Chapter 6 Biomechanics of the spine
Mechanical loading of the spine is an important factor in the etiology of spinal disorders. It also affects the outcome of orthotic treatments for spinal disorders. Loads on the human spine are produced by (1) gravitational forces due to the mass of body segments, (2) external forces and moments induced by a physical activity, and (3) muscle tension. These loads are shared by the osseoligamentous tissues and muscles of the spine. Tensile forces in the paraspinal muscles, which exert a compressive load on the spine, balance the moments created by gravitational and external loads (Fig. 6-1). Because these muscles have a small moment arm from the spinal segment, they amplify the compressive load on the osseoligamentous spine.
Adapted from Mow VC, Flatow EL, Ateshian GA: Biomechanics. In Buckwalter JA, Einhorn TA, Simon SR, editors: Orthopaedic basic science, Rosemont, Ill, 2000, American Academy of Orthopaedic Surgeons.
The human spine is subjected to large compressive preloads during activities of daily living. The internal compressive forces on the ligamentous spine have been estimated for different physical tasks using kinematic and electromyographic data in conjunction with three-dimensional biomechanical models. The compressive force on the human lumbar spine is estimated to range from 200 to 300 N during supine and recumbent postures to 1400 N during relaxed standing with the trunk flexed 30 degrees. The compressive force may be substantially larger when an individual is holding a weight in the hands in the static standing posture and even more so during dynamic lifting. The human cervical spine also withstands substantial compressive preloads in vivo. Cervical preload approaches three times the weight of the head due to muscle coactivation forces in balancing the head in the neutral posture. The compressive preload on the cervical spine increases during flexion, extension, and other activities of daily living and is estimated to reach 1200 N in activities involving maximal isometric muscle efforts. In normal individuals, the spine sustains these loads without causing injuries to bony, soft tissue, or neurologic structures.
In the absence of muscle forces, the osteoligamentous spine cannot support vertical compressive loads of in vivo magnitude. Experiments in which a vertical load was applied at the cephalic end of cervical, thoracolumbar, or lumbar spine specimens caused buckling of the spines at load levels well below those seen in vivo. The stability of the spine, characterized by a critical load (maximum load carrying capacity, or Euler buckling load of spinal column), was determined by these experiments. When the load exceeded the critical value, the spine, constrained to move only in the frontal plane in these experiments, became unstable and buckled. The cervical spine buckled at a vertical load of approximately 10 N, the thoracolumbar spine at 20 N, and the lumbar spine at 88 N, all well below the compressive loads expected in vivo during activities of daily living. When a compressive load is applied in a vertical direction to a multisegment spine specimen, segmental bending moments and shear forces are induced as a result of the inherent curvature of the spine. This load application causes large changes in the specimen’s posture at relatively small loads. Further loading can cause damage to the soft tissue or bony structures.
Some investigators have modeled the muscles as springs in order to explain their role in preventing a buckling instability of the spinal column. Simulation of active muscle forces in experiments on the ligamentous spine is difficult because of the large number of muscles and the uncertainty in load sharing among the various muscles during different activities. The simulated muscle actions must provide stability to the ligamentous spine to carry compressive loads while permitting the mobility needed to perform the activities of daily living.
Analysis using muscle models of the trunk support the argument that the individual spinal segments, often referred to as functional spinal units (FSUs), are subjected to nearly pure compressive loads in vivo. Attempts to determine joint loads based on the assumption of a vertical load on the spine have resulted in serious overprediction of shear forces on the FSU. Calculations of spine models, taking into consideration the activity of paraspinal and abdominal muscles, demonstrated that, in weight-holding tasks, the compressive force on the lumbosacral disc increased with increasing trunk inclination and the amount of weight lifted, whereas the maximum anteroposterior shear force remained small (approximately 20% to 25% of the compressive force). The obliquity of the short lumbar extensor muscles allows them to share anterior shear forces resulting from lifting a load. When these muscles are activated to contribute a balancing extensor moment, they help to offset the anterior shear force on the lumbar FSU.
It can be reasoned that coactivation of trunk muscles alters the direction of the internal compressive force vector such that its path follows the lordotic and kyphotic curves of the spine, passing through the instantaneous center of rotation of each segment (Fig. 6-2). This would minimize the segmental bending moments and shear forces induced by the compressive load, allowing the ligamentous spine to support loads that otherwise would cause buckling and providing a greater margin of safety against both instability and tissue injury. The load vector described is called a follower load.
Fig. 6-2 Depiction of compressive vertical (A) and follower (B) load vectors. The compressive follower load vector in the sagittal plane passes through the flexion–extension instantaneous center of rotation of each segment, minimizing the coupled flexion–extension angular changes.
Experiments on human cadaveric specimens of lumbar (L1–5), thoracolumbar (T2–sacrum), and cervical spines (C2–7) as well as mathematical models have demonstrated that (1) the ligamentous spine with multiple motion segments can withstand physiologic compressive loads without tissue injury or instability if the compressive load vector is applied along a follower load path approximating the curve of the ligamentous spine, (2) the ligamentous spine subjected to compressive preloads of in vivo magnitude along the follower load path permits physiologic mobility under flexion–extension moments, and (3) the follower preload simulates the resultant vector of muscles that allow the spine to support physiologic compressive loads. Intradiscal pressures in human cadaveric lumbar spines under a follower preload are comparable to those measured in vivo, and spinal stability is increased without compromising its mobility in flexion–extension and lateral bending. A superimposed follower preload renders more physiologic the in vitro loading of the ligamentous spine with pure moments.
The follower load concept suggests a new hypothesis for the role of muscle coactivation in providing in vivo spine stability. Coactivation of trunk muscles (e.g., lumbar multifidus, longissimus lumborum, iliocostalis lumborum) could alter the direction of the resultant internal force such that its path follows the curve of the spine (follower load path), allowing the ligamentous spine to support compressive loads that otherwise would cause buckling of the column. Muscle dysfunction can induce abnormal shear forces at the lumbar FSU, leading to segmental instability in the presence of disc degeneration. On the other hand, a compressive follower preload produced by coordinated muscle action could stabilize shear instability in a degenerative FSU. This suggests a role for muscle conditioning and therapy in treating degenerative spine conditions.
A spinal motion segment is the smallest functional unit of the osteoligamentous spine and exhibits the generic characteristics of the spine. The FSU consists of two vertebral bodies connected by an intervertebral disc, facet joints, and ligaments (except at the C1–2 segment, where no intervertebral disc is present). The FSU can be viewed as a three-joint complex consisting of the disc (a cartilaginous joint) and two facet joints (synovial joints). A dynamic relationship exists between the intervertebral disc and facet joints in sharing physiologic loads.
The intervertebral disc carries substantial loads as a result of gravitational and muscle forces. It is the major anterior load-bearing element in axial compression and flexion. In the young healthy spine, load transmission from vertebra to vertebra occurs primarily through the disc’s nucleus pulposus. As load is applied to the healthy disc, forces are distributed equally in all directions from within the nucleus, placing the annulus fibers in tension. The collagen fibers of the annulus fibrosis are well suited to resisting tension along the fiber direction. The pressure in the nucleus pulposus stretches the fibers in the annulus, and the resistance of the fibers to tensile loading allows the annulus to contribute to load sharing. The annulus fibrosis is well suited to resisting torsion as a result of the characteristic orientation of fibers in the each layer. The intervertebral disc provides most of the motion segment’s stiffness in compression, whereas ligaments and facets contribute significantly to resisting bending moments and axial torsion.
Facet joints provide a posterior load path and have an important role in determining the limits of motion in the FSU. Biomechanical studies demonstrated that facets in the lumbar spine carry 10% to 20% of the compressive load when a person is in the standing upright position. The proportion of the total load shared by the disc increases with flexion. Load transmission through the articular facet surfaces as well as through the tips of the inferior facets in extension relieves some of the load on the intervertebral disc. Maintenance of cervical and lumbar lordosis helps to reduce the load on the disc, whereas flexion increases disc loading. The contribution of the facet joints to the stability of an FSU is also dependent on the capsular ligament and the level within the spine. For example, thoracic facets have limited capsular reinforcement, which facilitates axial rotation, in contrast to the lumbar spine, where the facet capsule is well developed and capable of stabilizing the spine against rotation and lateral bending.