Patellar crepitus and clunk are tendofemoral-related complications predominantly associated with posterior-stabilizing (PS) total knee arthroplasty (TKA) designs [1]. Contact between the quadriceps tendon and the femoral component can cause irritation, pain, and catching of soft-tissue within the intercondylar notch (ICN). While the incidence of tendofemoral-related pathologies has been documented for some primary TKA designs, literature describing revision TKA is sparse. Revision components require a larger boss resection to accommodate a constrained post-cam and stem/sleeve attachments, which elevates the entrance to the ICN, potentially increasing the risk of crepitus. The objective of this study was to evaluate tendofemoral contact in primary and revision TKA designs, including designs susceptible to crepitus, and newer designs which aim to address design features associated with crepitus. Six PS TKA designs were evaluated during deep knee bend using a computational model of the Kansas knee simulator (Figure 1). Prior work has demonstrated that tendofemoral contact predictions from this model can differentiate between TKA patients with patellar crepitus and matched controls [2]. Incidence of crepitus of up to 14% has been reported in Insall-Burstein® II and PFC® Sigma® designs [3]. These designs, in addition to PFC® Sigma® TC3 (revision component), were included in the analyses. Primary and revision components of newer generation designs (NexGen®, Attune® and Attune® Revision) were also included. Designs were evaluated in a patient model with normal Insall-Salvati ratio and a modified model with patellar tendon length reduced by two standard deviations (13mm) to assess worst-case patient anatomy.Introduction
Methods
Within total hip replacement, articulation of the femoral head near the rim of the acetabular liner creates undesirable conditions leading to a propensity for dislocation[1], increased contact stresses[2], increased load and torque imparted on the acetabular component[3], and increased wear[4]. Propensity for rim loading is affected by prosthesis placement, as well as the kinematics and loading of the patient. The present study investigates these effects. CT scans from an average-sized patientwere segmented for the hemipelvis and femur of interest. DePuy Synthes implant models were aligned in a neutral position in Hypermesh. The acetabular liner was assigned deformable solid material properties, and the remainder of the model was assigned rigid properties. Joint reaction forces and kinematics of hip flexion were taken from the public Orthoload database to represent ADLs [5]: Active flexion lying on a table, gait, bending to lift and move a load, and sit-stand. The pelvis was fully constrained, while three-degree-of-freedom (3-DOF) forces were applied to the femur. Hip flexion was kinematically-prescribed while internal-external (I-E) and adduction-abduction (Ad-Ab) DOFs were constrained. Angles of acetabular implant positioning were based on published data by Rathod [6]. Femoral implant position was chosen based on cadaveric in vitro DePuy Synthes measurements of variation in femoral prosthesis position reported previously [7]. Acetabular and Femoral alignment angles were represented for nominal position, as well as positioning + 1σ and + 2σ from the mean in both anteversion and inclination for acetabular components, and both Varus/Valgus and Flexion (angle in sagittal plane) for the femoral component. The analyses were automated within Matlab to execute 68 finite element analyses in Abaqus Explicit and structured in a DOE style analysis with Cup inclination, Cup version, Stem Flexion, and Stem Varus/Valgus, and Activity as variables of interest (64 runs + 4 centerpoints = 68 analyses). From a previous study it was known that acetabular component inclination had the greatest effect on contact pressure location [7], so all data were analyzed relative to inclination, allowing other positioning variables to be represented as variation per inclination position. Results are presented as a percentage, with 0% being pole loading and 100% being rim loading, to normalize for head diameter.INTRODUCTION
METHODS
Computational models are the primary tools for efficient design-phase exploration of knee replacement concepts before in vitro testing. To improve design-phase efficiency, a subject-specific computational platform was developed that allows designers to assess devices in realistic conditions by directly integrating subject-specific experimental data in these models. Early in the design-phase of new implant design, numerous in vitro tests would be desirable to assess the influence of design parameters or component alignment on the performance of the device. However, cadaveric testing of knee replacement devices is a costly and time-consuming procedure, requiring manufacture of parts, preparation of cadaveric specimens, and personnel to carry of the experiments. Validated computational models are ideally suited for pre-clinical, high-volume design evaluation. Initial development of these models requires substantial time and expertise; once developed, however, computational simulations may be applied for comparative evaluation of devices in an extremely efficient manner [Baldwin et al. 2012]. Still, computational models are complementary of experimental testing and for this reason, computational models tuned with subject-specific experimental data, e.g. soft tissue parameters, could bring even more efficiency in the design phase. The objective of the current study was to develop a platform of tools that easily allows for subject-specific knee simulations. The system integrates with commercially available medical imaging and finite element software to allow for direct, efficient comparison of designs and surgical alignment under a host of different boundary conditions.Summary Statement
Introduction
Complex foot and ankle surgery and reconstruction require accurate preoperative planning. In the foot procedures are challenging and can be associated with a range of complications. The aim of planning is to correct only the deformity and prevent extensive surgery. Knowledge of foot and ankle morphometry is vital. For comparison between different subjects the coordinate system must be constant. To the authors knowledge there has been no previous description of a coordinate system for the foot and ankle. CT images of ten anatomically normal feet were segmented in a general purpose segmentation program for grey value images and imported to a shape analysis program for biomechanics. A coordinate frame was defined in a 3 × 3 identity matrix using the inter-malleolar axis and a fibular diaphyseal centroidal axis in the construction. Centroidal vectors were defined in the metatarsals. Correlation of metatarsal length, inter-metatarsal angles, inter-malleolar distance and height was carried out. The forefoot was examined in relation to the medial and lateral columns. Metatarsal length had a significant correlation within each column and between the two columns notably in the 3rd (0.525 – 0.965) metatarsal at the columns junction. The 3rd metatarsals also correlated significantly (−0.583) with the inter-metatarsal angles. There was a weak correlation between the 1st 3rd and the 3rd 5th inter-metatarsal angles directly however, each had a large correlation with the 1st 5th inter-metatarsal angle (0.734 – 0.950). There was also a large correlation between the individual’s stature and the metatarsal length and the inter-malleolar distance. We have presented a means defining a coordinate system for three dimensional analyses in the foot and ankle. This coordinate system can be used for meaningful comparison of data between multiple subjects. We have shown that this coordinate system to be effective in practice in the morphometrical analysis of the normal forefoot.
Many pedicle screw instrumentation systems are currently available to the spine surgeon. Each system has its unique characteristics. It is important for the surgeon to understand the differences in these pedicle screw systems1 Following the introduction of a new spinal instrumentation set to our clinical practice we encountered two cases of pedicle screw breakage. We thus decided to investigate the mechanism of this screw failure (screw A) in these particular cases and to compare the biomechanical properties, through independent analysis, of a variety of pedicle screws from different manufacturers. Samples of the broken pedicle screws were retrieved at surgery. Surface analysis of the fracture area using the electron microscope, demonstrated features consistent with fatigue fracture. Pedicle screws of comparable size from a variety of manufacturers were gathered for independent analysis. Shadowgraph analysis was performed of each screw allowing multiple measurements to be taken of the screw’s geometry. Using this data stress concentration factors were determined demonstrating screw A to have larger values than all the other screws ranging from 2 – 3.6 times the nominal stress. The smaller teeth of screw A, spaced further apart than in the other screws, means that the large proportion of the load which would be carried by the threads is distributed over a smaller area resulting in higher stresses in the threads. The sharp corner at the root of the thread, acting as a stress concentrator, would become the focal point of these high stresses, and magnify them by 2 to 3.6 times. These increased stresses most likely account for an increased susceptibility to fatigue fracture seen in screw A. In conclusion it is important to be careful with the introduction and use of new pedicle screw materials and designs, that all the standard biomechanical testing has been performed to a satisfactory standard. Knowing the physical characteristics of the available pedicle screw instrumentation systems may allow the choice of pedicle screw best suited for a given clinical situation.
