The goal of any exercise training program should be to improve the overall efficiency of human movement patterns. Whether the training program is designed to improve athletic performance, to rehabilitate an injury, or simply to improve overall physical fitness; the goal of performing repeated movements (exercises) should be to make movement patterns more efficient. The concept of efficiency as it relates to human movement has been defined several different ways. In terms of human movement energetics, mechanical efficiency is defined as the amount of mechanical work done divided by the metabolic cost needed to perform that work (39). Although this is the classic definition of mechanical efficiency, others have proposed more biomechanical definitions of human movement efficiency. These include patterns of movement that fulfill their tasks with minimal strain on the musculoskeletal system (27) or as movement that occurs without pain or discomfort and involves proper joint alignment, muscle coordination, and posture (18). These biomechanical definitions hypothesize that if we can develop more efficient movement patterns, this reduction in strain on the movement components could help in the rehabilitation process and in preventing the development of musculoskeletal pain altogether (31). Because musculoskeletal conditions accounted for $849 billion in health care costs in the U.S.A. in 2004 (7.7% of the gross domestic product), and with these costs expected to escalate as the population ages (16), developing noninvasive strategies to prevent these conditions is becoming increasingly important. Therefore, developing exercises aimed at correcting movement inefficiencies could help in delaying and preventing the onset of some musculoskeletal pathology.
The most common pathologic conditions in the lower extremity affect the knee joint (36,37), and women experience knee pathologic conditions much more commonly than men (2,13). It can be hypothesized that there may be differences in muscular activation strategies between men and women that might be one of several factors leading to these increased rates of knee injury. Studies that have attempted to determine the causes of knee pathologies have suggested that hip muscle weakness is associated with the development of knee conditions (5,6,20,26). Therefore, designing exercises aimed at preferentially strengthening the hip musculature is warranted, as is examining movement pattern differences between men and women.
One fundamental human movement pattern that is also commonly used as an exercise in many training programs is the squat. Variations on this basic human movement pattern have been used throughout time during various activities of daily living and have also been widely used as a core component of athletic, clinical, and fitness-related training programs (1,23). The squat is a complex multijoint movement requiring contributions from several muscles and joints to be performed optimally. It has been suggested that to perform the basic squat movement pattern efficiently, one requires mobility of the ankle, hip, and thoracic spine, while requiring stability of the foot, knee, and lumbar spine (18). This theory suggests that of the 2 largest lower limb joints producing most of the work for the squat movement (10), the hip joint should be most responsible for producing the movement.
During the squat movement, and during the loading response of many human movements, the ground reaction force falls anterior to the hip joint and posterior to the knee joint in the sagittal plane (8,35). This creates external flexion moments on both the hip and knee that must be balanced by the internal extension moments provided by the hip extensors (gluteus maximus [GM] and hamstring muscles) and knee extensors (quadriceps), respectively. Without considering the capability of the hamstring muscles to extend the hip, the GM alone has a larger cross sectional area (4,842 mm2) than the total combined area of all 4 quadriceps muscles together (4,317.5 mm2) (15). Because a muscle's cross sectional area is highly correlated to the amount of force it can generate (11), it can be hypothesized that the larger hip extensor musculature could be better able to handle external loading as compared with the knee extensor musculature. There is also evidence that improving the use of the hip extensors in the sagittal plane can help unload the knees by decreasing the amount of quadriceps force required (28). Therefore, it can be hypothesized that emphasizing the loading of the hip joint over the knee joint during a squat exercise could help create a more efficient pattern of movement.
It has been suggested that keeping the ground reaction force closer to the knee and further from the hip in the sagittal plane can help in emphasizing the loading of the hip extensors over the knee extensors during human movement (29). One potential method of increasing the contribution of the hip and decreasing the contribution of the knee during the squat movement would be to place additional mass in the hands and move this mass anteriorly so that it can counterbalance the hips moving posteriorly. Therefore, this study will investigate a counterbalanced squat (CBS) technique (Figure 1B) to determine whether this technique is better able to increase the sagittal plane loading of the hip while decreasing the same loading of the knee as compared with a regular squat (RS) technique (Figure 1A). It will also examine the differences in basic squat biomechanics between male and female subjects.
The subjects were 31 (15 men and 16 women) recreationally trained (involved in at least 2 h·wk−1 of sporting activities for a minimum of 2 years but had never regularly trained using loaded squats) healthy college-aged participants with no history of major musculoskeletal injury or surgery. They were 23.1 years of age (SD 2.1 years), had a height of 1.70 m (SD 11.4 m), and a mass of 71.0 kg (SD 17.3 kg). All the subjects signed a letter of informed consent approved by the University's Research Ethics Board before participation in the study.
Kinematic data were collected using a 9-camera Qualysis (Gothenburg, Sweden) Oqus 300 motion capture system. Retroreflective markers were fixed to rigid body clusters and securely attached bilaterally to the subject's foot, shank, and thigh using tensor bandages and athletic tape. There was also a single rigid body cluster attached to the sacrum and pelvis. The subject then stood with each foot on a separate AMTI (Newton, MA, USA) force platform to allow for the simultaneous collection of ground reaction force data along with the markers kinematics. Marker clusters were attached on both legs, and ground reaction force data were collected for both legs; however, for this current analysis, only the data for the subject's dominant leg (i.e., the leg they would preferentially use to kick a ball) was analyzed.
A Run-Technologies (Mission Viejo, CA, USA) Myopac 8 channel electromyographic (EMG) system was used to collect all EMG data (differentially amplified with a gain of 1,000, bandpass 20–540 Hz, common mode rejection ratio (CMRR) > 80 dB, input impedance >1015). Following standard skin preparation, silver-silver chloride electrodes were applied in a bipolar configuration (3.0 cm center-to-center) in line with the muscle fibers over the biceps femoris (BF) representing the hamstring muscle group, rectus femoris (RF) representing the quadriceps muscle group, and the GM muscles. It was also ensured that the motor point of each muscle was avoided. The reference-ground electrode was then attached to the tibial tuberosity of the participant's test (dominant) leg. The EMG electrodes were only attached unilaterally to the subject's dominant leg. Kinematic data (at 100 Hz), along with EMG and ground reaction force data (at 2,000 Hz) were collected simultaneously using Qualysis Track Manager software (Gothenburg, Sweden).
Before testing, it was ensured that the subjects had not significantly deviated from their normal daily habits of nutrition, hydration, and sleep. After the application of the EMG electrodes, the subjects performed 3 maximum voluntary isometric contractions (MVICs): knee flexion (BF), hip extension (GM), knee extension (RF). The knee flexion and hip extension MVICs were performed with the subject in a prone position, whereas the knee extension MVIC was performed with the subject in a seated position. The hip extension MVIC was also done with the knee flexed to 90° in an attempt to decrease the involvement of the hamstring muscle group. The subjects were asked to produce a maximum isometric exertion of each muscle group while being manually resisted. The peak muscle activity achieved for each muscle group during these 3 tasks was then used to normalize squat EMG data as has been done previously (9,12). There was also a 1-second resting EMG trace recording taken before data collection, and the average resting signal was subtracted off all EMG and squat trial EMG data.
After the MVIC trials and the application of motion tracking rigid bodies, the subjects stood with 1 foot on each force plate and were given dumbbells to hold in each hand. They were then instructed how to perform the 2 different methods of squats: the CBS and the RS. The RS involved the subjects squatting with the dumbbells held at shoulder level the whole time (Figure 1A).The CBS was performed with the dumbbells starting on the subject's thigh, then flexing 90° at the shoulder during the eccentric phase of the squat and extending the shoulder back to neutral during the concentric phase of the squat (Figure 1B). A metronome was also set to the preferred speed of movement for each subject (which was approximately 60 b·min−1 for all the subjects) to ensure that each subject performed all the trials at the same cadence and speed. They were given sufficient practice trials to ensure they were comfortable with the speed of movement and each type of squat movement. The subjects were asked to begin each squat on 1 beat, be at the bottom of the squat on the second beat, and have returned to a standing position on the third beat. Any trial in which the subject did not maintain a consistent pace was disregarded and the trial was repeated. This ensured that all trials for both squat conditions were performed at the same speed of movement. During testing, 5 trials were performed for each squat condition (RS and CBS) while marker motion, force plate, and EMG data were simultaneously recorded. There was at least 1 minute of rest given after practice trials and between test squats to ensure that fatigue was not a factor. The order of the 10 total trials was randomized for each subject.
After all the squat trials, the subject stood in view of the cameras while a anatomical landmark trial was completed. This involved using a specially designed probe to identify bony landmarks so that subject-specific segments could be created during processing. The subjects also performed 3 calibrated motions so the joint centers could be calculated. The hip calibrated motion involved performing 3 complete ‘hula-hoop’ hip rotations in each direction. The knee and ankle calibrated motion involved performing simple knee flexion and extension and ankle dorsi and plantar flexion movements, respectively.
Visual 3D (C-Motion Inc., Rockville, MD, USA) software was used to process all squat trial data. This software was used to filter raw marker data (fourth-order, low-pass, double-pass Butterworth filter with a 6-Hz cut-off frequency) (8) and to calculate joint angles at the hip and knee during the squat trials. These joint kinematics were calculated using the Cardan-Euler representation, which finds the orientation of the distal segment with respect to the reference proximal segment using x (flexion/extension), y (ab/adduction), z (axial rotation) sequence of rotations (40). It also combined motion, force, and anthropometric data using a standard inverse dynamics link segment model to calculate the net external moments at the hip and knee. Subject-specific anthropometric information used in these calculations was estimated from their height and weight (14). All the joint moments were normalized by body mass (newton meter per kilogram) and reported in the coordinate system of the distal segment of the joint. It should be noted that joint angles and moments were calculated in all 3 dimensions; however, this study only examined these angles and moments in the sagittal plane. Also, the peak knee flexion angle during the squat was used to define the concentric and eccentric phases of the squat movement for all analyses; and variables calculated from each of the 5 trials per condition were averaged together for each subject.
The EMG data were full-wave rectified and filtered using a fourth-order, low-pass, double-pass Butterworth filter with a 6-Hz cut-off frequency (39). Integrated EMG (IEMG) was then calculated for the concentric and eccentric phases of the 3 muscles during the squat using the following equation (17):
where n is the number of data points for the concentric and eccentric phases, yi is the EMG data at time (t), and Δt is the sampling interval (1/2,000 seconds). The IEMG is reported in millivolts.
All EMG waveforms were also converted into %MVIC units by dividing by the peak MVIC activity of each muscle (BF, RF, GM).
Angular work was also calculated at the hip and knee joints by multiplying the external joint moment (in newton meter per kilogram) by the joint angular displacement (in radians). At both the hip and knee joints, negative work corresponded to the eccentric phase of the squat and positive work corresponded with the concentric phase.
Key outcome measures were as follows: (a) the joint ranges of motion (ROMs) for the knee and hip during the entire squat movement; (b) the peak external net joint moments achieved at the bottom during the squat for the knee and hip joints; (c) total negative (eccentric) and positive (concentric) angular work performed at the knee and the hip; and (d) the peak EMG activity (in %MVIC) achieved during the entire squat and IEMG during the concentric and eccentric phases of the squat movement for the GM, BF, and RF muscles. There were 11 separate 2 × 2 (squat condition × sex) factorial mixed model analyses of variance run on each one of the variables, except for the IEMG data. The 6 IEMG variables were calculated using raw (nonnormalized) EMG data; therefore, the between subjects comparison were not examined on these variables and only the within-subjects comparison (squat type) were performed. For all analyses, statistical significance was set at p ≤ 0.05.
The results for the joint kinetic data (hip, knee) are reported in Table 1. It should be noted that there was a main effect (p < 0.05) for squat type in both the peak knee (F = 7.9, p = 0.009, effect size
) and hip (F = 12.5, p = 0.001, effect size
) moments. There was a larger magnitude of hip moment in the CBS, whereasthere was a larger magnitude of knee moment in the RS. Also, the hip moment displayed a main effect for sex (F = 7.1, p = 0.013, effect size
) as the male subjects had an increased magnitude of peak hip moment as compared with the female subjects. There was no main effect for sex (p < 0.05) in the peak knee moment (F = 0.9, p = 0.348, effect size
Table 1 also displays the joint work done by the hip and the knee in both the concentric and eccentric phases. There was a main effect for all of the joint work variables calculated: eccentric hip work (F = 126.4, p < 0.001, effect size
), concentric hip work (F = 70.2, p < 0.001, effect size
), eccentric knee work (F = 22.2, p < 0.001, effect size
), concentric knee work (F = 32.9, p = 0.001, effect size
). It should be noted that for both the concentric and eccentric phases, the knee does more work during the RS and the hip does more work during the CBS. There were also main effects for sex in the work done by the hip in both the eccentric (F = 10.9, p = 0.003, effect size
and concentric (F = 11.1, p = 0.002, effect size
phases; but no main effect for sex in the knee work done in either phase (eccentric—F = 6.2, p = 0.145, effect size
, concentric—F = 4.9, p = 0.153, effect size
Table 2 displays the results of the peak EMG activation data for the GM, RF, and BF. It should be noted that there was a main effect for squat type (p < 0.05) in the peak GM (F = 8.0, p = 0.008, effect size
) and RF (F = 13.7, p = 0.001, effect size
) EMG activation data but not in the BF (F = 2.6, p = 0.117, effect size
). The CBS squat produced a higher peak activation level of the GM, whereas the RS produced a higher peak activation level of the RF. There was also a main effect for sex in the BF peak EMG activation data (F = 5.1, p = 0.03, effect size
), as the women had an increased peak activation of the hamstring muscles as compared with the men across both types of squat; however, there was no main effect for sex in the GM (F = 0.49, p = 0.490, effect size
) or the RF (F = 1.7, p = 0.202, effect size
Table 2 also displays the IEMG data for the GM, RF, and BF during the concentric and eccentric phases. The IEMG data revealed main effects for squat type in RF activation during both the eccentric (F = 22.6, p < 0.001, effect size
) and concentric phases (F = 29.5, p < 0.001, effect size
). There were also main effect for squat type in the IEMG data of the concentric phases for the GM (F = 16.0, p < 0.001, effect size
) and BF (F = 7.7, p = 0.010, effect size
). These data revealed increased activation of the hip extensors (GM and BF) during the CBS in the concentric phase, whereas the knee extensors (RF) were more active in the RS during both the eccentric and concentric phases. There was no main effect for the IEMG data in the eccentric phases of the GM (F = 3.3, p = 0.082, effect size
) or the BF (F = 1.3, p = 0.260, effect size
The hip and knee joint ROM data are reported in Table 3. There was a main effect for squat type (p < 0.05) in both the hip (F = 47.9, p < 0.001, effect size
) and knee (F = 5.5, p = 0.026, effect size
) ROM data. The hip displayed a greater ROM during the RS, whereas the knee displayed a great ROM during the CBS. However, there was no main effect for sex in either the hip (F = 0.9, p = 0.342, effect size
) or knee (F = 2.3, p = 0.137, effect size
This study found that the CBS increased the sagittal plane loading of the hip joint while decreasing the sagittal plane loading on the knee as compared with the RS. This is most likely because of the increased mass placed anteriorly during the CBS when the dumbbells and arms are moved anteriorly with shoulder flexion. This coincides with the findings of previous work that has suggested that forward trunk lean during a drop jump task increases the demand on the hip while decreasing the demand on the knee as compared with landing with a more erect trunk (29), as forward leaning of trunk would also shift more mass anteriorly. This shift in external loading that increases the loading of the hip and decreases the loading of the knee during the CBS also led to a shift in muscular activation levels by increasing the peak activity of the larger proximal hip extensor (GM) and decreasing the peak activity of the relatively smaller distal knee extensors (RF). This coincides with the results of Blackburn and Padua (4), because they found that increasing the mass placed anteriorly during jump landing (i.e., forward trunk lean) caused decreased quadriceps activation; however, hip extensor muscle activity was not measured in this study.
Based on these results, it can be suggested that the CBS produces a more hip-dominant pattern of movement as compared with the RS, which produces a more knee-dominant pattern of movement. However, even though there are large effect sizes for most of the comparisons, the magnitude of these changes in external loading and muscular activation levels are small. It is unknown as to how these small alterations in muscle activation levels could change the overall patterns of movement if the CBS is performed as part of a long-term corrective exercise program. Future research should investigate both the acute and long-term effects of using a CBS in a training program to see if this can produce a more hip-dominant and less knee-dominant patterns of movement.
The use of a CBS as part of a training program could have a wide range of uses clinically. It has been shown that those with a torn anterior cruciate ligament (ACL) adopt a gait pattern where they attempt to reduce the contraction of the quadriceps (3) during gait. This is believed to be an attempt to decrease the anterior shear force produced on the proximal tibia when the quadriceps contracts. Therefore, performing a CBS exercise to decrease the recruitment of the quadriceps and increase the recruitment of the hip extensors may be advisable. This may allow for the subjects to begin rehabilitation exercise much sooner after ACL replacement surgery, because the graft could be spared some of the stresses of the anteriorly directed forces applied on the tibia with excessive quadriceps contraction. The CBS may also be useful in prevention of ACL injury as the decreased quadriceps activation could potentially help in reducing quadriceps to hamstring ratio and decreasing the “quad dominance” that has been shown to increase ACL injury risk (24).
Training a movement pattern that can reduce external knee moments and the resultant quadriceps activation may also have implications in the treatment and prevention of several other knee pathologic conditions. For example, it has been shown that increased quadriceps strength (force) increases the rate of cartilage wear over an 18-month period in those with malaligned knees and laxity in the knee joint (32). This may be because of the combination of compressive and shear forces placed on the knee with quadriceps contraction when the knee is misaligned (33), as it has been shown that shear stresses are highly detrimental to cartilage health (19,38). Also, training a movement pattern that decreases quadriceps force could also lead to smaller patellofemoral compressive forces, especially at larger angles of knee flexion, because this would decrease the forces in the quadriceps and patellar tendons (7,8). Therefore, using the CBS to train a pattern of movement that reduces the external knee moment and the resultant quadriceps activation may also be beneficial in prevention and treatment of knee osteoarthritis (OA) and patellofemoral joint pain.
It has been suggested that increased GM activation during any hip extension movement would lead to more precise control of the femur and less stress on the hip joint (21,22). More specifically, a weakness in the GM results in increased anterior hip joint forces during hip extension (21), which could lead to clinical conditions such as acetabular labrum tears and eventually hip OA. Because it has been discovered that the majority of hip OA cases have the cartilage wear occurring on the anterior surface of the acetabulum (25), it can be suggested that the CBS could be useful in increasing the activation of the GM during concentric hip extension movements. This could help in decreasing the anteriorly directed hip joint force and reduce the stress on the diseased part of the hip joint in those with hip OA on the anterior surface of the acetabulum. However, finding other ways to increase the activation of the GM during eccentric contraction is warranted, because the CBS did not increase total eccentric GM activity during the squat.
The CBS also altered the joint ROM at the knee and the hip as compared with the RS. The differences in the joint ROM between the 2 squat conditions were small as there was approximately 3.5° less flexion at the hip and 1.6° more flexion at the knee in the CBS. Although this is a small absolute change in angles, the effect sizes for these 2 comparisons were relatively strong, especially at the hip joint (effect size of 0.62). This is because all the subjects had same small changes in joint angles with the 2 different squat conditions, and this may be related to the fact that the CBS also had decreased peak (%MVIC) and total (IEMG) RSF activation. During the squat movement, the RSF eccentrically lengthens by an extremely small amount (<2%); therefore, one could characterize its role as being isometric in the dissipation of knee extensor torque (30). Because neural drive to the muscle is smaller in the CBS as compared with that in the RS, the RSF muscle would be less “stiff” in the CBS, which would produce less hip flexion and allow for more knee flexion in the CBS as compared with the RS. The origin of the RSF muscle on the anterior-inferior iliac spine could also result in an anterior pelvic tilt during the squat movement if this muscle has an increased neural drive and was not allowed to eccentrically lengthen (as would be the case in the RS). This anterior pelvic tilt would appear as an increase in hip flexion but would most likely also increase the lordotic curvature in the lumbar spine. Because increased lordosis (or spinal extension) has been shown to increase the contact forces in the facet joints at L4-5 (34), it can be hypothesized that reinforcing a movement pattern with increased RSF activation may increase the loading of the facet joints. In those with greatly reduced neural drive to the hip extensors, this could lead to facet joint arthritis and perhaps even spondilolisis and spondylolisthesis. However, this is strictly speculation because the curvature of the lumbar spine was not measured in this work. Therefore, future research should examine the effect of a CBS technique on the sagittal plane stability of the pelvis and lumbar spine during the squat.
Although it has been suggested that a more hip-dominant and less knee-dominant pattern of movement is beneficial, it should be noted that there are still relatively large quadriceps contractions even during the CBS (>50% MVIC). Therefore, the CBS does not completely unload the knee extensors, it simply shifts some of the loading onto the much larger hip extensors. The results of this study also suggest that if the goal of training is to strengthen the knee extensors in a weight-bearing closed chain exercise, positioning the mass more posteriorly (as was done in the RS) would be beneficial.
There were also sex differences in squat biomechanics across both conditions. The female subjects had decreased external hip loading (peak hip moment and concentric and eccentric hip joint work) and increased peak hamstring activity. It appears that the female subjects in this study adopt a less hip-dominant movement pattern that preferentially activates the 2 joint hamstrings rather than the 1 joint GM to overcome the much smaller (as compared with the male subjects) external hip moment created during the squat. This result is supported by the increased rate of noncontact ACL injury in female athletes (2,13). It can be hypothesized that this squat pattern observed in the female subjects may lack the frontal and transverse plane control of the femur that comes with activation of the large, proximally located gluteal muscles because this muscle is not needed to control the smaller external hip moment. This also results in a compensatory increase in hamstring activity, which could help decrease the anterior translation of the tibia and help relieve some of the stress on the ACL. However, if this were true, one would also expect decreased GM activation in the female subjects as compared with the male subjects, which we did not find. These results also contradict those of Youdas et al. (41), where it was found that female subjects had more quadriceps activation and less hamstring activation than did male subjects during single-limb squatting. These conflicting results indicate that the sex differences in movement pattern biomechanics are complex and require a much more thorough investigation.
The drawback of the CBS is that the amount of mass that can be used to counterbalance is limited by the amount of weight the shoulders can support through the 90° of flexion. This would also limit the ability to overload the squat pattern using the CBS and future research should examine whether it is possible to produce a training effect using the CBS. A CBS may also be contraindicated for those with certain shoulder or spine pathologic conditions. For example, placing more mass anteriorly could increase the moment on the lumbar spine and therefore, examining the “costs” of this exercise on other joints is warranted.
Perhaps the best use of a CBS technique would be in the initial stages of a training program to establish a more hip-dominant and less knee-dominant pattern of movement before progressing to more complicated movements. It could also be used as a warm-up or movement preparation routine to help establish a more hip-dominant pattern before performing other movements. It should also be pointed out that this study simply investigated the differences between 2 different squat techniques. Future research is needed to examine both the long-term and short-term training effects of using CBS movements to determine whether they have the ability to alter the mechanics of a subject's natural movement patterns.
If the goal of a squat training program is to increase the dominance of the hip extensors in producing a squat movement, there should be an attempt to increase the mass placed anteriorly. However, if the goal of a squat training program is to strengthen the knee extensors, there should be an attempt to increase the mass placed posteriorly.
The authors would like to thank their graduate students (Melinda Pittman, Yasuo Sakurai, and Lisa Wilson) for all of their help with data collection, processing, and organization. There was no funding received for this work.
1. Abelbeck KG. Biomechanical model and evaluation of a linear motion squat type exercise. J Strength Cond Res 16: 516–524, 2002.
2. Arendt E, Dick R. Knee injury patterns among men and women in collegiate basketball and soccer—NCAA data and review of literature. Am J Sports Med 23: 694–701, 1995.
3. Berchuck M, Andriacchi TP, Bach BR, Reider B. Gait adaptations by patients who have a deficient anterior cruciate ligament. J Bone Joint Surg Am 72A: 871–877, 1990.
4. Blackburn JT, Padua DA. Sagittal-plane trunk position, landing forces, and quadriceps electromyographic activity. J Athl Train 44: 174–179, 2009.
5. Boling MC, Padua DA, Creighton RA. Concentric and eccentric torque of the hip musculature in individuals with and without patellofemoral pain. J Athl Train 44: 7–13, 2009.
6. Chang A, Hayes K, Dunlop D, Song J, Hurwitz D, Cahue S, Sharma L. Hip abduction moment and protection against medial tibiofemoral osteoarthritis progression. Arthritis Rheum 52: 3515–3519, 2005.
7. Escamilla RF. Knee biomechanics of the dynamic squat exercise. Med Sci Sports Exerc 33: 127–141, 2001.
8. Escamilla RF, Fleisig GS, Lowry TM, Barrentine SW, Andrews JR. A three-dimensional biomechanical analysis of the squat during varying stance widths. Med Sci Sports Exerc 33: 984–998, 2001.
9. Escamilla RF, Fleisig GS, Zheng N, Barrentine SW, Wilk KE, Andrews JR. Biomechanics of the knee during closed kinetic chain and open kinetic chain exercises. Med Sci Sports Exerc 30: 556–569, 1998.
10. Flanagan SP, Salem GJ. Lower extremity joint kinetic responses to external resistance variations. J Appl Biomech 24: 58–68, 2008.
11. Gans C. Fiber architecture and muscle function. Exerc Sport Sci Rev 10: 160–207, 1982.
12. Gullett JC, Tillman MD, Gutierrez GM, Chow JW. A biomechanical comparison of back and front squats in healthy trained individuals. J Strength Cond Res 23: 284–292, 2009.
13. Gwinn DE, Wilckens JH, McDevitt ER, Ross G, Kao TC. The relative incidence of anterior cruciate ligament injury in men and women at the United States Naval Academy. Am J Sports Med 28: 98–102, 2000.
14. Hanavan EP. A mathematical model of the human body. Aero Medical Research Laboratory Technical Report. Air Force Base, Ohio, 1964.
15. Ito J, Moriyama H, Inokuchi S, Goto N. Human lower limb muscles: An evaluation of weight and fiber size. Okajimas Folia Anat Jpn 80: 47–55, 2003.
16. Jacobs JJ, King T. US bone and joint decade prepares for the future. Arth Care Res 61: 1470–1471, 2009.
17. Kamen G, Gabriel DA. Essential of Electromyography. Champaign, IL: Human Kinetics, 2010.
18. Kritz M, Cronin J, Hume P. The bodyweight squat: A movement screen for the squat pattern. Strength Cond J 31: 76–85, 2009.
19. Lee MS, Trindade MCD, Ikenoue T, Schurman DJ, Goodman SB, Smith RL. Effects of shear stress on nitric oxide and matrix protein gene expression in human osteoarthritic chondrocytes in vitro. J Orthop Res 20: 556–561, 2002.
20. Leetun DT, Ireland ML, Willson JD, Ballantyne BT, Davis IM. Core stability measures as risk factors for lower extremity injury in athletes. Med Sci Sports Exerc 36: 926–934, 2004.
21. Lewis CL, Sahrmann SA, Moran DW. Anterior hip joint force increases with hip extension, decreased gluteal force, or decreased iliopsoas force. J Biomech 40: 3725–3731, 2007.
22. Lewis CL, Sahrmann SA, Moran DW. Effect of position and alteration in synergist muscle force contribution on hip forces when performing hip strengthening exercises. Clin Biomech 24: 35–42, 2009.
23. McKean MR, Dunn PK, Burkett BJ. Quantifying the movement and the influence of load in the back squat exercise. J Strength Cond Res 24: 1671–1679, 2010.
24. Myer GD, Brent JL, Ford KR, Hewett TE. Real-time assessment and neuromuscular training feedback techniques to prevent anterior cruciate ligament injury in female athletes. Strength Cond J 33: 21–35, 2011.
25. Neumann G, Mendicuti AD, Zou KH, Minas T, Coblyn J, Winalski CS, Lang P. Prevalence of labral tears and cartilage loss in patients with mechanical symptoms of the hip: Evaluation using MR arthrography. Osteoarthritis Cartilage 15: 909–917, 2007.
26. Niemuth PE, Johnson RJ, Myers MJ, Thieman TJ. Hip muscle weakness and overuse injuries in recreational runners. Clin J Sport Med 15: 14–21, 2005.
27. Pitt-Brooke J, Reid A, Lockwood J, Kerr K. Rehabilitation
of Movement: Theoretical Basis of Clinical Practice. London, United Kingdom: WB Saunders, 1998.
28. Pollard CD, Sigward SM, Powers CM. Limited hip and knee flexion during landing is associated with increased frontal plane knee motion and moments. Clin Biomech 25: 142–146, 2010.
29. Powers CM. The influence of abnormal hip mechanics on knee injury: A biomechanical perspective. J Orthop Sports Phys Ther 40: 42–51, 2010.
30. Robertson DGE, Wilson JJ, Pierre TAS. Lower extremity muscle functions during full squats. J Appl Biomech 24: 333–339, 2008.
31. Sahrmann SA. Diagnosis and Treatment of Movement Impairment Syndromes. St. Louis, MO: Mosby, Inc., 2002.
32. Sharma L, Dunlop DD, Song J, Hayes KW. Quadriceps strength and osteoarthritis progression in maligned and lax knees. Ann Intern Med 138: 613–619, 2003.
33. Shin CS, Chaudhari AA, Dyrby CO, Andriacchi TP. Influence of patellar ligament insertion angle on quadriceps usage during walking in anterior cruciate ligament reconstructed subjects. J Orthop Res 27: 730–735, 2009.
34. Shirazi-Adl A, Drouin G. Load-bearing role of facets in a lumbar segment under sagittal plane loadings. J Biomech 20: 601–613, 1987.
35. Simonsen EB, DyhrePoulsen P, Voigt M, Aagaard P, Fallentin N. Mechanisms contributing to different joint moments observed during human walking. Scand J Med Sci Sports 7: 1–13, 1997.
36. Taanila H, Suni J, Pihlajamäki H, Mattila VM, Ohrankämmen O, Vuorinen P, Parkkari J. Musculoskeletal disorders in physically active conscripts: A one-year follow-up study in the Finnish Defence Forces. BMC Musculoskelet Disord 10: 89, 2009.
37. Taunton JE, Ryan MB, Clement DB, McKenzie DC, Lloyd-Smith DR, Zumbo BD. A retrospective case-control injuries analysis of 2002 running. Br J Sports Med 36: 95–101, 2002.
38. Tomatsu T, Imai N, Takeuchi N, Takahashi K, Kimura N. Experimentally produced fractures of articular-cartilage and bone–the effects of shear forces on the pig knee. J Bone Joint Surg Br 74: 457–462, 1992.
Winter DA. Biomechanics and Motor Control of Human Movement (2nd ed.): John Wiley and Sons, Inc., New York 205, 1990 .
40. Woltring HJ. 3-D attitude representation of human joints—A standardization proposal. J Biomech 27: 1399–1414, 1994 .
41. Youdas JW, Hollman JH, Hitchcock JR, Hoyme GJ, Johnsen JJ. Comparison of hamstring and quadriceps femoris electromyographic activity between men and women during a single-limb squat on both a stable and labile surface. J Strength Cond Res 21: 105–111, 2007.