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Effects of back support on intra-abdominal pressure and lumbar kinetics during heavy lifting.


Spine kinetic changes associated with the trunk anatomy have been linked with humans lifting near-maximal levels (Garg and Herrin, 1979; Morris, Lucas, and Bresler, 1961; Troup, Leskinen, Stalhammar and Kurinka, 1983). Obvious concerns center on the biomechanical vulnerability of the spine during lifting with the potential for herniation of the nucleus pulposus and annulus fibrosus complex. The nature of this problem has evolved in industrial populations with a need to reduce the occurrence of back injury. The impact of back injury and pain within the area of industrial medicine has also been demonstrated as significant (Kelsey, White, Pastides, and Bisbee, 1979).

Various preventive strategies inclusive of employee screening, redesigning the workplace, and lifting education programs have been employed to alter and reform lifting methods (Becker, 1961; Troup, 1977). More recently, the use of lumbosacral elastic supports has been incorporated, particularly in industry, as a means to reduce the number of missed days of work by minimizing both the occurrence and severity of back trauma (Walsh and Schwartz, 1990). The beneficial effects associated with lumbar spinal bracing have included containment of the abdominal viscera, restricted vertebral motion, and reduction of back pain (Norton and Brown, 1957; Perry, 1970).

Generally, lumbosacral supports contain the abdominal viscera and stabilize the trunk by limiting motion. Morris et al. (1961), who studied the effects of lumbosacral corsets on intra-abdominal pressure (IAP) and electromyographic (EMG) activity of the abdominal musculature, concluded that when large forces are imposed around the trunk, IAP and EMG tend to increase. When the lumbosacral area is contained by a corset, IAP tends to increase, with a significant decrease in EMG activity of the rectus abdominis and abdominal obliques. These effects were considered advantageous in providing a relieving force for the lumbosacral spine (Bartelink, 1957; Davis and Troup, 1964; Morris et al., 1961).

Harman, Rosenstein, Frykman, and Nigro (1989) conjectured that increasing abdominal surface area contact with a lumbosacral device would alter abdominal protrusion, thereby changing the effects of the abdominal musculature on lumbar spine kinetics. Their results support an ergogenic (work-enhancing) quality for lumbosacral supports; increases in IAP biomechanically reduce the compressive demands on vertebral bodies. Other investigations (Harman et al., 1989; Lander, Hundley, and Simonton, 1992; Lander, Simonton, and Giacobbe, 1990) have revealed similar results on the effects of weight-training belts during a squat-style, near-maximal lift. All conclude that weight-training belts increase IAP and aid in reducing the compressive forces of the spine during lifting.

Unfortunately, little investigation has been attempted to determine the biomechanical effects of elastic lumbosacral supports during lifting (Amendola, 1989; Jones, McEnvoy, Mills, Nash, and Perkins, 1985; Wu, 1985). Most of the supporting evidence for this type of lumbosacral support in industry remains anecdotal. Some investigations have isokinetically evaluated the effects of various types of rigid and semirigid lumbosacral supports with inconclusive results (Woodhouse, Heinen, Shall, and Bragg, 1990a, 1990b, 1993). These studies measured only external forces generated while wearing lumbosacral supports and did not report internal forces or IAP during lifting tasks.

Given the insidious nature of back injury, inconclusive results from past research, and the prevailing use of lumbosacral supports in industry, the present study was undertaken to ascertain some of the kinetic/kinematic properties associated with various types of low back support during near-maximal lifting. The purpose of the present investigation was to determine the effects on spine kinetics of wearing a weight training belt alone, wearing one with a rigid abdominal pad, or wearing a supportive elastic binder.

The rationale for providing a rigid abdominal pad was to create a surface opposing the abdominal wall (Woodhouse et al., 1990a, 1990b, 1993). Suppositions by Harman et al. (1989) and Lander et al. (1990), along with similar observations by the present authors, significantly influenced the formulation of one research question. Specifically, what effects on spine function would a rigid abdominal pad have during a near-maximal squat-style lift?

Intra-abdominal pressures were recorded and their corresponding relieving forces calculated during lifting to further determine the effects lumbosacral supports may have during a near-maximal squat-style lift.



This quasi-experimental study was performed on nine men (mean age 24.7 years; height 1.80 m; weight 824 N). The subjects were healthy and had no history of lower back pathology. A physical examination, which included a roentgenogram of the lower back, was performed prior to testing. Subjects provided informed voluntary consent and were paid for their participation.

Data Collection

Lifting task and support conditions. Subjects lifted a weighted box with handles (36.5 x 36.5 x 34.0 cm) from a squat to a standing position. They were instructed to lift with head erect, back straight, and arms extended in a squat-style lift with minimal horizontal movement and without holding their breath. The box weight equaled 90% of each subject's one-repetition maximum (mean = 545.3 N, SD = 56.0 N), which was determined 48 h prior to testing. Subjects lifted the box four times, once in each of the following four conditions: no support, weight belt, weight belt with a rigid abdominal pad [ILLUSTRATION FOR FIGURE 1 OMITTED], and an elastic abdominal binder [ILLUSTRATION FOR FIGURE 2 OMITTED]. The order of the conditions was counterbalanced, and a minimum of 5 min was allowed between lifts.

Videography. A Panasonic PV-330 VHS video camera was placed 8 m orthogonal to the subject's sagittal plane. The camera recorded at a rate of 60 Hz with a shutter speed of 1/500 s. A 300-W halogen light provided additional illumination. A meter stick was recorded prior to testing and was used to convert video units to meters. A reference point and trial numbers were placed in the field of view. The following sites were marked with white adhesive dots 2 cm in diameter to identify the axes of rotation of selected joints: head of the fifth metatarsal, lateral malleolus, lateral epicondyle of the femur, superior border of the greater trochanter, and head of the humerus.

Force platform. Subjects placed both feet on a force platform (Kistler 9281 B11) during the lifting motion. Signals from the force platform were amplified (Kistler type 9865) and recorded on a microcomputer at a sampling rate of 180 Hz using a 16-bit A/D board and Bio Ware Software (Kistler, Version 1.0).

Intra-abdominal pressure (IAP). IAP was measured by an MMS-200 system (Narco Bio-Systems) to a second microcomputer using data acquisition software (Easyest, Asyst, Inc.) and a 16-bit A/D board at a sampling rate of 1000 Hz. The IAP transducer was placed below the cardiac sphincter via the nasal cavity as per catheter length, body dimensions, and anatomical landmarks. The intra-abdominal relieving force was calculated by multiplying the recorded IAP by the estimated cross-sectional area of the diaphragm (Bartelink, 1957).

Synchronization. A light-emitting diode was placed in front of the camera lens and used to synchronize the kinematic data to the force and pressure data. When the light-emitting diode was activated, a 5-V signal was simultaneously recorded by the force platform and pressure-measuring computers.

Data Analysis

Videography. Joint markers and box handles were manually digitized from a minimum of 10 pictures before the box left the ground until the subject reached a standing position (Peak Performance Technologies, Englewood, CO). Because of scapular abduction during the lifting motion, the marker placed on the head of the humerus did not represent the true superior end of the trunk segment. Consequently, an adjacent point in line with and posterior to the shoulder marker, located along the trunk, was digitized to represent the trunk segment. The raw x,y coordinate data were smoothed using a Butterworth Digital Filter at a cutoff frequency of 8 Hz. The linear velocity of the box; the relative joint angles at the ankle, knee, and hip; and the absolute angles of the shank, thigh, and trunk were calculated for the entire lifting motion. Joint angles were evaluated at the time of peak compressive L5-S1 force to examine differences in lifting technique.

Joint forces and moments. With near-maximal lifts, the forces attributable to inertial quantities are minimal compared with the effect of the ground reaction forces (Lander et al., 1990, 1992). Thus a quasi-static model of the lifter [ILLUSTRATION FOR FIGURE 3, 5, AND 6 OMITTED] was used to calculate joint forces and moments of the lower extremity (Lander et al., 1990, 1992). Relieving abdominal force was determined by multiplying the estimated diaphragmatic surface area (0.0465 [m.sup.2]; Morris et al., 1961) by IAP. In addition, the IAP relieving force arm was dependent on trunk and thigh position (Freivalds, Chaffin, Garg, and Lee, 1984), and the back muscle force arm was estimated at 6.0 cm (Troup and Chapman, 1969). Finally, the compression and shear force values were adjusted by 40 deg to account for L5-S1 disc orientation (Freivalds et al., 1984). Position data from the videotape analysis, force plate data, and anthropometric estimates by Dempster (1955) were inputs to these calculations.

Statistics. A multivariate analysis of covariance (MANCOVA) with repeated measures using system weight as the covariate was performed to examine differences (p [less than] 0.05) between conditions for the dependent variables. These included the peak compression force, shear force, extension moment, and peak extension muscle force around L5-S1, as well as the peak IAP and the peak relieving force attributable to the IAP. Separate univariate analyses of variance (ANOVAs; p [less than] 0.05) were performed for all joint kinematic data.


Kinetics. Peak compressive force at the L5-S1 joint did not differ significantly, F(3,32) = 0.06, p = 0.9814, between the four lifting conditions (Table 1, [ILLUSTRATION FOR FIGURE 4 OMITTED]). The largest calculated compressive force occurred in the weight belt condition (7 kN), and the smallest force was observed in the condition utilizing the weight belt with a rigid abdominal pad (6.51 kN).

The back supports did not produce significant differences in the peak shear force at the L5-S1 joint, F(3,32) = 0.81, p = 0.4958. However, it is of interest that the unsupported condition had the highest shear force compared with the three supported conditions.

The peak forces calculated in the extensor muscles of the lower back were larger, F(3,32) = 0.13, p = 0.9427, than the peak compressive forces, with values exceeding 7 kN in all the lifting conditions. The net extensor moment around the L5-S1 joint was not significantly different, F(3,32) = 0.20, p = 0.8951, with magnitudes greater than 580 N [multiplied by] m.

The IAP was calculated at approximately 20 kPa in all the lifting conditions, with no trends or statistical differences between the back supports, F(3,32) = 0.40, p = 0.7513. The relieving force attributable to the IAP was demonstrated at approximately 1.0 kN for all lifting trials, F(3,32) = 0.12, p = 0.9473. This value represents 15% of the average compressive force at L5-S1.

Kinematics. Subjects' lifting strategies were investigated to examine the similar force calculations among the lifting conditions. These calculations were determined from the body position of the lower extremities and trunk at the time of peak L5-S1 compressive force (Table 2). When examining the joint angles at the time of peak compressive force, we observed no significant differences (p [less than] 0.05) among the lifting conditions. Thus lifting strategies did not change as a result of the supported conditions.

Subjects lifted the box at similar speeds in the different back support conditions, with an average velocity near 0.74 m/s. The total times for the lifts were more variable but still not significantly different, with ranges between 1.43 s for the elastic binder condition and 1.59 s for the weight belt with abdominal pad.


The biomechanical effects of manual lifting on the lumbosacral spine are considered significant (Chaffin, 1967; Freivalds et al., 1984). It remains possible that these effects may be altered when the trunk is supported by a weight-training belt and/or abdominal binder. Various parameters have been employed to study the effects of lumbosacral support. Trunk EMG (Lander et al., 1990, 1992), IAP (Harman et al., 1989), and intrarectal pressure (IRP) effects (Lander et al., 1990, 1992) have indicated some advantage (i.e., reduction in lumbosacral forces) when wearing lumbosacral supports during lifting.

Results from the present study noted no such positive statistical outcomes when examining the kinetic and kinematic effects of wearing a lumbosacral support. Differences in L-5 and S-1 vertebrae compressions, shear forces, and moments were not statistically significant among the various belt designs. Differences between the supported and unsupported conditions with regard to lower extremity joint kinematic effects were nonsignificant. These findings indicated that lumbosacral support neither alters nor hinders lifting technique or strategy. Intra-abdominal pressure and its relieving force did not differ significantly among the various lifting conditions.

Peak Kinetic Values during Lifting (Mean and Standard Deviation;
n = 9)

                                     Weight      Weight      Elastic
Forces                No Support     Belt     Belt and Pad    Binder

Compressive                6.78       7.00         6.51         6.60
(kN)                      (2.52)     (3.27)       (2.52)       (2.40)

Shear                      4.43       3.78         3.89         3.69
(kN)                      (1.36)     (0.89)       (1.00)       (1.11)

Extensor muscle            8.26       8.25         7.66         7.78
(kN)                      (2.68)     (3.18)       (2.21)       (2.35)

FIAP                       0.94       1.00         0.98         0.99
(kN)                      (0.17)     (0.21)       (0.21)       (0.25)

IAP                       20.1       21.0         20.4         20.2
(kPa)                     (3.3)      (4.6)        (4.6)        (5.4)

Extensor moment          606.0      639.0        580.0        591.0
(N[multiplied by]m)     (165.0)    (215.0)      (143.0)      (150.0)

Findings from the present study indicated that differences in IAP and its relieving force are minimal when comparing various belt designs inclusive of an abdominal pad. This is similar to the results of another study, which showed that wearing belts did not significantly increase IAP above that which could be achieved by holding [TABULAR DATA FOR TABLE 2 OMITTED] the breath (McGill, Norman, and Sharratt, 1990). One possible explanation for this could be that the young, healthy nature of the subjects resulted in the ability to develop sufficient contraction in the abdominal wall through activation of the truncal musculature, especially the transverse abdominis muscle. Investigators have shown that high levels of intra-abdominal pressures developed concurrently with near-maximal activation of the transverse abdominis muscle (Cresswell, Grundstrom, and Thorstensson, 1992). Because breath holding was not permitted in the present study, the absence of a Valsalva maneuver may have resulted in lower IAPs.

Interestingly, intrarectal pressure (IRP) has been hypothesized to be representative of intra-peritoneal pressure (Nordin, Elfstrom, and Dahlquist, 1984; Rushmer, 1946). Lander et al. (1990, 1992) utilized intrarectal pressures and interpreted their findings as indications of IAP. However, it should be noted that recordings of pressure estimates in work by Rushmer (1946) and by Nordin et al. (1984) were performed at rest and in an anatomically relaxed state. Generalizing these same effects of IRP as an indication of intraperitoneal pressure may be untenable during exertional activity. This may offer some rationale as to why different changes were demonstrated in the present study compared with those of Lander et al. (1990, 1992).

Moreover, Lander et al. (1990, 1992) recorded no significant differences when examining force platform and kinematic variables when subjects performed squat lifts at near-maximal levels. Harman et al. (1989) reported similar nonsignificant results when comparing peak ground reaction forces for subjects when lifting during supported and unsupported conditions. Both investigators employed controlled squat-lifting styles without significant deviation in lifting strategies by the subjects. Spine kinetic observations from the present study noted similar findings and resulted in no significant differences between subjects wearing various belt designs and subjects in the unsupported condition.

It appears from this investigation that lumbar kinetics are not augmented by the use of lumbosacral devices. This study did account for research on the kinetic effects of lumbosacral support in healthy young men during a specifically controlled lifting task employing restricted lifting strategies. Therefore, the results should not be advanced as definitive regarding the effectiveness of lumbosacral supports during lifting; rather, they should be taken as postulates in the ongoing resolution of back support theory.


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MICHAEL L. WOODHOUSE received his Ph.D. in exercise science from Old Dominion University. He is professor of exercise sciences and director of Exercise Science Laboratories at Norfolk State University and associate research professor of orthopedics at Eastern Virginia Medical School in Norfolk, Virginia. He serves as senior biomechanist for Forensic Technologies International Corporation in Annapolis, Maryland. His research focus has been primarily in the area of orthopedic biomechanics.

RAYMOND W. McCOY received his Ph.D. in biomechanics at the University of Southern California and is assistant professor of kinesiology and director of bio-mechanics research at the College of William and Mary in Williamsburg, Virginia. He also serves as assistant research professor of orthopedics at Eastern Virginia Medical School in Norfolk, Virginia. He has served as sports biomechanist for both the National and International Olympic Committees. His current research focus is in the area of orthopedic and occupational biomechanics.

DIEGO R. REDONDO received his Ph.D. in exercise science from the University of Southern Mississippi and has since been employed at Old Dominion University, where he is assistant professor of exercise science and sports medicine and director of the Laboratory of Kinesiological and Biomechanical Studies. His primary research interest is orthopedic and occupational biomechanics. Redondo is also assistant research professor of orthopedics at Eastern Virginia Medical School in Norfolk, Virginia.

LAWRENCE M. SHALL is a practicing orthopedic surgeon and associate professor of orthopedics at Eastern Virginia Medical School. He is a graduate of the Medical College of Ohio at Toledo. Shall's research emphasis has been primarily in the area of sports medicine.
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Author:Woodhouse, Michael L.
Publication:Human Factors
Date:Sep 1, 1995
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