Helical Axes of Skeletal Knee Joint Motion During Running

The purpose of this study was to determine the changes in the axis of rotation of the knee that occur during the stance phase of running. Using intracortical pins, the three-dimensional skeletal kinematics of three subjects were measured during the stance phase of five running trials. The stance phase was divided into equal motion increments for which the position and orientation of the finite helical axes (FHA) were calculated relative to a tibial reference frame. Results were consistent within and between subjects. At the beginning of stance, the FHA was located at the midepicondylar point and during the flexion phase moved 20mm posteriorly and 10mm distally. At the time of peak flexion, the FHA shifted rapidly by about 10–20mm in proximal and posterior direction. The angle between the FHA and the tibial transverse plane increased gradually during flexion, to about 15° of medial inclination, and then returned to zero at the start of the extension phase. These changes in position and orientation of FHA in the knee should be considered in analyses of muscle function during human movement, which require moment arms to be defined relative to a functional rotation axis. The finding that substantial changes in axis of rotation occurred independent of flexion angle suggests that musculoskeletal models must have more than one kinematic degree-of-freedom at the knee. The same applies to the design of knee prostheses, if the goal is to restore normal muscle function

Helical Axes of Skeletal Knee Joint Motion During Running

The purpose of this study was to determine the changes in the axis of rotation of the knee that occur during the stance phase of running. Using intracortical pins, the three-dimensional skeletal kinematics of three subjects were measured during the stance phase of five running trials. The stance phase was divided into equal motion increments for which the position and orientation of the finite helical axes (FHA) were calculated relative to a tibial reference frame. Results were consistent within and between subjects. At the beginning of stance, the FHA was located at the midepicondylar point and during the flexion phase moved 20mm posteriorly and 10mm distally. At the time of peak flexion, the FHA shifted rapidly by about 10–20mm in proximal and posterior direction. The angle between the FHA and the tibial transverse plane increased gradually during flexion, to about 15° of medial inclination, and then returned to zero at the start of the extension phase. These changes in position and orientation of FHA in the knee should be considered in analyses of muscle function during human movement, which require moment arms to be defined relative to a functional rotation axis. The finding that substantial changes in axis of rotation occurred independent of flexion angle suggests that musculoskeletal models must have more than one kinematic degree-of-freedom at the knee. The same applies to the design of knee prostheses, if the goal is to restore normal muscle function

Pre-Impact Lower Extremity Posture and Brake Pedal Force Predict Foot and Ankle Forces During an Automobile Collision

Background: The purpose of this study was to determine how a driver’s foot and ankle forces during a frontal vehicle collision depend on initial lower extremity posture and brake pedal force. Method of Approach: A 2D musculoskeletal model with seven segments and six right-side muscle groups was used. A simulation of a three-second braking task found 3647 sets of muscle activation levels that resulted in stable braking postures with realistic pedal force. These activation patterns were then used in impact simulations where vehicle deceleration was applied and driver movements and foot and ankle forces were simulated. Peak rearfoot ground reaction force ~FRF! , peak Achilles tendon force ~FAT! , peak calcaneal force ~FCF! and peak ankle joint force ~FAJ! were calculated.Results: Peak forces during the impact simulation were 4766687 N ~FRF! , 29346944 N ~FCF! and 24496918 N ~FAJ! . Many simulations resulted in force levels that could cause fractures. Multivariate quadratic regression determined that the pre-impact brake pedal force (PF), knee angle (KA) and heel distance (HD) explained 72% of the variance in peak FRF , 62% in peak FCF and 73% in peak FAJ . Conclusions: Foot and ankle forces during a collision depend on initial posture and pedal force. Braking postures with increased knee flexion, while keeping the seat position fixed, are associated with higher foot and ankle forces during a collision

Pre-Impact Lower Extremity Posture and Brake Pedal Force Predict Foot and Ankle Forces During an Automobile Collision

Background: The purpose of this study was to determine how a driver’s foot and ankle forces during a frontal vehicle collision depend on initial lower extremity posture and brake pedal force. Method of Approach: A 2D musculoskeletal model with seven segments and six right-side muscle groups was used. A simulation of a three-second braking task found 3647 sets of muscle activation levels that resulted in stable braking postures with realistic pedal force. These activation patterns were then used in impact simulations where vehicle deceleration was applied and driver movements and foot and ankle forces were simulated. Peak rearfoot ground reaction force ~FRF! , peak Achilles tendon force ~FAT! , peak calcaneal force ~FCF! and peak ankle joint force ~FAJ! were calculated.Results: Peak forces during the impact simulation were 4766687 N ~FRF! , 29346944 N ~FCF! and 24496918 N ~FAJ! . Many simulations resulted in force levels that could cause fractures. Multivariate quadratic regression determined that the pre-impact brake pedal force (PF), knee angle (KA) and heel distance (HD) explained 72% of the variance in peak FRF , 62% in peak FCF and 73% in peak FAJ . Conclusions: Foot and ankle forces during a collision depend on initial posture and pedal force. Braking postures with increased knee flexion, while keeping the seat position fixed, are associated with higher foot and ankle forces during a collision

Foot and Ankle Forces During an Automobile Collision: the Influence of Muscles

Muscles have a potentially important effect on lower extremity injuries during an automobile collision. Computational modeling can be a powerful tool to predict these effects and develop protective interventions. Our purpose was to determine how muscles influence peak foot and ankle forces during an automobile collision. A 2-D bilateral musculoskeletal model was constructed with seven segments. Six muscle groups were included in the right lower extremity, each represented by a Hill muscle model. Vehicle deceleration data were applied as input and the resulting movements were simulated. Three models were evaluated: no muscles (NM), minimal muscle activation at a brake pedal force of 400N (MN), and maximal muscle activation to simulate panic braking (MX). Muscle activation always resulted in large increases in peak joint force. Peak ankle joint force was greatest for MX (10120N), yet this model also had the lowest peak rearfoot force (629N). Peak force on the Achilles tendon was 4.5 times greater, during MX (6446N) compared to MN (1430N). We conclude that (1) external and internal forces are dependent on muscles, (2) muscle activation level could exacerbate axial loading injuries, (3) external and internal forces can be inversely related once muscle properties are included

Foot and Ankle Forces During an Automobile Collision: the Influence of Muscles

Muscles have a potentially important effect on lower extremity injuries during an automobile collision. Computational modeling can be a powerful tool to predict these effects and develop protective interventions. Our purpose was to determine how muscles influence peak foot and ankle forces during an automobile collision. A 2-D bilateral musculoskeletal model was constructed with seven segments. Six muscle groups were included in the right lower extremity, each represented by a Hill muscle model. Vehicle deceleration data were applied as input and the resulting movements were simulated. Three models were evaluated: no muscles (NM), minimal muscle activation at a brake pedal force of 400N (MN), and maximal muscle activation to simulate panic braking (MX). Muscle activation always resulted in large increases in peak joint force. Peak ankle joint force was greatest for MX (10120N), yet this model also had the lowest peak rearfoot force (629N). Peak force on the Achilles tendon was 4.5 times greater, during MX (6446N) compared to MN (1430N). We conclude that (1) external and internal forces are dependent on muscles, (2) muscle activation level could exacerbate axial loading injuries, (3) external and internal forces can be inversely related once muscle properties are included

The Influence of Orthotic Devices and Vastus Medialis Strength and Timing on Patellofemoral Loads During Running

Objective. To use a musculoskeletal model and simulation of running to examine: (1) the influence of two commonly prescribed treatments for patellofemoral pain (vastus medialis oblique strengthening and orthoses) and (2) the functional significance of timing differences between vastus medialis oblique and vastus lateralis on lateral patellofemoral joint loads. Design. A three-dimensional musculoskeletal model of the lower extremity was used to simulate running at 4 m/s. Background. Repetitive and excessive joint loading is often associated with overuse injuries that require clinical treatments to reduce pain and restore function. Affecting one in four runners, patellofemoral pain is one of the most common injuries in running. Although conservative treatments have been reported to successfully treat patellofemoral pain, the effectiveness is often based on subjective or empirical data, which have generated disagreement on the most effective treatment. Methods. Nine subject specific running simulations were generated and experiments were performed by applying the treatments and timing differences to the nominal simulations. Results. Both treatments significantly reduced the average patellofemoral joint load and the vastus medialis strengthening also significantly reduced the peak patellofemoral joint load. In addition, when the vastus medialis oblique timing was delayed and advanced relative to the vastus lateralis timing, a significant increase and decrease in the joint load was observed, respectively, during the loading response. Conclusions. Increasing vastus medialis oblique strength yielded more consistent results across subjects than the orthosis in reducing patellofemoral joint loads during running. The effect of orthoses was highly variable and sensitive to the individual subject\u27s running mechanics. Vastus medialis oblique activation timing is an important determinant of lateral patellofemoral joint loading during the impact phase. Relevance.These findings indicate that a reduction in patellofemoral pain may be achieved through techniques that selectively increase the vastus medialis oblique strength. Therefore, future studies should be directed towards identifying such techniques. Additionally, the functional significance of timing differences between the vastus medialis oblique and vastus lateralis is an important consideration in patellofemoral pain treatment and orthoses may be beneficial for some patients depending on their running mechanics

Mechanical Properties of the Tendinous Equine Interosseus Muscle are Affected by in Vivo Transducer Implantation

Liquid metal strain gauges (LMSGs) were implanted in the tendinous interosseous muscle, also called suspensory ligament (SL), in the forelimbs of 6 ponies in order to quantify in vivo strains and forces. Kinematics and ground reaction forces were recorded simultaneously with LMSG signals at the walk and the trot prior to implantation, and 3 and 4 days thereafter. The ponies were euthanised and tensile and failure tests were performed on the instrumented tendons and on the tendons of the contra lateral limb, which were instrumented post mortem. The origo–insertional (OI) strain of the SL was computed from pre- and post-operative kinematics, using a 2D geometrical model. The LMSG-recorded peak strain of the SL was 5.4±0.9% at the walk and 9.1±1.3% at the trot. Failure occurred at 15.4±2.1% (mean±S.D.). The LMSG strain was higher than the simultaneously recorded OI strain 0.5±0.7% strain at the walk and 2.2±1.1% strain at the trot. Post-operative OI strains were only slightly higher than pre-operative values. Failure strains of in vivo instrumented SLs were 2.0±1.2% strain higher, and failure forces were slightly lower, than those of the contra lateral SLs that were instrumented post mortem. SL strains appeared to be considerably higher than those found in earlier acute experiments. Differences between in vivo LMSG and OI strains, supported by lower failure strains comparing in vivo and post mortem instrumented SLs, revealed that local changes in tendon mechanical properties occurred within 3 to 4 days after transducer implantation. Therefore, measurements of normal physiological tendon strains should be performed as soon as possible after transducer implantation

The Influence of Foot Positioning on Ankle Sprains

The goal of this study was to examine the influence of changes in foot positioning at touch-down on ankle sprain occurrence. Muscle model driven computer simulations of 10 subjects performing the landing phase of a side-shuffle movement were performed. The relative subtalar joint and talocural joint angles at touchdown were varied, and each subject-specific simulation was exposed to a set of perturbed floor conditions. The touchdown subtalar joint angle was not found to have a considerable influence on sprain occurrence, while increased touchdown plantar flexion caused increased ankle sprain occurrences. Increased touchdown plantar flexion may be the mechanism which causes ankles with a history of ankle sprains to have an increased susceptibility to subsequent sprains. This finding may also reveal a mechanism by which taping of a sprained ankle or the application of an ankle brace leads to decreased ankle sprain susceptibility