Compensatory Stepping: The Biomechanics of a ...

Experimental Aging Research, 27: 361±376, 2001 Copyright # 2001 Taylor & Francis 0361-073X /01 $12.00 + .00

Compensatory Stepping: The Biomechanics of a Preferred Response Among Older Adults Jody L. Jensen Department of Kinesiology and Health Education, The University of Texas at Austin, Austin, Texas, USA

Lesley A. Brown Department of Kinesiology, University of Lethbridge, Lethbridge, Alberta, Canada

Marjorie H. Woollacott Department of Exercise and Movement Science, University of Oregon, Eugene, Oregon, USA The purpose of this study was to evaluate age-related differences in the mechanics of the compensatory stepping response to balance threats. A moving platform was used to disturb the balance of 16 younger (21 to 35 years) and 19 older (68 to 88 years) adults. Backward platform translations consisted of 15-cm displacements with peak accelerations ranging from 9.4 to 15.2 m=s2 . Older adults were more likely to use a step to recover balance and stepped at lower perturbation magnitudes than younger adults. Group differences were not found in time to step initiation or segmental momentum. The lack of group differences in momentum revealed that lower perturbation accelerations created an equivalent or greater magnitude of body motion in older adults compared to higher Received 22 June 2000; accepted 15 March 2001. The authors gratefully acknowledg e Mr. Dan Furgason (Department of Physics, University of Lethbridge) for programming contributions and Mr. Ryan Sleik (Department of Kinesiology, University of Lethbridge) for manuscript formatting. This research was supported by a National Institutes of Health Grant (AG05317 ) to M. Woollacott and a grant from the American Federation for Aging Research to J. L. Jensen. L. A. Brown was also supported by grants from the Alberta Heritage Foundation for Medical Research. Address correspondenc e to Jody L. Jensen, PhD, Department of Kinesiology and Health Education, Bellmont Hall 222, The University of Texas at Austin, Austin, TX 78712, USA. E-mail: [email protected]

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accelerations experienced by younger adults. Older adults also showed a reduced ability to attenuate the input acceleration and experienced significantly greater linear acceleration of the head. Loss of balance and falling is an unfortunate reality for many older adults. Why we fall and what we can do about reducing the incidence of falling are central themes in prevention and rehabilitation research arenas. One approach to understanding the propensity for instability among the aged is to explore how the strategies for balance recovery among older adults compare with those of younger adults who are not limited by balance dysfunction. For example, it has become a common observation that older adults adopt a compensatory step more readily than younger adults (Brown, Shumway-Cook, & Woollacott, 1999; McIlroy & Maki, 1996; Pai, Rogers, Patton, Cain, & Hanke, 1998; Luchies, Alexander, Schultz, & Ashton-Miller, 1994; Rogers, Hain, Hanke, & Janssen, 1996). It has also been noted that older adults adopt steps following less severe balance disturbances than younger adults (Pai et al., 1998; Brown et al., 1999). Given the prevalence with which older adults utilize compensatory stepping to prevent imbalance, it is surprising that the basis for this apparent preference remains largely undetermined. Here, we examine age-related differences in the compensatory stepping response to examine the possibilit y that the preference for stepping among older adults is mechanically based. The occurrence of a step typically has been regarded as an inevitable consequence of the horizontal movement of the center of mass (COM) beyond the limits of the base of support (BOS). Although steps of this type are considered to be necessary outcomes of the laws of physics, recent work implies that in both young and older adults, steps are also initiated when the COM is located within the base of support (Brown et al., 1999; McIlroy & Maki, 1993; Pai et al., 1998). Thus the occurrence of a compensatory step does not depend solely on the COM exceeding the base of support. Indeed recent work by Pai and colleagues (1998, 1997) demonstrated that the velocity of the COM is also a trigger for compensatory stepping. This work has led to the conceptualization of the ``dynamic stability limits’’ for feasible movement termination, which predicted that the occurrence of a step depends on the interaction between the COM position and its velocity. Following this model, stepping may be necessary if there is a sufficiently high velocity of COM displacement, even if the COM is located within the BOS at step initiation . Thus, if older adults adopt steps following less severe disturbances than younger adults, and if these steps are initiated before the COM exceeds the BOS, it would follow that a less intense balance disturbanc e would lead to an equivalent (or higher) COM velocity for older adults than for younger adults. Thus, one issue that remains to be explored is how a less severe perturbation might create the momentum (Mass £ Velocity) in an older adult that equates with that obtained following a larger perturbation experienced by a younger adult.

Age Differences in Rapid Step Initiation

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Insight into this issue may be gained by considering the momentum possessed by the body prior to the initiation of the step. When we experience a perturbation, the body is set in motion, and there is a change in momentum. In induced-momentum paradigms, momentum is ultimately constrained by the generation of muscle torques to reduce segmental motion, and hence the linear momentum of the whole body. One hypothesis for the increased incidence of stepping among older adults is that older adults may be unable to generate the necessary counterbalancing muscle torques, either in magnitude or in rate of muscle torque development, to control the body’s horizontal momentum. Evidence in support of this hypothesis comes from a number of studies showing an age-related reduction in voluntary muscle strength (Brown et al., 1999; McIlroy & Maki, 1993; Pai et al., 1998) and rate of muscle force production (Clarkson, Kroll, & Melchionda, 1981; Hakkinen & Hakkinen, 1991; Larsson, 1978). However, the results from voluntary muscle testing in the construction of explanations of function in reactive balance tasks should be viewed with some caution. Two cautionary points can be made. First, the initial response to a balance threat may not require a maximal muscle strength response (Gu, Schultz, Shepard, & Alexander, 1996). Secondly, there is evidence that the initial ankle muscle torque response to balance perturbation s (e.g., up to approximately 160 ms post perturbation ) does not differ between younger and older adults (Hall, Woollacott, & Jensen, 1999). These findings imply that the increased incidence of stepping among older adults may not be a function of limitations in initial muscle torque production . Although Hall et al. (1999) found the initial balance recovery response at the ankle to be no different between younger and older adults, the possibilit y remains that the differences may be found in how the motion induced by the perturbation is propagated up the linked system. The purpos e of this study was to determine if there are age-related differences in the mechanics of the balance response leading to step initiation . Our analyses were directed toward evaluating age-related differences in segmental momenta and accelerations at the time of step initiation.

METHOD Participants The data from 35 females were analyzed in this study. There were 16 younger adults and 19 older adults. The age of the younger participants ranged between 21 and 35 years (M ˆ 25.93, SD ˆ 4.62 years); the age of the older participants ranged between 68 and 88 years (M ˆ 72.1, SD ˆ 3.78 years). To qualify for participation in this study, all participants were screened to ensure that they were free from any orthopedic, neurological, and=or cardiovascular disorders; the older adults were required to receive medical

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clearance to participate in the study. In addition, only those individuals who had not experienced any fall episodes within the past year and did not have a fear of falling (participant self-report) were invited to participate. All participants were informed of the study protocol and indicated their willingness to participate through written consent. Study participation was restricted to females to control for potentially confounding gender-specific differences such as strength (Thelen, Schultz, & Alexander, 1996; Wolfson, Whipple, Derby, Amerman, & Nashner, 1994).

Protocol Balance was disturbed using a translating force platform (mounted flush with the surrounding surface). Each participant stood barefoot with arms folded across the chest. They wore a lightweight climbing harness secured overhead to ensure safety in the event of a stumble, misstep, or fall. Participants were asked to try to maintain their balance without moving their feet (i.e., steppingÐthe stepping area was unconstrained). A series of 18 consecutive perturbations were delivered using both backward- and forwarddirected platform displacements, ranging between 5 and 15 cm, with velocities ranging between 10 and 80 cm=s. Perturbation s were presented in a pseudorandomize d manner whereby less severe disturbances occurred within the early trials. This testing order served as a precautionary measure to ensure that participants would not be exposed to the more severe disturbances early in the test session. From these 18 trials, three backward-directed perturbation conditions were selected for analysis: 15 cm at 40 cm=s; 15 cm at 60 cm=s; 15 cm at 80 cm=s. In our previous work, these perturbation conditions were shown to elicit compensatory stepping in older and in younger adults (Brown et al., 1999). These perturbation s represent a progressive increase in severity, presenting peak accelerations of 9.4, l2.5, and 15.2 m=s2 for the 40-, 60-, and 80-cm=s conditions, respectively. In all cases, the platform was translated using a ramp onset with parabolic offset displacement waveform. Figure 1 illustrates the displacement waveform and resulting acceleration=deceleration characteristics for the mid-severity condition (peak acceleration ˆ 12.5 m=s2 ). Data were collected over a 5-s interval with platform translations programmed to commence 1 s from the start of data collection. Data collection was triggered after participants indicated that they were ready, though a variable wait period (approximately 1±3 s) was used so that participants could not predict when the balance disturbanc e would occur. In addition, the three trials used in this analysis represented trials numbered 27 (peak ˆ 9.4 m=s2 ), 21 (12.5 m=s2 ), and 18 (15.2 m=s2 ). Each trial was preceded by a different perturbation condition, making anticipation of the precise perturbation characteristics unlikely.


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FIGURE 1 Platform displacement and acceleration for the mid-severity condition of 15 cm at 60 cm=s condition (12.5 m=s2 ), showing the ramp onset and parabolic offset displacement waveform. NOTE: The perturbation conditions presented peak accelerations of 9.4 m=s2 , 12.5 m=s2 , and 15.2 m=s2 for the 40, 60, and 80 cm=s conditions respectively.

Instrumentation A hydraulically driven, strain-gauge force plate (30 £ 46 cm; University of Oregon, Institute of Neuroscience technical group) was used in this study. Analog outputs from a linear potentiometer, vertical and horizontal ground reaction forces, and the accompanying moment of force data were digitally sampled at 500 Hz. Kinematic data were obtained from a two-camera WATSMART motion analysis system (Northern Digital, Inc., Waterloo, Ontario). Light-emittin g diodes marked the following landmarks: (1) the head of the fifth metatarsal (`toe’), (2) the lateral calcaneus (`heel’), (3) the lateral malleolus (`ankle’), (4) the lateral femoral condyle (`knee’), (5) the greater trochanter (`hip’). (6) the acromion process (`shoulder’) , (7) the temporomandibula r joint (`ear’), (8) the temple (`head’), and (9) the platform. A calibrated volume measuring 1.7 £ 1.4 £ .55 m was defined. Average error (root mean square) for the reconstruction of markers was less than 5 mm. Kinematic data were sampled at 100 Hz.

Data Analyses The presence of a step was determined from video data and confirmed using center of pressure (COP) coordinates derived from the recorded vertical

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ground reaction force and moment of force data obtained from the force plate. Step initiation was defined as the first deviation beyond the mean in the lateral COP signal. The criterion for deviation was a consistent change in slope over a 20-ms interval, leading to lift-off of the swing limb. The body was modeled as a bilaterally symmetrical five-segment system composed of feet, shanks, thighs, trunk (with arms), and head. Definitions of segment length proportion s and inertial characteristics were based on anthropometric data reported by Winter (1990). Kinematic data were referenced to an inertial reference frame to compensate for plate motion. The location of the body’s COM relative to the base of support in the anterior-posterior plane was used to describe the type of step response that subjects used to regain balance. In particular, we were interested in quantifyin g the prevalence of steps that do not occur as a consequence of the COM exceeding the base of support. We defined the functiona l base of support as a distance equivalent to 80% of the length of the foot, measured from the heel. If the horizontal location of the whole body COM was found to be within this 80% boundary at the point of step initiation , then the step was regarded to occur while the COM remained within the BOS. Linear momentum (L) values at step initiation (for anterior-posterior plane motion) were determined for segmental and the whole body COM (L is the product of a body’s mass [m, segmental or whole body] and its linear velocity [v]; L ˆ mv). Angular momentum (H) was calculated for the trunk segment only (H is the product of a body’s inertia (I ) and its angular velocity (o); H ˆ Io). Trunk angular momentum (HTK ) was used to categorize those steps occurring while the COM remained within the BOS to be either `delayed’ or `rapid.’ Steps initiated after HTK had reached its maximum were categorized as `delayed’ steps (Figure 2, Graphs A and C); steps initiated before maximum HTK were classified as `rapid’ steps (Figure 2, Graphs B and D). Momentum represents the quantity of motion of a system at an instant in time. We were also interested in how the system was changing at the time of step initiation . Thus, we evaluated the horizontal linear acceleration (rate of change of velocity) of body segments relative to the acceleration of the toe to determine how the velocity (thus the momentum) of the system was changing at the time of step initiation . We compared the motion of the body segments across different perturbations by normalizing the magnitude of body landmark linear accelerations with respect to the peak acceleration associated with each perturbation .

Statistical Analyses The frequency of stepping was compared between younger and older adults using a chi-square test of association. Age differences in the measures of interest were assessed using independent samples t tests with alpha set to 0.05; a Bonferroni correction for multiple comparisons was used when appropriate.

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FIGURE 2 Graphs A and B show COM trajectories in the anterior=posterior (A=P) direction following perturbation onset (Ponset , marked by a vertical dashed line). Horizontal lines mark A=P location of the ankle joint and the base-of-support (BOS) limit. Graphs C and D show trunk angular momentum (H) values. Negative (positive) values are associated with flexion (extension) of the trunk on the thigh. Step initiation is marked by a solid vertical line.

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RESULTS Our results indicate that across the range of perturbation s selected for analysis older adults were more likely than younger adults to use a step to recover balance (w2 ˆ 6:85; p < :01). In particular, older adults (OA) stepped in 95% of the trials (n ˆ 54 out of 57 trials), whereas younger adults (YA) stepped in 62% of the trials (n ˆ 30 out of 48 trials). In addition, the severity of the perturbation had a differential effect on OA compared to YA. As illustrated in Figure 3, 84% (16 of 19) of OA, but less than 15% (2 of 16) of YA, stepped following the low acceleration perturbation (9.4 m=s2 ) (w2 ˆ 10.89, p < :01). The medium acceleration condition (12.5 m=s2 ) elicited steps from 100% of OA and 75% of YA (w2 ˆ 1.58, p > :01). All participants stepped following the most demanding condition (15.2 m=s2 ). These results show the reliance of these older adults on the stepping response to regain equilibrium compared to these younger adults. Furthermore, these findings demonstrate the propensity among these older adults, compared to these younger adults, to use the stepping response following less demanding perturbations. Of the step trials available for analysis (n ˆ 25=30 for YA; n ˆ 40=54 for OA), the COM was located within the base of support on 76% (19 of 25) of the step trials for YA and on 83% (33 of 40) of trials for OA. Of the steps taken by YA with COM within BOS, 16% (3 of 19) were characterized as delayed steps; that is, the steps were initiated after the angular momentum of the trunk had reached its maximum value (Figure 4, left). By definition, delayed steps are characterized by reduction in the motion of the trunk prior to

FIGURE 3 Prevalence (%) of stepping among young and older adults at each perturbation condition.


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step initiation. For OA, 12% (4 of 33) of the steps were classified as delayed steps (Figure 4, left). The remaining 84% (16 of 19) of steps by YA and 88% (29 of 33) of the steps by OA were classed as rapid steps (Figure 4, right) steps initiated before the angular momentum of the trunk had reached a maximum value. Further analysis was restricted to only those responses classified as ``rapid’’ stepsÐthose steps taken while the COM was within the base of support and before trunk segment angular momentum had reached its maximum value (YA: n ˆ 10; OA: n ˆ 11). Rapid steps describe those responses where insufficient constraints have been placed on the trunk to slow its rotation prior to step initiation (see again Figure 2, Graph D). These steps contrast with ``delayed steps’’ where step initiation does not occur until after active muscle or passive tissue constraints have been applied to constrain trunk rotation (see Figure 2, Graph C). Mean time of step initiation across conditions was 144 § 32 ms post perturbation for YA and 155 § 30 ms for OA. These means were not significantly different (t(19) ˆ .82, p > :05). Likewise, there were minimal body configuration differences between the groups at step onset. The horizontal location of the COM was not different between groups (t(19) ˆ 1.17, p > 0:05). Tests of joint angle differences at the ankle, knee, and hip revealed significant differences only at the ankle with means of 5.2¯ § 1.7¯ and 3.6¯ § 0.8¯ for YA and OA respectively (t(19) ˆ 7 2.761,

FIGURE 4 Prevalence of rapid versus delayed steps for each age group. NOTE: Data are expressed relative to the number of steps taken within each perturbation condition. Young adults did not initiate steps when the COM was inside the BOS during the low acceleration (9.4 m=s2 ) condition. Delayed steps are defined as steps initiated after the local angular momentum of the trunk reaches its maximum; steps initiated before maximum local trunk angular momentum are classified as `rapid’ steps.

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p ˆ .012). Knee joint angle excursions of 2.3¯ ( § 1.7¯ ) and 1.2¯ (§ 1.0¯ ) for YA and OA were nonsignificant (t(19) ˆ 7 2.18, p ˆ 0.04; Bonferroni correction for multiple t tests), as were differences in hip joint excursions of 2.6¯ ( § 2.1¯ ) and 3.0¯ ( § 1.3¯ ) for YA and OA, respectively (t(19)ˆ ¡0:38, p ˆ .71). Analysis of the perturbation conditions that first led to the execution of rapid steps revealed that 70% of OA produced a rapid step at the 9.4 m=s2 condition while the remaining 30% of OA first made rapid steps following the 12.5 m=s2 condition (Figure 5). No young adults produced a rapid step at the 9.4 m=s2 condition (though 2 of 13 YA produced delayed steps at this condition) . Younger adults did not begin to produce rapid steps until the 12.5 m=s2 condition. Fifty percent of YA first produced rapid steps at the 12.5 m=s2 condition. The remaining 50% first produced rapid steps at the 15.2 m=s2 condition. The interesting observation in Figure 5 is that OA adopted the rapid step following lower perturbation accelerations than YA. Figure 6A, shows the comparison of group mean horizontal segmental momenta at the time of step initiation for the transition condition (the lowest perturbation acceleration that elicited a rapid step). The data have been normalized to perturbation impulse (time integral of the horizontal force between the onset of the perturbation and the initiation of the rapid step) to account for age group differences in the input acceleration producing the rapid step. Thus the data in Figure 6 (A or B) may be interpreted as the momentum achieved as a consequence of each unit of

FIGURE 5 Percent of each age group employing a rapid step in each acceleration condition. NOTE: The bars represent the smallest acceleration to elicit the rapid step in each group. Rapid steps were characterized by (1) the location of the COM being within the BOS at step initiation and (2) that steps were initiated before the local angular momentum of the trunk reached its maximum value.


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input acceleration delivered by the perturbation. Independent samples t tests failed to reveal age group differences for the lower extremity segments (t(19) ˆ 1.99, p ˆ .06; t(19) ˆ 1.9, p ˆ .07; t(19) ˆ 1.92, p ˆ .07 for the foot, shank, and thigh, respectively). Nor were age group differences found for the head-arms-trunk (HAT) segment (t(19) ˆ 1.51, p ˆ .15) or for the whole body COM (t(19) ˆ 1.10, p ˆ .28) (Figure 6B). This finding indicates that an

FIGURE 6 (A) Horizontal linear momenta for the foot, shank, thigh, and HAT segments. (B) Whole body COM horizontal linear momentum at the point of step initiation for young and older adults. NOTE: Data presented have been normalized to the perturbation impulse (time integral of the horizontal force between the onset of the perturbation and the initiation of the rapid step) and are expressed in absolute value terms.

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equivalent magnitude of body motion between younger and older adults is achieved following different input acceleration values. In particular, a lower perturbation acceleration creates an equivalent magnitude of body motion in an older adult that a higher perturbation acceleration creates for a younger adult. At step initiation , the groups were not distinguishe d by any of the horizontal segmental momentum terms. There was a consistent pattern, however, for the older adults to have larger momentum values. Given the rapidity with which the steps were initiated, it is conceivable that there simply was insufficient time for the motion differences to become apparent between the groups. Velocity terms (upon which momentum is based) reveal the instantaneous state of motion of a body or system. Acceleration terms reveal the rate at which velocity is changing. Thus, we evaluated the horizontal linear acceleration of segments to determine how the system, specifically segment velocities, were changing. At the onset of the step, segmental velocities could be quite similar between the groups, but the rate at which the velocities were changing could be quite different. Figure 7 presents the horizonta l linear acceleration of segmental endpoints at step initiation. The data reveal how the input acceleration was reduced, or augmented, across segments as the consequences of the perturbation were propagated through the linked system. As illustrated, there was a noticeable decrease in acceleration between the ankle and the head in YA. In contrast, this trend of attenuation was not seen among OA. For OA, there was a slight decrease in the resultant acceleration from the ankle to the hip; however, acceleration magnitudes subsequently increased from hip to head. As the

FIGURE 7 The horizontal acceleration of each joint relative to the linear acceleration of the toe. NOTE: Data presented were obtained at the point of step initiation and have been normalized for the magnitude of perturbation acceleration.


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head contains the apparatus for acceleration detection, we did an analysis of the differences in acceleration at the head. Indeed, the independent sample t test confirmed that there was a significant effect of age for horizontal acceleration at the head, with OA showing greater resultant accelerations than YA (t(19) ˆ 2.41, p ˆ .02).

DISCUSSION All participants were instructed to recover balance without moving their feet (i.e., stepping). Despite that instructiona l constraint, both young and older adults employ a rapid step as a balance recovery response. Our results also reveal that older adults employ a rapid stepping response more often when compared under similar condition s to younger adults. Older adults, compared to young adults, also resort to the step as a balance recovery strategy under less threatening conditions . The propensity of older adults to use the rapid step was not due to these adults having a larger momentum at the time of step initiation , though a trend for greater momenta in the older adults was evident. What did distinguis h the older adults from their younger adult counterparts was their failure to attenuate the destabilizing accelerations introduced by the perturbation . The difference between younger and older adults leading to rapid stepping under less threatening conditions is not a matter of older adults adopting a different behavioral response compared to younger adults. For example, in both groups, steps were initiated (on average) within the first 150 ms following onset of the perturbation. The older adults were neither quicker nor slower in responding to the balance threat. Further, the older adults were not in a more tenuous position at the time of step initiation . The body configuration at the moment of the step showed minimal differences between younger and older adults, with the older adults showing smaller joint excursions at the ankle and knee and slightly larger excursions at the hip. The issue here is also not one of older adults being unable to contain the center of mass trajectory to excursions within the limits of the base of support. Our results clearly show that rapid steps, those initiated well before the whole body COM has exceeded the base of support and before trunk angular momentum has reached its maximum value, are a typical response among both younger and older adults. The interesting observation is that older adults employ these rapid step responses under less threatening circumstances than younger adultsÐand not because they are thrown into a more precarious posture. Thus if the positional constraints of the COM did not require establishing a new base of support, was there another mechanical demand requiring the step? Such a mechanical demand could arise from a system that has gathered a great deal of momentum. It has been suggested that controllin g the horizontal momentum of the COM is a strategy employed to maintain balance during

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voluntary movements (e.g., gait: Kaya, Krebs, & Riley, 1998; sit-to-stand : Kaya et al., 1998; Pai & Lee, 1994; Pai & Rogers, l990; 1991). The strategy would be to limit the segmental or whole body momenta to those levels that can be managed by the application of an opposing impulse (the application of muscle force over time). The data presented did not strictly support this hypothesis in that we found no significant group differences in horizontal COM momentum at the time of step initiation . We did, however, observe a tendency for older adults to be described by larger segmental and whole body momenta. It was this observation that led to the consideration of not just the instantaneous state of the system as represented by the velocity-based momentum variable, but how the system was changing as revealed in the acceleration analysis. Based on the acceleration analysis, the production of rapid steps by older adults under relatively less threatening conditions is related to the rate at which the body is gaining momentum. Although the groups were not distinguishe d by differences in the momentum of the body possessed at step onset, the older adults were experiencing larger accelerations. Older adults demonstrated only a very modest reduction in the horizontal acceleration across the lower extremity segments. The trunk and head, in fact, experienced augmented accelerationsÐthus increasing gains in momentum. This finding is consistent with previous reports of age-related differences in the attenuation of horizontal segmental accelerations during gait. Winter (1991) found that younger adults were able to reduce head acceleration to 23% of the pelvic acceleration magnitude. Older adults attenuated the acceleration to only 42% of pelvic magnitudes. Our results reveal the same tendency. Both younger and older adults reduced the horizontal acceleration of segments up to the level of the hip joint. It is at the hip joint, however, where age-related differences in the pattern of attenuation emerge (see Figure 7). Although the acceleration at the shoulder is greater than that at the hip for both age groups, the magnitude of acceleration increases substantially among older adults, with the cumulative effect that older adults experience larger horizontal accelerations of the head than younger adults per unit of perturbation acceleration. Thus it is possible that older adults perceive smaller challenges to balance (relative to threats experienced by younger adults) as more threatening because the undamped accelerations experienced at the head imply a larger threat to balance than actually exists. Allum and Pfaltz (1985) have shown that acceleration of the head following a platform perturbation activates the vestibular system and leads to early (within 150 ms) postural adjustments serving to stabilize the body (activation of ankle musculature) and head (activation of head musculature). Given the timing of this vestibular-base d response, one may infer that the vestibular system also underlies the origin of the rapid step response. Further work investigatin g the postural compensations used by vestibular-deficien t patients following rapid platform accelerations is needed to elucidate this possibility.


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It is in the performance of voluntary tasks that we acquire a sense of force production limits and set thresholds that define our strategic choices. In the face of an unexpected perturbation, it is these limits, defined under voluntary conditions and past experience, that provide the first approximation s for how to perform the `novel’ task. What we observed was the possibility of a reduced acceleration attenuation response in the older adults. Whether this is a physical limitation or a learned response is not clear from these results. Changing the consequences of taking a step (e.g., Brown & Frank, 1997) would allow us to probe limitations in physical capacity versus learned limits and examine in greater detail the age-related behavioral response differences now typically observed during balance-control tasks.

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Pai, Y. C., & Lee, W. A. (1994). Effect of a terminal constraint on control of balance during sitto-stand. Journal of Motor Behavior, 26, 247±256. Pai, Y. C., & Patton, J. L. (1997). Center of mass velocity-displacemen t predictions for balance control. Journal of Biomechanics, 30, 347±354. Pai, Y. C., & Rogers, M. W. (1990). Control of body mass transfer as a function of speed of ascent in sit-to-stand. Medicine and Science in Sports and Exercise, 22, 378±384. Pai, Y. C., & Rogers, M. W. (1991). Segmental contributions to total body momentum in sit-tostand. Medicine and Science in Sports and Exercise, 23, 225±230. Rogers, M. W., Hain, T. C., Hanke, T. A., & Janssen, I. (1996). Stimulus parameters and inertial load: Effects on the incidence of protective stepping responses in healthy human subjects. Archives of Physical Medicine and Rehabilitatio n, 77, 363±368. Thelen, D. G., Schultz, A. B., & Alexander, N. B. (1996). Effects of age on rapid ankle torque development. Journals of Gerontology : Medical Sciences, 51, 5, M226±M232. Winter, D. A. (1990) . Biomechanics and motor control of human movement (2nd ed.). Toronto : John Wiley & Sons. Winter, D. A. (1991). The biomechanics and motor control human gait: Normal, elderly, and pathological (2nd ed.). Waterloo, Ontario: University of Waterloo Press. Wolfson, L., Whipple, R., Derby, C. A., Amerman, P., & Nashner, L. (1994). Gender differences in the balance of healthy elderly as demonstrated by dynamic posturography . Journal of Gerontology : Medical Sciences, 49, M160±M167.