Ankle sprains are one of the most common injuries in competitive and recreational athletics (12,16,18). In collegiate sports, injuries to an ankle ligament account for 15% to 45% of all injuries (8,16,27), and the athletes most commonly affected are participants in jumping and landing activities such as volleyball and basketball (16,18).
Prevention of ankle sprains is important to retain health and encourage activity adhesion in trained and recreational athletes. As a result of the frequent occurrence of ankle injury, several methods of preventing sprains have been developed including joint bracing. Brace manufacturers design ankle braces in an attempt to prevent sprains or to provide added support to an injured ligament. The braces available use various designs intended to essentially restrict the inversion range of motion (2,10) or improve proprioception in the ankle joint (9,22).
Ideally, the mechanical goal of an ankle brace is to externally support the ligamentous structures of the ankle and reduce inversion and eversion at the end ranges of motion only, just before ligament failure. In turn, the brace should not restrict the normal frontal and sagittal plane motions (12) that contribute to energy absorption and generation. Thus, the ideal ankle brace would protect the ankle from lateral ligament injury without hindering the joint to perform its required functions.
Many athletes and even whole sports teams have taken the idea of the general usage of ankle braces after an injury a step further and use braces as protective measures or prophylactic devices to prevent sprains from occurring. Teams wear the braces throughout an entire season of practices and games even if players have never experienced a previous injury. Studies have shown that the use of prophylactic ankle braces effectively reduce the frequency of ankle inversion sprains in athletes (14,21,25,26,29). However, although it is generally accepted that the ankle braces are effective at reducing frontal plane motion, some researchers suggest that the braces may also reduce ankle sagittal plane motion as well (2,3,19,24,25). The primary role of sagittal plane ankle movements during landing, specifically dorsiflexion, is to absorb kinetic energy, attenuating the forces at impact (15). Essentially, the muscles that eccentrically control ankle dorsiflexion and knee and hip flexion during drop landings act to absorb the majority of the kinetic energy present in the moving body (5). The remaining energy may be absorbed by various other structures including articular cartilage, bone, and other soft tissue structures.
The three joints that primarily contribute to attenuating a typical landing impact are the ankle, knee, and hip. These are followed to a lesser extent by the spine, intervertebral disks, and musculature of the trunk. Because the human body is a kinetic chain, it is thought that if one of the major joint contributors was compromised during a landing task, the energy would be transferred proximally up the leg to the next available joint (5–7,15). Therefore, it is important to study energy absorption by calculating the amount of work done by the joints to discover how the application of an ankle brace affects the absorption characteristics of the ankles, knees, and hips. Furthermore, it is important to calculate impulse to determine whether the braces have an effect on the pattern of joint contributions to energy absorption.
The purpose of this study was to quantify lower extremity joint contributions to energy absorption during single-legged drop landings in two ankle-braced conditions and one nonbraced condition. We hypothesized that the total energy absorbed by the lower extremities would not be different between the braced and unbraced conditions. However, when braced, the ankle would absorb less energy during drop landings compared with the unbraced condition. Subsequently, the energy absorbed by the knee and/or hip joints would be increased to compensate for the reduced energy absorption at the ankle in the braced conditions. Furthermore, we hypothesized that between the two ankle braces, the brace with less restriction in the sagittal plane would allow greater energy absorption by the ankle and therefore less energy absorption by the knee and hip.
Eleven female college students (age = 22.3 ± 1.7 yr, height = 1.66 ± 0.04 m, mass = 58.43 ± 5.83 kg) participated in this study. All subjects were free of lower extremity injury within the past 6 months, were not diagnosed with chronic lateral ankle instability, and had never sustained an injury to the lower extremities that required surgery. All subjects had a history of participation in recreational landing sports such as basketball or volleyball, defined as having participated in high school athletics and/or current athletic participation at least 2 d·wk−1. The experimental tasks and risks were explained, and each participant read and signed an informed consent before participation. All methods were approved by the university institutional review board for testing of human subjects.
Ground reaction force (GRF) data were sampled at 1000 Hz using an Advanced Mechanical Technology Incorporated (Newton, MA) force plate. GRF data were synchronized with the kinematics using the Vicon Nexus platform (Vicon Motion Systems, Oxford, United Kingdom). Kinematic data were recorded using 10 Vicon high-speed infrared cameras sampling at 200 Hz and digitized using the Plug-in Gait module in Vicon Nexus (Vicon Motion Systems).
Data collection took place in a university biomechanics laboratory. Each participant wore athletic shorts, a sports bra (or tied up her T-shirt), and jumped barefoot during data collection. The participant’s height and mass were recorded, and lower extremity anthropometric measurements were taken including knee width, ankle width, and the length of both legs. The participants performed a self-selected warm-up with stretching before the application of reflective markers.
The participant’s skin was prepped at specific landmarks for the reflective markers (1.4-cm diameter) using 70% isopropyl rubbing alcohol pads. The reflective markers were placed by the same researcher on both the right and left sides of the body according to the Vicon Nexus Plug-in Gait model (Vicon Motion Systems). The toe markers were placed on the surface of the participant’s skin, whereas the lateral ankle markers were placed on the skin in the unbraced condition and on the braces in the braced conditions. To ensure consistent replacement, the bony landmarks were palpated through the brace material and placed on the outside of the brace as closely as possible to the point of palpation.
Participants drop jumped off a platform 0.33 m high. Verbal instructions and a jump landing demonstration were given to each participant. The participants were instructed to hold their arms straight out in front throughout the duration of the landing so that the arms did not impede the camera’s view of the reflective markers. They were instructed to stand at the edge of the platform and to lift their nonpreferred leg to stand on one leg. They waited for the investigator to say “Go,” then leaned their center of mass forward until they began to fall. At that point, they were to quickly remove their preferred foot from the drop platform and land with their preferred foot on the force plate (Fig. 1). Participants practiced two landings before data collection. Leg preference was determined by asking the participants to perform single-leg drops off a small height at their convenience before reporting to the laboratory for data collection and identifying the leg on which they felt most comfortable landing.
The verbal instructions for the landing consisted of telling the participants to land and then extend their leg at the end of the jump so that they were standing up straight. They were instructed to remain standing on one leg until the investigator said “Stop,” which occurred at the end of a 3-s capture time. Each participant performed 10 successful drop jumps per condition—one unbraced and two braced conditions—for a total of 30 landings. Braced conditions were randomized for each participant, and each trial was separated by a rest period selected by the participant to minimize the effects of fatigue. The two braces worn by the participants were the DonJoy (DJO Incorporated, Vista, CA) Stabilizing Ankle, a soft lace-up boot (“boot”), and the DonJoy (DJO Incorporated) Velocity, a brace hinged at the ankle with a locking strap that surrounds the calf and a soft cuff with laces that cover the anterior surface of the joint (“hinge”). Each brace was placed on the participant’s ankles by the same researcher as per the brace manufacturer’s instructions.
Both GRF and kinematic data were smoothed using a Butterworth filter set at 10 Hz (1). Ankle, knee, and hip joint moments were calculated using standard inverse dynamics techniques from the digitized data using the Vicon Nexus Plug-in Gait module. Joint powers in the sagittal plane were calculated using custom-written software in QuickBasic (joint power = joint moment of force × joint angular velocity). The landing phase was defined as the period from initial contact with the force plate until the participant reached maximum knee flexion (32). The moment–time and power–time curves for each landing were integrated to calculate angular impulse (a measure of the change in angular momentum) and work (a measure of energy generation or absorption), respectively. A negative work value indicates eccentric muscle activity across the three joints with the knee and hip flexing against an extensor joint moment and the ankle dorsiflexing against a plantarflexion joint moment. Relative work at each joint was calculated by dividing the individual joint work values by the total work of the lower extremity then multiplying by 100.
Each participant’s 10-trial mean value of dependent variables at the ankle, knee, and hip joints (hip, knee, and ankle joint impulse; hip, knee, ankle, and total work; hip, knee, and ankle joint relative work) for each stabilizer condition was entered into a one-way repeated-measures ANOVA (α = 0.05). A Bonferroni post hoc analysis was used to determine the source of significance when appropriate. Data were analyzed using SPSS 16.0 (SPSS Incorporated, Chicago, IL). In addition, Cohen d effect sizes were calculated between each of the conditions as a measure of the strength of relationship between variables. In accordance with Cohen (4), the levels of effect size were deemed small (d = 0.2), medium (d = 0.5), or large (d = 0.8).
For presentation of the joint moment and joint power data, grand ensembles were created for the variable–time curves. Ensemble curves for each subject at each joint were calculated by using custom software to interpolate 101 data points between initial ground contact and maximum knee flexion and 10 data points during the 50 ms before ground contact. Grand ensemble curves were then calculated as the mean value of each of the 111 points across all trials of all subjects. The support moment was calculated as the sum of the ankle, knee, and hip moment ensemble curves and reflects the overall performance of the lower extremity in providing support (30). Similarly, support power was calculated as the sum of the ensemble power curves of the individual joints to reflect the overall performance of the lower extremity as an energy absorber. The term “support power” was used to remain consistent with the term “support moment” proposed by Winter (30).
Descriptive statistics (mean and SD) for individual joint impulse values during the landing are presented in Table 1, and Cohen d effect sizes for the condition comparisons are presented in Table 2. The grand ensemble joint moment–time curves are presented in Figure 2. Lower extremity joint impulse values were not significantly different across the three bracing conditions: ankle (P = 0.49), knee (P = 0.70), and hip (P = 0.57) (Table 1).
Descriptive statistics (mean and SD) for individual joint work values during the landing are presented in Table 1, and Cohen d effect sizes for the condition comparisons are presented in Table 2. The grand ensemble joint power–time curves are presented in Figure 3. As hypothesized, there were no significant differences among the conditions for total work performed by the lower extremity (P = 0.06). There was a moderate effect size between the two bracing conditions (P = 0.17, Cohen d = 0.73) and a small effect size between the hinged-brace condition and the unbraced condition (P = 0.21, Cohen d = 0.43).
A significant main effect of the brace condition was identified for the eccentric ankle joint work (P = 0.01). A Bonferroni post hoc analysis revealed that the ankle joint performed significantly less work in the boot brace condition (1.10 ± 0.30 J·kg−1) than in both the unbraced (1.28 ± 0.34 J·kg−1, P = 0.03, Cohen d = 0.56) and hinged-brace (1.27 ± 0.34 J·kg−1, P = 0.04, Cohen d = 0.53) conditions. There was no significant difference between the unbraced and hinged-brace conditions (P = 1.0, Cohen d = 0.03).
There was not a significant main effect of ankle bracing identified (P = 0.26) for knee work. In addition, the effect sizes for each variable pair were small (Table 2) according to Cohen d.
A significant main effect of the brace condition was identified for hip work (P = 0.01). A Bonferroni post hoc analysis revealed that in the hinged-brace condition, the hip performed significantly more negative work (0.18 ± 0.14 J·kg−1) than in the unbraced condition (0.13 ± 0.12 J·kg−1, P = 0.01, Cohen d = 0.38). However, the energy absorbed by the hip in the boot condition (0.16 ± 0.13 J·kg−1) was not statistically different from that in the unbraced condition (P = 0.26, Cohen d = 0.24).
Relative joint contributions
Descriptive statistics (mean and SD) of the relative work contributed by each joint are presented in Table 1. There was a significant effect of ankle bracing on the relative work performed at the ankle (P = 0.02), the knee (P = 0.04), and the hip (P = 0.02) joints. Post hoc analyses revealed that the ankle plantar flexors performed approximately 5% less relative work in the boot brace (37.72% ± 11.75%) than in the unbraced condition (42.69% ± 11.22%, P = 0.04, Cohen d = 0.43). The relative work of the knee extensors (56.95% ± 12.95%) was increased compared with the unbraced condition (52.87% ± 12.55%); however, the difference was not statistically significant (P = 0.08), and there was a small effect size (Cohen d = 0.32). When compared with the unbraced condition (4.45% ± 4.05%), the relative energy absorbed by the hip extensors in the boot brace condition (5.33% ± 4.69%, mean ± SD) was not significantly different (P = 0.14, Cohen d = 0.20).
In the hinged-brace condition, the ankle plantar flexors performed similar relative work (40.32% ± 10.62%) compared with the unbraced condition (42.69% ± 11.22%, mean ± SD). The difference was not statistically significant (P = 0.40), and there was a small effect size (Cohen d = 0.22). Relative energy absorption by the knee extensors was not different in the hinged-brace (53.95% ± 11.75%) compared with the unbraced condition (52.87% ± 12.55%, P = 1.0, Cohen d = 0.09). The relative work of the hip extensors in the hinged-brace condition (5.75% ± 4.95%) increased by about 1.3% when compared with the unbraced landing condition (4.45% ± 4.05%), a difference statistically significant (P = 0.04) but with a small Cohen d effect size of 0.29.
A direct comparison of the two bracing conditions revealed no statistically significant differences in energy absorption by the joints of the lower extremity (P values > 0.32, ankle Cohen d = 0.23, knee Cohen d = 0.24).
In 1973, Garrick and Requa (12) emphasized the importance of a prophylactic ankle brace to restrict frontal plane motions at the end range but to not adversely affect sagittal plane motion. This design of a brace would help to prevent injury to the ankle ligaments in inversion and eversion but would not restrict the ankle’s contribution to the attenuation of kinetic energy when landing. The purpose of this study was to quantify lower extremity joint contributions to energy absorption during single-legged drop landings in three conditions (unbraced, a brace with straps that surround the ankle joint, and a hinged brace that allows the ankle joint free movement in the sagittal plane). It was hypothesized that the total energy absorbed by the lower extremity would not change across the conditions but that the energy absorbed by individual joints would be affected. Specifically, we expected to see the ankle perform less eccentric work in the braced conditions with the knee, hip, or both performing more eccentric work to compensate for the reduced work by the ankle.
Our hypothesis that the total energy absorbed by the lower extremity would not be different across the conditions was supported with the lower extremities absorbing 3.02 ± 0.35 J·kg−1 in the unbraced condition, 2.95 ± 0.26 J·kg−1 with the boot brace, and 3.17 ± 0.34 J·kg−1 with the hinged brace (P = 0.06). These results suggest consistent performance of the lower extremity as an energy absorber regardless of ankle bracing.
The total energy absorbed by the joints was about 25% larger than the 2.42 J·kg−1 reported in a similar landing protocol reported by Schmitz et al. (23). The differences in energy absorption may arise from the data filtering techniques. Their GRF data were filtered with a fourth-order Butterworth at 60 Hz, whereas their kinematic data were filtered using the same method at 12 Hz. In accordance with the recommendations of Bisseling and Hof (1), our kinetic and kinematic data were both filtered at the same frequency (10 Hz) with a fourth-order recursive Butterworth filter. The discussion of the effects of data filtering on the calculation of joint kinetics and energetics is ongoing in current biomechanics research.
Our hypothesis that the ankle would absorb less energy in the braced conditions was partially supported. The largest significant reduction at the ankle was in the boot brace condition, which absorbed about 5% less energy than in the unbraced condition. There was no significant difference in ankle energy absorption when landing with the hinged-brace compared with the unbraced condition. The results of the boot brace compared with the no-brace condition supports the hypothesis of McCaw and Cerullo (19), who reported that the use of some ankle braces when landing reduced ankle joint sagittal plane range of motion about 5° and reduced the dorsiflexion angular velocity by about 110°·s−1 compared with an unbraced landing. The authors suggested that a reduction in joint angular velocity reflects reduced energy absorption at the ankle joint. Although we did not report range of motion or angular velocity, we did find that an application of a moderately restrictive boot brace did reduce the amount of eccentric energy absorbed by the ankle. In our study, less negative work was performed at the ankle in the boot brace compared with the hinged brace. A plausible explanation is that a hinged brace is intended to maintain movement in the sagittal plane. Less restriction of ankle dorsiflexion in a hinged brace compared with a boot brace would allow the ankle to dorsiflex through a greater range of motion and not overly restrict ankle joint energy absorption as suggested by McCaw and Cerullo (19).
Our hypothesis that the knee, hip, or both the knee and hip joints would absorb more energy during braced conditions was partially supported, however, not quite as expected. Generally, there was a trend for an increase in knee and hip work in the braced conditions, but the differences were statistically significant only in the hinged-brace condition. We expected the hip to absorb more energy in the boot brace condition than in the hinged-brace condition to compensate for the restricted ankle dorsiflexion in a boot brace. However, it is also important to note that there were fairly large SD associated with each variable, especially at the hip joint.
The high variability in landings may be confirmed by a study that compared lower limb EMG in males and females when landing on one leg (13). The researchers found no differences between genders, although visually, women tend to land with more knee valgus than men. Although knee valgus has been attributed to a multitude of different factors, EMG activity is a plausible explanation. However, the authors attributed their lack of significance to a small sample size and “inherent” variability in the EMG data. It is plausible that EMG data do contain inherent variations such as cross talk or external noise, but much effort has been placed in equipment and skin preparation that these types of problems should be very minimal. We propose that the variations are related to subtle individual changes from landing to landing, resulting in marked fluctuations from the mean value.
It is also possible that the large SD at the knee and hip are present because there is more muscle redundancy at the knee and hip joints compared with the ankle joint. The ankle joint is primarily controlled by the plantar flexors, whereas both the hip and knee joints are controlled by multiple biarticulating muscles such as the rectus femoris and sartorius that conversely affect both joints at the same time (31). Van Ingen Schenau et al. (28) showed that biarticular muscles play an important role in energy transfer from proximal to distal joints in vertical jumping. The authors give an example of the gastrocnemius, which, because of its biarticular characteristic, allows the knee extensors to deliver work during the jump and acts to transfer the work distally to the ankle to be used for plantarflexion.
In the current study, the energy transfer is occurring in the opposite direction, from distal to proximal. As the person is falling and his/her feet begin to touch the ground, the ankle plantar flexors must act eccentrically to keep the ankle from collapsing in dorsiflexion. However, only contracting the muscles that cross the ankle is simply not enough. The knee and the hip must be involved, and if necessary, the musculature of the trunk may need to be used as well. Biarticular muscles that act to plantar flex the ankle also act to transfer the energy proximally to allow the knee extensors to produce work and keep the knee joint from collapsing. This chain of events continues up the body in a coordinated manner until the person’s mass has stabilized.
Although statistically significant differences were not identified in the absolute energy absorbed by the knee, the nearly moderate Cohen d effect sizes in relative energy absorption suggest that there is an effect of brace use. The boot brace condition exhibited a 5% decrease in relative energy absorbed at the ankle joint and a 4% increase in relative energy absorption at the knee. A similar change was evident with the hinged brace, but the differences were smaller (2.3% decrease at the ankle and a 1% increase at the knee). The results indicate that the use of ankle braces alters the relative contributions of the lower extremity joints to energy absorption during landing, but brace design can influence the effects.
On average across the conditions, the ankle performed about 40% of the total work, whereas the knee and hip performed 55% and 5%, respectively. This pattern across the joints is different from that reported in a previous study of one-legged landing (17). Although Kulas et al. (17) did not incorporate ankle braces, their control group exhibited an energy absorption pattern of the ankle, knee, and hip joints each contributing approximately one-third of the total work. Although their participants landed from a similar height (0.30 m), the different landing techniques used in the studies may account for the observed differences. Kulas et al. (17) asked participants to focus more on their trunk flexion while landing, whereas we allowed the participants to self-select landing technique. Because energy absorption at the ankle, knee, and hip joints is affected by the orientation of body segments relative to the GRF during landing (19), the more upright posture of the subjects in our study may have reduced the energy absorbed at the hip joint during eccentrically controlled flexion (Fig. 1).
Our results suggest that brace design affects the energy absorption characteristics of the lower extremity joints in the sagittal plane. The use of a boot brace decreased the absolute and relative contributions of the ankle joint and increased the work of the knee and hip joints during one-leg landing. Effects of the braces on frontal plane motion were not investigated, and therefore, no inferences should be drawn regarding the effect of the braces in this plane. Furthermore, this study was limited by low statistical power because only 11 subjects were analyzed because of laboratory constraints. The low number of subjects should be kept in mind for the interpretation of the results. Further study using more subjects is warranted to evaluate alternative brace designs and to further quantify the effects of ankle bracing on joint energetics in the frontal and transverse planes.
This study demonstrated that ankle, knee, and hip energetics may be altered depending on the design of the ankle brace. More specifically, when braced with a moderately restrictive support such as the boot brace, the ankle’s contribution to energy absorption in the sagittal plane is reduced, and the knee and hip joints’ relative contribution may be increased.
The authors thank Jenny Knutson for her laboratory assistance. Completion of this project was supported by a Jump Rope for Heart grant from the Illinois Association for Health, Physical Education, Recreation and Dance. DonJoy (DJO Incorporated, Vista, CA) provided the ankle braces used in the study.
This study presents no conflicts of interest.
The results of the present study do not constitute endorsement by the American College of Sports Medicine.
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