Energy cost and muscular activity required for propulsion during walking.

We reasoned that with an optimal aiding horizontal force, the reduction in metabolic rate would reflect the cost of generating propulsive forces during normal walking. Furthermore, the reductions in ankle extensor electromyographic (EMG) activity would indicate the propulsive muscle actions. We applied horizontal forces at the waist, ranging from 15% body weight aiding to 15% body weight impeding, while subjects walked at 1.25 m/s. With an aiding horizontal force of 10% body weight, 1) the net metabolic cost of walking decreased to a minimum of 53% of normal walking, 2) the mean EMG of the medial gastrocnemius (MG) during the propulsive phase decreased to 59% of the normal walking magnitude, and yet 3) the mean EMG of the soleus (Sol) did not decrease significantly. Our data indicate that generating horizontal propulsive forces constitutes nearly half of the metabolic cost of normal walking. Additionally, it appears that the MG plays an important role in forward propulsion, whereas the Sol does not.     

Metabolic energy and muscular activity required for leg swing in running.

The metabolic cost of leg swing in running is highly controversial. We investigated the cost of initiating and propagating leg swing at a moderate running speed and some of the muscular actions involved. We constructed an external swing assist (ESA) device that applied small anterior pulling forces to each foot during the first part of the swing phase. Subjects ran on a treadmill at 3.0 m/s normally and with ESA forces up to 4% body weight. With the greatest ESA force, net metabolic rate was 20.5% less than during normal running. Thus we infer that the metabolic cost of initiating and propagating leg swing comprises approximately 20% of the net cost of normal running. Even with the greatest ESA, mean electromyograph (mEMG) of the medial gastrocnemius and soleus muscles during later portions of stance phase did not change significantly compared with normal running, indicating that these muscles are not responsible for the initiation of leg swing. However, with ESA, rectus femoris mEMG during the early portions of swing phase was as much as 74% less than during normal running, confirming that it is responsible for the propagation of leg swing.     

Biomechanical mechanism for transitions in phase and frequency of arm and leg swing during walking

As humans increase walking speed, there are concurrent transitions in the frequency ratio between arm and leg movements from 2:1 to 1:1 and in the phase relationship between the movements of the two arms from in-phase to out-of-phase. Superharmonic resonance of a pendulum with monofrequency excitation had been proposed as a potential model for this phenomenon. In this study, an alternative model of paired pendulums with multiple-frequency excitations is explored. It was predicted that the occurrence of the concurrent transitions was a function of (1) changes in the magnitude ratio of shoulder accelerations at step and stride frequencies that accompany changes in walking speed and (2) proximity of these frequencies to the natural resonance frequencies of the arms modeled as a pair of passive pendulums. Model predictions were compared with data collected from 14 healthy young subjects who were instructed to walk on a treadmill. Walking speeds were manipulated between 0.18 and 1.52 m/s in steps of 0.22 m/s. Kinematic data for the arms and shoulders were collected using a 3D motion analysis system, and simulations were conducted in which the movements of a double-pendulum system excited by the accelerations at the suspension point were analyzed to determine the extent to which the arms acted as passive pendulums. It was confirmed that the acceleration waveforms at the shoulder are composed primarily of stride and step frequency components. Between the shoulders, the stride frequency components were out-of-phase, while the step frequency components were in-phase. The amplitude ratio of the acceleration waveform components at the step and stride frequencies changed as a function of walking speed and were associated with the occurrence of the transitions. Simulation results using these summed components as excitatory inputs to the double-pendulum system were in agreement with actual transitions in 80% of the cases. The potential role of state-dependent active muscle contraction at shoulder joints on the occurrence of the transitions was discussed. Due to the tendency of arm movements to stay in the vicinity of their primary resonance frequency, these active muscle forces were hypothesized to function as escapements that created limit cycle oscillations at the shoulders resonant frequency.     

Linear approximations for swing leg motion during gait

The applicability of a linear systems analysis of two-dimensional swing leg motion was investigated. Two different linear systems were developed. A linear time-varying system was developed by linearizing the nonlinear equations describing swing leg motion about a set of nominal system and control trajectories. Linear time invariant systems were developed by linearizing about three different fixed limb positions. Simulations of swing leg motion were performed with each of these linear systems. These simulations were compared to previously performed nonlinear simulations of two-dimensional swing leg motion and the actual subject motion. Additionally, a linear system analysis was used to gain some insight into the interdependency of the state variables and controls. It was shown that the linear time varying approximation yielded an accurate representation of limb motion for the thigh and shank but with diminished accuracy for the foot. In contrast, all the linear time invariant systems, if used to simulate more than a quarter of the swing phase, yielded generally inaccurate results for thigh shank and foot motion.     

Lumbar back muscle activity during walking with a leg inequality

The influence of an artificial leg length discrepancy (= ALLD) on stride times, pelvic rotations and activity of the intrinsic lumbar back muscles (= ILBM) was investigated for 20 subjects. An ALLD was created by shoes with a raised sole. Walking with an ALLD produced an increase of the swing phase time and a decrease of the stance phase time for both feet. The influence of an ALLD on pelvic rotations in the sagittal and frontal plane and on ILBM-activity was small. Changes in pelvic rotations in the sagittal plane were too small to observe. The mean pelvic rotation angle in the frontal plane was changed 1.52 degrees when walking with an ALLD of 40 mm (6.9 degrees while standing with an ALLD of 40 mm with extended knees). Only small changes were found in activity time due to an ALLD (not in EMG-amplitude). The activity time of the ILBM around heel strike of the raised limb was increased and unilaterally shifted from toe off in the direction of heel strike with the raised limb.     

Energy cost of balance control during walking decreases with external stabilizer stiffness independent of walking speed

Human walking requires active neuromuscular control to ensure stability in the lateral direction, which inflicts a certain metabolic load. The magnitude of this metabolic load has previously been investigated by means of passive external lateral stabilization via spring-like cords. In the present study, we applied this method to test two hypotheses: (1) the effect of external stabilization on energy cost depends on the stiffness of the stabilizing springs, and (2) the energy cost for balance control, and consequently the effect of external stabilization on energy cost, depends on walking speed. Fourteen healthy young adults walked on a motor driven treadmill without stabilization and with stabilization with four different spring stiffnesses (between 760 and 1820 N m⁻¹) at three walking speeds (70%, 100%, and 130% of preferred speed). Energy cost was calculated from breath-by-breath oxygen consumption. Gait parameters (mean and variability of step width and stride length, and variability of trunk accelerations) were calculated from kinematic data. On average external stabilization led to a decrease in energy cost of 6% (p<0.005) as well as a decrease in step width (24%; p<0.001), step width variability (41%; p<0.001) and variability of medio-lateral trunk acceleration (12.5%; p<0.005). Increasing stabilizer stiffness increased the effects on both energy cost and medio-lateral gait parameters up to a stiffness of 1260 N m⁻¹. Contrary to expectations, the effect of stabilization was independent of walking speed (p=0.111). These results show that active lateral stabilization during walking involves an energetic cost, which is independent of walking speed.     

Energy cost of walking and running at extreme uphill and downhill slopes.

The costs of walking (Cw) and running (Cr) were measured on 10 runners on a treadmill inclined between -0.45 to +0.45 at different speeds. The minimum Cw was 1.64 +/- 0.50 J. kg(-1). m(-1) at a 1.0 +/- 0.3 m/s speed on the level. It increased on positive slopes, attained 17.33 +/- 1.11 J. kg(-1). m(-1) at +0.45, and was reduced to 0.81 +/- 0.37 J. kg(-1). m(-1) at -0.10. At steeper slopes, it increased to reach 3.46 +/- 0.95 J. kg(-1). m(-1) at -0.45. Cr was 3.40 +/- 0.24 J. kg(-1). m(-1) on the level, independent of speed. It increased on positive slopes, attained 18.93 +/- 1.74 J. kg(-1). m(-1) at +0.45, and was reduced to 1.73 +/- 0.36 J. kg(-1). m(-1) at -0.20. At steeper slopes, it increased to reach 3.92 +/- 0.81 J. kg(-1). m(-1) at -0.45. The mechanical efficiencies of walking and running above +0.15 and below -0.15 attained those of concentric and eccentric muscular contraction, respectively. The optimum gradients for mountain paths approximated 0.20-0.30 for both gaits. Downhill, Cr was some 40% lower than reported in the literature for sedentary subjects. The estimated maximum running speeds on positive gradients corresponded to those adopted in uphill races; on negative gradients they were well above those attained in downhill competitions.     

Segment interactions within the swing leg during unloaded and loaded running

In this study, designed to determine the effect of lower extremity inertia manipulation on joint kinetics and segment energetics during the swing phase, 15 male distance runners were filmed as they performed treadmill running (3.35 m s-1) under five load conditions: no added load and loads of 0.25 kg and 0.50 kg added to each thigh or each foot. Results of this study demonstrated that the energetics of the lower extremity movements during the swing phase of the running cycle were dominated by mechanical energy transfers between adjacent segments attributed to the joint reaction forces, which acted to redistribute mechanical energy within the system. These contributions were considerably greater than those of the net joint moments, which primarily reflected muscular generation and dissipation of mechanical energy. Lower extremity loading caused little change in the movement pattern of the swing leg. However, increases in the joint reaction forces and net moments and in the amount of work done and the energy transfer attributed to the reaction forces and moments were observed, but were limited to the joints proximal to the location of the added load. These results were consistent with the increased aerobic demand associated with increases in lower extremity inertia that have been reported elsewhere and also have implications for the manner in which the neuromuscular system controls the motion of the legs during running.     

Effects of aging and arm swing on the metabolic cost of stability in human walking

To gain insight into the mechanical determinants of walking energetics, we investigated the effects of aging and arm swing on the metabolic cost of stabilization. We tested two hypotheses: (1) elderly adults consume more metabolic energy during walking than young adults because they consume more metabolic energy for lateral stabilization, and (2) arm swing reduces the metabolic cost of stabilization during walking in young and elderly adults. To test these hypotheses, we provided external lateral stabilization by applying bilateral forces (10% body weight) to a waist belt via elastic cords while young and elderly subjects walked at 1.3m/s on a motorized treadmill with arm swing and with no arm swing. We found that the external stabilizer reduced the net rate of metabolic energy consumption to a similar extent in elderly and young subjects. This reduction was greater (6-7%) when subjects walked with no arm swing than when they walked normally (3-4%). When young or elderly subjects eliminated arm swing while walking with no external stabilization, net metabolic power increased by 5-6%. We conclude that the greater metabolic cost of walking in elderly adults is not caused by a greater cost of lateral stabilization. Moreover, arm swing reduces the metabolic cost of walking in both young and elderly adults likely by contributing to stability.     

Energy cost of treadmill and floor walking at self-selected paces

Oxygen uptake-velocity regression equations were developed for floor and level treadmill walking by having two groups of men, aged 19-29 years (n = 20) and 55-66 years (n = 22), walk at four self-selected paces, from "rather slowly" to "as fast as possible". A two-variable quadratic model relating VO2 (ml X kg-1 X min-1) to velocity (m X s-1) was adopted for prediction purposes. However, age and fatness significantly (P less than 0.05) interacted with treadmill walking speed, while age alone significantly interacted with floor speed. In addition, a significant difference was found between the energy cost of floor and treadmill walking. For example at the normal walking speed of 1.33 m X s-1, the energy cost for the treadmill (age 55-66 years) was 10.58 ml X kg-1 X min-1 and for the floor, 11.04 ml X kg-1 X min-1 (P less than 0.05). Four quadratic equations are therefore presented, one each for floor and treadmill in each of the two age-groups. The percent variance explained was between 87 and 95% for each of these equations.     

Predicting the metabolic cost of incline walking from muscle activity and walking mechanics

The goal of this study was to identify which muscle activation patterns and gait features best predict the metabolic cost of inclined walking. We measured muscle activation patterns, joint kinematics and kinetics, and metabolic cost in sixteen subjects during treadmill walking at inclines of 0%, 5%, and 10%. Multivariate regression models were developed to predict the net metabolic cost from selected groups of the measured variables. A linear regression model including incline and the squared integrated electromyographic signals of the soleus and vastus lateralis explained 96% of the variance in metabolic cost, suggesting that the activation patterns of these large muscles have a high predictive value for metabolic cost. A regression model including only the peak knee flexion angle during stance phase, peak knee extension moment, peak ankle plantarflexion moment, and peak hip flexion moment explained 89% of the variance in metabolic cost; this finding indicates that kinematics and kinetics alone can predict metabolic cost during incline walking. The ability of these models to predict metabolic cost from muscle activation patterns and gait features points the way toward future work aimed at predicting metabolic cost when gait is altered by changes in neuromuscular control or the use of an assistive technology.     

Energy cost of walking and gait instability in healthy 65- and 80-yr-olds.

This study tested whether the lower economy of walking in healthy elderly subjects is due to greater gait instability. We compared the energy cost of walking and gait instability (assessed by stride to stride changes in the stride time) in octogenarians (G80, n = 10), 65-yr-olds (G65, n = 10), and young controls (G25, n = 10) walking on a treadmill at six different speeds. The energy cost of walking was higher for G80 than for G25 across the different walking speeds (P < 0.05). Stride time variability at preferred walking speed was significantly greater in G80 (2.31 +/- 0.68%) and G65 (1.93 +/- 0.39%) compared with G25 (1.40 +/- 0.30%; P < 0.05). There was no significant correlation between gait instability and energy cost of walking at preferred walking speed. These findings demonstrated greater energy expenditure in healthy elderly subjects while walking and increased gait instability. However, no relationship was noted between these two variables. The increase in energy cost is probably multifactorial, and our results suggest that gait instability is probably not the main contributing factor in this population. We thus concluded that other mechanisms, such as the energy expenditure associated with walking movements and related to mechanical work, or neuromuscular factors, are more likely involved in the higher cost of walking in elderly people.     

Metabolic cost of generating muscular force in human walking: insights from load-carrying and speed experiments.

We sought to understand how leg muscle function determines the metabolic cost of walking. We first indirectly assessed the metabolic cost of swinging the legs and then examined the cost of generating muscular force during the stance phase. Four men and four women walked at 0.5, 1.0, 1.5, and 2.0 m/s carrying loads equal to 0, 10, 20, and 30% body mass positioned symmetrically about the waist. The net metabolic rate increased in nearly direct proportion to the external mechanical power during moderate-speed (0.5-1.5 m/s) load carrying, suggesting that the cost of swinging the legs is relatively small. The active muscle volume required to generate force on the ground and the rate of generating this force accounted for >85% of the increase in net metabolic rate across moderate speeds and most loading conditions. Although these factors explained less of the increase in metabolic rate between 1.5 and 2.0 m/s ( approximately 50%), the cost of generating force per unit volume of active muscle [i.e., the cost coefficient (k)] was similar across all conditions [k = 0.11 +/- 0.03 (SD) J/cm3]. These data indicate that, regardless of the work muscles do, the metabolic cost of walking can be largely explained by the cost of generating muscular force during the stance phase.     

Energy exchange between subject and belt during treadmill walking

Treadmill walking aims to simulate overground walking, but intra-stride belt speed variations of treadmills result in some interaction between treadmill and subject, possibly obstructing this aim. Especially in self-paced treadmill walking, in which the belt speed constantly adjusts to the subject, these interactions might affect the gait pattern significantly. The aim of this study was to quantify the energy exchange between subject and treadmill, during the fixed speed (FS) and self-paced (SP) modes of treadmill walking. Eighteen subjects walked on a dual-belt instrumented treadmill at both modes. The energy exchange was calculated as the integration of the product of the belt speed deviation and the fore-aft ground reaction force over the stride cycle. The total positive energy exchange was 0.44 J/stride and the negative exchange was 0.11 J/stride, which was both less than 1.6% of the performed work on the center of mass. Energy was mainly exchanged from subject to treadmill during both the braking and propulsive phase of gait. The two treadmill modes showed a similar pattern of energy exchange, with a slightly increased energy exchange during the braking phase of SP walking. It is concluded that treadmill walking is only mildly disturbed by subject-belt interactions when using instrumented treadmills with adequate belt control.     

Contributions of the individual ankle plantar flexors to support, forward progression and swing initiation during walking.

Walking is a motor task requiring coordination of many muscles. Previous biomechanical studies, based primarily on analyses of the net ankle moment during stance, have concluded different functional roles for the plantar flexors. We hypothesize that some of the disparities in interpretation arise because of the effects of the uniarticular and biarticular muscles that comprise the plantar flexor group have not been separated. Furthermore, we believe that an accurate determination of muscle function requires quantification of the contributions of individual plantar flexor muscles to the energetics of individual body segments. In this study, we examined the individual contributions of the ankle plantar flexors (gastrocnemius (GAS); soleus (SOL)) to the body segment energetics using a musculoskeletal model and optimization framework to generate a forward dynamics simulation of normal walking at 1.5 m/s. At any instant in the gait cycle, the contribution of a muscle to support and forward progression was defined by its contribution to trunk vertical and horizontal acceleration, respectively, and its contribution to swing initiation by the mechanical energy it delivers to the leg in pre-swing (i.e., double-leg stance prior to toe-off). GAS and SOL were both found to provide trunk support during single-leg stance and pre-swing. In early single-leg stance, undergoing eccentric and isometric activity, they accelerate the trunk vertically but decelerate forward trunk progression. In mid single-leg stance, while isometric, GAS delivers energy to the leg while SOL decelerates it, and SOL delivers energy to the trunk while GAS decelerates it. In late single-leg stance through pre-swing, though GAS and SOL both undergo concentric activity and accelerate the trunk forward while decelerating the downward motion of the trunk (i.e., providing forward progression and support), they execute different energetic functions. The energy produced from SOL accelerates the trunk forward, whereas GAS delivers almost all its energy to accelerate the leg to initiate swing. Although GAS and SOL maintain or accelerate forward motion in mid single-leg stance through pre-swing, other muscles acting at the beginning of stance contribute comparably to forward progression. In summary, throughout single-leg stance both SOL and GAS provide vertical support, in mid single-leg stance SOL and GAS have opposite energetic effects on the leg and trunk to ensure support and forward progression of both the leg and trunk, and in pre-swing only GAS contributes to swing initiation.     

Control of flexor motoneuron activity during single leg walking of the stick insect on an electronically controlled treadwheel

In the present study, motoneurons innervating the flexor tibiae muscle of the stick insect (Cuniculina impigra) middle leg were recorded intracellularly while the single leg performed walking-like movements on a treadwheel. Different levels of belt friction (equivalent to a change in load) were used to study the control of activity of flexor motoneurons. During slow leg movements no fast motoneurons were active, but a recruitment of these neurons could be observed during faster leg movements. The firing rate of slow and fast motoneurons increased with incremented belt friction. Also, the force applied to the treadwheel at different frictional levels was adapted closely to the friction of the treadwheel to be overcome. The motoneurons innervating the flexor tibiae were recruited progressively during the stance phase, with the slow motoneurons being active earlier than the fast (half-maximal spike frequency after 10-15% and 50-60% of the stance phase, respectively). The resting membrane potential was more hyperpolarized in fast motoneurons (64.6 +/- 6.5 mV) than in slow motoneurons (-52.9 +/- 5.4 mV). However, the threshold for the initiation of action potentials was not statistically significantly different in both types of flexor motoneurons. Therefore, action potentials were generated in fast motoneurons after a longer period of depolarization and thus later during the stance phase than in slow motoneurons. We show that motoneurons of the flexor tibiae receive substantial common excitatory inputs during the stance phase and that the difference in resting membrane potential between slow and fast motoneurons is likely to play a crucial role in their consecutive recruitment.