Postural steadiness is maintained by a control mechanism in which the body's center of mass (COM) is regulated through vestibular and neuromuscular responses to the movement of the center of pressure (COP).1 These spontaneous changes in the COP during quiet standing are known as postural sway and can be influenced by external forces such as backpack loads.2,3 The increased weight on the back can result in unfavorable forces on the spine and surrounding tissues, leading to pathological back problems and compensatory postural adjustments that may lead to increased stresses on bodily structures and hamper development.4,5 The current backpack load often borne by schoolchildren exceeds, in terms of percentage of weight, some legal load-bearing limits set for adults.6
Balance impairments caused by altered sensory, motor, or central nervous function related to factors such as a loss of a limb will be reflected in altered characteristics of postural sway.7–10 Functional balance is significantly decreased in children with lower-limb amputation (LLA) when compared with that in typically developing children.11 Children with LLA have been found to compensate for the functional loss of one or more joints by increasing COP displacements more than able-bodied subjects during quiet standing.7,12
Although it has been documented that backpacks have an effect on the posture and mobility of able-bodied children, there is not sufficient evidence of how the posture of children with unilateral LLA will change during backpack carriage due to their altered afferent feedback and lower-limb musculature. Therefore, the aim of this study was to assess the effect of backpack loads on postural sway in pediatric subjects with LLA compared with those of able-bodied subjects of the same age.
MATERIALS AND METHODS
Four children with unilateral Syme amputation (mean age, 12.25 ± 2.2 years) along with five age-matched typically developing controls (mean age, 12.8 ± 4.1 years) participated in this study (Table 1). The cause of amputation was congenital fibular deficiency in three subjects and trauma in one.
Potential subjects were screened through the use of a questionnaire to give an assessment of health and fitness. Individuals with current musculoskeletal injuries were excluded. Subjects with amputations had stable volume of their residual limb and used the same prosthesis for the previous year adjusted by a certified prosthetist within the previous 6 months.
The study protocol was approved by the Georgia State University Human Subjects Review Board, and informed assent and parental permission were obtained before the data collection.
Two AMTI force platforms (Advanced Mechanical Technology, Newton, MA, USA) were used to collect ground reaction force (GRF) data during quiet standing at a sampling rate of 600 HZ.
Vicon Peak Motus software (Vicon Motion Systems, Oxford, UK, Version 8.0) was used to convert analog signals to digital signals and calculate individual force plate COP data.
All subjects were asked to perform five 40-second trials of quiet standing with incremental changes in load using a custom-designed backpack. The backpack contained an adjustable double-strap suspension system to center the bag between the T11/T12 vertebrae and a centrally located cylinder upon which barbell weights were placed (Figure 1). Trial loads started at 0 N then were randomly changed to the levels of 0%, 10%, 20%, and 25% of the subject's total body weight (% BW). The maximum weight of 25% BW was set below the load that is normally defined as the weight in which backpack load induces back pain.13
Each subject's shoulder width was measured, and two markers were placed at that width, one on each force platform. Subjects were all required to wear soft-soled tennis shoes with each foot centered on a marker, and foot rotation was monitored to ensure a natural lateral foot progression angle (Figure 1). All the participants were instructed to look straight ahead at a marker on the wall during data collection. They were not notified when data collection began or ended to reduce the possibility of subject bias.
To evaluate the standing balance control, the middle 20 seconds (12,000 data points) of each trial were used to calculate a series of time-domain area and distance measures of COP.14
Among the time area measures, total excursion of COP (TOTEX) was defined as the total length of the COP path and calculated by the sum of the distances between consecutive points on their respective paths in the resultant distance time series.14 The total excursion gives a measurement that can display the total sway during each backpack weight. In addition, mean velocity of COP (MVELO) was defined as the average velocity of the COP excursion for the totality of the trial using Equation 114:
TOTEX and MVELO were also calculated separately for anterior-posterior (AP) and medial-lateral (ML) directions. Table 2 demonstrates a list of study parameters and their abbreviations.
To assess the time-domain area measures of COP, the 95% confidence ellipse area (AREA-CE) was determined by the method described by Prieto et al.14 This was defined as an ellipse with diameters that contain 95% of the points on the postural sway path.
Asymmetries in quiet standing were measured by the weight-bearing contribution of each limb for the children with limb loss group and control group, which was determined by dividing the mean of the vertical GRFs of one leg by the sum of the mean vertical GRFs of both legs. The weight-bearing contribution of the ipsilateral leg (amputated) is defined by Van Asseldonk et al.15
Values of mean and standard deviation (SD) for all parameters are represented in Table 3. The COP distance variables increased in both groups as backpack weight increased. Although the control group showed the expected increase in COP excursion and velocity with each successive weight level, the limb loss group did not. Instead, for most variables, the children with limb loss increased from 0% BW to 10% BW, decreased as the backpack weight increased from 10% BW to 20% BW, and increased again as the backpack weight increased from 20% BW to 25% BW (Figures 2–4).
In addition, the 95% confidence ellipse showed differences in shape. The sway path of the control group moved in the AP direction as backpack weight increased (Figure 5A), whereas the sway path of the limb loss group moved in the ML direction as backpack weight increased (Figure 5B).
Weight-bearing asymmetry was small. The mean weight-bearing contribution was in the range of 45% to 54% for all subjects. Subjects maintained their level of weight-bearing symmetry as backpack weight was increased (Figure 6).
The purpose of this study was to understand static balance strategies in children with unilateral LLA while carrying heavy backpacks. As indicated by the model of quiet standing presented by Winter et al.,16 the normal body maintains balance via two distinct control mechanisms: AP COM position is maintained by the action of the ankle dorsiflexion/plantar flexion mechanism, and ML COM position is maintained by the hip adductor/abductor mechanism. A posterior load (such as a weighted backpack) introduces an external joint moment at the ankle and hip, and both joints can contribute internal moments to counteract the load. Results from this study indicate that children with limb loss may adopt a different strategy to maintain balance in the presence of a heavy posterior load, possibly due to the altered proprioception or musculature in the residual lower limb.
Findings showed that the COP distance measures during static balance increased in both groups with increasing carriage weight. However, nonlinearity was seen in parameters with respect to increasing backpack load in the limb loss group. The local maximum at 10% BW load was present in almost every measure of sway, even though the order of loads was randomized. This could possibly be an indication of changing control strategies due to a change in the proprioceptive capabilities of the neuromuscular control system. A backpack weight of 10% may not cause a definitive shift in AP COM position, yet it may still be enough of a load to cause an increased reaction of the ML control mechanism; therefore, the postural control system that is used to maintain COP position may still be solely dependent on the ML hip abduction/adduction system. This may cause an overreaction of the system and a larger increase in COP motion. A backpack weight of 20% might cause enough of a shift of the COM in the AP direction to initiate the AP control mechanisms of the contralateral limb and therefore reduce the dependence on the ML control system, thus causing more controlled posture and a subsequent decrease in the COP motion. A further increase in backpack weight (25%) might then cause more instability and the consequential increase in COP motion witnessed during these trials. It is not known when each system initiates control; therefore, there may be a certain threshold of COM motion that causes a reaction in the AP or ML control systems.
Figure 5 shows a difference in the direction of postural sway displacement and subsequent area of the 95% confidence ellipse. In the control group, the area of the 95% confidence ellipse extends along the AP axis at backpack weights of 0% BW, 10% BW, and 25% BW, whereas the 95% confidence ellipse increases in the ML direction for these same backpack weights in the limb loss group. It is possible that the increase in ML COP movement in the limb loss group indicates a greater dependence on the hip abductor/adductor muscle pair to maintain COP position and control COM position. It may be the case that the limb loss group uses a completely different postural control system than that seen in typically developing individuals. This larger movement of the control group COP in the AP direction may be a consequence of the “search process” of sway, indicating a more active search in the direction that the subject is expected to move, which in quiet standing is in the AP direction.2 The larger ML range of motion and increased sway in that direction during quiet standing trials indicate a more active search of the postural control mechanism in this direction.
Children with unilateral LLA showed bilateral symmetry in maintaining static balance, even in the presence of heavy backpack loads. The children in the present study had amputations between the ages of 6 months and 3 years. Therefore, they have been accustomed to their residual limb either before or near the onset of walking. Static balance is a learned motor skill that begins at the age of two and continues to develop through young adulthood.17 Although children with limb loss display a lack of certain anatomical structures and requisite proprioception that have been shown to affect both gait and quiet standing,1,18 they appear to be able to adapt and compensate for the single anatomical limb loss, which allows them to function similar to their able-bodied peers. Another possible reason could be the effect of the long residual limb length (Syme amputation) of the pediatric subjects in this study. Lenka et al.19 found that postural stability in terms of COP measures is higher in adults with amputation with longer residual limbs.
This study was limited to the small number of pediatric subjects with transtibial Syme amputations. Results are difficult to generalize given the small sample size and the focus on a specific level of amputation.
Children with limb loss adapt to increasing posterior loads differently than able-bodied counterparts by relying on ML shifts in COP as opposed to the expected AP shifts. In addition, the unexpected local maximum in postural sway outcome measures at 10% BW warrants further investigation to determine underlying neuromuscular balance control mechanisms.
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