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The impact of subject weight and activity level on over-inserted cemented acetabular cups after Total HIP Arthroplasty (THA)
, Jose F. Del-Valle-Mojicaa, Israel Miguel-Andresb, Colin G. Baileya, Tim N. Boardc
⁎Corresponding author: Teresa Alonso-Rasgado. t.alonso@qmul.ac.uk
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Received: ,
Accepted: ,
This article was originally published by Reed Elsevier India Pvt. Ltd. and was migrated to Scientific Scholar after the change of Publisher.
Abstract
Abstract
In cemented Total Hip Arthroplasty (THA), several risk factors have been identified with the failure of the acetabular component, including body mass index (BMI), exercise, femoral head size, cup placement, and cement mantle integrity. Elevated BMI and larger femoral heads increase bone and cement stresses, accelerating wear and predisposing to aseptic loosening. This study evaluates the effect of over-inserted acetabular cups on cement mantle interfaces and pelvic bone stresses using two femoral head sizes (28 mm and 36 mm), four body weights (normal, overweight, obese, and morbidly obese), and three activities (one-leg stand, stair descent, and stumbling). Results show that von Mises stresses rise with increasing body weight, activity intensity, and femoral head size, with the highest values observed in the superior periacetabular region of the pelvic bone for the morbidly obese subject during stumbling with the 36 mm head. Stresses at the bone-cement interface exceeded those at the cement-cup interface, particularly in the superior quadrant of over-inserted cups. Our findings suggest that morbidly obese subjects are at a higher risk of aseptic loosening due to the stresses induced in the bone-cement interface during physical activities, resulting in higher peak hip reaction forces. This risk is increased in the case of over-insertion of the acetabular cup, leading to a thinner cement layer. These findings highlight the combined influence of implant design, patient characteristics, and surgical technique on long-term THA performance.
Keywords
Finite element
Total hip arthroplasty
Femoral head
Cement mantle
1 Introduction
Risk factors for failure of the acetabular component in cemented Total Hip Arthroplasty (THA) include exercise, obesity, femoral head size, acetabular cup placement, and cement mantle thickness 1–7. Studies have suggested that low-impact sports activities and increased daily activity levels can loosen the acetabular cup.1,2 However, other studies argue that a sedentary lifestyle poses a greater risk of loosening of the acetabular cup, which is also associated with a high Body Mass Index (BMI).8,9
BMI is associated with a higher failure rate of THA 3–5. Wagner et al. found that although increasing BMI was associated with an overall increasing risk of implant revision, it was not associated with a higher risk of revision for aseptic loosening and mechanical failure of the implant.5 Kumar et al. reported that increased body weight increases bone and cement stress, affecting the performance of the acetabular component in obese patients.10,11 Toh et al. determined that higher body weight correlates with higher wear rates, which may lead to failure.12 Eletricwala et al. reported that obesity increased the relative risk of early revision THA due to aseptic loosening.4
Larger femoral heads in THA are being used more frequently due to their potential to reduce dislocation and improve joint stability 9,10,13–17. However, data suggest that larger femoral heads may be related to higher rates of joint failure.6 Miyamoto et al.18 reported that femoral head size was significantly associated with the risk of development of radiolucent lines (an indicator of subsequent loosening of the acetabular component) at the bone-cement interface one-year post-operative. Annanto et al.19 indicated that von Mises and contact stress of the cement layer increased with head size by up to 10 %.
Cement mantle integrity is essential for long-term fixation in cemented implants,14 and key to this is proper acetabular placement, cement mantle thickness, and uniformity.7 Concentric placement of the acetabular cup in a uniform cement mantle of 3 mm thickness is often cited as the ideal to reduce the risk of elevated stress and crack formation and growth in the cement layer.7 However, over-insertion of the acetabular component, resulting in a non-uniform, thinner than ideal cement layer, is commonplace in practice,7,20 potentially resulting in mechanical overload at the cement–bone interface.21
It has been found that the cement mantle stress is influenced by femoral head implant size, acetabulum morphology, bone quality,22,23 and acetabular cup outer diameter.24 Del-Valle-Mojica et al.13 showed that von Mises stress at the bone‐cement interface increased with femoral head size, subject weight, and activity level, with the highest stresses in the cement mantle occurring during stumbling for an obese subject. However, those critical scenarios have not been evaluated with an over-inserted acetabular cup (O-IAC).
This work describes an investigation into the effect of O-IACs on the acetabular cement mantle interfaces and pelvic bone stress for two femoral head sizes; 28 mm and 36 mm considering three daily activities (one leg stand, descending stairs and stumbling) and a range of subject weights (for a 182 cm tall subject); normal weight (82 kg, BMI 24), overweight (95 kg, BMI 29), obese (111 kg, BMI 34) and morbidly obese (143 kg, BMI 43). In addition, a comparison for over-inserted and non-over-inserted13 cups was carried out for the two femoral head sizes, the critical activity, stumbling, and the range of subject body weights. This work is an extension of a recently published study.25
2 Materials and methods
Institutional Review Board (IRB)/ethical approval was not required as the work was undertaken using synthetic hemipelves. Twenty-four 3D Finite Element (FE) models (2 femoral head diameters × 3 activities × 4 wt) were developed to simulate the conditions analysed in this study (Fig. 1). The models were generated from CT data of two synthetic hemipelves implanted with over-inserted cemented cups (O-IAC) accommodating 28 mm and 36 mm femoral heads, Fig. 2a–b. Cement mantle thickness ranged from 1.04 to 3.71 mm for the 28 mm head, and 1.16 to 5.68 mm for the 36 mm head, Fig. 2a.
Experimentally measured surface strains from a synthetic hemipelvis (Sawbones®) were compared with model predictions to validate the FE simulations. Once validated, the models were used to evaluate stresses in key regions: the superior periacetabular area (Z1), the anterior periacetabular region (Z2), both primary zones of force transfer to the hip joint,26 the bone-cement mantle interface (Z3), and the cement-cup interface (Z4), Fig. 2c–d. Finally, stresses for O-IAC configurations were compared with published data for non-over-inserted cups (NON-O-IAC) reported by Del-Valle-Mojica et al..13
2.1 Finite Element model
CT data from Sawbones hemipelves implanted with over-inserted cemented all-polyethylene cups for 28 mm and 36 mm femoral heads were used to generate the 3D FE models. A total of 475 CT slices were segmented to obtain solid geometries of cortical bone, cancellous bone, and the cement mantle (Fig. 2a–c). These geometries were imported into Abaqus 6.13 (Dassault Systemes, RI, USA) for assembly and meshing using C3D10 tetrahedral elements, while the femoral head and acetabular cup were meshed with C3D20R hexahedral elements, Fig. 2c.
3 Materials
Material properties from the Sawbones® manufacturer were used in the FE models, along with published values for acrylic cement, the UHMWPE acetabular cup, and the steel femoral head (Table 1) 27,28. All materials were modelled as linear, elastic, and isotropic.
| Material | Young's Modulus (MPa) | Poisson's ratio (v) |
| Trabecular bone | 155 | 0.3 |
| Cortical bone | 16,000 | 0.3 |
| UHMWPE (acetabular cup) | 800 | 0.4 |
| Acrylic bone cement (cement mantle) | 2000 | 0.3 |
| Steel (femoral head) | 207,000 | 0.3 |
3.1 Boundary conditions
All rotations and displacements were constrained at the pubic symphysis and sacroiliac joint,21Fig. 2c. Cement–cup and cement-bone interfaces were modelled as bonded,24 while cup-head contact was frictionless.26 The reaction force magnitude (FR) and loading angle for each activity and body weight are summarized in Table 2.
| Activities | Load Applied at Hip Joint (% of Body Weight, BW) | Loading Angle |
| One Leg Stance | 231 % BW | 6.5o |
| Descending stairs | 260 % BW | 10.8o |
| Critical Case: Stumbling | 1.25x(870 % BW) | 13.1 |
3.2 Mesh sensitivity analysis
A mesh sensitivity analysis comparing three mesh densities was performed to ensure FE model accuracy. Bone and cement were meshed with C3D10 elements, and femoral heads (28 mm and 36 mm) and cups with C3D20R elements. The medium mesh was selected, as stress differences with the fine mesh were below 1 % in target regions. Mesh details are provided in Table 3.
| Acetabular cup Internal diameter | Region | Mesh Size | Elements | Nodes | Mises Stress (MPa) | % Error |
| 28 mm | Z1 | Coarse | 9649 | 17,589 | 3.35712 | – |
| Medium | 10,805 | 19,391 | 3.349592 | 0.2a | ||
| Fine | 17,107 | 29,222 | 3.3559 | 0.2b | ||
| Z2 | Coarse | 2800 | 5563 | 3.03812 | – | |
| Medium | 4150 | 7708 | 3.036487 | 0.1c | ||
| Fine | 8312 | 14,368 | 3.041005 | 0.2d | ||
| 36 mm | Z1 | Coarse | 2144 | 4474 | 2.918 | – |
| Medium | 2351 | 4880 | 2.922 | 0.1a | ||
| Fine | 3290 | 6530 | 2.917 | 0.2b | ||
| Z2 | Coarse | 1160 | 2501 | 2.534 | – | |
| Medium | 1501 | 3156 | 2.556 | 0.9c | ||
| Fine | 2268 | 4612 | 2.553 | 0.1d |
3.3 Experimental work: FE model validation
The twenty-four 3D FE models were validated using a hemipelvis manufactured from synthetic bone. The hemipelvis was implanted by an orthopaedic surgeon with an over-inserted cemented, 36 mm internal diameter, 50 mm outer diameter, acetabular cup. A load corresponding to a subject with a body weight of 82 kg undertaking a one-leg stand was used for the experimentation. Digital Image Correlation (DIC) was used to measure vertical strains in two areas on the surface of the hemipelvis for validation purposes, Fig. 3a–d.
3.4 Experimental setup
A 36 mm stainless-steel ball mounted on an Instron machine (5900R84 MA USA) applied the load to the hemipelvis and acetabular cup. The hemipelvis was fixed at the sacroiliac joint and supported at the pubic symphysis. Two 5-megapixel cameras with 12 mm lenses captured vertical surface strains, which were processed using dedicated software, as shown in Fig. 3a–c.
3.5 Experimental procedure
A speckle pattern was applied to the hemipelvis by spraying black paint over a white base to enable DIC tracking of surface deformation. The painted specimen was then secured in the Instron machine after DIC calibration. During testing, the cameras were positioned 280 mm from the hemipelvis to record the strain data, Fig. 3b.
3.6 Data processing
After measurements were taken, image data were processed in Aramis V6.1 (GOM Ltd, Athlone, Ireland). Linear vertical strains were computed using a 20 × 20-pixels quadrangular mask on the hemipelvis surface, and a Gaussian filter was applied to reduce noise. The software's statistics tool provided average surface strains in the target regions, which were then compared with FE model predictions.
4 Results
Fig. 4 presents the average von Mises stresses (AvMS) in trabecular and cortical bone for regions Z1 and Z2. For the cement-cup and bone-cement interfaces, the AvMS distribution is shown across four quadrants: anterosuperior (Q1), posterosuperior (Q2), anteroinferior (Q3), and posteroinferior (Q4), as illustrated in Figs. 5 and 6.
4.1 FE model validation: vertical surface strains in zones EW1 and EW2
Fig. 3d shows average strains obtained from the predicted and experimental results for the two areas considered (EW1 and EW2). These two regions were selected as they are found in the region where most of the force is transmitted to the pelvis. The difference obtained from comparing the numerical and experimental results was 3 %, so it was accepted as validation of the FE numerical model. The value of predicted strain is 367 με compared to experimental work 360 με.
4.2 Periacetabular bone in zones: Z1 and Z2
Fig. 4 shows the predicted average von Mises stresses (AvMS) in the periacetabular bone for both femoral head sizes across all simulated scenarios. The results indicate that body weight (BMI) influences AvMS in both cortical and trabecular bone in the two regions analysed. Cortical bone consistently exhibited higher stresses in zone 1 (Z1) than in zone 2 (Z2). In contrast, trabecular bone showed markedly lower stresses (0.2–2.1 MPa) compared with cortical bone (2.5–21.2 MPa) across all BMI levels, activities, femoral head sizes, and regions. Cortical (124 MPa) and trabecular (6.7 MPa) yield strengths were used as a failure criterion.
In trabecular bone, AvMS generally increased with both activity intensity and BMI for both femoral head sizes. Similarly, cortical bone stresses rose with BMI for each activity in both zones. Stumbling produced the highest cortical stresses: for the 28 mm head, stresses in morbidly obese subjects were 3.5 times higher in Z1 and 3.3 times higher in Z2 compared with descending stairs; for the 36 mm head, the corresponding increases were 3.6 and 3.3 times. For one-leg stand and descending stairs, cortical stress levels were comparable between head sizes, whereas stumbling produced up to 0.7 MPa higher stresses for the 36 mm head.
4.3 Interface: bone-cement
Fig. 5 compares the average von Mises stress (AvMS) in the superior (Q1, Q2) and inferior (Q3, Q4) quadrants of the bone-cement interface for both femoral head sizes across all scenarios. AvMS was influenced by body weight in every quadrant, activity, and head size. Stress magnitudes followed a consistent pattern: Q1 showed the highest values (2.6–19.2 MPa), followed by Q2 (1.8–14.8 MPa), Q3 (0.9–9.5 MPa), and Q4 (0.7–6.3 MPa). In all quadrants, AvMS increased with activity intensity and BMI.
For the 28 mm head, Q1 stresses in morbidly obese subjects were 75 %, 70 %, and 70 % higher than in normal-weight subjects during one-leg stance, descending stairs, and stumbling, respectively. In morbidly obese subjects, Q1 stresses for stumbling were 3.1 times higher than for descending stairs and 3.5 times higher than for one-leg stance. For the 36 mm head, Q1 stresses were 117 %, 69 %, and 68 % higher in morbidly obese subjects for the same activities, with stumbling producing 3.1 times and 3.4 times higher Q1 stresses than descending stairs and one-leg stance, respectively. Overall, the 36 mm head produced higher stresses than the 28 mm femoral head.
4.4 Interface: cement-cup
Fig. 6 compares the AvMS across the four quadrants (Q1-Q4) of the cement-cup interface for both femoral head sizes and all scenarios. Stress magnitudes followed a consistent pattern: Q1 showed the highest values (2.7–16.3 MPa), followed by Q2 (2.1–10.9 MPa), Q3 (0.8–5.2 MPa), and Q4 (0.8–6.4 MPa). In all quadrants, AvMS increased with activity intensity and BMI for both head sizes.
For the 28 mm head, Q1 stresses in morbidly obese subjects were 14 %, 63 %, and 82 % higher than in normal-weight subjects during one-leg stance, descending stairs, and stumbling, respectively. In morbidly obese subjects, Q1 stresses during stumbling were 3.5 times higher than during descending stairs and 3.8 times higher than during one-leg stance. For the 36 mm head, Q1 stresses were 33 %, 63 %, and 80 % higher for morbidly obese subjects across the same activities. In this group, stumbling produced Q1 stresses 3.4 times greater than descending stairs and 3.7 times greater than one-leg stance. Overall, the 36 mm head generated higher stresses than the 28 mm head.
A comparison was undertaken of AvMS on the bone-cement and cement-cup interfaces between O-IAC and NON-O-IAC femoral heads for quadrants Q1, Q2, Q3, and Q4 for the normal weight, overweight, obese, and morbidly obese subjects and activity (stumbling) for femoral head sizes of 28 and 36 mm.
Fig. 7 shows that AvMS were consistently highest in quadrant Q1 for both femoral head sizes and for both the bone-cement and cement-cup interfaces. Overall, stresses in Q1 were higher in the over-inserted than in the non-over-inserted condition, but lower in Q2. In Q1, over-inserted heads exhibited up to 35 % higher stresses at the bone-cement interface and up to 10 % higher stresses at the cement-cup interface. AvMS increased with BMI for all configurations.
For the 28 mm and 36 mm femoral heads, stresses in quadrants Q1 and Q2 were 1.6–1.8 times higher in morbidly obese subjects compared with normal-weight subjects for both over-inserted and non-over-inserted cups. At the bone-cement and cement-cup interfaces, the 36 mm over-inserted head produced higher Q1 stresses than the 28 mm head but lower stresses in Q2; in the non-over-inserted case, the 36 mm head showed higher stresses than the 28 mm head in both quadrants. In the inferior quadrants (Q3 and Q4), AvMS were generally higher for over-inserted than for non-over-inserted heads for both sizes. For both interfaces, stresses in Q3 and Q4 were higher for the 36 mm than for the 28 mm head in the over-inserted and non-over-inserted cases.
5 Discussion
Aseptic loosening remains a primary cause of construct failure and revision in cemented total hip arthroplasty (THA).13,21,26 Risk factors for acetabular component failure include physical activity, obesity, femoral head size, cup positioning, and cement mantle thickness 1–7. Mechanical and biological mechanisms contribute to loosening.21,26,29 Mechanically, insufficient initial fixation and postoperative degradation can compromise stability.29 In cemented THA, the bone-cement and cement-cup interfaces are essential to construct integrity, with the bone-cement interface being the weakest and most susceptible to failure.30 Elevated stresses at this interface under physiological loads may induce cement mantle damage and lead to mechanical failure.
This study investigates the influence of over-inserted acetabular cups on stresses within the cement mantle interfaces and pelvic bone for two femoral head sizes, across three daily activities, and a range of body weights, and compares these outcomes with non-over-inserted cups.13
The highest stresses in periacetabular bone were in cortical bone, and these were around ten times those of the corresponding values in the trabecular bone. Stress increased in cortical bone as subject weight and activity level increased, and tended to be higher for the 36 mm femoral case than for the 28 mm head. The highest overall stress was found in the superior periacetabular region in the morbidly obese subject, during stumbling for the 36 mm femoral head case. This value was 3.6 times higher than for the same subject during the descending stairs activity and 7.3 times higher than the normal weight subject when undertaking a one-leg stance. We found that the highest stresses are in the superior periacetabular region, which is consistent with previous research that determined that this is the primary area of load transfer for the hip joint force.31 Stresses in both cortical and trabecular bone were well below their respective yield strength, indicating that mechanical overload of the periarticular bone would not be a factor affecting the structural integrity of the prosthesis.
Similar patterns of behaviour were observed for stresses at the bone-cement and cement-cup interfaces with respect to body weight, activity level, and femoral head size. In both interfaces, the highest stresses were consistently located in the antero-superior quadrant. As with the periacetabular bone, stresses increased with rising body weight and activity level, and were generally higher for the 36 mm femoral head than for the 28 mm head. In the morbidly obese subject, stumbling produced bone-cement interface stresses up to 3.5 times greater than those during one-leg stance and stair descent; at the cement-cup interface, stresses were up to 3.8 times higher.
For a given activity, increasing subject weight resulted in stress elevations of up to 117 % at the bone-cement interface and 82 % at the cement-cup interface in the antero-superior quadrant. Notably, bone-cement interface stresses exceeded cement-cup values by up to 14.8 %. Comparisons with non-over-inserted femoral heads13 showed that over-insertion markedly increased stresses during stumbling, by up to 35 % at the bone-cement interface and 10 % at the cement-cup interface.
Predicted stresses in the antero-superior quadrant approached the bone-cement yield strength (21.7 MPa32) for the over-inserted 36-mm head in the morbidly obese subject during stumbling. Moreover, stresses exceeded the cement endurance limit (8.1 MPa33) for all BMI groups, including normal weight. Such elevated stresses can promote crack initiation and propagation within the cement, contributing to mantle failure.21 Fatigue-induced debonding at the bone-cement interface has been identified as the primary failure mechanism,34 and while individual load cycles may be insufficient to cause immediate failure, repeated loading can precipitate failure in the presence of initial defects.34
A number of assumptions and limitations apply to this study. All materials in the models were assumed to behave as linear elastic and isotropic, and the bone-cement and cement-cup interfaces were treated as perfectly bonded. Moreover, loading was applied statically rather than cyclically. Nonetheless, the model was validated against experimentally measured surface strains obtained from an instrumented hemipelvis fabricated from synthetic bone, providing a strong level of confidence in the numerical predictions.
In summary, stresses in the periacetabular bone and at the bone-cement and cement-cup interfaces increased with subject weight, activity level, and femoral head size. Morbidly obese individuals showed a heightened risk of aseptic loosening due to elevated bone-cement interface stresses, further exacerbated by larger femoral heads. Over-insertion of the acetabular cup also increased the risk by reducing cement mantle thickness. Importantly, our models did not represent a worst-case scenario, as clinical mantles may be thinner and generate even higher stresses.
Author contribution
All the authors contributed equally to this paper.
Ethical statement
Institutional Review Board (IRB)/ethical approval was not required as the work was undertaken using synthetic hemipelves.
Patient-guardian consent
Not applicable.
Credit author statement
Teresa Alonso-Rasgado: Conceptualization, Formal analysis, Methodology, Supervision, Validation, Visualization, Writing - original draft, Writing - review & editing.
Tim N. Board: Conceptualization, Formal analysis, Methodology, Supervision, Validation, Visualization, Writing - original draft, Writing - review & editing.
Colin G. Bailey: Conceptualization, Formal analysis, Investigation, Methodology, Supervision, Validation, Visualization, Writing - original draft, Writing - review & editing.
Israel Miguel-Andres: Conceptualization, Formal analysis, Methodology, Validation, Visualization, Writing - original draft, Writing - review & editing.
Jose F. Del-Valle-Mojica: Conceptualization, Formal analysis, Investigation, Methodology, Validation, Visualization, Writing - original draft, Writing - review & editing.
Funding statement
There were no funding sources for this work.
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