David A. Winter

Human balance and posture control during standing and walking

The common denominator in the assessment of human balance and posture is the inverted pendulum model. If we focus on appropriate versions of the model we can use it to identify the gravitational and acceleration perturbations and pinpoint the motor mechanisms that can defend against any perturbation. We saw that in quiet standing an ankle strategy applies only in the AP direction and that a separate hip load/unload strategy by the hip abd/adductors is the totally dominant defence in the ML direction when standing with feet side by side. In other standing positions (tandem, or intermediate) the two mechanisms still work separately, but their roles reverse. In the tandem position ML balance is an ankle mechanism (invertors/evertors) while in the AP direction a hip load/unloading mechanism dominates. During initiation and termination of gait these two separate mechanisms control the trajectory of the COP to ensure the desired acceleration and deceleration of the COM. During initiation the initial acceleration of the COM forward towards the stance limb is achieved by a posterior and lateral movement of the COP towards the swing limb. After this release phase there is a sudden loading of the stance limb which shifts the COP to the stance limb. The COM is now accelerated forward and laterally towards the future position of the swinging foot. Also ML shifts of the COP were controlled by the hip abductors/adductors and all AP shifts were under the control of the ankle plantar/dorsiflexors. During termination the trajectory of both COM and COP reverse. As the final weight-bearing on the stance foot takes place the COM is passing forward along the medial border of that foot. Hyperactivity of that foots plantarflexors takes the COP forward and when the final foot begins to bear weight the COP moves rapidly across and suddenly stops at a position ahead of the future position of the COM. Then the plantarflexors of both feet release and allow the COP to move posteriorly and approach the COM and meet it as quiet stance is achieved. The inverted pendulum model permitted us to understand the separate roles of the two mechanisms during these critical unbalancing and rebalancing periods. During walking the inverted pendulum model explained the dynamics of the balance of HAT in both the AP and ML directions. Here the model includes the couple due to the acceleration of the weight-bearing hip as well as gravitational perturbations. The exclusive control of AP balance and posture are the hip extensors and flexors, while in the ML direction the dominant control is with the hip abductors with very minor adductor involvement. At the ankle the inverted pendulum model sees the COM passing forward along the medial border to the weight-bearing foot. The model predicts that during single support the body is falling forward and being accelerated medially towards the future position of the swing foot. The model predicts an insignificant role of the ankle invertors/evertors in the ML control. Rather, the future position of the swing foot is the critical variable or more specifically the lateral displacement from the COM at the start of single support. The position is actually under the control of the hip abd/adductors during the previous early swing phase. The critical importance of the hip abductors/adductors in balance during all phases of standing and walking is now evident. This separate mechanism is important from a neural control perspective and clinically it focuses major attention on therapy and potential problems with some surgical procedures. On the other hand the minuscule role of the ankle invertors/evertors is important to note. Except for the tandem standing position these muscles have negligible involvement in balance control.

Kinematic and kinetic patterns in human gait: Variability and compensating effects

Abstract In the presence of fairly well defined kinematic patterns in human walking there was considerable variability at the kinetic level. Intra-subject variability of joint moment patterns over the stride period was high at the knee and hip, but low at the ankle and in a recently defined total limb pattern, called support moment. A similar profile of variability was evident for inter-subject trials at slow, natural and fast cadences, with the percentage variability at the knee and hip decreasing as cadence increases. These moment of force patterns were not random, but were highly correlated. Such a finding points to compensating mechanisms by the biarticulate muscles crossing these joints. Also shown was the fact that these compensating patterns were highly predictable from link segment theory.

Control of whole body balance in the frontal plane during human walking

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.

An integrated biomechanical analysis of normal stair ascent and descent

Three normal males of similar height and weight ascended and descended a five step staircase with a riser height of 22 cm and a tread of 28 cm. EMG, force plate and cine data were collected for the stride over the second to fourth step during each mode. Kinematic and kinetic analyses were integrated with EMG to yield an interpretation of the mechanics of normal stair walking. Movement from one step to the next involved simultaneous lifting and horizontal translation of the body, and each stride showed specific phases for progression. The extensor muscles about the knee played a dominant role in progression from one step to the next in both modes coupled with the ankle plantar flexors. The total lower limb extensor pattern, called the support moment, was highly correlated between subjects and to level walking. Intra- and inter-subject variability of the motor patterns were also determined. The greatest variability was seen at the hip, while stereotypic kinetic patterns emerged at the ankle and knee for all subjects across the 24 trials of each mode.

EMG profiles during normal human walking; stride-to-stride and inter-subject variability

The EMG patterns for 16 muscles involved in human walking are reported along with stride-to-stride and inter-subject variability measures. These profiles and measures were developed for basic researchers and clinical investigators as a baseline reference of motor patterns and for use in the diagnosis of gait pathologies. Evident from a comparison of these patterns were some fundamental aspects of the neuromuscular control and the mechanical demands of walking. These comparisons can be summarized as follows: (1) The distal support muscles (soleus, tibialis anterior, gastrocnemii) are the most active muscles, the more proximal muscles are least active. (2) The least variable EMG patterns, as quantified by the normalized inter-subject variability measures, are seen in the most distal single joint muscles, the most variable are the more proximal muscles. The EMGs of the biarticulate muscles, both proximal and distal, exhibit higher variability than the EMGs of the single joint muscles. (3) The detailed patterns and levels of EMG activity demonstrate the different mechanical tasks of each muscle over the gait cycle.

Biomechanical Motor Patterns in Normal Walking

Motor patterns in normal human gait are evident in several biomechanical and EMG analyses over the stride period. Some of these patterns are invariant over the stride period with changes of cadence, whole others are closely correlated with speed changes. The findings for slow, natural, and fast walking are summarized: 1. Joint angle patterns over the stride period are quite invariant, and do not change with cadence; 2. Moment of force patterns at the ankle are least variable and quite consistent at all speeds; 3. A recently defined support moment is quite consistent at all speeds. 4. Moments at the knee and hip are highly variable at all cadences but decrease their variability as cadence increases; 5. Mechanical power patterns at all joints show consistent timing over the stride period; 6. EMG profiles of 5 muscles show consistent timing over the stride, but the amplitude increases as walking speed increases. Arguments are presented to support the concept that walking speed is largely controlled by gain and that the timing of the motor patterns, which is extremely tightly synchronized with the anatomical position, is under major afferent control.

Biomechanics of below-knee amputee gait.

Sagittal plane biomechanical and EMG analyses from eight below knee (B/K) amputee trials demonstrate considerably modified motor patterns from the residual muscles at the hip and knee. Five SACH fittings, two Uniaxial and one Gressinger prostheses were analysed. Moments of force and mechanical power were analysed on all eight trials and EMG profiles are reported for three of the amputees fitted with SACH prostheses. The findings can be summarized as follows: 1. All eight trials had similar internal moment of force patterns at the ankle. A dorsiflexor moment commenced at heel contact and continued for the first third of stance. The prostheses generated a plantarflexor moment for the balance of stance which increased in late stance to about 2/3 that seen in normals. 2. The two Uniaxial prostheses showed a 20% recovery of stored energy which was returned at push-off. The recovery by the Gressinger fitting was 30%. 3. For all but the Gressinger prosthesis the knee moment of force was negligible during early stance (when normals have an extensor moment), below normal in late stance and fairly normal during swing. The amputee wearing the Gressinger prosthesis had a normal but slightly reduced pattern of moments of force over the entire stride. 4. All eight trials had hyperactive hip extensors during early and mid-stance which resulted in above-normal energy generation by these concentrically contracting muscles. This compensation makes up for the loss of the major energy generation by the plantarflexors at push-off. 5. The moment of force and power patterns at the hip for all eight trials during late stance and swing were fairly normal.(ABSTRACT TRUNCATED AT 250 WORDS)

Energy Generation and Absorption at the Ankle and Knee during Fast, Natural, and Slow Cadences

In 15 normal adults an advanced biomechanical analysis of walking patterns at slow, natural, and fast cadences showed that the ankle had two mechanical power phases: a negative work phase during weight acceptance, followed by a dominant burst of positive work at push-off. The knee had four power phases: a negative work phase at weight acceptance, a small positive work phase during mid-stance, a major negative work burst at push-off and early swing, and a final energy-absorbing phase at the end of swing. The power phases at the hip are quite irregular and somewhat lower than those at the knee and ankle. The dominant positive work burst by the plantarflexors drops as speed decreases, but less rapidly than the positive work by the knee muscles during midstance. The energy absorption by the quadriceps during weight acceptance decreases rapidly as speed decreases and at late stance decreases moderately. The energy absorption by the ankle plantarflexors during weight acceptance remains fairly constant at all walking speeds, and the absorption by the knee flexors at end of swing drops only slightly as cadence decreases.

Kinetic analysis of the lower limbs during walking: What information can be gained from a three-dimensional model?

Kinetic analyses (joint moments, powers and work) of the lower limbs were performed during normal walking to determine what further information can be gained from a three-dimensional model over planar models. It was to be determined whether characteristic moment and power profiles exist in the frontal and transverse planes across subjects and how much work was performed in these planes. Kinetic profiles from nine subjects were derived using a three-dimensional inverse dynamics model of the lower limbs and power profiles were then calculated by a dot product of the angular velocities and joint moments resolved in a global reference system. Characteristic joint moment profiles across subjects were found for the hip, knee and ankle joints in all planes except for the ankle frontal moment. As expected, the major portion of work was performed in the plane of progression since the goal of locomotion is to support the body against gravity while generating movements which propel the body forward. However, the results also showed that substantial work was done in the frontal plane by the hip during walking (23% of the total work at that joint). The characteristic joint profiles suggest defined motor patterns and functional roles in the frontal and transverse planes. Kinetic analysis in three dimensions is necessary particularly if the hip joint is being examined as a substantial amount of work was done in the frontal plane of the hip to control the pelvis and trunk against gravitational forces.

Overall principle of lower limb support during stance phase of gait

The basic function of the lower limb during stance is to resist collapse and to extend sufficiently to achieve the required push-off. Collapse of the lower limb requires a flexion at all three joints (knee, ankle and hip), thus support of the body requires net extensor activity at these joints. A new support moment, Ms, is defined as the algebraic sum of the extensor moments at the three joints, and for 24 subjects, nine patient and three jogging trials Ms was calculated and found to be positive (net extension) over the stance period. Normalizing the peak amplitude of Ms to 100% and averaging these curves over 100% of stance revealed a significant basic pattern. The ensemble average of 12 subjects walking at their natural cadence was very similar to the ensemble average of a mixed group of 24 subjects (walking at fast, natural and slow cadences) and 9 patients. Examination of individual subject and patient joint moment histories revealed considerable variability at the knee and hip in spite of consistent Ms patterns. For example, one knee replacement patient had a moderate knee flexion for the entire stance period but compensated and prevented knee collapse by large hip extension during that time. The three joggers also showed a consistent Ms pattern in the presence of individual variations at the knee and hip; however, the shape of Ms curve had a single peak compared with a double peak for the walking trials.