ESCAMILLA, RAFAEL F.1; ZHENG, NAIQUAN2; IMAMURA, RODNEY3; MACLEOD, TORAN D.4; EDWARDS, W. BRENT5; HRELJAC, ALAN3; FLEISIG, GLENN S.6; WILK, KEVIN E.7; MOORMAN, CLAUDE T. III8; ANDREWS, JAMES R.6,9
Weight-bearing exercises, such as the squat, are commonly used by athletes to train the hip and the thigh musculature. Physical therapists and trainers also have their patients or clients use squatting-type exercises during anterior cruciate ligament (ACL) and posterior cruciate ligament (PCL) rehabilitation to allow them to recover faster and return to function earlier (6,29,37).
Several studies involving barbell and body weight squat exercises reported PCL forces between 300 and 2700 N and no ACL forces throughout the knee range of motion (8,11,12,31,35). In contrast, other squat studies reported relatively low magnitude peak ACL forces between 30 and 500 N approximately between 0° and 60° knee angles and PCL forces approximately between 60° and 120° knee angles (3,14,25,32). These data are supported by other weight-bearing knee flexion studies, with ACL strain occurring at lower knee angles and PCL strain occurring at higher knee angles (9,17). What is consistent in the squat literature is that PCL loading occurs at higher knee angles typically greater than approximately 60°. What is inconsistent in the squat literature is whether or not ACL strain always occurs at smaller knee angles. Part of the inconsistencies in ACL strain during the squat is that some studies estimated ACL strain in vivo using strain sensors inserted within the ACL (3,15), whereas other studies used biomechanical musculoskeletal models to estimated ACL strain (8,11,12,31,35). However, it is clear that when ACL strain does occur, it occurs at smaller knee angles and its strain or force magnitudes are relatively low. Using dynamic optimization techniques, peak ACL forces have been reported to be less than 20 N during body weight squatting (30).
Although the effects of exercise technique variations on cruciate ligament strain while performing the barbell squat have been examined (11,12), there are no studies that have examined the effects of technique variations on cruciate ligament loading while performing the one-leg squat and wall squat. One-leg squat and wall squat exercises are both performed in training and rehabilitation settings. Wall squats can be performed with varying techniques, such as positioning the heels farther or closer to the wall. Positioning the heels farther from the wall typically results in the knees being maintained over the feet at the lowest position of the squat, whereas positioning the heels closer to the wall typically results in the knees moving anterior beyond the toes at the lowest position of the squat. Performing a one-leg squat also causes the knees to move forward beyond the toes at maximum knee flexion. Clinicians and trainers often believe that anterior movement of the knees beyond the toes during the wall squat or one-leg squat may increase cruciate ligament loading, although there are very limited data that support this belief (1). Moreover, it is unclear if the ACL or the PCL is loaded when anterior knee movement occurs.
The purpose of this study was to compare cruciate ligament tensile forces among squat types (wall squat with the feet farther away from the wall-wall squat long; wall squat with the feet closer to the wall-wall squat short; and the one-leg squat) and squat phases (squat descent and squat ascent) at specific knee angles (0°, 10°, 20°, 30°, 40°, 50°, 60°, 70°, 80°, and 90°). It was hypothesized that 1) ACL tensile force would occur at knee angles 30° or less in the one-leg squat and wall squat short; 2) PCL tensile force would occur throughout the knee angle range of motion in the wall squat long; 3) PCL forces would generally be greater in the wall squat long compared with the wall squat short and one-leg squat; 4) PCL forces would generally not be significantly different between the wall squat short and the one-leg squat; and 5) for all three squat types, ACL and PCL forces would generally be greater at specific knee angles during the squat ascent compared with the corresponding knee angles during the squat descent. Quadriceps and hamstrings muscle force magnitudes will also be presented to help better understand ACL and PCL force magnitudes. Understanding how cruciate ligament tensile forces vary among squatting techniques will allow physical therapists, physicians, and trainers to prescribe safer and more effective knee rehabilitation to patients during ACL or PCL rehabilitation.
Eighteen healthy individuals (nine males and nine females) without a history of cruciate ligament pathology participated with an average age, mass, and height of 29 ± 7 yr, 77 ± 9 kg, and 177 ± 6 cm, respectively, for males, and 25 ± 2 yr, 60 ± 4 kg, and 164 ± 6 cm, respectively, for females. All subjects were required to perform wall squat and one-leg squat exercises pain-free and with proper form and technique for 12 consecutive repetitions using their 12 repetition maximum (12 RM) weight.
To control the EMG signal quality, the current study was limited to males and females that had average or below average body fat, which was assessed by Baseline skinfold calipers (Model 68900; Country Technology, Inc., Gays Mill, WI) and body fat standards set by the American College of Sports Medicine. Average body fat was 12% ± 4% for males and 18% ± 1% for females. All subjects provided written informed consent in accordance with the Institutional Review Board at California State University, Sacramento, which approved the research conducted and informed consent form.
Wall squat (Figs. 1 and 2).
The wall squat began with the right foot on a force platform and their left foot on the ground, both knees fully extended (0° knee angle), the back flat against the wall, and a dumbbell weight held in both hands with the arms straight and at the subject's side. From this position, the subject slowly flexed both knees and squatted down until the thighs were approximately parallel to the ground with the knees flexed approximately 90°-110°, and in a continuous motion the subject returned back to the starting position. A metronome was used to help ensure that the knees flexed and extended at approximately 45°·s−1. The surface of the wall was smooth, and a towel was positioned between the wall and the subject to minimize friction as the subject slid down and up the wall. The stance width (distance between inside heels) was 32 ± 6 cm for males and 28 ± 7 cm for females, and the foot angle was approximately 0° (feet pointing approximately straight ahead), and both stance and foot angle were according to subject preference.
The wall squat was performed with two technique variations, wall squat long (Fig. 1) and wall squat short (Fig. 2). The foot position relative to the wall for the wall squat long was determined using a heel-to-wall distance that resulted in the legs being approximately vertical and the knees positioned above the ankles when the thighs were parallel with the ground (Fig. 1), which is commonly recommended by clinicians and trainers. The average heel-to-wall distance for the wall squat long was 45 ± 3 cm for males and 41 ± 3 cm for females. The heel-to-wall distance for the wall squat short was one half the distance of the heel-to-wall distance for the wall squat long. The shorter heel-to-wall distance for the wall squat short resulted in the anterior surface of the knee moving beyond the distal end of the toes 9 ± 2 cm at the lowest position of the wall squat short (Fig. 2).
The one-leg squat started with the subject standing on one leg with the right foot on a force platform, the right knee fully extended, the left knee bent approximately 90°, and a single dumbbell weight held with both hands in front of the chest. From this position, the subject slowly flexed the right knee and squatted down until the right knee was flexed approximately 90-100° with the trunk tilted forward approximately 30-40° (Fig. 3), and in a continuous motion the subject returned back to the starting position. A metronome was used to help ensure that the right knee flexed and extended at approximately 45°·s−1. At the lowest position of the one-leg squat, the anterior surface of the knee moved 10 ± 2 cm beyond the distal end of the toes (Fig. 3).
FIGURE 3-One-leg squ...Image Tools
Each subject came in for a pretest 1 wk before the testing session. The experimental protocol was reviewed, the subject was given the opportunity to practice the one-leg squat and wall squat exercises, and each subject's heel-to-wall distances for the wall squat short and wall squat long were determined. In addition, to normalize intensity between the wall squat and the one-leg squat exercises, each subject's 12 RM was determined. To determine the weight used for the wall squat short and long, each subject used their 12 RM weight while performing the wall squat using a heel-to-wall distance that was halfway between the heel-to-wall distances for the wall squat short and wall squat long, and this weight was used for both the wall squat short and the wall squat long during the testing session. The mean total dumbbell mass used was 56 ± 9 kg for males and 36 ± 8 kg for females for the wall squat short and wall squat long and 15 ± 3 kg for males and 10 ± 3 kg for females for the one-leg squat.
Blue Sensor (Ambu Inc., Linthicum, MD) disposable surface electrodes (type M-00-S) were used to collect EMG data. These oval-shaped electrodes (22 mm wide and 30 mm long) were placed in a bipolar electrode configuration along the longitudinal axis of each muscle, with a center-to-center distance of approximately 3 cm. Before positioning the electrodes over each muscle, the skin was prepared by shaving, abrading, and cleaning with isopropyl alcohol wipes to reduce skin impedance. As previously described (2), electrode pairs were then placed on the subject's right side for the following muscles: a) rectus femoris, b) vastus lateralis, c) vastus medialis, d) medial hamstrings (semimembranosus and semitendinous), e) lateral hamstrings (biceps femoris), and f) gastrocnemius.
Spheres (3.8 cm in diameter) were attached to adhesives and positioned over the following bony landmarks: a) third metatarsal head of the right foot, b) medial and lateral malleoli of the right leg, c) upper edges of the medial and the lateral tibial plateaus of the right knee, d) posterosuperior greater trochanters of the left and the right femurs, and e) lateral acromion of the right shoulder.
Once the electrodes and the spheres were positioned, the subject warmed up and practiced the exercises as needed, and data collection was commenced. A six-camera peak performance motion analysis system (Vicon-Peak Performance Technologies, Inc., Englewood, CO) was used to collect 60-Hz video data. Force data were collected at 960 Hz using a force platform (Model OR6-6-2000; Advanced Mechanical Technologies, Inc.). EMG data were collected at 960 Hz using a Noraxon Myosystem unit (Noraxon USA, Inc., Scottsdale, AZ). The EMG amplifier bandwidth frequency was 10-500 Hz. Video, EMG, and force data were electronically synchronized and simultaneously collected as each subject performed in a randomized manner one set of three continuous repetitions (trials) during the wall squat short, wall squat long, and one-leg squat.
After completing all exercise trials, EMG data were collected during maximum voluntary isometric contractions (MVIC) to normalize the EMG data collected during each exercise (11). The MVIC for the rectus femoris, vastus lateralis, and vastus medialis were collected in a seated position at 90° knee and hip flexion with a maximum effort knee extension. The MVIC for the lateral and the medial hamstrings were collected in a seated position at 90° knee and hip flexion with a maximum effort knee flexion. MVIC for the gastrocnemius was collected during a maximum effort standing one-leg toe raise with the ankle positioned approximately halfway between neutral and full plantarflexion. Two 5-s trials were randomly collected for each MVIC.
Video images for each marker were tracked and digitized in three-dimensional space with peak performance software. Ankle, knee, and hip joint centers were mathematically determined using the external markers and appropriate equations as previously described (11). Testing of the accuracy of the calibration system resulted in markers that could be located in three-dimensional space with an error less than 4-7 mm. The raw position data were smoothed with a double-pass fourth-order Butterworth low-pass filter with a cutoff frequency of 6 Hz (11). Joint angles, linear and angular velocities, and linear and angular accelerations were calculated using appropriate kinematic equations (11).
Raw EMG signals were full-waved rectified and smoothed with a 10-ms moving average window throughout the knee range of motion for each repetition. These EMG data were then normalized for each muscle and were expressed as a percentage of each subject's highest corresponding MVIC trial. The MVIC trials were calculated using the highest EMG signal over a 1-s time interval throughout the 5-s MVIC. Normalized EMG data were then averaged over the three repetition trials performed for each exercise as a function of knee angle and were used in the biomechanical model described below.
As previously described (11,41), a biomechanical model of the knee (Figs. 4 and 5) was used to continuously estimate cruciate ligament forces throughout a 90° knee range of motion during the knee flexing (squat descent) phase (0°-90°) and the knee extending (squat ascent) phase (90°-0°) of the lunge. Resultant force and torque equilibrium equations were calculated using the inverse dynamics and the biomechanical knee model (11,41). Anteroposterior shear forces in the knee were calculated and adjusted to ligament orientations to estimate ACL or PCL forces (16). Moment arms of muscle forces and angles for the line of action for the muscles and the cruciate ligaments were expressed as polynomial functions of knee angle using data from Herzog and Read (16). Knee torques from cruciate and collateral ligament forces and bony contact were assumed to be negligible as were forces and torques out of the sagittal plane.
Quadriceps, hamstrings, and gastrocnemius muscle forces were estimated as previously described (11,41). Because the accuracy of estimating muscle forces depends on accurate estimations of a muscle's physiological cross-sectional area (PCSA), maximum voluntary contraction force per unit PCSA, and the EMG-force relationship, resultant force and torque equilibrium equations may not be satisfied. Therefore, each muscle force Fm (i) was modified by the following equation at each knee angle:
Equation (Uncited)Image Tools
where Ai is the PCSA of the ith muscle, σm(i) is the MVIC force per unit PCSA of the ith muscle, EMGi and MVICi are the EMG window averages of the ith muscle EMG during exercise and MVIC trials, ci is a weight factor (values given below) adjusted in a computer optimization program to minimize the difference between the resultant torque from the inverse dynamics (Tres) and the resultant torque calculation from the biomechanical model (Tmi) (Fig. 4), kli represents each muscle's force-length relationship as function of hip and knee angles (based on muscle length, fiber length, sarcomere length, pennation angle, and cross-sectional area) (33), and kvi represents each muscle's force-velocity relationship based on a Hill-type model for eccentric and concentric muscle actions using the following equations from Zajac (38) and Epstein and Herzog (10):
Equation (Uncited)Image Tools
with F0 representing the isometric muscle force, l0 is the muscle fiber length at rest, v is the velocity, and a = 0.32F0, b = 3.2l0·s−1, and C = 1.8.
PCSA data from Wickiewicz et al. (33) were used to determine the ratios of PCSA between muscle groups (41). According to Narici et al. (24), the total PCSA of the quadriceps was approximately 160 cm2 for a 75-kg man. Total PCSA of the quadriceps was scaled up or down by individual body mass (41). Forces generated by the knee flexors and extensors at MVIC were assumed to be linearly proportional to their PCSA (41). Muscle force per unit PCSA at MVIC was 35 N·cm−2 for the knee flexors and 40 N·cm−2 for the quadriceps (7,23,24,34).
The objective function used to determine each ith muscle's coefficient ci was as follows:
Equation (Uncited)Image Tools
subject to clow ≤ ci ≤ chigh, where clow and chigh are the lower and the upper limits for ci, and λ is a constant. The weight factor c was to adjust the final muscle force calculation. The bounds on c were set between 0.5 and 1.5. The torques predicted by the EMG driven model matched well (<2%) with the torques generated from the inverse dynamics.
To determine the effects of squat type (wall squat long, wall squat short, and one-leg squat), squat phase (squat descent and squat ascent), and knee angles (0°-90° in 10° intervals) on cruciate ligament forces, a three-factor repeated-measure ANOVA with planned comparisons was used. Bonferroni t-tests were used to evaluate the significance of pairwise comparisons. The level of significance used was P < 0.05.
Mean cruciate ligament force curves are shown in Figure 6. Main effect differences were identified among the three squat types (P < 0.001), between the two squat phases (P < 0.001), and among the 10 knee angles (P < 0.001). When examined at each knee angle, a significant squat type by squat phase interaction was identified at 0° (P = 0.039), 10° (P = 0.002), 20° (P = 0.003), 30° (P = 0.011), 40° (P = 0.010), 50° (P < 0.001), 60° (P = 0.048), 80° (0.003), and 90° (P < 0.001). Pairwise comparisons of mean cruciate ligament forces at specific knee angles (0°-90°) between squat exercises and between squat descent and ascent phases are shown in Table 1. During the squat descent phase, mean PCL forces were significantly greater in the wall squat long (259-757 N range) compared with the wall squat short (100-786 N range) between 0° and 80° knee angles, significantly greater in the wall squat long compared with the one-leg squat (64-414 N) between 0° and 90° knee angles, and significantly greater in the wall squat short compared with the one-leg squat between 0°-20° and 90° knee angles. During the squat ascent phase, mean PCL forces were significantly greater in the wall squat long compared with the wall squat short between 70° and 0° knee angles, significantly greater in the wall squat long compared with the one-leg squat between 90°-60° and 20-0° knee angles, and significantly greater in the wall squat short compared with the one-leg squat between 90°-70° and 0° knee angles. For all three squat exercises, mean peak PCL force magnitudes occurred between 80° and 90° knee angles during the squat ascent and were 723 ± 127 N for the wall squat long, 786 ± 197 N for the wall squat short, and 414 ± 133 N for the one-leg squat. ACL forces, which were generated only during the one-leg squat (31-59 N range), occurred between 0° and 40° knee angles during the squat descent and at 0° knee angle during the squat ascent. The mean peak ACL force magnitude during the one-leg squat was 59 ± 52 N and occurred at 30° knee angle during the squat descent.
Significant differences (P < 0.05) in cruciate ligament force at specified knee angles between the descent and the ascent phases of each squat exercise are shown in Table 1. Mean PCL force was significantly greater in the ascent phase compared with the descent phase between 60° and 80° knee angles for the wall squat long, 70°-90° knee angles for the wall squat short, and 20°-70° knee angles for the one-leg squat. Descriptive data of mean quadriceps and hamstrings force values during wall squat and one-leg squat exercises are shown in Table 2. Quadriceps force ranged approximately between 30 and 720 N and generally increased with knee flexion, whereas hamstring force ranged approximately between 15 and 190 N. At each knee angle, quadriceps and hamstrings forces were generally greater during the ascent compared with the descent.
It is not well understood what PCL or ACL force magnitudes become injurious to the healthy or reconstructed ACL and PCL. In healthy adults, the ultimate strength of the ACL and PCL is approximately 2000 N (36) and 4000 N (27), respectively, although these values depend on age and anatomical factors. Therefore, the ACL and the PCL loads generated during the one-leg squat and the wall squat exercises appear to be well within a safe limit for the healthy ACL and PCL. The reconstructed ACL and PCL have similar ultimate strengths compared with the healthy ACL or PCL, although these values can change considerably depending on graft type and donor characteristics (e.g., autograft vs allograft; patellar tendon vs hamstrings graft) (4,28). However, the healing graft site may be injured with considerably less force compared with the ultimate strength of the graft, although it is not well understood how much force to the graft site is too much and how soon force can be applied after reconstruction. Therefore, the mean peak PCL forces of approximately 400 N during the one-leg squat and approximately 750 N during the wall squat exercises may be problematic early after PCL reconstruction when the graft site is still healing. Moreover, during PCL reconstruction, at the same relative intensity, it may be appropriate to use the one-leg squat before wall squat exercises due to less PCL loading during the one-leg squat, especially compared with the wall squat long. In addition, it may be prudent to use smaller knee angles (e.g., 0°-50°) before progressing to larger knee angles (e.g., 50°-100°) because PCL forces generally increase as knee angle increases. In contrast, wall squat exercises may be a better choice compared with the one-leg squat early after ACL reconstruction due to ACL forces generated during the one-leg squat. However, because peak ACL force during the one-leg squat were only approximately 60 N, it is not likely that the one-leg squat will produce forces that would be injurious to the healing ACL graft, and mild strain to the graft may enhance the healing process (13). Nevertheless, after ACL reconstruction, it may be safer to start with wall squat exercises and progress to the one-leg squat and use larger knee angles (e.g., 50°-100°) before progressing to smaller knee angles (e.g., 0°-50°) because ACL forces may be generated at smaller knee angles less than 50°.
As hypothesized, ACL forces were greater in the one-leg squat compared with the wall squat long and occurred at knee angles between 0° and 40° with a peak magnitude of approximately 60 N at 30° knee angle. During the one-leg sit-to-stand, which is similar to ascent phase of the one-leg squat, Heijne et al. (15) reported a peak 2.8% ACL strain (calibrated to approximately 100 N) at 30° knee angle. Moreover, Kvist and Gillquist (19) reported a peak anterior shear ACL force of less than 90 N at 30° knee angle during the one-leg bodyweight squat, which is similar to the results in the current study. Butler et al. (5) demonstrated that the ACL provides 86% of the total resistance to anterior drawer (caused by an anterior shear force) and the PCL provides approximately 95% of the total restraining force to posterior drawer (caused by a posterior shear force). Therefore, the anterior shear force is resisted primarily by the ACL, and posterior shear force is resisted primarily by the PCL. Moreover, ACL forces as a function of knee angle in the current study are similar to ACL forces and knee angles in the squat literature (3,15,25,32). However, both the ACL and the PCL forces that are generated while performing squatting exercises are dependent on which exercise technique is used and whether external resistance is used. For example, in Beynnon et al. (3), it appears that the subjects may have squatted using a more upright trunk position with relatively little forward trunk tilt, which suggests that these subjects may use their quadriceps to a greater extent than their hamstrings (26). This is important because hamstrings force has been shown to unload the ACL and to load the PCL during the weight-bearing squat exercise (11,21,26). Ohkoshi et al. (26) reported no ACL strain at all knee angles tested (15°, 30°, 60°, and 90°) while maintaining a squat position with the trunk tilted forward, with 30° or more forward trunk tilt being optimal for eliminating or minimizing ACL strain throughout the knee range of motion and recruiting relatively high hamstrings activity.
The exercises that had the greatest amount of anterior knee movement beyond the knees, the one-leg squat (10 ± 2 cm) and wall squat short (9 ± 2 cm), also generated the greatest ACL forces and least PCL forces. These exercises may be preferable to the wall squat long during PCL rehabilitation. In contrast, as hypothesized, the wall squat long, in which the knees did not move beyond the toes, generated the highest PCL forces and no ACL forces and may be problematic during PCL rehabilitation. Anterior knee movement beyond the toes can influence quadriceps activity and patellar tendon force, which in turn can influence cruciate ligament loading. Zernicke et al. (40) estimated the force in the patellar tendon at approximately 17 times bodyweight in a subject that used a considerable external load during a squat descent with excessive anterior knee movement beyond the toes. Although 17 times bodyweight may be an over estimate of the actual force in the patella tendon, large patellar tendon forces tend to load the ACL at smaller knee angles less than approximately 60° (primarily between 0° and 30°) but load the PCL at larger knee angles greater than approximately 60° (9,17,18). Although patellar tendon force from quadriceps activity can load either the ACL or the PCL depending on knee angle, it is difficult to make definite conclusions regarding how quadriceps activity and anterior knee movement may influence cruciate ligament loading while performing squat exercises, and additional research in this area is needed.
Although the wall squat short and one-leg squat both resulted in similar amounts of anterior knee movement at maximum knee flexion, PCL forces were significantly lower in the one-leg squat compared with the wall squat short between 90° and 70° knee angles during the squat ascent (Table 1 and Fig. 6). One explanation of the greater PCL forces between 90° and 70° knee angles in the wall squat short compared with the one-leg squat is greater quadriceps forces that are generated during the wall squat short because quadriceps forces at knee angles greater than 60° load the PCL (9,17,18). Between 90° and 70° knee angles during the ascent, the estimated quadriceps forces in the current study were approximately 30-50% greater in the wall squat short compared with the one-leg squat. Although hamstrings forces between 90° and 70° knee angles also load the PCL, hamstrings forces were only 20-30 N greater in the one-leg squat compared the wall squat short. In contrast, quadriceps force magnitudes were approximately 150 N greater in the wall squat short compared with the one-leg squat, therefore loading the PCL to a great extent compared with the hamstrings.
Although hamstrings forces were greatest in the one-leg squat between 0° and 30° knee angles, the hamstrings are not effective in either unloading the ACL or loading the PCL due to a small insertion angle into the tibia that results in most of the hamstrings force being directed parallel instead of perpendicular to the tibia. Hamstrings force is most effective in generating posterior shear force and in loading the PCL when the knee is flexed approximately 90° (20). The relatively low hamstrings force (typically less than 50 N) generated during the wall squat exercises throughout the knee range of motion implies that wall squat exercises primarily target the quadriceps and not the hamstrings, whereas the one-leg squat is more effective in recruiting the hamstrings. One reason for greater quadriceps force and less hamstrings force in the wall squat short compared with the one-leg squat is because the trunk is erect in the wall squat short (greater knee extensor torque and less hip extensor torque needed to overcome the effects of gravity) but tilted forward 30°-40° in the one-leg squat (less knee extensor torque and greater hip extensor torque needed to overcome the effects of gravity).
The friction and the normal forces that the wall applied to the subject may also help explain why quadriceps forces were greater in the wall squat short compared with the one-leg squat during the squat ascent. Although friction was minimized during the wall squat by using a smooth wall, the normal force that the wall exerted on the subject's back during the wall squat exercises resulted in an increased friction force on the subject as they slid down and up the wall. Because the friction force opposes motion, it acted opposite the force of gravity during the squat descent but in the same direction as the force of gravity during the squat ascent. Therefore, the friction force made it easier for the subject to control the rate of sliding down the wall by producing a knee extensor torque but made it more difficult for the subject to slide up the wall by producing a knee flexor torque. Because the one-leg squat did not have a friction force compare to the wall squat, this provides one plausible explanation why quadriceps force and PCL force were greater in the ascent phase of the wall squat exercises compared with the one-leg squat.
The friction force also differed between the wall squat long and short. Because during the wall squat long the heels were twice as far from the wall compared with the wall squat short, the normal force must be greater in the wall squat long. Because friction force is directly proportional to the normal force, the downward-acting friction force on the subject during the squat ascent was greater in the wall squat long compared with the wall squat short, which makes the wall squat long more difficult to perform. This may partially explain why PCL forces were greater in the wall squat long compared with the wall squat short.
Cruciate ligament forces tended to be higher in the ascent phase compared with the descent phase, in part because quadriceps and hamstrings forces were also greater during the ascent phase. For the wall squat exercises, significant PCL force differences between squat descent and ascent occurred only at higher knee angles between 60° and 90°. As previously mentioned, quadriceps force at knee angles greater than 60° loads the PCL, and the greater quadriceps force was greater during the ascent than the descent in part due to having to overcome gravity and the downward-acting friction force. A different pattern occurred during the one-leg squat, in which between 20° and 70° knee angles PCL forces were significantly greater during the squat ascent compared with the squat descent. These findings are in agreement with the squat literature, in which cruciate forces have been reported to be greater in the squat ascent compared with the squat descent (11,12).
There are limitations in this study. Firstly, muscle and cruciate ligament forces were estimated from biomechanical modeling techniques and not measured directly because it is currently not possible to measure cruciate ligament forces in vivo while performing wall squat and one-leg squat exercises in healthy subjects. However, both Beynnon et al. (3) and Heijne (15), who implanted strain sensors in patients within the anteromedial bundle of an ACL during arthroscopic surgery for partial minisectomies or capsule/patellofemoral joint debridement, after surgery had these patients perform one- and two-leg squat-type exercises. These authors reported a peak ACL strain of approximately 2.8-4% (approximately 100-150 N) at knee angles between 0° and 30°. These ACL force magnitudes and knee angles from Beynnon et al. (3) and Heijne (15) are similar to the current study. Unfortunately, there are no studies that have quantified PCL forces in vivo while performing a squat exercise, so it is not possible to compare the modeled PCL force results in the current study to in vivo PCL forces.
The current study was limited to sagittal plane motion only, and only subjects who could perform all exercises without discernable frontal or transverse plane movements were used in this study. Future three-dimensional biomechanical analyses of the knee during squatting are needed to investigate the effects of transverse plane rotary motions and frontal plane valgus and varus motions on cruciate ligament loading. Slightly different cruciate ligament loading patterns during squatting may occur between two- and three-dimensional analyses, although normal squatting is primarily sagittal plane movements. A normal range of motion of 5°-7° knee valgus and 6°-14° of knee varus has been reported during the one-leg squat (39), although these relatively small amounts of valgus and varus may not affect cruciate ligament loading. However, excessive knee valgus has been shown to be associated with an increased risk of ACL ruptures (22,39). Transverse and frontal plane hip joint motions have also been shown to be associated with an increased risk of ACL ruptures and are relatively common in individuals with weak hip abductors and external rotations (22).
In conclusion, throughout the 0°-90° knee angles, the wall squat long generally exhibited significantly greater PCL forces compared with the wall squat short and the one-leg squat. There was generally no significant difference in PCL force between the wall squat short and the one-leg squat, except at 80° and 90° knee angles, where PCL forces were greater in the wall squat short. Throughout the 0°-90° knee angles, the wall squat exercises generated PCL force magnitudes ranging approximately from 100 to 790 N, with PCL magnitudes generally decreasing between 0° and 30° knee angles and increasing between 40° and 90° knee angles. Moreover, the one-leg squat generated PCL force magnitudes ranging approximately from 60 to 410 N, with PCL magnitudes generally increasing between 50° and 90° knee angles during the descent and 10°-90° knee angles during the ascent. ACL forces were only found in the one-leg squat, which generated relatively small magnitudes of approximately 20-60 N between 0° and 40° knee angles. The one-leg squat, the wall squat long, and the wall squat short all appear to load the ACL and the PCL within a safe range in healthy individuals.
The authors would like to thank Lisa Bonacci, Toni Burnham, Juliann Busch, Kristen D'Anna, Pete Eliopoulos, and Ryan Mowbray for all their assistance during data collection and analyses.
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