawbones, Vashon, WA, USA) were potted in cement blocks in 7° of adduction to simulate the single-leg stance of walking [17,19]. The femurs had an intramedullary canal with a 16 mm diameter, a solid cancellous foam with a 0.27 g/cm3 density; and a cortical bone with a 1.64 g/cm3 density. The working length of the femurs was 330 mm from top of the cement block to the apex of the greater trochanter. Three femurs were left unimplanted to serve as control specimens. Six femurs were implanted with a size 5 Short Modular Femoral (SMF) hip stem (Smith & Nephew, Cordova, TN, USA): 3 in neutral alignment and 3 in maximum anteversion. Six femurs were implanted with a size 17 Synergy hip stem (Smith & Nephew): 3 in neutral alignment and 3 in maximum anteversion. Implant sizes were selected by preoperative templating by an orthopaedic resident and visually confirmed by a second observer. Measurements and photographs of the implants were confirmed as acceptable by an experienced orthopaedic surgeon. 2.2.2. Implant configuration Five implant configurations were initially considered for testing: 3 modular necks for the SMF system, 1 monolithic Synergy stem,
and 1 intact femur. For the SMF system, the “Standard” modular neck produced no additional anteversion (i.e. “SMF-Neutral”) in the intact left femur, the “Left” modular neck produced additional anterior offset in the intact left femur (i.e. “SMF-Anteverted”), and the “Right” modular neck produced posterior offset in the intact left femur (i.e. “SMF-Retroverted”) (Fig. 1). Consequently, only stem and neck combinations which produced an overall neutral or anteverted head position were used in testing. Thus, 7 configurations were tested in total: 1) intact femur; 2) SMF-Neutral, neutral stem; 3) SMF-Anteverted, neutral stem; 4) SMF-Neutral, anteverted stem; 5) SMF-Retroverted, anteverted stem; 6) Synergy, neutral stem; 7) Synergy, anteverted stem. Each configuration had 3 specimens tested 3 times each. 2.2.3. Strain gauges Each femur was instrumented with 5 linear strain gauges and 1 rosette strain gauge (Models #125UW and #062UR, Vishay Precision Group, Malvern, PA, USA). For the SMF-implanted specimens, 3 linear gauges were mounted along the lateral cortex at 1.5, 3.5, and 6 inches below the widest point of the greater trochanter, thus corresponding to the middle of the stem, the tip of the stem, and just distal to tip of the stem. Two linear gauges were mounted on the medial cortex to be located at the same height as the 2
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to an 8-channel Cronos-PL data acquisition unit (IMC MeβSysteme GmbH, Berlin, Germany). This unit was connected to a laptop running IMC Device Control Software V2.6, which collected and stored the data for later analysis using FAMOS V5.0 software (IMC MeβSysteme GmbH). Strain data was collected at 10 Hz, and stored in 30-s segments. 2.2.4. Loading regime Femurs were secured distally to an angle vice mounted in a mechanical tester (Model 8874, Instron, Norwood, MA, USA). Three load angles were tested which represent a full gait cycle of singleleg-stance loading: flexion (15°), neutral (0°), and extension (−15°). Axial compression (preload = 50 N, rate = 25, 50, 75, or 100 N/s, max = 250, 500, 750, and 1000 N) was applied to the femoral head and maintained for 90 s to collect strain gauge data. These subclinical loads prevented specimen damage. Load was applied via a flat plate in order to transmit vertical load only by allowing lateral sliding of the femoral head under the plate; this produced pure bending and torsion in the femoral shaft. An acetabular cup-shaped load applicator was avoided, since it might have constrained the femoral head to produce buckling (rather than bending) of the femoral shaft and also a moment in the load cell. While cup loading has been used [19,20], others acknowledge its drawbacks and use flat plates instead [17,18,21]. 2.3. Phase 2: Finite element modeling
Fig. 2. Strain gauge placement pattern for Phase 1 experiments.
distal lateral gauges. The rosette gauge was mounted in the calcar region, directly anterior to the lesser trochanter (Fig. 2). Strain gauges were mounted in identical locations for the unimplanted control femurs. For the 6 Synergy-implanted femurs, strain gauges were located in the same pattern as above, but at different locations: 3.5, 7.5, and 9.5 inches below the greater trochanter. Wire leads were soldered to the strain gauge contacts and connected
2.3.1. CAD models A previously validated FE model of synthetic femur geometry was used (Table 1) [20,22,23]. FE models of SMF and Synergy stems were created using a HandySCAN 3D laser scanner (Creaform Inc., Levis, Quebec, Canada) (Table 1). The scanned files were imported into Geomagic Studio 12 (Geomagic, Morrisville, NC, USA) where the model geometry was virtually repaired. SMF implant head and neck geometries were created in Solidworks 2013 (Dassault Systemes, Waltham, MA, USA) using caliper measurements and dimensional drawings available from the implant manufacturer [24]. Head and neck geometries were merged with stem geometry in Geomagic, creating a single part for each implant configuration. 2.3.2. Material properties Material properties were assigned to the various solid bodies as follows (Table 1) [25]: Femurs were modeled as having cortical bone (Young’s Modulus, 16.7 GPa, Poisson’s ratio, 0.26) and cancellous bone (Young’s Modulus, 155 MPa, Poisson’s ratio,
Table 1 FE model geometric dimensions and material properties. Feature
Sawbone femur
SMF implant
Synergy implant
Max length (mm) Max diameter (mm)
485 52 (head) 37 (neck) 32 (shaft) Glass fibers in epoxy (cortical) Polyurethane (cancellous)
135 28 (head) 13.5 (neck) 38 (stem) CoCrMo (head) CoCrMo (neck) Ti alloy (stem) 240 (head) 240 (neck) 114 (stem) 0.30 (head) 0.30 (neck) 0.30 (stem) 8589
223 28 (head) 12.9 (neck) 37 (stem) CoCrMo (head) CoCrMo (neck) Ti alloy (stem) 240 (head) 240 (neck) 114 (stem) 0.30 (head) 0.30 (neck) 0.30 (stem) 15,345
13,387
23,490
Material
Young’s Modulus (GPa)
16.7 (cortical) 0.155 (cancellous)
Poisson’s Ratio (–)
0.26 (cortical) 0.30 (cancellous)
FE Elements
55,939 (intact) 52,344 (with SMF implant) 48,137 (with Synergy implant) 89,650 (intact) 84,531 (with SMF implant) 77,143 (with Synergy implant)
FE Nodes
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0.3). Implant stems were modeled as being made from titanium alloy (ASTM F1472) (Young’s Modulus, 114 GPa, Poisson’s ratio, 0.3). Implant heads and necks were modeled as being made from cobalt-chromium-molybdenum alloy (ASTM F799) (Young’s Modulus, 240 GPa, Poisson’s ratio, 0.3). All specimen components were modeled to behave as linear, elastic, homogenous, and isotropic materials [20,23,26]. 2.3.3. Elements and nodes All contact between bodies was “bonded” to simulate full bony ongrowth around implants. All bodies were meshed with patchconforming tetrahedral 10-node elements (SOLID187) with a maximum size of 4 mm. A mesh relevance of 80% was determined by iterative simulations to produce convergence in cortical bone Von Mises peak strain of less than 1% in successive simulations. The total number of elements and nodes, respectively, depended on the particular item being modeled (Table 1). 2.3.4. Simulating experiments Femurs and implants were modeled using 14 adjustable design parameters to account for variations in implant orientation. Parameters were measured for each experimental specimen using calipers or photographs using MB-Ruler (Markus Bader, www. markus-bader.de) and verified to ±0.5° or ±0.5 mm. Femur models were imported into Ansys Workbench 15.0 (Ansys Inc., Canonsburg, PA, USA) and virtually positioned. Distally, femur condyles were removed and constrained to replicate the cement potting blocks using 2 design parameters (i.e. axial working length and adduction angle). Proximally, femurs were osteotomized to remove proximal bone, followed by implant positioning using body transformations in 3 translational (superior– inferior, anterior–posterior, and medial–lateral) and 2 rotational (anteversion–retroversion and varus–valgus) design parameters. Boolean subtraction was used to remove cortical and cancellous bone from regions occupied by the implant. All strain gauges were modeled using 1-mm square surfaces projected onto the cortical bone surface to match locations of the experimental strain gauges. Specifically, linear strain gauges along the medial and lateral femur were given an anterior–posterior design parameter. This helped match patch location to the strain gauge center by measuring the strain gauge’s position relative to the anterior–posterior femoral shaft surface on medial or lateral photographs using MB-Ruler. These measurements varied by <3 mm across specimens and were entered directly into FE software to locate the strain gauge patch. Superior–inferior positioning had <0.5 mm variation across experimental specimens, measured by Vernier caliper from the greater trochanter and confirmed with medial and lateral photographs measured with MBRuler. Similarly, rosette gauges were given superior–inferior and medial–lateral design parameters. Anterior photographs were measured using MB-Ruler, landmarked from the greater trochanter level (superior–inferior measurement) and medial and lateral shaft surface (medial–lateral measurement). Measurements varied by <4 mm across specimens. Finally, a vertical vector load was applied directly to the implant femoral head, or to a patch on the superior surface of the intact femoral head for 250, 500, 750, and 1000 N. The load vector’s direction could be varied to match the flexion, neutral, or extension stances used experimentally. Strain data were then gathered from the 6 strain gauge locations on the femurs to compare to the experimental strain gauge data. 2.4. Phase 3: Stress shielding analysis and optimization 2.4.1. Finite element modeling For 5 implant configurations (intact femur, SMF-Neutral, SMFAnteverted, SMF-Retroverted, and Synergy), 3 important regions
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were defined. The calcar region was the primary region of interest, since prior studies have shown the highest stress shielding to occur here [6–8]. The medial femur was also chosen, since preliminary simulations revealed intact femurs to have a stress concentration in the calcar region, while implanted femurs showed a distal shift of this stress concentration, which could be monitored more easily in the medial femur. The “overall femur” was also considered for the intact femur and all implant configurations, since this allowed stress to be compared in the same volume of bone. Load was applied as in Phase 2, except at a clinical-level of 3000 N, i.e. 4 × body weight of a 75 kg patient, as done previously [19,27]. 2.4.2. Design of experiments (DOE) strategy Two design parameters were selected for DOE analysis across the 3 SMF configurations (i.e. anteversion–retroversion angle and anterior–posterior position). The other 5 design parameters from Phase 2 were held constant as they were irrelevant to the implant (i.e. femoral working length and adduction), or indicated by the manufacturer to be held constant (implant superior–inferior, medial–lateral, and varus–valgus) [24]. The two parameters were assigned levels which produce physiologically-possible implant orientations, where the implant does not excessively impinge on the cortical bone or protrude through the outer cortical surface. The range for anteversion was 0°–15°, relative to the femur’s existing angle of torsion. In the synthetic femur geometry, this angle is 12°, making the total anteversion 12°–27° with respect to the knee’s transcondylar axis. The remaining anterior–posterior ranges were found by trial-and-error, and varied based on anteversion. Once the allowable range was established for each parameter, 5 DOE levels were evenly distributed through this range. Since 2 inputs were tested at 5 levels each, the total number of experiments for each SMF stem was 25. For the SMF-Anteverted and SMF-Retroverted implants (± 6° anteversion caused by the modular neck), stem anteversion ranges were chosen such that both the implant stem and the resulting head position were within the 0°–15° overall anteversion envelope. Due to the large size of the Synergy stem, the range of anterior–posterior positions was small enough to eliminate this parameter, leaving only anteversion to be measured at 5 levels. The intact femur required no DOE. 2.4.3. DOE analysis of stress shielding The degree of stress shielding caused by an implant can be measured as a change in stress magnitude, with respect to an intact femur [10,18,28]. Four outcome measures were collected from each DOE run: 1) Mean Calcar Stress, which is an average of equivalent Von Mises stress in the calcar region nodes, indicating overall load transmitted to this region; 2) Standard Deviation of Calcar Stress, which is the standard deviation of equivalent Von Mises stress across calcar region nodes, approximately indicating the range of stress magnitudes and the presence or absence of stress concentrations; 3) Peak Stress Location, which is the distal shift in load distribution along the medial shaft; and 4) Peak Overall Stress, which is the maximum stress across the entire femoral cortical bone, indicating stress concentrations and potential failure sites. Mean and Standard Deviation of stress were calculated as percent changes in stress from the intact femur. Each outcome measure was collected for loading in 3 phases of gait (flexion, neutral, and extension). These results were combined, creating 4 aggregate stress results for each implant orientation and neck configuration. These aggregate results were used as the primary analysis of stress shielding, both in analyzing various orientations and comparing entire implants. Analysis was performed in Minitab 16 (Minitab Inc., State College, PA, USA), using ANOVA at a 95% confidence interval. Pairwise comparisons were calculated using a General Linear Model, and p-values were calculated using the Tukey test; statistical significance was set at p = 0.05.
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2.4.4. DOE optimization of implant orientation DOE results were used as an input for Minitab’s “Result Optimizer” function. For each implant configuration, the optimal level of anteversion and anterior–posterior offset was calculated, which reduced the change in calcar stress (Mean and Standard Deviation outputs; target = 0%, lower = −100%, upper = 100%, weight = 1) while keeping stress below the fracture limit (Peak Overall Stress output; target = 106 MPa [29], upper = 500 MPa, weight = 10). Peak Stress Location was not used, since variation was only present between implants. These 3 inputs were used for each of the 3 stances (flexion, neutral, and extension) to find a single orientation for each implant that optimized results across all stances. 3. Results 3.1. FE model validation (Phases 1 and 2) Correlation plots from the FE vs. experimental strain data are shown (Fig. 3). Correlation was strong for all models (Intact femur: slope = 0.898, R = 0.943; SMF implants: slope = 0.731, R = 0.948; Synergy implants: slope = 0.743, R = 0.859). Correction factors were calculated (as 1/slope) to adjust FE model data to unity versus experimental data. 3.2. Calcar stress shielding (Phase 3) Boxplots were first generated using one-way ANOVA in Minitab, P-values for pairwise comparisons were calculated with the Tukey test at 95% confidence interval, and post-hoc two-tailed power analysis of each pairwise comparison yielded 100% statistical power for all significantly different pairs. Primary DOE analysis of mean stress show greater stress shielding in the Synergy implant than in all SMF implants with all pairwise comparisons significantly different (p > 0.001), except for SMF-Anteverted and SMFRetroverted (p > 0.05) (Figs. 4 and 5). Similar results were obtained for standard deviation of stress. Calcar stress concentration was shifted distally for both implants (SMF: 38-47 mm, Synergy: 141 mm). Overall peak stresses were below the 106 MPa ultimate tensile strength of cortical bone in neutral stance, but above this limit when loaded in flexion and extension. 3.3. Optimal implant orientation (Phase 3) The “Response Optimizer” function yielded stem angles (SMFNeutral: 0.0°; SMF-Anteverted: 0.0°; SMF-Retroverted: 15.0°, Synergy: 10.3°) to minimize change in stress and maintain safe peak stresses. Anteverted and retroverted necks produced an additional ±6° of anteversion, when measured from the lateral aspect of the stem. Anterior–posterior positioning was unlikely to be considered clinically relevant owing to varying patient anatomy. To determine the magnitude of this anteversion optimization effect, 5 levels of anteversion for each implant were compared using the Tukey test for pairwise comparisons (Fig. 4). No significant differences among any anteversion level were present, except for SMF-Neutral at 15.0° compared to 0°–7.5° (p < 0.05). 4. Discussion 4.1. Overall stress maps Several aspects of the stress magnitude and distribution for the cases examined are worth highlighting. The current intact femur showed peak compressive stress of 50–60 MPa on the inferior side of the neck, indicating that this is an area of high risk for fracture
Fig. 3. FE versus experimental strain validation plots of a) intact femur, b) SMFimplanted femurs, and c) Synergy-implanted femurs.
during over-loading (Fig. 5). This is consistent with a prior thermographic imaging study of synthetic intact femurs during 5 Hz cyclic vertical loading at 1500, 1800, or 2100 N with the femur in 15° adduction, which indicated peak stresses were on the inferior neck and could reach 91, 96, and 104 MPa [19]. Similarly, a previous investigation of quasi-static vertical load-to-failure tests on the same model of the intact synthetic femur showed failure always occurred through the neck [30]. However, at present, once a hip implant was inserted, regardless of type or orientation, peak stresses shifted distally toward the level of the hip stem’s tip because of the impingement of the metal tip against the surrounding bone, thus creating a stress riser (Fig. 5). This is consistent with a prior biomechanical examination of short-stem hip stems, which illustrated that experimental strains increased at the level of the hip stem tip compared to more proximal regions of bone [18].
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Fig. 4. Average change in mean calcar stress (%) compared to the intact femur. A greater decrease indicates greater stress shielding. Error bars indicate ±1 standard deviation. Individual levels of anteversion are given (0° to 15°) as well as an aggregate average. A preliminary statistical analysis showed no significant difference in calcar stress (with exception of “Short Stem, Neutral Neck” at 15° versus 0°– 7.50°) caused by anteversion angle; this justified combining all anteversion points into aggregate averages. Each (non-aggregate) point represents n = 5 simulations for Short Stem and n = 1 simulation for Standard Stem configurations. Asterisks above Short Stem Anteverted and Retroverted aggregates indicate no significant difference between these configurations (p > 0.05); all other pairwise comparisons yielded significant differences (p < 0.001).
Fig. 5. Calcar Von Mises stress (MPa) for anterior (top) anterior and medial (bottom) views in a) Intact, b) SMF-Neutral, c) SMF-Anteverted, d) SMF-Retroverted, and e) Synergy implanted femurs.
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Finally, the current stress peak FE stresses were predicted to be 60 MPa (Fig. 5), which are far below the ultimate tensile stress of synthetic cortical bone (i.e. 106 MPa) [31], indicating that intact and implanted femurs would never be in danger of fracture at the applied loads used. 4.2. Calcar stress shielding Calcar stress was significantly reduced in the short-stem SMF-implanted femurs and further reduced in the conventional Synergy-implanted femur (Figs. 4 and 5). Direct evidence for stress shielding caused by reduced calcar stress has been shown in previous biomechanical studies [18,28,32–34]. Indirect evidence has been suggested for stress shielding by the 8%–24% reductions in bone mineral density in standard length stemmed hip implant patients at 2 years post-surgery [8,21]; however, long-term results for short-stem implants have not yet been published. Moreover, SMFAnteverted and SMF-Retroverted necks created significantly greater resistance to stress shielding than the SMF-Neutral implant, due to the additional 4 mm anterior–posterior and 6 mm medial offset in head position. This confirms a previous study in modular neck selection [35], which found higher strains in longer or anteverted necks than in shorter or neutral necks. 4.3. Implant stem anteversion The secondary goal of this study was to determine the effect of implant anteversion on stress shielding in the calcar region (Fig. 4). The results of the optimization procedure were not statistically important, since pairwise comparisons revealed very few statistically significant differences in stress shielding caused by anteversion for an implant type. As well, the effect of anteversion is not as large as that of implant selection, since there is a large degree of overlap in data groups of the same implant. These conclusions match a previous study on implant anteversion in short-stem implants [34], which concluded that small changes in implant anteversion had no significant effect on cortical strains. 4.4. Clinical relevance Several findings of clinical relevance are worth highlighting. First, the short-stem SMF implant with 3 of its various modular necks produced stress shielding at about half the magnitude of the conventional long-stem Synergy implant. Second, modular necks that increase the implant head’s distance from the stem (e.g. the SMF-Anteverted and SMF-Retroverted necks) present no disadvantage in terms of stress shielding. This may be particularly important for patients with unique anatomies who may require anteverted, retroverted, or otherwise offset modular necks. Third, current predictions of long-term stress shielding may be helpful to orthopaedic surgeons and biomechanical engineers because there are currently no long-term clinical analyses evaluating the success of short-stem implants. Fourth, current stress results could potentially be correlated with BMD measurements, as previously investigated [21,36,37]. If successful, this correlation could yield more accurate predictors of long-term bone integrity. It must be cautioned that too high a stress level, especially in the calcar, may result in short-term destabilization and micromotion of the implant. Fifth, the degree of anteversion had no significant effect on stress shielding, regardless of the implant selected. However, this was only applicable to a 0°–15° anteversion range (in addition to the femur’s natural 12° anteversion with respect to the knee) and assumed the implant is inserted according to manufacturer instructions.
Finally, one possible disadvantage to reducing stress shielding is the loss of a clinical indicator of osseointegration. When an implant is fully osseointegrated and stable, some stress shielding will occur over time, which can be detected via radiographs. If an implant is not osseointegrated, no stress shielding will occur, though this may ultimately lead to aseptic loosening. Therefore, some degree of stress shielding is beneficial as an indicator of proper fixation. 4.5. Limitations Synthetic femurs in present experiments cannot fully predict the in-vivo implantation outcomes or bony ongrowth over time in human femurs. However, synthetic bones have been validated previously against human cadaveric bone and have enjoyed increasing use in biomechanical studies by reducing inter-specimen variation [20]. Experimental tests were carried out at subclinical loads of 250, 500, 750, and 1000 N, which is acceptable for the following reasons. Subclinical loads were used only to experimentally validate the stresses (and strains) generated by the FE model, rather than to provide absolute stresses (and strains) for higher clinical-level loads. But, the final FE stresses (and strains) used to evaluate the various implants, in fact, were generated using the realistic clinical-level load of 3000 N, i.e. 4 x body weight of a 75 kg patient. Moreover, since the FE model is linearly elastic, therefore, the slope and linearity results from the validation graphs up to 1000 N (Fig. 3) apply equally to the final FE analysis at 3000 N. Furthermore, subclinical load experiments prevented gross damage to the physical specimens, since specimens needed to be tested multiple times to generate enough strain data for FE validation purposes. Finally, the above strategy is commonly used in similar biomechanics studies [20,23,38–40]. The change in hip contact force orientation and magnitude were not examined in experiments or FE models due to change in femoral head center caused by stem anteversion and medialized offset. Rather, hip contact force was applied only vertically at a fixed magnitude to simulate the stance phase of walking, as often done to simplify experiments and FE computations in other biomechanical hip implant studies [38–42]. In reality, with respect to an intact femur’s head center during walking, a medialized hip implant offset can produce 3D hip contact force components that are increased by 3% (medial), while simultaneously reduced by 15% (anterior) and 2% (superior), causing more cortical strain on the anterior, posterior, medial, and lateral surfaces in the proximal femur down to the stem tip [34]. But, stem anteversion can produce 3D hip contact force components that are increased by 11% (medial), 23% (anterior), and 6% (superior), causing increased medial and lateral cortical strain, but not medial and lateral cortical strain [34]. However, even with such considerations, the current overall comparative trends between configurations would have remained similar. Muscle forces were not replicated for experiments or FE models which would reduce superior neck stresses, but this is a common and acceptable simplification in biomechanical studies that are primarily focused on comparative analysis of a series of implant/bone constructs. Varus–valgus orientation was currently not investigated for experiments or FE models, since it has already been well-established biomechanically that some valgus angulation of bone conserving hip implants is preferable with respect to the native femoral neck angle [41–45]. Thus, it was felt that varus–valgus analysis would not have yielded any novel results at present. Fully bonded contact was chosen to reduce the time required for FE solution convergence, thus the current model replicated full bony ongrowth around the hip implant, rather than the immediate
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post-operative situation. This has also been done in some previous studies [5,28]. The FE model did not produce a 1:1 correspondence to experimental strain gauge data. However, the reasonable slopes (= 0.731– 0.898) and strong correlations R (= 0.859–948) of the data allow minor slope correction factors to be used. More importantly, the relative trends in stress shielding would still be maintained even for perfect 1:1 correlation. 5. Conclusions This study presents the first biomechanical analysis of stress shielding, stem anteversion, anterior–posterior translation, and modular neck options for the SMF short-stem hip system versus a conventional long-stem implant. Mechanical tests on synthetic femurs validated an FE model used to analyze a range of implant stem anteversion angles and neck options. The intact femur case showed peak stress occurring at the inferior neck, suggesting that any over-loading would begin to cause injury in this region first. However, once femurs were implanted with a hip stem, regardless of type or configuration, the peak stress shifted distally towards the level of the hip stem tip, indicating that this region was now more vulnerable to failure than the inferior neck. Moreover, while stress shielding was not eliminated entirely, short-stem implants significantly reduced stress shielding in the calcar region compared to conventional long-stem implants. Changes in implant stem anteversion and anterior–posterior translation had no significant effect on stress shielding. These results have important implications for implant selection in the young patient undergoing stemmed total hip arthroplasty. Author contributions statement Peter Goshulak – Study concept, experimentation, FE modeling and analysis, design of experiments, data interpretation, and article preparation. Saeid Samiezadeh – Design, analysis, and interpretation of FE model. Mina S.R. Aziz – Specimen preparation, and design and execution of experimental analysis. Habiba Bougherara – Design, analysis, and interpretation of FE model. Radovan Zdero – Study concept, FE model and design of experiments, data interpretation, and article preparation. Emil H. Schemitsch – Supervised first author, study concept, FE model and design of experiments, data interpretation, and article preparation. Funding Mr. and Mrs. Frank and Barbara Milligan Graduate Fellowship, the University of Toronto, and St. Michael’s Hospital. Ethical approval Not required. Competing interests None declared. References [1] Polkowski G, Gallaghan J, Mont M, Clohisy J. Total hip arthroplasty in the very young patient. J Am Acad Orthop Surg 2012;20(8):487–97.
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[2] Launay F, Jouve J, Guillaume J, Viehweger E, Jacquemier M, Bollini G. Total hip arthroplasty without cement in children and adolescents: 17 cases. Revue de Chirurgie Orthopedique et Reparatrice de l’Appareil Moteur 2002;88(5): 460–466. [3] Smith & Nephew, “SMF Short Modular Femoral Hip,” Smith & Nephew Corporate, 2011. [Online]. Available: http://www.smith-nephew.com/key-products/ orthopaedic-reconstruction/smf-short-modular-femoral-hip/ [4] Mulholland S, Wyss U. Activities of daily living in non-Western cultures: range of motion requirements for hip and knee joint implants. Int J Rehabil Res 2001;24(3):191–8. [5] Huiskes R, Weinans H, Rietbergen BV. The relationship between stress shielding and bone resorption around total hip stems and the effects of flexible materials. Clin Orthop Relat Res 1992;274:124–34. [6] Kroger H, Venesmaa P, Jurvelin J, Miettinen H, Suomalainen O, Alhava E. Bone density at the proximal femur after total hip arthroplasty. Clin Orthop Relat Res 1998;352:66–74. [7] Engh C, McGovern T, Bobyn J, Harris W. A quantitative evaluation of periprosthetic bone-remodeling after cementless total hip arthroplasty. J Bone Joint Surg (Am) 1992;74(7):1009–20. [8] Karachalios T, Tsatsaronis C, Efraimis G, Papadelis P, Lyritis G, Diakoumopoulos G. The long-term clinical relevance of calcar atrophy caused by stress shielding in total hip arthroplasty. J Arthroplast 2004;19(4):469–75. [9] Lombardi AVJ, Berend KR, Ng VY. Stubby stems: good things come in small packages. Orthopedics 2011;34(9):464–6. [10] Pettersen SH, Wik TS TS, Skallerud B. Subject specific finite element analysis of stress shielding around a cementless femoral stem. Clin Biomech 2009;24(2):196–202. [11] Wang YS, Liu M, Li JW, Hao YJ, Li JF, Yang J, et al. Application of CFP short-stem prosthesis in the treatment of osteonecrosis of the femoral head. Zhonghua Yi Xue Za Zhi 2011;91(47):3320–3. [12] Braun A, Sabah A. Two-year results of a modular short hip stem prosthesis – a prospective study. Z für Orthopädie und Unfallchirurgie 2009;147(6):700–6. [13] Soderman P, Malchau H. Is the Harris hip score useful to study the outcome of total hip replacement? Clin Orthop Relat Res 2001;384:189–97. [14] Amenabar T, Marimuthu K, Hawdon G, Gildone A, McMahon S. Total hip arthroplasty using a short-stem prosthesis: restoration of hip anatomy. J Orthop Surg 2015;23(1):90–4. [15] Dettmer M, Pourmoghaddam A, Kreuzer SW. Comparison of patient-reported outcome from neck-preserving, short-stem arthroplasty and resurfacing arthroplasty in younger osteoarthritis patients. Adv Orthop 2015:1–7 ID# 817689. [16] Von Lewinski G, Floerkemeier T. 10-year experience with short stem total hip arthroplasty. Orthopedics 2015;38(3):S51–6. [17] Bieger R, Ignatius A, Reichel H, Durselen L. Biomechanics of a short stem: in vitro primary stability and stress shielding of a conservative cementless hip stem. J Orthop Res 2013;31(8):1180–6. [18] Gronewold J, Berner S, Olender G, Hurschler C, Windhagen H, von Lewinski G, et al. Changes in strain patterns after implantation of a short stem with metaphyseal anchorage compared to a standard stem: and experimental study in synthetic bone. Orthop Rev 2014;6(5211):25–32. [19] Shah S, Bougherara H, Schemitsch E, Zdero R. Biomechanical stress maps of an artificial femur obtained using a new infrared thermography technique validated by strain gauges. Med Eng Phys 2012;34(10):1496–502. [20] Papini M, Zdero R, Schemitsch EH, Zalzal P. The biomechanics of human femurs in axial and torsional loading: comparison of finite element analysis, human cadaveric femurs, and synthetic femurs. J Biomech Eng 2007;129(1):12– 19. [21] Ike H, Inaba Y, Kobayashi N, Hirata Y, Yukizawa Y, Aoki C, et al. Comparison between mechanical stress and bone mineral density in the femur after total hip arthroplasty by using subject-specific finite element analyses. Comput Methods Biomech Biomed Eng 2015;8(10):1056–65. [22] M. Papini, and P. Zalzal, 2013. [Online]. Available: http://www.tecno.ior.it/ VRLAB/researchers/repository/BEL_repository.html. [23] Cheung G, Zalzal P, Bhandari M, Spelt J, Papini M. Finite element analysis of a femoral retrograde intramedullary nail subject to gait loading. Med Eng Phys 2004;26(2):93–108. [24] Smith & Nephew Inc., “SMF surgical technique,” 2011. [Online]. Available: www.smith-nephew.com/global/assets/pdf/products/surgical/71381476_ rev0_smf_stem_st_0610_lores(1).pdf. [25] Matweb, 2014. [Online]. Available: http://www.matweb.com/search/DataSheet. aspx?MatGUID=42cd25dc7f414bfcb4432a8bcb969889. [26] Cheal E, Spector M, Hayes W. Role of loads and prosthesis material properties on the mechanics of the proximal femur after total hip arthroplasty. J Orthop Res 1992;10(3):405–22. [27] Bergmann G, Deuretzbacher G, Heller M, Graichen F, Rohlmann A, Strauss J. Hip contact forces and gait patterns from routine activities. J Biomech 2001;34(7):859–71. [28] Baharuddin M, Salleh S, Zulkifly A, Lee M, Noor A, Harris A, et al. Design process of cementless femoral stem using a nonlinear three dimensional finite element analysis. BMC Musculoskelet Disord 2014;15:30. [29] MatWeb, 2014. [Online]. Available: http://www.matweb.com/search/DataSheet. aspx?MatGUID=14f152aec09e43e8a7861742be6c6cd3. [30] Nicayenzi B, Shah S, Schemitsch EH, Bougherara H, Zdero R. The biomechanical effect of changes in cancellous bone density on synthetic femur behavior. Proc Inst Mech Eng (Part H): J Eng Med 2011;225(11):1050–60. [31] Sawbone Worldwide, http://www.sawbones.com
240
P. Goshulak et al. / Medical Engineering and Physics 38 (2016) 232–240
[32] Pettersen S, Wik T, Skallerud B. Subject specific finite element analysis of stress shielding around a cementless femoral stem. Clin Biomech 2009;24(2):196–202. [33] Yamako Y, Chosa E, Zhao X, Totoribe K, Watanabe S, Sakamoto T, et al. Loadtransfer analysis after insertion of cementless anatomical femoral stem using pre- and post-operative CT images based patient-specific finite element analysis. Med Eng Phys 2014;36(6):694–700. [34] Speirs A, Heller M, Taylor W, Duda G, Perka C. Influence of changes in stem positioning on femoral loading after THR using a short-stemmed hip implant. Clin Biomech 2007;22(4):431–9. [35] Politis A, Siogkas G, Gelalis I, Xenakis T. Patterns of stress distribution at the proximal femur after implantation of a modular neck prosthesis. A biomechanical study. Clin Biomech 2013;28(4):415–22. [36] Herrera A, Panisello J, Ibarz E, Cegonino J, Puertolas J, Gracia L. Comparison between DEXA and finite element studies in the long-term bone remodeling of an anatomical femoral stem. J Biomech Eng 2009;131(4):041013-1-10. [37] Kwon J, Naito H, Matsumoto T, Tanaka M. Estimation of change of bone structures after total hip replacement using bone remodeling simulation. Clin Biomech 2013;28(5):514–18. [38] Ebrahimi H, Rabinovich M, Vuleta V, Zalcman D, Shah S, Dubov A, et al. Biomechanical properties of an intact, injured, repaired, and healed femur: an experimental and computational study. J Mech Behav Biomed Mater 2012;16:121– 35.
[39] Shah S, Kim SYR, Dubov A, Schemitsch EH, Bougherara H, Zdero R. The biomechanics of plate fixation of periprosthetic femoral fractures near the tip of a total hip implant: cables, screws, or both? Proc Inst Mech Eng (Part H): J Eng Med 2011;225(9):845–56. [40] Bougherara H, Zdero R, Dubov A, Shah S, Khurshid S, Schemitsch EH. A preliminary biomechanical study of a novel carbon-fibre hip implant versus standard metallic hip implants. Med Eng Phys 2011;33(1):121–8. [41] Anglin C, Masri B, Tonetti J, Hodgson A, Greidanus N. Hip resurfacing femoral neck fracture influenced by valgus placement. Clin Orthop Relat Res 2007;465:71–9. [42] Davis E, Olsen M, Zdero R, Waddell J, Schemitsch E. Femoral neck fracture following hip resurfacing: the effect of alignment of the femoral component. J Bone Jt Surg (Br) 2008;90(11):1522–7. [43] Nabavi A, Yeoh K, Shidiac L, Appleyard R, Gillies R, Turnbull A. Effects of positioning and notching of resurfaced femurs on femoral neck strength: a biomechanical test. J Orthop Surg (Hong Kong) 2009;17(1):47–50. [44] Olsen M, Lewis P, Waddell J, Schemitsch E. A biomechanical investigation of implant alignment and femoral neck notching with the Birmingham Mid-Head Resection. J Arthroplast 2010;25(6 (Suppl)):112–17. [45] Radcliffe I, Taylor M. Investigation into the effect of varus–valgus orientation on load transfer in the resurfaced femoral head: a multi-femur finite element analysis. Clin Biomech 2007;22(7):780–6.