Humans walk and run at a range of speeds. While steady locomotion at a given speed requires no net mechanical work, moving faster does demand both more positive and negative mechanical work per stride. Is this increased demand met by increasing power output at all lower limb joints or just some of them? Does running rely on different joints for power output than walking? How does this contribute to the metabolic cost of locomotion? This study examined the effects of walking and running speed on lower limb joint mechanics and metabolic cost of transport in humans. Kinematic and kinetic data for 10 participants were collected for a range of walking (0.75, 1.25, 1.75, 2.0 m s−1) and running (2.0, 2.25, 2.75, 3.25 m s−1) speeds. Net metabolic power was measured by indirect calorimetry. Within each gait, there was no difference in the proportion of power contributed by each joint (hip, knee, ankle) to total power across speeds. Changing from walking to running resulted in a significant (p = 0.02) shift in power production from the hip to the ankle which may explain the higher efficiency of running at speeds above 2.0 m s−1 and shed light on a potential mechanism behind the walk–run transition.

Humans walk and run over a wide range of speeds with remarkable efficiency. For steady locomotion, moving at different speeds requires the muscle–tendon units of the leg to modulate the amount of mechanical power the limb absorbs and outputs in each step. How individual muscles adapt their behavior to modulate limb power output has been examined using computer simulation and animal models, but has not been studied in vivo in humans. In this study, we used a combination of ultrasound imaging and motion analysis to examine how medial gastrocnemius (MG) muscle–tendon unit behavior is adjusted to meet the varying mechanical demands of different locomotor speeds during walking and running in humans. The results highlighted key differences in MG fascicle-shortening velocity with both locomotor speed and gait. Fascicle-shortening velocity at the time of peak muscle force production increased with walking speed, impairing the ability of the muscle to produce high peak forces. Switching to a running gait at 2.0 m⋅s ⁻¹ caused fascicle shortening at the time of peak force production to shift to much slower velocities. This velocity shift facilitated a large increase in peak muscle force and an increase in MG power output. MG fascicle velocity may be a key factor that limits the speeds humans choose to walk at, and may explain the transition from walking to running. This finding is consistent with previous modeling studies.

Dynamic measurements of human muscle fascicle length from sequences of B-mode ultrasound images have become increasingly prevalent in biomedical research. Manual digitisation of these images is time consuming and algorithms for automating the process have been developed. Here we present a freely available software implementation of a previously validated algorithm for semi-automated tracking of muscle fascicle length in dynamic ultrasound image recordings, "UltraTrack". UltraTrack implements an affine extension to an optic flow algorithm to track movement of the muscle fascicle end-points throughout dynamically recorded sequences of images. The underlying algorithm has been previously described and its reliability tested, but here we present the software implementation with features for: tracking multiple fascicles in multiple muscles simultaneously; correcting temporal drift in measurements; manually adjusting tracking results; saving and re-loading of tracking results and loading a range of file formats. Two example runs of the software are presented detailing the tracking of fascicles from several lower limb muscles during a squatting and walking activity. We have presented a software implementation of a validated fascicle-tracking algorithm and made the source code and standalone versions freely available for download.

Human feet have evolved to facilitate bipedal locomotion, losing an opposable digit that grasped branches in favor of a longitudinal arch (LA) that stiffens the foot and aids bipedal gait. Passive elastic structures are credited with supporting the LA, but recent evidence suggests that plantar intrinsic muscles (PIMs) within the foot actively contribute to foot stiffness. To test the functional significance of the PIMs, we compared foot and lower limb mechanics with and without a tibial nerve block that prevented contraction of these muscles. Comparisons were made during controlled limb loading, walking, and running in healthy humans. An inability to activate the PIMs caused slightly greater compression of the LA when controlled loads were applied to the lower limb by a linear actuator. However, when greater loads were experienced during ground contact in walking and running, the stiffness of the LA was not altered by the block, indicating that the PIMs’ contribution to LA stiffness is minimal, probably because of their small size. With the PIMs blocked, the distal joints of the foot could not be stiffened sufficiently to provide normal push-off against the ground during late stance. This led to an increase in stride rate and compensatory power generated by the hip musculature, but no increase in the metabolic cost of transport. The results reveal that the PIMs have a minimal effect on the stiffness of the LA when absorbing high loads, but help stiffen the distal foot to aid push-off against the ground when walking or running bipedally.

Previous reports of the mechanics and energetics of post-stroke hemiparetic walking have either not combined estimates of mechanical and metabolic energy or computed external mechanical work based on the limited combined limbs method. Here we present a comparison of the mechanics and energetics of hemiparetic and unimpaired walking at a matched speed. Mechanical work done on the body centre of mass (COM) was computed by the individual limbs method and work done at individual leg joints was computed with an inverse dynamics analysis. Both estimates were converted to average powers and related to simultaneous estimates of net metabolic power, determined via indirect calorimetry. Efficiency of positive work was calculated as the ratio of average positive mechanical power P ¯ + $$ {\overline{P}}^{+} $$ to net metabolic power. Total P ¯ + $$ {\overline{P}}^{+} $$ was 20% greater for the hemiparetic group (H) than for the unimpaired control group (C) (0.49 vs. 0.41 W · kg−1). The greater P ¯ + $$ {\overline{P}}^{+} $$ was partly attributed to the paretic limb of hemiparetic walkers not providing appropriately timed push-off P ¯ + $$ {\overline{P}}^{+} $$ in the step-to-step transition. This led to compensatory non-paretic limb hip and knee P ¯ + $$ {\overline{P}}^{+} $$ which resulted in greater total mechanical work. Efficiency of positive work was not different between H and C. Increased work, not decreased efficiency, explains the greater metabolic cost of hemiparetic walking post-stroke. Our results highlighted the need to target improving paretic ankle push-off via therapy or assistive technology in order to reduce the metabolic cost of hemiparetic walking.

The rigidity of the human foot is often described as a feature of our evolution for upright walking and is bolstered by a thick plantar aponeurosis that connects the heel to the toes. Previous descriptions of human foot function consider stretch of the plantar aponeurosis via toe extension (windlass mechanism) to stiffen the foot as it is levered against the ground for push-off during walking. In this study, we applied controlled loading to human feet in vivo , and studied foot function during the push-off phase of walking, with the aim of carefully testing how the foot is tensioned during contact with the ground. Both experimental paradigms revealed that plantar aponeurosis strain via the ‘windlass mechanism' could not explain the tensioning and stiffening of the foot that is observed with increased foot-ground contact forces and push-off effort. Instead, electromyographic recordings suggested that active contractions of ankle plantar flexors provide the source of tension in the plantar aponeurosis. Furthermore, plantar intrinsic foot muscles were also contributing to the developed tension along the plantar aspect of the foot. We conclude that active muscular contraction, not the passive windlass mechanism, is the foot's primary source of rigidity for push-off against the ground during bipedal walking.