journal of the mechanical behavior of biomedical materials 32 (2014) 257–269

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Research Paper

Bone remodelling around cementless composite acetabular components: The effects of implant geometry and implant–bone interfacial conditions Rajesh Ghosh, Sanjay Guptan Department of Mechanical Engineering, Indian Institute of Technology Kharagpur, Kharagpur 721 302, West Bengal, India

art i cle i nfo

ab st rac t

Article history:

Recent developments in acetabular implants suggest flexible, alternative bearing material

Received 11 September 2013

that may reduce wear and peri-prosthetic bone resorption. The goal of this study was to

Received in revised form

investigate the deviations in load transfer and the extent of bone remodelling around

10 January 2014

composite acetabular components having different geometries, material properties and

Accepted 13 January 2014

implant–bone interface conditions, using 3-D FE analysis and bone remodelling algorithm.

Available online 23 January 2014

Variation in prosthesis type and implant–bone interface conditions affected peri-prosthetic

Keywords:

strain distribution and bone remodelling. Strain shielding was considerably higher for bonded

Pelvic bone

implant–bone interface condition as compared to debonded implant–bone interface condition.

Hip prosthesis

The average bone deformation (0.133 mm) for horseshoe-shaped CFR-PEEK (resembling

Composite acetabular component

MITCH PCRTM cup) was very close to that of the intact acetabulum (0.135 mm) at comparable

Finite element analysis

locations. A reduction in bone density of 21–50% was predicted within the acetabulum for the

Bone remodelling

implant resembling Cambridge cup, having bonded interface. For debonded interface condition, bone density increase of
55% was observed in the supero-posterior part of acetabulum, whereas bone density reductions were low (1–20%) in other locations. Bone density reductions were considerably less (2–4%) for horseshoe-shaped CFR-PEEK component. Moreover, an increase in bone density of 1–87% was predicted around the acetabulum. Compared to the horseshoe-shaped design, the hemispherical design exacerbated bone resorption. Results indicated that the thickness of the acetabular component played a crucial role in the implant induced bone adaptation. The horseshoe-shaped CFR-PEEK component of 3 mm thickness seemed a better alternative bearing surface than other designs, with regard to strain shielding, bone deformation and bone remodelling. & 2014 Elsevier Ltd. All rights reserved.

1.

Introduction

The most common causes of failure of the acetabular component are either due to stress shielding induced adverse n

Corresponding author. Tel.: þ91 3222 282958; fax: þ91 3222 282277. E-mail address: [email protected] (S. Gupta).

1751-6161/$ - see front matter & 2014 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.jmbbm.2014.01.010

bone remodelling or accumulation of wear particle debris of the implant material in peri-prosthetic tissue (Schmalzried et al., 1997; Dumbleton et al., 2002). Excessively stiff acetabular implants, made of metallic alloy or ceramic, potentially cause

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journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

loosening by stress shielding and peri-prosthetic bone resorption (Lewis et al., 1984; Huiskes et al., 1992; Bobyn et al., 1992; Hedia et al., 2000). In comparison, although low stiffness implant material, such as ultra high molecular weight polyethylene (UHMWPE), has been widely used during the past decades, excessive generation of wear debris has been a limiting factor for long-term success of the prosthesis (Harris, 1995). Although, development of highly cross-linked UHMWPE has shown improved performance, more recently, composite materials evolved as an alternative to metallic, ceramic or polyethylene acetabular components, since it offers better wear resistance properties, has the potential to fabricate components with specific requirements and its elastic modulus (E) is more close to that of the host cortical bone. Recent developments in acetabular component suggest flexible, wear resistant, anatomic shaped acetabular components fabricated from polymer composites such as, carbonfiber reinforced Polybutyleneterephthalate (CFR-PBT) and carbon-fiber reinforced polyetheretherketone (CFR-PEEK), which would allow for more deformation and less periprosthetic bone loss than metal backed polyethylene, metallic and ceramic hemispherical components (Morscher et al., 1997; Brooks et al., 2004; Field and Rushton, 2005; Field et al., 2006, 2008; Manley et al., 2006; Manley and Sutton, 2008; Latif et al., 2008; Dickinson et al., 2012). It should be noted that intact acetabulum is subjected to large amount of elastic deformation due to the action of musculoskeletal loads during physiological activities (Konrath et al., 1998). Considering the requirements of flexible, wear resistance and anatomic shaped component, the horseshoe-shaped Cambridge cup, made of CFR-PBT with UHMWPE articulating surface, was designed to replace the articular cartilage of the acetabulum and the underlying subchondral bone. The Cambridge cup design was modified later, in which UHMWPE layer was removed from the articulating surface to avoid interface debonding between two materials and high volumetric wear rate of UHMWPE (Latif et al., 2008). The modified acetabular component, known as MITCH PCRTM cup, consisted of two parallel fins and made of only CFR-PEEK material, was reported to have better wear resistant properties (Scholes and Unsworth, 2007; Scholes et al., 2008). Despite the predictions of less adverse bone adaptation and low wear rates for composite acetabular components, there are only a few clinical and Finite Element (FE) studies that may be useful for investigating the extent of bone remodelling. The clinical study by Field et al. (2006) on Cambridge cup reported decrease in peri-prosthetic bone mineral density (BMD) up to six months, post-operatively. The FE study by Manley et al. (2006) on acetabular components predicted inevitable bone adaptation that cannot be resolved by changes in implant material properties alone. However, the FE model of the implanted pelvis assumed constant cortical bone thickness, and the predicted results corresponded to the bonded implant–bone interface condition only (Manley et al., 2006), which might not be entirely representative of the physiological conditions. Moreover, the changes (positive/negative) in strain energy density before and after implantation were assumed to cause changes in bone density (formation/resorption), without

actually simulating the bone remodelling algorithm. A recent experimental study by Dickinson et al. (2012), revealed pelvis cortex strain was close to natural pelvis for implantation with CFR-PEEK than metallic (CoCrMo alloy) and UHMWPE. However, the effects of these composite materials and geometry of the acetabular components on bone remodelling within the acetabulum are not clearly understood yet. It is therefore, hypothesised that the geometry and implant–bone interface conditions of composite cups, one made of CFR-PBT with UHMWPE linear and another CFR-PEEK, affect strain shielding and bone remodelling. The objectives of this study are, (1) to investigate the extent of peri-prosthetic bone remodelling for the composite acetabular components having variable implant–bone interface conditions and subsequently (2) to evaluate the effect of implant geometry on acetabular strains and bone remodelling.

2.

Materials and methods

The 3-D FE models of the intact and the implanted pelvic bones were developed using the CT-scan data set (500  410 pixels; pixel size of 0.781 mm; slice thickness of 2.5 mm) of the same subject (a 62 years old female) and following similar procedure described in an earlier study by Ghosh et al. (2013a, b, in press-c). Only a brief outline is presented here. A linear relationship between the bone apparent density (ρ in g cm  3) and CT-grey value (HU) given by, ρ ¼ 0:022 þ 0:001038  HU

ð1Þ

was derived using the CT number of water, i.e. zero corresponding to a bone density of 0.022 g cm  3, and the highest CT number of cortical bone, 1646 corresponding to the highest cortical bone density of 1.73 g cm  3. In order to correct for the partial-volume effect at the periosteal and endocortical boundaries, a thresholding method was used to determine the cortical bone thickness from the CT-scan data set (Anderson et al., 2005; Zhang et al., 2010; Varghese et al., 2011; Ghosh et al., 2013b). Based on a threshold density value of 1.3 g cm  3 for the endocortical boundary (Zioupos et al., 2008), which corresponded to 1231 HU in our CT data set, the cortical thickness ranged between 0.781 and 3.124 mm (Ghosh et al., 2013b). The cortical bone was assumed to be homogeneous with Young's modulus of 17 GPa (Anderson et al., 2005; Thompson et al., 2002; Dalstra and Huiskes, 1995; Ghosh et al., 2013b, in press-c), and heterogeneous isotropic material property distribution was used for the cancellous bone. The Young's modulus of cancellous bone elements was extracted from the CT-scan images using the public domain software Bonemat_V2 (Taddei et al., 2004) and the following density-elastic modulus relationship empirically derived by Dalstra et al. (1993), E ¼ 2017:3ρ2:46

ð2Þ

The resulting values of Young's modulus of cancellous bone ranged from 2.1 to 3846.54 MPa. The Poisson's ratio for bone was taken as 0.3. A horseshoe-shaped component, resembling Cambridge cup, having 54 mm outer diameter and 45 mm bearing diameter (Prosthesis 1) was virtually implanted in the

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

acetabulum in consultation with an experienced orthopaedic surgeon, maintaining an angle of 451 inclination and 151 anteversion (Janssen et al., 2010), using Solid Works software (DS Solid Works Corp., Concord, MA, USA) and NURBS modelling software Rhinoceros (Robert McNeel & Associates, Seattle, WA, USA). The geometry and material combinations of the Cambridge cup were similar to earlier published study (Field et al., 2006), wherein 3 mm thick UHMWPE bearing surface was used, interlocked with 1.5 mm thick 30% CFR-PBT (Fig. 1). Two more implants, one horseshoe-shaped resembling MITCH PCRTM cup (Prosthesis 2) and the other hemi-

259

spherical shaped (Prosthesis 3), both made of 30% CFR-PEEK and both having 54 mm outer diameter and 48 mm bearing diameter (Latif et al., 2008) were virtually implanted using similar procedure (Fig. 1). Mesh generation was carried out using ten-node tetrahedral element with maximum edge length of 3 mm and ANSYS FE analysis software (ANSYS Inc., Pennsylvania, USA). The validity of the FE model generation procedure for the intact and the implanted pelvises was assessed in earlier studies employing experimental measurements on composite hemi-pelvises, using digital image correlation and strain

Fixed boundary conditions

Z

Y

Hip joint force

X

Cut-boundaries (nodal displacements are

Mitch acetabular component

transferred at the cut-boundaries from implanted full model)

Fig. 1 – Finite element models intact and implanted pelvises; (a) full model of the intact pelvis, (b) full model of the implanted pelvis, (c) submodel of the implanted pelvis, (d) Prosthesis 1 resembling Cambridge cup, (e) Prosthesis 2 resembling MITCH PCR™ cup, and (f) Prosthesis 3 hemispherical CFR-PEEK cup.

260

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

gauge techniques (Ghosh et al., 2012, 2013a). Moreover, a mesh convergence study was performed based on stress generated in the implant and the bone, for accurate mesh density of bone and implant. The FE models of the intact and the implanted pelvises consisted of 231,532 and
270,000 elements, respectively, and are shown in Fig. 1. A single cartilage layer was assigned between intact acetabulum and femoral head with Young's modulus of 10.35 MPa and Poisson's ratio of 0.45 (Kempson, 1980). The prosthesis designs and material properties are summarised in Table 1.

2.1.

contact between the intact femoral head and the single layer acetabular cartilage was modelled as Coulomb frictional with a friction coefficient of 0.01 (Unsworth et al., 1975; Leung et al., 2009). Frictionless contact with 100 mm diametric clearance was assumed between the femoral head and the acetabular component (Manley et al., 2006). Augmented Lagrange contact algorithm was used for the contact simulation. A normal contact stiffness of 100 N mm  1 and a penetration tolerance factor of 0.1 were chosen for the convergence of the non-linear solution (ANSYS user's manual).

Implant–bone interface condition: contact simulations 2.2.

Two limiting implant–bone interface conditions were analysed; one with fully bonded (best case scenario) and other debonded interface (exact fit and worst case scenario) having coefficient of friction 0.5 (Shirazi-Adl et al., 1993; Janssen et al., 2010). Six-node asymmetric surface-to-surface contact elements were chosen for the contact simulation. The

Applied loading conditions

In this study, 21 muscle forces and hip joint force of a normal walking cycle were used as applied loading conditions. Magnitude of hip joint forces and muscle forces were based on the data reported by Dalstra and Huiskes (1995), whereas the directions of the hip joint force were calculated based on

Table 1 – Acetabular component designs, thicknesses and material properties used in the study. Acetabular component

Design

Prosthesis 1

Resembling Cambridge cup

Prosthesis 2 Prosthesis 3

Resembling MITCH PCR™ cup Hemispherical cup

Material

30% CFR-PBT UHMWPE 30% CFR-PEEK 30% CFR-PEEK

Thickness (in mm)

Young's modulus (in GPa)

Poisson's ratio

1.5 3 3 3

16.6 1.174 13 13

0.4 0.4 0.4 0.4

Table 2 – Magnitudes (N) of hip joint force and twenty one muscle (m.) forces corresponding to two load cases during a normal walking cycle.Data adopted from Dalstra and Huiskes (1995). Force

Hip joint force m. Adductor brevis m. Adductor longus m. Adductor magnus m. Biceps femoris m. Gemellus inferior m. Gemellus superior m. Gluteus maximus m. Gluteus medius m. Gluteus minimus m. Gracilis m. Iliopsoas m. Obturator externus m. Obturator internus m. Pectineus m. Piriformis m. Quadratus femoris m. Rectus femoris m. Sartorius m. Semimembranosus m. Semitendinosus m. Tensor fasciae latae

Load cases Load case 1 Beginning right single support (13% of cycle)

Load case 2 Double support, end right stance (52% of cycle)

2158 114 88 0 202 0 88 930 1053 140 0 0 0 123 0 175 96 123 88 368 140 132

1180 0 88 132 123 0 0 456 1412 175 88 395 123 61 0 0 88 0 35 421 316 149

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

the data reported by Bergmann et al. (2001). Two static load cases, one beginning of right single support (13% of cycle) and the other double support, end right stance (52% of cycle), were used as presented in Table 2. The directions of the muscle forces were found by subtracting the coordinates of the distal and proximal insertion points (Dostal and Andrews, 1981) and were resolved with respect to our FE coordinate system. The attachment sites of muscles were identified as patched areas on the pelvis cortex (Ghosh et al., 2013b). In the FE models, these muscle forces were distributed on the set of nodes contained in the respective patched areas of attachment (Ghosh et al., 2013b). Fixed constraints were prescribed at the pubic symphysis and the sacroiliac joint (Thompson et al., 2002; Dalstra and Huiskes, 1995; Clarke et al., 2013; Ghosh et al., 2013b).

2.3.

Submodelling technique

An overall full model of the implanted pelvis corresponded to musculoskeletal loading conditions of hip joint force and twenty one muscle forces during a gait cycle. A submodel of the implanted acetabulum, with a fine mesh around the prosthesis, was developed to focus our investigation around the acetabulum, without neglecting the effect of the forces acting on bony structures connected to the acetabulum (Ghosh et al., 2013b). A link between the submodel and the overall model was established by transferring the displacements (at nodes) from the implanted full model to the submodel at the cut-boundaries, located sufficiently far away from the implanted acetabulum (Fig. 1c). The submodel of the implanted pelvises contained
120,000 elements with element edge length ranging between 0.5 and 2.5 mm (Fig. 1c).

2.4.

The remodelling algorithm

In the present study, only internal morphological changes of bone were considered (Carter et al., 1989; Huiskes et al., 1987; Pal et al., 2010), since the cancellous bone has a higher metabolic activity and appears to respond more rapidly to changes in mechanical loads, as compared to the cortical bone (Garcia et al., 2002). The adaptive bone remodelling theory was based on strain energy density and used a site specific formulation (van Rietbergen et al., 1993; Weinans et al., 1993; Huiskes and van Rietbergen, 1995). The reference stimulus (Sref) and the actual stimulus (S) are the local (per element) elastic strain energy per unit of bone mass averaged over a loading history (n), for an intact and implanted pelvis, respectively, given by the following: S¼

1 n Ui Ua ¼ ∑ ni¼1 ρ ρ

ð3Þ

The dead zone (s) was taken as 70.75 of the reference stimulus (Huiskes et al., 1992; Pal et al., 2010). The mathematical expression for the change in apparent density is described as follows: ( aðρÞfS ð17sÞSref gτ Δt if Srð1 sÞSref or SZ ð1 þ sÞSref Δρ ¼ 0 if Sref ð1 sÞoSoSref ð1 þ sÞ ð4Þ where a(ρ) is the free surface area per unit volume for bone [mm2/mm3] in the internal bone structure (Martin, 1972),

261

estimated as a function of apparent density (Ghosh et al., 2013b), aðρÞ ¼ 0:0293 þ 8:5124ρ 4:887ρ2  1:568ρ3 þ3:7182ρ4 1:6352ρ5

ð5Þ

The reference stimulus for each bone element was obtained from the intact pelvis model, which was compared with the remodelling stimulus of the bone element at comparable location in the submodel of the implanted pelvis. The iterative bone remodelling algorithm was based on earlier investigations with metallic and ceramic acetabular components (Ghosh et al., 2013b). The integration was carried out in steps of ‘simulation time scale’ τΔt (Suarez et al., 2012). The adaptation rate (τ) was assumed as 129.6 g. mm  2 (J/g) months for calculating Δt (Weinans et al., 1993). The lower and upper bounds of bone density were set to 0.01 g cm  3 and 1.73 g cm  3, respectively. The convergence criterion was chosen when the density did not change more than 0.005 g cm  3 per element per time increment (τ Δt).

2.5.

Interpretation of results

Although load transfer and strength of a material are most commonly expressed in terms of stress, recent experimental study revealed that strain-based descriptions may be more mathematically simple and statistically powerful for human cancellous bone (Morgan and Keaveney, 2001). Morgan and Keaveney (2001) compared the uniaxial tensile and compressive yield properties for human trabecular bone specimens taken from different anatomic site and confirmed that it can be considered uniform within a single site, due to the weak dependence of it on the apparent density, despite substantial variation in elastic modulus and yield stress. Accordingly, equivalent strain distribution in the cancellous bone was chosen for assessment (Verhulp et al., 2008; Pal et al., 2010; Davis et al., 2009; Ghosh et al., 2013b). Four regions of interest (ROI) in the acetabulum were defined for detail evaluation of results (Fig. 2). Strain distributions in the intact and the implanted pelvis were obtained for both the load cases, results corresponding to beginning right single support (13% of cycle) phase were considered for more detailed evaluations, since the magnitude of hip joint force is higher for this load case.

3.

Results

Variation in prosthesis type and implant–bone interface conditions causes differences in strain distributions within the acetabular cancellous bone, after implantation (Fig. 2). High strains (0.7–1% strain) were generated in the acetabular fossa for the horseshoe-shaped components (Prostheses 1 and 2) for both the implant–bone interface conditions (Fig. 2). Post-operatively, Prosthesis 1 led to a reduction in equivalent (von Mises) strain of 20–60% in ROIs 1 and 3, when bonded implant–bone interface was assumed (Fig. 2b). Considering debonded implant–bone interface condition, an increase in strain of 40–80% was observed at ROI 1, whereas a strain reduction of 10–30% was observed at ROI 3 (Fig. 2c). Assuming debonded condition, 15–30% higher strains were observed for the Prosthesis 2 in ROIs 1 and 3, as compared to Prosthesis 1

262

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

1

2

3

4

1

2

1

2

3

4

3

4

1

2

1

2

3

4

3

4

1

2

1

2

3

4

3

4

Strain

Fig. 2 – Equivalent strain distribution in the cancellous bone for post-operative condition, lateral view; (a) intact acetabulum, (b) Prosthesis 1 having bonded implant–bone interface, (c) Prosthesis 1 having debonded implant–bone interface, (d) Prosthesis 2 having bonded implant–bone interface, (e) Prosthesis 2 having debonded implant–bone interface, (f) Prosthesis 3 having bonded implant–bone interface, and (g) Prosthesis 3 having debonded implant–bone interface. (Fig. 2c and e). The effect of strain shielding was maximum for Prosthesis 3 with fully bonded interface condition; a strain reduction of 10–60%, was observed in ROIs 1–4 (Fig. 2f). After attainment of equilibrium in bone remodelling, considerable reduction in strain of 50–80% was observed in all ROIs 1–4, for Prosthesis 1 having fully bonded interface condition (Figs. 2b and 3a). Whereas for debonded interface condition, an increase in strain of 30–60% was observed in ROI 3 and reduction in strain of 50–70% was observed in ROIs 1, 2 and 4 (Figs. 2c and 3b). In the case of Prosthesis 2, a reduction in strain of 30–60%

was observed in ROIs 1–4 due to bone remodelling, for fully bonded interface condition (Figs. 2d and 3c). However, when debonded interface condition was assumed, a decrease in strain of 20–60% was observed in ROIs 1–4 (Figs. 2e and 3d). For Prosthesis 3, bone remodelling led to an increase in strain of 10–25% in ROIs 1 and 2, whereas strain reduction of 5–20% was observed in ROIs 3 and 4, for fully bonded interface condition (Figs. 2f and 3e). Considering debonded interface condition, an increase in strain of 25–40% was observed in ROI 3, and 50–70% decrease in strain was observed in ROIs 1, 2 and 4 (Figs. 2g and 3).

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

1

2

1

2

3

4

3

4

1

2

1

2

3

4

3

4

263

Strain

1

2

1

2

3

4

3

4

Fig. 3 – Equivalent strain distribution in the implanted pelvis after equilibrium in bone remodelling; (a) Prosthesis 1 having bonded implant–bone interface, (b) Prosthesis 1 having debonded implant–bone interface, (c) Prosthesis 2 having bonded implant–bone interface, (d) Prosthesis 2 having debonded implant–bone interface, (e) Prosthesis 3 having bonded implant– bone interface, and (f) Prosthesis 3 having debonded implant–bone interface. The maximum equivalent (von Mises) stress generated within the acetabular components varied between 6 and 30 MPa. For Prosthesis 1, the maximum stress generated was
30 MPa around the antero-superior part of the horseshoe-shaped region. At similar location, lower peak stress of
20 MPa was generated in Prosthesis 2. The peak stress was considerably reduced to
6 MPa for the hemispherical CFR-PEEK cup (Prosthesis 3). Since the composite cups offer more flexibility than the metallic or ceramic cups, the average bone deformation underlying a prosthesis with regard to that of the intact acetabulum might be useful for design evaluations. The average bone deformations underlying Prostheses 1 and 2 were found to be 0.126 mm and 0.133 mm, respectively, whereas for Prosthesis 3 it was 0.129 mm. The bone deformation for Prosthesis 2 was very close to the average bone deformation of 0.135 mm at comparable locations in the intact acetabulum.

3.1.

Bone remodelling

Strain shielding led to bone remodelling within the implanted acetabulum. The changes in bone density distribution, immediate post-operative and after equilibrium in bone remodelling for

Prostheses 1, 2 and 3 are presented in Figs. 4–6, respectively. Moreover, the quantitative changes in average bone density in the four ROIs for the acetabular components are listed in Table 3. The implant material, geometry and implant–bone interface condition affect changes in bone density within the implanted acetabulum. In the case of Prosthesis 1 having fully bonded condition, bone resorption of 21–50% density reduction were predominantly observed in the cancellous bone underlying the implant (Fig. 4, Table 3). However, considering debonded interfacial condition, bone apposition of
55% increase in bone density was predicted in ROI 1, whereas bone resorption (1–20% density reduction) was predicted in the other three ROIs (Fig. 4, Table 3). Bone density reduction was less for Prosthesis 2 for both the implant–bone interface conditions (Fig. 5, Table 3). Bone apposition of 1–68% increase in bone density was observed in ROIs 1, 2 and 4, whereas only 2% reduction in bone density was observed in ROI 3, when implant–bone interface condition was fully bonded (Fig. 5, Table 3). For debonded implant– bone interface condition, increase in bone density of 5–87% was observed at ROI 1 and 2 (Fig. 5, Table 3). However, bone density reduction of 3–4% was observed at ROIs 3 and

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journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

1 2 3

1 2 3

Section 1 - 1

Section 1 - 1

Section 2 - 2

Section 1 - 1

Section 2 - 2

Section 2 - 2 Density (g·cm-3)

Section 3 - 3

Section 3 - 3

Section 3 - 3

1

2

1

2

1

2

3

4

3

4

3

4

Time Increment = 0

Time Increment = 780

Time Increment = 780

Fig. 4 – Bone density distribution in the implanted acetabulum for Prosthesis 1, lateral view and sectional views; (a) postoperative, (b) bonded interface condition: after equilibrium in bone remodelling, and (c) debonded interface condition: after equilibrium in bone remodelling.

4 (Fig. 5, Table 3). For Prosthesis 3, bone resorption of 5–21% decrease in bone density was predicted in ROIs 2–4, whereas bone apposition of 58–65% increase in bone density was predicted only in ROI 1, for both the implant–bone interface conditions (Fig. 6, Table 3).

4.

Discussion

Recent developments in acetabular implants suggest more flexible, wear resistant alternative bearing material with anatomical shaped acetabular component and material property close to cortical bone, in order to reduce wear and periprosthetic bone loss. Despite a few clinical and FE studies (Field et al., 2006; Manley et al., 2006; Latif et al., 2008), to the

authors’ knowledge, there is scarcity of data on bone remodelling around composite acetabular components. The goal of this study was to investigate the deviations in load transfer and to gain an insight into the adaptive bone remodelling around composite acetabular components having different geometries, material properties and implant–bone interface conditions. For this purpose, a time-dependent bone remodelling algorithm was developed and used in combination with an experimentally validated FE model (Ghosh et al., 2013b). Variations in geometries and implant–bone interface condition affect peri-prosthetic strain distributions and bone remodelling around implanted acetabulam. It is evident from the results that the effect of strain shielding was more for the implant–bone bonded interface as compared to the debonded

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

265

1 2 3

1 2 3

Section 1 - 1

Section 1 - 1

Section 1 - 1

Section 2 - 2

Section 2 - 2

Section 2 - 2

Density (g·cm-3 )

Section 3 - 3

Section 3 - 3

Section 3 - 3

1

2

1

2

1

2

3

4

3

4

3

4

Time Increment = 0

Time Increment = 780

Time Increment = 780

Fig. 5 – Bone density distribution in the implanted acetabulum for Prosthesis 2, lateral view and sectional views; (a) postoperative, (b) bonded interface condition: after equilibrium in bone remodelling, and (c) debonded interface condition: after equilibrium in bone remodelling.

interface condition, indicated by more load transfer through the cortex than the subchondral bone, similar to the findings of Thompson et al. (2002). Experiments conducted with maximum hip joint force of 1500 N, showed an increase of around 12% peri-acetabular cortex strain superior to the acetabulum due to implantation with CFR-PEEK material (Dickinson et al., 2012). Whereas, the increase in cortex strain in this study was found to be around 16% at similar locations for implantation with Prosthesis 2 and 3. The occurrence of high strains (0.7–1% strain) in the acetabular fossa for the horseshoe-shaped components (Prostheses 1 and 2) was comparable to the results reported earlier (Manley et al., 2006). The FE study by Manley et al. (2006) reported hemispherical acetabular component produces unphysiological strain distributions as compared to anatomical shaped designs. This observation was consistent with our results, wherein only strain shielding, a reduction in strain of 10–60%,

was observed for Prosthesis 3 in ROIs 1–4 for bonded implant– bone interface condition. Results of this study may be compared with available clinical findings (Field et al., 2006). The clinical study on Cambridge acetabular component implanted in 50 patients showed 13% reduction in bone density after six months, post-operatively (Field et al., 2006). Whereas in our study
15% reduction in bone density was observed in similar locations and time scale, when debonded interface condition was assumed. Furthermore, the clinical study reported BMD decrease in regions superior and medial to the acetabulum during the first six months, post-operatively. At the inferior side of the acetabulum, BMD decrease was observed until one year after surgery, with no significant changes in bone density thereafter. Our FE results predicted a decline in bone density at the inferior part of acetabulum for both the implant–bone interfacial conditions, which may be perhaps

266

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

1 2 3

1 2 3

Section 1 - 1

Section 1 - 1

Section 2 - 2

Section 1 - 1

Section 2 - 2

Section 2 - 2

Density (g·cm-3 ) Section 3 - 3

Section 3 - 3

Section 3 - 3

1

2

1

2

1

2

3

4

3

4

3

4

Time Increment = 0

Time Increment = 780

Time Increment = 780

Fig. 6 – Bone density distribution in the implanted acetabulum for Prosthesis 3, lateral view and sectional views; (a) postoperative, (b) bonded interface condition: after equilibrium in bone remodelling, and (c) debonded interface condition: after equilibrium in bone remodelling. Table 3 – Predicted changes in bone density after equilibrium in bone remodelling in four ROIs for different implants and implant–bone interface conditions. The positive sign (þ) indicates bone densification, whereas negative sign ( ) indicates bone resorption. Acetabular component

Implant–bone interface condition

Change in bone density (in percentage) ROI 1 (%)

ROI 2 (%)

ROI 3 (%)

ROI 4 (%)

Prosthesis 1

Fully bonded Debonded

 32 þ55

 50  17

 34  20

 21 1

Prosthesis 2

Fully bonded Debonded

þ68 þ87

þ1 þ5

2 4

þ14 3

Prosthesis 3

Fully bonded Debonded

þ58 þ65

5 7

 21  18

 10  18

due to the cup thickness or the cut-out of the horseshoeshaped Cambridge cup (Field et al., 2006). However, BMD increase after six month, as reported by Field et al. (2006), was

not predicted by our simulations, perhaps due to modelling inaccuracy as compared to in vivo condition. Stress shielding and bone resorption were found to be minimal for the

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

CFR-PEEK acetabular component, which is similar to the trends reported by earlier studies with different orthopaedic implants made of composite materials (Aitasalo et al., 2001; Ballo, 2008; Tuusa et al., 2008; Zhao et al., 2009). The thickness of the acetabular component also affects the stiffness of the implant and thereby, influences the changes in bone density distribution around the implanted acetabulum. Moreover, variation in component design affected the average bone deformation underlying a prosthesis. The average bone deformation for the thin horseshoe cup (Prosthesis 2) was very close to that of the intact acetabulum at comparable location, similar to the finding of Latif et al. (2008). Whereas, for the thicker cup (Prosthesis 1) and hemispherical design (Prosthesis 3), the bone deformations were less. Results predicted that peri-prosthetic bone density reductions were more for Prosthesis 1 than Prosthesis 2 (Figs. 4 and 5, Table 3). Considering bonded interface condition, only bone density reductions of 21–50% were observed in all ROIs 1–4 for pelvis implanted with Prosthesis 1. Whereas for the Prosthesis 2 having bonded interface, bone density increase of 1–68% in ROI 1, 2 and 4 and bone density decrease of
2% in ROI 3, were observed. For components having same thickness, the hemispherical design exacerbated bone resorption (Table 3). Hence it appears that the Prosthesis 2 is a better suited alternative bearing surface with regard to strain shielding, bone deformation and bone remodelling than other designs. The study, however, have a number of limitations, which are mostly similar to those reported previously (Ghosh et al., 2013b). In this study, bone geometry and density distribution were based on CT-scan data of one representative right hemipelvis. Assumption of isotropic cancellous bone material property seems justified, similar to other studies (Dalstra and Huiskes, 1995; Anderson et al., 2005; Zhang et al., 2010), since pelvic trabecular bone was shown not to be highly anisotropic (Dalstra et al., 1993). An adaptation rate (τ) of 129.6 g/mm2 (J/g) months for the bone remodelling simulation was assumed similar to that of the femur (Weinans et al., 1993). Results are presented using simulation time scale i.e. τ Δt (Suarez et al., 2012), in which the total time was calculated approximately as six months. Bone remodelling is known to be dependent on loading conditions; our study is limited to two static load cases during a normal walking cycle. The effect of damage accumulation in the bone tissue has not been considered in this study. Moreover, bone ingrowth on the porous coated implant was not actually simulated and instead, variable implant–bone interfacial condition was assumed (Suarez et al., 2012). A further study, combining mechano-regulatory bone-ingrowth and bone remodelling algorithms (Liu and Niebur, 2008), might be more useful for better understanding of the load transfer and bone density changes around the acetabular components. The FE models of the metallic and composite acetabular components were based on the same pelvis (CT-scan data set), enabling one-to-one comparisons of predicted results. Adaptive changes in bone density distributions for Prosthesis 1 were comparable to those of the metallic component (Ghosh et al., 2013b). In the case of metallic component having fully bonded implant–bone interface condition, 20–50% bone density reductions was observed in the cancellous bone

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underlying the metallic implant. At similar locations, a reduction of 21–50% bone density was observed for Prosthesis 1. However, an increase in 50–60% bone density in ROI 1 and a decrease in bone density of
1–20% were observed in other three ROIs for both the prostheses, having debonded implant–bone interfacial condition. A comparison between the results predicted by Prosthesis 1 and the metallic cup indicated similar changes in bone density distributions for the same implant–bone interface condition. However, Prosthesis 2 (CFR-PEEK with thickness 3 mm) offers more flexibility and considerably less peri-prosthetic bone resorption. The thickness of the acetabular component therefore, seemed to play a crucial role in the implant induced bone remodelling process.

5.

Conclusions

Based on the FE analysis of implanted pelvic bone models coupled with bone remodelling simulation, the following conclusions may be drawn. The study has been useful in understanding the deviations in load transfer and adaptive changes in bone density distributions due to implantation with composite acetabular components. Variations in geometry and implant– bone interfacial condition affected peri-prosthetic strain distributions and bone remodelling around the implanted acetabulum. Strain shielding was considerably higher for bonded implant– bone interface condition as compared to debonded implant– bone interface condition. Compared to the horseshoe-shaped design, the hemispherical design exacerbated bone resorption. Moreover, the thickness of the acetabular component played a crucial role in the implant induced bone remodelling process. A composite cup with thickness of 4.5 mm (resembling Cambridge cup) evoked similar changes in bone density distribution to those predicted by a 3 mm metallic cup. Whereas the CFR-PEEK component with thickness 3 mm (resembling MITCH PCRTM cup) offered more flexibility and considerably less periprosthetic bone resorption. This prosthesis is therefore, a better suited alternative bearing surface than other designs with regard to strain shielding, bone deformation and bone remodelling.

Acknowledgements The authors sincerely wish to thank Department of Biotechnology, New Delhi, for supporting this study. The authors are thankful to University of Southampton for providing the CT-scan data of the pelvic bone under the UK–India collaborative research project, sponsored by the UKIERI British Council. The authors are also thankful to Dr. Bidyut Pal, Mr. Kaushik Mukherjee and Mr. Souptick Chanda for their help and suggestions.

r e f e r e nc e s

Aitasalo, K., Kinnunen, I., Palmgren, J., Varpula, M., 2001. Repair of orbital floor fractures with bioactive glass implants. J. Oral Maxillofacial Surg. 59 (12), 1390–1395. Anderson, A.E., Peters, C.L., Tuttle, B.D., Weiss, J.A., 2005. Subjectspecific finite element model of the pelvis: development,

268

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

validation, sensitive studies. Trans. ASME J. Biomech. Eng. 127 (3), 364–373. Ballo, A.M., 2008. Fiber-reinforced Composite as Oral Implant Material: Experimental Studies of Glass Fiber and Bioactive Glass In Vitro and In Vivo. A Ph.D Thesis. University of Turku, Finland. Bergmann, G., Deuretzbacher, G., Heller, M., Graichen, F., Rohlmann, A., Strauss, J., Duda, G.N., 2001. Hip contact forces and gait patterns from routine activities. J. Biomech. 34 (7), 859–871. Bobyn, J.D., Mortimer, E.S., Glassman, A.H., Engh, C.A., Miller, J.E., Brooks, C.E., 1992. Producing and avoiding stress shielding: laboratory and clinical observations of noncemented total hip arthroplasty. Clin. Orthop. Relat. Res. 274, 79–96. Brooks, R.A., Jones, E., Storer, A., Rushton, N., 2004. Biological evaluation of carbon-fibre-reinforced polybutyleneterephthalate (CFR-PBT) employed in a novel acetabular cup. Biomaterials 25 (17), 3429–3438. Carter, D., Orr, T.E., Fyhrie, D., 1989. Relationships between bone loading history and femoral cancellous bone architecture. J. Biomech. 22 (3), 231–244. Clarke, S.G., Phillips, A.T. M., Bull, A.M. J., 2013. Evaluating a suitable level of model complexity for finite element analysis of the intact acetabulum. Comput. Methods Biomech. Biomed. Eng. 16 (7), 717–724. Dalstra, M., Huiskes, R., Odgaard, A., van Erning, L., 1993. Mechanical and textural properties of pelvic trabecular bone. J. Biomech. 26 (4–5), 523–535. Dalstra, M., Huiskes, R., 1995. Load transfer across the pelvis bone. J. Biomech. 28 (6), 715–724. Davis, E.T., Olsen, M., Zdero, R., Papini, M., Wadell, J.P., Schemetisch, E.H., 2009. A biomechanical and finite element analysis of femoral neck notching during hip resurfacing. Trans. ASME J. Biomech. Eng. 131 (4), 1–8 (041002). Dickinson, A.S., Taylor, A.C., Browne, M., 2012. The influence of acetabular cup material on pelvis cortex surface strains, measured using digital image correlation. J. Biomech. 45 (4), 719–723. Dostal, W.F., Andrews, J.G., 1981. A three-dimensional biomechanical model of hip musculature. J. Biomech. 14 (11), 803–812. Dumbleton, J.H., Manley, M.T., Edidin, A.A., 2002. A literature review of the association between wear rate and osteolysis in total hip arthroplasty. J. Arthroplasty 17 (5), 649–661. Field, R.E., Rushton, N., 2005. Five-year clinical, radiological and postmortem results of the Cambridge cup in patients with displaced fractures of the neck of the femur. J. Bone Joint Surg. Br. 87 (10), 1344–1351. Field, R.E., Cronin, M.D., Singh, P.J., Burtenshaw, C., Rushton, N., 2006. Bone remodelling around Cambridge cup: a DEXA study of 50 hips over 2 years. Acta Orthop. 77 (5), 726–732. Field, R.E., Jones, E., Nuijten, P., Storer, A., Cronin, M.D., Rushton, N., 2008. Pre-clinical evaluation of the Cambridge acetabular cup. J. Mater. Sci. Mater. Med. 19 (8), 2791–2798. Garcia, J.M., Doblare, M., Cegonino, J., 2002. Bone remodelling simulation: a tool for implant design. Comput. Mater. Sci. 25 (1–2), 100–114. Ghosh, R, Gupta, S, Dickinson, A, Browne, M., 2012. Experimental validation of finite element models of intact and implanted composite hemi-pelvises using digital image correlation. Trans. ASME J. Biomech. Eng. 134 (8), 1–9 (081003). Ghosh, R., Gupta, S., Dickinson, A., Browne, M., 2013a. Experimental validation of numerically predicted strain and micromotion in intact and implanted composite hemipelvises. J. Eng. Med. Proc. Inst. Mech. Eng. H 227 (2), 162–174. Ghosh, R., Mukherjee, K., Gupta, S., 2013b. Bone remodelling around uncemented metallic and ceramic acetabular

components. J. Eng. Med. Proc. Inst. Mech. Eng. H 227 (5), 490–502. Ghosh, R., Pal, B., Ghosh, D., Gupta. Finite element analysis of a hemi-pelvis: the effect of inclusion of cartilage layer on acetabular stresses and strain. Comput. Methods Biomech. Biomed. Eng. in press-c, http://dxdoi.org/10.1080/ 10255842.2013.843674. Harris, W.H., 1995. The problem is osteolysis. Clin. Orthop. Relat. Res. 311, 46–53. Hedia, H.S., Abdl-Shafi, A.A., Fouda, N., 2000. The effect of elastic modulus of the backing material on the fatigue notch factor and stress. Biomed. Mater. Eng. 10 (3–4), 141–156. Huiskes, R., Weinans, H., Grootenboer, H.J., Dalstra, M., Fudala, B., Sloof, T.J., 1987. Adaptive bone remodelling theory applied to prosthetic design analysis. J. Biomech. 20 (11–12), 1135–1150. Huiskes, R., Weinans, H., van Rietbergen, B., 1992. The relationship between stress shielding and bone resorption around total hip stems and the effects of flexible materials. Clin. Orthop. Relat. Res. 274, 124–134. Huiskes, R., van Rietbergen, B., 1995. Preclinical testing of total hip stem: the effects of coating placement. Clin. Orthop. Relat. Res. 319, 64–76. Janssen, D., Zwartele, R.E., Doets, H.C., Verdonschot, N., 2010. Computational assessment of press-fit acetabular implant fixation: the effect of implant design, interference fit, bone quality, and frictional properties. J. Eng. Med. Proc. Inst. Mech. Eng. H 224 (1), 65–75. Kempson, G., 1980. The mechanical properties of articular cartilage. In: Fluid, L. Sokoloff (Ed.), The Joint and Synovial. Academic, New York, pp. 385–409 (Chapter 5). Konrath, G.A., Hamel, A.J., Olson, S.A., Bay, B., Sharkey, N.A., 1998. The role of the acetabular labrum and the transverse acetabular ligament in load transmission of the hip. J. Bone Joint Surg. Am. 80 (12), 1781–1788. Latif, A.M. H., Mehats, A., Elcocks, M., Rushton, N., Field, R.E., Jones, E., 2008. Pre-clinical studies to validate the MITCH PCRTM cup: a flexible and anatomical shaped acetabular component with novel bearing characteristics. J. Mater. Sci. Mater. Med. 19 (4), 1729–1736. Leung, A.S. O., Gordon, L.M., Skrinskas, T., Szwedowski, T., Whyne, C.M., 2009. Effects of bone density alterations on strain patterns in the pelvis: application of a finite element model. J. Eng. Med. Proc. Inst. Mech. Eng. H 223 (8), 965–976. Lewis, J.L., Askew, M.J., Wixson, R.L., Kramer, G.M., Tarr, R.R., 1984. The influence of prosthetic stem stiffness and of a calcar collar on stresses in the proximal end of the femur with a cemented femoral component. J. Bone Joint Surg. Am. 66 (2), 280–286. Liu, X., Niebur, G.L., 2008. Bone ingrowth into a porous coated implant predicted by a mechano-regulatory tissue differentiation algorithm. Biomech. Model. Mechanobiol. 7 (4), 335–344. Manley, M.T., Ong, K.L., Kurtz, S.M., 2006. The potential for bone loss in acetabular structures following THA. Clin. Orthop. Relat. Res. 453, 246–253. Manley, M.T., Sutton, K., 2008. Bearings of the future for total hip arthroplasty. J. Arthroplasty 23 (7), 47–50. Martin, R.B., 1972. Effects of geometric feedback in development of osteoporosis. J. Biomech. 5 (5), 447–455. Morgan, E.F., Keaveney, T.M., 2001. Dependence of yield strain of human trabecular bone on anatomic site. J. Biomech. 34 (5), 569–577. Morscher, E., Berli, B., Jockers, W., Schenk, R., 1997. Rationale of a flexible press fit cup in total hip replacement: 5-year followup in 280 procedures. Clin. Orthop. Relat. Res. 341, 42–50. Pal, B., Gupta, S., New, A.M. R., 2010. Influence of the change in stem length on the load transfer and bone remodelling for a cemented resurfaced femur. J. Biomech. 43 (15), 2908–2914.

journal of the mechanical behavior of biomedical materials 32 (2014) 257 –269

Schmalzried, T.P., Akizuki, K.H., Fedenko, A.N., Mirra, J., 1997. The role of access of joint fluid to bone in periarticular osteolysis: a report of four cases. J. Bone Joint Surg. Am. 79 (3), 447–452. Scholes, S.C., Unsworth, A., 2007. The wear properties of CFR-PEEK-OPTIMA articulating against ceramic assessed on a multidirectional pin-on-plate machine. J. Eng. Med. Proc. Inst. Mech. Eng. H 221 (3), 281–289. Scholes, S.C., Inman, I.A., Unsworth, A., Jones, E., 2008. Tribological assessment of a flexible carbon-fibre-reinforced poly (ether-ether-ketone) acetabular cup articulating against an alumina femoral head. J. Eng. Med. Proc. Inst. Mech. Eng. H 222 (3), 273–383. Shirazi-Adl, A., Dammak, M., Paiement, G., 1993. Experimental determination of friction characteristics at the trabecular bone/porous coated metal interface in cementless implants. J. Biomed. Mater. Res. 27, 167–175. Suarez, D.R., Weinans, H., van Keulen, F., 2012. Bone remodelling around a cementless glenoid component. Biomech. Model. Mechnobiol. 11 (6), 903–913. Taddei, F., Pancanti, A., Viceconti, M., 2004. An improved method for the automatic mapping of computed tomography numbers onto finite element models. Med. Eng. Phys. 26 (1), 61–69. Thompson, M.S., Northmore-Ball, M.D., Tanner, K.E., 2002. Effect of acetabular resurfacing component material and fixation on the strain distribution in the pelvis. J. Eng. Med. Proc. Inst. Mech. Eng. H 216 (4), 237–245. Tuusa, S.M. R., Peltola, M.J., Tirri, T., Puska, M.A., Roytta, M., Aho, H., Sandholm, J., Lassila, L.V. J., Vallittu, P.K., 2008. Reconstruction of critical size calvarial bone defects in rabbit with glass-fiber-reinforced composite with bioactive glass granule coating. J. Biomed. Mater. Res. B: Appl. Biomater. 84 (2), 510–519.

269

Unsworth, A., Dowson, D., Wright, V., 1975. The frictional behaviour of human synovial joints – Part I: natural joints. J. Lubr. Technol. 97 (3), 369–376. van Rietbergen, B., Huiskes, R., Weinans, H., Sumner, D.R., Turner, T.M., Galante, J.O., 1993. The mechanism of bone remodelling and resorption around press-fitted THA stems. J. Biomech. 26 (4–5), 369–382. Varghese, B., Short, D., Penmetsa, R., Goswami, T., Hangartner, T., 2011. Computer-tomography-based finite-element models of long bones can accurate capture strain response to bending and torsion. J. Biomech. 44 (7), 1374–1379. Verhulp, E., van Rietbergen, B., Muller, R., Huiskes, R., 2008. Indirect determination of trabecular bone effective tissue failure properties using micro-finite element simulations. J. Biomech. 41 (7), 1479–1485. Weinans, H., Huiskes, R., van Reitbergen, B., Sumner, D.R., Turner, T.M., Galante, J.O., 1993. Adaptive bone remodelling around bonded noncemented total hip arthroplasty: a comparison between animal experiments and computer simulation. J. Orthop. Res. 11 (4), 500–513. Zioupos, P., Cook, R.B., Hutchinson, J.R., 2008. Some basic relationships between density values in cancellous and cortical bone. J. Biomech. 41 (9), 1961–1968. Zhang, Q.H., Wang, J.Y., Lupton, C., Adegbile, P.H., Guo, Z.X., Liu, Q., Tong, J., 2010. A subject-specific pelvic bone model and its application to cemented acetabular replacements. J. Biomech. 43 (14), 2722–2727. Zhao, D.S., Moritz, N., Laurila, P., Mattila, R., Lassila, L.V. J., Strandberg, N., Mantyla, T., Vallittu, P.K., Aro, H.T., 2009. Development of a multi-component fiber-reinforced composite implants for load shearing conditions. Med. Eng. Phys. 31 (4), 461–469.