REMODELING OF FEMORAL STEM1.docx

7/30/2019 REMODELING OF FEMORAL STEM1.docx

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FINITE ELEMENT METHOD

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ABSTRACT:

The most commonly reported complications related to femoral stems are loosening and thighpain; both of these have been attributed to high levels of relative micro motion at the boneimplant interface due to insufficient primary fixation. Primary fixation is believed by many to relyon achieving a sufficient interference fit between the implant and the bone. However, attemptingto achieve a high interference fit not infrequently leads to femoral canal fracture either intra-operatively or soon after. The appropriate range of diametrical interference fit that ensuresprimary stability without risking femoral fracture is not well understood. In this study, a finiteelement model was constructed to predict micro motion and, therefore, instability of femoralstems. The model was correlated with an in vitro micro motion experiment carried out on fourcadaver femurs. It was confirmed that interference fit has a very significant effect on micromotion and ignoring this parameter in an analysis of primary stability is likely to underestimatethe stability of the stem. Furthermore, it was predicted that the optimal level of interference fit isaround 50 mm as this is sufficient to achieve good primary fixation while having a safety factorof 2 against femoral canal fracture. This result is of clinical relevance as it indicates arecommendation for the surgeon to err on the side of a low interference fit rather than riskingfemoral fracture.


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INTRODUCTION

Femur

The femur is the longest and strongest bone in the skeleton, is almost perfectly cylindrical in the greater

part of its extent. In the erect posture it is not vertical, being separated above from its fellow by a

considerable interval, which corresponds to the breadth of the pelvis, but inclining gradually downwardand medial ward, so as to approach its fellow toward its lower part, for the purpose of bringing the

knee-joint near the line of gravity of the body. The degree of this inclination varies in different persons,

and is greater in the female than in the male, on account of the greater breadth of the pelvis.

Figure 1: Femur

Fractures

A femoral fracture that involves the femoral head, femoral neck or the shaft of the femur immediately

below the lesser trochanter may be classified as a hip fracture, especially when associated with

osteoporosis.

Figure 2; Points of Fractures of Femur
http://en.wikipedia.org/wiki/Femoral_fracturehttp://en.wikipedia.org/wiki/Femoral_headhttp://en.wikipedia.org/wiki/Femoral_neckhttp://en.wikipedia.org/wiki/Shaft_of_the_femurhttp://en.wikipedia.org/wiki/Hip_fracturehttp://en.wikipedia.org/wiki/Osteoporosishttp://en.wikipedia.org/wiki/Osteoporosishttp://en.wikipedia.org/wiki/Hip_fracturehttp://en.wikipedia.org/wiki/Shaft_of_the_femurhttp://en.wikipedia.org/wiki/Femoral_neckhttp://en.wikipedia.org/wiki/Femoral_headhttp://en.wikipedia.org/wiki/Femoral_fracture


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Femoral Stem

The femoral stem component replaces a large portion of bone in the femur, and this is therefore the

load-bearing part of the implant. To bear this load, it must have a Youngs Modulus comparable to that

of cortical bone. If the implant is not as stiff as bone, then the remaining bone surrounding the implantwill be put under increased stress. If it is stiffer than bone, then a phenomenon known as stress

shielding will occur.

Figure 3; Femoral stem

DESIGN OF FEMORAL STEM

Design of the femoral stem is an important issue in the field of total hip arthroplasty, but design is just

one component in the success or failure of the operation. Other components are surgical technique,

cement technique or press-fit technique, bone quality, as well as patient related factors.

The quality of design may not also be matched with quality of manufacturing and machining of the stem.

The ultimate outcome of the arthroplasty obviously depends also on a matching acetabular component.

Currently the femoral stem revision rate at 10-15 years is reported to be between 0% and 4.8% and does

not correlate well with the radiographic stem loosening.

Femoral stem design options are related to whether the stem is curved or straight, the presence or

absence of collar support on the calcar, the stem cross section, the stem offset, the surface finish, as

well as the value of stem modularity and some metallurgical issues.


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Stem Offset?

The offset is the transverse distance between the centre of the head and the vertical line representing

mid-stem or mid-femur (fig.4). Variability of offset helps to replicate the anatomy by insuring proper softtissue tension (fig.5) which balances the hip bearings. Although a high offset stem relatively increases its

bending moment, various reports show that a high offset does not increase cement strain on medial

cement mantle.

Figure 4; Stem offset is the distance between the head centre and vertical line representing the mid-stem.

Figure 5; The offset of the stem helps to replicate normal soft tissue tension.

Surface Finish?

How smooth should be the surface of the stem! Is a feature of great variation as it comes in five

different ranges? Any surface will show peaks and valleys when examined by scanning electron

microscopy, the average between Peak and Valley is known as the Roughness Average (Ra); according to

Ra the surface finish of femoral stems may be classified as:

1. Highly polished

2. Satin

3. Matt

4. Rough

5. Textured


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A polished surface will show less fixation strength to cement, to the contrary of rougher surfaces which

show greater fixation strength to cement.

Debonding is the loss of fixation between metal and cement. When debonding happens rough surfaces

behave badly, as it will abrade the adjacent cement and will cause microfractures in cement mantle,

ultimately leading to loss of fixation. This may lead also to the release into the effective joint space of

abrasive wear debris from cement and metal, which when ground inside the bearings will act by 3rd body

wear mechanism to release submicron poly wear particles initiating the process of osteolysis.

Surface Features?

These are any irregularities present on the stem surface apart from its finish discussed above, like

flanges, serrations, centralizers, pre-coated beads and knobsetc..

The only surface features that may be beneficial are flanges and centralizers.

Flanges are a part of the stem popularized in later Charnley design stems (fig.6) to help pressurize the

cement as the proximal stem part is pushed into the femur.

Figure 6; The flanged design followed the round back design in the Charnley stem series.

The stem centralizer (fig.16) is also beneficial as it prevents the stem from deviating in the canal,insuring even cement mantle and perhaps preventing an unwanted varus position of the stem. Non end-

bearing centralizers may prevent cement fracture below the stem when subsidence occurs.

Figure 7; Stem centraliser insures a regular cement mantle and a centrally located stem.


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Pre-coating with PMMA was a good idea assuming better cement bonding to the PMMA pre-coat as

compared to metal. This did not seem to work, as there were reports by Mohler, 1995 of early femoral

loosening in 2-10 years, other reported 15% stem failure rate over 6 years due to poor cement mantle

and centralization.

Modularity?

Modularity helps intra-operative adjustment of components, most designs allow neck length (fig.8) and

head size modularity, and a select few allow modularity in anteversion and CCD angle.

Figure 8; Modular neck length, the short, standard and long heads can vary the neck length, allowing adjustments during

surgery.

The questions of increased wear and corrosion due to micro motion between the different pieces of the

modular stem remain to be proven to assume a clinical disadvantage to these designs, however; the

clinical problems of impingement / dislocation (e.g. by using a skirted extra-long head, or a very shorthead on a broad conical neck) and of undue lengthening fall under the technical control of the surgeon,

who must be aware of design and limitations of the stem he is implanting.

The modular stem costs more than the mono block sibling, and adds to the logistics of the hospital

creating more stock control overload on the administrator.

Metallurgical Issues

The current concept in hip arthroplasty prefers Cobalt-Chrome or Stainless steel for the cemented stems

and Titanium for the cementless. Other ideas are also available; but the majority of surgeons world wide

support this current concept.

The Scope of Stem Design

This presentation stressed mainly on the standard cemented stem, but the scope of stem design is much

larger, the cementless stem may share many of the above points of discussion apart from those related


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to cement mantle and bonding. The surface of the cementless stem and its coating may warrant a

separate article.

The recent evolution of special stems used in femoral reconstruction and revisions is also not covered in

this article, the author believe that these are better understood when discussed among topics related to

complex femoral reconstruction and revision arthroplasty

PURPOSE OF FEMORAL STEM REMODELING

A hip replacement with a femoral stem produces an effect on the bone called adaptive remodeling,

attributable to mechanical and biological factors. All of the prostheses designs try to achieve an optimal

load transfer in order to avoid stress-shielding, which produces an osteopenia.

INTRODUCTION

The implantation of a cemented or cementless femoral stem implies an important change inthe

physiological load distribution. The bone reacts to the new situation, in accordance withWolff 's law,

undergoing a process of adaptive remodeling, related to both mechanicaland biological factors, beingthe most important the initial bone mass.

Achieving good primary fixation is of crucial importance in cementless hip arthroplasty to ensure good

short-term and long-term results. Lack of primary stability leads to thigh pain and eventual loosening of

the prosthesis because of a continuous disruption of the bone formation process around the implant

(Kim et al., 2003; Knight et al., 1998;Mont and Hungerford, 1997; Petersilge et al., 1997). The stability, or

the lack of it, is commonly measured as the amount of relative motion at the interface between the

bone and the stem under physiological load. Large interfacial relative movements reduce the chance of

osseointegration, and cause the formation of a fibrous tissue layer at the boneimplant interface (Pilliar

et al.,1986), which may eventually lead to loosening and failure of the arthroplasty.

The threshold value of micro motion, above which a fibrous tissue layer forms, has been studied in both

animals and humans. In a review of dental implants in animals, a threshold micro motion value between

50 and 150 mm was found (Szmukler-Moncler et al., 1998). A similar range of values was reported for

orthopaedic implants in humans. In a retrieval study of cementless femoral components, Engh et al.

(1992) found indications that micro motions less than 40 mm had resulted in osseointegration while

micro-motions of 150 mm had caused the interposition of a fibrous tissue layer at the stembone

interface. It can be concluded from these reports that the value of micro motion, above which

osseointegration is disrupted, ranges from 50 to 150 mm, possibly skewed towards the lower end of this

range.

While many believe a sufficiently high interference fit is essential to achieve good primary stability, it is

also clear that introducing a interference fit has caused a clinically significant increase in intra-operativefemoral canal fractures (Cameron, 2004; Meek et al., 2004), an effect which has also been demonstrated

during in vitro testing (Jastiet al., 1993; Monti et al., 2001). The appropriate range of interference fit that

ensures primary stability without risking femoral fracture is not well understood.

There are in principle two parts to this study. In order to get a rough idea of the interference fit

introduced using current surgical practise, in the first part of this study, finite element predictions were

correlated with in vitro micro motion measurements. The aim of this was to enable back calculation of

the real interference fit introduced by the surgeon during the in vitro experiment. In the second part


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of the study, the effect of a range of interference fits on micromotion predictions was investigated using

finite element models of a more physiologically realistic loading scenario than was possible during the

first part of the study.

METHODOLOGY

In the first part of the study, the finite element models were based on CT scans from the specific bones

used in the experiment. In the second part of the study, the CT scans from the visible human dataset

were used. Also in the first part of the study, the purpose was simply to compare finite element

predictions and experiments and to simplify the experiments, a simple load configuration was chosen. In

the second part of the study, physiological loads including muscle loads were used.

In vitro experimental set-up

The experiment was designed for direct comparison of micromotion values between experiment and FE

analyses. Four cadaver femurs and Alloclassic (Zimmer GmbH, Winterthur, Switzerland) hip stems were

used, and two points, one in the proximal part and another in the distal part of the stem (Fig. 9), werechosen for micromotion measurement. In order to avoid damaging the stembone interface during

drilling action, the two points on the implant were drilled before implantation. A guide jig ensured that

the bone, subsequent to stem insertion, was drilled in the position matching these same two points on

the stem. Finally, steel pegs were glued into the holes in the stem and protruding through the bone (Fig.

9, right). A linear variable differential transducer (LVDT Model DFg5, DC Miniature series, Solartron

Metrology, UK), was rigidly fixed to the outside of the femur (Fig. 9, right). The connecting rod of the

LVDT core rested on the free-end of the steel peg. When the implant was loaded, the implant and hence

the peg moved relative to the bone and the LVDT measured the axial movement of the peg relative to

the transducer, thus providing an estimate of the relative axial movement between bone and stem.

Implantation was carried out by an experienced orthopaedic surgeon (D.L.). The neck of the femur was

first resected, and the femur was then reamed with firm impaction using a series of reamers to open the

canal. A femoral stem was then implanted in the femur.

Figure 9; The jig used to position the holes in the bone and the pegs in the implant, respectively (left). The implant bone

specimen with LVDT attached to the femur loaded in compression in the mechanical testing machine (right).


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The femur was sectioned 250mm distal to the lesser trochanter and its distal end fixed inside a

cylindrical metal container using polymethyl- methacrylate (PMMA). These were then placed onto the

table bed of auniversal materials testing machine (Instron 5565, Instron Corp., Canton, MA). The

specimen was adjusted so that the long axis of the stem was coaxial to the direction of loading. A cyclical

axial compression load of 02 kN and triangular waveform was applied to the shoulder of the stem for

50 cycles at a rate of 1 kN/min using a 5 kN load cell. Micro motion readings via the LVDT were taken

manually at maximum load of 2 kN and when fully unloaded at each cycle.

Finite element methodology for correlation study

A 3D model of a hip stem (Alloclassic, Zimmer GmbH) was constructed from CAD files received from the

manufacturer (Fig. 10).

Figure 10; The hip stem used in the study indicating the FE mesh used (left) and the implant inserted in the femur (right).

In the correlation part of the study, the finite element model needs to be as accurate a representation of

the experimental set-up as possible. Hence, the FE simulations of this part of the study were based on

CT scans of the specific bones used in the experiments. There were two sets of scans: one scan prior to

inserting the implant in the femur and a subsequent scan after implantation. The first set of scans was

used to derive bone geometry and material properties from the Hounds field units of the scan, while the

second set of scans was used to ensure that the implant position and orientation in the FE model

precisely matched the implant position within the femur in the experiment. The reason for this two-stepprocedure is that it would be inappropriate to use the CT datasets from the implanted femur for bone

property assignment due to artefacts in these datasets caused by the metal stem.

The construction of 3D models of the hip was done using AMIRA software (Mercury Computer Systems,

Inc., San Diego, CA). Segmentation was compiled automatically using the softwares marching cubes

algorithm which generates a 3D triangular surface mesh. The completed model was then converted to

solid linear tetrahedral elements using Marc. Mentat (MSC.Software, Santa Ana, CA) software. The mesh


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was inspected to ensure it was reasonably shaped throughout. The Marc finite element software

package was used in this study.

Material properties for the bone were assigned based on the grey-scalevalue of the CT images on an

element-by-element basis. The grey-level ofthe CT images were related to the apparent density using a

linear correlation (Cann and Genant, 1980; McBroom et al., 1985). This allowed for the transformation

of the spatial radiological description into thedescription of bone density. The modulus of elasticity of

individualelements was then calculated from the assigned apparent densities using the cubic

relationship proposed by Carter and Hayes (1977). The materialproperties were assumed to be linear

elastic and isotropic with Poissonsratio set to 0.35. The FE model was loaded at the centre of the

shoulder of the stem with 2 kN, the stem being coaxial to the direction of loading, hence, matching the

loading configuration in the experiment.

Mesh convergence is a standard issue in any finite element analysis and in a contact analysis, there are

many other numerical parameters that affect the predicted micro motions. The default contact strategy

inMarc is a direct constraint algorithm (MSC.Marc-Manual, 2004)which most importantly requires the

input of a contact zone size (CZS).Furthermore, Bernakiewicz and Viceconti (2002) described the

importanceof the convergence tolerance (CTol) in non-linear analyses. They alsosuggested that the

appropriate parameter settings should be such that theresultant change in predicted micro motion

between models with differentparameter settings should be small relative to 150 mm. A sequentialsensitivity analysis involving mesh density, CZS and CTol was carried out and a model with 12,078 nodes,

CZS 0.025mm and CTol 1% was found to be sufficient for an accurate solution.

We also chose a Coulomb friction model which in Marc requires the input of the friction coefficient (m)

as well as a parameter (SL). The Marcsoftware has introduced the parameter SL, which describes a

smoothing ofthe step-function of the Coulomb model, only in order to deal with anotherwise

numerically difficult to handle discontinuity. However, not onlydoes this parameter dramatically affect

the predicted micro motion (Fig. 3)it also has an important physical interpretation. Shirazi-Adl et al.

(1993)showed that the boneimplant interface friction curve is highly non-linear,exhibiting micro

motion on the order of 150 mm (that is in the order of the critical level for osseointegration) before the

slip load predicted by theCoulomb model is reached. The implication of Shirazi-Adl et al.s work isthat

adopting the ideal Coulomb model is inadequate. However, the SLparameter can be interpreted andused to represent this non-linear behavior.

Figure 11; Contour plots of micromotion over the surface of the Alloclassic stem under stairclimbing loads and for different

values of the SL parameter (SL describes the non-linear friction characteristics of the interface).


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To establish the appropriate setting of the SL parameter, we simulated Shirazi-Adl et al.s relatively

simple experiment consisting of a bone cubeexposed to normal and tangential loads moving on a metal

plate. In Fig. 12 is shown Shirazi-Adls experimental curve of tangential load versus tangential

displacement. The tangential load that would initiate slipaccording to the Coulomb model is 30.6. The

finite element predictedcurves for various settings of SL is also shown and a setting of SL 0.1 predicts

the experimental curve well. Hence, in the rest of this study, thissetting was used.

Figure 12; Tangential load versus tangential displacement of bone cube sliding on metal plate. The finite element predicted

non-linear friction behavior for different levels of the parameter SL is shown as well as the experimental curve reported by

Shirazi-Adl et. The critical value at which sliding would initiate according to an ideal Coulomb friction model is also indicated.

The effect of friction coefficient on micro motion is relatively minor forfriction coefficients higher than

0.15 (Kuiper and Huiskes, 1996). Viceconti et al. (2000) found that a friction coefficient between 0.2 and

0.5 led to the best correlation with experiments. Rancourt et al. (1990)measured friction coefficients

experimentally and found a coefficient of0.4. Based on these previous studies, a friction coefficient of

0.4 was used in this study. The objective of this study was to estimate the effective interference fit.

Hence, we varied the interference fit in the finite element models. Thepredictions were then compared

to the experimentally measured values toestimate which level of interference best matched the

experiment.


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implantbone interface and prevent osseointegration. Hence, it is the converged values of Fig. 5 which

arerelevant. Based on the data of Fig. 5, the converged average value in the distal and proximal regions

were 1872 and 1975 mm, respectively.

Figure 13; Distal micro motion (top) and proximal micro motion (bottom) results from the experiment.

The results of the FE analyses using different levels ofinterference fit and simulating the experiment are

shown inFig. 6. The figure shows that with just 1 mm of interference, the level of micro motion is

predicted to be in the range of2030 mm. With 2 mm of interference, this drops to 1020 mm.

Comparing this to the experimental values of18 and 19 mm also shown in the figure, this implies thatthe interference fit introduced by the surgeon is only 1 or 2 mm.


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Figure 14; Contour plots of micro motion over the surface of the stem under an axial load of 2 kN, using interference fits of

(from left to right) 0, 1, 2 and 5 mm, respectively. The experimentally determined proximal and distal micro motion is also

indicated.

This seems perhaps unrealistically low. Shultz et al. (2006) considered an interference fit of100 mm tocause bone interface damage and reported thislevel of interference as a threshold value. Therefore, we

included an interference fit of 100 mm in one of the finite element models and inspected the resulting

tensile hoopstresses (Fig. 7). This model was not exposed to any other loads. As can be seen from the

figure, interference inducedhoop stresses are on the order of 50MPa on the surface of the bone

(internally the stresses are somewhat higher).Comparing this stress level with the transverse tensile

strength of cortical bone of approximately 50MPa (Reillyand Burstein, 1975), it would seem that 100

mm representsthe critical level of effective interference fit above which the femoral canal will fracture.

The location of high hoopstresses towards the distal end of the implant seen in Fig. 7 also matches the

location of 77% of intra-operative fractures (Meek et al., 2004).

Figure 15; Hoop stresses in femoral bone caused by an interference fit of 100 mm. No external loads are applied in this

model.


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Considering that femoral canal fractures are not infrequently occurring intra-operatively (Cameron,

2004; Meeket al., 2004), it would seem that surgeons are introducingclose to the critical level of

interference fit of 100 mm.

Assuming that surgeons are able to control the insertion process within a factor of 2, perhaps a realistic

rangeof interference fit can be argued to be in the range of 50100 mm.In summary, this first part of the

study indicates that therange of realistic interference fits may be within a range of very low levels (just a

few microns) and up to 100 mm.

The effect of interference fit on micro motion

Fig. 16 shows contour plots of predicted micro motion over the stem surface under stair climbing loads

and for four different levels of interference fit. Fig. 17 shows the change in micro motion with levels of

interference fit for the two points labeled P (proximal) and D (distal) shown on the left model of Fig. 16.

Also in Fig. 17 is indicated, by the grey-colored region, the threshold range of micro motion above which

soft tissue formation will be predicted and below which osseointegration would be expected. From

these two figures, it is clear that the interferencefit had a very large effect on micro motion predictions.

In the case of no interference fit, the entire surface area of theimplant was in or above the grey area

indicating that theprimary stability of the implant is at risk. In contrast, with50 mm of interference, allbut the most proximal part of theimplant was predicted to osseointegrate. Interestingly,increasing the

level of interference beyond 50 mm hadnegligible effect. Also, it is clear that the effect of

theinterference fit was most dramatic at low levels ofinterference. Including just 5 mm of interference

causesalmost a 50% reduction in micro motion and including more interference only has a relatively

small effect.

Figure 16; Contour plots of micro motion over the surface of the stem under stair climbing loads and with interference fits of

(from left to right) 0, 5, 25 and 50 mm, respectively.


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Figure 17; Micro motion at points P (proximal) and D (distal) as a function of the level of interference fit. Locations of point P

and D are shown in Fig. 16 (left). The grey area indicates the range of the critical micro motion threshold. Above this level,

fibrous tissue formation would be expected; below, osseointegration is anticipated.

CONCLUSION

This study has shown that modeling the interference fit characteristic of hip stems is crucial for

quantitative predictions of micro-motion. Ignoring the interference fit will probably lead to an under

estimation of the stability of the stem. In contrast, ignoring the non-linear friction behavior reported by

Shirazi-Adl et al. (1993) and reproduced in Fig. 4, will probably to lead to too optimistic predictions of

stem stability. The magnitude of interference fit is fundamentally unknown and may be the reason most

previous works have omitted this parameter from their finite element analyses. Indeed, during this

study it became clear just how difficult it is to estimate this parameter. Nevertheless, this study

demonstrates the importance of the interference fit as including only a small level of interference

changed the evaluation of the investigated stem from that of an unstable stem to that of a stable stem.

Our predictions showed high levels of micro-motion distally and proximally while micro-motion at the

stem midsection was lower (Fig. 8, left). This is qualitatively consistent with the finite element

predictions by Keaveny and Bartel (1993). Keaveny and Bartel did not include an interference fit and

predicted very high absolute values of micro motion (0550 mm). Keaveny and Bartel simulated a

cylindrical stem which is likely to be less resistant to torsional loads and that may explain the higher

levels of Micro motion as compared to our results. Viceconti et al. (2000) did simulate a press-fit

although it is notpossible to quantify this press-fit in a manner that allows a direct correlation with our

results. Vicecontiet al. predicted micro motions ranging from 17 to 49 mm across the surface of the

implant which is reasonably consistent with our results simulating interference fit of 25 mm (Fig. 8).

The results of Fig. 6 indicate that surgeons introduce very low interference fits, on the order of 12 mm.

Apart from any aspects of the model that may cause inaccurate predictions, it is of course also possible

that the experimental results are inaccurate. Notably, our experiment, like the vast majority of other

experimental micro motion studies, does not measure the actual interface micro motion but instead

measures the motion between the LVDT fixation point on the bone and the point of the peg insertion on

the implant. The motion measured, therefore, includes other flexibilities such as bone deformation and

will tend to overestimate micro motion (Bu hler et al., 1997). If these flexibilities are substantial


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compared to the true interface micro motion, it would cause our methodology to predict very small

levels of interference which is of course what seems to be the case.

In connection with Fig. 7, we proposed that surgeons are in fact more likely to introduce interference fits

of 50100 mm. Shultz et al. (2006) predicted that with an interference fit of 100 mm, the hoop stresses

in the bone would visco-elastically relax by approximately 50%. In other words, if a surgeon introduces

an interference fit of 100 mm, this would relax and represent an effective interference of 50 mm. Shultz

reported that interference fits lower than 100 mm would relax less than 50%. Therefore, even if a

surgeon only achieves the lower range of the50100 mm interference, we have estimated, there should

be at least 25 mm of effective interference left after relaxation, well above the 12 mm estimated from

the experiment. We have no evidence to explain the small levels of interference fit predicted from the

experiments but we are inclined to believe that the experiment overestimated the micro motion, for the

reasons noted above.

We have assumed a uniform interference fit over the entire surface of the implant. Accordingly, the

press-fit (pressure) varied considerably from the proximal cancellous femur to the cortical distal femur

as modelled through the variation in the local Youngs modulus of the bone adjacent to the implant. This

variation in press-fit between the proximal and distal region is undoubtedly qualitatively correct.However, our study was not set up to investigate variation in interference fit. This was not included due

to the practical difficulty in quantifying the variation and generalizing such variation that is likely to vary

between implants. It is also probable, given the very small interferences calculated, that surgeons

cannot create implant cavities with uniform interference across the interface area, so that clinical cases

would include variations from the micro motions predicted. The effects of a more realistic scenario are

not yet known.

The results of this study support the suggestion made earlier (Shirazi-Adl et al., 1994) that the cavity that

is created in the femur is larger than is indicated by the nominal interference of 0.30.5mm (Otani et al.,

1995; Ramamurti et al., 1997); such a large interference would cause the femur to fracture, according to

our results.

Perhaps the most important result of the study and the result with direct clinical relevance relates to

Figs. 7 and 9. Fig. 7 predicts that surgery is safe against femoral canal fracture at interference fits lower

than 100 mm. Fig. 9 predicts that the stem would osseointegrate at interference levels of 50 mm.

Therefore, the recommendation is for the surgeon to err on the side of a low interference fit during

surgery as only 50 mm is enough to achieve stability and provides a safety factor of 2 against femoral

canal fracture. If considering a stem likely to be successful as long as just the distal part of the stem

(embedded in the strong cortical bone) osseointegrates, Fig. 9 indicates that just 10 mm of interference

fit is necessary for stability and provides a safety factor of 10 against femoral canalfracture.

Of course, our computational predictions should befurther investigated before being applied in clinical

practice. It is likely, that stems with different geometry ormaterial will behave differently. The Alloclassic

stem in this study, for example, has a rectangular cross-section, whichmight be advantageous in

resisting torsional loading during the stair climbing simulated. Nevertheless, the predictionsclearly

indicate a recommendation to modify surgical practice thereby reducing or even eliminating the

7%intra-operative femoral canal fractures during primary hipsurgery reported by Cameron, (2004) and

the 650%fracture rates reported by Meek et al. (2004) in connectionwith revision hip surgery.

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REMODELING OF FEMORAL STEM1.docx