Computer Methods in Biomechanics and Biomedical Engineering
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Hip joint degeneration due to cam impingement: a finite element analysis F.L. Hellwig, J. Tong & J.G. Hussell To cite this article: F.L. Hellwig, J. Tong & J.G. Hussell (2016) Hip joint degeneration due to cam impingement: a finite element analysis, Computer Methods in Biomechanics and Biomedical Engineering, 19:1, 41-48, DOI: 10.1080/10255842.2014.983490 To link to this article: http://dx.doi.org/10.1080/10255842.2014.983490
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Date: 05 November 2015, At: 14:27
Computer Methods in Biomechanics and Biomedical Engineering, 2016 Vol. 19, No. 1, 41–48, http://dx.doi.org/10.1080/10255842.2014.983490
Hip joint degeneration due to cam impingement: a finite element analysis F.L. Hellwiga, J. Tonga* and J.G. Hussellb a
Mechanical Behaviour of Materials Group, School of Engineering, University of Portsmouth, Portsmouth, UK; bOrthopaedics and Trauma Centre, Queen Alexandra Hospital, Portsmouth, UK
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(Received 20 August 2012; accepted 30 October 2014) The goal of this study was to investigate the impact of cam impingement, a biomechanical risk factor, on hip joint degeneration and ultimately coxarthrosis. 3D finite element solid models of a healthy and a pathologic hip were developed based on clinical reports. The biphasic characteristics of cartilaginous tissues were considered to identify localised solid matrix overloading during normal walking and sitting down (SD). Localised femoral intrusion at the anterior-superior pelvic horn was revealed in the pathologic hip during SD, where the radial and meridional solid stresses in the acetabular cartilage and circumferential solid stresses within the acetabular labrum increased by 3.7, 1.5 and 2.7 times, respectively. The increased solid-on-solid stresses, reduction in fluid-load support and associated higher friction during articulation may result in joint wear and other degenerative changes in the hip. Keywords: hip joint; cartilage degeneration; biphasic model; contact pressure; finite element; impingement
Introduction Excessive contact stresses in the hip joint due to obesity (Recnik et al. 2009) or decreased weight-bearing area typical in a dysplastic hip (Mavcˇicˇ et al. 2002; Russel et al. 2006) are thought to result in coxarthrosis. Recent clinical studies hypnotised a link between abnormal anatomical conditions of femoral head – neck junction and early degenerative changes of the hip joint, such as osteochondral defects and labral tear leading ultimately to idiopathic osteoarthritis (Ito et al. 2001). The pathologic condition, named cam-type femoroacetabular impingement, is caused by jamming of a non-spherical extension of the femoral head into the acetabular cavity (Ganz et al. 2003, 2008). No¨tzli et al. (2002) characterised this anatomic abnormality by introducing an alpha angle (Figure 1). This angle is measured between the femoral neck axis and a line extended from the point C, the femoral head centre, to point D, which is the point of deviation from femoral head sphericity. The abnormality might emerge as an anterior bump, spigot or thickened femoral neck, limiting the range of motion (ROM; Ito et al. 2001; Ganz et al. 2003, 2008). Recently, a finite element study (Chegini et al. 2009) of normal and impinged articulation has been presented to examine the joint mechanics during daily activities of normal walking (NW) and sitting down (SD), for which the maximum contact force was reported as 233% and 156% bodyweight, respectively (Bergmann et al. 2001). Chegini et al. (2009) focused on the total stress responses of the cartilaginous tissues, which were represented as isotropic linear elastic materials. The cartilage and labrum, however, are fully fluid saturated porous mediums (Mow
*Corresponding author. Email: [email protected] q 2015 Taylor & Francis
et al. 1980); hence, the load is shared between the interstitial fluid and a solid extracellular matrix (ECM):
stot ¼ ssolid þ Dpl;
ð1Þ
where stot is the total stress of the tissue, ssolid is the stress acting on the ECM and Dpl is the pore pressure of the interstitial fluid (Mow et al. 1980). Early macroscopic changes in articular cartilage were reported to begin with fraying at the superficial zone of the ECM (Pritzker et al. 2006), a sign of joint wear accelerated with increased friction (Mow and Huiskes 2005) and associated with decreased fluid-load support (FLS; Krishnan et al. 2004). Ferguson et al. (2000a) utilised an isotropic poroelastic material model in ABAQUS (Wu et al. 1998) to simulate the sealing function of the acetabular labrum. They correlated alterations in solid-on-solid stress with the changes in FLS for astatic load up to 75% bodyweight held constant for up to 10,000 s. More recently, in order to study the cartilage response to elevated loading during unstable motion, Goreham-Voss et al. (2007) modelled the articulating cartilage layers in knee joints, considering the directional-dependent material properties of the ECM. The aim of this work is to study the stresses in the cartilage in the principal directions of an orthotropic ECM and the FLS during normal and ‘cam’-type impinged articulations under selected physiological loading conditions. To do so, the cartilaginous tissues were modelled as biphasic orthotropic materials, and the results for normal and ‘cam’-type impinged articulation were compared.
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F.L. Hellwig et al. and 4.98 £ 1024 mm4/Ns, respectively (Athanasiou et al. 1994; Ferguson et al. 2001), and they were converted from the biphasic, k, to the poroelastic model, k0 , via:
Anterior Anterior Hump
D
k0 ¼ gk;
C
Medial
Lateral
α
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Posterior
Figure 1. Illustration of a pathologic femur (anterior hump): the deviation of femoral head sphericity is measured by the a angle between the femoral neck axis, passing though the femoral neck and head centre C, and a line extended from the point C– D, the point of deviation from femoral head sphericity (edge of hump).
Numerical methods Geometry and materials A normal hip joint model with a ¼ 408 was developed based on morphologic simplifications and dimensions of Chegini et al. (2009; Figure 2(a)), while a typical camimpinged hip joint with a ¼ 748 was also modelled based on No¨tzli et al. (2002) (Figure 2(b)), using ABQUS CAE. The finite element model, depicted in Figure 2(c), of the normal hip joint (c) consists of acetabular labrum (1), acetabular cartilage (2), femoral cartilage (3), impermeable membrane (4) and femoral head neck section (5), following (Chegini et al. 2009). The pathologic case was identical except the anterior hump at the head–neck junction. Femoral and pelvic bones were considered rigid (Anderson et al. 2008) and the cartilaginous tissues were considered as fully fluid-saturated porous mediums (Mow et al. 1980), taking into account the orthotropic structure of the ECM following Goreham-Voss et al. (2007). A spherical coordinate system was used to allocate the directionaldependent material properties. The permeabilities of the femoral and acetabular cartilage and labrum are 9.15, 8.98
(a)
(b)
ð2Þ
where the volumetric weight of the interstitial water, g, was taken as 9.81 £ 1026 N mm23 (Wu et al. 1998). The solid matrix compressive radial moduli ‘1’of femoral and acetabular cartilage and labrum were averaged to 1.23, 1.18 and 0.157 MPa (Athanasiou et al. 1994; Ferguson et al. 2001), respectively. For cartilage, the tangential moduli in circumferential ‘2’ and meridional ‘3’ directions were extracted from the toe region as 8.5 MPa (Roth and Mow 1980). To permit opening between the acetabular labrum and femoral cartilage under load, a tangential moduli of 26 MPa was chosen in direction ‘2’, close to that of the toe region (20 MPa) reported in Ferguson et al. (2001). According to the ratio of meridional to circumferential strength of the meniscus (Tissakht and Ahmed 1995), the tangential modulus of the labrum in direction ‘3’ was set to 3 MPa, since labrum and meniscus exhibit a similar ultra structure (Ferguson et al. 2001). As shear stresses peak at the cartilage–bone interface (Goreham-Voss et al. 2007), a shear modulus of 2.5 MPa (Wong et al. 2008) was chosen. Shear modulus and volumetric response (n12, n23; Athanasiou et al. 1994; Federico et al. 2005) are given for cartilage in Table 1, similar values were assigned to the labrum due to the lack of data. A summary of all material properties used in the models is given in Table 1. As tissue properties were taken from different sources due to limited availability, a pre-study on the sensitivity to material property variation (^ 50%) was performed. The peak contact force for NW, taken from Bergmann et al. (2001), was applied to the hip joint centre during a 1 s ramp phase for the normal articulation.
Finite element modelling The cartilaginous tissues were meshed with first-order pore pressure brick elements, and the femoral-head neck
(c)
(1) (2) (3) (4) (5)
Figure 2. The solid models for a normal (a) and a pathologic femur with hump (b). The finite element model of the normal hip joint (c) consists of acetabular labrum (1), acetabular cartilage (2), femoral cartilage (3), an impermeable membrane (4) and femoral head neck section (5), following (Chegini et al. 2009).
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Computer Methods in Biomechanics and Biomedical Engineering Table 1.
Material properties used in this study. Biphasic formulation
Tissue Acetabular labrum Acetabular cartilage Femoral cartilage
E1 (MPa)
E2 (MPa)
E3(MPa)
n12
n13
n23
G (MPa)
k0 (mm/s)
F
0.157a 1.23b 1.18b
26a 8.5c 8.5c
3e 8.5c 8.5c
0.04 0.044b 0.046b
0.04 0.044b 0.046b
0.1 0.146d 0.146d
1.5 2.5f 2.5f
4.89b £ 10209 8.52d £ 10209 8.98d £ 10209
4 4 4
a
Ferguson et al. (2001). Athanasiou et al. (1994). c Roth and Mow (1980). d Federico et al. (2005). e Tissakht and Ahmed (1995). f Wong et al. (2008).
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b
section and the impermeable membrane were meshed with quadrilateral bilinear rigid and quadrilateral membrane elements, respectively. Reduced integration was used for deformable elements to reduce the calculation time. The selected grid sizes and number of elements are given in Table 2. Because the pelvis was assumed rigid, the nodes of the acetabular cartilaginous tissue on the bone-cartilage interface were fully constrained. The hip joint forces for NW and SD were taken from Bergmann et al. (2001) and applied incrementally as a point force to the centre of the femoral head. The hip joint force was synchronised with the locomotion data prescribed as rotational movements of the femur. The analysis was performed by first establishing contact between the femur and the acetabulum, then applying an initial load of 33.31% or 55.51% of the peak load for NW and SD, respectively, ramped over 1 s, and finally simulating the locomotion cycle in 60 sequences. The bodyweight was assumed 86 kg. Contact was implemented using the penalty method assuming frictionless finite sliding, due to small coefficient of friction (Pawaskar et al. 2007). Throughout the analysis, the femur was the master and the acetabulum the slave surface. Further, for poroelastic mediums, the contact model should satisfy the biphasic jump condition, where in areas of contact fluid velocity should be continuous or otherwise interstitial fluid should be free-draining (Hou et al. 1989). However, as contact pressure varies, local rehydration occurs (Pawaskar et al. 2007) and fluid efflux and cross-flow are minimal for loadings such as NW and SD (Pawaskar et al. 2011); hence, the biphasic jump
Table 2.
condition was not considered. Because in a biphasic medium load is shared between the fluid and the solid phase (Mow et al. 1980), the total contact pressure is carried by both phases. However, in the ABAQUS model for porous mediums in contact, the fluid phase contribution to the total contact pressure is ignored to ensure point-wise continuity of the pore pressure at opposing sides (Hibbitt et al. 2007). On the other hand, contact pressure between a porous and an impermeable material includes contributions from both solid and fluid phase (Hibbitt et al. 2007). In this work, the total contact pressure was obtained by inserting an impermeable membrane (E ¼ 1 MPa; n ¼ 0.49) between the cartilaginous tissues (Wilson et al. 2003; Manda et al. 2011). This membrane has only a minor effect on contact pressure results (Wilson et al. 2003).
Results Normal articulation For both NW and SD, the weight-bearing area, peak contact and peak pore pressure (P-POR) vary during the course of the simulated activity, and the variation is consistent with that of the applied contact force. The peak contact pressures (P-CPRESSs) of 2.86 and 3.66 MPa are located in the superior region for NW and the posterior region for SD, respectively. However, weight-bearing area changes from superior – anterior to superior – posterior for NW and from superior to posterior during SD; hence, the contact and pore pressure are distributed asymmetrically about the applied load. The timing of the occurrence of
The number of elements and grid sizes used in the FE models of the normal and the abnormal hip joint. Healthy joint
Acetabular cartilage and labrum Femoral cartilage Femoral head – neck section Impermeable membrane
Abnormal joint
Element type
Number
Grid size
Number
Grid size
C3D8RP C3D8RP R3D4 M3D4R
21,460 13,666 9956 4004
0.70 0.90 0.90 1.15
21,460 12,420 8767 4166
0.70 1.00 1.00 0.95
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Table 3.
Peak biphasic response for NW and SD for normal articulation. Biphasic response Location
P-CPRES (MPa)
CAREA (mm )
P-POR (MPa)
S11
Time of cycle (%)
Superior Posterior– medial
2.87 3.58
1798 1574
2.84 3.54
2 0.03 2 0.04
54 46.23
Activity
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NW SD
2
‘Cam’-type impinged articulation The pathologic hip joint revealed localised impingement at the anterior –superior pelvic horn during deep flexion. Hence, pore pressure increased significantly by eight times from 0.42 to 3.76 MPa at the anterior – superior horn due to the intrusion of the femoral hump (Figure 4). Compressive radial stresses, within the acetabular cartilage, increased at the impinged location by 3.7 times from 0.07 to 0.33 MPa. Tensile circumferential and meridional stresses are at peaks within the acetabular labrum and at the cartilage – labral interface, respectively (Figure 4(a),(b)), and increased by a factor of 2.7 and 1.5, respectively. S22 and S33 rose from 0.84 to 2.26 MPa and 0.92 to 1.37 MPa (Figure 4(a),(b)). FLS in the impinged zone dropped to 91.9%, as opposed to 98.4% in the posterior region. Shear stress also doubled to 2.61 MPa on the acetabular surface and surpassed the maximum at the acetabular cartilage – bone interface in the normal articulation.
peak contact and P-POR coincided with the peak contact force. The P-CPRESS, the related contact area (CAREA), the timing and the location of their occurrence for NW and SD are summarised in Table 3. In addition, the P-POR, which corresponded to the location of the P-CPRESS, and the related compressive radial stress (S11) are also summarised in Table 3. For NW, a plot of the results, including the stress components of the solid matrix, is given in Figure 3. The locations of maximum P-CPRESS (2.87 MPa) and maximum P-POR (2.85 MPa) are shown in Figure 3(a), (b), where the distribution patterns are identical. Stresses within the ECM in the three directions are highest on the articulating surface. The highest stress concentrations, radial stress 2 0.05 MPa and meridional stress 0.50 MPa, are found on the edges of the acetabular cartilage, where the pore pressure is minimal (Figure 3(c),(e)). Circumferential stress is the highest in the acetabular labral ring (0.33 MPa, Figure 3(e)). Shear stress, however, peaked at the cartilage – bone interface along the edge of the acetabular cartilage, rising up to 1.35 MPa (Figure 3(f)). The ratio of fluid load-support is 99.3% and 98.4% for NW and SD, respectively, indicating a similarity in fluid –solid interaction in both cases.
(a)
contact pressure (CPRESS)
A
max = 2.87 MPa
(b)
Effect of material property variation The influence of material property variation on some of the key parameters such as femoral head displacement, contact
max = 2.84 MPa
pore pressure (POR)
(c)
fluid velocity (FLVEL)
A-A
A (d) radial stress (S11)
A-A
(e) circumferential stress (S22)
A-A
(f)
meridional stress (S11)
A-A
(g) Tresca stress (S, Tresca)
A-A
Figure 3. The P-CPRESS (MPa) (a) during NW; the related pore pressure (MPa) (b) and the stress components (MPa) in the solid matrix: radial (c), circumferential (d), meridional (e) and Tresca stress (f), section A – A as indicated in (a).
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Computer Methods in Biomechanics and Biomedical Engineering (a) 'cam'-type impinged articulation pore pressure (POR)
radial stress (S11)
circumferential stress (S22)
meridional stress (S33)
Tresca stress (S, Tresca)
radial stress (S11)
circumferential stress (S22)
meridional stress (S33)
Tresca stress (S, Tresca)
(b) normal articulation
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pore pressure (POR)
Figure 4. Pore pressure for normal (a) and CAM-impinged (b) acetabular labral and cartilage exposure during SD: Radial, circumferential, meridional and Tresca stresses (MPa) within the ECM.
pressure and pore pressure is summarised in Figure 5. The highest model sensitivity was recorded for a reduction of 50% in shear modulus, with an increase of 46.18% of femoral head displacement, although the contact pressure and the pore pressure dropped only marginally by 2.89% and 3.68%. Reducing the tangential modulus by 50% resulted in only marginal influence on femoral head displacement (2 4.28%), contact (2.7%) and pore pressure (1.97%), respectively. For all other variations the changes in the key parameters are insignificant , 1%. This gives us confidence that the observed enhanced responses in impinged case are indeed due to the abnormal articulation, and the influence of model parameters are insignificant except the shear modulus.
Discussion The goal of this study was to explore how ‘cam’-type impinged articulation can affect cartilage stresses compared with those under normal articulation; and whether or not cartilage structural integrity may be assessed based on a comparison of the total contact pressure, fluid FLS and generated stress components within the solid ECM in these two conditions. Special emphasis was placed on the consideration of orthotropic properties of the cartilaginous tissue structures. Comparisons with previous work seem to suggest that the locations identified for P-CPRESS during NW and SD are consistent with those published (Yoshida et al. 2006; Chegini et al. 2009). The maximum contact pressures in a normal joint are found to be 2.87 and 3.66 MPa during NW
and SD, compared well with those (2.35 and 3.34 MPa, respectively) reported by Chegini et al. (2009). Other subject-specific models, however, predicted higher PCPRESS, such as 6.6 MPa in Jorge et al. (2014) and 10.78 MPa in Anderson et al. (2008). According to Anderson et al. (2010), cartilage thickness variation accounted for more than doubled P-CPRESS compared with a constant cartilage thickness. Harris et al. (2012) reported a P-CPRESS rage of 7.51 ^ 2.11 MPa for walking, and suggested that the P-CPRESS varies with subjects. Hence, it seems reasonable that our results are within the range of reported data, given the variation of the joint geometries. Experimentally measured P-CPRESSs seem to be higher than predicted by the simplified FE models. According to Afroke et al. (1987), P-CPRESS varied strongly with subjects (4.9 –10.2 MPa). Anderson et al. (2008) reported a rather inhomogeneous contact pressure distribution with a peak of 10 MPa for NW. In contrast to that Brown and Shaw (1982) reported a P-CPRESS of 3.45 MPa, close to our findings. To the best of the authors’ knowledge, the stresses within the solid phase have not been measured in a cadaveric hip joint. Before this study, ECM of articular cartilage was assumed to deform linear elastic isotropically (Macirowski et al. 1994; Ferguson et al. 2000a, 2000b) or hyperelastic isotropically (Haemer et al. 2012), as the material properties of the cartilage were taken from uniaxial compression tests (Mow et al. 1980; Macirowski et al. 1994; Wu and Herzog 2000) where shear cannot be measured. To accommodate the multi-directional cartilage
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F.L. Hellwig et al. A
Femoral Head Displacement G
Ec
v12/13
v23
k –0.11%
–0.09%
2.5%
High
Low
–0.77%
–3.58%
–0.67%
–17.99%
0.07%
0.19%
4.23%
46.18%
–2.5%
0.08%
0.0%
0%
0.86%
Deviation
25%
–25%
Et
5.0%
50%
–5.0%
–50% Low
High
Low
High
Low
High
Low
High
Low
High
Contact Pressure B Et
Ec
v12/13
v23
k
0.00%
–0.02%
–0.19%
–1.79%
–0.13%
–2.89%
High
Low
0.02%
0.14%
0.11%
0.33%
–2.5%
2.70%
0.0% 1.95%
Deviation
2.5%
–5.0% Low
High
Low
High
Low
High
Low
High
Low
High
Pore Pressure
C G
Et
Ec
v12/13
v23
k
–0.05%
–0.02%
High
Low
High
–0.01%
–1.64%
–3.68%
2.5%
–0.75%
5.0%
Low
High
0.01%
0.03%
Low
0.14%
High
1.90%
–2.5%
0.97%
0.0% 2.18%
Deviation
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G 5.0%
–5.0% Low
High
Low
Low
High
Figure 5. The variation in the femoral head displacement, contact pressure and pore pressure due to the variation in the material parameters (low: 2 50%; high: þ50%).
deformation while using linear or hyperelastic isotropic properties, either lateral extension was permitted (Macirowski et al. 1994) and applied load reduced (Ferguson et al. 2000a), or the stress-bearing area increased, neglecting the lunate shape of the acetabular cartilage (Haemer et al. 2012). None of these modelling techniques seem to be sufficient to answer the current research questions. Taking the different modelling techniques into account, the high FLS of 99.3% during NW, compares reasonable well with 98.0% and 96.0% obtained by Haemer et al. (2012) and Macirowski et al. (1994) for a ramp load of 1 s, but less favourably with 91.7% (following 1000 static load) reported by Ferguson et al. (2000a).
In the pathologic case, a ¼ 748, the anterior protrusion produced an intrusion into the anterior – superior acetabular cavity during SD, which is consistent with earlier observations (Chegini et al. 2009; Jorge et al. 2014). Although the intrusion of the abnormal protrusion did not yield significant elevated P-CPRESS, the FLS was affected and the stress components of the ECM elevated. Although the FLS was still above 90%, the radial compressive stress (or solid-on-solid) increased significantly (3.7 times) and the ECM was not shielded from elevated solid-on-solid stress (Athesian and Mow 2005). This may explain some of the early macroscopic changes in the cartilage found in clinical observation (Ganz et al.
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Computer Methods in Biomechanics and Biomedical Engineering 2003, 2008; Beck et al. 2005), leading to severe osteoarthritis (Pritzker et al. 2006). Circumferential stress in the labrum was doubled, although this might not be as significant due to the labral design to withstand high exposure in this direction (Ferguson et al. 2001). The elevated stresses are mainly found in the cartilage-ECM, as opposed to those reported by Chegini et al. (2009) in cartilage and labrum simultaneously. The increased meridional stresses within the cartilage at the cartilage–labral interface seems to support the hypothesis that cam impingement results in cartilage defects and subsequent separation of the acetabular cartilage from the labrum (Ganz et al. 2003, 2008; Beck et al. 2005), and not otherwise as discussed (MacCarthy et al. 2001). The emphasis of earlier numerical work (Chegini et al. 2009; Jorge et al. 2014) was on severe abnormal femoral protrusions at high loads. Chegini et al. (2009) predicted, dependent on the morphology of the femoral neck, PCPRESSs are in the range of 3.68–12.84 MPa for ‘cam’type impingement. Jorge et al. (2014) also reported high PCPRESSs up to 16.4 MPa. Due to the varied abnormal conditions presented, the results cannot be readily compared with the current ones. Nonetheless, Ganz et al. (2003) suggested that mild to unrecognisable, in the eye of an observer, abnormal femora are the genesis of impingement. To the best of the authors knowledge, tissue stresses during ‘cam’-type impinged articulation have not been measured. Admittedly, we only considered a typical case of impingement and limited loading scenarios. The geometries considered are simplified versions based on CT images and reported anatomic characteristics, although we believe they contain the essence of a typical case of ‘cam’-type impingement. The finite element mesh was optimised to obtain convergence during ‘cam’-type impinged articulation; the chosen grid size was the optimum from a sensitivity study. Although the normal and the abnormal joints were meshed slight differently, it should not affect the results. The grid-ratios on both acetabular and femoral side were chosen carefully to obtain contact convergence. The slope of the contact-pressure over closure relationship was also adjusted to minimise nodal penetration and contact pressure errors. Further, loading scenarios considered here are limited to NW and SD. In certain sporting activities, such as martial arts, a higher ROM is required hence higher stresses due to impingement may be expected. As in all numerical work, experimental verification would be highly desirable, but it is not yet possible to quantify the load sharing between the fluid and solid phase or measure the strains within the ECM directly. Nonetheless, the key parameters were found to be generally insensitive to the variation of material properties even by variation of ^50% (Figure 5). Because linear behaviour was considered for the fluid (permeability) and the solid phase (elastic orthotropic), the error during a cam-impinged articulation should be similar to that determined during static loading.
47 7
The shear modulus, however, does seem to affect the femoral head displacement, thus cartilage consolidation. Mansours (2009) argued that, during cartilage compaction, elevated shear stresses will occur at the cartilage–bone interface, see Figure 3(f), as the cartilage layers are next to the much stiffer underlying subchondral bone layer, and high lateral contraction is a result of compression of the opposing cartilage layers. Hence, it is important to consider the orthotropic properties of the cartilage ECM when simulating multi-directional deformation of cartilage. Conclusions Results from a biphasic analysis of a healthy and a camimpinged joint show that ‘cam’-type impingement can elevate stresses significantly in the ECM of cartilaginous tissues. During SD, the peak radial and the peak meridional stresses within the cartilage and the circumferential stress within the acetabular labrum increased by 3.7, 2.7 and 1.5 times, respectively, compared with those in the normal joint. P-POR also increased at the anterior-superior ‘cam’-type impinged zone. The increased solid-on-solid stresses and associated decrease in FLS would result in higher friction during articulation and hence joint wear. In summary, ‘cam’-type impingement is a risk factor that could increase cartilaginous tissue strains leading to joint wear and other degenerative changes in the hip, consistent with clinical findings.
Acknowledgements The authors gratefully acknowledge access to the Sciama High Performance Computer (HPC) Cluster, which was supported by the ICG, SEPNet and the University of Portsmouth and support by Sciama technician Gary Burton.
Conflict of interest disclosure statement No potential conflict of interest was reported by the authors.
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Beck M, Kahlor M, Leunig M, Ganz R. 2005. Hip morphology influences the pattern of damage to the acetabular cartilage. J Bone Joint Surg Br. 87:1012– 1018. Bergmann G, Deuretzbacher G, Heller M, Graichen F, Rohlmann A, Strauss J, Duda GN. 2001. Hip contact forces and gait patterns from routine activities. J Biomech. 34(7):859– 871. Brown TD, Shaw DTA. 1982. A technique for measuring instantaneous in vitro contact stress distributions in articular joints. J Biomech. 15:329 – 331. Chegini S, Beck M, Ferguson SJ. 2009. The effects of impingement and dysplasia on stress distribution in the hip joint during sitting and walking: a finite element analysis. J Orthop Res. 27:195 – 201. Federico S, Grillo A, La Rosa G, Giaquinta G, Herzog W. 2005. A transversely isotropic, transversely homogeneous microstructural statistical model of articular cartilage. J Biomech. 38:2008 – 2018. Ferguson SJ, Bryant JT, Ganz R, Ito K. 2000a. The influence of the acetabular labrum on hip joint cartilage consolidation: a poroelastic finite element model. J Biomech. 33:953 – 960. Ferguson SJ, Bryant JT, Ganz R, Ito K. 2000b. The acetabular labrum seal: a poroelastic finite element model. Clin Biomech. 15:463 – 468. Ferguson SJ, Bryant JT, Ito K. 2001. The material properties of the bovine acetabular labrum. J Orthop Res. 19:887 – 896. Ganz R, Leunig M, Leunig-Ganz K, Harris WH. 2008. The etiology of osteoarthritis of the hip: an integrated mechanical concept. Clin Orthop Relat Res. 466:264 – 272. Ganz R, Parvizi J, Beck M, Leunig M, No¨tzli H, Siebenrock KA. 2003. Femoroacetabular impingement: a cause for osteoarthritis of the hip. Clin Orthop Relat Res. 417:112 –120. Goreham-Voss CM, McKinley TO, Brown TD. 2007. A finite element exploration of cartilage stress near an articular incongruity during unstable motion. J Biomech. 40:3438 – 3447. Haemer JM, Carter DR, Giori N. 2012. The low permeablity of healthy meniscus and labrum limit articular cartilage consolidation and maintain fluid load support in the knee and hip. J Biomech. 45:1450– 1456. Harris MD, Anderson AE, Henak CR, Ellis BJ, Peters CL, Weiss JA. 2012. Finite element predicitons of cartilage contact stresses in normal human hips. Journal of Orthop Res. 30:1133 – 1139. Hibbitt D, Karlsson B, Sorensen P. 2007. ABAQUS / Theory Manual Version 6.7. Providence (RI): Hibbitt, Karlsson & Sorensen, Inc. Hou JS, Holmes MH, Lai WM, Mow VC. 1989. Boundary conditions at the cartilage – synovial fluid interface for joint lubrication and theoretical verifications. J Biomech Eng-T ASME. 111:78 – 87. Ito K, Minka MA, Leunig M, Werlen S, Ganz R. 2001. Femoroacetabular impingement and the cam-effect. A MRIbased quantitative anatomical study of the femoral head– neck offset. J Bone Joint Surg Br. 82:171– 176. Jorge JP, Simo˜es FM, Pires EB, Rego PA, Tavares DG, Lopes DS, Gaspar A. 2014. Finite element simulations of a hip with femoroacetabular impingement. Comput Methods Biomech Biomed Eng. 17(11):1275 –1284. Krishnan R, Kopacz M, Ateshian GA. 2004. Experimental verification of the role of interstitial fluid pressurisation in cartilage lubrication. J Orthop Res. 22:565– 570. MacCarthy JC, Noble PC, Schuck MR. 2001. The role of labral lesions to development of early degenerative hip disease. Clin Orthop. 393:25 – 37.
Macirowski T, Tepic S, Mann RW. 1994. Cartilage stresses in the human hip joint. J Biomech Eng. 116:10 –18. Manda K, Leif R, Erikson A. 2011. Finite element simulations of focal knee resurfacing implant applied to localized cartilage defects in a sheep model. J Biomech. 44:794 – 801. Mansours JM. 2009. Kinesiology – The mechanics and pathomechanics of human movement. In: Oatis CA, editor. Chapter 6. Biomechanics of cartilage. Philadelphia (PA): Lippincott Williams & Wilkins; p. 69– 83. Mavcˇicˇ B, Pompe B, Antolicˇ V, Daniel V, Iglicˇ V, Kralj-Iglicˇ V. 2002. Mathematical estimation of stress distribution in normal and dysplactic human hips. J Orthop Res. 5:1025–1030. Mow VC, Huiskes R. 2005. Basic orthopaedic biomechanics and mechano-biology. 3rd ed. Philadelphia (PA): Lippincott Williams & Wilkins. Mow CV, Kuei SC, Lai WM, Armstrong CG. 1980. Biphasic creep and stress relaxation of articular cartilage in compression: theory and experiments. J Biomech Eng-T ASME. 102:73 – 84. No¨tzli HP, Wyss TF, Stoecklin CH, Schmid MR, Treiber K, Hodler J. 2002. The contour of the femoral head– neck junction as a predictor for the risk of anterior impingement. J Bone Joint Surg Br. 84:556 – 560. Pawaskar SS, Ingham E, Fisher J, Jin Z. 2011. Fluid load support and contact mechanics of hemiarthroplasty in the natural hip joint. Med Eng Phys. 33:96 –105. Pawaskar SS, Jin ZM, Fisher J. 2007. Modelling of fluid support inside articular cartilage during sliding. Proc Inst Mech Eng J J Eng Tribol. 221:165 – 174. Pritzker KPH, Gay S, Jimenez SA, Ostergaard K, Pelletier JP, Pevell PA, Salter D, Path FRC, van den Berg WB. 2006. Osteoarthritis cartilage histology: grading and staging. Osteoarthritis Cartilage. 1:13– 29. Recnik G, Kralj-Iglicˇ V, Iglicˇ A, Antolicˇ V, Kramberger S, Rigler I, Pompe B, Vengust I. 2009. The role of obesity, biomechanical constitiution of the pelvis and contact joint stress in progression of hip osteoarthritis. Osteoarthritis Cartilage. 7:879 – 882. Roth V, Mow VC. 1980. The intrinsic tensile behaviour of the matrix of bovine articular cartilage and its variation with age. J Bone Joint Surg Am. 62:1002 – 1117. Russel ME, Shivanna KH, Grosland NM, Pedersen DR. 2006. Cartilage contact pressure elevations in dysplastic hips: a chronic overload model. J Orthop Surg Res. 6:1– 11. Tissakht M, Ahmed AM. 1995. Tensile stress–strain characteristics of the human meniscal material. J Biomech. 28:411–422. Wilson W, van Rietbergen B, van Donkelaar CC, Huiskes R. 2003. Pathways of load-induced cartilage damage causing cartilage degeneration in the knee after meniscectomy. J Biomech. 36:845– 851. Wong BL, Won BC, Chun J, Gratz KR, Lotz M, Sah RL. 2008. Biomechanics of cartilage articulation: effects of lubrication and degeneration on shear deformation. Arthritis Rheum. 58:2065 – 2074. Wu JZ, Herzog W. 2000. Finite element simulation of location and time-dependent mechanical behavior of chondrocyt in unconfined compression tests. Ann Biomed Eng. 28:318 – 380. Wu JZ, Herzog W, Epstein M. 1998. Evaluation of the finite element software ABAQUS for biomechanical modelling of biphasic tissues. J Biomech. 31:165 – 169. Yoshida H, Faust A, Wilckens J, Kitagawa M, Fetto J, Chao EYS. 2006. Three-dimensional dynamic hip contact area and pressure distribution during activities of daily living. J Biomech. 39:1996– 2004.
ISSN: 1025-5842 (Print) 1476-8259 (Online) Journal homepage: http://www.tandfonline.com/loi/gcmb20
Hip joint degeneration due to cam impingement: a finite element analysis F.L. Hellwig, J. Tong & J.G. Hussell To cite this article: F.L. Hellwig, J. Tong & J.G. Hussell (2016) Hip joint degeneration due to cam impingement: a finite element analysis, Computer Methods in Biomechanics and Biomedical Engineering, 19:1, 41-48, DOI: 10.1080/10255842.2014.983490 To link to this article: http://dx.doi.org/10.1080/10255842.2014.983490
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Date: 05 November 2015, At: 14:27
Computer Methods in Biomechanics and Biomedical Engineering, 2016 Vol. 19, No. 1, 41–48, http://dx.doi.org/10.1080/10255842.2014.983490
Hip joint degeneration due to cam impingement: a finite element analysis F.L. Hellwiga, J. Tonga* and J.G. Hussellb a
Mechanical Behaviour of Materials Group, School of Engineering, University of Portsmouth, Portsmouth, UK; bOrthopaedics and Trauma Centre, Queen Alexandra Hospital, Portsmouth, UK
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(Received 20 August 2012; accepted 30 October 2014) The goal of this study was to investigate the impact of cam impingement, a biomechanical risk factor, on hip joint degeneration and ultimately coxarthrosis. 3D finite element solid models of a healthy and a pathologic hip were developed based on clinical reports. The biphasic characteristics of cartilaginous tissues were considered to identify localised solid matrix overloading during normal walking and sitting down (SD). Localised femoral intrusion at the anterior-superior pelvic horn was revealed in the pathologic hip during SD, where the radial and meridional solid stresses in the acetabular cartilage and circumferential solid stresses within the acetabular labrum increased by 3.7, 1.5 and 2.7 times, respectively. The increased solid-on-solid stresses, reduction in fluid-load support and associated higher friction during articulation may result in joint wear and other degenerative changes in the hip. Keywords: hip joint; cartilage degeneration; biphasic model; contact pressure; finite element; impingement
Introduction Excessive contact stresses in the hip joint due to obesity (Recnik et al. 2009) or decreased weight-bearing area typical in a dysplastic hip (Mavcˇicˇ et al. 2002; Russel et al. 2006) are thought to result in coxarthrosis. Recent clinical studies hypnotised a link between abnormal anatomical conditions of femoral head – neck junction and early degenerative changes of the hip joint, such as osteochondral defects and labral tear leading ultimately to idiopathic osteoarthritis (Ito et al. 2001). The pathologic condition, named cam-type femoroacetabular impingement, is caused by jamming of a non-spherical extension of the femoral head into the acetabular cavity (Ganz et al. 2003, 2008). No¨tzli et al. (2002) characterised this anatomic abnormality by introducing an alpha angle (Figure 1). This angle is measured between the femoral neck axis and a line extended from the point C, the femoral head centre, to point D, which is the point of deviation from femoral head sphericity. The abnormality might emerge as an anterior bump, spigot or thickened femoral neck, limiting the range of motion (ROM; Ito et al. 2001; Ganz et al. 2003, 2008). Recently, a finite element study (Chegini et al. 2009) of normal and impinged articulation has been presented to examine the joint mechanics during daily activities of normal walking (NW) and sitting down (SD), for which the maximum contact force was reported as 233% and 156% bodyweight, respectively (Bergmann et al. 2001). Chegini et al. (2009) focused on the total stress responses of the cartilaginous tissues, which were represented as isotropic linear elastic materials. The cartilage and labrum, however, are fully fluid saturated porous mediums (Mow
*Corresponding author. Email: [email protected] q 2015 Taylor & Francis
et al. 1980); hence, the load is shared between the interstitial fluid and a solid extracellular matrix (ECM):
stot ¼ ssolid þ Dpl;
ð1Þ
where stot is the total stress of the tissue, ssolid is the stress acting on the ECM and Dpl is the pore pressure of the interstitial fluid (Mow et al. 1980). Early macroscopic changes in articular cartilage were reported to begin with fraying at the superficial zone of the ECM (Pritzker et al. 2006), a sign of joint wear accelerated with increased friction (Mow and Huiskes 2005) and associated with decreased fluid-load support (FLS; Krishnan et al. 2004). Ferguson et al. (2000a) utilised an isotropic poroelastic material model in ABAQUS (Wu et al. 1998) to simulate the sealing function of the acetabular labrum. They correlated alterations in solid-on-solid stress with the changes in FLS for astatic load up to 75% bodyweight held constant for up to 10,000 s. More recently, in order to study the cartilage response to elevated loading during unstable motion, Goreham-Voss et al. (2007) modelled the articulating cartilage layers in knee joints, considering the directional-dependent material properties of the ECM. The aim of this work is to study the stresses in the cartilage in the principal directions of an orthotropic ECM and the FLS during normal and ‘cam’-type impinged articulations under selected physiological loading conditions. To do so, the cartilaginous tissues were modelled as biphasic orthotropic materials, and the results for normal and ‘cam’-type impinged articulation were compared.
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F.L. Hellwig et al. and 4.98 £ 1024 mm4/Ns, respectively (Athanasiou et al. 1994; Ferguson et al. 2001), and they were converted from the biphasic, k, to the poroelastic model, k0 , via:
Anterior Anterior Hump
D
k0 ¼ gk;
C
Medial
Lateral
α
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Posterior
Figure 1. Illustration of a pathologic femur (anterior hump): the deviation of femoral head sphericity is measured by the a angle between the femoral neck axis, passing though the femoral neck and head centre C, and a line extended from the point C– D, the point of deviation from femoral head sphericity (edge of hump).
Numerical methods Geometry and materials A normal hip joint model with a ¼ 408 was developed based on morphologic simplifications and dimensions of Chegini et al. (2009; Figure 2(a)), while a typical camimpinged hip joint with a ¼ 748 was also modelled based on No¨tzli et al. (2002) (Figure 2(b)), using ABQUS CAE. The finite element model, depicted in Figure 2(c), of the normal hip joint (c) consists of acetabular labrum (1), acetabular cartilage (2), femoral cartilage (3), impermeable membrane (4) and femoral head neck section (5), following (Chegini et al. 2009). The pathologic case was identical except the anterior hump at the head–neck junction. Femoral and pelvic bones were considered rigid (Anderson et al. 2008) and the cartilaginous tissues were considered as fully fluid-saturated porous mediums (Mow et al. 1980), taking into account the orthotropic structure of the ECM following Goreham-Voss et al. (2007). A spherical coordinate system was used to allocate the directionaldependent material properties. The permeabilities of the femoral and acetabular cartilage and labrum are 9.15, 8.98
(a)
(b)
ð2Þ
where the volumetric weight of the interstitial water, g, was taken as 9.81 £ 1026 N mm23 (Wu et al. 1998). The solid matrix compressive radial moduli ‘1’of femoral and acetabular cartilage and labrum were averaged to 1.23, 1.18 and 0.157 MPa (Athanasiou et al. 1994; Ferguson et al. 2001), respectively. For cartilage, the tangential moduli in circumferential ‘2’ and meridional ‘3’ directions were extracted from the toe region as 8.5 MPa (Roth and Mow 1980). To permit opening between the acetabular labrum and femoral cartilage under load, a tangential moduli of 26 MPa was chosen in direction ‘2’, close to that of the toe region (20 MPa) reported in Ferguson et al. (2001). According to the ratio of meridional to circumferential strength of the meniscus (Tissakht and Ahmed 1995), the tangential modulus of the labrum in direction ‘3’ was set to 3 MPa, since labrum and meniscus exhibit a similar ultra structure (Ferguson et al. 2001). As shear stresses peak at the cartilage–bone interface (Goreham-Voss et al. 2007), a shear modulus of 2.5 MPa (Wong et al. 2008) was chosen. Shear modulus and volumetric response (n12, n23; Athanasiou et al. 1994; Federico et al. 2005) are given for cartilage in Table 1, similar values were assigned to the labrum due to the lack of data. A summary of all material properties used in the models is given in Table 1. As tissue properties were taken from different sources due to limited availability, a pre-study on the sensitivity to material property variation (^ 50%) was performed. The peak contact force for NW, taken from Bergmann et al. (2001), was applied to the hip joint centre during a 1 s ramp phase for the normal articulation.
Finite element modelling The cartilaginous tissues were meshed with first-order pore pressure brick elements, and the femoral-head neck
(c)
(1) (2) (3) (4) (5)
Figure 2. The solid models for a normal (a) and a pathologic femur with hump (b). The finite element model of the normal hip joint (c) consists of acetabular labrum (1), acetabular cartilage (2), femoral cartilage (3), an impermeable membrane (4) and femoral head neck section (5), following (Chegini et al. 2009).
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Computer Methods in Biomechanics and Biomedical Engineering Table 1.
Material properties used in this study. Biphasic formulation
Tissue Acetabular labrum Acetabular cartilage Femoral cartilage
E1 (MPa)
E2 (MPa)
E3(MPa)
n12
n13
n23
G (MPa)
k0 (mm/s)
F
0.157a 1.23b 1.18b
26a 8.5c 8.5c
3e 8.5c 8.5c
0.04 0.044b 0.046b
0.04 0.044b 0.046b
0.1 0.146d 0.146d
1.5 2.5f 2.5f
4.89b £ 10209 8.52d £ 10209 8.98d £ 10209
4 4 4
a
Ferguson et al. (2001). Athanasiou et al. (1994). c Roth and Mow (1980). d Federico et al. (2005). e Tissakht and Ahmed (1995). f Wong et al. (2008).
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b
section and the impermeable membrane were meshed with quadrilateral bilinear rigid and quadrilateral membrane elements, respectively. Reduced integration was used for deformable elements to reduce the calculation time. The selected grid sizes and number of elements are given in Table 2. Because the pelvis was assumed rigid, the nodes of the acetabular cartilaginous tissue on the bone-cartilage interface were fully constrained. The hip joint forces for NW and SD were taken from Bergmann et al. (2001) and applied incrementally as a point force to the centre of the femoral head. The hip joint force was synchronised with the locomotion data prescribed as rotational movements of the femur. The analysis was performed by first establishing contact between the femur and the acetabulum, then applying an initial load of 33.31% or 55.51% of the peak load for NW and SD, respectively, ramped over 1 s, and finally simulating the locomotion cycle in 60 sequences. The bodyweight was assumed 86 kg. Contact was implemented using the penalty method assuming frictionless finite sliding, due to small coefficient of friction (Pawaskar et al. 2007). Throughout the analysis, the femur was the master and the acetabulum the slave surface. Further, for poroelastic mediums, the contact model should satisfy the biphasic jump condition, where in areas of contact fluid velocity should be continuous or otherwise interstitial fluid should be free-draining (Hou et al. 1989). However, as contact pressure varies, local rehydration occurs (Pawaskar et al. 2007) and fluid efflux and cross-flow are minimal for loadings such as NW and SD (Pawaskar et al. 2011); hence, the biphasic jump
Table 2.
condition was not considered. Because in a biphasic medium load is shared between the fluid and the solid phase (Mow et al. 1980), the total contact pressure is carried by both phases. However, in the ABAQUS model for porous mediums in contact, the fluid phase contribution to the total contact pressure is ignored to ensure point-wise continuity of the pore pressure at opposing sides (Hibbitt et al. 2007). On the other hand, contact pressure between a porous and an impermeable material includes contributions from both solid and fluid phase (Hibbitt et al. 2007). In this work, the total contact pressure was obtained by inserting an impermeable membrane (E ¼ 1 MPa; n ¼ 0.49) between the cartilaginous tissues (Wilson et al. 2003; Manda et al. 2011). This membrane has only a minor effect on contact pressure results (Wilson et al. 2003).
Results Normal articulation For both NW and SD, the weight-bearing area, peak contact and peak pore pressure (P-POR) vary during the course of the simulated activity, and the variation is consistent with that of the applied contact force. The peak contact pressures (P-CPRESSs) of 2.86 and 3.66 MPa are located in the superior region for NW and the posterior region for SD, respectively. However, weight-bearing area changes from superior – anterior to superior – posterior for NW and from superior to posterior during SD; hence, the contact and pore pressure are distributed asymmetrically about the applied load. The timing of the occurrence of
The number of elements and grid sizes used in the FE models of the normal and the abnormal hip joint. Healthy joint
Acetabular cartilage and labrum Femoral cartilage Femoral head – neck section Impermeable membrane
Abnormal joint
Element type
Number
Grid size
Number
Grid size
C3D8RP C3D8RP R3D4 M3D4R
21,460 13,666 9956 4004
0.70 0.90 0.90 1.15
21,460 12,420 8767 4166
0.70 1.00 1.00 0.95
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Table 3.
Peak biphasic response for NW and SD for normal articulation. Biphasic response Location
P-CPRES (MPa)
CAREA (mm )
P-POR (MPa)
S11
Time of cycle (%)
Superior Posterior– medial
2.87 3.58
1798 1574
2.84 3.54
2 0.03 2 0.04
54 46.23
Activity
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NW SD
2
‘Cam’-type impinged articulation The pathologic hip joint revealed localised impingement at the anterior –superior pelvic horn during deep flexion. Hence, pore pressure increased significantly by eight times from 0.42 to 3.76 MPa at the anterior – superior horn due to the intrusion of the femoral hump (Figure 4). Compressive radial stresses, within the acetabular cartilage, increased at the impinged location by 3.7 times from 0.07 to 0.33 MPa. Tensile circumferential and meridional stresses are at peaks within the acetabular labrum and at the cartilage – labral interface, respectively (Figure 4(a),(b)), and increased by a factor of 2.7 and 1.5, respectively. S22 and S33 rose from 0.84 to 2.26 MPa and 0.92 to 1.37 MPa (Figure 4(a),(b)). FLS in the impinged zone dropped to 91.9%, as opposed to 98.4% in the posterior region. Shear stress also doubled to 2.61 MPa on the acetabular surface and surpassed the maximum at the acetabular cartilage – bone interface in the normal articulation.
peak contact and P-POR coincided with the peak contact force. The P-CPRESS, the related contact area (CAREA), the timing and the location of their occurrence for NW and SD are summarised in Table 3. In addition, the P-POR, which corresponded to the location of the P-CPRESS, and the related compressive radial stress (S11) are also summarised in Table 3. For NW, a plot of the results, including the stress components of the solid matrix, is given in Figure 3. The locations of maximum P-CPRESS (2.87 MPa) and maximum P-POR (2.85 MPa) are shown in Figure 3(a), (b), where the distribution patterns are identical. Stresses within the ECM in the three directions are highest on the articulating surface. The highest stress concentrations, radial stress 2 0.05 MPa and meridional stress 0.50 MPa, are found on the edges of the acetabular cartilage, where the pore pressure is minimal (Figure 3(c),(e)). Circumferential stress is the highest in the acetabular labral ring (0.33 MPa, Figure 3(e)). Shear stress, however, peaked at the cartilage – bone interface along the edge of the acetabular cartilage, rising up to 1.35 MPa (Figure 3(f)). The ratio of fluid load-support is 99.3% and 98.4% for NW and SD, respectively, indicating a similarity in fluid –solid interaction in both cases.
(a)
contact pressure (CPRESS)
A
max = 2.87 MPa
(b)
Effect of material property variation The influence of material property variation on some of the key parameters such as femoral head displacement, contact
max = 2.84 MPa
pore pressure (POR)
(c)
fluid velocity (FLVEL)
A-A
A (d) radial stress (S11)
A-A
(e) circumferential stress (S22)
A-A
(f)
meridional stress (S11)
A-A
(g) Tresca stress (S, Tresca)
A-A
Figure 3. The P-CPRESS (MPa) (a) during NW; the related pore pressure (MPa) (b) and the stress components (MPa) in the solid matrix: radial (c), circumferential (d), meridional (e) and Tresca stress (f), section A – A as indicated in (a).
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Computer Methods in Biomechanics and Biomedical Engineering (a) 'cam'-type impinged articulation pore pressure (POR)
radial stress (S11)
circumferential stress (S22)
meridional stress (S33)
Tresca stress (S, Tresca)
radial stress (S11)
circumferential stress (S22)
meridional stress (S33)
Tresca stress (S, Tresca)
(b) normal articulation
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pore pressure (POR)
Figure 4. Pore pressure for normal (a) and CAM-impinged (b) acetabular labral and cartilage exposure during SD: Radial, circumferential, meridional and Tresca stresses (MPa) within the ECM.
pressure and pore pressure is summarised in Figure 5. The highest model sensitivity was recorded for a reduction of 50% in shear modulus, with an increase of 46.18% of femoral head displacement, although the contact pressure and the pore pressure dropped only marginally by 2.89% and 3.68%. Reducing the tangential modulus by 50% resulted in only marginal influence on femoral head displacement (2 4.28%), contact (2.7%) and pore pressure (1.97%), respectively. For all other variations the changes in the key parameters are insignificant , 1%. This gives us confidence that the observed enhanced responses in impinged case are indeed due to the abnormal articulation, and the influence of model parameters are insignificant except the shear modulus.
Discussion The goal of this study was to explore how ‘cam’-type impinged articulation can affect cartilage stresses compared with those under normal articulation; and whether or not cartilage structural integrity may be assessed based on a comparison of the total contact pressure, fluid FLS and generated stress components within the solid ECM in these two conditions. Special emphasis was placed on the consideration of orthotropic properties of the cartilaginous tissue structures. Comparisons with previous work seem to suggest that the locations identified for P-CPRESS during NW and SD are consistent with those published (Yoshida et al. 2006; Chegini et al. 2009). The maximum contact pressures in a normal joint are found to be 2.87 and 3.66 MPa during NW
and SD, compared well with those (2.35 and 3.34 MPa, respectively) reported by Chegini et al. (2009). Other subject-specific models, however, predicted higher PCPRESS, such as 6.6 MPa in Jorge et al. (2014) and 10.78 MPa in Anderson et al. (2008). According to Anderson et al. (2010), cartilage thickness variation accounted for more than doubled P-CPRESS compared with a constant cartilage thickness. Harris et al. (2012) reported a P-CPRESS rage of 7.51 ^ 2.11 MPa for walking, and suggested that the P-CPRESS varies with subjects. Hence, it seems reasonable that our results are within the range of reported data, given the variation of the joint geometries. Experimentally measured P-CPRESSs seem to be higher than predicted by the simplified FE models. According to Afroke et al. (1987), P-CPRESS varied strongly with subjects (4.9 –10.2 MPa). Anderson et al. (2008) reported a rather inhomogeneous contact pressure distribution with a peak of 10 MPa for NW. In contrast to that Brown and Shaw (1982) reported a P-CPRESS of 3.45 MPa, close to our findings. To the best of the authors’ knowledge, the stresses within the solid phase have not been measured in a cadaveric hip joint. Before this study, ECM of articular cartilage was assumed to deform linear elastic isotropically (Macirowski et al. 1994; Ferguson et al. 2000a, 2000b) or hyperelastic isotropically (Haemer et al. 2012), as the material properties of the cartilage were taken from uniaxial compression tests (Mow et al. 1980; Macirowski et al. 1994; Wu and Herzog 2000) where shear cannot be measured. To accommodate the multi-directional cartilage
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Femoral Head Displacement G
Ec
v12/13
v23
k –0.11%
–0.09%
2.5%
High
Low
–0.77%
–3.58%
–0.67%
–17.99%
0.07%
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46.18%
–2.5%
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0.0%
0%
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Deviation
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0.0% 1.95%
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0.0% 2.18%
Deviation
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G 5.0%
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High
Low
Low
High
Figure 5. The variation in the femoral head displacement, contact pressure and pore pressure due to the variation in the material parameters (low: 2 50%; high: þ50%).
deformation while using linear or hyperelastic isotropic properties, either lateral extension was permitted (Macirowski et al. 1994) and applied load reduced (Ferguson et al. 2000a), or the stress-bearing area increased, neglecting the lunate shape of the acetabular cartilage (Haemer et al. 2012). None of these modelling techniques seem to be sufficient to answer the current research questions. Taking the different modelling techniques into account, the high FLS of 99.3% during NW, compares reasonable well with 98.0% and 96.0% obtained by Haemer et al. (2012) and Macirowski et al. (1994) for a ramp load of 1 s, but less favourably with 91.7% (following 1000 static load) reported by Ferguson et al. (2000a).
In the pathologic case, a ¼ 748, the anterior protrusion produced an intrusion into the anterior – superior acetabular cavity during SD, which is consistent with earlier observations (Chegini et al. 2009; Jorge et al. 2014). Although the intrusion of the abnormal protrusion did not yield significant elevated P-CPRESS, the FLS was affected and the stress components of the ECM elevated. Although the FLS was still above 90%, the radial compressive stress (or solid-on-solid) increased significantly (3.7 times) and the ECM was not shielded from elevated solid-on-solid stress (Athesian and Mow 2005). This may explain some of the early macroscopic changes in the cartilage found in clinical observation (Ganz et al.
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Computer Methods in Biomechanics and Biomedical Engineering 2003, 2008; Beck et al. 2005), leading to severe osteoarthritis (Pritzker et al. 2006). Circumferential stress in the labrum was doubled, although this might not be as significant due to the labral design to withstand high exposure in this direction (Ferguson et al. 2001). The elevated stresses are mainly found in the cartilage-ECM, as opposed to those reported by Chegini et al. (2009) in cartilage and labrum simultaneously. The increased meridional stresses within the cartilage at the cartilage–labral interface seems to support the hypothesis that cam impingement results in cartilage defects and subsequent separation of the acetabular cartilage from the labrum (Ganz et al. 2003, 2008; Beck et al. 2005), and not otherwise as discussed (MacCarthy et al. 2001). The emphasis of earlier numerical work (Chegini et al. 2009; Jorge et al. 2014) was on severe abnormal femoral protrusions at high loads. Chegini et al. (2009) predicted, dependent on the morphology of the femoral neck, PCPRESSs are in the range of 3.68–12.84 MPa for ‘cam’type impingement. Jorge et al. (2014) also reported high PCPRESSs up to 16.4 MPa. Due to the varied abnormal conditions presented, the results cannot be readily compared with the current ones. Nonetheless, Ganz et al. (2003) suggested that mild to unrecognisable, in the eye of an observer, abnormal femora are the genesis of impingement. To the best of the authors knowledge, tissue stresses during ‘cam’-type impinged articulation have not been measured. Admittedly, we only considered a typical case of impingement and limited loading scenarios. The geometries considered are simplified versions based on CT images and reported anatomic characteristics, although we believe they contain the essence of a typical case of ‘cam’-type impingement. The finite element mesh was optimised to obtain convergence during ‘cam’-type impinged articulation; the chosen grid size was the optimum from a sensitivity study. Although the normal and the abnormal joints were meshed slight differently, it should not affect the results. The grid-ratios on both acetabular and femoral side were chosen carefully to obtain contact convergence. The slope of the contact-pressure over closure relationship was also adjusted to minimise nodal penetration and contact pressure errors. Further, loading scenarios considered here are limited to NW and SD. In certain sporting activities, such as martial arts, a higher ROM is required hence higher stresses due to impingement may be expected. As in all numerical work, experimental verification would be highly desirable, but it is not yet possible to quantify the load sharing between the fluid and solid phase or measure the strains within the ECM directly. Nonetheless, the key parameters were found to be generally insensitive to the variation of material properties even by variation of ^50% (Figure 5). Because linear behaviour was considered for the fluid (permeability) and the solid phase (elastic orthotropic), the error during a cam-impinged articulation should be similar to that determined during static loading.
47 7
The shear modulus, however, does seem to affect the femoral head displacement, thus cartilage consolidation. Mansours (2009) argued that, during cartilage compaction, elevated shear stresses will occur at the cartilage–bone interface, see Figure 3(f), as the cartilage layers are next to the much stiffer underlying subchondral bone layer, and high lateral contraction is a result of compression of the opposing cartilage layers. Hence, it is important to consider the orthotropic properties of the cartilage ECM when simulating multi-directional deformation of cartilage. Conclusions Results from a biphasic analysis of a healthy and a camimpinged joint show that ‘cam’-type impingement can elevate stresses significantly in the ECM of cartilaginous tissues. During SD, the peak radial and the peak meridional stresses within the cartilage and the circumferential stress within the acetabular labrum increased by 3.7, 2.7 and 1.5 times, respectively, compared with those in the normal joint. P-POR also increased at the anterior-superior ‘cam’-type impinged zone. The increased solid-on-solid stresses and associated decrease in FLS would result in higher friction during articulation and hence joint wear. In summary, ‘cam’-type impingement is a risk factor that could increase cartilaginous tissue strains leading to joint wear and other degenerative changes in the hip, consistent with clinical findings.
Acknowledgements The authors gratefully acknowledge access to the Sciama High Performance Computer (HPC) Cluster, which was supported by the ICG, SEPNet and the University of Portsmouth and support by Sciama technician Gary Burton.
Conflict of interest disclosure statement No potential conflict of interest was reported by the authors.
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