Balancing in total knee replacement is generally carried out using the feel and experience of the surgeon, using spacer blocks or distractors. However, such a method is not generally applicable to all surgeons and nor does it provide quantitative data of the balancing itself. One approach is the use of instrumented distractors, which have been used to monitor soft tissue releases or indicate a flexion cut for equal lateral and medial forces. More recently an instrumented tibial trial has been introduced which measures and displays the magnitude and location of the loads on the lateral and medial plateaus, during various manoeuvres carried out at surgery. The data set is then used by the surgeon to determine options, whether soft tissue releases or bone cut adjustments, to achieve lateral-medial equality.

The testing method consisted of mounting the femoral component rigidly in a fixture on the vertical arm of an MTS machine. The tibial component was fixed on to a platform which allowed varus-valgus correction, and where the component could be displaced or rotated in a horizontal plane. Two of each size times 4 sizes of production components were tested. Compressive forces from 0–400N in steps of 50N were applied and the readings taken. There were strong correlations between applied and measured forces with mean Pearson's Correlation Coefficient of 0.958.

The special tests under different conditions did not have any effect on the output values. The output data proved to be repeatable under Central Loading with a maximum standard deviation of ± 15.36N at the highest applied force of 400N. “Low battery” did not adversely affect the data. Applying the load steadily to maximum versus load-unload-zero tests produced similar results. Lubrication versus no lubrication tests produced no changes to the results. There was no cross talk of the electronics within the device when loaded on one condyle. For both central and anterior-posterior loading, the contact points were centered medial-lateral on the GUI display, and tracked contact point translation appropriately.

Anterior-posterior loading did create output load variance at the extremes. However, it enabled the validation of the relationship of the femur on the trial surface. In addition, malrotation would be indicated by the femur riding up on the anterior or posterior tibial edges, important for soft tissue tension in all flexion angles.

In conclusion, the sensors provided data which was accurate to well within a practical range for surgical conditions. In our separate experiments on 10 cadaveric leg specimens, even the same test under controlled conditions could produce variations of up to ± 30N. Hence the sensor outputs indicated whether or not the knee was balanced to that level of tolerance, while the contact point data would indicate contacts too close to the anterior or posterior of the tibial surface.

The analysis of hip joint vibrations (phonoarthrography, vibration arthrometry, vibroarthrography, hip auscultation) has been explored as a means to assess joint pathologies, disease status and recently, incipient prosthesis failure. Frequencies < 100Hz have been used to diagnose gross pathology and wear in knee prostheses, frequencies from 1k to 10k Hz for progression of osteoarthritis, and frequencies > 10k Hz for loosening of cemented hip prostheses. It is possible that detailed analysis of higher frequencies could detect and quantify the smaller geometric changes (asperities) that develop in articular prosthetic wear.

We examined the ultrasound emission generated by various types of hip prostheses and native hips of 98 patients. The ultrasonic transducer was attached to the skin over the greater trochanter with a hypoallergenic, transparent dressing using a standard acoustic coupling gel layer on the microphone face to improve skin contact. The transducer was attached by a 2m cord to a battery operated, data recorder/logger. The patients were asked to sit in a chair, rise, sit again and then rise and take 5 steps while recording the acoustic data from these two movements of sitting and walking. This procedure was repeated for the opposite hip in each patient as well. Acoustic emission analysis examined frequency distributions and power spectrums of the recorded signals and their relations to prosthesis type and implantation time. Review of x-rays of prosthetic and native hips was carried out with OA grading and prosthetic wear quantification.

We have obtained data on 79 metal-polyethylene (average duration of 8.5 years; 0.1–28), 20 ceramic-ceramic (average duration of 8.5 years; 0.5–10), 17 metal-metal (average duration of 1.2 years; 0.1–5.5) and 15 ceramic-polyethylene (average duration of 0.6 years; 0.1–1) hip arthroplasties as well as 75 native hips.

Analysis of the data enabled us to tell the difference between patients whose native hips did not cause them any discomfort and those patients with painful osteoarthritis (initial findings indicate that OA severity can be quantified as well). The measurements of wear of the metal-polyethylene prostheses obtained from patients’ x-rays were compared to an analysis of the ultrasonic emissions, a homogeneity showed no significant differences (all p’s > 0.24) between the curve type and amount of wear of the prosthesis polyethylene.

Our data suggests that we are capable of assessing the status of OA by acoustic emission. Further analysis of wear data coupled to ultrasonic emission is needed for accurate quantification of THA wear.

The site of initiation of failure of a cemented femoral component is usually the prosthesis-cement interface. Strengthening this interface with porosity reduction may improve survivorship. Cement pores which propagate crack formation can be reduced by vacuum mixing or centrifugation, but this does not effect interface porosity.

Utilising simulated stems cemented into a “Sawbones” femur in a manner replicating surgery, we determined the effect of stem warming on various parameters. Maximum temperature and time of polymerisation, mechanical strength, porosity reduction and pore distribution in the cement mantle were measured with stems at room temperature (RT), 37, 44, and 50 degrees Celcius. Mechanical testing included initial “push-out” tests, tests after agiing in 37 degrees Celcius saline for two weeks,and fatigue testing (3 HZ at 90% initial failure load). Porosity distribution was measured by the percentage area of pores on the interface surfaces and the transverse plane.

Polymerisation time decreased as the stems were heated. The time decreased from 8.1 minutes at RT to 5.9 minutes at 50 degrees Celcius. The maximum temperature in the cement mantle rose from 50.2 to 56.4 degrees Celcius comparing stems at RT to those at 37 degrees Celcius, and did not elevate further as stems were preheated to 44 and 50 degrees Celcius. Similarly, static and fatigue interface strength improved by preheating stems, but no significant gain compared to RT stems was realised by heating above 37 degrees Celcius.

A dramatic reduction in porosity at the prosthesis-cement region was found with the heated stems, with no additional benefit to heating beyond 37 degrees Celcius. An increase in porosity at the cement-bone interface was noted as stems were heated. This may be due to the direction of polymerisation shrinkage in the cement mantle as influenced by stem temperature.