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. Author manuscript; available in PMC: 2011 Nov 30.
Published in final edited form as: Exp Brain Res. 2006 Jan 18;171(4):459–468. doi: 10.1007/s00221-005-0293-6

Altered Timing of Postural Reflexes Contributes to Falling in Persons with Chronic Stroke

Daniel S Marigold 1,3, Janice J Eng 1,2,3
PMCID: PMC3226801  CAMSID: CAMS1999  PMID: 16418855

Abstract

The purpose of this study was to determine differences in the timing of postural reflexes and changes in kinematics between those who fell (Fallers) in response to standing platform translations and those who did not (Non-fallers). Forty-four persons with stroke were exposed to unexpected forward and backward platform translations while standing. Surface electromyography from bilateral tibialis anterior, gastrocnemius, rectus femoris, and biceps femoris were recorded along with kinematic data. Those that fell in response to the translations were compared to those who did not fall in terms of (1) postural reflex onset latency, (2) the time interval between the activation of distal and proximal muscles (i.e. intralimb coupling), and (3) changes in joint angles and trunk motion. Approximately 85% of falls occurred in response to the forward translations. Postural reflex onset latencies were delayed and intralimb coupling durations were longer in the Faller versus Non-faller group. At the time that the platform completed the translating motion (300 ms), the Faller group demonstrated higher trunk velocity, greater change in paretic ankle angle, and the trunk was further behind the ankle compared to the Non-faller group. This study suggests that following platform translations, delays in the timing of postural reflexes and disturbed intralimb coupling result in changes in kinematics, which contribute to falls in persons with stroke.

Keywords: falls, cerebrovascular accident, reflex, perturbation, postural control

Introduction

Falls and fall-related injuries occur frequently among community-dwelling stroke survivors with many of these individuals falling several times within a given year (Forster and Young, 1995; Harris et al. 2005; Hyndman et al., 2002; Jorgensen et al. 2002). In fact, persons with stroke are at a much greater risk of fractures (i.e. sevenfold increase) than healthy adults (Kanis et al. 2001). The primary mechanisms behind a falling episode in persons with stroke are not clear. To date, the few studies on falls in persons with chronic stroke have focussed on predictors (Forster and Young 1995; Harris et al. 2005; Jorgensen et al. 2002), which can identify those individuals at high risk for falls, but do not delineate the specific strategies necessary to prevent the occurrence of a fall. An understanding of fall prevention strategies is essential in designing effective rehabilitation programs that may lead to improved quality of life for these individuals and a reduced burden on the health care system.

Following an unexpected balance-threatening event (i.e. perturbation), automatic postural reactions (or postural reflexes) are initiated to prevent the occurrence of a fall and maintain upright stability. The central nervous system must utilize and integrate the available sensory and environmental information to select an appropriate response strategy. During a stance perturbation, these postural reflexes are initiated with onset latencies between 100 and 120 ms in healthy older adults (Manchester et al. 1989; Marigold et al. 2004). In addition to the rapid timing of these responses, postural reflexes are coordinated into muscle synergies such that muscles act together as a functional unit through neuronal coupling (Shumway-Cook and Woollacott 2001). A typical muscle synergy observed during standing platform translations involves the activation of musculature surrounding the ankle joint followed by more proximal leg and trunk muscles, termed the ‘ankle strategy’ (Horak and Nashner 1986). Under more challenging conditions, a ‘hip strategy’ can also be observed where muscles are coordinated to act predominantly at the hip joint (Horak and Nashner 1986; Runge et al. 1999) or a compensatory step may be used to re-position the base of support (Maki and McIlroy 1997).

Studies have shown that persons with stroke have delayed paretic limb muscle onset latencies following perturbations while standing on a moveable platform compared to their non-paretic limb and to healthy older adults (Berger et al. 1988; Dietz and Berger 1984; Di Fabio and Badke 1988; Di Fabio et al. 1986; Di Fabio 1987; Marigold et al. 2004). In addition, the timing between the activation of the distal and proximal muscles (i.e. intralimb coupling) appears to be disturbed in stroke (Badke and Duncan 1983; Di Fabio et al. 1986). Furthermore, ankle torque responses are reduced on the paretic side following platform translations (Ikai et al. 2003). Whether delays in postural reflex muscle onset latency or disturbed intralimb coupling contribute to falls in persons with stroke is unknown. Investigation into the kinematics following the perturbation would also facilitate the interpretation of the muscle activation patterns and be useful in better understanding falls.

Falls generated in a laboratory setting can be induced in such a way that every individual is exposed to the same destabilizing force. Consequently, the neurophysiological mechanisms associated with falls can be quantified. Therefore, the purpose of this study was to determine differences in postural reflex muscle onset latency and intralimb coupling durations between those who fell in response to standing platform translations and those who did not. Furthermore, we also investigated the changes in kinematics that took place between the start and end of the platform movement. Persons with chronic stroke were exposed to a series of standing platform translations that had the capability to induce a fall. We hypothesized that individuals who fall in response to the external perturbation would have slower postural reflex onset latencies, greater delays between the activation of the distal and proximal muscles (i.e. intralimb coupling durations), and demonstrate larger changes in kinematics over the duration of the translation (first 300 ms) compared to those who do not fall.

Methods

Participants

Forty-four persons with unilateral hemiparesis due to stroke were recruited from the community. The inclusion criteria for the persons with stroke were: (1) over 50 years of age, (2) a single stroke at least one-year from onset (i.e. chronic), (3) able to stand independently for at least 5 minutes without an assistive device, and (4) able to follow two consecutive instructions. Persons with musculoskeletal conditions (e.g. severe arthritis, recent joint replacement surgery) or neurological disorders (e.g. Parkinson’s disease) in addition to their stroke were excluded. Following university and hospital ethics approval, informed consent was received from all participants prior to their participation.

Information on the type and location of the participants’ stroke was collected through medical records and/or physician notes. This information along with other characteristics of the participant sample is described in Table 1. The American Heart Association Stroke Functional Classification (AHASFC) was used to provide an indication of the level of disability of the participants. The AHASFC is based on the level of independence of an individual, where level I represents complete independence in basic and instrumental daily activities of living and level V represents complete dependence (Kelly-Hayes et al. 1998). The Berg Balance scale assessed functional balance and the Timed Up and Go test assessed functional mobility. The Berg Balance scale consists of 14 balance-related tasks including stepping, reaching, and turning (Berg et al. 1992). The Timed Up and Go test (Podsiadlo and Richardson 1991) measures the time to stand up from an arm chair, walk a distance of 3 m, turn, and walk back to the chair and sit down again. A pressure aesthesiometer kit (8 monofilaments) was used to determine the tactile sensitivity (i.e. cutaneous sensation) of the persons with stroke. Participants faced away from the tester and the monofilaments were applied for less than a second and deformed to half their length against the plantar surface of the feet (heel pads). A sequence of thicker to thinner filaments was applied and the number of the monofilament that was last able to be felt was recorded. The thickest monofilament (i.e. poor tactile sensitivity) was number 8, while number 9 indicated no tactile sensation.

Table 1.

Participant characteristics of the Non-faller and Faller groups.

Non-fallers (n = 33)
Mean (SD) or n		 Fallers (n = 11)
Mean (SD) or n
Gender, M/F 24/9 6/5
Age, yrs 66.1 (8.7) 68.9 (6.7)
Height, cm 171.3 (9.3) 165.3 (9.8)
Mass, kg 79.9 (17.7) 79.2 (14.9)
Stroke Duration, yrs 3.6 (2.2) 3.4 (2.2)
Hemiparetic Side, R/L 11/22 4/7
AHASFC, I/II/III/IV/V 8/10/13/2/0 1/2/6/1/1
Cutaneous Sensation - Paretic Heel Pad 5.5 (3.0)* 5.0 (1.0)*
Cutaneous Sensation - Non-paretic Heel Pad 5.0 (2.0)* 6.0 (1.0)*
Berg Balance, max. 56 46.4 (4.9) 37.6 (7.4)
Timed Up and Go, sec 17.7 (12.0) 26.0 (11.2)
Type of Stroke 17 ischemic
16 hemorrhagic		 8 ischemic
3 hemorrhagic
Stroke Location 10 cortical
23 subcortical/brainstem/cerebellum		 6 cortical
5 subcortical/brainstem/cerebellum
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Abbreviations: M = male; F = female; R = right; L = left; AHASFC = American

Heart Association Stroke Functional Classification

*

Median (IQR)

Protocol

A total of 20 platform translations (0.300 s duration, 0.08 m ramp displacement, 0.27 m/s velocity, and 3.00 m/s2 acceleration and deceleration), separated by 15–30 second intervals, were induced while participants stood on two force plates (Bertec Corp., Columbus, OH, USA) embedded in a custom built platform with one limb on each force plate. To prevent the occurrence of a fall to the ground, participants wore a full-body harness that was attached to a beam in the ceiling via a dynamic rock-climbing rope and at least one spotter stood beside them. The harness was adjusted so as not to take any weight of the individual unless they fell. The experimenter ensured that the participant stood straight with feet shoulder width apart and that no noticeable pre-leaning occurred in the sagittal plane prior to platform movement.

Participants were instructed to try and maintain an upright standing posture following the translation. Participants were told that the platform could move at any time but the onset, number, and direction of the translations were unexpected in nature. The direction of the translation was counterbalanced across participants so that either 10 consecutive backward translations followed 10 consecutive forward translations or vice versa. Thus, the perturbations were only truly random in the time applied.

Surface electromyography (EMG) (Bortec Electronics, Calgary, Canada) were collected from bilateral tibialis anterior (TA), the medial head of gastrocnemius (MG), rectus femoris (RF), and biceps femoris (BF) using self-adhesive electrodes (Ag/AgCl) (Kendall Medi-Trace®, Chicopee, MA, USA) placed approximately 2 cm apart and longitudinally on the belly of the muscle. The signals were band-pass filtered (10–1000 Hz) and sampled at 600 Hz for 6 seconds (2 seconds prior to platform movement and 4 seconds following) along with force plate data. The common mode rejection and input impedance of the EMG system was 115 db and 10 GΩ, respectively.

One 3-dimensional optoelectronic camera (OPTOTRAK, Northern Digital, Inc., Waterloo, Canada) was used to collect whole body kinematics (sampling frequency of 60 Hz) and a video camera (Panasonic Canada Inc., Mississauga, Canada) recorded the perturbation trials from the left side of the participant’s body for qualitative observations. A total of 11 infrared emitting diodes (IREDs) were placed bilaterally on the fifth metatarsal, ankle, knee, greater trochanter, acromion, and the xyphoid process of each participant and viewed by the 3D OPTOTRAK camera placed in the frontal plane.

Data Analyses

All data was processed using custom written MATLAB (Mathworks, Natick, MA, USA) programs. EMG was full-wave rectified and low-pass filtered at 100 Hz (zero lag, second-order, Butterworth algorithm). The mean EMG signal for one second prior to the onset of platform movement was determined along with the standard deviation. Postural reflex muscle onset latency (see Figure 1) was defined as an increase in muscle activity that exceeded two standard deviations of the mean background activity for at least 30 ms following the onset of platform movement and was determined by a combination of visual inspection and computer algorithm via an interactive program (Marigold et al. 2004). The person performing the visual inspection was blinded as to whether the person had fallen in response to the platform translation. Less than 1% of trials were eliminated through visual inspection due to problems such as noisy or missing data. The time interval between the activation of the distal (i.e. TA) and proximal (i.e. RF) muscles within the same limb was also determined (see Figure 1) and measured the intralimb coupling duration.

Figure 1.

Figure 1

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Diagrammatic definitions of postural reflex onset latency and intralimb coupling. Postural reflexes are from the paretic limb of one individual.

Position data for all IRED markers were low-pass filtered at 6 Hz using a second-order, dual-pass, Butterworth algorithm (Winter 2005). Subsequently, bilateral ankle, knee, and hip joint angles were calculated (see Figure 2). In addition, trunk pitch angle (i.e. forward flexion/backward extension) from a line joining the bisection of the acromion markers and the bisection of the greater trochanter markers in the saggital plane was calculated. The anterior-posterior (AP) trunk velocity (estimated from the displacement of the xyphoid marker) was also calculated. In addition, the horizontal distance between the xyphoid marker (representing the trunk) and the paretic and non-paretic ankle markers (a close approximation of the axis of body rotation if an ankle strategy is used) were determined and are referred to as the paretic horizontal Trunk-Ankle distance and non-paretic horizontal Trunk-Ankle distance, respectively.

Figure 2.

Figure 2

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Kinematic measures. A: Bilateral ankle, knee, and hip joint angles along with trunk pitch angle were determined. Note that for the ankle angle, 90° is vertical and less than 90° indicates increasing dorsiflexion; for the knee angle, 180° is full extension with decreasing values indicating flexion; for the hip angle, positive angles are hip flexion and negative angles are hip extension; and for trunk pitch angle, less than 0° is trunk pitch backwards (i.e. backward trunk extension). B: Relationship between xyphoid and ankle marker (representing the trunk and close approximation of the axis of body rotation if an ankle strategy is used), respectively) defining the paretic and non-paretic Trunk-Ankle measures. Note that a negative value represents the xyphoid marker posterior to the ankle.

Kinematic measures (i.e. joint angles and horizontal Trunk-Ankle distance) were extracted at the start of platform movement and at the end of the platform movement. Subsequently, the changes in value (i.e. the difference in the kinematic variable between the end and start of the platform movement) were calculated. The interval between the start and end of platform movement was approximately 300ms.

Force plate data was calibrated and filtered (second-order, 50 Hz low-pass filter). The average centre of pressure (COP) location in both the AP and medial-lateral (ML) directions was determined over a 500 ms period prior to platform movement from the net COP of the two force plates. The COP location was used to determine whether participants were altering their initial standing posture over the consecutive trials (e.g. leaning forward).

The number of falls induced by the platform translations was also recorded using qualitative observations during testing and through video analysis by the experimenter and an individual blinded to the purpose of the study. A fall was defined as a participant who lost their balance and required the harness to prevent them from falling to the floor (e.g. participants applied force through the rope and harness system as evident from the rope and harness going taut). Trials were additionally checked through video analysis to confirm a fall episode. The spotter was instructed to assist the participant only after they had fallen (i.e. to prevent swinging caused by the rope). Subsequently, those that fell at least once during testing were labeled as ‘Fallers’ and those that did not were labeled ‘Non-fallers’.

Statistical Analyses

Clinical measures of functional balance (Berg Balance) and mobility (Timed Up and Go) were compared between the Non-faller and Faller groups using Mann-Whitney U tests.

The majority of falls occurred during forward platform translations (see Results section below). Thus, only these trials were analyzed. Since platform translations were presented consecutively for each direction, the first step in our analyses was to determine the habituation to consecutive trials, if any, of the Faller and Non-faller groups before data was pooled across trials. Postural reflex onset latency, kinematic variables, and COP location were assessed for habituation. Thus, in both groups, a repeated measures analysis of variance (ANOVA) was performed for each variable (after it was rank transformed due to the fact that the data was not normally distributed) to examine a trial order effect. When a significant ANOVA was found, a contrast analysis compared the first trial versus the average of the remainder trials to further explore the habituation response. In addition, a Student-Newman-Keuls (SNK) post-hoc test was also performed to potentially separate trials. No characteristic startle responses were observed for either group. For the Non-faller group there were several measures that demonstrated a trial order effect including bilateral TA and non-paretic RF postural reflex onset latency, bilateral ankle angles (just prior to the start of the translation), paretic and non-paretic horizontal Trunk-Ankle distance (just prior to the start of the translation), and the changes in non-paretic ankle angle, non-paretic knee angle, and bilateral hip angle (over entire translation displacement) values (p < 0.05). The a priori contrast analysis demonstrated that the first translation trial was significantly different from subsequent trials for all of these measures except for the paretic ankle angle at translation onset: the latencies were slower and there were greater changes in kinematics (from the start to the end of platform movement) in the first trial. SNK results separated the first translation trial out for the non-paretic RF onset latency, paretic and non-paretic horizontal Trunk-Ankle distance (just prior to the start of the translation), and changes in the non-paretic ankle, knee, and hip angle between the start and end of platform movement. In contrast, there were no trial order effects for any measure (p > 0.05) for the Faller group except the paretic horizontal Trunk-Ankle distance (just prior to the start of the translation) (p < 0.05) but the SNK post-hoc tests failed to separate trials.

As the majority of variables for the Non-faller group demonstrated a clear difference between the first trial and the remainder trials we felt there was sufficient evidence to remove this trial and assess the remaining trials to ensure homogeneity of the responses. There were no trial effects in the Non-faller group once the first trial was removed. Thus, the first forward translation trial was removed from the subsequent analyses for the Non-faller group such that nine trials per participant were used. Since the vast majority of measures did not show a trial order effect for the Faller group, all ten trials per participant were used in the analysis for this group.

Mann-Whitney U tests compared the responses of the Non-faller group with the fall trials of the Faller group for postural reflex muscle onset latencies, intralimb coupling, and kinematic variables. An alpha level of 0.05 was used for all statistical analyses.

Results

Eleven (25 %) persons with stroke fell at least once in response to forward platform translations, for a total of 41 falls. In contrast, only three (7 %) persons with stroke fell in response to backward translations for a total of seven falls. As a result, only forward platform translations were analyzed. Among the Faller group, there was an average of 3.7 falls/person in response to the forward translations. Although more of the falls (65.9 %) occurred within the first half of platform translation trials, there were no significant order effects.

While there were a smaller number of individuals who fell compared to individuals who did not fall, we felt it was preferable to use the entire sample rather than match the number of Non-fallers to Fallers. The proportion of fallers (i.e. 25%) within the entire sample agrees with proportions of fallers from other studies, which examine prospective community-based falls in persons with chronic stroke (Jorgensen et al. 2002).

Scores on the Berg Balance scale were significantly reduced in the Faller group compared to the Non-faller group (Fallers = 37.6 ± 7.4 versus Non-fallers = 46.4 ± 4.9, p = 0.004). In addition, the Faller group took more time to complete the Timed Up and Go test compared to the Non-faller group (Fallers = 26.0 ± 11.2 seconds versus Non-fallers = 17.7 ± 12.0 seconds, p = 0.017).

Both groups activated the MG muscle less than 20% of the trials for the forward platform translations. Thus, the primary recovery muscles (TA and RF) for this direction (Horak and Nashner 1986) as well as the antagonist muscle (BF) were analyzed.

Comparison of postural reflexes between the Non-faller and Faller groups in response to forward translations

Typical postural reflex profiles in response to the forward platform translation for one individual are shown in Figure 3. A distal to proximal sequence of muscle activation in both paretic and non-paretic limbs and both the Non-faller and Faller groups can be seen. The results demonstrate that the Faller group had significantly slower postural reflex onset latencies compared to the Non-faller group over consecutive trials (see also Table 2). The paretic TA, but not non-paretic TA, was delayed in the Faller group (Table 2). Proximal muscles of the lower limb were particularly altered in the Faller group. The RF of the Faller group was delayed by approximately 35 ms and 22 ms for the paretic and non-paretic limbs, respectively, compared to the Non-Faller group. In addition, the antagonist, paretic BF, was delayed approximately 30 ms in the Faller group. Both the paretic and non-paretic intralimb coupling durations were 16 ms longer for the Faller group compared to the Non-faller group (Table 2).

Figure 3.

Figure 3

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Typical postural reflex responses for the paretic and non-paretic tibialis anterior (TA) and rectus femoris (RF) muscles during forward platform translations for a (A) Non-faller and (B) Faller. Vertical lines are drawn at the onset latency of the Non-faller to show delay in onset latency with the Faller. Platform translation commenced at time zero.

Table 2.

Mean (SD) postural reflex muscle onset latencies and intralimb coupling durations for the Non-faller (excluding the first forward translation trial) and Faller (Fall trials only) groups.

Measure Non-fallers (n = 33) Fallers (n = 11) P-value
Onset Latency (ms)
 Paretic TA 119.0 (21.6) 131.5 (21.9) <0.0001
 Paretic RF 149.5 (35.8) 184.4 (45.8) 0.001
 Paretic BF 158.8 (39.2) 189.3 (34.1) <0.0001
 Non-paretic TA 108.5 (20.0) 112.5 (18.1) 0.149
 Non-paretic RF 137.3 (36.2) 159.3 (34.3) <0.0001
 Non-paretic BF 136.9 (31.4) 139.2 (39.0) 0.854
Intralimb Coupling (ms)
 Paretic limb 33.6 (30.6) 49.5 (38.6) 0.039
 Non-paretic limb 31.0 (34.4) 46.8 (28.3) <0.0001
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Comparison of kinematics between the Non-faller and Faller groups in response to forward translations

Table 3 shows the results of the kinematic measures. At translation onset (i.e. immediately preceding platform movement) the non-paretic ankle of the Non-faller group was significantly more dorsiflexed and the paretic hip was more flexed (angle greater than 180°) than the Faller group. Furthermore, the Non-faller group’s trunk was more anterior to their non-paretic ankle compared to the Faller group.

Table 3.

Mean (SD) kinematic variables for the Non-faller (excluding the first forward translation trial) and Faller (Fall trials only) groups.

Measurea Non-fallers (n = 33) Fallers (n = 11) P-value
Value at translation onset
 Paretic Ankle Angle (°) 76.0 (6.3) 76.1 (6.8) 0.906
 Non-paretic Ankle Angle (°) 75.6 (5.8) 78.6 (3.7) <0.0001
 Paretic Knee Angle (°) 171.3 (13.5) 168.2 (10.9) 0.194
 Non-paretic Knee Angle (°) 171.2 (11.5) 173.0 (6.3) 0.255
 Paretic Hip Angle (°) −2.2 (10.8) 1.5 (6.5) 0.002
 Non-paretic Hip Angle (°) −1.8 (9.4) 0.6 (4.9) 0.070
 Trunk Pitch Angle (°) −1.2 (5.3) −0.2 (4.6) 0.761
 Paretic Trunk-Ankle distance (cm) 6.0 (4.5) 5.1 (3.0) 0.459
 Non-paretic Trunk-Ankle distance (cm) 6.3 (5.0) 4.6 (2.9) 0.011
Value at max. translation displacement
 Trunk Velocity (cm/s) −6.6 (7.3) −10.2 (6.3) 0.001
Change over translation displacement
 Paretic Ankle Angle (°) 1.1 (3.3) 2.3 (2.0) 0.003
 Non-paretic Ankle Angle (°) 0.3 (4.4) 1.2 (2.4) 0.572
 Paretic Knee Angle (°) −6.5 (6.3) −5.2 (4.0) 0.131
 Non-paretic Knee Angle (°) −7.4 (7.3) −7.4 (4.6) 0.316
 Paretic Hip Angle (°) 5.5 (4.0) 6.0 (2.9) 0.543
 Non-paretic Hip Angle (°) 5.9 (4.0) 6.2 (2.8) 0.323
 Trunk Pitch Angle (°) −2.1 (2.6) −2.4 (1.7) 0.298
 Paretic Trunk-Ankle distance (cm) −8.2 (1.7) −9.4 (1.2) <0.0001
 Non-paretic Trunk-Ankle distance (cm) −7.9 (1.7) −8.8 (0.9) <0.0001
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a

See Figure 2 and accompanying legend for definitions of kinematic variables

In response to the forward platform translations, there was an immediate change in ankle angle followed closely by changes in knee and hip angles (see Figure 4). In particular, both groups demonstrated ankle plantarflexion initially after the perturbation commenced. This was presumably from the platform sliding forward under the feet. This plantarflexion was then followed by ankle dorsiflexion among the Non-faller group; however, Fallers failed to show this switch to ankle dorsiflexion (particularly on the paretic side). In addition, both groups demonstrated knee and hip flexion, and trunk backward extension in response to the forward platform translation. The trunk velocity of the Faller group was significantly higher at the end of the translation than the Non-faller group. Furthermore, the trunk was more posterior relative to the ankles compared to the Non-faller group at the end of the translation period.

Figure 4.

Figure 4

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Kinematic profiles of paretic and non-paretic ankle, knee, and hip joint angles for a representative Non-faller (solid line) and Faller (dashed line). Platform translation profile is shown in the bottom panel under both the paretic limb and non-paretic limb columns. Platform movement commenced at time zero and denoted by the vertical dashed lines in the joint angle graphs above. Note that for the ankle angle, 90° is vertical and less than 90° indicates increasing dorsiflexion; for the knee angle, 180° is full extension with decreasing values indicating flexion; for the hip angle, positive angles are hip flexion and negative angles are hip extension.

Discussion

Although others have reported delays in paretic limb onset latency and disturbed intralimb coupling in response to platform perturbations in persons with stroke (Badke and Duncan 1983; Berger et al. 1988; Dietz and Berger 1984; Di Fabio and Badke 1988; Di Fabio et al. 1986; Di Fabio 1987; Marigold et al. 2004), none have attempted to relate these observations to falls. The results of this study suggest that upper motoneuron lesions affect the neuronal circuitry involved in the generation and organization of a postural reflex following a balance-threatening event during stance, which contributes to falls in this population. Specifically, we found (1) delayed paretic TA and BF and bilateral RF onset latency and (2) longer paretic and non-paretic intralimb coupling durations for the Faller group in response to forward platform translations. In addition, the Faller group had greater trunk velocity (backward) at the end of the platform translation, greater difference in paretic ankle angle over the translation (i.e. between end and start of platform movement), and greater change in horizontal distance between the trunk and ankle markers such that the xyphoid marker was further behind the ankle compared to the Non-faller group.

The majority of falls occurred in response to the forward platform translations. The greater difficulty with forward translations may stem from the biomechanics of the foot as the centre of pressure under the foot has less distance of possible backward travel (as compared to forward travel) to control the backward falling COM. In addition, while applying transient treadmill accelerations to standing, Berger et al. (1988) found ankle joint angular velocities were faster in response to forward treadmill accelerations compared to backward accelerations, which might contribute to the greater difficulty in recovering from forward platform translations. Furthermore, individuals with supraspinal lesions often exhibit dorsiflexor muscle weakness reflected in foot drop during gait and the induced backward sway caused by the forward translation requires ankle dorsiflexion for recovering.

Disrupted timing of postural reflexes contributes to falls

Postural reflexes in all muscles were elicited between approximately 110 and 190 ms following the onset of platform movement (see Tables 2). Postural reflex onset latencies of healthy older adults in response to similar platform translations range from 100–115 ms (Marigold et al. 2004). Studies using transcranial magnetic stimulation have suggested that reflex responses after 80–90 ms may be partially mediated by a transcortical reflex loop (Nielsen et al. 1997; Petersen et al. 1998; van Doornik et al. 2004). Therefore, the greater postural reflex delay in the Faller group may represent a greater deficit in cortical functioning, which was exhibited by the lower functional scores in this group. Cortical reorganization is a critical component of motor recovery following a stroke (Thickbroom et al. 2004).

A coordinated muscle synergy, which links distal and proximal musculature, is fundamental in providing an efficient postural response and in maintaining upright stability following a platform translation. Proximal muscles were delayed by 30 ms compared to the distal muscles in the Non-fallers, while this delay was 170% longer (50 ms) in the Fallers. Intralimb coupling durations during forward platform translations in healthy young and older adults have been reported to be between 7 and 18 ms (Badke and Duncan 1983; Horak and Nashner 1986).

Spinal interneuronal circuitry has been suggested to be impaired following stroke (Dietz and Berger 1984) and may have contributed to the longer intralimb coupling durations particularly for those that fell during platform translations. However, lower extremity intralimb coordination has contributions from both spinal and supraspinal mechanisms (Forssberg et al. 1975; Forssberg et al. 1980; Hiebert et al. 1994). Given the cortical contribution to the latency delays, we postulate that disrupted supraspinal centres are mainly responsible for the longer intralimb coupling duration.

Differences in kinematics between the Non-faller and Faller groups and their relation to muscle activation patterns

The abrupt forward platform translation results in the body swaying backward. This detrimental motion must be overcome to maintain an upright posture and prevent falling. In particular, the platform sliding forward causes an initial ankle plantarflexion (see Figure 4), which must be slowed via TA activation and subsequently reversed to dorsiflexion (this reversal was absent in many of the falling trials). As the rest of the body sways backward, RF activation serves to facilitate hip flexion and reduce the trunk backward motion while BF activation flexes the knee joint (reflected in the kinematics: Figure 4), which can lower the body and thereby increase stability. RF activation may also act to control the knee flexion caused by the BF muscle and the long head of the BF may participate in the hip movement observed due to their bi-articular roles. In view of these actions, Fallers exhibited delays in bilateral RF and paretic TA postural reflexes. Consequently, these altered muscle reflexes resulted in greater backward trunk velocity at the end of the translation with the trunk further behind the ankle. The greater change in paretic ankle angle over the translation among the Fallers may be due to poor control of the ankle dorsiflexor muscles. For example, Marigold et al. (2004) found persons with chronic stroke have greater difficulty in modulating the activity of this muscle to different amounts of weight-bearing load following platform translations compared to healthy older adults. The delay in paretic BF exhibited by the Faller group may explain the lack of knee flexion (see Figure 4) observed in this group. In contrast, the large knee flexion among the Non-fallers may have facilitated their recovery response. Thus, differences in the latency of postural reflexes and the resulting changes in kinematics among the Faller group contributed to the large number of fall episodes in this group.

The initial standing posture adopted by the Non-faller group prior to platform translation demonstrated that they were more dorsiflexed at the non-paretic limb, more extended at the paretic hip joint, and had the trunk positioned more anterior relative to the ankle compared to the Faller group. This posterior shift in resting standing before the platform translation for the Faller group resembles the tendency of persons with Parkinson’s disease to have their centre of pressure shifted backwards compared to healthy individuals during a quiet standing task (Schieppati and Nardone 1991). Further work on the relationship between muscle activation patterns and kinematics is required to extend our understanding of falls in persons with stroke.

Limitations

Limitations of this study include the nature of the presentation of the platform translations and the variability of stroke lesion site among our participants. Platform translations were presented in a consecutive fashion rather than randomly. It is possible that the postural reflex onset latencies would be different when trials are presented in a random manner. However, muscle onset latencies were not different between trials after the first translation (which was removed from the analyses in the Non-faller group).

Since our sample had a large amount of variability in the location of their lesion, the influence of specific brain regions involved in the postural reflex could not be determined. However, we had similar responses despite varying lesion volume and sites, which suggests that upper motoneuron lesions may produce similar deficits in postural control. It is possible that we might have been able to attribute specific postural control functions to brain regions if patients with targeted lesions had been assessed.

In addition, the magnitude of the postural reflexes was not examined in this study due to the inherent problems of comparing EMG magnitude across different subject pools (e.g., skin impedance, varying fat content, inability to establish a true maximum voluntary contraction in stroke to normalize data). It is known that persons with stroke have decreased motor unit recruitment and disturbed motor unit firing rates (Dietz et al. 1986; Gemperline et al. 1995), which may contribute to the reduced motor output of these individuals. Thus, differences in postural reflex magnitude may have also contributed to falls in this population.

Conclusion

Understanding the neural mechanisms of postural control and falls will form the basis of interventions to improve postural control. We have found several differences between Non-fallers and Fallers in our study; however, future research needs to determine the relationship of laboratory-generated falls compared to those occurring in the community or during activities of daily living. Nonetheless, our results suggest that rehabilitation of persons with chronic stroke should include a focus on exercises that elicit rapid movements and/or perturbation training. In fact, an agility training program with an emphasis on fast dynamic movements and perturbation training in persons with chronic stroke has found exercise can lead to faster paretic limb postural reflexes following platform translations, improved functional balance and mobility, and reduced falls (Marigold et al. 2005).

Acknowledgments

The authors wish to thank Erica Botner, Cheryl Louis, and Craig Tokuno for their help. This study was supported by an operating grant from the Canadian Institutes of Health Research (CIHR), salary support to JJE from CIHR (MSH-63617) and the Michael Smith Foundation for Health Research (MSFHR), and trainee support to DSM from MSFHR and the Natural Sciences and Engineering Research Council of Canada.

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