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The Effect of Constraining Mediolateral Ankle Moments and Foot Placement on the Use of the Counter-Rotation Mechanism During Walking
... Walking without falling requires active stabilization of gait in the mediolateral direction (Bauby and Kuo, 2000). The main mechanism to achieve mediolateral stability during steady-state walking is center-ofpressure control (Hof, 2007;van den Bogaart et al., 2021). Generally, the center-of-pressure is controlled through foot placement (Bruijn and van Dieën, 2018;van Leeuwen et al., 2020) with a complementary role of ankle moment modulation after the foot is placed van Leeuwen et al., 2021). ...
... The main mechanism to achieve mediolateral stability during steady-state walking is center-ofpressure control (Hof, 2007;van den Bogaart et al., 2021). Generally, the center-of-pressure is controlled through foot placement (Bruijn and van Dieën, 2018;van Leeuwen et al., 2020) with a complementary role of ankle moment modulation after the foot is placed van Leeuwen et al., 2021). Both foot placement and ankle moment control are actively driven (Fettrow et al., 2019;Rankin et al., 2014;van Leeuwen et al., 2021;van Leeuwen et al., 2020), and coordinate the center-of-pressure with respect to variations in the center-of-mass kinematic state (Hurt et al., 2010;Wang and Srinivasan, 2014). ...
... Generally, the center-of-pressure is controlled through foot placement (Bruijn and van Dieën, 2018;van Leeuwen et al., 2020) with a complementary role of ankle moment modulation after the foot is placed van Leeuwen et al., 2021). Both foot placement and ankle moment control are actively driven (Fettrow et al., 2019;Rankin et al., 2014;van Leeuwen et al., 2021;van Leeuwen et al., 2020), and coordinate the center-of-pressure with respect to variations in the center-of-mass kinematic state (Hurt et al., 2010;Wang and Srinivasan, 2014). ...
External lateral stabilization can help identify stability control mechanisms during steady-state walking. The degree of step-by-step foot placement control and step width are known to decrease when walking with external lateral stabilization. Here, we investigated the effect of external lateral stabilization on ankle moment control in healthy participants. Ankle moment control complements foot placement, by allowing a corrective center-of-pressure shift once the foot has been placed. This is reflected by a model predicting this center-of-pressure shift based on the preceding foot placement error. Here, the absolute explained variance accounted for by this model decreased when walking with external lateral stabilization. In other words, we found a reduction in the contribution of step-by-step ankle moment control to mediolateral gait stability when externally stabilized. Concurrently, foot placement error and the average center-of-pressure shift remained unchanged.
During steady-state walking, mediolateral gait stability can be maintained by controlling the center of pressure (CoP). The CoP modulates the moment of the ground reaction force, which brakes and reverses movement of the center of mass (CoM) towards the lateral border of the base of support. In addition to foot placement, ankle moments serve to control the CoP. We hypothesized that, during steady-state walking, single stance ankle moments establish a CoP shift to correct for errors in foot placement. We expected ankle muscle activity to be associated with this complementary CoP shift. During treadmill walking, full-body kinematics, ground reaction forces and electromyography were recorded in thirty healthy participants. We found a negative relationship between preceding foot placement error and CoP displacement during single stance; steps that were too medial were compensated for by a lateral CoP shift and vice versa, steps that were too lateral were compensated for by a medial CoP shift. Peroneus longus, soleus and tibialis anterior activity correlated with these CoP shifts. As such, we identified an (active) ankle strategy during steady-state walking. As expected, absolute explained CoP variance by foot placement error decreased when walking with shoes constraining ankle moments. Yet, contrary to our expectations that ankle moment control would compensate for constrained foot placement, the absolute explained CoP variance by foot placement error did not increase when foot placement was constrained. We argue that this lack of compensation reflects the interdependent nature of ankle moment and foot placement control. We suggest that single stance ankle moments do not only compensate for preceding foot placement errors, but also assist control of the subsequent foot placement. Foot placement and ankle moment control are ‘caught’ in a circular relationship, in which constraints imposed on one will also influence the other.
Step-by-step foot placement control, relative to the center of mass (CoM) kinematic state, is generally considered a dominant mechanism for maintenance of gait stability. By adequate (mediolateral) positioning of the center of pressure with respect to the CoM, the ground reaction force generates a moment that prevents falling. In healthy individuals, foot placement is complemented mainly by ankle moment control ensuring stability. To evaluate possible compensatory relationships between step-by-step foot placement and complementary ankle moments, we investigated the degree of (active) foot placement control during steady-state walking, and under either foot placement-, or ankle moment constraints. Thirty healthy participants walked on a treadmill, while full-body kinematics, ground reaction forces and EMG activities were recorded. As a replication of earlier findings, we first showed step-by-step foot placement is associated with preceding CoM state and hip ab-/adductor activity during steady-state walking. Tight control of foot placement appears to be important at normal walking speed because there was a limited change in the degree of foot placement control despite the presence of a foot placement constraint. At slow speed, the degree of foot placement control decreased substantially, suggesting that tight control of foot placement is less essential when walking slowly. Step-by-step foot placement control was not tightened to compensate for constrained ankle moments. Instead compensation was achieved through increases in step width and stride frequency.
During human walking, the centre of mass (CoM) is outside the base of support for most of the time, which poses a challenge to stabilizing the gait pattern. Nevertheless, most of us are able to walk without substantial problems. In this review, we aim to provide an integrative overview of how humans cope with an underactuated gait pattern. A central idea that emerges from the literature is that foot placement is crucial in maintaining a stable gait pattern. In this review, we explore this idea; we first describe mechanical models and concepts that have been used to predict how foot placement can be used to control gait stability. These concepts, such as for instance the extrapolated CoM concept, the foot placement estimator concept and the capture point concept, provide explicit predictions on where to place the foot relative to the body at each step, such that gait is stabilized. Next, we describe empirical findings on foot placement during human gait in unperturbed and perturbed conditions. We conclude that humans show behaviour that is largely in accordance with the aforementioned concepts, with foot placement being actively coordinated to body CoM kinematics during the preceding step. In this section, we also address the requirements for such control in terms of the sensory information and the motor strategies that can implement such control, as well as the parts of the central nervous system that may be involved. We show that visual, vestibular and proprioceptive information contribute to estimation of the state of the CoM. Foot placement is adjusted to variations in CoM state mainly by modulation of hip abductor muscle activity during the swing phase of gait, and this process appears to be under spinal and supraspinal, including cortical, control. We conclude with a description of how control of foot placement can be impaired in humans, using ageing as a primary example and with some reference to pathology, and we address alternative strategies available to stabilize gait, which include modulation of ankle moments in the stance leg and changes in body angular momentum, such as rapid trunk tilts. Finally, for future research, we believe that especially the integration of consideration of environmental constraints on foot placement with balance control deserves attention.
Lateral balance is a critical factor in keeping the human body upright during walking. Two important mechanisms for balance control are the stepping strategy, in which the foot placement is changed in the direction of a sensed fall to modulate how the gravitational force acts on the body, and the lateral ankle strategy, in which the body mass is actively accelerated by an ankle torque. Currently, there is minimal evidence about how these two strategies complement one another to achieve upright balance during locomotion. We use Galvanic vestibular stimulation (GVS) to induce the sensation of a fall at heel-off during gait initiation. We found that young healthy adults respond to the illusory fall using both the lateral ankle strategy and the stepping strategy. The stance foot center of pressure (CoP) is shifted in the direction of the perceived fall by ≈2.5 mm, starting ≈247 ms after stimulus onset. The foot placement of the following step is shifted by ≈15 mm in the same direction. The temporal delay between these two mechanisms suggests that they independently contribute to upright balance during locomotion, potentially in a serially coordinated manner. Modeling results indicate that without the lateral ankle strategy, a much larger step width is required to maintain upright balance, suggesting that the small but early CoP shift induced by the lateral ankle strategy is critical for upright stability during locomotion. The relative importance of each mechanism and how neurological disorders may affect their implementation remain an open question.
During human walking, perturbations to the upper body can be partly corrected by placing the foot appropriately on the next step. Here, we infer aspects of such foot placement dynamics using step-to-step variability over hundreds of steps of steady-state walking data. In particular, we infer dependence of the 'next' foot position on upper body state at different phases during the 'current' step. We show that a linear function of the hip position and velocity state (approximating the body center of mass state) during mid-stance explains over 80% of the next lateral foot position variance, consistent with (but not proving) lateral stabilization using foot placement. This linear function implies that a rightward pelvic deviation during a left stance results in a larger step width and smaller step length than average on the next foot placement. The absolute position on the treadmill does not add significant information about the next foot relative to current stance foot over that already available in the pelvis position and velocity. Such walking dynamics inference with steady-state data may allow diagnostics of stability and inform biomimetic exoskeleton or robot design.
Background
Whole-body angular momentum (H) influences fall risk, is tightly regulated during walking, and is primarily controlled by muscle force generation. People with transtibial amputations using passive-elastic prostheses typically have greater H compared with nonamputees.
Questions/purposes
(1) Do people with unilateral transtibial amputations using passive-elastic prostheses have greater sagittal and frontal plane H ranges of motion during walking compared with nonamputees and compared with using powered prostheses? (2) Does use of powered ankle-foot prostheses result in equivalent H ranges in all planes of motion compared with nonamputees during walking as a result of normative prosthetic ankle power generation?
Methods
Eight patients with a unilateral transtibial amputation and eight nonamputees walked 0.75, 1.00, 1.25, 1.50, and 1.75 m/s while we measured kinematics and ground reaction forces. We calculated H for participants using their passive-elastic prosthesis and a powered ankle-foot prosthesis and for nonamputees at each speed.
Results
Patients using passive-elastic prostheses had 32% to 59% greater sagittal H ranges during the affected leg stance phase compared with nonamputees at 1.00 to 1.75 m/s (p
Falling poses a major threat to the steadily growing population of the elderly in modern-day society. A major challenge in the prevention of falls is the identification of individuals who are at risk of falling owing to an unstable gait. At present, several methods are available for estimating gait stability, each with its own advantages and disadvantages. In this paper, we review the currently available measures: the maximum Lyapunov exponent (λS and λL), the maximum Floquet multiplier, variability measures, long-range correlations, extrapolated centre of mass, stabilizing and destabilizing forces, foot placement estimator, gait sensitivity norm and maximum allowable perturbation. We explain what these measures represent and how they are calculated, and we assess their validity, divided up into construct validity, predictive validity in simple models, convergent validity in experimental studies, and predictive validity in observational studies. We conclude that (i) the validity of variability measures and λS is best supported across all levels, (ii) the maximum Floquet multiplier and λL have good construct validity, but negative predictive validity in models, negative convergent validity and (for λL) negative predictive validity in observational studies, (iii) long-range correlations lack construct validity and predictive validity in models and have negative convergent validity, and (iv) measures derived from perturbation experiments have good construct validity, but data are lacking on convergent validity in experimental studies and predictive validity in observational studies. In closing, directions for future research on dynamic gait stability are discussed.
We studied the extent to which automatic postural actions in standing human subjects are organized by a limited repertoire of central motor programs. Subjects stood on support surfaces of various lengths, which forced them to adopt different postural movement strategies to compensate for the same external perturbations. We assessed whether a continuum or a limited set of muscle activation patterns was used to produce different movement patterns and the extent to which movement patterns were influenced by prior experience. Exposing subjects standing on a normal support surface to brief forward and backward horizontal surface perturbations elicited relatively stereotyped patterns of leg and trunk muscle activation with 73- to 110-ms latencies. Activity began in the ankle joint muscles and then radiated in sequence to thigh and then trunk muscles on the same dorsal or ventral aspect of the body. This activation pattern exerted compensatory torques about the ankle joints, which restored equilibrium by moving the body center of mass forward or backward. This pattern has been termed the ankle strategy because it restores equilibrium by moving the body primarily around the ankle joints. To successfully maintain balance while standing on a support surface short in relation to foot length, subjects activated leg and trunk muscles at similar latencies but organized the activity differently. The trunk and thigh muscles antagonistic to those used in the ankle strategy were activated in the opposite proximal-to-distal sequence, whereas the ankle muscles were generally unresponsive. This activation pattern produced a compensatory horizontal shear force against the support surface but little, if any, ankle torque. This pattern has been termed the hip strategy, because the resulting motion is focused primarily about the hip joints. Exposing subjects to horizontal surface perturbations while standing on support surfaces intermediate in length between the shortest and longest elicited more complex postural movements and associated muscle activation patterns that resembled ankle and hip strategies combined in different temporal relations. These complex postural movements were executed with combinations of torque and horizontal shear forces and motions of ankle and hip joints. During the first 5-20 practice trials immediately following changes from one support surface length to another, response latencies were unchanged. The activation patterns, however, were complex and resembled the patterns observed during well-practiced stance on surfaces of intermediate lengths.(ABSTRACT TRUNCATED AT 400 WORDS)
The balance of standing humans is usually explained by the inverted pendulum model. The subject invokes a horizontal ground-reaction force in this model and controls it by changing the location of the centre of pressure under the foot or feet. In experiments I showed that humans are able to stand on a ridge of only a few millimetres wide on one foot for a few minutes. In the present paper I investigate whether the inverted pendulum model is able to explain this achievement. I found that the centre of mass of the subjects sways beyond the surface of support, rendering the inverted pendulum model inadequate. Using inverse simulations of the dynamics of the human body, I found that hip-joint moments of the stance leg are used to vary the horizontal component of the ground-reaction force. This force brings the centre of mass back over the surface of support. The subjects generate moments of force at the hip-joint of the swing leg, at the shoulder-joints and at the neck. These moments work in conjunction with a hip strategy of the stance leg to limit the angular acceleration of the head-arms-trunk complex. The synchrony of the variation in moments suggests that subjects use a motor programme rather than long latency reflexes.
The equations of motion for a standing multi-segment human model are derived. Output quantity of these equations is the horizontal acceleration of the whole-body centre of mass (CoM). There are three input terms and they can be identified as the three mechanisms by which balance can be maintained: (1) by moving the centre of pressure with respect to the vertical projection of the CoM, (2) by counter-rotating segments around the CoM, and (3) by applying an external force, other than the ground reaction force. For the first two mechanisms the respective contributions to CoM acceleration can be obtained from force plate recordings. This is illustrated by some example data from experiments, which show that the contribution from mechanism 2 can be considerable, e.g. in one-legged standing.
In walking the human body is never in balance. Most of the time the trunk is supported by one leg and the centre of mass (CoM) 'falls' to the contralateral side. In dynamical situations the velocity of the CoM should be acknowledged as well in the 'extrapolated centre of mass' (XcoM). Centre of pressure (CoP) position was recorded by a treadmill with built-in force transducers. Lateral CoM and XcoM position were computed by filtering the CoP data. Subjects were six above-knee amputees and six matched healthy controls. They walked at approximately 0.75, 1, and 1.25m/s for 2min. Amputees showed asymmetric gait with shorter stance (60%) at the prosthetic side versus 68% at the non-prosthetic side and a wider stride (13+/-4cm, mean+/-S.D.) compared to controls (9+/-3cm). At foot placement CoP was just lateral to the XcoM. The margin between average CoP and XcoM at foot contact was only 1.6+/-0.7cm in controls, 2.7+/-0.5cm in amputees at the prosthetic side and 1.9+/-0.6cm at the non-prosthetic side. Next to this 'stepping strategy', CoP position was corrected after initial contact by modulating the lateral foot roll-off ('lateral ankle strategy') in non-prosthetic legs up to about 2cm. A simple mechanical model, the inverted pendulum model, can explain that: (1) a less precise foot placement (greater CoP-XcoM margin) results in a wider stride, (2) this effect can be reduced by walking with a higher cadence, and (3) a greater margin at one side, as with a leg prosthesis, should be compensated by a shorter stance duration at the same side to achieve a straight path. This suggests that not in all cases symmetric gait should be an aim of rehabilitation.
Shifts of the center of pressure (CoP) through modulation of foot placement and ankle moments (CoP-mechanism) cause accelerations of the center of mass (CoM) that can be used to stabilize gait. An additional mechanism that can be used to stabilize gait, is the counter-rotation mechanism, i.e., changing the angular momentum of segments around the CoM to change the direction of the ground reaction force. The relative contribution of these mechanisms to the control of the CoM is unknown. Therefore, we aimed to determine the relative contribution of these mechanisms to control the CoM in the anteroposterior (AP) direction during a normal step and the first recovery step after perturbation in healthy adults. Nineteen healthy subjects walked on a split-belt treadmill and received unexpected belt acceleration perturbations of various magnitudes applied immediately after right heel-strike. Full-body kinematic and force plate data were obtained to calculate the contributions of the CoP-mechanism and the counter-rotation mechanism to control the CoM. We found that the CoP-mechanism contributed to corrections of the CoM acceleration after the AP perturbations, while the counter-rotation mechanism actually counteracted the CoM acceleration after perturbation, but only in the initial phases of the first step after the perturbation. The counter-rotation mechanism appeared to prevent interference with the gait pattern, rather than using it to control the CoM after the perturbation. Understanding the mechanisms used to stabilize gait may have implications for the design of therapeutic interventions that aim to decrease fall incidence.
Lower-limb amputees have a higher risk of falling compared to non-amputees. Proper regulation of whole-body angular momentum is necessary to prevent falls, particularly in the frontal plane where individuals are most unstable. However, the balance recovery mechanisms used by lower-limb amputees when recovering from a perturbation are not well-understood. This study sought to understand the balance recovery mechanisms used by lower-limb amputees in response to mediolateral foot perturbations by examining changes to frontal plane whole-body angular momentum and hip joint work. These metrics provide a quantitative measure of frontal plane dynamic balance and associated joint contributions required to maintain balance during gait. Nine amputees and 11 non-amputees participated in this study where an unexpected medial or lateral foot placement perturbation occurred immediately prior to heel strike on the residual, sound or non-amputee limbs. Lateral perturbations of all limbs resulted in a reduced range of whole-body angular momentum and increased positive frontal plane hip work in the first half of single limb support. Medial perturbations for all limbs resulted in increased range of whole-body angular momentum and decreased positive frontal plane hip work, also in the first half of single limb support. These results suggest that medial foot placement perturbations are particularly challenging and that hip strategies play an important role in balance recovery. Thus, rehabilitation interventions that focus on hip muscles that regulate mediolateral balance, particularly the hip abductors, and the use of prostheses with active ankle control, may reduce the risk of falls.
Humans perform a variety of feedback adjustments to maintain balance during walking. These include lateral footfall placement, and center of pressure adjustment under the stance foot, to stabilize lateral balance. A less appreciated possibility would be to steer for balance like a bicycle, whose front wheel may be turned toward the direction of a lean to capture the center of mass. Humans could potentially combine steering with other strategies to distribute balance adjustments across multiple degrees of freedom. We tested whether human balance can theoretically benefit from steering, and experimentally tested for evidence of steering for balance. We first developed a simple dynamic walking model, which shows that bipedal walking may indeed be stabilized through steering—externally rotating the foot about vertical toward the direction of lateral lean for each footfall—governed by linear feedback control. Moreover, least effort (mean-square control torque) is required if steering is combined with lateral foot placement. If humans use such control, footfall variability should show a statistical coupling between external rotation with lateral placement. We therefore examined the spontaneous fluctuations of hundreds of strides of normal overground walking in healthy adults (N=26). We found significant coupling ( ), of 0.54 rad of external rotation per meter of lateral foot deviation. Successive footfalls showed a weaker, negative correlation with each other, similar to how a bicycle's steering adjustment made for balance must be followed by gradual corrections to resume the original travel direction. Steering may be one of multiple strategies to stabilize balance during walking.
Over 50% of individuals with lower limb amputation fall at least once each year. These individuals also exhibit reduced ability to effectively respond to challenges to frontal plane stability. The range of whole body angular momentum has been correlated with stability and fall risk. This study determined how lateral walking surface perturbations affected the regulation of whole body and individual leg angular momentum in able-bodied controls and individuals with unilateral transtibial amputation. Participants walked at fixed speed in a Computer Assisted Rehabilitation Environment with no perturbations and continuous, pseudo-random, mediolateral platform oscillations. Both the ranges and variability of angular momentum for both the whole body and both legs were significantly greater (p<0.001) during platform oscillations. There were no significant differences between groups in whole body angular momentum range or variability during unperturbed walking. The range of frontal plane angular momentum was significantly greater for those with amputation than for controls for all segments (p<0.05). For the whole body and intact leg, angular momentum ranges were greater for patients with amputation. However, for the prosthetic leg, angular momentum ranges were less for patients than controls. Patients with amputation were significantly more affected by the perturbations. Though patients with amputation were able to maintain similar patterns of whole body angular momentum during unperturbed walking, they were more highly destabilized by the walking surface perturbations. Individuals with transtibial amputation appear to predominantly use altered motion of the intact limb to maintain mediolateral stability.
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Individuals with a unilateral transtibial amputation have a greater risk of falling compared to able-bodied individuals, and falling on stairs can lead to serious injuries. Individuals with transtibial amputations have lost ankle plantarflexor muscle function, which is critical for regulating whole-body angular momentum to maintain dynamic balance. Recently, powered prostheses have been designed to provide active ankle power generation with the goal of restoring biological ankle function. However, the effects of using a powered prosthesis on the regulation of whole-body angular momentum are unknown. The purpose of this study was to use angular momentum to evaluate dynamic balance in individuals with a transtibial amputation using powered and passive prostheses relative to able-bodied individuals during stair ascent and descent. Ground reaction forces, external moment arms, and joint powers were also investigated to interpret the angular momentum results. A key result was that individuals with an amputation had a larger range of sagittal-plane angular momentum during prosthetic limb stance compared to able-bodied individuals during stair ascent. There were no significant differences in the frontal, transverse, or sagittal-plane ranges of angular momentum or maximum magnitude of the angular momentum vector between the passive and powered prostheses during stair ascent or descent. These results indicate that individuals with an amputation have altered angular momentum trajectories during stair walking compared to able-bodied individuals, which may contribute to an increased fall risk. The results also suggest that a powered prosthesis provides no distinct advantage over a passive prosthesis in maintaining dynamic balance during stair walking.
Stability is an important concern during human walking, and can limit mobility in clinical populations. Mediolateral stability can be efficiently controlled through appropriate foot placement, although the underlying neuromechanical strategy is unclear. We hypothesized that humans control mediolateral foot placement through swing leg muscle activity, basing this control on the mechanical state of the contralateral stance leg. Participants walked under Unperturbed and Perturbed conditions, in which foot placement was intermittently perturbed by moving the right leg medially or laterally during the swing phase (by ~50-100 mm). We quantified mediolateral foot placement, electromyographic activity of frontal plane hip muscles, and stance leg mechanical state. During Unperturbed walking, greater swing phase gluteus medius (GM) activity was associated with more lateral foot placement. Increases in GM activity were most strongly predicted by increased mediolateral displacement between the center of mass (CoM) and contralateral stance foot. The Perturbed walking results indicated a causal relationship between stance leg mechanics and swing phase GM activity. Perturbations which reduced the mediolateral CoM displacement from the stance foot caused reductions in swing phase GM activity and more medial foot placement. Conversely, increases in mediolateral CoM displacement caused increased swing phase GM activity and more lateral foot placement. Under both Unperturbed and Perturbed conditions, humans controlled their mediolateral foot placement by modulating swing phase muscle activity in response to the mechanical state of the contralateral leg. This strategy may be disrupted in clinical populations with a reduced ability to modulate muscle activity or sense their body's mechanical state.
Gait research and clinical gait training may benefit from movement-dependent event control, that is, technical applications in which events such as obstacle appearance or visual/acoustic cueing are (co)determined online on the basis of current gait properties. A prerequisite for successful gait-dependent event control is accurate online detection of gait events such as foot contact (FC) and foot off (FO). The objective of the present study was to assess the feasibility of online FC and FO detection using a single large force platform embedded in a treadmill. Center-of-pressure, total force output and kinematic data were recorded simultaneously in 12 healthy participants. Online FC and FO estimates and spatial and temporal gait parameters estimated from the force platform data--i.e., center-of-pressure profiles--were compared to offline kinematic counterparts, which served as the gold standard. Good correspondence was achieved between online FC detections using center-of-pressure profiles and those derived offline from kinematic data, whereas FO was detected 31 ms too late. A good relative and absolute agreement was achieved for both spatial and temporal gait parameters, which was improved further by applying more fine-grained FO estimation procedures using characteristic local minima in the total force output time series. These positive results suggest that the proposed system for gait-dependent event control may be successfully implemented in gait research as well as gait interventions in clinical practice.
Unilateral, below-knee amputees have an increased risk of falling compared to non-amputees. The regulation of whole-body angular momentum is important for preventing falls, but little is known about how amputees regulate angular momentum during walking. This study analyzed three-dimensional, whole-body angular momentum at four walking speeds in 12 amputees and 10 non-amputees. The range of angular momentum in all planes significantly decreased with increasing walking speed for both groups. However, the range of frontal-plane angular momentum was greater in amputees compared to non-amputees at the first three walking speeds. This range was correlated with a reduced second vertical ground reaction force peak in both the intact and residual legs. In the sagittal plane, the amputee range of angular momentum in the first half of the residual leg gait cycle was significantly larger than in the non-amputees at the three highest speeds. In the second half of the gait cycle, the range of sagittal-plane angular momentum was significantly smaller in amputees compared to the non-amputees at all speeds. Correlation analyses suggested that the greater range of angular momentum in the first half of the amputee gait cycle is associated with reduced residual leg braking and that the smaller range of angular momentum in the second half of the gait cycle is associated with reduced residual leg propulsion. Thus, reducing residual leg braking appears to be a compensatory mechanism to help regulate sagittal-plane angular momentum over the gait cycle, but may lead to an increased risk of falling.
A whole-body inverted pendulum model was used to investigate the control of balance and posture in the frontal plane during human walking. The model assessed the effects of net joint moments, joint accelerations and gravitational forces acting about the supporting foot and hip. Three video cameras and two force platforms were used to collect kinematic and kinetic data from repeat trials on four subjects during natural walking. An inverse solution was used to calculate net joint moments and powers. Whole body balance was ensured by the centre of mass (CM) passing medial to the supporting foot, thus creating a continual state of dynamic imbalance towards the centerline of the plane of progression. The medial acceleration of the CM was primarily generated by a gravitational moment about the supporting foot, whose magnitude was established at initial contact by the lateral placement of the new supporting foot relative to the horizontal location of the CM. Balance of the trunk and swing leg about the supporting hip was maintained by an active hip abduction moment, which recognized the contribution of the passive accelerational moment, and countered a large destabilizing gravitational moment. Posture of the upper trunk was regulated by the spinal lateral flexors. Interactions between the supporting foot and hip musculature to permit variability in strategies used to maintain balance were identified. Possible control strategies and muscle activation synergies are discussed.
We measured variability of foot placement during gait to test whether lateral balance must be actively controlled against dynamic instability. The hypothesis was developed using a simple dynamical model that can walk down a slight incline with a periodic gait resembling that of humans. This gait is entirely passive except that it requires active control for a single unstable mode, confined mainly to lateral motion. An especially efficient means of controlling this instability is to adjust lateral foot placement. We hypothesized that similar active feedback control is performed by humans, with fore-aft dynamics stabilized either passively or by very low-level control. The model predicts that uncertainty within the active feedback loop should result in variability in foot placement that is larger laterally than fore-aft. In addition, loss of sensory information such as by closing the eyes should result in larger increases in lateral variability. The control model also predicts a slight coupling between step width and length. We tested 15 young normal human subjects and found that lateral variability was 79% larger than fore-aft variability with eyes open, and a larger increase in lateral variability (53% vs. 21%) with eyes closed, consistent with the model's predictions. We also found that the coupling between lateral and fore-aft foot placements was consistent with a value of 0.13 predicted by the control model. Our results imply that humans may harness passive dynamic properties of the limbs in the sagittal plane, but must provide significant active control in order to stabilize lateral motion.
Assessing the stability of human
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Control of lateral balance in walking
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Active foot placement control ensures stable gait: Effect of constraints on foot placement and ankle moments
- M Van Den Bogaart