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The effect of abductor muscle and anterior-posterior hip contact load simulation on the in-vitro primary… Park, Youngbae; Albert, Carolyne; Yoon, Yong-San; Fernlund, Göran; Frei, Hanspeter; Oxland, Thomas R Jun 24, 2010

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Park et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Open AccessR E S E A R C H  A R T I C L EResearch articleThe effect of abductor muscle and anterior-posterior hip contact load simulation on the in-vitro primary stability of a cementless hip stemYoungbae Park*†1, Carolyne Albert†2, Yong-San Yoon1, Göran Fernlund3, Hanspeter Frei4 and Thomas R Oxland5AbstractBackground: In-vitro mechanical tests are commonly performed to assess pre-clinically the effect of implant design on the stability of hip endoprostheses. There is no standard protocol for these tests, and the forces applied vary between studies. This study examines the effect of the abductor force with and without application of the anterior-posterior hip contact force in the in-vitro assessment of cementless hip implant stability.Methods: Cementless stems (VerSys Fiber Metal) were implanted in twelve composite femurs which were divided into two groups: group 1 (N = 6) was loaded with the hip contact force only, whereas group 2 (N = 6) was additionally subjected to an abductor force. Both groups were subjected to the same cranial-caudal hip contact force component, 2.3 times body weight (BW) and each specimen was subjected to three levels of anterior-posterior hip contact load: 0, -0.1 to 0.3 BW (walking), and -0.1 to 0.6 BW (stair climbing). The implant migration and micromotion relative to the femur was measured using a custom-built system comprised of 6 LVDT sensors.Results: Substantially higher implant motion was observed when the anterior-posterior force was 0.6BW compared to the lower anterior-posterior load levels, particularly distally and in retroversion. The abductor load had little effect on implant motion when simulating walking, but resulted in significantly less motion than the hip contact force alone when simulating stair climbing.Conclusions: The anterior-posterior component of the hip contact load has a significant effect on the axial motion of the stem relative to the bone. Inclusion of the abductor force had a stabilizing effect on the implant motion when simulating stair climbing.BackgroundLoosening of femoral hip implants is a major problem intotal hip arthroplasty [1]. Clinical studies have shown thatearly implant migration negatively affects the long termperformance of cementless femoral stems [2-4]. Excessivemicromotion at the bone-implant interface inhibits suc-cessful bone ingrowth in cementless implants and maytherefore result in early implant loosening [5-7]. Theimmediate post operative migration and micromotion(primary stability) of different femoral stems have beenevaluated under simulated physiological loading in in-vitro experiments [8-12]. Although it has not yet beendemonstrated for cementless stems, some cementedstems with inferior clinical results have been shown toalso result in higher in-vitro micromotions [13], whichdemonstrates the clinical relevance of these in-vitro tests.The physiological loads acting on the head of a femoralstem have been established by telemetric measurementsfor daily activities such as walking and stair climbing [14-17], while the muscle forces for these activities have beenestimated by numerical models [15,18-20]. It is challeng-ing to include all hip contact and muscle forces acting onthe femur in an in-vitro test and simplified test setups* Correspondence: ybpark@kaist.ac.kr1 Department of Mechanical Engineering, Korean Advanced Institute of © 2010 Park et al; licensee BioMed Central Ltd. This is an Open Access article distributed under the terms of the Creative Commons At-tribution License (http://creativecommons.org/licenses/by/2.0), which permits unrestricted use, distribution, and reproduction in anymedium, provided the original work is properly cited.have therefore been used to simulate the biomechanicalScience and Technology, Daejeon, Republic of Korea† Contributed equallyFull list of author information is available at the end of the articlePark et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 2 of 14environment to which hip implants are subjected to post-operatively. Some in-vitro studies have simulated the hipcontact force alone [9,11,12,21-23], while others includedone [24,25] or many muscle forces [8,10]. However, it isnot clear how these variations affect stem migration andmicromotion.In particular, the precise effect of the abductor muscleload (Fabd) on the primary stability of uncemented stemshas not been demonstrated. Of all muscle groups, theabductors have been shown to have the most pronouncedeffect on femoral strains, increasing medial bending inthe proximal femur during gait [26-28]. There are, how-ever, contradictory results concerning the effects ofincluding muscle loading on primary stability, and thesestudies also incorporated more than one muscle groupsuch that the effect of the abductor muscles has not beenisolated. In an in-vitro study of cemented stems, the sim-ulation of muscle forces (abductor, vastus lateralis, andtensor fascia latae) resulted in a small and non-significantreduction in migration compared with the hip contactforce applied alone [8]. On the other hand, in another in-vitro study, the inclusion of muscle loads (abductor, ten-sor fascia latae, ilio-tibial tract, vastus lateralis and vastusmedialis) increased migration and micromotion of acementless stem [10]. We hypothesise that simulation ofan abductor muscle force increases implant micromotionand migration of cementless stems compared with hipcontact forces alone.The effect of the anterior-posterior component of thehip contact force (Fap) on implant primary stability hasalso not been established definitely. In-vitro studies havemeasured the torsional strength of cementless implantfixation [29-31] and these values were found to approachthe torque levels measured in-vivo during stair climbing[32] Physiological cranial-caudal loads, however, were notapplied in these in-vitro studies, which may underesti-mate the torsional strength of the stem-femur constructs.Studies have measured implant migration and micromo-tion under varying Fap loads [10,24]. One study reportedhigher distal migration and micromotion when simulat-ing stair climbing compared to walking loads [10],whereas the other did not observe a difference in distalmicromotion between stair climbing and single-legstance, a configuration without Fap [24]. These studies,however, also varied muscular loading such that the effectof Fap was not isolated. We hypothesise that the higher Fapload observed during stair climbing generates greaterimplant-bone micromotion and migration comparedwith walking.To test our hypotheses, we conducted in-vitro tests onMethodsA cementless femoral stem (VerSys collarless size 14,Zimmer Co., Dover, Ohio, USA) was implanted in twelvecomposite femurs (Model 3303, Third Generation,Pacific Research Laboratories, Vashon, Washington,USA). The femoral cavity was prepared manually accord-ing to the implant manufacturer's instructions, usingstraight reamers and broaches. Visual inspection of thecavity after preparation revealed that the regions of con-tact between the stem and the cortical component of thecomposite bones were consistent between specimens.The specimens were cut at 27 cm from the proximal endand the distal 6 cm were potted in dental stone (Tru-Stone, Heraeus Kulzer, Armonk, New York). The speci-mens were then loaded cyclically on a biaxial servohy-draulic testing machine (Instron Model 8874, Instron,Canton, Massachusetts). The loads applied were designedto mimic walking and stair climbing loads as measured byBergmann et al. [15].The specimens were divided into two groups for bio-mechanical testing. Group 1 (N = 6, Figure 1a-b) wasloaded with the hip contact force only. A cranial-caudalforce (Fcc) of 2.3 times body weight (BW) was applied bythe linear actuator, with the femur potted at 13° of adduc-tion (Figure 1a), generating a proximal-distal componentof 2.2 BW and a medial-lateral component of 0.5 BW. Abody weight of 75 kg was used for the simulations. Thepotted distal femur was fastened to a linear guide to avoida horizontal reaction force in the frontal plane. Group 2(N = 6, Figure 1c-d) was additionally loaded with anabductor muscle load (Fabd). The Fabd was applied with asteel cable using a lever that was joined to the actuatorthrough a hinge (Figure 1c). The steel cable was attachedto the greater trochanter through a custom-mouldedpolymethylmethacrylate (PMMA) cap. The cable passedthrough a copper tube that was embedded into thePMMA cap, and the cap was attached to the bone with a4 mm diameter steel pin inserted anterior-posteriorlythrough the greater trochanter. The same muscle attach-ment cap was used for all specimens to obtain a repeat-able muscle orientation relative to the femur. An Fabd of1.1 BW [20] was applied by adjusting the offset betweenthe actuator and the femoral head, doff, in proportion tothe muscle-to-femoral head lever arm, dm, see Figure 1c.The measured dm varied between 46 and 50 mm, and doffwas adjusted in proportion to dm to maintain the same Fccand Fabd values between specimens. Based on equilibriumcalculations (shown in Figure 2), the same Fcc orientationas group 1 was achieved for group 2 by potting the femursat 4° of abduction.composite femurs, in which we examined the effect of theabductor on the motion of a cementless implant at threelevels of anterior-posterior hip contact load.For both groups, the anterior-posterior hip contact load(Fap) was applied by the rotary actuator (M = Fap*doff). ForPark et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 3 of 14group 1 doff was 32 mm, whereas in group 2 it was set at0.83*dm, (and since dm ranged between 46 and 49 mm, dofftherefore ranged between 38 and 41 mm). The Fap wasapplied in three phases of 1000 cycles each. The first loadphase simulated walking without Fap (Fap = 0), the secondsimulated walking with Fap (Fap = -0.1 to 0.3 BW), and thethird simulated stair climbing (Fap = -0.1 to 0.6 BW).These peak Fap loads are based on published results of in-vivo measurements [15]. During stair climbing, an actua-tor rotation of approximately 1° in amplitude wasobserved in the muscle group. Based on the geometry ofthe implant and loading set-up, we estimate that thisrotation would have affected the orientation of theabductor load relative to the femur by approximately 1°.The applied peak loads for both groups are summarizedin Table 1. The loads were sinusoidal with a frequency of1 Hz with in-phase peak loads.The relative motion between stem and bone was mea-sured with a custom-built system similar to previouslypublished designs [33-35]. The system, illustrated in Fig-ure 3, was comprised of six linear variable differentialtransformers (LVDTs) mounted on a frame that was rig-surface of the implant through a hole in the cortex. Theimplant motion was calculated from the motion of thetriangle using a custom program implemented in Matlab(MathWorks, Natick, Massachussetts). The measurementresolution was smaller than 0.7 μm in all translationaldirections, and smaller than 0.001° in rotation. The accu-racy of the system in measuring translation was evaluatedagainst a micrometer precision dial gauge (Kafer, Ger-many). Translation along each of the three axes wasapplied to the implant, with the sensors attached to anover-reamed composite femur. The maximum translationerror observed was 2 μm over a range of 30 μm (mean 0.8μm, stdev 0.8 μm for 9 measurements), and 10 μm over arange of 300 μm (mean 5.6 μm, stdev 3.0 μm for 9 mea-surements). The accuracy of each sensor was also mea-sured with a dial gauge (Kafer, Germany), where amaximum error of 1.7 μm was observed over a range of200 μm (mean 0.6 μm, stdev 0.4 μm for 60 measure-ments). The rotation accuracy was evaluated analyticallyfrom the maximum individual LVDT errors, yielding amaximum rotation error of 0.0026°.Migration was defined as the difference in stem meanposition (translations and rotations) between cycle 100Figure 1 Loading set-ups. (a) Group 1 - no abductor, i.e. hip contact force alone. Axial and torsional loading of the actuator produced distal (Fd), me-dial (Fm) and anterior-posterior (Fap) loading of the femoral head due to the mounting geometry and the offset between the femoral head and the central axis of the actuator, doff (32 mm). (b) Resulting forces on the femur for group 1. (c) Group 2 - hip contact force and abductor. (d) Resulting forces on the femur for group 2 (equilibrium calculations are presented in Figure 2).(a) (b) (c) (d) doffbiaxial  actuatorload cell cappotting linear guide Instron table cable clips steel cable wire PMMA cap Fabd(1.1BW) Fcc(2.3BW) Fap(0 to 0.6BW)Fap(0 to 0.6BW) Fcc(2.3BW) 131334doffdm4FccFmFd13idly attached to the femur with seven set screws. The sen-sors measured the three dimensional motion of atriangular plate that was rigidly attached to the lateraland the last cycle of each loading step, i.e. cycle 1000 (Fap= 0), cycle 2000 (Fap = 0.3 BW) and cycle 3000 (Fap = 0.6BW), see Figure 4. The first 100 cycles were used for pre-Park et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 4 of 14Figure 2 Equilibrium calculations for group 2 (abductor).Where: Fa force applied by the linear actuator   Fcc  cranial-caudal hip contact force on the femoral head,     i.e. resultant of the distal and the medial force components   Fabd abductor force We want: Fcc = 2.3BW at 13° from the femur long axis, i.e. cc=13°- b  Fabd =1.1BW at 34° from the femur long axis, i.e. abd=34°- bEquilibrium on lever plate: Fx = 0    Fabd sin(34°-b) – Fcc sin(13°-b) = 0    ... b = -4° Fy = 0    Fabd cos(34°-b) + Fa – Fcc cos(13°-b) = 0    … Fa = 1.33BW M = 0 (with femoral head as reference point)    Fabd dm = Fa doff   … doff = 0.83 dmxyFaFccFabdccabddoffdmbabdFaFa MPark et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 5 of 14conditioning [10,12]. Micromotion was defined as theaverage reversible motion of the stem during the last 200cycles of each loading step, i.e. cycles 800-1000, 1800-2000, and 2800-3000 (Figure 4). The migration andmicromotion were each comprised of 6 components:translation along the medial, anterior and distal axes (atthe reference point shown in Figure 3), as well as rota-tions projected in the frontal, sagittal and transverseplanes. The resultants of the three translational migrationand micromotion components are presented as 'totaltranslational migration' and 'total translational micromo-tion'. Similarly, the terms 'total rotational migration' and'total rotational micromotion' were used to represent theresultant of all rotational components, and were definedas the rotations about the helical axis [36].The effects of Fabd and Fap on each migration and micro-motion component and their resultants were examinedwith a two-way ANOVA, with Fap as a repeated measure,followed by Student Newman Keuls post hoc analysiswith a significance level of 95%.ResultsThe implant-bone migration and micromotion compo-nents for both groups at all loading conditions are sum-marized in Tables 2 and 3. The resultants of thesecomponents, i.e. total translational and rotational migra-tions and micromotions, are presented in Figures 5 and 6.Migration occurred primarily along and about theimplant axis. Distal migration accounted for 94 to 99% ofthe total translational migration. The average absoluterotational migration was smaller than 0.04° in the sagittaland frontal planes, but much larger in the transverseplane (rotation about the implant axis) where it reachedStatistically, the abductor force Fabd did not have a sig-nificant main effect on the total translational migration (p= 0.13), however, the total translational micromotion andthe total rotational migration and micromotion were onaverage smaller with Fabd than without Fabd (p < 0.01), seeFigures 5 and 6. In contrast, the anterior-posterior hipcontact force component Fap had a clear significant maineffect on the total translational and rotational migrationsand micromotions (p < 0.01).There was, however, a strong interaction between theabductor and the Fap present in all the motion resultants(p < 0.01) and all the components (p < 0.05), except therotational migration in the frontal plane (p = 0.38). Ingeneral the abductor was only observed to affect theimplant motion at Fap 0.6 BW. With this Fap, all compo-nents of migration and micromotion were significantlygreater without the abductor (Tables 2 and 3). The onlymotion components that were significantly affected bythe abductor at all Fap levels were the rotational migrationin the frontal plane, opposite in direction between thetwo groups, and the translational micromotion in the lat-eral axis, which was smaller for the abductor group.Similarly, the effect of increasing Fap was mainly seen inthe no abductor group. Without the abductor, increasingFap from 0 to 0.3 BW increased the translational micro-motion only in the lateral direction (p < 0.02). IncreasingFap to 0.6 BW, however, led to significantly higher micro-motion in all directions (p ≤ 0.01), higher translationalmigration in all directions (p < 0.01), as well as higherrotational migration in the transverse plane (p < 0.01).With the abductor set-up, increasing Fap from 0 to 0.3 BWdid not significantly affect implant motion, and increas-Table 1: Loads applied to the hip systemLoading step Cycles Hip contact force (xBW) FabdFcc Fap(xBW)Group 1 No abductor 1 1-1000 0.4 to 2.3 0 n/a2 1001-2000 0.4 to 2.3 -0.1 to +0.3 n/a3 2001-3000 0.4 to 2.3 -0.1 to +0.6 n/aGroup 2 Abductor 1 1-1000 0.4 to 2.3 0 1.12 1001-2000 0.4 to 2.3 -0.1 to +0.3 1.13 2001-3000 0.4 to 2.3 -0.1 to +0.6 1.1Fcc is in the caudal direction, and a positive Fap is in the posterior direction. The peak loads (maximum and minimum) were defined based on published data [15], and scaled for a 70 kg individual.an average of -1.2° and 0.4° for groups 1 and 2, respec-tively. Micromotion, on the other hand, was generally notdominated by motion in a specific direction.ing the Fap to 0.6 BW only gave a significant increase intranslational migration in the lateral and distal directionsPark et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 6 of 14Figure 3 Motion measurement set-up. (a) LVDT set-up. (b) Coordinate system and sensor diagram. The reference point is located on the lateral side of the stem, 113 mm proximal from the stem tip. The arrows show the location and direction of each sensor.(a) (b) (c) XZPark et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 7 of 14(p < 0.05), and rotational migration in the sagittal plane (p< 0.01).DiscussionIn-vitro mechanical tests are commonly performed toassess the effect of implant design on the stability of hipendoprostheses pre-clinically. There is no standard pro-tocol for these tests, and the loading conditions used varygreatly. Efforts have been made to standardize the testconditions [37], however, it is not clear how the abductormuscle and the anterior-posterior hip contact force influ-ence the translational and rotational stability of theimplant. The present study examined the effect of thesetwo parameters in the in-vitro assessment of cementlesship implant primary stability.As any biomechanical investigation this study has somedetail in the following paragraphs. In addition, differentload magnitudes were applied in sequence to each speci-men,. To minimize this effect on subsequent migration,the study was designed such that the load magnitude wasapplied in increasing increments simulating postopera-tive rehabilitation. However, during a pilot test, themicromotion observed during simulated walking wassimilar whether these loads were applied before or afterthe stair climbing cycles.Composite femurs were used to minimize experimentalvariability, as was done in other studies for the same rea-son [13,23,38]. Their structural stiffness has been shownto approximate that of natural bone, but with less vari-ability [39,40]. No comprehensive study comparingimplant stability in composite versus cadaveric femurswas found in the literature, however, in-vitro tests withFigure 4 Distal movement of the stem relative to the bone. Micromotion was calculated as the average amplitude of the cyclic motion during the last 200 cycles of each loading step (Fap = 0, Fap = 0.3 BW, and Fap = 0.6 BW). Migration was the cumulative stem displacement at the end of each step, with respect to its position at cycle 100.0 500 1000 1500 2000 2500 30000100200300400500600cyclesmicrometerWalkingFap=0WalkingFap=0.3BWStair ClimbingFap=0.6BW100th cycleMigrationMicromotionMigrationMicromotionDistal displacement (m)CycleWalkingFap = 0WalkingFap = 0.3BWStair climbingFap = 0.6BWmicrometerDistal displacement (m)Distal displacement (m)limitations. Composite femurs were used instead ofhuman femurs, and the implant motion was measured atonly one location. These two limitations are discussed incomposite femurs [23] have yielded axial migration com-parable to cadaveric femurs [41] for the CLS and press-fitMuller implants.Park et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 8 of 14In our tests, the implant motion was measured at a sin-gle location. With the magnitude of physiological loadsapplied, the stem and the bone could not be consideredrigid bodies; therefore the motion at other locationscould not be determined from our experimental data.Some in-vitro studies have measured bone-implantmotion at multiple locations, as reviewed by Britton et al[42], but the individual measurements are often limitedto a single axis (e.g [23,43]). Experimentally, space restric-tions generally translate into having to choose betweenmeasuring three-dimensional motion at limited locationsand measuring uniaxial motion at several locations. Witha single axis motion measurement approach, however,rotational motions between the implant and the bone canfreedom motion measurement device enabled us to avoidsuch error, however, our motion measurements were lim-ited to one location.Two common testing set-ups were selected for thisstudy: the first set-up applied the hip contact force alonewhile the second applied the hip contact force togetherwith the abductor force. The abductor force is oftenincluded rather than other muscle groups because theabductors were demonstrated to have the most importanteffect of all muscle groups on stresses and strains in theproximal femur [26,28]. More complex set-ups have beenused in the literature, but they are less common. Forexample, in one study several muscle forces (abductor,ilio-tibial band, tensor fascia latae, vastus lateralis andTable 2: Migration resultsGroup 1 No abductor Group 2 AbductorComponent FapAverage 95% CI Average 95% CILateral translation (μm) 0 7 ± 4 5 ± 80.3 BW 17 ± 10 9 ± 130.6 BW 62 ab ± 18 20* a ± 14Anterior translation (μm) 0 -2 ± 7 0 ± 20.3 BW 7 ± 7 10 ± 200.6 BW 39 ab ± 26 19 * ± 16Distal translation (μm) 0 50 ± 28 63 ± 430.3 BW 100 ± 52 103 ± 670.6 BW 385 ab ± 147 191* ab ± 123Sagittal plane rotation (×10-3°) 0 12 ± 7 5 ± 50.3 BW 30 ± 13 14* ± 70.6 BW -8 ab ± 24 32* ab ± 22Frontal plane rotation (×10-3°) 0 15 ± 6 -2* ± 50.3 BW 34 ± 16 -10* ± 160.6 BW 25 ± 56 -14* ± 19Transverse plane rotation (×10-3°) 0 34 ± 53 43 ± 550.3 BW -19 ± 144 175 ± 890.6 BW -1175 ab ± 567 359* ± 172* p < 0.05 compared to other group at same Fap levela p < 0.05 compared to same group at Fap = 0 BWb p < 0.05 compared to same group at Fap = 0.3 BWincur large errors in translational motion measurement,which are proportional to the distance between the bone-implant interface and the sensor axis. A six-degree ofvastus medialis) were simulated with multiple indepen-dent actuators [10]. A set-up modeling the hip contactforce alone, on the other hand, is advocated for its sim-Park et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 9 of 14plicity and reproducibility. In a previous study [13], theuse of this simpler model was justified based on thereported small effect of muscles on cement stresses incemented constructs [28].Our measured distal migration/micromotion magni-tudes for the VerSys FMT stem (walking: ~100 μm/10 μmwith both set-ups; and stair climbing: 191 μm/8 μm and385 μm/16 μm with and without the abductor force,respectively) were within the range of values reported forother cementless implants tested in composite or cadaverfemurs. Distal migration/micromotion in the order of 150μm/10 μm, 70 μm/30 μm, and 400 μm/50 μm werereported in other studies [9,10,23] for the CLS stem, aage in the order of 0.7 mm at 6 months) than the reportedvalues from in-vitro experiments [2,44]. This may be inpart due to the limited number of gait cycles modeled in-vitro (usually 1000 or 5000 cycles) and/or the use of sim-pler and lower loads compared to those sometimes seenin-vivo, which may reach as high as eight times the bodyweight during stumbling, for example [45]. Furthermore,adaptation of the bone, i.e. remodelling and local boneresorption, may also affect post-operative implantmotion. In-vitro tests could at best simulate resorption bymilling the bone interface at a predetermined locationprior to testing [46]. Nonetheless, the objective of in-vitroprimary stability tests for cementless stems is not to pro-Table 3: Micromotion resultsGroup 1 No abductor Group 2 AbductorComponent FapAverage 95% CI Average 95% CILateral translation (μm) 0 7 ± 5 2* ± 20.3 BW 10 a ± 4 2* ± 30.6 BW 16 ab ± 5 2* ± 2Anterior translation (μm) 0 4 ± 5 0 ± 40.3 BW 7 ± 9 -4 ± 160.6 BW 25 ab ± 8 -1* ± 19Distal translation (μm) 0 11 ± 6 9 ± 30.3 BW 12 ± 6 9* ± 30.6 BW 16 ab ± 7 8* ± 2Sagittal plane rotation (×10-3°) 0 0 ± 1 3 ± 50.3 BW 5 ± 5 3 ± 60.6 BW 11 ab ± 9 1* ± 5Frontal plane rotation (×10-3°) 0 14 ± 6 10 ± 120.3 BW 20 ± 6 12 ± 160.6 BW 35 ab ± 11 10* ± 12Transverse plane rotation (×10-3°) 0 -0 ± 2 -1 ± 20.3 BW -6 ± 14 9 ± 160.6 BW -61 ab ± 16 14* ± 37* p < 0.05 compared to other group at same Fap levela p < 0.05 compared to same group at Fap = 0 BWb p < 0.05 compared to same group at Fap = 0.3 BWpress-fit cementless implant similarly intended for proxi-mal fixation. Stem migration measured clinically for theCLS stem, however, is substantially larger (with an aver-vide an estimate of in-vivo migration, but to ensure that afavourable environment for successful bone ingrowth willbe achieved post-operatively. It has been proposed thatPark et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 10 of 14Figure 5 Implant migration resultants as a function of Fap for each group. (top) Total translational migration, i.e. (medial2 + anterior2 + distal2)1/2. (bottom) Total rotational migration (about the helical axis). Results shown are means (N = 6) and 95% confidence intervals. * p < 0.05 compared to the other group at the same Fap value. a p < 0.05 compared to the same group at Fap = 0 BW. b p < 0.05 compared to the same group at Fap = 0.3 BW.Total translational migration01002003004005006000 0.3 0.6Fap (x BW)Migration (microns)Group 1- no abductorGroup 2 - abductorTotal rotational migration05001000150020000 0.3 0.6Fap (x BW)Rotation (x10-3 degrees)Group 1- no abductorGroup 2 - abductorFap (xBW)Fap (xBW)ab*abab*abMigration (microns)Rotation (x10-3 degrees)Migration (microns)Rotation (x10-3 degrees)Park et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 11 of 14Figure 6 Implant micromotion resultants as a function of Fap for each group. (a) Total translational micromotion (b) Total rotational micromotion (about the helical axis). Results shown are means (N = 6) and 95% confidence intervals. * p < 0.05 compared to the other group at the same Fap value. a p < 0.05 compared to the same group at Fap = 0 BW. b p < 0.05 compared to the same group at Fap = 0.3 BW.Total translational micromotion010203040500 0.3 0.6Fap (x BW)Micromotion (microns)Group 1 - no abductorGroup 2 - abductorTotal rotational micromotion0204060801001201400 0.3 0.6Fap (x BW)Rotation (x10-3 degrees)Group 1 - no abductorGroup 2 - abductorFap (xBW)Fap (xBW)ab*aab*aa*Micromotion (microns)Rotation (x10-3 degrees)Micromotion (microns)Rotation (x10-3 degrees)Park et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 12 of 14micromotion may be a better predictor than migrationfor the long-term performance of femoral implants [47],however, no clinical data was available to compare withour micromotion results.The high torsional Fap loads experienced by the proxi-mal femur during stair climbing are well documented andhave been shown to occur during other activities such asjogging, fast walking, and rising from a chair [32,48,49].Concerns have been raised that these forces may exceedthe stem's torsional fixation strength [32]. However, theseconcerns were based on comparisons with in-vitro tor-sional strength assessments obtained without cranical-caudal loading on the implant [29-31], which may haveunderestimated the torsional strength under more physi-ological loading. Torsional loading has been said to affectthe rotational motion of femoral hip implant [24,50]. Oneof these studies, however, did not apply a cranial-caudalload or measure the translational motion [49], while theother varied not only the torsional load applied, but alsothe muscle loads [24]. Our results indicate that for a col-larless, cementless implant, increasing Fap not onlyincreases the axial rotation of the implant but that themotion increases in other directions as well, particularlydistally. A similar finding was reported in another study,in which stair climbing loads generated approximately150 μm of distal migration, compared to 30 μm of proxi-mal migration when simulating walking loads for the CLSimplant [10]. In their study, however, the Fap (~200N, i.e.~0.3 BW for a 70 kg individual) was smaller than the val-ues reported for stair climbing in-vivo, i.e. 0.6 BW [32]and muscle forces also varied between their walking andstair climbing set-ups [10]. Moreover, proximal migrationwas observed under walking loads, which the authorsattributed to errors inherent in their motion measure-ment system. The current study, on the other hand,looked at the effect of Fap in isolation from other parame-ters. Increasing Fap from 0 to 0.3 BW did not have a signif-icant effect on implant motion, but a significant increasein migration (mainly in the distal direction) was observedwhen increasing Fap from 0.3 BW (walking) to 0.6 BW(stair climbing) - this effect was largest without theabductor. The micromotion also increased with increas-ing Fap (mainly in the anterior direction), but this effectwas only seen without the abductor. Rotation was pri-marily in the transverse plane, i.e. about the implant longaxis; without the abductor stair climbing produced onaverage 10 times higher rotational micromotion (Table 3)and 62 times higher rotational migration (Table 2) aboutthis axis compared to walking loads. Our results thereforesupport our first hypothesis: the higher Fap loadsWe found that inclusion of the abductor muscle forcestabilized the implant both in translation and rotation,particularly when simulating stair climbing. This doesnot support our second hypothesis. This observation,however, is similar to another study in which inclusion ofmuscles (abductor, tensor fascia latae and vastus lateralis)resulted in less migration than did the hip contact forcealone for a cemented implant [8]. Nonetheless, there areseemingly conflicting results in the literature; anotherstudy reported that including muscle forces (abductor,tensor fascia latae, vastus lateralis, and vastus medialis)resulted in much greater motion than did the hip contactforce alone for the CLS cementless implant [10].Although related debates[51], there is no clear explana-tion on this conflicting result. We suspect that these dif-fering observations may be related to differences inmedial-lateral bending moments in the femur, which are.not only affected by the abductors, but also in great partby the orientation of the hip contact force. In the study byKassi et al. [10], the hip contact force was applied at a 20°angle from the long axis of the femur in the frontal plane,whereas in the current study and that of Britton el al. [8]it was applied at 13°. These two angles are within therange reported from in vivo measurements [15,44,52], yetthey generate different bending moment distributions. At13° from the femur axis [15], the hip contact force gener-ates medial bending in the femur, which tapers to roughlyneutral bending around the implant tip, whereas at anangle of 20° [44] it generates medial bending in the femuraround the proximal stem, but substantial lateral bendingat the implant tip. The abductor load generates an addi-tional medial bending moment, which, when superposedwith the effect of the hip contact force, results in a morepronounced medial moment when the hip contact forceis applied at 13° compared with when the force is appliedat an angle of 20°. Differences in implant-bone interfacecontact stresses from the resulting bending momentsmay explain why the muscle forces affected implantmotion differently between these studies. If this is thecase, the orientation of the hip contact force may be moreimportant than whether or not the abductor force isincluded in in-vitro primary stability studies. Nonethe-less, it is also possible that the effect of muscles onimplant motion is sensitive to the implant design.The muscle attachment technique may also haveaffected the implant motion. In one study [10] the femurswere machined at the muscle insertion site which mayhave artificially weakened the femur, possibly increasingin the bone-implant motion. In the current study, theabductor attachment was done through a polymethyl-methacrylate that was fitted onto the greater trochanter,observed during stair climbing result in greater implant-bone micromotion and migration compared with walk-ing.and which may have reduced the motion by stiffening thebone locally. Britton et al., however, also observed areduction in implant motion when adding muscle forcesPark et al. Journal of Orthopaedic Surgery and Research 2010, 5:40http://www.josr-online.com/content/5/1/40Page 13 of 14with woven polyethylene straps glued to the greater tro-chanter, which is unlikely to have stiffened the bone [8].Whether it is better to include or exclude the abductorand/or other muscles during pre-clinical testing is debat-able. It can reasonably be argued that including all mus-cles provides a more physiologically representativeloading scenario. However, the question of how muchbending occurs physiologically is still being argued, e.g.[53]. Inclusion of muscle forces also introduces a poten-tial source of inter-specimen variability which could over-shadow the effect of the variable being studied. Sincemigration measured in-vitro is typically lower thanreported clinically, a set-up yielding higher bone-implantmotion could be considered as favourable for pre-clinicaltesting. Based on our results, with the hip contact forceapplied at 13° from the femur axis in the frontal plane,maximum implant motion was observed when simulatingstair climbing without the abductor force.ConclusionsSubstantially higher rotational and translational implantmotion was observed when applying an anterior-poste-rior hip contact force representative of stair climbingloads versus walking loads. This difference, however, wasmost prominent in the absence of the abductor muscleforce. We believe that the current study improves uponprevious research by examining the effect of the abductorforce and the anterior-posterior hip contact force onimplant primary stability under physiological cranial-caudal loading and in isolation from other muscle groups.Competing interestsThe authors declare that they have no competing interests.Authors' contributionsYP performed the design and execution of the experimental setup and analy-sis, as well as drafted the manuscript. CA executed and analyzed the experi-ment, performed statistical analsys as well as drafted the manuscript. YYprovided the design of the experimental setup, and participated in the intro-duction and study design. GF provided important feedback on the experimen-tal setup and partipated in the discussion. HF provided the design of theexperimental setup and participated in the discussion. TO provided importantfeedback on the statistical analysis and participated in the discussion.Author Details1Department of Mechanical Engineering, Korean Advanced Institute of Science and Technology, Daejeon, Republic of Korea, 2Orthopaedic and Rehabilitation Engineering Center, Marquette University, Milwaukee, Wisconsin, USA, 3Department of Materials Engineering, University of British Columbia, Vancouver, Canada, 4Department of Mechanical and Aerospace Engineering, Carleton University, Ottawa, Canada and 5Department of Mechanical Engineering, University of British Columbia, Vancouver, CanadaReferences1. Malchau H, Herberts P, Eisler T, Garellick G, Soderman P: The Swedish 2. Freeman MA, Plante-Bordeneuve P: Early migration and late aseptic failure of proximal femoral prostheses.  J Bone Joint Surg Br 1994, 76(3):432-8.3. Kobayashi A, Donnelly WJ, Scott G, Freeman MA: Early radiological observations may predict the long-term survival of femoral hip prostheses.  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