Introduction. Femoral neck impingement occurs clinically in total hip replacements (THR) when the acetabular liner articulates against the neck of a femoral stem prosthesis. This may occur in vivo due to factors such as prostheses design, patient anatomical variation, and/or surgical malpositioning, and may be linked to joint instability, unexplained pain, and dislocation. The Standard Test Method for Impingement of Acetabular Prostheses, ASTM F2582 −14, may be used to evaluate acetabular component fatigue and deformation under repeated impingement conditions. It is worth noting that while femoral neck impingement is a clinical observation, relative motions and loading conditions used in ASTM F2582-14 do not replicate in vivo mechanisms. As written, ASTM F2582-14 covers failure mechanism assessment for acetabular liners of multiple designs, materials, and sizes. This study investigates differences observed in the implied and executed kinematics described in ASTM F2582-14 using a Prosim electromechanical hip simulator (Simulation Solutions, Stockport, Greater Manchester) and an AMTI hydraulic 12-station hip simulator (AMTI, Watertown, MA). Method. Neck impingement testing per ASTM F2582-14 was carried out on four groups of artificially aged acetabular liners (per ASTM F2003-15) made from GUR 1020 UHMWPE which was re-melted and cross-linked at 7.5 Mrad. Group A (n=3) and B (n=3) consisted of 28mm diameter femoral heads articulating on 28mm ID × 44mm OD acetabular liners. Group C (n=3) and D (n=3) consisted of 40mm diameter femoral heads articulating on lipped 40mm ID × 56mm OD 10° face changing acetabular liners. All acetabular liners were tested in production equivalent shell-fixtures mounted at 0° initial inclination angle. Femoral stems were potted in resin to fit respective simulator test fixtures. Testing was conducted in bovine serum diluted to 18mg/mL protein content supplemented with sodium azide and EDTA. Groups A and C were tested on a Prosim; Groups B and D were tested on an AMTI. Physical examination and coordination measurement machine (CMM) analyses were conducted on all liners pre-test and at 0.2 million cycle intervals to monitor possible failure mechanisms. Testing was conducted for 1.0 million cycles or until failure. An Abaqus/Explicit model was created to investigate relative motions and contact areas resulting from initial impingement kinematics for each test group. Results. Effects of kinematic differences in the execution of ASTM F2582-14 were observed in the four groups based on simulator type (Figure 1) and liner design. The Abaqus/Explicit FEA model revealed notable differences in relative motions and contact points (Figure 2) between specimen components i.e. acetabular liner, femoral head, and femoral stem throughout range of motion. Acetabular liner angular change within shell-fixtures, rim deformation, crack propagation, and metal-on-metal contact between acetabular shell-fixtures and femoral stems were observed as potential failure mechanisms (Figure 3) throughout testing. These mechanisms varied in severity by group due to differing contact stresses and simulator constraints. Significance. Investigating failure mechanisms caused by altered kinematics of in-vitro neck impingement testing, due to influences of simulator type and acetabular liner design, may aid understanding of failure mechanisms involved when assessing complaints/retrievals and influence future prosthetic designs. For any figures or tables, please contact the authors directly

Introduction. Instability continues to be the number one reason for revision in primary total hip arthroplasty (THA). Commonly, impingement precedes dislocation, inducing a levering out the prosthetic head from the liner. Impingement can be prosthetic, bony or soft tissue, depending on component positioning and anatomy. The aim of this virtual study was to investigate whether bony or prosthetic impingement occurred first in well positioned THAs, with the hip placed in deep flexion and hyperextension. Methods. Twenty-three patients requiring THA were planned for a TriFit/Trinity ceramic-on-poly cementless construct using the OPS. TM. dynamic planning software (Corin, UK). The cups were sized to best fit the anatomy, medialised to sit on the acetabular fossa and orientated at 45° inclination and 25° anteversion when standing. Femoral components and head lengths were then positioned to reproduce the native anteversion and match the contralateral leg length and offset. The planned constructs were flexed and internally rotated until anterior impingement occurred in deep flexion [Fig. 1]. The type (bony or prosthetic), and location, of impingement was then recorded. Similarly, the hips were extended and externally rotated until posterior impingement occurred, and the type and location of impingement recorded [Fig. 2]. Patients with minimal pre-operative osteophyte were selected as a best-case scenario for bony impingement. Results. 6/23 (26%) patients were planned with only a 32mm articulation (<50mm cup size), with the remaining 17 patients all planned with both 32mm and 36mm articulations (≥50mm cup size). Anterior impingement was 26% prosthetic and 74% bony with the 32mm articulations, and 100% bony with the 36mm articulations. Bony impingement in deep flexion was exclusively anterior neck on anterior inferior iliac spine. Posterior impingement was 57% prosthetic and 43% bony with the 32mm articulations, and 41% prosthetic and 59% bony with the 36mm articulations. Bony impingement in hyperextension was exclusively lesser trochanter (LT) on ischium. Of the patients planned with both 32mm and 36mm articulations, there was a 14% increase in prosthetic impingement when a 32mm head was planned (35% and 21% respectively). Discussion. Impingement in THA usually precedes dislocation and should be avoided with appropriate component positioning. We found that in hyperextension, prosthetic and bony impingement were equally common. In deep flexion, impingement was almost exclusively bony. Further studies should investigate the effects of stem version, cup orientation, liner design, cup depth, native offset and retained osteophytes on the type of impingement in THA. For any figures or tables, please contact the authors directly

Introduction. Lipped liners have the potential to decrease the rate of revision for instability after total hip replacement since they increase the jumping distance in the direction of the lip. However, the elevated lip also may reduce the Range of Motion and may lead to early impingement of the femoral stem on the liner. It is unclear whether the use of a lipped liner has an impact on the level of lever-out moments or the contact stresses. Therefore, the aim of the current study was to calculate these values for lipped liners and compare these results to a conventional liner geometry. Materials and Methods. 3D Finite Element studies were conducted comparing a ceramic lipped liner prototype and a ceramic conventional liner both made from BIOLOX. ®. delta. The bearing diameter was 36 mm. To apply loading, a test taper made of titanium alloy was bonded to a femoral head, also made from BIOLOX. ®. delta. Titanium was modeled with a bilinear isotropic hardening law. For the bearing contact a coefficient of friction of both 0.09 or 0.3 was assumed to model a well and poorly lubricated system. Frictionless contact was modeled between taper and liner. Pre-load was varied between 500 N and 1500 N and applied along the taper axis. While keeping pre-load constant, lever-out force was applied perpendicular to the taper axis until subluxation occurred. Liners were fixed at the taper region. Lever-out moment, equivalent plastic strain and von Mises stress of the taper, bearing contact area and contact area between taper and liner was evaluated. Results. With increasing pre-load, larger lever-out moment, equivalent plastic strain, contact area between taper and liner and bearing contact area was found for both liner designs. However, von Mises stresses were nearly constant but slightly exceeded yield strength of titanium. For all evaluated parameters almost no differences were found between the liner designs. Lever-out moments were comparable for both designs ranging from 4.5–10.5 Nm for the lipped liner and 4.4–10.2 Nm for the conventional liner. The increase of the coefficient of friction strongly affected lever-out moments, equivalent plastic strain and contact area between taper and liner. The other parameters were not affected by varying the coefficient of friction. Discussion. This study did not find significant differences in the lever-out behavior of the lipped acetabular liner compared to the conventional liner design. The inner geometry of the lipped liner is comparable to the conventional liner inner geometry. Therefore, contact area showed no significant differences and contact mechanics are identical in the current setup leading to similar results of both liner designs. For both designs small plastic deformations in the contact point of the taper were found at the contact region between liner and taper. However, the investigated mechanical parameters did not differ between the two investigated liner types. For any figures or tables, please contact authors directly

Introduction. Dual-mobility (DM) liners provide increased range of motion and stability. However, large head diameters have been associated with anterior hip pain due to impingement with surrounding soft-tissues, particularly the iliopsoas. Further, during hip extension the liner can get trapped due to anterior soft-tissue impingement that resists rotation being imparted to the liner from posterior stem-liner contact. Over time this can cause liner rim damage, leading to intra-prosthetic dislocation of the small diameter inner head. To address this, an anatomically contoured dual mobility (ACDM) liner was designed to reduce the volume of the liner below the equator that can interact with soft-tissues (Fig. 1). In this study, we utilized finite element analysis to evaluate tendon-liner contact pressure and tendon stresses with ACDM and conventional designs during hip extension, wherein the posterior edge of liner is in contact with the stem while the anterior edge is exposed to the soft-tissue. Methods. The average uniaxial stiffness (350 N/mm), and average dimensions (width × thickness = 14mm × 4mm) of 10 cadaver psoas tendon samples were determined in a separate study. The iliopsoas tendon was modelled as a Yeoh hyper-elastic material, and the material constants were tuned to match the experimental uniaxial test data. Cadaver specific FEA models were created for 5 specimens (10 hips) using computed tomography (CT) scans. The implant components were modeled as being rigid relative to the iliopsoas tendon. The iliopsoas tendon was modelled as extending from its insertion point on the lesser trochanter to the psoas notch on the pelvis for hip flexion angles of −15°, 0°, 15° and 30°. Appropriately sized DM components were implanted virtually for each specimen. Once placed in its proper position, the liner was rotated about the flexion axis until it contacted the stem posteriorly to represent its orientation during hip extension (Fig. 2). A 500N tensile load was applied to the iliopsoas tendon and the average/max stresses within the tendon, and average/max contact pressures between the tendon and liner were measured. Results. At all hip flexion angles from −15° to 30°, the tendon-liner contact pressure and tendon stresses were lower with the ACDM liners compared to the conventional liner. Contact pressure and tendon stress decreased for both liner designs with increasing hip flexion angle. At −15° flexion angle, the average contact pressure was 42.3% lower (0.36Mpa), and the maximum contact pressure was 45.1% (8.5Mpa lower), with the ACDM compared to conventional liner design. Similarly, at −15° flexion angle the average vonMises pressure in the tendon was 32.5% lower (14.8Mpa), and the maximum vonMises stress in the tendon was 55.7% (159Mpa lower) with the ACDM design. (Fig 3). Discussion. This study utilized cadaver specific FEA models to evaluate interaction between the iliopsoas tendon and conventional and ACDM liners during hip extension. The results showed a notable reduction in contact pressure and tendon stress resulting from reduced volume and more soft-tissue friendly profile of the ACDM design. Thus, the ACDM design may be able to reduce undesirable soft-tissue interaction with dual mobility liners

Introduction. Impingement of total hip arthroplasties (THAs) has been reported to cause rim damage of polyethylene liners, and in some instances has led to dislocation and/or mechanical failure of liner locking mechanisms in modular designs. Elevated rim liners are used to improve stability and reduce the risk of dislocation, however they restrict the possible range of motion of the joint, and retrieval studies have found impingement related damage on lipped liners. The aim of this study was to develop a tool for assessing the occurrence of impingement under different activities, and use it to evaluate the effects a lipped liner and position of the lip has on the impingement-free range of motion. MATERIALS & METHOD. A geometrical model incorporated a hemi-pelvis and femur geometries of one individual with a THA (DePuy Pinnacle® acetabular cup with neutral and lipped liners; size 12 Corail® stem with 32mm diameter head) was created in SOLIDWORKS (Dassault Systèmes). Joint motions were taken from kinematic data of activities of daily living that were associated with dislocation of THA, such as stooping to pick an object off the floor and rolling over. The femoral component was positioned to conform within the geometry of the femur, and the acetabular component was orientated in a clinically acceptable position (45° inclination and 20° anteversion). Variation in orientation of the apex of the lip was investigated by rotating about the acetabular axes from the superior (0°) in increments of 45° (0°−315°), and compared to a neutral liner. Results. When a lipped liner was used, implant (neck on acetabular rim) impingement was found to occur when performing sit-to-stand from a normal seat, leg cross and pivot, whereas no impingement occurred with a neutral liner. The presence and position of the lip reduced the impingement-free range of motion, compared to the neutral liner. Impingement occurred when the lip was positioned superiorly and anteriorly, when performing most of the activities that were prone to posterior dislocation, and posteriorly, posterior-superiorly and posterior-inferiorly when performing activities prone to anterior dislocation. During sit-to-stand from a normal seat no impingement occurred when a lipped or neutral liner was used. Bone impingement was observed when the performing the roll activity with both lipped and neutral liners. DISCUSSION. Impingement was observed more with lipped liners compared to neutral liners, this agrees with the findings of some clinical studies. The results indicate that the positioning of the lip influences the possible range of impingement-free motion. Considering this and the improved joint stability of using a lipped liner, a balance is required to achieve an optimal range of motion without increasing the risk of dislocation. This tool could potentially to be used to optimise lipped liner design and position, and could assist with the liner selection for patients based on their activities

Young patients have been reported to have a higher risk of revision following total hip arthroplasty than older cohorts. This was attributed to the higher activity level which led to increased wear, osteolysis, and component fracture. We prospectively assessed the clinical results, wear and osteolysis, the incidence of squeaking, and the survivorship of ceramic on ceramic THA in patients younger than 50 years (mean age of 42 [18–50] years). The series included 425 THAs in 370 patients with 368 hips followed for a minimum of 2 years (mean 7.1 years, range 2–14 years). All patients received uncemented acetabular components with flush-mounted acetabular liners using an 18 degree taper. No osteolysis was observed in any uncemented construct. There was osteolysis around one loose cemented femoral component. The survivorship for reoperation for implant revision was 96.7%. There were only two acetabular liner fractures (0.47%) and one femoral head fracture (0.24%). Two of the three fractures involved a fall from a significant height. There were no hip dislocations. Five patients (1.17%) noted rare or occasional squeaking. None had reproducible squeaking. In summary, the current study shows that ceramic-on-ceramic THAs in the young patient population are extremely reliable with a very low revision rate and an absence of wear-induced osteolysis. In addition, it shows that both bearing fracture in this young patient population typically occurs with polytrauma and squeaking issues that have been raised relative to ceramic bearings occur very rarely with the flush-mounted ceramic liner design used in this study

Introduction. Dual Mobility (DM) implants have gained popularity for the treatment and prevention of hip dislocation, with increased stability provided by a large diameter mobile liner. However, distal regions of the liner can impinge on soft-tissues like hip capsule and iliopsoas, leading to anterior hip pain. Additionally, soft-tissue impingement may trap the mobile liner, leading to excessive loading of the liner rim, from engagement with the femoral stem, and subsequent intra-prosthetic dislocation. The hypothesis of this study was that reducing the liner profile below the equator (contoured design) can mitigate soft-tissue impingement without compromising inner-head pull-out resistance and overall hip joint stability (Fig. 1). Methods. The interaction of conventional and contoured liners with anterior soft-tissues was evaluated in 10 cadaveric hips (5 specimens; 2 male, 3 female; age 65 ± 10 yrs; liner diameter 42–48mm) via visual observation and fluoroscopic imaging. A metal wire was sutured to the deep fibers of the iliopsoas tendon/muscle, and metal wires were embedded in the mobile liners for fluoroscopic visualization (Fig. 2). All soft-tissue except the anterior hip capsule and iliopsoas was removed, and a rope was attached to the iliopsoas to apply tension along its natural orientation. Resistance to inner-head pull-out was evaluated via Finite Element Analysis (FEA) by simulating a full cycle of insertion of the inner head into the mobile liner and subsequent pullout. The femoral head, acetabular shell, and stem were modeled as rigid, while the mobile liner was modeled as plastically deformable. Hip joint stability was evaluated by dynamic simulations in for two dislocation modes: (A) Posterior dislocation (at 90° hip flexion) with internal hip rotation; (B) Posterior dislocation (starting at 90° flexion) with combined hip flexion and adduction. A 44 mm diameter conventional and a 44 mm contoured liner were evaluated during these tests. Results. The cadaver experiments showed that distal portion of conventional liners impinge on anterior hip capsule and iliopsoas at low flexion angles (<30°). Additionally, when the hip moved from flexion into extension, the liner motion was blocked between posterior neck engagement, and anterior soft-tissue impingement. In all hips, the soft-tissue impingement / tenting was significantly reduced with contoured liners (Fig. 7). The change in tenting could be visualized as change in distance between the iliopsoas wire, and the contoured/conventional liners on sequential fluoroscopic images. The maximum reduction in iliopsoas tenting for a given specimen ranged from 1.8 mm to 5.5 mm. Additionally, the contoured and conventional liners had identical inner-head pull-out resistance (901N vs. 909N), jump distance (9.4 mm mode-A, 11.7 mm mode-B) and impingement-free range of motion (47° mode-A, 29° mode-B). Conclusion. This study showed that distal portions of conventional DM liners can impinge against iliopsoas and hip capsule in low flexion leading to functional impediment of liner motion. Additionally, reducing the liner profile below the equator led to significant reduction in soft-tissue impingement/tenting without affecting mechanical performance. Thus, a contoured dual mobility liner design may reduce the risk of anterior hip pain and intra-prosthetic dislocation resulting from soft-tissue impingement and liner entrapment. To view tables/figures, please contact authors directly

Introduction. Dual-mobility (DM) liners have increased popularity due to the range of motion and stability provided by these implants. However, larger head diameters have been associated with anterior hip pain, due to surrounding soft-tissue impingement, particularly the iliopsoas. To address this, an anatomically contoured dual mobility (ACDM) liner was designed by reducing the volume of the liner below the equator (Fig1). Previous cadaver studies have shown that the ACDM significantly reduces iliopsoas tenting and trapping of the liner compared to conventional designs. We created a finite element study based on previous cadaver testing to further analyze the effectiveness of the ACDM design in reducing soft-tissue impingement, specifically the tendon-liner contact pressure and the tendon stress. Methods. The finite element model was developed within COMSOL 4.3b. The psoas tendon was modelled as a Yeoh hyper-elastic Material, which uses 3 constants (c1-c3), density (1.73g/cm3) and a bulk modulus (26GPa)[Hirokawa,2000]. In a previous, separate study, the average stiffness of 10 psoas tendon samples (5 cadavers), were measured to be 339[N/mm] in the linear region with average width and thickness of 14mmX4mm. The 3 constants were tuned to match experimental uniaxial test data, and were 5[GPa], 0[Gpa], and 46[GPa] for c1, c2, and c3 respectively. The implant components were rigidly modeled relative to the psoas. Cadaver specific CT models were used to create the FEA geometry. The insertion points for the Psoas were digitally determined on the proximal end of the lesser trochanter, and the psoas notch on the pelvis for hip flexion angles of −15°, 0°, 15° and 30°. These insertion points determined the length of the psoas and its relative position to the femoral head in 3D. The specific liner size and position for each cadaver was determined by implant planning with the CT models. In this abstract, we only present data for 2 specimens (left/right hips) with 44mm conventional DM, and 44mm ACDM, matching specimen anatomy. A 500N tensile load was applied to the psoas tendon proximally to simulate moderate physiological loading, the average/max stresses and contact pressures between the psoas and the two liner designs were determined. Results. At all flexion angles from −15° to 30°, the ACDM had lower psoas-liner contact pressure and stress compared to the conventional liner. Both contact pressure and tendon stress decreased for both liners with increasing hip flexion. At −15° flexion angle, there was an average contact pressure difference of .51MPa between the conventional and ACDM designs, or 37% decrease in pressure when using the ACDM. The average difference in tendon stress was 67.9MPa, or a 59% decrease in stress when using the ACDM (fig2, fig3). Conclusion. This study utilized cadaver specific FEA models to evaluate interaction between the iliopsoas tendon and conventional and ACDM liners. Although this abstract presented FEA models for only four hips (two specimens), the results show a notable reduction in contact pressure and tendon stress with ACDM designs. This validates findings from previous cadaver studies, suggesting that anatomically contoured designs could reduce anterior hip pain and soft tissue impingement