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UNLABELLED: Spinal curvatures alter measured stature and may influence the evaluation of skeletal maturity and growth based on stature measurements. METHODS: A dataset of calibrated measurements of vertebral positions of 407 radiographs in the frontal plane, together with clinically measured Cobb angles was used to determine the difference between spinal length and spinal height ('height loss') as a function of Cobb angles for radiographs indicating both single (N=182) and double (N=225) curves. RESULTS: An apparently quadratic relationship: Height loss (mm)=1.0+0.066*Cobb+0.0084*Cobb*Cobb was found between height loss and each patient's mean Cobb angle for double curves. There was close agreement of the regression coefficients for single and double curves, and the present findings were very similar to the relationship reported by Ylikoski (Eur Spine J, 2003, 12:288-291). The relationships differed substantially from those proposed by Bjure (Clin Orthop, 1973 93:44-52) and by Brookenthal (SRS Exhibit 15, 2002). DISCUSSION AND CONCLUSIONS: The findings of the present study indicate that height loss (in mm) occurring with a 10 degrees increase in mean Cobb angle (for two curves) would be 1.1+0.16 times the mean Cobb angle (in degrees). For example, for a Cobb angle change from 30 to 40 degrees, the expected height loss would be 1.1+35*0.16 mm=6.7 mm. This assumes that height loss occurs only as a result of altered curvature, without alteration in disc height associated with an increase in scoliosis.

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The intervertebral discs become wedged and narrowed in a scoliosis curve, and this may be due in part to altered biomechanical environment. To study this, external rings were attached by percutaneous pins transfixing adjacent vertebrae in 5-week-old Sprague-Dawley rats and four permutations of mechanical conditions (4 groups of animals) were compared: (A) 15 degrees Angulation, (B) Angulation with 0.1 MPa Compression, (C) 0.1 MPa Compression, and (D) Reduced mobility. These altered mechanical conditions were applied for 5 weeks. After 5 weeks, disc narrowing at the intervention levels was evident in micro-CT images. Average disc space loss as a percent of the initial values over the 5 weeks was 19%, 28%, 22% and 20% four groups listed above. Increased lateral bending stiffness relative to within-animal controls was also observed at all groups. The minimum stiffness was recorded at an angle close to the in vivo value, indicating that angulated discs had adapted to the imposed deformity. In the angulated and compressed discs there was a small difference in the amount of collagen crimping in the disc annuli between concave and convex sides. All experimental interventions produced substantial changes in the intervertebral discs of these growing animals. 'Reduced mobility' was present in all interventions, and the changes in the discs with reduced mobility alone were comparable with those in loaded and angulated discs. This suggests that imposed reduced mobility is the major source of disc changes, and may be a factor in disc degeneration in scoliosis. Further studies are in progress to characterize gene expression, matrix protein synthesis and composition in these discs.

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Human intervertebral disc specimens were tested to determine the regions of largest maximum shear strain (MSS) experienced by disc tissues in each of three principal displacements and three rotations, and to identify the physiological rotations and displacements that may place the disc at greatest risk for large tissue strains and injury. Tearing of disc annulus may be initiated by large interlamellar shear strains. Nine human lumbar discs were tagged with radiographic markers on the endplates, disc periphery and with a grid of wires in the mid-transverse plane and subjected to each of the six principal displacements and rotations. Stereo-radiographs were taken in each position and digitized for reconstruction of the three-dimensional position of each marker. Maximum tissue shear strains were calculated from relative marker displacements and normalized by the input displacement or rotation. Lateral shear, compression, and lateral bending were the motions that produced the mean (95% confidence interval) largest mean MSS of 9.6 (0.7)%/mm, 9.0 (0.5)%/mm, and 5.8 (1.6)%/ degrees , respectively, and which occurred in the posterior, posterolateral and lateral peripheral regions of the disc. After taking into account the reported maximum physiological range of motion for each degree of freedom, motions producing the highest physiological MSS were lateral bending (57.8 (16.2)%) and flexion (38.3 (3.3)%), followed by lateral shear (14.4 (1.1)%) and compression (12.6 (0.7)%).

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Scoliosis is a three-dimensional deformation of the spine that can be treated by vertebral fusion using surgical instrumentation. However, the optimal configuration of instrumentation remains controversial. Simulating the surgical maneuvers with personalized biomechanical models may provide an analytical tool to determine instrumentation configuration during the pre-operative planning. Finite element models used in surgical simulations display convergence difficulties as a result of discontinuities and stiffness differences between elements. A kinetic model using flexible mechanisms has been developed to address this problem, and this study presents its use in the simulation of Cotrel-Dubousset Horizon surgical maneuvers. The model of the spine is composed of rigid bodies corresponding to the thoracic and lumbar vertebrae, and flexible elements representing the intervertebral structures. The model was personalized to the geometry of three scoliotic patients (with a thoracic Cobb angle of 45 degrees, 49 degrees and 39 degrees ). Binary joints and kinematic constraints were used to represent the rod-implant-vertebra joints. The correction procedure was simulated using three steps: (1) Translation of hooks and screws on the first rod; (2) 90 degrees rod rotation; (3) Hooks and screws look-up on the rod. After the simulation, slight differences of 0-6 degrees were found for the thoracic spine scoliosis and the kyphosis, and of 1-8 degrees for the axial rotation of the apical vertebra and for the orientation of the plane of maximum deformity, compared to the real post-operative shape of the patient. Reaction loads at the vertebra-implant link were mostly below 1000 N, while reaction loads at the boundary conditions (representing the overall action of the surgeon) were in the range 7-470 N and maximum torque applied to the rod was 1.8 Nm. This kinetic modeling approach using flexible mechanisms provided a realistic representation of the surgical maneuvers. It may offer a tool to predict spinal geometry correction and assist in the pre-operative planning of surgical instrumentation of the scoliotic spine.

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The growth (i.e. increase of external dimensions) of long bones and vertebrae occurs longitudinally by endochondral ossification at the growth plates, and radially by apposition of bone at the periosteum. It is thought that mechanical loading influences the rate of longitudinal growth. The 'Hueter-Volkmann Law' proposes that growth is retarded by increased mechanical compression, and accelerated by reduced loading in comparison with normal values. The present understanding of this mechanism of bone growth modulation comes from a combination of clinical observation (where altered loading and growth is implicated in some skeletal deformities) and animal experiments in which growth plates of growing animals have been loaded. The gross effect of growth modulation has been demonstrated qualitatively and semi-quantitatively. Sustained compression of physiological magnitude inhibits growth by 40% or more. Distraction increases growth rate by a much smaller amount. Experimental studies are underway to determine how data from animal studies can be scaled to other growth plates. Variables include: differing sizes of growth plate, different anatomical locations, different species and variable growth rate at different stages of skeletal maturity. The two major determinants of longitudinal growth are the rate of chondrocytic proliferation and the amount of chondrocytic enlargement (hypertrophy) in the growth direction. It is largely unknown what are the relative changes in these key variables in mechanically modulated growth, and what are the signaling pathways that produce these changes.

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A retrospective longitudinal radiographic study of patients with progressive scoliosis was conducted to determine the relative amount of wedging between vertebrae and discs as a function of progression of the scoliosis curve, cause of the scoliosis, and anatomic curve region. Posteroanterior radiographs of 27 patients with idiopathic scoliosis and of 17 patients with scoliosis associated with cerebral palsy were studied. The amount of wedging of vertebrae and discs at the curve apex was measured by the Cobb method and expressed as a proportion of the curve's Cobb angle. On average, the relative amount of vertebral and disc wedging did not differ significantly between initial and follow-up radiographs made after progression of the scoliosis. In both groups of patients, the mean vertebral wedging was more than the disc wedging in the thoracic region; the converse was found in curves in the lumbar and thoracolumbar regions. The patients with scoliosis associated with cerebral palsy had curves that were longer and more commonly in the thoracolumbar and lumbar regions. The relative wedging did not change significantly with curve progression and did not appear to differ by diagnosis. In the management of scoliosis, including small curves, it should be recognized that both the vertebrae and discs have a wedging deformity.

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The hypothesis that control of lumbar spinal muscle synergies is biomechanically optimized was studied by comparing EMG data with an analytical model with a multi-component cost function that could include (1) trunk displacements, (2) intervertebral displacements, (3) intervertebral forces; (4) sum of cubed muscle stresses, and (5) eigenvalues for the first two spinal buckling modes. The model's independent variables were 180 muscle forces. The 36 displacements of 6 vertebrae were calculated from muscle forces and the spinal stiffness. Calculated muscle activation was compared with EMG data from 14 healthy human subjects who performed isometric voluntary ramped maximum efforts at angles of 0 degrees, 45 degrees, 90 degrees, 135 degrees and 180 degrees to the right from the anterior direction. Muscle activation at each angle was quantified as the linear regression slope of the RMS EMG versus external force relationship, normalized by the maximum observed EMG.There was good agreement between the analytical model and EMG data for the dorsal muscles when the model included either minimization of intervertebral displacements or minimization of intervertebral forces in its cost function, but the model did not predict a realistic level of abdominal muscles activation. Agreement with EMG data was improved with the sum of the cubed muscle stresses added to the cost function. Addition of a cost function component to maximize the trunk stability produced higher levels of antagonistic muscle activation at low efforts than at greater efforts. It was concluded that the muscle activation strategy efficiently limits intervertebral forces and displacements, and that costs of higher muscle stresses are taken into account, but stability does not appear to be maximized. Trunk muscles are apparently not controlled solely to optimize any one of the biomechanical costs considered here.

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Trunk stiffness was measured in healthy human subjects as a function of steady-state preload efforts in different horizontal loading directions. Since muscle stiffness increases with increased muscle activation associated with increasing effort, it is believed that coactivation of muscles helps to stiffen and stabilize the trunk. This paper tested whether increased steady-state preload effort increases trunk stiffness. Fourteen young healthy subjects each stood in an apparatus with the pelvis immobilized. They were loaded horizontally at directions of 0, 45, 90, 135 and 180 degrees to the forward direction via a thoracic harness. Subjects first equilibrated with a steady-state load of 20 or 40% of their maximum extension effort. Then a sine-wave force perturbation of nominal amplitude of 7.5 or 15% of maximum effort and nominal period of 250ms was applied. Both the applied force and subsequent motion were recorded. Effective trunk mass and trunk-driving point stiffness were estimated by fitting the experimental data to a second-order differential equation of the trunk dynamic behavior. The mean effective trunk mass was 14.1kg (s.d.=4.7). The trunk-driving point stiffness increased on average 36.8% (from 14.5 to 19.8N/mm) with an increase in the nominal steady-state preload effort from 20 to 40% (F(1,13)=204.96, p<0.001). There was a smaller, but significant variation in trunk stiffness with loading direction. The measured increase in trunk stiffness probably results from increased muscle stiffness with increased muscle activation at higher steady-state efforts.

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STUDY DESIGN: An experimental study of healthy subjects' trunk muscle responses to force perturbations at differing angles and steady state efforts. OBJECTIVES: To determine whether increased preactivation of muscles was associated with decreased likelihood of muscular activation in response to a transient force perturbation. SUMMARY OF BACKGROUND DATA: Trunk stability (ability to return to equilibrium position after a perturbation) requires the stiffness of appropriately activated muscles to prevent buckling and consequent "self-injury." Therefore, greater trunk muscle preactivation might decrease the likelihood of reflex muscle responses to small perturbations. METHODS: Each of 13 subjects stood in an apparatus with the pelvis immobilized. A harness around the thorax provided a preload and a force perturbation by a horizontal cable and a movable pulley attached to one of five anchorage points on a wall track surrounding the subject at angles of 0 degrees, 45 degrees, 90 degrees, 135 degrees, and 180 degrees to the forward direction. Subjects first equilibrated with a preload effort of nominally 20% or 40% of their maximum extension effort. Then a single full sine-wave force perturbation pulse of nominal amplitude, 7.5% or 15% of maximum effort, duration 80 milliseconds or 300 milliseconds, was applied at a random time, with three repeated trials of each test condition. The applied force (via a load cell) and the electromyographic activity of six right and left pairs of trunk muscles were recorded. Muscle responses were detected by two methods. 1) Shewhart method: electromyographic signal greater than "baseline" values by more than three standard deviations, and 2) Mean Electromyographic Difference method: mean electromyographic signal in a time window 25 to 150 milliseconds after the force perturbation greater than that in a 25- to 150-millisecond window before the perturbation. RESULTS: Lower preload efforts were associated with more muscle responses (overall mean response detection rate = 33% at low preload and 25% at high preload). Using the Shewhart method, there were significant differences by effort (P<0.05) for all abdominal muscles and for all left dorsal muscles except multifidus. Using the Mean Electromyographic Difference method, there were significant differences by effort (P<0.05) for the same dorsal muscles, but only for one of the abdominal muscles. CONCLUSIONS: Findings are consistent with the hypothesis that the spine can be stabilized by the stiffness of activated muscles, obviating the need for active muscle responses to perturbations.

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The stiffness of activated muscles may stabilize a loaded joint by preventing perturbations from causing large displacements and injuring the joint. Here the elbow muscle recruitment patterns were compared with the forearm loaded vertically (a potentially unstable inverted pendulum configuration) and with horizontal loading. Eighteen healthy subjects were studied with the forearm vertical and supinated and the elbow flexed approximately 90 degrees. In the first experiment EMG electrodes recorded activity of biceps, triceps, and brachioradialis muscles for joint torques produced (a) by voluntarily exerting a horizontal force isometrically (b) by voluntarily flexing and extending the elbow while the forearm was loaded vertically with 135N. The relationship between the EMG and the torque generated was quantified by the linear regression slope and zero-torque intercept. In a second experiment a vertical load increasing linearly with time up to 300N was applied. In experiment 1 the EMG-torque relationships for biceps and triceps had an intercept about 10% of maximum voluntary effort greater with the vertical compared to the horizontal force, the inverse was found for Brachioradialis, but the EMG-torque slopes for both agonist and antagonistic muscles were not different. In experiment 2 there were 29 trials with minimal elbow displacement and all the three muscles activated on the order of 11% of maximum activation to stabilize the elbow; 19 trials had small elbow extension and 14 trials small flexion requiring altered muscle forces for equilibrium; 7 trials ended in large unstable displacement or early termination of the test. An analysis indicate that the observed levels of muscle activation would only provide stability if the muscles' short-range stiffness was at the high end of the published range, hence the elbow was marginally stable. The stability analysis also indicated that the small elbow extension increased stability and flexion decreased stability.

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This study tested the following hypotheses: (a) a vertebral wedge deformity created by chronic static asymmetrical loading will be corrected by reversal of the load asymmetry; (b) a vertebral wedge deformity created by chronic static asymmetrical loading will remain if the load is simply removed; and (c) vertebral longitudinal growth rates, altered by chronic static loading, will return to normal after removal of the load. An external fixator was used to impose an angular deformity (Cobb angle of 30 degrees) and an axial compression force (60% body weight) on the ninth caudal (apical) vertebra in two groups of 12 5-week-old Sprague-Dawley rats. This asymmetrical loading was applied to all rats for 4 weeks to create an initial wedge deformity in the apical vertebra. The rats from group I (load reversal) then underwent 1 week of distraction loading followed by 4 weeks of asymmetrical compressive loading with the imposed 30 degree Cobb angle reversed. The rats from group II (load removal) had the apparatus removed and were followed for 5 weeks with no external loading. Weekly radiographs were obtained and serial fluorochrome labels were administered to follow vertebral wedging. After the initial 4-week loading period, the combined average wedge deformity that developed in the apical vertebra of the animals in both groups was 10.7 +/- 4.4 degrees. The group that underwent load reversal showed significant correction of the deformity with the wedging of the apical vertebra decreasing to, on average, 0.1 +/- 1.4 degrees during the 4 weeks of load reversal. Wedging of the apical vertebra in the group that underwent load removal significantly decreased to 7.3 +/- 3.9 degrees during the first week after removal of the load, but no significant changes in wedging occurred after that week. This indicated a return to a normal growth pattern following the removal of the asymmetrically applied loading. The longitudinal growth rate of the apical vertebra also returned to normal following removal of the load. Vertebrae maintained under a load of 60% body weight grew at a rate that was 59.4 +/- 17.0% lower than that of the control vertebrae, whereas after vertebrae were unloaded their growth averaged 102.4 +/- 31.8%. These findings show that a vertebral wedge deformity can be corrected by reversing the load used to create it and that vertebral growth is not permanently affected by applied loading.

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STUDY DESIGN: An Ilizarov-type apparatus was applied to the tails of rats to assess the influence of immobilization, chronically applied compression, and sham intervention on intervertebral discs of mature rats. OBJECTIVES: To test the hypothesis that chronically applied compressive forces and immobilization cause changes in the biomechanical behavior and biochemical composition of rat tail intervertebral discs. SUMMARY OF BACKGROUND DATA: Mechanical factors are associated with degenerative disc disease and low back pain, yet there have been few controlled studies in which the effects of compressive forces on the structure and function of the disc have been isolated. METHODS: The tails of 16 Sprague-Dawley rats were instrumented with an Ilizarov-type apparatus. Animals were separated into sham, immobilization, and compression groups based on the mechanical conditions imposed. In vivo biomechanical measurements of disc thickness, angular laxity, and axial and angular compliance were made at 14-day intervals during the course of the 56-day experiment, after which discs were harvested for measurement of water, proteoglycan, and collagen contents. RESULTS: Application of pins and rings alone (sham group) resulted in relatively small changes of in vivo biomechanical behavior. Immobilization resulted in decreased disc thickness, axial compliance, and angular laxity. Chronically applied compression had effects similar to those of immobilization alone but induced those changes earlier and in larger magnitudes. Application of external compressive forces also caused an increase in proteoglycan content of the intervertebral discs. CONCLUSIONS: The well-controlled loading environment applied to the discs in this model provides a means of isolating the influence of joint-loading conditions on the response of the intervertebral disc. Results indicate that chronically applied compressive forces, in the absence of any disease process, caused changes in mechanical properties and composition of tail discs. These changes have similarities and differences in comparison with human spinal disc degeneration.

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The purpose of this study was to determine whether compression and distraction applied to adjacent vertebrae in the calf tail could modulate vertebral growth. Seven 6-week-old calves had two pairs of adjacent tail vertebrae instrumented with an Ilizarov external fixator with calibrated springs designed to apply a 30-50-N axial load to the vertebrae. Data were obtained from 17 vertebrae loaded in compression and 4 vertebrae loaded in distraction. Vertebrae adjacent (cranial and caudal) to the instrumented vertebrae served as controls. The length of each vertebra on the postoperative radiograph was subtracted from the length of the same vertebra on the radiograph taken 6 months after the operation to calculate vertebral growth. The vertebrae loaded in compression had a growth rate of 68 +/- 42% of that of the controls. In contrast, the vertebrae loaded in distraction had a growth rate of 123 +/- 78% of that of the controls.

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This paper describes the anatomy of the musculature crossing the lumbar spine in a standardized form to provide data generally suitable for static biomechanical analyses of muscle and spinal forces. The muscular anatomy from several sources was quantified and transformed to the mean bony anatomy of four young healthy adults measured from standing stereo-radiographs. The origins, insertions and physiological cross-sectional area (PCSA) of 180 muscle slips which act on the lumbar spine are given relative to the bony anatomy defined by the locations of 12 thoracic and five lumbar vertebrae, and the sacrum, and the shape and positions of the 24 ribs. The broad oblique abdominal muscles are each represented by six vectors and an appropriate proportion of the total PCSA was assigned to each to represent the muscle biomechanics.

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Surgical instrumentation of the scoliotic spine is a complex procedure with many parameters, such as the spinal segment to operate on, the number and position of the hooks and screws, etc. Biomechanical modeling is a tool which can be used to determine the influence of these parameters. However, technical difficulties due to the large stiffness range of involved components and the large deformations associated with surgical maneuvers are encountered when using the finite elements method. Thus, the objective of this study is to adapt a modeling approach using analysis of flexible mechanisms and evaluate its feasibility. The model combines rigid bodies for the vertebrae and flexible elements representing intervertebral structures. The mechanical properties were calculated from published data and the geometry was personalized with intraoperative measurements. Following the installation of the hooks and screws on the modeled spine, two steps were used to simulate the surgical maneuvers: 1) translation and attachment of the hooks/screws on the first rod; 2) rod rotation. The feasibility of this modeling approach was evaluated by simulating the surgical maneuvers on 2 cases: 1) a physical model; 2) a clinical case. The agreement between intraoperative measurements and simulation results (frontal curvatures are reproduced with over 80% accuracy) shows the feasibility of the modeling approach. This approach also reduces computational convergence problems because of its limited sensitivity to stiffness differences between elements, which demonstrates the advantage of flexible mechanism modeling relative to finite element modeling. Long term goals of subsequent refinements are the development of a tool for surgical correction predictions and for the design of more efficient instrumentation.

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Spinal instrumentation has a primarily mechanical function, but mechanical testing procedures are designed more by expediency than by their ability to evaluate whether instrumentation will produce a good clinical outcome. Differing testing protocols preclude direct comparisons between measurements made in different laboratories. However, standardization of testing may not resolve the fact that much of the information from laboratory testing cannot be used in surgical decision-making. Researchers and journals that publish their work should focus on determining the mechanical requirements of instrumentation, including in vivo loading, the biologic response to instrumentation in the presence of pathology, and how this information can assist surgeons in selecting instrumentation in individual cases.

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Progression of scoliosis deformity during growth is thought to be caused by asymmetrical loading, resulting in asymmetrical growth with vertebral and disc wedging in a "vicious cycle." The purpose of this study was to quantify the changes in disc thickness during growth in rat tails subjected to compression or distraction loading for 6 or 9 weeks, to investigate the hypothesis that disc growth is mechanically modulated. Six-week-old Sprague-Dawley rats were studied with compression loading (13 animals) or distraction loading (15 animals) applied to their tails, and there were 8 sham animals. Loading was applied to tail segments by means of an external ring fixator. Radiographic measurements of disc thickness were made at biweekly intervals. From the initial to final radiograph, compressed discs had reduced thickness averaging (+/-SD) 0.50 +/- 0.28 mm, distraction discs had average increased thickness of 0.20 +/- 0.42 mm, and sham discs lost an average of 0.21 +/- 0.18 mm of thickness (analysis of variance p < 0.001). There was an "initial change" in disc thickness averaging 0.18 +/- 0.32 mm in nonloaded discs, which was similar in magnitude to the elastic deformation and was attributed to disc swelling under anesthesia. These results indicate that growth in disc thickness is mechanically modulated by axial loading in growing rats.

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STUDY DESIGN: A biomechanical model of the lumbar spine was used to calculate the effects of abdominal muscle coactivation on spinal stability. OBJECTIVES: To estimate the effects of abdominal muscle coactivation on lumbar spine stability, muscle fatigue rate, and lumbar spine compression forces. SUMMARY OF BACKGROUND DATA: The activation of human trunk muscles has been found to involve coactivation of antagonistic muscles, which has not been adequately predicted by biomechanical models. Antagonistic activation of abdominal muscles might produce flexion moments resulting from abdominal pressurization. Qualitatively, antagonistic activity also has been attributed to the need to stabilize the spine. METHODS: Spinal loads and spinal stability were calculated for maximum and submaximum (40%, 60% and 80%) efforts in extension and lateral bending using a previously published, anatomically realistic biomechanical model of the lumbar spine and its musculature. Three different antagonistic abdominal muscle coactivation patterns were imposed, and results were compared with those found in a model with no imposed coactivation. RESULTS: Results were quantified in terms of the sum of cubed muscle stresses (sigma sigma m3, which is related to the muscle fatigue rate), the maximum compressive loading on the lumbar spine, and the critical value of the muscle stiffness parameter (q) required for the spine to be stable. Forcing antagonistic coactivation increased stability, but at the cost of an increase in sigma sigma m3 and a small increase in maximum spinal compression. CONCLUSIONS: These analyses provide estimates of the effects of antagonistic abdominal muscle coactivation, indicating that its probable role is to stabilize the spine.

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STUDY DESIGN: A biomechanical model was used to calculate muscle and intervertebral forces in a spine with and without a lumbar scoliosis. OBJECTIVES: To quantify the loading of the motion segments in a lumbar scoliosis. SUMMARY OF BACKGROUND INFORMATION: Scoliosis is thought to cause asymmetric loading of vertebral physes, causing asymmetric growth according to the Hueter-Volkmann principle. The magnitude of vertebral loading asymmetry as a function of scoliosis magnitude is unknown, however, as is the sensitivity of growth to asymmetric loading. METHODS: The analysis included five lumbar vertebrae, the thorax, and the sacrum/pelvis and 90 pairs of multijoint muscles. Five spinal geometries were analyzed: the mean spinal shape of 15 patients with left lumbar scoliosis (38 degrees Cobb angle, apex at L1-L2, the reference or "100%" geometry), and the geometry scaled to 0%, 33%, 67%, and 132% of the asymmetry of the reference shape. The muscle and intervertebral forces for maximum efforts opposing moments applied to the T12 vertebra in each of the three principal directions were calculated. The loading at each intervertebral level was expressed as the resultant force (P), the axial torque, the lateral and anteroposterior offset of P from the disc center, and the angle of P from the axial direction. RESULTS: With increasing scoliosis, there was a weak trend of increasing lateral offset of P, but not consistently to either the convex or concave direction. There was a much stronger trend of increasing angle between the force P and the motion segment longitudinal axis with increasing Cobb angle. Typically, this angle was 10-30 degrees for the largest scoliosis (51 degrees Cobb) and in a direction tending to increase the scoliosis. This angulation of the force results from shear loading of the disc. Axial torques tending to increase the transverse plane deformity increased with scoliosis for extension efforts. CONCLUSIONS: These analyses indicate that lumbar scoliosis produces asymmetric spinal loading characterized by shear forces tending to increase the scoliosis, but with little increase in the asymmetric compression of motion segments. If scoliosis progression results from asymmetric loading, it appears that the shear force component is responsible.

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STUDY DESIGN: A rat tail model was used to test the hypothesis that angulation and asymmetric axial compressive loading would lead to vertebral wedging because of asymmetric longitudinal growth in the physes. OBJECTIVES: To study the effect of angulation and asymmetric loading on the progression of spinal curvature in a rat tail model. SUMMARY OF BACKGROUND DATA: Large idiopathic scoliotic curves in children with significant growth remaining are the curves most likely to progress. The mechanism of progression of skeletal deformities is thought to be controlled by the Hueter-Volkmann law, whereby additional axial compression decelerates growth, and reduced axial compression accelerates growth. It has been hypothesized that spinal curvature leads to asymmetric loading transversely along the vertebral growth plate, causing progressive vertebral wedging by means of a vicious cycle. METHODS: Two 32-mm diameter external ring fixators were glued to 0.7-mm pins that had been inserted percutaneously through the eighth and 10th caudal vertebra of 10 6-week-old Sprague-Dawley rats. Calibrated springs and 15 degrees wedges, mounted on stainless steel threaded rods passing through holes distributed around the rings, imposed a 30 degrees Cobb angle and axially compressed the instrumented vertebrae. Fluorochrome labels and radiographs were used to document the progression of vertebral wedging. RESULTS: The wedging initially was entirely in the intervertebral discs, but by 6 weeks the wedging of the discs and vertebrae were approximately equal. Fluorochrome labeling confirmed that the vertebral wedging resulted from asymmetric growth in the physes. CONCLUSIONS: This study shows that vertebrae, when asymmetrically loaded, become wedged. This is consistent with the concept of mechanically provoked progression of scoliotic deformities according to the Hueter-Volkmann law.

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