One-third of the adults in the United States are obese (1). Obesity is linked to a number of gait alterations, including slower gait speed (2,3), shorter step length (4), longer duration stance phase and double support (3,5,6), and difficulty walking long distances (7). Possible reasons for these gait alterations include range-of-motion limitations (3,6), knee extensor dysfunction (8), and neuromuscular adaptations to reduce energy expenditure (2,4), reduce knee joint loading (2,5,6), or maintain balance (3,4).
Another possible reason for obesity-related gait alterations is that individuals who are obese walk with greater relative effort. Relative effort has been calculated by expressing joint torques during an activity as a percentage of strength, or maximum voluntary torque (9–11). Relative effort during gait may be higher among individuals who are obese because they exhibit higher joint torques at the ankle (5,6), knee (5), and hip (5). Although lower limb strength is also higher among individuals who are obese in knee extension (12–14), hip flexion (13), and ankle dorsiflexion (13), relative effort may still be higher if obesity-related increases in joint torques surpass increases in strength. For example, Browning and Kram (5) reported a 51% obesity-related increase in peak knee extension torque during 1.5 m·s−1 gait, while Koushyar et al. (13) reported a 30% obesity-related increase in knee extension strength, which would together suggest an 11% obesity-related increase in knee extension relative effort during gait. Prior studies have quantified relative effort during sit-to-stand (9–11), stair ascent and descent (10,15), and gait (15–17). However, we are not aware of any studies that have investigated the effects of obesity on relative effort during gait. Identifying any such effects could help clarify the underlying factors by which obesity compromises walking ability and inform interventions to improve walking capacity among individuals who are obese.
Aging has also been associated with gait alterations, including a lower gait speed (18), shorter step length (16,18), and smaller ranges of motion at the ankle and hip (16,18). These alterations have been attributed, in part, to walking with greater relative effort (15,16). For example, Anderson and Madigan (16) reported relative effort of the ankle plantar flexors to be close to 100% among older adults walking at a hurried speed, indicating that near full effort was being exerted. Older adults who are obese would seem to be particularly susceptible to greater relative effort during gait due to obesity-related increases in joint torques, and age-related loss of strength (13,19).
The goal of this study was to investigate obesity-related and age-related differences in relative effort during the support phase of gait. We focused on women because they have a higher prevalence of obesity (1) and obesity-related mobility limitations (20) than men. We hypothesized that: 1) relative effort would be higher among women who were obese compared with healthy-weight women, 2) relative effort would be higher among older women compared with young women, and 3) obesity-related differences in relative effort would be larger among older women than young women. We evaluated subjects during self-selected gait, as well as during a constrained gait (fixed speed and step length across all subjects), due to the potential for obesity-related differences in gait characteristics from confounding any obesity-related differences in relative effort.
Thirty-nine women were recruited to form four groups: 10 young (18–30 yr) healthy-weight (body mass index, 18–24.9 kg·m−2), 10 young obese (body mass index >30 kg·m−2), 10 older (65–80 yr) healthy-weight, and 9 older obese. A medical screening was used to exclude individuals with self-reported neurological, cardiac, or musculoskeletal conditions that affected balance or walking ability, or >2.3-kg change in body mass over the prior 6 months. Subjects also completed the Godin leisure-time exercise questionnaire (21) to quantify their physical activity level, which was viewed to potentially influence relative effort. Body fat percentage was estimated using a Lange skinfold caliper (Cambridge Scientific Industries, Cambridge, MA) and the manufacturer’s recommended usage. All but one young healthy-weight and one older healthy-weight subject were right-foot dominant (preferred to kick a ball). The study was approved by the university institutional review board, and written informed consent was obtained from all subjects prior to participation.
Two experimental sessions were used to minimize potential fatigue development from completing all strength testing during a single session. In the first session, subjects completed all gait trials as described below, and knee strength measurements. In the second session, subjects completed ankle and hip strength measurements.
At the beginning of the first session, subjects donned compression shorts and walking shoes (same make/model for all subjects). All gait trials were performed along a 10-m level walkway, with self-selected trials performed before constrained gait trials. During self-selected trials, subjects were instructed to walk naturally and to look straight ahead. Up to seven practice trials were performed to acclimatize subjects to the experimental surroundings and to identify the proper starting position along the walkway so their right foot naturally and consistently landed on a force platform (Model 9090; Bertec Corporation, Columbus, OH) integrated into the middle of the walkway. Five trials were then performed with proper foot placement on the force platform. During constrained trials, mean speed during each trial was constrained to 1.5 ± 0.05 m·s−1, via verbal feedback after each trial, and measured using a reflective marker on the subject’s back. Step length was constrained to 0.65 m, by asking subjects to step on markings spaced at this distance along the walkway. The chosen speed and step length were intended as representative of self-selected values among young adults (18,22,23). Five constrained trials were performed.
After completing all gait trials in the first session, knee strength was measured in extension and flexion using a commercial dynamometer (System 3; Biodex Medical Systems, Inc., Shirley, NY). Measurements were collected from the right knee while in a seated position with the hip flexed 70°. Relaxed trials were first performed to measure the passive elastic/gravitational torque over the entire range of motion. During these trials, subjects were instructed to remain relaxed while the dynamometer attachment moved at 5°·s−1 throughout their individual joint range of motion at least three times. Next, subjects performed isometric maximum voluntary contractions in knee extension and flexion at four joint angles individualized to each subject’s ROM. We found the range of motion (max flexion to min flexion), subtracted 10° from it, and then divided that by 3. For example, if we consider the range of motion to be 130°, the difference in the angles would be (120/3 = 40°). The angles would be 5°, 45°, and 85°, and 125° knee flexion. For all strength measurements, one practice trial was performed first, followed by three actual trials with ~30 s of rest between trials. The actual trial with the highest torque was used for further analysis.
In the second session, ankle and hip strength in extension and flexion were measured. The testing protocol was similar to knee measurements, except for differences in posture. Ankle measurements were completed first while subjects were in a seated position, with the knee flexed 50° and the hip flexed 80°. Hip measurements were taken while subjects were in a standing position, using a custom setup (19), with the knee held near full extension by a knee immobilizer. A 10-min rest period was given between ankle and hip measurements.
Strength measurements were used to individualize joint-specific models of strength for each subject, and were then used to predict strength as a function of joint angle. By doing so, we accounted for known variations in strength with joint angle. During strength measurements, dynamometer attachment angle, angular velocity, and torque were sampled at 200 Hz and low-pass filtered at 5 Hz (fourth-order zero-phase shift Butterworth filter). Passive elastic and gravitational torque was estimated by fitting a line (least squares) to torque data from relaxed trials throughout the range of motion, and was subtracted from each maximum voluntary contraction trial (24) to find the active component of joint torque. Both active and passive components of the joint torque/strength measurements were used to model strength as a function of joint angle (19) for each joint (hip, knee, and ankle) and exertion (flexion and extension) combination.
During gait trials, ground reaction forces were sampled at 1000 Hz from the force platform, low-pass filtered at 40 Hz (fourth-order zero-phase-shift Butterworth filter), and down sampled to 100 Hz. Segmental kinematics were collected using reflective markers placed bilaterally at the acromion process, anterior superior iliac spine, posterior superior iliac spine, greater trochanter, lateral femoral epicondyle, lateral malleolus, calcaneus, and head of the fifth metatarsal. In addition, clusters of three markers were placed mid-shank and mid-thigh. Markers positions were sampled at 100 Hz using a six-camera motion capture system (MX-T10; Vicon Motion Systems Inc., Los Angeles, CA) and low-pass filtered at 7 Hz (fourth-order zero-phase-shift Butterworth filter).
Sagittal plane joint torques were calculated at the ankle, knee, and hip using a two-dimensional inverse dynamics analysis and custom-written code (Matlab 2013a; The Mathworks Inc., Natick, MA). Segmental inertial parameters were estimated using established methods (25), and hip, knee, and ankle joint centers were identified using a functional approach (26). For each subject, one representative self-selected trial and one constrained trial was used for the analysis. The former was identified as having the minimum deviation from the subject’s mean speed across their self-selected trials, and the latter as having the minimum deviation from the target speed (1.5 m·s−1). Relative effort at the ankle, knee, and hip was determined at each sample (100 Hz) during stance phase, as the ratio of the joint torque to strength, the latter being predicted from the aforementioned model of strength based on joint angle (16). For example, when knee torque indicated an extensor-dominant torque, relative effort was calculated using this torque value and the subject’s knee extensor strength at the corresponding knee angle that was predicted by their individualized knee extensor strength model. The same approach was used to calculate relative effort at the hip and ankle using separate individualized strength models for these joints.
Two-way analyses of covariance were used to investigate differences in hip, knee, and ankle peak relative effort between obesity groups (healthy-weight or obese), age groups (young or older), and their interaction during the stance phase of gait. Godin scores of leisure time physical activity level were used as a covariate to account for variations in physical activity that may influence relative effort. To provide insight on potential mechanisms underlying any obesity or age-related differences in peak relative effort, we also performed an identical analysis on torque at peak relative effort (the numerator of the relative effort calculation), and strength at peak relative effort (the denominator of the relative effort calculation that was predicted by the strength model). Because we hypothesized (and anticipated) obesity–age interactions, we performed separate but identical analyses for self-selected and constrained trials to avoid three-way interactions that may have occurred if both gait conditions were included in the same analyses. In the event of a significant obesity–age interaction, post hoc tests were performed using Tukey’s honestly significant difference test. Effect sizes were quantified using partial eta squared. Statistical analyses were performed using JMP Pro 13 (SAS Institute, Inc., Cary, NC) with a significance level of P < 0.05.
Body mass index and body fat were higher among women who were obese compared with healthy-weight women, and among older women compared with young women (Table 1). Results for self-selected gait are reported here, whereas those for constrained gait are available elsewhere for conciseness (see Document, Supplemental Digital Content 1, text summary and tables of results, http://links.lww.com/MSS/B690). Gait speed did not exhibit an obesity–age interaction, but did exhibit a main effect of obesity in that it was 0.10 m·s−1 slower among women who were obese. Step length did not exhibit an obesity–age interaction, or main effects of obesity or age (Table 2).
Relative effort exhibited similar time-varying trends over stance across subject groups (Fig. 1). Two local maxima/minima (i.e., peaks) occurred at the hip, four at the knee, and two at the ankle. These peaks were the focus of our analysis. Peak relative effort exhibited no obesity–age interactions, but did exhibit main effects of obesity and age (Table 2). Regarding main effects of obesity, peak relative effort among women who were obese was 23% higher in knee extension and 35% higher in ankle plantar flexion compared with healthy-weight women. Regarding main effects of age, peak relative effort among older women was 22% higher in hip flexion and 15% higher in knee extension compared with young women.
Joint torques also exhibited similar time-varying trends over stance across subject groups and exhibited the same local peaks as relative effort (Fig. 1). Joint torque at peak relative effort exhibited an obesity–age interaction for ankle plantar flexion torque (Table 3). This interaction was due to a 47.0-N·m obesity-related increase among young women, and a 27.5-N·m obesity-related increase among older women. Joint torque at peak relative effort also exhibited main effects of obesity and age. Regarding main effects of obesity, joint torque at peak relative effort among women who were obese was 20% to 24% higher at the hip, and 29% to 58% higher at the knee compared with healthy-weight women. Regarding the main effect of age, joint torque at peak relative effort among older women was 38% lower in knee flexion compared with young women.
Strength at peak relative effort exhibited no obesity–age interactions, but did exhibit main effects of obesity and age (Table 4). Regarding main effects of obesity, strength at peak relative effort among women who were obese was 17% to 18% higher at the hip, and 15% higher in ankle dorsiflexion compared with healthy-weight women. Regarding main effects of age, strength at peak relative effort among older women was 18% to 26% lower at the hip, 25% to 38% lower at the knee, and 20% lower in ankle dorsiflexion compared with young women.
Results from constrained gait were generally similar to those from self-selected gait, particularly for torque and strength at peak relative effort. Peak relative effort during both gait conditions was higher among women who were obese in ankle plantar flexion compared with healthy-weight women, and higher among older women in hip flexion and knee extension compared with young women. Torque at peak relative effort during both gait conditions was higher among women who were obese in hip extension and flexion, knee extension and flexion, and ankle plantar flexion. Torque at peak relative effort during both gait conditions was also lower older women in ankle plantar flexion compared with young women. Strength at peak relative effort during both gait conditions was higher among women who were obese in hip flexion and ankle dorsiflexion compared with healthy-weight women, and lower among older women in hip extension and flexion, knee extension, and ankle dorsiflexion compared with young women. Results are summarized, and effect sizes are reported, for self-selected gait in Table 5, and for constrained gait in Table S4, http://links.lww.com/MSS/B690.
The goal of this study was to investigate obesity-related and age-related differences in relative effort during the support phase of gait. Our first hypothesis was that relative effort would be higher among women who were obese compared with healthy-weight women. This hypothesis was supported because peak relative effort was 20% to 48% higher among women who were obese in hip flexion, knee flexion, knee extension, and ankle plantar flexion. This greater relative effort appeared to generally result from joint torques exhibiting greater obesity-related increases than strength (Tables 5 and S4, http://links.lww.com/MSS/B690). Our second hypothesis was that relative effort would be higher among older women compared with young women. This hypothesis was supported because relative effort was higher among older women in hip flexion and knee extension. In contrast to obesity-related increases in relative effort, age-related increases in relative effort seemed to generally result from age-related reductions in strength at peak relative effort, and less due to age-related differences in joint torques. Our third hypothesis was that obesity-related differences in relative effort would be larger among older women than young women. This hypothesis was not supported because peak relative effort did not exhibit any obesity–age interaction effects. Given that obesity-related differences during constrained gait, when speed and step length were fixed, were generally similar as during self-selected gait, the greater relative effort among women who were obese was not attributed to differences in these gait characteristics. Overall, our results suggest that women who are obese use a greater proportion of their strength while walking than healthy-weight women, and this obesity-related effect occurred similarly among young and older women.
Several limitations to this study should be noted. Relative effort exceeded 100% for 62 of 296 (21%) peak relative effort values estimated during self-selected gait. Most of these were for ankle plantar flexion (58%) and knee extension (18%). Several sources of error may have contributed. First, the torque-producing capability of biarticular muscles (e.g., gastrocnemius) is dependent upon kinematics at both spanned joints (e.g., ankle and knee), yet the models we used to predict strength only accounted for kinematics at one joint (e.g., ankle). Second, inaccuracies in joint torque estimates using a two-dimensional inverse dynamics analyses, such as those resulting from the use of segment inertial estimates and skin surface markers to represent joint axes. Third, inaccuracies in strength measurements, such as those resulting from misalignment between axis of rotation of dynamometer and joint and submaximal effort. The consequence of relative effort exceeding 100% was that the relative effort at which strength was fully utilized was unclear. Nevertheless, relative effort still provided a quantitative measure of effort relative to maximum capacity, and thus is still considered relevant to evaluate obesity-related differences. Another limitation of our methodology was that we did not account for the effects of joint angular velocity on strength in our strength models that were used to estimate relative effort. We chose this approach because our initial attempts to account for both seemed overly sensitive to joint angular velocity, particularly for higher velocities at which our strength model predicted a relatively small strength value. Lastly, as with any cross-sectional experimental design, differences between subject groups other than obesity and age may have contributed to the differences identified.
Some differences in relative effort are apparent between the current values and those reported elsewhere. Peak relative effort among young adults found here during self-selected gait was consistently higher than values reported elsewhere. In particular, mean peak relative effort in hip extension was 42% here versus 30% elsewhere (16,17), in hip flexion was 57% here and ~30% elsewhere (16), and in ankle plantar flexion was 124% here and 58% to 86% elsewhere (16,17). By contrast, peak relative effort among older adults found here during self-selected gait was generally lower than values reported by Samuel et al. (15) In particular, mean peak relative effort for hip extension was 51% here and 127% elsewhere, for hip flexion was 80% here and 68% elsewhere (15), for knee flexion was 58% here and 75% elsewhere (15), and for knee extension was 88% here and 101% elsewhere (15). These differences in peak relative effort between studies were likely due to methodological differences, including differences in gait speed, subject sex distributions, strength measuring protocols, and methods to account for the effects of joint angle and angular velocity on strength.
Obesity and age-related increases in relative effort occurred at important times during the gait cycle when mechanical energy is being generated. For example, the ankle plantar flexors provide support and propulsion during push-off (27,28) during late stance and produce over two thirds of the total mechanical energy generation by the lower extremities during a gait cycle (6). The hip flexors generate mechanical energy during late stance and early swing. Moreover, increasing gait speed requires an increase in mechanical energy generation in hip flexion, ankle plantar flexion, knee flexion, and knee extension (29).
Walking with greater relative effort can compromise walking in multiple ways, even if relative effort does not reach its theoretical maximum. First, given that joint torques (and therefore relative effort) increase as gait speed increases (5), walking with greater relative effort suggests less reserve capacity to increase gait speed, or to perhaps execute a rapid stepping response to prevent a fall after balance perturbations such as tripping or slipping. Second, walking with greater relative effort can result in earlier fatigue (10,30), which may, at least partly, explain why individuals who are obese adopt a slower gait speed and have greater self-reported difficulty in walking longer distances (2,7). Third, a higher level of muscle force exertion than may occur with greater relative effort can increase in muscle force variability (31,32). Such variability has been considered as deleterious noise in the neuromuscular system (33) that can lead to undesirable inaccuracies in goal-oriented movements (34,35). Our results therefore provide mechanistic support of the potential limiting effect of strength on gait among women who are obese.
Women who were obese walked with greater relative effort than healthy-weight women. Although both joint torques and strength were higher among women who were obese, the greater relative effort was attributed to greater obesity-related increases in joint torques than strength. Additionally, older women walked with greater relative effort than young women, with these age-related increase generally resulting from age-related reductions in strength and not age-related increases in joint torques. These results may help to explain the compromised walking ability among women who are obese, as well as among older women.
This work was supported by award R01OH009880 from the Centers for Disease Control and Prevention (CDC). The contents of this article are solely the responsibility of the authors, and do not necessarily represent the official views of CDC.
Conflict of Interest: The authors have no conflicts of interest to report. The results of the present study are presented clearly, honestly, and without fabrication, falsification, or inappropriate data manipulation. In addition, these results do not constitute endorsement by ACSM.
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