In the outpatient setting patellofemoral pain syndrome (PFPS) is the most common type of knee pain and accounts for 25-30% of all knee pathologies treated (10,11,16). Because the etiology of PFPS is poorly understood and multifaceted, it remains one of the most difficult clinical challenges in rehabilitative medicine (39). PFPS primarily affects younger active individuals approximately 18-40 yr old (although older individuals can also be affected), athletes and nonathletes (11,16,21), and males and females (9). Although patellofemoral rehabilitation can be a long and arduous process, the use of appropriate exercises can improve this process by decreasing rehabilitation time and improving function (4,18,26,40,41).
High patellofemoral joint compressive force (patellofemoral force) can result in PFPS from numerous soft tissues, such as synovial plicae, infrapatellar fat pad, retinacula, joint capsule, and patellofemoral ligaments (3). Patellofemoral force can also elevate subchondral bone stress (patellofemoral force per unit patella contact area) in the patellofemoral joint (2). Because the subchondral bone plate is rich in pain receptors (42), increased subchondral bone stress from high patellofemoral force may also result in PFPS (3). Patellofemoral joint stress can result in a cartilage degeneration and a decrease in the ability of the cartilage to distribute patellofemoral force (2). Therefore, understanding what patellofemoral force and stress magnitudes are generated among patellofemoral rehabilitation exercises may be helpful to clinicians when prescribing therapeutic exercises to individuals with PFPS.
Weight-bearing exercises, such as the squat, are frequently used during patellofemoral rehabilitation and are specific to many functional activities such as walking, running, and jumping (4,18,26,40,41). The use of weight-bearing exercises have been shown to be effective, both in short- and long-term outcomes, in decreasing PFPS and in enhancing functional performance (4,18,26,40,41). Therapists use these types of exercises to minimize PFPS and muscle loss, to strengthen hip and thigh musculature, to enhance balance and stability, and to minimize the risk of future injuries and associated costs of health care (35). However, all weight-bearing exercises may not produce similar magnitudes of patellofemoral force and stress. Moreover, using varying techniques within a weight-bearing exercise may also affect patellofemoral force and stress.
Wall squats and one-leg squats are common weight-bearing exercises used by athletes and other individuals with healthy knees to train the hip and the thigh musculature. Therapists and trainers also use wall squats, one-leg squats, and other similar weight-bearing exercises during patellofemoral rehabilitation for PFPS patients to allow patients to recover faster and return to function earlier (4,18,26,40,41). Wall squats can be performed with varying techniques, such as positioning the heels farther from 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, while positioning the heels closer to the wall typically results in anterior knee translation beyond the toes at the lowest position of the squat. Performing a one-leg squat also causes the stance knee to translate forward beyond the toes at the lowest position of the squat. Clinicians and trainers often believe that anterior translation of the lead knee beyond the toes during squatting type exercises increases patellofemoral force and stress, but there is currently no evidence to support this belief.
Understanding how patellofemoral force and stress vary among weight-bearing exercises will allow clinicians and trainers to prescribe safer and more effective knee rehabilitation treatment to patients with PFPS or to athletes during training. For example, if performing the wall squat with the heels closer to the wall (causing greater anterior translation of the lead knee over the toes) results in greater patellofemoral force and stress compared with performing the wall squat with the heels farther away from the wall (causing the lead knee to be maintained over the foot), a wall squat with the heels closer to the wall may be discouraged during training and rehabilitation. There may also be differences in patellofemoral force and stress over a specific knee flexion range of motion between wall squat and one-leg squat exercises. Excess patellofemoral force and stress over time may lead to PFPS in individuals with asymptomatic patellofemoral joints or may exacerbate PFPS in patients with patellofemoral pathology and retard the rehabilitation process.
Currently, patellofemoral force and stress magnitudes during the wall squat or one-leg squat are unknown. Therefore, the purpose of this study was to compare patellofemoral force and stress during the wall squat with the feet farther away from the wall (wall squat long), the wall squat with the feet closer to the wall (wall squat short), and the one-leg squat. It was hypothesized that patellofemoral force and stress would be greater in the one-leg squat compared with the wall squat long, similar between the one-leg squat and the wall squat short and greater in the wall squat short compared with the wall squat long.
Eighteen healthy individuals (nine males and nine females) without a history of patellofemoral 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 the wall squat and the 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 appropriate regression equations 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.
The wall squat began with the right foot on an AMTI force platform (Model OR6-6-2000; Advanced Mechanical Technologies, Inc.) 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 with the ground (resulting in approximately 90-100° of knee flexion in the wall squat long and approximately 100-110° of knee flexion in the wall squat short), 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. 1A) and wall squat short (Fig. 1B). 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 at the lowest position of the squat (Fig. 1A), with the knees above the ankles, 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. This distance was chosen because the shorter heel-to-wall distance for the wall squat short resulted in the anterior surface of the knee translating beyond the distal end of the toes at the lowest position of the wall squat short (Fig. 1B), which is typically discouraged by clinicians and trainers.
The one-leg squat started with the subject standing on one leg with the right foot on the AMTI 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 (subject preference). From this position, the subject slowly flexed the right knee and squatted down until the right thigh was approximately parallel with the ground (resulting in approximately 100-110° of knee flexion) with the trunk tilted forward approximately 30-40° (Fig. 2). 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. Like the wall squat short, at the lowest position of the one-leg squat, the distal surface of the knee translated beyond the distal end of the toes (Fig. 2).
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 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 the wall squat short and 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 between electrodes. 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 (1), 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 semitendinosus); e) lateral hamstrings (biceps femoris); and f) gastrocnemius (midpoint between medial and lateral heads).
Spheres (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.
Once the electrodes and the spheres were positioned, the subject warmed up and practiced the exercises as needed, and data collection 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 an AMTI 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, 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 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.
Subsequent to completing all exercise trials, EMG data were collected during maximum voluntary isometric contractions (MVIC) to normalize the EMG data collected during each exercise (14). 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 (14). 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 (14). 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 plantar flexion (14). Two 5-s trials were randomly collected for each MVIC.
Video images for each reflective marker were tracked and digitized in three-dimensional space with Peak Performance software, using the direct linear transformation calibration method (31). Testing of the accuracy of the calibration system resulted in reflective balls 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 (14). 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 (14).
Raw EMG signals were full-waved rectified, smoothed with a 10-ms moving average window, and linear enveloped throughout the knee range of motion for each repetition. These EMG data were then normalized for each muscle and 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 for the three repetitions (trials) were then averaged at corresponding knee angles between 0° and 90° and were used in the biomechanical model described below.
As previously described (14,44), a biomechanical model of the knee was used to continuously calculate patellofemoral forces throughout a 90° knee range of motion during the knee flexing (descent) phase (0-90°) and knee extending (ascent) phase (90-0°) of the wall squat and one-leg squat (Fig. 3). Resultant force and torque equilibrium equations were calculated using inverse dynamics and the biomechanical knee model (14,44). Moment arms for muscle forces lines of action angles for muscles were represented as polynomial functions of the knee flexion angle using data from Herzog and Read (19).
Quadriceps, hamstrings, and gastrocnemius muscle forces were calculated as previously described (14,44). Because the accuracy of calculating muscle forces depends on accurate calculations of a muscle's physiological cross-sectional area (PCSA), a 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 F m(i) was modified by the following equation:
where A i was the PCSA of the ith muscle, σ m(i) was the MVIC force per unit PCSA of the ith muscle, EMGi and MVICi were the EMG window averages of the ith muscle EMG during exercise and MVIC trials, c i was a weight factor (values given below) adjusted in a computer optimization program to minimize the difference between the resultant torque from the inverse dynamics (T res) and the resultant torque calculation from the biomechanical model (T mi), k li represented 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) (37), and k vi represented each muscle's force-velocity relationship based on a Hill-type model for eccentric and concentric muscle actions using the following equations from Zajac (43) and Epstein and Herzog (13):
with F 0 representing isometric muscle force, v = velocity, a = 0.32F 0, b = 3.2l 0 ·s−1, and C = 1.8. Muscle force from eccentric contractions was scaled up by 1.8 times the isometric muscle force F 0. Forces generated by the knee flexors and extensors at MVIC were assumed to be linearly proportional to their PCSA. Muscle force per unit PCSA at MVIC was 35 N·cm−2 for the knee flexors and 40 N·cm−2 for the quadriceps (6,24,25,38).
The objective function used to determine each ith muscle's coefficient c i was as follows:
subject to c low ≤ c i ≤ c high, where c low and c high were lower and upper limits for c i, 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. The torques predicted by the EMG-driven model matched well (<2%) with the torques generated from the inverse dynamics. The assumptions associated with this model are 1) that the torque from cruciate ligament forces was ignored and 2) that other forces and torques out of the sagittal plane were ignored.
Patellofemoral force was a function of patellar tendon force and quadriceps tendon force. Patellar tendon force was calculated by the quadriceps tendon force and the ratio of the patellar tendon force and the quadriceps tendon force, as previously described (33,34). The angles between the patellar tendon, quadriceps tendon, and patellofemoral joint were expressed as functions of knee angle (33,34).
Patellofemoral stress, which was calculated every 10° between 0° and 90° knee angles, was expressed as the ratio of patellofemoral force, calculated from the biomechanical model described above (14,44), and the patellar contact area. Patellar contact areas were determined at 10° intervals between 0° and 90° knee angles. Contact areas from in vivo MRI data from Salsich et al. (30), who also used both male and female subjects with healthy knees and had them perform weight-bearing exercise using resistance, were used at 0° (146 mm2), 20° (184 mm2), 40° (290 mm2), and 60° (347 mm2) knee angles. These four contact area values formed a near linear relationship as a function of knee angle, resulting in a line of best fit equation of y = 3.55x + 135 (r = 0.98), with y = contact area and x = knee angle. This line of best fit equation was used to determine contact areas at 10° knee angle (171 mm2), 30° knee angle (242 mm2), and 50° knee angle (313 mm2). The contact areas above at 40°, 50°, and 60° knee angles were used to develop the line of best fit equation, y = 2.81x + 176 (r = 0.99), which was used to determine contact areas at 70° knee angle (373mm2), 80° knee angle (401 mm2), and 90° knee angle (429 mm2). Like the current study, a near linear relationship between patellar contact area and knee angles has been reported between 0° and 90° knee angles in several studies involving weight-bearing exercises (2,8,20,27,30).
To determine significant differences among the wall squat long, wall squat short, and one-leg squat, patellofemoral force and stress were analyzed every 10° during the 0-90° descent phase and the 90-0° ascent phase using a one-factor repeated-measure ANOVA. Bonferroni t-tests were used to assess pairwise comparisons. To minimize the probability of type I errors secondary to the use of a separate ANOVA for each knee angle, a Bonferroni adjustment was performed with the level of significance established at 0.0025 (0.05/20 knee angles). A separate set of analyses was not performed for patellofemoral joint stress values because stress values for each knee angle were derived from dividing force data by a constant, therefore not affecting statistical results.
Descriptive patellofemoral force and stress data during the wall squat and the one-leg squat are shown in Figures 4 and 5. Visual observation of the data indicates that patellofemoral force and stress progressively increased during the squat descent and progressively decreased during the squat ascent, except between 90° and 70° during the squat ascent in which patellofemoral force and stress progressively increased. During the squat descent, there were no significant differences in patellofemoral force and stress among the three squat exercises. During the squat ascent, there were significant differences in patellofemoral force and stress among the three squat exercises at 90° knee angle (P = 0.002), 80° knee angle (P = 0.002), 70° knee angle (P < 0.001), and 60° knee angle (P = 0.001). Patellofemoral force and stress were significantly greater at 90° knee angle in the wall squat short compared with wall squat long and one-leg squat, significantly greater at 70° and 80° knee angles in the wall squat short and long compared with the one-leg squat, and significantly greater at 60° knee angle in the wall squat long compared with the wall squat short and one-leg squat. At the lowest position of the wall squat short, the knees translated beyond the toes 9± 2 cm, whereas at the lowest position of the one-leg squat the knee translated beyond the toes 10 ± 2 cm.
As hypothesized, patellofemoral force and stress were greater during the wall squat short compared with the wall squat long, but only at 90° knee angle during the squat ascent. At 90° knee angle, the knees translated beyond the toes in the wall squat short, but the knees remained over the feet in the wall squat long. Also, as the knees translated forward beyond the toes in the wall squat short, the orientation of the leg tilted forward (Fig. 1B), changing the direction of the patellar tendon force, which potentially may increase patellofemoral force compared with the vertical leg position in the wall squat long at 90° knee angle (Fig. 1A). Therefore, anterior knee translation and forward tilt of the leg may be related to increased patellofemoral force and stress. The results of the current study support the belief of many clinicians that anterior knee translation beyond the toes while performing squatting type exercises increases patellofemoral force and stress compared with maintaining the knees over the feet.
The wall squat short and the one-leg squat both resulted in similar amounts of anterior knee translation at maximum knee flexion, but patellofemoral force and stress were significantly lower in the one-leg squat compared with the wall squat short between 90° and 70° knee angles during the squat ascent (Figs. 4 and 5). The primary cause of the greater patellofemoral force and stress between 90° and 70° knee angles in the wall squat short compared with the one-leg squat was greater quadriceps force during the wall squat short. Between 90° and 70° knee angles during the squat ascent, the estimated quadriceps forces that were calculated in the current study using our EMG-driven knee model were approximately 30-40% greater in the wall squat short compared with the one-leg squat. In contrast, between 90° and 70° knee angles during the squat ascent, the estimated hamstring forces that were calculated were approximately 60-70% greater in the one-leg squat compared with the wall squat short. 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 (Fig. 1A) and tilted forward 30-40° in the one-leg squat (Fig. 2). The erect trunk in the wall squat short produced a line of force from the center of the mass of the lifter-dumbbell system (lifter's mass plus dumbbell mass) that resulted in a relatively small hip moment arm and hip extensor muscle moment and a relatively large knee moment arm and knee extensor muscle moment (Fig. 1B). In contrast, the forward trunk tilt in the one-leg squat produced a line of force from the center of the mass of the lifter-dumbbell system that resulted in a relatively large hip moment arm and hip extensor muscle moment and a relatively small knee moment arm and knee extensor muscle moment (Fig. 2).
Although friction was minimized during the wall squat by using a smooth wall and a towel between the subject and the wall, the normal force that the wall applied to 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. This friction force acted opposite the force of gravity during the squat descent but acted 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 patellofemoral force and stress were greater in the ascent phase of the wall squat exercises compared with the one-leg squat.
If acceleration, whose magnitudes were very small, was discounted during the wall squat and a static analysis was performed at a point on the feet where the ground reaction force acted, gravity acting on the center of mass of the lifter-dumbbell system would produce a torque about the feet that must be countered by an equal and opposite torque generated by the normal force and static friction force that the wall exerts on the lifter's back (Fig. 1A and B). Because during the wall squat long the heels were twice as far from the wall compared with the wall squat short (Fig. 1A and B), the normal force must be greater in the wall squat long compared with the wall squat short. Because friction force is directly proportional to the normal force, the downward-acting friction force on the subject while sliding up the wall was greater in the wall squat long compared with the wall squat short, making it more difficult for a subject to slide up the wall during the wall squat long. However, because the normal force generates a knee extensor torque during thewall squat exercises, the greater normal force during the wall squat long compared with the wall squat short may have made it easier for the subject to slide up the wall during the wall squat long. Therefore, compared with the wall squat short, during the wall squat long, the greater friction force made it more difficult to slide up the wall but the greater normal force made it easier to slide up the wall. The varying and apposing actions of the normal force and friction force during the wall squat long and the wall squat short may help explain why the quadriceps force and resulting patellofemoral force and stress were generally similar between these two exercises, with the only exceptions at 60° and 90° knee angles during the squat ascent. It is unclear why patellofemoral force and stress were significantly greater in the wall squat long at 60° knee angle and significantly greater in the wall squat short at 90° knee angle.
The wall squat short and long as defined in the current study may represent two extremes in heel-to-wall distances that can be used while performing wall squat exercises. It is unlikely that the heel-to-wall distances used in the current study would ever be greater when performing the wall squat long or less when performing the wall squat short. It is alsopossible that patellofemoral force and stress may be different than the results of the current study if the wall squat were performed with heel-to-wall distances somewhere between those used for the wall squat short and long, and this should be the focus of additional studies. There may be an optimal heel-to-wall distance that minimizes patellofemoral force and stress.
Another consideration during patellofemoral rehabilitation is what knee flexion range of motion to use while performing squat exercises. Because patellofemoral force and stress generally increased with greater knee angles and decreased with smaller knee angles, a more functional knee flexion range between 0° and 50° may be more appropriate for patellofemoral patients compared with higher knee angles between 60° and 90°. For example, during the squat ascent phase of the wall squat short, patellofemoral force ranged from approximately 75 to 1400 N between 0° and 50° knee angle and from approximately 2100 to 3650 N between 60° and 90° knee angles, and patellofemoral stress ranged from approximately 0.5 to 4.4 MPa between 0° and 50° knee angles and from approximately 5.9 to 8.9 MPa between 60° and 90° knee angles. This same pattern of increased patellofemoral force and stress with larger knee angles has been reported during the barbell squat and leg press (14,15,29,32,36). These authors reported that patellofemoral force and stress progressively increased from 0° to approximately 90°, peaking at approximately 90°, and then progressively decreasing from approximately 90° to 0°. Computer optimization techniques demonstrated similar results during a simulated squat (7).
Peak patellofemoral force and stress magnitudes from the current study are less than some weight-bearing exercises, such as the barbell squat and the leg press (14), but more than some weight-bearing functional activities, such as walking (17) and going up and down the stairs (5). Escamilla et al. (14) reported peak patellofemoral force magnitudes between 4500 and 4700 N at 90° knee angle during the 12 RM barbell squat and leg press using healthy subjects, resulting in patellofemoral stress magnitudes between 11 and 12 MPa. Having healthy subjects squat with a barbell using a 35% bodyweight load, Wallace et al. (36) reported peak patellofemoral force magnitudes near 2500 N and patellofemoral stress magnitudes near 13 MPa, both occurring at 90° knee angle. Peak force and stress magnitudes during the barbell squat and leg press are approximately 25-50% greater compared with peak force and stress magnitudes in the current study, which also occurred at 90° knee angle. In contrast, peak patellofemoral force and stress in healthy subjects during fast walking reportedly are approximately 900 N and 3.13 MPa, respectively (17), which are approximately two to three times lower than the peak force and stress magnitudes in the current study. However, peak patellofemoral force and stress magnitudes in healthy subjects going up and down the stairs reportedly are approximately 2500 N and 7 MPa, respectively (17), which are only 15-30% less compared with the peak force and stress magnitudes in the current study. Understanding patellofemoral force and stress magnitudes among varying resistance exercises and functional activities is helpful to clinicians and trainers when deciding which interventions to use for patients with PFPS.
Unlike healthy subjects, patients with PFPS exhibit smaller patellar contact areas and greater patellofemoral stress during some weight-bearing functional activities (5,17). Compared with healthy individuals, patients with PFPS had 40% smaller patellar contact areas and 110% greater peak patellofemoral stress (6.61 MPa) during fast walking (17). Moreover, Hinterwimmer et al. (20) reported that patients with patellar subluxation had 40-55% smaller patellar contact areas compared with healthy individuals, which implies greater patellofemoral stress in these patients compared with healthy individuals. These patellofemoral stress data involving patients with PFPS implies that PFPS patients may be at higher risk of experiencing pain and discomfort while performing wall squat and one-leg squat exercises, especially at higher knee angles where patellofemoral force and stress are greatest, and this should be the focus of future studies.
Unfortunately, it is currently unknown what patellofemoral force or stress magnitudes and over what time duration can ultimately lead to patellofemoral pathology. There are many factors that may contribute to patellofemoral pathology, such as 1) overuse or trauma; 2) imbalance or malalignment of the extensor mechanism, which can lead to lateral patellar subluxation or tilt; 3) muscle weakness, such as weak quadriceps and hip external rotators; 4) muscle tightness, such as tight quadriceps, hamstrings, or iliotibial band; and 5) lower extremity malalignment, such as patella alta, genu valgum, femoral neck anteversion, excessive Q-angle, and excessive rearfoot pronation. It can only be surmised that relatively large patellofemoral force and stress magnitudes over time may lead to patellofemoral pathology, especially in individuals that exhibit some of the above factors and thus are predisposed to patellofemoral problems. Nevertheless, clinicians can use information regarding patellofemoral force and stress magnitudes among different weight-bearing exercises, technique variations, and functional activities to be able to make more informed decisions regarding which exercise they choose to use during patellofemoral rehabilitation.
Patellofemoral force and stress curves were similar in shape due to proportional increases in patellofemoral force and patellar contact area with increased knee angles. One exception was at higher knee angles between 70° and 90°, in which patellofemoral stress began to plateau or decrease to a greater extent than patellofemoral force. This occurred because although patellar contact area increased between 70° and 90°, patellofemoral force did not increase proportionally but instead began to plateau or decrease. These findings are consistent with patellofemoral force and stress data during the barbell squat from Escamilla et al. (14) and Salem and Powers (29). Escamilla et al. (14) reported that patellofemoral forces increases until 75-80° knee flexion and then began to plateau or slightly decrease. Salem and Powers (29) reported no significant differences in patellofemoral force or stress at 75°, 100°, and 110° knee flexion. It can be concluded from these squat data that injury risk to the patellofemoral joint may not increase with knee angles between 75° and 110° due to similar magnitudes in patellofemoral stress during these knee angles, with the benefit of increased quadriceps, hamstrings, and gastrocnemius activity when training at higher knee angles (75-110°) compared with training at lower knee angles (0-70°) (14).
There are limitations in the current study. Firstly, MRI knee kinematic data have shown during the weight-bearing squat that the femur moves and rotates underneath a relatively stationary patella, and if femoral rotation is excessive, it may result in an increase in patellofemoral contact area, force, and stress on the contralateral patellar facets (12,22,28). This implies that excessive medial femoral rotation during the squat ascent may place more stress on the lateral patellar facets, whereas excessive lateral femoral rotation during the squat descent may place more stress on the medial patellar facets. Unfortunately, collecting MRI knee kinematic data while performing the wall squat or the one-leg squat is not currently possible due to limitations in equipment design, so it is unknown how much femoral rotation occurs during the wall squat or the one-leg squat, how this rotation varies among healthy versus pathologic individuals while performing these exercises, and if femoral rotation occurs in the wall squat and the one-leg squat similarly to how it occurs in other weight-bearing exercises.
Another limitation is the effect of Q-angle on patellofemoral force and stress. From cadaveric data during a simulated squat, it was shown that an increased Q-angle significantly caused a lateral shift and medial tilt and rotation of the patella, which may increase patellofemoral force and stress (23). Unfortunately, it is currently difficult or impossible to effectively measure lateral shift and medial tilt and rotation of the patellar while performing squat type exercises. Moreover, increased medial femoral rotation, which also increases Q-angle, is also difficult to measure accurately while squatting.
There are also limitations in the biomechanical model. First, muscle and patellofemoral forces were estimated from modeling techniques and not measured directly, which is currently not possible in vivo. Second, patellofemoral stress magnitudes were measured using patellar contact area values from MRI data from the literature and were not measured directly. However, the contact areas used from the literature were determined during loaded weight-bearing exercise in healthy male and female subjects, similar to the current study. Moreover, the near linear and direct relationship between contact area and knee angle has been shown to be similar among studies (2,8,20,27,30). This implies that the patellofemoral stress curve patterns shown in Figure 5 using contact areas from the literature will be similar to patellofemoral stress curve patterns if contact areas were measured directly using MRI. The patellofemoral stress patterns are important to clinicians in determining what knee range of motions stress increases or decreases.
There are limitations regarding the magnitude of patellofemoral contact areas (and concomitant stress magnitudes), in which the literature reports a wide array. For example, both Patel et al. (27) and Besier et al. (2), who also used loaded weight-bearing exercise, reported approximately 40-50% higher patellofemoral contact areas compared with contact areas data from Salsich et al. (30). Using these larger contact areas from Patel et al. (27) and Besier et al. (2) would have resulted in smaller patellofemoral stress magnitudes than those reported in the current study. Differences in patellar contact area magnitudes and concomitant patellofemoral stress magnitudes among weight-bearing studies are due to many factors, such as sex (males generally have greater contact areas than females), mass (larger individuals generally have greater contact areas than smaller individuals), measuring techniques, and loading magnitudes (greater loading and quadriceps contraction results in increased contact areas and less patellofemoral stress (2,30). Nevertheless, although patellofemoral stress magnitudes during weight-bearing exercises or functional activities are approximations only and not exact values, understanding how magnitudes vary among exercises may be helpful to clinicians in deciding interventions to use during patellofemoral rehabilitation.
Finally, the current study was limited to healthy subjects who were able to perform the wall squat and the one-leg squat in the sagittal plane of motion without transverse and frontal plane motions. Future studies are needed during the wall squat and the one-leg squat to investigate the effects of transverse plane rotary motions and frontal plane valgus/varus motions on patellofemoral force and stress magnitudes, which may occur with individuals with patellofemoral pathology. Moreover, the results of this study are specific to performing a one-leg squat holding a single dumbbell at the chest with a forward trunk tilt of approximately 30-40° and performing the wall squat exercises with a dumbbell in each hand with the arms at the side. Using different technique variations during the one-leg squat and the wall squat exercises may affect patellofemoral force and stress magnitudes, and this should also be the focus of future studies.
The efforts of Dr. Bonnie Raingruber and the 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.
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Keywords:©2009The American College of Sports Medicine
BIOMECHANICS; KINETICS; CLOSED CHAIN EXERCISES; KNEE