Forces and Moments on the 

Knee During Kneeling and Squatting
�

Jonisha P. Pollard, William L. Porter,

Jonisha P. Pollard (Corresponding Author) and William L. 

Porter are with the National Institute for Occupational Safety 

and Health, Office of Mine Safety and Health Research, Pittsburgh, PA.

 and Mark S. Redfern 

 Mark S. Redfern is with the Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA. 

Euler angle decomposition and inverse dynamics were used to determine the knee angles and net forces and 

moments applied to the tibia during kneeling and squatting with and without kneepads for 10 subjects in four 

postures: squatting (Squat), kneeling on the right knee (One Knee), bilateral kneeling near full flexion (Near 

Full) and bilateral kneeling near 90° flexion (Near 90). Kneepads affected the knee flexion (p = .002), medial 

forces (p = .035), and internal rotation moments (p = .006). Squat created loading conditions that had higher 

varus (p < .001) and resultant moments (p = .027) than kneeling. One Knee resulted in the highest force 

magnitudes and net moments (p < .001) of the kneeling postures. Thigh-calf and heel-gluteus contact forces 

decreased the flexion moment on average by 48% during Squat and Near Full. 

Keywords: restricted workspace, high flexion, thigh-calf contact, heel-gluteus contact 

Low-seam coal miners are required to work in low 

seam heights (1.1 m [42 inches] or less), which restrict 

their postures, forcing them to kneel, crawl, and squat 

to perform work. Low-seam coalmines report 10 times 

as many lost-time injuries related to these postures as 

high-seam mines, with 75% of these injuries being to 

the knee (Peters et al., 2001). Knee injuries are also the 

most prevalent, accounting for 17% of total injury costs 

to mining companies—more than any other body part, 

including the lower back (Gallagher et al., 2009). Many 

low-seam coal miners wear kneepads to protect their 

knees from the uneven mine floor, which may contain 

sharp, jagged rocks. Although kneepads may protect the 

knees from cuts and scrapes, their effect on the internal 

knee structures is unknown. Castagno (2004) found no 

indications that kneepads affect long-term gait patterns. 

However, their effect on knee biomechanics is not known, 

and it is likely that kneepads affect kneeling posture and 

thereby knee loading. 

Other occupations, such as carpet and floor layers, 

also require prolonged kneeling and report increased 

knee injuries (Tanaka et al., 1982). Thun et al. (1987) 

concluded that “workers who kneel to perform their jobs 

inflict chronic trauma to their knee joints.” Carpet layers 

are at an increased risk for a variety of knee disorders, 

including cartilage damage due to frequent and prolonged 

kneeling and squatting in addition to the use of a knee 

kicker to stretch carpet (NIOSH, 1990; Habes et al., 1994; 

Thun et al., 1987; Rytter et al., 2008; Holmstrom et al., 

1995). The risk for development of knee osteoarthritis is 

also increased in occupations requiring kneeling, squatting, climbing, and heavy physical workloads (Manninen 

et al., 2002; Schneider, 2001). 

Previous studies have shown increased knee forces 

when kneeling in high flexion and squatting. Due to variability in computation methods, knee forces reported in 

literature exhibit wide variability and have been reported 

to be as high as 7 times body weight in deep flexion 

(Nagura et al., 2002; Dahlkvist et al., 1982; Perry et al., 

1975; Escamilla et al., 1998; Komistek et al., 2005). 

However, many of these studies have not examined fullrange flexion, and all have neglected the contact between 

the thigh and shank (thigh-calf contact) that occurs in 

high flexion. 

Thigh-calf contact may have a significant effect on 

knee forces and moments. Caruntu et al. (2003) modeled 

the behavior of the knee in a deep squat (165° flexion) 

to determine the effect of thigh-calf contact. The authors 

found that neglecting thigh-calf contact resulted in a 700 

N overestimation in quadriceps force and a 50% underestimation in medial collateral ligament forces. However, 

the authors never reported the magnitude or location of 

the thigh-calf contact force used in their model. Zelle et 

al. (2007) measured thigh-calf contact pressure distributions using a pressure mapping system. Subjects were 

instructed to perform two activities. In the first activity, 

subjects descended from erect standing to a squatting 

position. In the second, subjects descended from erect 

kneeling to full flexion with their ankles dorsal flexed. 
 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

Figure 1 — Kneepads used in study. Outer shell, inner padding, and side views of articulated kneepad (A) and nonarticulated 

kneepad (B). 

Mean knee angles for the final squatting and kneeling 

postures were 151.8° and 156.4° respectively. Results 

showed exponentially increasing contact pressures with 

increasing knee flexion and maximal forces at terminal 

flexion. Resultant thigh-calf contact forces were shown 

to be >30% body weight and located within 17 cm of the 

epicondylar axis with contact occurring more proximal 

to the tibia when squatting compared with kneeling. In a 

later investigation, Zelle et al. (2009) suggested that thighcalf contact considerably reduced the joint forces during 

high-flexion squatting. Results showed compressive force 

to decrease from 4.89 to 2.90 times body weight when 

accounting for thigh-calf contact. In addition, authors 

found maximum forces to occur at the initiation angle 

of thigh-calf contact instead of at maximal flexion. This 

implies that knee forces may be greater in lower flexion 

kneeling postures than in full-range kneeling and squatting due to thigh-calf contact. 

The purpose of this study was to quantify the net 

forces and moments applied to the tibia and to determine 

the knee angles associated with kneeling and squatting. 

In addition, the impact of different kneeling and squatting postures and kneepad use on the tibial forces and 

moments were examined. Thigh-calf and heel-gluteus 

contact forces were also measured and incorporated into 

the biomechanical model. This information may be useful 

in developing guidelines for mining and other occupations 

requiring frequent and prolonged kneeling. 

Methods 

Subjects 

Ten subjects (7 male, 3 female) recruited from the 

National Institute for Occupational Safety and Health 

(NIOSH) in Pittsburgh, PA, participated in this study. 

The average ± SD of age, weight, and height were 34 

± 17 years, 683 ± 98 N, and 169 ± 8 cm, respectively. 

Subjects were healthy, with no history of knee surgery. 

Before participating, each subject read and signed an 

informed consent form approved by the NIOSH Human 

Subjects Review Board. 

Experimental Design 

A split-plot design was used to evaluate the effect of 

posture and kneepad on the angles, forces, and moments 

at the right knee joint. Subjects performed four postures: 1—Squatting (Squat), 2—bilateral kneeling near 

90° flexion (Near 90), 3—bilateral kneeling near full 

flexion (Near Full), and 4—kneeling on the right knee 

(One Knee) during three kneepad conditions: (1) no 

kneepad, (2) articulated, and (3) nonarticulated (Figure 

1). Testing was blocked by kneepad condition, which was 

randomized, and, within these blocks, the testing order 

of postures was also randomized. One trial per kneepad 

condition was collected for each posture. 
 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

Preliminary Measurements 

Preliminary measurements were taken before subject 

testing, including subject height, weight, and thigh-calf 

and heel-gluteus contact forces. Pilot testing showed 

standard repeatability error within 5% body weight for 

repeated measures of contact forces with and without 

kneepads. Therefore, contact force measurements were 

suitable for collection before the start of testing. Pressure 

profiles of the forces applied normal to the tibia and heel 

were obtained using the ClinSeat pressure sensing system 

(Tekscan, South Boston, MA). The sensor pad was placed 

on the subject’s lower leg (from popliteal cavity to heel), 

and the subject was instructed to squat and kneel near full 

flexion. The distance from the top of the uppermost sensing unit (placed in the popliteal) to the lateral epicondyle 

of the femur was then measured and recorded. Pressure 

data were then collected for 5 s at 4 Hz. Using the supplied 

software (Advanced ClinSeat, Tekscan, South Boston, 

MA), the thigh-calf and heel-gluteus contact areas were 

selected and the average total force and center of pressure 

(COP) for a 3 s period were determined. The moment 

arms (distances from knee joint center to COPs along the 

long axis of tibia) of these contact forces were calculated 

by adding the distance from the COP to the uppermost 

sensing unit to the distance from the uppermost sensing 

unit to the lateral epicondyle of the femur. Thus, these 

contact forces only contributed to anterior forces and 

sagittal moments. 

Procedure 

To ensure kneeling, work area height was restricted to 

1.2 m (48 inches). Four postures were selected based 

upon those observed to be commonly adopted in lowseam mining: Squat, Near 90, Near Full, and One Knee 

(Figure 2). Subjects performed each posture for three 

kneepad conditions: no kneepads, articulated kneepads, 

and nonarticulated kneepads. The kneepads tested were 

those most frequently purchased from mining distributors 

in 2007. The articulated kneepads consisted of a hard 

outer shell with firm inner padding. The nonarticulated 

kneepads consisted of a soft outer rubber shell and soft 

inner foam padding. 

Figure 2 — Schematic showing each posture tested during 

the experiment. All postures were simulated in a 1.2 m (48 

inch) working height. 

A motion capture system (Eagle; Motion Analysis 

Corporation, Santa Rosa, CA) was used to determine the 

location of the body segments as the subject assumed the 

four postures at a sampling rate of 60 Hz. With the use 

of kneepads, it was not possible to place markers on the 

medial and lateral epicondyles of the knee. Instead, the 

locations of these markers were determined from a calibration trial. Two marker sets were used: an anatomical 

marker set and a measured marker set (Figure 3), which 

were modified versions of the Cleveland Clinic marker set 

 

 

Figure 3 — Anatomical marker set (A) and measured marker 

set (B) in which eight markers were removed (R.Knee.Medial, 

R.Knee.Lateral, R.Ankle.Medial, R.ASIS (anterior superior 

iliac spine), L.ASIS, L.Knee.Medial, L.Knee.Lateral, and 

L.Ankle.Medial) (R = Right, L = Left). 
 

 

 

 

 

 

 

 

 

 

 

(Kirtley, 2006). The anatomical marker set was composed 

of the measured marker set with eight additional anatomical markers. Calibration data of the combined marker sets 

were collected to acquire the anatomical marker locations 

relative to the measured markers. During actual testing, 

the anatomical markers were removed. The estimated 

locations of the anatomic landmarks were accurate to 

within an average of 9% for all postures. 

Two force plates (Advanced Mechanical Technology, 

Inc., Watertown, MA) were used to capture the ground 

reaction forces on the right leg: one plate under the knee 

and one under the foot. Force plate data were collected 

at 1020 Hz using the software provided by the Motion 

Analysis Corporation (EvaRT 5.0, Motion Analysis Corporation, Santa Rosa, CA) through an analog-to-digital 

board (PCI-6071E, National Instruments, Austin, TX). 

Before each trial, the subject was shown a diagram 

(similar to Figure 2) and asked to assume the posture. 

Minimal verbal feedback was given unless the subject 

was assuming a posture with knee angles grossly different 

from what was expected. Motion capture and force plate 

data were collected for 10 s. Each trial was performed 

once and the subject was provided seated rests between 

trials. Motion data were low-pass filtered using a 4thorder Butterworth to 15 Hz. 

Data Analysis 

Anatomical and measured coordinate systems were created for the thigh and shank from the calibration trial 

using the methods from Pollard (2008). The anatomical 

system allowed the location of the ankle joint center 

(AJC), knee joint center (KJC), hip joint center (HJC), 

and lower leg center of mass to be determined. The 

measured coordinate system was used to estimate the 

location of the anatomical markers during testing. The 

AJC was assumed to be midway between the medial 

and lateral malleoli of the tibia (measured by the medial 

and lateral ankle markers), the KJC was assumed to be 

midway between the medial and lateral epicondyles of the 

femur (measured by the medial and lateral knee markers) 

and the HJC was approximated using regression equations proposed by Bell et al. (1990). 

The anatomical coordinate system of the shank 

was constructed such that the Z-axis was oriented in the 

superior direction of the right tibia (AJC to KJC vector), 

the Y-axis was in the anterior direction (from the cross 

product of the AJC to KJC vector and medial epicondyle 

to lateral epicondyle vector), and the X-axis was in the 

lateral direction (from the cross product of the Y and Z 

axis vectors). Similarly, the anatomical coordinate system 

of the thigh was constructed. The rotation matrix from the 

anatomical thigh to the anatomical shank coordinates was 

created to determine the Euler angles. Knee angles were 

estimated for each posture using Euler angle decomposition, with the decomposition order being first about the 

X-axis (included flexion angle), followed by the Y-axis 

(varus/valgus), and the Z-axis (internal/external rotation). 

Forces and moments at the knee were calculated 

using ground reaction forces, lower leg weight, thighcalf contact force (Ft/c), and the heel-gluteus contact 

force (Fh/g). The center of mass location and weight of 

the shank (Wshank) were determined using equations from 

Clauser et al. (1969) adjusted to use the KJC and AJC in 

this model (Hinrichs, 1990). Ground reaction forces at 

the foot (F1) and knee (F2) were also included. External 

force diagrams for all postures are shown in Figure 4. 

Figure 4 — External force diagrams with respect to the anatomical shank coordinate system (ASCS). The ASCS is oriented such 

that the X-axis is in the lateral direction, the Y-axis is in the anterior direction, and the Z-axis is in the superior direction of the right 

tibia. Note that squatting includes the thigh-calf contact force, as does kneeling near full flexion in addition to the heel-gluteus 

contact force. 
 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

Forces and moments applied to the tibia were calculated 

in three dimensions with respect to the anatomical shank 

coordinate system (ASCS). All resulting forces were 

normalized to body weight (BW) and moments were 

normalized to the product of body weight and height 

(BW*Ht) (Moisio et al., 2003). 

Statistical Analyses 

Statistical analyses were performed using Statistix 8.0 

for Windows. A two-way (3 kneepads × 4 postures) 

split-plot ANOVA was performed to determine whether 

significant differences existed in knee angles, forces, and 

moments among the postures and kneepad conditions. A 

priori orthogonal contrasts included comparisons of the 

no-kneepad to kneepad-present (articulated and nonarticulated) conditions, high-flexion (Squat and Near Full) 

and lower-flexion (Near 90 and One Knee) postures, and 

Squat versus kneeling (Near 90, One Knee and Near Full). 

All orthogonal contrasts were tested using a t statistic. 

An alpha value of 0.05 was used. A priori nonorthogonal 

contrasts were tested to determine whether significant 

differences existed between the kneepad conditions using 

the Scheffé F test (p < .05). 

Results 

Thigh-calf and heel-gluteus contact forces and moments 

are shown in Table 1. All subjects had thigh-calf contact 

forces greater than 20%BW for Squat with a mean of 

39 ± 14%BW. In the Near Full posture, seven subjects 

had contact between the heel and gluteal muscles, with 

a mean of 11 ± 6%BW. Pressure distribution profiles for 

Squat and Near Full are shown in Figure 5. 

Significant differences were seen between postures 

for the included (p < .001) and internal rotation angles (p 

< .001) (Figure 6). When compared with the kneeling postures, Squat showed higher flexion and internal rotation (p 

< .001). A priori orthogonal contrasts showed significant 

differences in knee flexion between the high-flexion and 

lower-flexion postures (p < .001) and between Near 90 

and One Knee (p < .001). Differences in internal rotation 

angles were seen between high-flexion and lower-flexion 

postures (p < .001) and between Squat and Near Full (p 

< .001). The no-kneepad condition showed higher knee 

flexion (included angles = 35 ± 24°) than the nonarticulated kneepad condition (included angles = 41 ± 26°) (p 

= .033). No significant differences were seen between 

the articulated (40 ± 26°) and nonarticulated kneepads. 

Posture had a significant effect on tibial forces 

(Figure 7). When compared with the kneeling postures, 

Squat showed significantly higher superior and lower 

posterior and resultant forces (p < .001). Lower-flexion 

postures showed greater posterior, inferior, and resultant 

forces than the high-flexion postures (p < .001). Squatting 

showed significantly higher medial (p = .002), superior 

(p < .001), and resultant forces (p < .001) and lower 

posterior forces than Near Full (p < .001). One Knee 

showed significantly higher medial, posterior, inferior, 

and resultant force magnitudes than Near 90 (p < .001). 

The use of kneepads affected the medial forces (p = .036) 

with the no-kneepad condition (–6.1 ± 8.1%BW) having 

Table 1 Thigh-calf and heel-gluteus contact forces, as a function of 

percent body weight (%BW), and sagittal moment, as a function of 

body weight times height (%BW*Ht) contributions for the Squat and 

Near Full postures 

Subject Squat Near Full 

Thigh-Calf Contact Thigh-Calf Contact Heel-Gluteus Contact 

Force Moment Force Moment Force Moment 

1 45 4.9 31 2.7 13 3.4 

2 24 2.8 19 1.5 8 2.0 

3 20 1.8 32 3.0 3 0.8 

4 62 8.6 53 6.7 No contact 

5 57 5.1 8 0.5 16 3.7 

6 50 5.1 39 3.9 5 1.3 

7 36 3.2 26 2.5 10 2.4 

8 33 3.2 13 1.2 20 4.8 

9 27 2.4 23 2.4 No contact 

10 33 2.3 37 2.6 No contact 

Mean (SD) 39 (14) 3.9 (2) 28 (13) 2.7 (1.7) 11 (6) 2.6 (1.4) 
 

Figure 5 — Typical pressure distribution (psi) for Squat (A) and Near Full (B). 

Figure 6 — Knee angles (degrees) for all postures. Varus and internal rotation angles are positive. 

 

 

 

 

 

 

 

 

 

 

 

higher medial forces than the articulated kneepad condition (–4.3 ± 7.6%BW). 

Posture also had a significant affect on tibial moments 

(Figure 8). Flexion moments had the greatest magnitudes 

for all postures. Of the postures examined, One Knee had 

the greatest flexion moments, Squat had the greatest varus 

moments and Near 90 had the greatest internal rotation 

moments. One Knee and Squat had the highest resultant 

moments with comparable magnitudes of 5 ± 4%BW*Ht 

and 5 ± 2%BW*Ht, respectively. When compared with 

the kneeling postures, Squat showed significantly lower 

internal rotation (p = .027) and increased varus (p < .001) 

and resultant moments (p = .027). High-flexion postures 

showed significantly higher varus (p < .001), internal 

rotation (p = .009), and resultant moments compared 

with the lower-flexion postures (p = .007). Squat showed 

significantly higher varus moments than Near Full (p < 

.001). One Knee showed significantly higher flexion (p 

< .001), varus (p = .012), and resultant moments and 

significantly lower (p < .001) internal rotation moments 

than Near 90 (p = .01). Kneepad use affected the internal 

rotation moments (p < .001), with the no-kneepad condition (0.23 ± 0.32%BW*Ht) having significantly smaller 

internal rotation moments than the nonarticulated (0.36 

± 0.36%BW*Ht) (p = .049) and articulated kneepad 

conditions (0.36 ± 0.35%BW*Ht) (p = .008). 
 

Figure 7 — Mean tibial forces normalized to body weight 

(BW). Lateral, superior, and anterior are positive. 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

Figure 8 — Mean tibial moments normalized to body weight 

times height (BW*Ht). Extension, varus, and internal rotation 

moments are positive. 

Discussion 

In this study, knee angles and net forces and moments 

applied to the right tibia were determined while in kneeling and squatting postures with and without kneepads. 

Significant differences were found between squatting and 

kneeling postures as well as between the high-flexion 

(Squat and Near Full) and lower-flexion (Near 90 and 

One Knee) postures. The results present several important 

findings: (1) tibia loading when squatting is significantly 

different than when kneeling; (2) the highest shear loading across the knee joint (i.e., medial forces) occurred 

when kneeling on one knee; (3) heel-gluteus contact force 

significantly affects tibial forces and moments; and (4) 

the use of kneepads affect joint loading when kneeling. 

In all the kneeling postures examined, the forces 

acting on the tibia occurred in the medial, posterior, 

and inferior directions. Squatting created a different 

loading pattern with mean forces acting in the medial, 

posterior, and superior directions of the tibia. When 

squatting, subjects were found to exhibit significant 

internal tibial rotation, which may be the primary cause 

of the increased varus moments. These varus moments 

increase the loading to the medial compartment of the 

knee. In high flexion, tibiofemoral contact occurs at the 

posterior aspects of the femur and tibia with greater 

loading on the medial compartment (Hefzy et al., 1998; 

Thambyah et al., 2005). Although the articular cartilage 

and meniscus are designed for high compressive loads 

in the central portion, the effects of high loading on the 

posterior portions of the medial compartments remains 

unclear. An increased risk for medial tibiofemoral narrowing has been associated with varus alignment (Sharma 

et al., 2001). The increased varus moments (increasing 

loading on the medial compartment of the tibia), high 

knee flexion (decreasing tibiofemoral contact area), and 

high posterior and inferior forces found in this study 

may explain the greater effect of squatting on medial 

compartment osteoarthritis and the significant relationship between varus moments and osteoarthritis severity 

(Zhang et al., 2004; Sharma et al., 1998; Cooper et al., 

1994; Anderson & Felson, 1988). These previous findings, in conjunction with the current study’s results, imply 

that deep squatting is likely to be the most detrimental 

high-flexion postureOf all postures examined, One Knee 

had the highest resultant force magnitudes. It was noted 
 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

 

that when kneeling on the right knee, a predominant 

portion of the subjects’ weight was transmitted through 

their right knee and foot. Therefore, it is believed that in 

the One Knee posture, the knee making contact with the 

floor is likely to be primarily weight bearing and the other 

leg may act solely to maintain balance. In effect, high 

posterior forces, as much as 60%BW are transmitted to 

the tibia in addition to flexion moments up to 5%BW*Ht. 

During flexion, the weight-bearing surfaces of the knee 

move posterior to the tibial plateaus and decrease in area 

which may increase the stresses transmitted to the tibia. 

(Maquet et al., 1975; Thambyah et al., 2005) This may 

create a detrimental loading condition for the articular 

cartilage, which is thought to be damaged by maximum 

shear stress (Atkinson et al., 1998). High shear forces 

combined with high flexion moments are believed to 

make kneeling on one knee the most detrimental of the 

kneeling postures examined. 

Previous research has shown thigh-calf contact to 

have a significant effect on model predictions of joint 

forces and moments (Caruntu et al., 2003; Zelle et al., 

2009). To the authors’ knowledge, Zelle et al. (2007) was 

the first study to quantify thigh-calf contact forces. The 

current study expanded on this knowledge, quantifying 

the additional heel-gluteus contact force found when 

kneeling near full flexion.Although this force had smaller 

magnitudes than thigh-calf contact, its effect on the flexion moment was comparable. The larger moment arm of 

the heel-gluteus contact force acted to create an extension 

moment that counteracted the flexion moment created by 

the ground reaction forces at the knee and foot. The results 

of the current study imply that heel-gluteus contact may 

also reduce loading of knee tissues as has been shown 

for thigh-calf contact (Zelle et al., 2009). Although knee 

kinematics were determined in three dimensions, due 

to the limitations of the pressure sensor, thigh-calf and 

heel-gluteus contact forces were obtained in 2-D. These 

contact forces were measured normal to the tibia and 

thus the medial-lateral and superior-inferior shear forces 

caused by tissue deformation were neglected, thereby 

affecting the accuracy of the estimated shear forces and 

internal rotation and varus moments. 

High-flexion kneeling and squatting create a significant change in internal joint structure orientations, 

such as femoral rollback and varying segment contact 

forces. During increased knee flexion, the tibia internally 

rotates changing the orientation of the knee joint center 

to a position that may only be determined using imaging techniques not employed in this study. In effect, this 

shift in knee joint center location may have caused inaccuracies in moment calculations, as the net moment was 

assumed to act at a point in the center of the epicondylar 

axis. However, the 3-D model used in this study may 

minimize this error by accounting for tibial rotations. 

The current study is the first to imply that kneeling 

with kneepads affects posture and joint loading. While 

the smooth, even laboratory floor did not create a kneeling environment comparable to that of a coalmine, it is 

believed that the general findings within this study hold 

for other environments. However, the increase in transverse moments associated with the hard contoured outer 

shell kneepad may create deleterious effects that may be 

reduced when wearing kneepads with a softer outer shell. 

The thigh-calf contact forces measured in this 

study (39%BW for Squat and 28%BW for Near Full) 

are comparable to the values found in literature. Zelle 

et al. (2007) reported average thigh-calf contact forces 

of 34.2%BW and 30.9%BW for squatting and kneeling 

near full flexion, respectively. However, subjects knelt 

with ankles dorsally flexed, which was not used by all 

subjects in the current study and may explain the presence of heel-gluteus contact. Foot posture has been shown 

to significantly affect varus knee moments during gait 

(Teichtahl et al., 2006; Lynn et al., 2008), and it is possible that the differences in foot posture across subjects 

may have increased the internal rotation of the tibia and 

thereby the varus moments. 

Although flexion angles may be of similar magnitude 

in different postures when kneeling in restricted workspaces, the biomechanical stresses on the knee differ 

significantly. Squatting and kneeling near full flexion 

have similar joint flexion requirements; however, squatting was found to have much great forces and moments 

than kneeling near full flexion. Thus, posture of the entire 

body, as well as simple knee flexion angle, are important 

determinants of knee loading characteristics. 

Reducing the loads and moments on the knee are 

assumed to provide a reduction in risk of injury. The 

results of this study lead to three key recommendations 

toward reducing forces and moments on the joint during 

kneeing. (1) Equipment that increase thigh-calf or heelgluteus contact forces will lead to decreased loading at 

the knee. (2) When kneeling in high flexion, internal 

tibial rotations should be minimized to reduce loading 

on the medial meniscus and articular cartilage, thereby 

reducing the risk for medial compartment osteoarthritis. 

This could be accomplished by limiting ankle rotation. 

(3) Postural rotation strategies (i.e., changing postures 

throughout the day) can reduce any cumulative effect 

on tissues. Kneeling on one knee and squatting have 

particularly high forces and moments, and so should not 

be sustained for long periods of time. 

Acknowledgments and Disclaimer 

The authors would like to acknowledge the contributions of the 

following individuals in the conduct of this study:Al Cook, Sean 

Gallagher,Alan Mayton, Susan Moore, Mary Ellen Nelson, and 

NIOSH Pittsburgh Researchers. 

The findings and conclusions in this paper are those of 

the authors and do not necessarily represent the views of the 

National Institute for Occupational Safety and Health. 

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