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Quantifying the stability of the spinal column offers a perspective on the effectiveness of the motor control strategy to ensure a stable spine--and minimize the risk of injury from experiencing an unstable event. There are essentially three energy based methods of calculating a stability index for the lumbar spine. All three methods involve mathematical manipulation of an 18 x 18 Hessian matrix. The purpose of this paper was to consider the mathematical implications for the three methods of determining a single stability index, and examine the effects of biological factors such as muscle activation in each of these methods. The first approach computes the Hessian's determinant and is thought to represent a more global or "average" perspective on stability. A second approach computes the smallest eigenvalue of the Hessian matrix to determine the weakest link of the spine. The final method determines an average critical stiffness difference for the spine and is intended to effectively determines how far a human spine is from instability, and allows comparison between tasks. This study shows that the same interpretation of stability is achieved via all three computational approaches--they agree as to whether the spine is stable or not. However they appear to differ in their sensitivity to the effect of muscle activation patterns.

This work describes a dynamic model of the low back that incorporates extensive anatomical detail of the musculo-ligamentous-skeletal system to predict the load time histories of individual tissues. The dynamic reaction moment about L4/L5 was determined during lateral bending from a linked-segment model. This reaction moment was partitioned into restorative components provided by the disc, ligament strain, and active-muscle contraction using a second model of the spine that incorporated a detailed representation of the anatomy. Muscle contraction forces were estimated using both information from surface electromyograms, collected from 12 sites, and consideration of the modulating effects of muscle length, cross-sectional area and passive elasticity. This modelling technique is sensitive to the different ways in which individuals recruit their musculature to satisfy moment constraints. Time histories of muscle forces are provided. High muscle loads are consistent with the common clinical observation of muscle strain often produced by load handling. Furthermore, the coactivation measured in muscles on both sides of the trunk suggests that muscles are recruited to satisfy the lateral bending reaction torque in addition to performing other mechanical roles such as spine stabilization. If an estimate of the intervertebral joint compression is desired for assessment of lateral bends in industry, then a single equivalent lateral muscle with a moment arm of approximately 3.0-4.0 cm would conservatively capture the effects of muscle co-contraction quantified in this study.

The mechanical role of the lumbar posterior ligaments during lifting tasks remains controversial. This study was designed to assess the ligament and disc contribution in resisting trunk flexion moment during extremely heavy lifts performed by national class powerlifters. Direct measurements of lumbar vertebrae kinematics in sagittal plane were obtained from videofluoroscopy utilizing multiple digitizing, correction for optical distortions and digital filtering. Four experienced powerlifters executed three trials, resulting in about 72 mA s of total radiation exposure. In the first trial, joint angles were measured when subjects fully flexed their spines to a point where the passive tissues resisted the flexor moment creating myoelectric silence in the extensor musculature. Next, two conventional deadlift style lifts were executed with the barbell load ranging from 183.7 to 210.9 kg. Four vertebral corners were digitized at a sampling rate of 30 Hz. The relative intervertebral joint angles, distance between the ligament attachment points, shearing and compressive displacements were calculated from a rigid body motion approach. Analysis revealed that except for one trial of one subject, they accomplished their lifts with an amount of lumbar flexion between 1.5 and 13 degrees less than they demonstrated during full flexion. Resultant ligament lengths at the beginning of the lifts ranged from 56.1 to 99.8% of their lengths when the trunk was fully flexed. It was concluded that ligaments did not strain sufficiently to contribute substantial resistance to the trunk flexion moment, relegating this responsibility to the musculature.

Biomechanical models utilized for analysis of tasks that load the lumbar spine often predict the resultant moment, disc compression and sometimes shear. Usually the extensor muscular and ligament forces of the lumbar spine are assumed to act 5 cm posterior to a disc centre of rotation. This study has re-examined the generation and pathways of muscular force transmission within the extensor musculature. The effects on L4/L5 disc compression and shear estimates of an anatomically and biomechanically justifiable range of tissue moment arms, lines of force and force generating capacity of muscle, input to a computer model, have been determined. Results indicated that L4/L5 compression estimates could be reduced by up to 35% when the output from a more realistic anatomical model of the erector spinae muscle group was compared with that from the frequently reported and simplified single muscle equivalent with a 5 cm moment arm. The shear force estimates could be altered from more than 500 N (L4 tending to shear anteriorly on L5) to less than 200 N with L4 tending to shear posteriorly on L5. Using the combination of input variables considered by the authors to be most feasible to estimate compression, a single 'equivalent' extensor soft tissue moment arm of 7.5 rather than 5 cm would be needed to equate the compression. This simplification of course, does not accommodate the shear force estimate problem.

Assessment of the effects of lifting on the low back has most frequently been done with the aid of static models. Many lifting movements appear to have substantial inertial components. It was of interest, therefore, to determine the size of the difference between statically and dynamically calculated lumbar moments during a demanding but not unusual manual lift observed in a metal fabrication industry. The results of several trials by four young men showed that the dynamic model resulted in peak L4/L5 moments 19% higher on average, with a maximum difference of 52%, than those determined from the static model. The technique adopted in the lift could minimize the difference. When the inertial forces of the load itself and the load weight were incorporated into an otherwise static model (quasi-dynamic) then the resulting L4/L5 moments exceeded those of the fully dynamic model by 25%. In many industrial tasks static analyses may severely underestimate the demands of dynamic lifts. These results show that a reasonably inexpensive approach in lifting task analysis is to measure the dynamic forces of the load on the hands and to use these in an otherwise static model. This results in a conservative assessment of the injury risk of lifts at least of the type reported in this study.

The purpose of this study was to add to the growing database of cross-sectional areas and moment arm lengths of trunk musculature using the methods of computerized tomographic scanning. An attempt was also made to estimate muscle force and moment generating capacity under various reported values of muscle force per unit cross-sectional area. The data were obtained on 13 active men 40.5 +/- 11.9 years of age, 173.8 +/- 5.9 cm tall and 89.1 +/- 11.7 kg body mass. Transverse CT scans were taken at the level of the L4/L5 disc with the subjects supine. Muscle cross-sectional areas were measured from 35 mm slides of the scans using a planimeter and moment arm length in the transverse plane were taken from the centroid of the L4/L5 disc to the centroid of the muscle section. Prior to estimating force and moment generating capacity, areas were corrected, where necessary, for fibre pennation angle to produce a physiological cross-sectional area. The physiological cross-sectional areas (cm2) for one side of the body were (mean +/- S.D.): sacrospinalis (SS) 15.9 +/- 2.5; multifidus (Mu) 4.2 +/- 0.7; psoas (Ps) 17.6 +/- 4.0; rectus abdominis (RA) 7.9 +/- 2.5; external oblique (EO) 9.4 +/- 2.7; internal oblique (IO) 8.1 +/- 2.3; transverse abdominus (TA) 2.9 +/- 1.3. The anterior posterior moment arm lengths were: erector mass (SS and Mu combined) 5.90 +/- 0.52; Ps 0.58 +/- 0.40; R.A. 10.28 +/- 2.07; E.O. (anterior portion) 5.94 +/- 1.39; E.O. (posterior portion) 2.08 +/- 1.39; I.O. (anterior portion) 6.92 +/- 1.63; I.O. (posterior portion) 3.85 +/- 1.54. The corresponding lateral moment arm lengths were: 3.26 +/- 0.36; 4.88 +/- 0.36; 4.35 +/- 1.31; 12.86 +/- 1.93; 13.95 +/- 1.16; 10.77 +/- 2.02; 12.52 +/- 1.26. The maximum force per unit cross-section that human muscles are capable of generating is not well defined. However, assuming an intermediate value of 50 N cm-2 of physiological cross-section, the erector musculature observed at the L4/L5 level should be capable of generating an extensor moment of about 118 N.m. At a muscle stress of 30 or 90 N cm-2, values also reported on human muscle, the moment would be 71 and 213 Nm, respectively. It must be remembered, however, that muscles not observable at the L4/L5 level can create moments around that center of rotation.(ABSTRACT TRUNCATED AT 400 WORDS)