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In an effort to prevent degeneration of articular cartilage associated with meniscectomies, both meniscal allografts and synthetic replacements are subjects of current interest and investigation. The objectives of the current study were to (1) determine whether a transversely isotropic, linearly elastic, homogeneous material model of the meniscal tissue is necessary to achieve a normal contact pressure distribution on the tibial plateau, (2) determine which material and boundary condition (attachments) parameters affect the contact pressure distribution most strongly, and (3) set tolerances on these parameters to restore the contact pressure distribution to within a specified error. To satisfy these objectives, a finite element model of the tibio-femoral joint of a human cadaveric knee (including both menisci) was used to study the contact pressure distribution on the tibial plateau. To validate the model, the contact pressure distribution on the tibial plateau was measured experimentally in the same knee used to create the model. Within physiologically reasonable bounds on five material parameters and four attachment parameters associated with a meniscal replacement, an optimization was performed under 1200 N of compressive load on the set of nine parameters to minimize the difference between the experimental and model results. The error between the experimental and model contact variables was minimized to 5.4%. The contact pressure distribution of the tibial plateau was sensitive to the circumferential modulus, axial/radial modulus, and horn stiffness, but relatively insensitive to the remaining six parameters. Consequently, both the circumferential and axial/radial moduli are important determinants of the contact pressure distribution, and hence should be matched in the design and/or selection of meniscal replacements. In addition, during surgical implantation of a meniscal replacement, the horns should be attached with high stiffness bone plugs, and the attachments of the transverse ligament and deep medial collateral ligament should be restored to minimize changes in the contact pressure distribution, and thereby possibly prevent the degradation of articular cartilage.

Knowledge of the anterior-posterior (AP) tibial contact locations is useful in assessing wear of tibial inserts and detecting posterior rim loading. The objectives of this study were to 1) create a new 2D planar model to determine AP tibial contact locations, 2) use the 2D planar model to determine AP tibial contact locations for cadaveric TKA knees, and 3) determine whether errors of the 2D planar model are lower than those of the penetration method. A slopes-of-sagittal profiles (SSP) model was created using mathematical functions to simulate articular surfaces of the tibial insert and femoral condyles. AP tibial contact locations were computed using the model and the penetration method and simultaneously measured with a custom tibial force sensor in 10 cadaveric TKA knees at 0°, 30°, 60°, and 90° of flexion in each compartment during passive motion. For each method, the overall bias, overall precision, and overall root mean square error (RMSE) were calculated from the differences between the computed AP tibial contact locations and the measured locations. The SSP model had an overall bias of 0.6 mm and precision of 2.8 mm which were significantly greater than the overall bias of −0.1 mm (p = 0.0369) and overall precision of 2.0 mm (p = 0.0021) of the penetration method. A planar model based on the analysis of single-plane radiographs did not decrease overall errors in AP tibial contact locations compared to the penetration method.

A common method used to study tibiofemoral joint biomechanics following total knee arthroplasty (TKA) is the lowest point method, which finds the lowest points of each femoral condyle in relation to the plane of the resected tibia. The objectives of this paper were twofold: 1) to use a circle-based model to demonstrate the large inherent error introduced when the lowest points are used to indicate anterior-posterior (AP) positions of contact by the femur on the tibial insert, 2) to use the circle-based model to estimate the magnitude of error. A circle-based model was created to simulate articular surfaces of the tibial insert and condyles of the femoral component and to demonstrate the error. Equations relating the error to radii of tibial and femoral articular surfaces were derived. The magnitude of the error was estimated for common low-conforming TKA components by determining radii using best-fit circles to approximate curvature of articular surfaces. Error in AP tibial insert contact locations is caused by the slope of the tibial articular surface and the magnitude increases with increasing slope and increasing radius of the femoral condyle. For radii approximating articular surfaces of common low-conforming components, relative errors range from 45% to 109%. The circle-based model effectively demonstrates the cause of the large error in using lowest points to indicate AP tibial insert contact locations and enables an estimate of relative error. Because relative error exceeds 45%, the lowest point method should not be used to indicate the AP tibial insert contact locations.

When used in in vitro studies, soft tissues such as the meniscus and articular cartilage are susceptible to dehydration and its effects, such as changes in size and shape as well as changes in structural and material properties. To quantify the effect of dehydration on the meniscus and articular cartilage, the first two objectives of this study were to (1) determine the percent change in meniscal dimensions over time due to dehydration, and (2) determine the percent change in articular cartilage thickness due to dehydration. To satisfy these two objectives, the third objective was to develop a new laser-based three-dimensional coordinate digitizing system (3-DCDS II) that can scan either the meniscus or articular cartilage surface within a time such that there is less than a 5% change in measurements due to dehydration. The new instrument was used to measure changes in meniscal and articular cartilage dimensions of six cadaveric specimens, which were exposed to air for 120 and 130 min, respectively. While there was no change in meniscal width, meniscal height decreased linearly by 4.5% per hour. Articular cartilage thickness decreased nonlinearly at a rate of 6% per hour after 10 min, and at a rate of 16% per hour after 130 min. The system bias and precision of the new instrument at 0° slope of the surface being scanned were 0.0 and 2.6 μ m , respectively, while at 45° slope the bias and precision were 31.1 and 22.6 μ m , respectively. The resolution ranged between 200 and 500 μ m . Scanning an area of 60 × 80 mm (approximately the depth and width of a human tibial plateau) took 8 min and a complete scan of all five sides of a meniscus took 24 min. Thus, the 3-DCDS II can scan an entire meniscus with less than 2% change in dimensions due to dehydration and articular cartilage with less than 0.4% change. This study provides new information on the amount of time that meniscal tissue and articular cartilage can be exposed to air before marked changes in size and shape, and possibly biomechanical, structural and material properties, occur. The new 3-DCDS II designed for this study provides fast and accurate dimensional measurements of both soft and hard tissues.

In the commonly used SIMM software, which includes a complete musculoskeletal model of the lower limbs, the reaction forces at the knee are computed. These reaction forces represent the bone-on-bone contact forces and the soft tissue forces (e.g. ligaments) other than muscles acting at the joint. In the knee model integrated into this software, a patellotibial joint rather than a patellofemoral joint is defined, and a force acting along the direction of the patellar ligament is not included. Although this knee model results in valid kinematics and muscle moment arms, the reaction forces at the knee calculated do not represent physiologic knee joint reaction forces. Hence our objectives were to develop a method of calculating physiologic knee joint reaction forces using the knee model incorporated into the SIMM software and to demonstrate the differences in the forces returned by SIMM and the physiologic forces in an example. Our method converts the anatomically fictional patellotibial joint into a patellofemoral joint and computes the force in an inextensible patellar ligament. In our example, the rectus femoris was fully excited isometrically, with the knee and hip flexed to 90°. The resulting SIMM tibiofemoral joint reaction force was primarily shear, because the quadriceps force was applied to the tibia via the fictional patellotibial joint. In contrast the physiologic tibiofemoral joint reaction force was primarily compression, because the quadriceps force was applied through the patellar ligament. This result illustrates that the physiologic knee joint reaction forces are profoundly different than the forces returned by SIMM. However physiologic knee joint reaction forces can be computed with postprocessing of SIMM results.

Mathematical models of the inter-relationship of muscle force, velocity, and activation are useful in forward dynamic simulations of human movement tasks. The objective of this work was to determine whether the parameters (maximum shortening velocity V max and shape parameter k) of a Hill-type muscle model, interrelating muscle force, velocity, and activation, are themselves dependent on the activation. To fulfill this objective, surface EMG signals from four muscles, as well as the kinematics and kinetics of the arm, were recorded from 14 subjects who performed rapid-release elbow extension tasks at 25%, 50%, 75%, and 100% activation (MVC). The experimental elbow flexion angle was tracked by a forward dynamic simulation of the task in which V max and k of the triceps brachii were varied at each activation level to minimize the difference between the simulated and experimental elbow flexion angle. Because a preliminary analysis demonstrated no dependency of k on activation, additional simulations were performed with constant k values of 0.15, 0.20, and 0.25. The optimized values of V max normalized to the average value within a subject were then regressed onto the activation. Normalized V max depended significantly on the activation ( p<0.001) for all values of k. Furthermore, the estimated V max values were not sensitive to the selected k value. The results support the use of Hill-type models in which V max depends on activation in forward dynamic simulations modeling muscles with mixed fiber-type composition recruited in the range of 25–100% activation. The use of more accurate models will lend greater confidence to the results of forward dynamic simulations.

The primary purpose of this investigation was to test the hypothesis that cycling economy, as measured by rate of oxygen consumption ( V ˙ O 2 ) in healthy, young, competitive cyclists pedaling at a constant workrate, increases (i.e. V ˙ O 2 decreases) when the attachment point of the foot to the pedal is moved posteriorly on the foot. The V ˙ O 2 of 11 competitive cyclists (age 26.8±8.9 years) was evaluated on three separate days with three anterior–posterior attachment points of the foot to the pedal (forward=traditional; rear=cleat halfway between the head of the first metatarsal and the posterior end of the calcaneous; and mid=halfway between the rear and forward positions) on each day. With a randomly selected foot position, V ˙ O 2 was measured as each cyclist pedaled at steady state with a cadence of 90 rpm and with a power output corresponding to approximately 90% of their ventilatory threshold (VT) (mean power output 203.3±20.8 W). After heart rate returned to baseline, V ˙ O 2 was measured again as the subject pedaled with a different anterior–posterior foot position, followed by another rest period and then V ˙ O 2 was measured at the final foot position. The key finding of this investigation was that V ˙ O 2 was not affected by the anterior–posterior foot position either for the group ( p = 0.311 ) or for any individual subject ( p⩾0.156). The V ˙ O 2 for the group was 2705±324, 2696±337, and 2747±297 ml/min for the forward, mid, and rear foot positions, respectively. The practical implication of these findings is that adjusting the anterior–posterior foot position on the pedal does not affect cycling economy in competitive cyclists pedaling at a steady-state power output eliciting approximately 90% of VT.

Assessing the importance of non-driving intersegmental knee moments (i.e. varus/valgus and internal/external axial moments) on over-use knee injuries in cycling requires the use of a three-dimensional (3-D) model to compute these loads. The objectives of this study were: (1) to develop a complete, 3-D model of the lower limb to calculate the 3-D knee loads during pedaling for a sample of the competitive cycling population, and (2) to examine the effects of simplifying assumptions on the calculations of the non-driving knee moments. The non-driving knee moments were computed using a complete 3-D model that allowed three rotational degrees of freedom at the knee joint, included the 3-D inertial loads of the shank/foot, and computed knee loads in a shank-fixed coordinate system. All input data, which included the 3-D segment kinematics and the six pedal load components, were collected from the right limb of 15 competitive cyclists while pedaling at 225 W and 90 rpm. On average, the peak varus and internal axial moments of 7.8 and 1.5 N m respectively occurred during the power stroke whereas the peak valgus and external axial moments of 8.1 and 2.5 N m respectively occurred during the recovery stroke. However, the non-driving knee moments were highly variable between subjects; the coefficients of variability in the peak values ranged from 38.7% to 72.6%. When it was assumed that the inertial loads of the shank/foot for motion out of the sagittal plane were zero, the root-mean-squared difference (RMSD) in the non-driving knee moments relative to those for the complete model was 12% of the peak varus/valgus moment and 25% of the peak axial moment. When it was also assumed that the knee joint was revolute with the flexion/extension axis perpendicular to the sagittal plane, the RMSD increased to 24% of the peak varus/valgus moment and 204% of the peak axial moment. Thus, the 3-D orientation of the shank segment has a major affect on the computation of the non-driving knee moments, while the inertial contributions to these loads for motions out of the sagittal plane are less important.

Mathematical models of the muscle excitation are useful in forward dynamic simulations of human movement tasks. One objective was to demonstrate that sloped as opposed to rectangular excitation waveforms improve the accuracy of forward dynamic simulations. A second objective was to demonstrate the differences in simulated muscle forces using sloped versus rectangular waveforms. To fulfill these objectives, surface EMG signals from the triceps brachii and elbow joint angle were recorded and the intersegmental moment of the elbow joint was computed from 14 subjects who performed two cyclic elbow extension experiments at 200 and 300 deg/s. Additionally, the surface EMG signals from the leg musculature, joint angles, and pedal forces were recorded and joint intersegmental moments were computed during a more complex pedaling task (90 rpm at 250 W). Using forward dynamic simulations, four optimizations were performed in which the experimental intersegmental moment was tracked for the elbow extension tasks and four optimizations were performed in which the experimental pedal angle, pedal forces, and joint intersegmental moments were tracked for the pedaling task. In these optimizations the three parameters (onset and offset time, and peak excitation) defining the sloped (triangular, quadratic, and Hanning) and rectangular excitation waveforms were varied to minimize the difference between the simulated and experimentally tracked quantities. For the elbow extension task, the intersegmental elbow moment root mean squared error, onset timing error, and offset timing error were less from simulations using a sloped excitation waveform compared to a rectangular excitation waveform ( p < 0.001 ). The average and peak muscle forces were from 7% to 16% larger and 20–28% larger, respectively, when using a rectangular excitation waveform. The tracking error for pedaling also decreased when using a sloped excitation waveform, with the quadratic waveform generating the smallest tracking errors for both tasks. These results support the use of sloped over rectangular excitation waveforms to establish greater confidence in the results of forward dynamic simulations.

In an effort to prevent degeneration of articular cartilage associated with meniscectomies, both meniscal allografts and synthetic replacements have been studied. A number of biomechanical criteria may be important for a meniscal replacement to restore normal tibiofemoral contact pressure in the knee joint and hence be clinically successful. One of these criteria is geometric similarity. The objectives of the current study were to: determine the sensitivity of the contact variables of the tibial plateau to the transverse depth and width of both the lateral and medial menisci; determine the sensitivity of the contact variables of the tibial plateau to the cross-sectional width and height of the lateral and medial menisci; and determine the tolerances on each of the four parameters for both menisci. To satisfy these objectives, a previously developed finite element model of the tibiofemoral joint was used to compute the contact pressure distribution on the tibial plateau. The effect of the above-mentioned geometric parameters on the contact behavior was studied by perturbing the finite element model. Results showed that the contact variables are similarly sensitive to both the transverse and cross-sectional parameters of the menisci. Additionally the medial meniscal parameters have a greater effect on the contact variables than do the lateral meniscal parameters. Finally, less than a 0.5 mm change in the medial meniscal height and greater than a 1 mm change in the lateral meniscal height could be tolerated before the relative difference in the contact variables from those for the original geometry exceeded 10%. Thus in the design or selection of meniscal replacements, each of the four parameters should be measured when sizing a replacement tissue. Also tighter tolerances should be placed on the medial meniscal parameters compared to the lateral meniscal parameters.