ESCAMILLA, RAFAEL F.1; ZHENG, NAIQUAN2; MACLEOD, TORAN D.3; IMAMURA, RODNEY4; EDWARDS, W. BRENT5; HRELJAC, ALAN4; FLEISIG, GLENN S.6; WILK, KEVIN E.7; MOORMAN, CLAUDE T. III8; PAULOS, LONNIE1; ANDREWS, JAMES R.1,6
Closed chain weight-bearing exercises such as the squat, leg press, and forward lunge are commonly used in rehabilitation settings, such as after anterior cruciate ligament (ACL) or posterior cruciate ligament (PCL) reconstruction surgery (10,39). These exercises can be performed with technique variations, which may affect ACL and PCL loading. Although the effects of exercise technique variations on cruciate ligament loading have been examined while performing the squat and the leg press (13,14), there are no studies that have examined the effects of technique variations on cruciate ligament loading while performing the forward lunge. Because patients use the forward lunge after ACL and PCL reconstruction, it is important to understand how the cruciate ligaments are loaded, especially during the early phases of rehabilitation when the goal is to minimize ACL or PCL loading.
There are multiple techniques that individuals can use during the forward lunge, such as lunging using a short step length or a long step length. Lunging forward using a long step typically results in the lead knee being maintained over the lead foot throughout the knee range of motion, whereas lunging forward using a short step length lunge typically results in the lead knee translating beyond the toes throughout much of the knee range of motion. Some clinicians believe that anterior movement of the lead knee beyond the toes during a short step forward lunge increases cruciate ligament loading, although there are very limited data that support this belief (2). Moreover, it is unclear if the ACL or the PCL is loaded when anterior knee movement occurs.
The forward lunge can be performed and progressed using varying techniques. One technique involves starting in an upright position, stepping forward with the lead leg and flexing the lead knee until the rear knee touches the ground, and then pushing back to the starting upright position. Because a stride is taken by the lead leg during each repetition, this technique may be called a forward lunge with a stride. Another technique involves first stepping forward with the lead leg and starting with both knees fully extended. From this position, the individual flexes the lead knee until the rear knee touches the ground, and then both knees are extended back to the starting position. In this technique, which has been previously described (16), both feet remain stationary as the individual lunges up and down. Because this technique does not involve striding forward during each repetition, this technique may be called a forward lunge without a stride. The lunge with a stride can be a progression of the lunge without a stride, with the lunge without a stride being a beginning exercise and the lunge with a stride being more difficult to perform and advanced. The lunge with a stride requires higher levels of lower body strength and coordination compared with the lunge without a stride.
Understanding how cruciate ligaments are loaded differently among these technique variations of the forward lunge may allow clinicians to prescribe safer and more effective knee rehabilitation treatment to patients during ACL or PCL rehabilitation. For example, during the forward lunge, if ACL loading occurs when using a short step but not when using a long step, the forward lunge with a long step may be more appropriate for the patient if the clinician's immediate goal for the patient is to minimize ACL loading. Similarly, during the forward lunge, if PCL loading occurs with a stride but not without a stride, the forward lunge without a stride may be more appropriate for the patient if the clinician's immediate goal for the patient is to minimize PCL loading.
Our purpose was to compare cruciate ligament tensile forces while performing the forward lunge with a long step (forward lunge long), with a short step (forward lunge short), with a stride, and without a stride. It was hypothesized that ACL tensile forces would be greater in the forward lunge short compared with the forward lunge long, PCL tensile forces would be greater in the forward lunge long compared with the forward lunge short, and PCL tensile forces would be greater during the forward lunge with a stride compared with without a stride. Muscle force magnitudes in each subject's quadriceps and hamstrings will also be estimated to help better understand ACL and PCL force magnitudes.
Eighteen healthy individuals (nine men and nine women) 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 men and 25 ± 2 yr, 60 ± 4 kg, and 164 ± 6 cm, respectively, for women. All subjects were required to perform the forward lunge exercises pain free and with proper form and technique for 12 consecutive repetitions using their 12-repetition maximum (12RM) weight.
To control the EMG signal quality, this study was limited to men and women who had average or below average body fat, which was assessed by the Baseline skinfold calipers (Model 68900; Country Technology, Inc., Gays Mills, WI), and appropriate regression equations and body fat standards set by the American College of Sports Medicine (3). Average body fat was 12% ± 4% for men and 18% ± 1% for women. The protocol used in the current study was approved by the institutional review board at the California State University, Sacramento, CA, and all subjects provided written informed consent.
Each subject performed the forward lunge long (Fig. 1) and the forward lunge short (Fig. 2) with and without a stride. The starting and the ending positions of the forward lunge long with stride and forward lunge short with stride were the same, which involved standing upright with both feet together and the knees fully extended (full knee extension = 0° knee angle). From this position, the subject held a dumbbell weight in each hand and lunged forward with the right leg toward a force platform at ground level. At right foot contact, the right knee flexed at approximately 45°·s−1 until approximately 90°-100° knee angle, at which time the left knee made contact with the ground. From this position, the subject immediately pushed backward off the force platform and returned to the upright standing position with feet together.
During the forward lunge long, each subject used a long step length that resulted in the right leg (tibia) being approximately vertical at the lowest position of the lunge (Fig. 1), thus maintaining the knee over the foot. The average step length (measured from left toe to right heel) for the forward lunge long was 89 ± 4 cm for men and 79 ± 6 cm for women. The step length for the forward lunge short was one half the distance of the step length of the forward lunge long. The shorter step length for the forward lunge short caused the anterior surface of the knee to translate beyond the distal end of the toes, as shown in Figure 2.
The forward lunge long and short without stride was performed the same as the forward lunge long and short with stride, with the exception that during the forward lunge long and short without stride both feet remained stationary throughout each repetition. That is, from the lowest position of the forward lunge long and short shown in Figures 1 and 2, the subject fully extended both knees and then flexed both knees returning back to the lowest position of the lunge. For all lunge variations, a metronome was used to help ensure the right knee flexed and extended at a normal rate of approximately 45°·s−1. During the forward lunge long and short with and without a stride, maximum forward trunk tilt (which occurred near maximum lead knee flexion) was approximately 10°-20° for all subjects.
Each subject came in for a familiarization session 1 wk before the testing session. The experimental protocol was reviewed, the subject was given the opportunity to practice the lunge variations, and each subject's step length for the forward lunge long was determined. In addition, each subject's 12RM was determined while performing the forward lunge with stride using a step length halfway between the forward lunge long and the forward lunge short. Subjects used their 12RM weight for the four lunge variations during data collection. The mean total dumbbell mass used was 49 ± 11 kg for men and 32 ± 8 kg for women.
To collect EMG data, Blue Sensor (Ambu Inc., Linthicum, MD) disposable surface electrodes (type M-00-S) 22 mm wide and 30 mm long were positioned in a bipolar configuration along the longitudinal axis of each muscle, with a center-to-center distance of approximately 3 cm between electrodes. Before applying the electrodes, the skin was prepared by shaving, abrading, and cleaning with isopropyl alcohol wipes to reduce skin impedance. Each subject had electrode pairs positioned on the right side using previously described locations (4) for the following muscles: a) rectus femoris, b) vastus lateralis, c) vastus medialis, d) medial hamstrings (semimembranosus and semitendinosus), e) lateral hamstrings (biceps femoris), and f) gastrocnemius (middle portion between medial and lateral bellies).
Spherical markers (3.8 cm in diameter) covered with 3M™ reflective tape 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 lateral tibial plateaus of the right knee, d) posterosuperior greater trochanters of the left and right femurs, and e) lateral acromion of the right shoulder.
After the subject warmed up and practiced the exercises as needed, data collection commenced. A six-camera Peak Performance motion analysis system (Vicon-Peak Performance Technologies, Inc., Englewood, CO) collected 60 Hz of video data. A force platform (Model OR6-6-2000; Advanced Mechanical Technologies, Inc., Watertown, MA) collected 960 Hz of force data, while a Noraxon EMG system (Noraxon USA, Inc., Scottsdale, AZ) collected 960 Hz of EMG data. The EMG amplifier bandwidth frequency was 10-500 Hz with an input impedance of 20,000 kΩ, and the common-mode rejection ratio was 130 dB. Video, EMG, and force data were electronically synchronized and collected simultaneously as each subject performed one set of three repetitions using their 12RM weight of the forward lunge long with stride, forward lunge long without stride, forward lunge short with stride, and forward lunge short without stride, assigned in a random order. Each subject rested approximately 2-3 min between lunge variations. Tape markers were used to help each subject identify the proper stride length distance between their rear and lead foot for each lunge variation.
After completing all lunge variations, each subject performed maximum voluntary isometric contractions (MVIC) to normalize the EMG data collected during each lunge variation. The MVIC for the rectus femoris, the vastus lateralis, and the vastus medialis were collected in a seated position at 90° knee and hip flexion with a maximum effort knee extension (13). The MVIC for the lateral and medial hamstrings were collected in a seated position at 90° knee and hip flexion with a maximum effort knee flexion (13), with the ankle maintained in a neutral position. 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 (13). Two 5-s trials were randomly collected for each MVIC, with 1-2 min of rest given between trials.
Video images for each reflective marker were tracked and digitized in three-dimensional space with Peak Performance software (version 5.0), using the direct linear transformation calibration method (34). Ankle, knee, and hip joint centers from the link segment model were mathematically determined using the external markers and appropriate equations as previously described (7,13). Testing of the accuracy of the calibration system resulted in reflective markers that could be located in three-dimensional space within our laboratory with an error less than 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 (13). Joint angles, linear and angular velocities, and linear and angular accelerations were calculated in a two-dimensional sagittal plane of the knee using appropriate kinematic equations (13).
Raw EMG signals were full-waved rectified, smoothed with a 10-ms moving average window, linear enveloped (5) throughout the knee range of motion for each repetition, and normalization by expressing the data as a percentage of each subject's highest corresponding MVIC trial. The highest EMG signal over a 1-s time interval throughout the 5-s MVIC was determined to calculate MVIC trials. Normalized EMG data for the three repetitions (trials) were then averaged at corresponding knee flexion angles between 0° and 90° and were used in the biomechanical model described below.
As previously described (13,41), a biomechanical model of the knee (Figs. 3 and 4) 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 inverse dynamics and the biomechanical knee model (13,41). Anteroposterior shear forces in the knee were calculated and adjusted to ligament orientations to estimate ACL or PCL forces, while moment arms of muscle forces and angles for the line of action for the muscles and cruciate ligaments were expressed as polynomial functions of knee angle (23). 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 an EMG-driven biomechanical knee model, as previously described (13,41). 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 EMG-force relationship, resultant force and torque equilibrium equations may not be satisfied. Therefore, the modified muscle force Fm(i) equation at each knee angle is as follows:
where Ai is the PCSA of the ith muscle, σm(i) is the MVIC force per unit PCSA of the ith muscle, EMG 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. 3), kli represents each muscle's force-length relationship as function of hip and knee flexion angles (on the basis of muscle length, fiber length, sarcomere length, pennation angle, and cross-sectional area) (36), and kvi represents each muscle's force-velocity relationship on the basis of a Hill-type model for eccentric and concentric muscle actions using the following equations from Zajac (40) and Epstein and Herzog (12):
Equation (Uncited)Image Tools
with F0 representing isometric muscle force; l0, muscle fiber length at rest; v, velocity; a = 0.32 F0; b = 3.2l0 per second; and C = 1.8.
Ratios of PCSA between muscle groups (41) were determined from the PCSA data from Wickiewicz et al. (36). According to Narici et al. (29), the total PCSA of the quadriceps was approximately 160 cm2 for a 75-kg man, and the 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 was 35 N·cm−2 for the hamstrings and gastrocnemius and 40 N·cm−2 for the quadriceps (11,28,29,37).
The objective function used to determine each ith muscle's coefficient ci was as follows:
subject to clow ≤ ci ≤ chigh, where clow and chigh were lower and upper limits for ci, and λ was 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, which resulted in the torques predicted by the EMG-driven model matching well (less than 2% of the knee torques) with the torques generated from the inverse dynamics.
Equation (Uncited)Image Tools
To determine significant differences in cruciate ligament forces between the two step length variations (forward lunge long and forward lunge short) and two stride variations (with stride and without stride), cruciate ligament forces were statistically analyzed every 10° during the 0°-90° knee flexing (descent) phase and the 90°-0° knee extending (ascent) phase using a two-factor (step length variations and stride variations) repeated-measures ANOVA. To minimize the probability of type I errors secondary to the use of a separate ANOVA for each knee angle, the Bonferroni adjustment had a level of significance set at P < 0.0025 (0.05/20 knee angles).
Table 1 and Figures 5-8 display cruciate ligament force magnitudes and patterns. Comparing the forward lunge long with the forward lunge short across stride variations (Table 1), mean PCL forces ranged between 69 and 765 N and were significantly greater (P < 0.001) in the forward lunge long at 0°, 10°, 20°, 30°, 40°, 50°, 60°, 70°, and 80° knee flexion angles of the descent phase and at 70°, 60°, 50°, 40°, 30°, 20°, 10°, and 0° knee flexion angles of the ascent phase. Comparing with and without stride differences across step length variations (Table 1), mean PCL forces ranged between 86 and 691 N and were significantly greater (P < 0.001) without a stride at 0°, 10°, and 20° knee flexion angles during the descent phase. There were no significant interactions between levels of step length (forward lunge long and forward lunge short) and stride (with stride and without stride).
Visual observation of the data (Figs. 5-8) indicates that PCL force generally increased progressively as knee angle increased and decreased progressively as knee angle decreased. Moreover, for a given knee angle, cruciate ligament forces were greater during the ascent phase compared with the descent phase (Table 1).
ACL forces were observed only during the forward lunge short with stride, occurring between 0° and 10° knee flexion angles during the descent phase and ranging from 0 to 50 N (Figs. 6 and 8). Compared with the forward lunge short with stride, between 0° and 10° knee flexion angles during the descent phase, mean PCL forces occurred during both the forward lunge long with stride (approximately 250-300 N from Fig. 6) and the forward lunge short without stride (approximately 250 N from Fig. 8).
Table 2 displays descriptive data of mean quadriceps and hamstrings force values during the forward lunge exercises. Quadriceps force ranged between approximately 65 and 680 N and generally increased with knee flexion, whereas hamstring force ranged between approximately 20 and 145 N and remained relatively constant throughout the descent phase and throughout the ascent phase. At each knee angle, quadriceps and hamstrings forces were generally greater during the ascent phase compared with the descent phase.
Our results demonstrate that performing the forward lunge with varying techniques does affect cruciate ligament loading. For healthy individuals or during the early phases of ACL rehabilitation when the goal is to minimize ACL loading (such as after ACL reconstruction), all four forward lunge variations may be appropriate because relatively low ACL forces were generated (<50 N). However, it is unknown how much loading can safely occur in the reconstructed ACL (and graft type must also be considered), although some ACL loading during rehabilitation is probably desirable (18). Although the ultimate strength of the healthy ACL is in excess of 2000 N (38) and the ultimate strength of the reconstructed ACL has been estimated between approximately 2500 and 4000 N (8,33), it is unclear how much ACL loading may become injurious to the graft healing site during ACL rehabilitation.
As hypothesized, the forward lunge short, which resulted in the lead knee translating forward beyond the toes 8 ± 3 cm at maximum knee flexion, generated greater ACL forces and smaller PCL forces compared with the forward lunge long, which maintained the lead knee over the foot throughout the knee range of motion. These results support the beliefs of some clinicians that cruciate ligament loading is different between the forward lunge short and the forward lunge long.
When the goal is to minimize ACL loading, the forward lunge long may be a more appropriate and safer choice compared with the forward lunge short, especially the forward lunge short with stride, which was the only lunge variation that generated ACL loading. In addition, lunging without a stride may be a safer choice compared with lunging with a stride because on the basis of the results of this study, the ACL is less likely to be loaded without a stride compared to with a stride. Moreover, performing the lunge with more knee flexion may be preferred compared with less knee flexion because ACL forces occurred only when the knee was flexed between 0° and 10°.
In contrast to ACL rehabilitation, during the early phases of PCL rehabilitation, when the goal is to minimize PCL loading (such as after PCL reconstruction), all lunge variations should be used cautiously, especially at higher knee flexion angles between 60° and 90° where mean PCL forces ranged between approximately 475 and 775 N for the forward lunge long, between 250 and 600 N for the forward lunge short, between 375 and 675 N for lunging with a stride, and between 350 and 700 N for lunging without a stride. Like the reconstructed ACL, it is unknown how much loading can safely occur in the reconstructed PCL, and graft type is important to consider. Because the ultimate strength of the healthy PCL is approximately 4000 N (32), all four lunge variations appear appropriate for healthy individuals. However, although the reconstructed PCL typically has equal or greater ultimate strength compared with the healthy PCL, it is unclear how much PCL loading may become injurious to the healing graft site during PCL rehabilitation.
As hypothesized, PCL forces were greater while performing the forward lunge long compared with the forward lunge short. Therefore, when the goal is to minimize PCL loading, the forward lunge short may be a more appropriate and safer choice compared with the forward lunge long. In addition, lunging with a stride may be a safer choice compared with lunging without a stride, but only when the knee was flexed at lower angles between 0° and 20°. Moreover, performing the lunge throughout a lower knee flexion range may be preferred compared with a higher knee flexion range because PCL forces were generally greater at higher knee flexion angles.
An unexpected finding was that at lower knee flexion angles, lunging without a stride loaded the PCL to a greater extent compared with lunging with a stride. In contrast, at lower knee flexion angles, lunging with a stride loaded the ACL to a greater extent compared with lunging without a stride but only during the forward lunge short (Fig. 8). One possible explanation on why cruciate ligament forces were different only at lower knee flexion angles (0°-20°) between lunging with and without a stride is because compared with lunging without a stride, lunging with a stride produced 15%-30% greater quadriceps forces when the knee was flexed between 0° and 20° during the descent. Higher quadriceps force at these lower knee flexion angles has been shown to result in greater ACL loading (13,15). At these lower knee flexion angles, force from the patellar tendon via the quadriceps attempts to pull the tibia anterior, which is restrained primarily by the ACL (9). Because the lines of pull of the cruciate ligaments change at different knee flexion angles (23), which affects cruciate ligament loading, this should be investigated more thoroughly in future studies.
One reason quadriceps forces were greater with a stride compared to without a stride is that the peak resultant ground reaction forces acting on the lead foot were approximately 15%-20% greater with a stride between 0° and 20° knee flexion angles of the descent. The resultant ground reaction force vector acting on the lead foot produced a flexor torque on the lead knee throughout the lunge, opposed by the knee extensors. Just after lead foot contact during the descent, when the knee was flexed 0°-20°, peak resultant ground reaction forces acting on the lead foot were greater with a stride because the center of mass of the body had more forward and downward acceleration compared with without a stride. Therefore, with a stride, the lead foot had to push harder into the ground to slow down the forward and the downward accelerating center of mass of the body and to control the rate of lead knee flexion, which was 45°·s−1 for both with and without stride conditions.
Like the current study, several studies have reported primarily PCL loading and not ACL loading while performing weight-bearing closed chain exercises. Escamilla et al. (13,14) reported PCL loading only throughout the knee range of motion during the barbell squat and leg press using a 12RM load. Stuart et al. (35) reported tibial posterior shear loads only (PCL loading) throughout the knee range of motion while performing a forward lunge exercise using a 50-N barbell, which support the results of the current study. Moreover, the subjects in the aforementioned studies all used external resistance while performing the squat, the leg press, and the lunge.
In contrast to PCL-only loading during closed chain exercises, Beynnon et al. (6) and Heijne et al. (22) reported a peak ACL strain of approximately 4% (estimated to be approximately 150 N on the basis of the finding that a 150-N Lachman test produced 3.7% strain at 30° knee flexion angle) at knee flexion angles between 0° and 60° during squatting with and without a low-resistance sport cord and no ACL strain at knee flexion angles greater than 60°. During the lunge with a stride (no external resistance and step length not reported), Heijne et al. (22) reported a mean ACL strain of approximately 1% or less (estimated to be approximately 40 N or less) at knee flexion angles less than 60° (no ACL strain at knee flexion angles greater than 60°) and a peak ACL strain of 1.8% (estimated to be approximately 75 N) between a 0° and 30° knee flexion angle range. By comparison, the peak ACL force in the current study was approximately 50 N in the forward lunge short with stride between a 0° and 10° knee flexion angles. This demonstrates a remarkable similarity between the ACL lunge data in the current study, calculated by knee modeling techniques, and the ACL strain lunge data by Heijne et al. (22), calculated by direct measurement using force sensors within the ACL. The subjects in Heijne et al. (22) were patients that had force sensors implanted within the anteromedial bundle of a healthy ACL during arthroscopic surgery to repair damaged knee structures (partial meniscectomies; capsule and patellofemoral joint debridement). Immediately after surgery, these patients performed a variety of exercises, including the lunge, with strain measured within the anteromedial bundle of the ACL and referenced to an instrumented Lachman test. Unfortunately, Heijne et al. (22) did not measure PCL strain, so we cannot compare PCL loads between studies.
What is consistent in closed chain exercise studies is that PCL loading occurred at knee flexion angles greater than 60°. Although the current study only investigated cruciate ligament loading between 0° and 90° knee flexion angles, it is likely, on the basis of the results of both the current study and the previous studies (13,15), that PCL loading would continue to increase at knee flexion angles greater than 90°. What is inconsistent in closed chain exercise studies is that ACL loading occurred at knee flexion angles between 0° and 60° in some studies, and no ACL loading occurred throughout the knee range of motion in other studies. The conflicting findings in ACL loading among weight-bearing exercises may be due to differences in the exercise technique used, differences in the external resistance used, or methodological differences. For example, in Beynnon et al. (6), the subjects appeared to have squatted using an upright trunk with relatively little forward trunk tilt. This suggests that these subjects may have used their quadriceps to a greater extent than their hamstrings because it has been demonstrated during the squat that the hamstrings are recruited more and the quadriceps are recruited less as the trunk tilts forward (30). However, the marker set used in the current study is unable to discriminate between pelvis and trunk positions, and future studies are needed to investigate the effects of pelvis and trunk positions on cruciate ligament loading. This is important because trunk position has been shown to affect hamstrings activity, and hamstrings force has been shown to unload the ACL and load the PCL during the weight-bearing squat exercise (13,26,30). For example, Ohkoshi et al. (30) reported no ACL strain at all knee flexion angles tested (15°, 30°, 60°, and 90°) while maintaining a squat position with the trunk tilted forward from 0° to 90°, with 30° or more forward trunk tilt being optimal for eliminating or minimizing ACL strain throughout the knee range of motion. In the current study, although trunk positions were similar among all four lunge variations, both hamstring force and PCL force were greater in the forward lunge long compared with the forward lunge short (Tables 1 and 2). In fact, the estimated hamstring forces calculated in the current study were 50%-60% greater in the forward lunge long compared with the forward lunge short, which helps explain the greater PCL forces generated in the forward lunge long.
Compared with quadriceps force (peak force near 700 N), hamstrings force (peak force near 150 N) was relatively low during all four lunge variations. This relatively low hamstrings force compared with quadriceps force may occur by a relatively erect trunk position during all lunge variations. Farrokhi et al. (17) demonstrated that compared with performing a forward lunge with a relatively erect trunk (similar to the forward trunk tilt in the current study), performing the lunge with the trunk tilted forward approximately 30°-45° resulted in significantly greater hip extensor impulse and significantly greater hamstrings and gluteus maximum activity. The greater hamstrings activity during the lunge with the excessive forward trunk tilt would likely result in an increase in hamstrings force compared with performing the lunge with a more erect trunk, which may result in greater PCL loading and less ACL loading.
In addition to forward trunk tilt, increasing external resistance during weight-bearing exercise increases hamstrings involvement. For example, Escamilla et al. (13) reported no ACL loading throughout the knee range of motion in power lifters who squatted with a 12RM external resistance and a forward trunk tilt of approximately 30°. Moreover, these subjects had relatively high hamstrings activity, ranging between 40% and 80% of an MVIC for the lateral hamstrings and between 30% and 60% of an MVIC for the medial hamstrings. Other studies involving the forward lunge exercise have also reported relatively high hamstring activity (1,21,31) and hip extensor torque (19), which implies hamstring involvement.
This initial lunge study examined healthy individuals without cruciate ligament pathology because cruciate ligament forces during forward lunge variations are currently unknown using the healthy population. Additional research is needed using patients with cruciate ligament pathology or reconstruction to determine if cruciate ligament forces generated during forward lunge variations are similar between healthy individuals and patients with cruciate ligament deficiencies or reconstruction. Additional studies are also needed using other techniques variations, such as using a lunge step length somewhere between the forward lunge long and the forward lunge short, to determine what optimal step length minimizes cruciate ligament loading.
There are some limitations to the current study. First, there is no practical way to validate our knee model against the gold standard of measuring ACL strain (force) directly. This is because currently in the United States, committees that regulate and ensure the protection of human subjects in research endeavors do not approve invasive techniques of inserting force sensors within the ACL in normal healthy subjects for the purposes of exercise research. However, as previously discussed, force sensors have been inserted within the anteromedial bundle of the ACL in patients who underwent arthroscopic surgery to repair damaged knee structures, and some of these patients were asked to perform the lunge exercise immediately after surgery (22). As previously mentioned, the ACL strain (force) that was reported during the lunge by Heijne et al. (22) (measured directly by force sensors within the ACL) is similar to the ACL forces in the current study and occurred at similar knee flexion angles, which provides some validation for our modeled data. Another limitation is that the current study was limited to sagittal plane motion only and only included subjects that could perform the exercises without transverse plane rotary motions and frontal plane valgus/varus motions. Future studies should investigate the effects of transverse plane rotary motions and frontal plane valgus/varus motions on cruciate ligament loading as well as investigate the effects of performing lunging exercises in individuals with cruciate ligament deficiencies. Individuals that perform the lunge with excessive transverse or frontal plane rotary motions may result in increased loading of the ACL, and this needs to be investigated. During drop landing, Kernozek and Ragan (25) reported that rotational moments were small in drop landing and contributed little to ACL tension. These authors reported that the factors that contributed most to ACL loading were patellar tendon force and the tibial slope as well as joint axial loads. Sex differences should also be examined in future studies because knee biomechanical differences between sex have been shown to occur during jumping and landing (20) and likely also would occur during lunging exercises, especially in women that have poor hip and weak hip external rotators and abductors (24,27).
Lunge technique variations do affect cruciate ligament loading. All lunge variations appear appropriate and safe during ACL rehabilitation because of minimal ACL loading, especially the forward lunge long and lunging without a stride. However, clinicians should be cautious in prescribing the forward lunge exercises during the early phases of PCL rehabilitation when the graft site is still healing because of relatively high PCL forces, especially at higher knee flexion angles during the forward lunge long. PCL forces were greater in the forward lunge long compared with the forward lunge short throughout most of the descent and ascent phases. Relatively low ACL forces occurred during the forward lunge short at small knee flexion angles, but no ACL loading occurred during the forward lunge long. The only difference in PCL force between with stride and without a stride was at 0°-20° knee flexion angles during the descent phase, in which PCL forces were significantly greater without a stride. PCL forces generally progressively increased as knee angle increased and decreased as knee angle decreased and were greater during the ascent phase compared with the descent phase.
The efforts of Dr. Bonnie Raingruber and funding from the National Institute of Child Health and Human Development's Extramural Associates Research Development Award program made this research possible. Also acknowledged are Lisa Bonacci, Toni Burnham, Juliann Busch, Kristen D'Anna, Pete Eliopoulos, and Ryan Mowbray for their assistance in data collection and analyses.
The results of the current study do not constitute endorsement by the American College of Sports Medicine.
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