Orthopaedic impaction-instruments are used to drive implants into the bone of the patient. Pre-clinical experimental testing protocols and computer models of those are used to assess robustness and functional efficiency of such instruments. This generally involves impaction of the instrument mounted on a substrate that should represent the mechanics of the patient. In this study, the effects of the substrate on stressing of the impaction-instruments were investigated using dynamic finite element analysis. Model results were compared with experimental data from lab protocols, which have been derived to recreate the mechanics of cadaveric implantations, which represent clinical conditions. FEA models of selected experimental protocols were created in which a simplified instrument was impacted on substrates with varying material properties and boundary conditions. After impaction, the instrument settled into a modal vibration which then decayed over time. The resulting axial strain data from the computational model was compared to strain-gauge data collected from experimental measurements. Strain signal amplitude, frequency and decay were compared. The damping-ratio was derived from the decay of the strain signal. The computational model slightly over-predicted the initial experimental strain amplitudes in all cases, but the frequency of the cyclic strain signals matched. However, the model underestimated the experimentally measured rate of signal decay. Inclusion of implant seating and soft-tissue conditions had little effect on decay. Clinical failures of impaction-instruments may be related to multiple fatigue cycles for each impaction and should be modelled accurately to allow failure prediction. Any soft substrate results in an impedance mismatch at the instrument interface, which reflects the pressure wave and causes vibration with a frequency related to the speed-of-sound in the instrument, and its geometry. While this could be accurately modelled computationally, signal decay was underestimated. Further experimental quantification of energy losses will be important to understand vibration decay.
Taper corrosion in Total Hip Arthroplasty has surfaced as a clinically relevant problem and has recently also been reported for metal heads against polyethylene. Low neck stiffness is a critical contributing factor. Catastrophic taper failures have been reported for one particular stem design with a small V-40 taper made from a less stiff titanium-alloy. The purpose of this study was to identify factors involved in the failure process. 31 revised CoCr heads ranging from 32 to 44m diameter combined with TMZF-Titanium alloy stem with a V-40 taper (Accolade I) were analysed. Stems were only available for catastrophic failure cases with dis-association (n=8) or taper fracture (n=1). Clinical data were limited to time-in-situ, patient gender and age. Head material loss increased with time in situ (r²=0.49, p<0.001). Longer heads and material loss exceeding 15mm³ showed bottoming out and consecutive catastrophic stem taper failure. Heads with failed stem tapers were all 36mm diameter. The head starts rotating on the stem taper after bottoming out, causing major abrasive wear, ultimately resulting in catastrophic failure; it is surprising that these catastrophic cases did not exhibit clinical symptoms due to raised Co and Cr metal ions, which must have resulted from the large amount of CoCr lost from the female head taper. This would have attracted medical attention and prevented catastrophic failure by taper dis-association. Control exams of patients treated with the respective stem type in combination with large CoCr heads should include metal ion determination in blood or serum, even if no clinical symptoms are present, in order to detect taper corrosion before catastrophic failure occurs.
Wear and corrosion between head and stem tapers of modular hip implants have recently been related to clinical failures, possibly due to high friction moments in poorly lubricated joints [1–2]. In-vivo measurements have revealed reversing joint friction moments in the hip during a gait cycle [3], which may foster relative motion between the modular components. Blood, soft tissue or bone debris at the taper interface during assembly can lead to decreased stability or increased stress concentrations due to non-uniform loading [4]. The purpose of this study is to investigate the influence of taper contamination and the assembly force on the seating characteristic of the head on the stem incorporating realistic reversing joint friction moments. Cobalt chrome heads (M-SPEC, 36mm, +1.5mm; n=5) were assembled on titanium femoral stems (Corail 12/14, both components Depuy Synthes; n=5) by quasistatic axial push-on forces (F=0.5kN, 1kN, 2kN). Heads were modified by milling a flat plane, to which the joint load was applied alternately to point A and point B for 20 cycles to provide reversing moments (heel-strike FA=1971N, MA=5.4Nm; toe-off FB=807N, MB=4.6Nm; Fig. 1). All 6 degrees of freedom of relative displacement between head and stem were determined in the unloaded state and after each loading cycle. A coordinate measurement machine (accuracy ±2µm) was used to determine the components positions. Pull-off forces were measured after the last loading cycle. Each taper was tested in pristine condition and then contaminated with a bone chip (1.7±0.2mg).Introduction
Methods
The early failure and revision of bimodular primary
total hip arthroplasty prostheses requires the identification of the
risk factors for material loss and wear at the taper junctions through
taper wear analysis. Deviations in taper geometries between revised
and pristine modular neck tapers were determined using high resolution
tactile measurements. A new algorithm was developed and validated
to allow the quantitative analysis of material loss, complementing
the standard visual inspection currently used. The algorithm was applied to a sample of 27 retrievals ( Cite this article:
Nucleotomy almost doubles the transmitted forces on the facet joints in human lumbar spine, regardless of the amount of removed nucleus pulposus. Low back pain involves the lumbar facet joints in 15% to 45% of the cases. The surgical intervention, nucleotomy, might also lead to painful facets with a high risk; however, its mechanism is yet to be fully understood. The aim of this study is to reveal how a small amount of nucleus removal can change the force transmission on the facets.Summary
Introduction
Head sizes used in total hip arthroplasty (THA) has increased drastically from the original 22mm used by Charnley. This is due to two factors: the use of hard-on-hard materials for the bearing articulation and the increasing problem of dislocation. Hard-on-hard materials enable mixed or fluid film lubrication due to their good wettability. The development of a fluid film layer is encouraged by smaller surface pressures (larger area) and higher velocity at the articulating interface (larger radius), suggesting that larger diameters exhibit better lubrication and such less wear. This was effectivly proven in pre-clinical simulator studies and used as argument to increase the diameters of metal-on-metall and ceramic-on-ceramic bearings. Clinically the tribological advantage of larger diameters has not yet been shown. For hard-on-soft bearings the situation is different. Due to the bad wettability of Polyethylene (PE), the abrasive wear regime is dominant. This means that the longer wear path of a larger diameter will inevitably carry a larger amount of wear debris. Despite this relation, the heads used in combination with PE were also increased up to 40mm diameter, justified by the overall greatly reduced wear amount of the new generation(s) of cross-linked PE and favourable simulator results. First in-vivo studies have shown that larger heads carry larger amounts of wear particles. Whether this increase is relevant with respect to osteolysis is still unclear and will have to be shown in longer term studies. Larger heads require a larger “jumping” distance until they dislocate. Consequently the use of larger heads reduces dislocation rates, which was shown in multiple clinical studies. However, the reduction in dislocation rate achieved by increasing diameters varies greatly. Some centres achieve dislocation rates below 1% with 28mm heads, other centres require 36mm heads to achieve the same result. No study shows any further advantage with head diameters larger than 36mm. Despite their obvious biomechanical advantage with regard to stability, larger heads also have large disadvantages. Larger heads carry inevitably larger friction moments, requiring better anchoring of the components. In unfavourable conditions (start-up, break-down of lubrication film), friction moments of hard-on-hard bearings can get very high and reach or even exceed the losening torque of the head on the taper. Depending on the head impaction foce during assembly, the loosening torques amount to 8 to 17Nm. Movement at the head-taper connection possibly causes wear and increased corrosion at this interface. Larger head diameters also require thinner shells and/or liners, leading to problems with liner chipping or incomplete seating. Large head diameters have also lead to the use of sub-hemispherical cups with reduced covering surface, increasing the risk of fluid film break down due to edge loading if not well positioned. Finally, larger heads might give the surgeon a wrong feeling of security regarding a sub-optimal positioned cup. The question regarding “the optimal” head diameter is open for discussion and needs to consider the bearing material used. Head size should be limited to a reasonable compromise, which based on the information currently available, could be 36mm. Join the “36 and under” club.The tribological aspect
The biomechanical aspect
There was a double-curve in 10 cases, single-curve in 5. The mean Cobb-angle was 72°. The standard construct was a double rod with pedicle/pelvic screws at the base, double claw at the proximal end and sublaminar wires at intervening levels. The most proximal level was T1–T4 in 13 cases. Instrumentation was carried to the pelvis in 7 cases. Intra-operative fractures occurred in 5 cases. The mean blood loss was 999 mls (range 295–5500). Spinal cord monitoring was abnormal in 3 cases. 1 case resulted in postoperative lower limb paralysis, which recovered. The mean hospital stay was 7.5 days. Serious postoperative complications included one case of bilateral anterior compartment syndrome and one tibial fracture. The mean curve correction was 31%. Two cases required revision surgery: extension of fusion to the pelvis. The mean follow-up was 22.7 months (range 4–40). There was no measurable change in position over time.
