Anatomy and biomechanics of the spinal column and cord

Handbook of Clinical Neurology, Vol. 109 (3rd series) Spinal Cord Injury J. Verhaagen and J.W. McDonald III, Editors # 2012 Elsevier B.V. All rights reserved

Chapter 2

Anatomy and biomechanics of the spinal column and cord VINCENT J. MIELE 1, MANOHAR M. PANJABI 2, AND EDWARD C. BENZEL 1* Department of Neurosurgery, Neurological Institute, Cleveland Clinic, Cleveland, OH, USA

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Section of Orthopedic Surgery, Yale University School of Medicine, New Haven, CT, USA

INTRODUCTION This chapter breaks the field of biomechanics down into three sections: (1) the fundamentals, including the laws of physics and physical principles; (2) the biomechanics of spinal column failure; (3) the biomechanics of spinal cord injury. They are presented in sequence in order to provide a foundation, followed by the development of biomechanical principles based on the foundations.

BIOMECHANICS: THE BASIC CONCEPTS Definitions Embarking on a discussion of the biomechanics of biological systems is best begun by understanding the definition of various terms commonly used in this field. Kinematics is the study of motion of objects without considering the factors that cause or affect the motion. The latter is the subject of dynamics. Momentum is the product of mass and velocity. Moment is a circular force creating a rotational vector around an axis. Torque is the magnitude of a force moment and is equal to the magnitude of the force multiplied by the perpendicular distance from the axis. Coupling occurs when more than one noncollinear force acts about the same axis and the resultant force moment is the sum of the individual forces. Example – when holding a weight out from the body, the compressive forces upon the spinal cord are offset/coupled with an extensor muscle force supplied by the erector spinal musculature. Stress is the force/load applied to an object divided by its cross-sectional area. Strain is the change in length of an object secondary to a deforming force.

Stress/strain behavior helps define an object’s intrinsic material properties. Modulus of elasticity is stress/strain. Stiffness is the relationship of stress/force and strain/ deformation. Deformation is a change in shape or size secondary to stress and strain on an object from applied forces and moments and is a structural property of a material that depends on the shape, size, and intrinsic material properties. The study of deformation characteristics aids in the understanding of modes of failure. This usually causes a change in the object in both the x and y axis. Elastic deformation occurs when strain on a material is totally recovered when the stress is removed. Plastic deformation occurs at the point where stress is no longer proportional to strain. Yield point is the point at which elastic deformation becomes plastic deformation. Ultimate tensile strength/breaking point is the point at which an object fails. Strength is the maximum stress that a material can sustain and coincides with the area under the stress/strain curve to the point of its ultimate tensile strength. Intrinsic material properties are independent of an object’s shape and size, thus their study requires that the effect of the object’s shape and size (geometry) are eliminated. Ductile – materials with intrinsic properties that allow permanent deformation before failure. Brittle – materials with intrinsic properties that cause failure before permanent deformation. Hooke’s law – the degree of elastic deformation of a solid object is proportional to the deforming force and the elastic modulus is a measure of “deformability” of a solid object.

*Correspondence to: Edward C. Benzel, Department of Neurosurgery, Neurological Institute, Cleveland Clinic, 9500 Euclid Avenue, Desk S40, Cleveland, Ohio, OH 44195, USA, E-mail: [email protected]

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Isotropic objects have intrinsic material properties independent of the direction of loading and a randomly dispersed internal structure (metal, glass, plastic). Anisotropic objects have intrinsic material properties dependent of the direction of loading and an orderly internal structural arrangement (bone, intervertebral discs, and ligaments/tendons).

Stability and kinematics Spinal stability is one of the most fundamental concepts required to characterize and evaluate the spinal column and is critical for proper function. In biological systems, stability is circumstance-dependent, rather than being an all-or-none phenomenon. It must be defined both for static conditions, in which the system is in equilibrium, and for dynamic situations, in which the system is moving along some trajectory. Whether the system is in equilibrium (static) or changing with time (dynamic), stability may be assessed by the presence or absence of novel behavior secondary to small perturbations of force acting on the system. A sign of stability is if the new behavior is approximately the same as the old or if the changed behavior becomes indistinguishable from the old behavior after a period of time. A sign of instability would be if the disturbed behavior were to differ significantly from the old behavior. With this in mind, the classic definition of clinical stability is: “the ability of the spine under physiological loads to limit patterns of displacement so as not to damage or irritate the spinal cord or nerve roots and, in addition, to prevent incapacitating deformity or pain caused by structural changes” (White and Panjabi, 1990). Stability is maintained by three mechanisms: (1) the active subsystem (musculoskeletal system); (2) the passive subsystem (the spinal column); (3) the neural system (activation of the active system through neurological control). Under normal conditions, the three subsystems maintain mechanical stability while the spinal column translates and rotates about the three cardinal anatomical axes (Fig. 2.1). This provides six potential movements referred to as degrees of motion. Kinematics is the study of the motion of bodies. Segmental motions at the various spinal levels are generally determined by facet orientation, bony anatomy, associated soft tissue support (muscles and ligaments), and supporting structures such as the rib cage. This complex interrelationship can be simplified by dividing the spinal column into smaller units known as functional spinal units (FSU) and multilevel spinal units (MSU). Several concepts have been effective in characterizing the complex nonlinear load–displacement relationship between spinal units. When a force is applied to a FSU or MSU, the unit will displace from a neutral

Fig. 2.1. The Cartesian coordinate system with the instantaneous axis of rotation (IAR) as the center. Translation and rotation can occur in both of their respective directions about each axis. (From Benzel (2001), with permission.)

Fig. 2.2. A typical load/deformation curve depicting the neutral and elastic zones (deformation or strain versus load or stress). (From Benzel (2001), with permission.)

position to a position where an appreciable resistance is first encountered (Fig. 2.2). This initial region of “laxity” is termed the neutral zone (NZ), and allows the spine to undergo relatively large motions with very little muscular effort. If this area begins to increase in

ANATOMY AND BIOMECHANICS OF THE SPINAL COLUMN AND CORD size (increased laxity), it could represent decreased stability. When the maximum strain capacity of the NZ is reached, the tissues are then deformed according to Hooke’s law. This law states that for small displacements, the size of deformation is proportional to the deforming force. This region is known as the elastic zone (EZ). Once the elastic limit is reached, any further stress application results in permanent deformation, which is known as the plastic zone. Finally, range of motion (ROM) is the displacement at the largest applied load or at the limit of motion for an activity. Several other concepts are important when discussing spinal kinematics and stability. The motion pattern describes the displacement path a vertebral body follows under load. When this pattern begins to deviate from its historical norm, it could be a sign of instability. Motion about or along axes secondary to those of the axis of applied load is known as coupling. This is seen in the cervical spine when lateral bending produces a concomitant axial torsion due to the orientation of the articulating surfaces of the facets. These motions can also change if the spine begins to become unstable. The axis about which a vertebral body rotates at some instant of time is known as the instantaneous axis of rotation (IAR). While in normally functioning spinal units the IAR is confined to a relatively small area somewhere within the spinal unit, it can shift outside of the physical space of the unit and noticeably enlarge if the area becomes unstable.

BIOMECHANICS OF SPINAL COLUMN FAILURE Relevant osteoligamentous anatomy Stability of the osteoligamentous spinal column is maintained by interdependent systems composed of discrete bony elements (vertebrae) separated by intervertebral discs and articulating joints, which are joined together by passive ligamentous restraints and dynamically controlled muscular activation. A functional spinal unit (FSU) is the basic unit of study of the spine and consists of a superior vertebra-intervertebral disc-inferior vertebra osteoligamentous unit.

OSSEOUS

STRUCTURES

A typical vertebral body consists of an anterior cylindrical vertebral body (or centrum) and a dorsal segment (vertebral arch/neural arch). The anterior vertebral body is the main axial load-bearing structure of the spine and is primarily composed of cancellous bone (an anisotropic viscoelastic material) encased peripherally by an outer shell of cortical bone and rostrocaudally by end plates of compacted cancellous (or trabecular) bone. For a wide

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range of strain rates, cancellous bone behaves elastically and the elastic moduli and strength of cancellous bone is dependent on its density to the second power. The width and depth of vertebral bodies increase as one descends in the spine due to increasing axial loads. Primarily owing to this increase in size, the absolute failure load in normal adults increases from the cervical down to the lumbar spine. While the uncovertebral joints are poor at axial load resistance, they are ideal at regulating extension and lateral bending motion and torsion resistance. The vertebral arch begins bilaterally with the pedicles whose axes are oriented anteroposteriorly and mediolaterally. These structures form a junction with the laminae, which extend around the spinal canal, the superior and inferior facets, and the transverse processes. The articulating facets limit motion, transmit direct compressive forces, and bear compressive forces from bending and rotation. The thoracic section of vertebrae also possesses costovertebral (rib-vertebra) facets anterior to the transverse processes.

INTERVERTEBRAL

DISCS

The intervertebral discs comprise the firm, structural annulus fibrosus and the softer, pliable, shock absorbing nucleus pulposus. The normal nucleus pulposus is located posterocentral in the disc where it can take up 30–50% of the cross-sectional area of the disc. As would be expected, the water content of the nucleus decreases as it degenerates. For example, the water content in the normal nucleus of lumbar discs decreases from about 90% of its total volume during the 1st year of life to 74% in the 80th year. The annulus fibrosus is designed more for structural support and is composed of concentric layers of collagen fiber bundles wound in a helicoid manner. This arrangement results in equally distributed forces within the disc from concentric axial loads. Eccentrically placed loads create bulging of the annulus on the side of the applied force, with associated displacement of the nucleus to the opposite side. The annular fibers’ orientation alternates from layer to layer, with the fibers generally oriented at an angle of approximately
30 with respect to the horizontal plane and in any two adjacent layers at 120 with respect to each other. This improves resistance to shearing and rotational forces. The fibers in the inner third of the annulus interconnect with the cartilaginous end plate and the fibers in the outer portion are firmly bonded to the epiphyseal ring of the vertebral body.

FACET

JOINTS

The facet joints, like the intervertebral disc, provide articulation between segmental levels. Orientation of these joints serves to facilitate or limit degrees of

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motion. For example, because the cervical facets are coronally oriented, they resist translation and facilitate flexion, extension, and rotation. On the other hand, the facet joints in the lumbar spine are sagittally oriented, with the exception of L5–S1, resisting rotation and allowing flexion and extension. The thoracic facets are intermediately oriented and provide a middle of the road restriction in translation and rotation.

LIGAMENTS Excluding the upper cervical spine, a FSU is connected by numerous viscoelastic ligaments with nonlinear elastic responses. They include: the intertransverse ligaments (ITL) and interspinous ligaments (ISL), which attach to the transverse and spinous processes, respectively, of adjacent vertebrae; the supraspinous ligament (SSL), which originates as the ligamentum nuchae (LN) of the neck, extends the length of the spine posterior to the ISL, and attaches firmly to the tip of each spinous process; the capsular ligaments (CL) that surround each facet joint; and the ligamentum flavum (LF), which originates bilaterally on the anteroinferior aspect of the lamina of the superior vertebral body and inserts on the posterosuperior aspect of the lamina of the inferior vertebra. Ligaments are considered passive stabilizers of the spinal column, providing both tension-band and translational support. The tension band support is the result of the ligament’s tensile strength and the moment arm through which it acts. As discussed previously, the moment arm is the perpendicular distance from the instantaneous axis of rotation (IAR) to the applied force vector and the amount of resistance (counter-bending moment) a ligament provides is proportional to its distance from the IAR. The vertebral bodies are stabilized anteriorly by the anterior longitudinal ligament (ALL) and posteriorly by the posterior longitudinal ligament (PLL). The ALL originates at the base of the occiput and extends the entire length of the spine into the sacral region along the anterior aspect of the spine. It is attached to the vertebral body edges at each segmental level. In addition to the strong physical characteristics of the ALL, its position ventral to the IAR provides a moment arm that resists extension. The PLL also extends the length of the spine along the posterior aspect of each vertebral body. It has far less biomechanical strength than the ALL. This is primarily due to the position of the PLL dorsal to the IAR, which provides a short moment arm and, in combination with its weak intrinsic mechanical properties, far less resistance to flexion than the dorsal elements.

MUSCLES The spinal musculature may be divided into five major classifications based on location. These include: the

posterior wall musculature (erector spinae of paravertebral muscles), the respiratory or intercostal muscles, the abdominal wall muscles (intertransversus, interior and exterior obliques, rectus abdominis), the superficial trunk musculature (rhomboids, latissimus dorsi, pectoralis, and trapezius), and the lower trunk musculature (transversus abdominis). With the exception of the erector spinae muscles, the primary function of the muscles immediately surrounding the spinal column attached to the vertebrae are to stabilize the spinal column, rather than to affect motion.

RIB

CAGE

The rib cage adds a significant amount of stability to the upper and middle thoracic segments. This stability is achieved via the costovertebral and costosternal joints.

Spinal instability CLASSIFICATION Instability is referred to as either acute or chronic and is rarely an all-or-none phenomenon, more commonly occurring on a spectrum ranging from stable to grossly unstable. Acute instability is often the result of trauma and can be further subdivided into either overt or limited. Overt instability is the inability of the spine to support the torso during normal activity. For such instability to occur, a loss of vertebral body or disc integrity must be combined with a loss of integrity of the dorsal elements, resulting in a circumferential loss of spinal integrity. Fractures involving both ventral and dorsal columns should be considered overtly unstable. Overt instability is synonymous with gross instability and commonly requires surgical stabilization. Limited instability is defined as the loss of either ventral or dorsal spinal integrity, with the preservation of the other. Examples would include isolated laminar fractures or ligamentous disruption with intact ventral elements. Such instability is sufficient to support most normal activities. Chronic instability may be subdivided into glacial instability (in which the deformity progresses slowly, like the motion of a glacier) and/or dysfunctional segment motion. In the latter, there is no progression of deformity, but rather a pain syndrome generated by “dysfunctional motion.”

Specific failure mechanisms The area of structural failure of the spinal column is dependent on the orientation and magnitude of force applied and the structure’s vulnerable areas. Vulnerability is a function of the material properties of the tissue composing the area as well as the way the tissues interact with each other. For example, the dorsal components of

ANATOMY AND BIOMECHANICS OF THE SPINAL COLUMN AND CORD the spinal column are very vulnerable to compressive forces and are best at resisting and transmitting tensile forces. Conversely, the ventral components are most capable of resisting compression. In combination, the column is best able to resist forces that result in bending in ventral flexion, which both loads the dorsal components in tension and the ventral components in compression. Its strength is more limited in resisting flexure in lateral bending, extension, and torsion. The configuration and mechanism of a spinal column fracture can be predicted by understanding the magnitude and direction of the force vector in relationship to the IAR. The extent of the fracture would depend on the magnitude of the bending moment, which is proportional to both the magnitude and the perpendicular distance of force application in relation to the IAR (moment arm). If the IAR is altered, as in a kyphotic deformity, the bending moment will be significantly affected. The actual amount of bone disruption can be predicted by the stress/strain curve (Fig. 2.3). Most of the initial strain is dissipated through the ligaments and disc, which is the neutral zone (NZ). When the maximum strain capacity of the NZ is reached, the tissues are then deformed according to Hooke’s law. This is known as the elastic zone (EZ). The size of the elastic zone is dependent on the elastic modulus of each specific tissue, and accordingly is greater for ligaments than bone. Once the elastic limit is reached, any further stress application results in permanent deformation, which is known as the plastic zone (PZ). If damage occurs at the upper limit of the EZ, the segment is left in a state of relative laxity, with an expanded NZ. This increase in the NZ is synonymous with segmental instability.

Fig. 2.3. A typical stress/strain curve for a biological tissue, such as a ligament. AB, the neutral zone. BC, the elastic zone. When the elastic limit (yield point) (C) is reached, permanent deformation can occur (permanent set). CD, the plastic zone where a permanent set occurs. Past D, failure occurs and the load diminishes. Hashed plus dotted area represents strength, whereas the dotted area represents resilience. (From Benzel (2001), with permission.)

FLEXION–COMPRESSION/VENTRAL

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WEDGE

COMPRESSION FRACTURES

A compressive force vector placed ventral to the IAR can result in forced flexion with axial loading. The resulting bending moment generates compression of the vertebral elements ventral to the IAR with preservation, and sometimes distraction, of the dorsal elements (Fig. 2.4). The severity of the resultant injury is defined by both the degree of anterior vertebral body damage and posterior ligamentous disruption. Generally, if only the anterior column is injured and the vertebral body height loss is less than 50%, the fracture is usually stable. If a loss of vertebral body height greater than 50% occurs there is an increased risk of instability. Obviously, if all three columns fail the fracture is unstable. These injuries are often called anterior teardrop or quadrangular fractures. The teardrop fracture is an injury in which the severe forward-bending forces fracture off a triangular piece of the anterior lip of the rostral vertebral body, often with retrolisthesis of the remaining body into the central canal. Similarly, the quadrangular fracture is an injury in which a large piece of anterior vertebral lip is broken off. It is associated with retrolisthesis, kyphosis, and circumferential soft-tissue disruption. Areas of the spine that are in a naturally kyphotic posture, such as thoracic and thoracolumbar regions, are biomechanically disadvantaged by a ventral gravitational bending moment and are predisposed to this type of fracture.

Fig. 2.4. A depiction of the injury force vector causing a ventral wedge compression fracture. F: applied force vector; D: length of moment arm (from IAR to plane of F); M: bending moment. (From Benzel (2001), with permission.)

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AXIAL

V.J. MIELE ET AL. COMPRESSION (BURST) FRACTURES

This pattern of injury results from an axial loading force with no associated eccentric load outside of the IAR – in other words, no bending moment (Fig. 2.5). If there is an eccentric load, an angular deformity such as a ventral or lateral compression wedge fracture occurs. Burst fractures are most commonly observed in the upper and middle cervical and lumbar spine, since the IAR is located along the anatomic plumb-line. Although this type of fracture disrupts both the anterior and middle columns of Denis (Denis, 1983), which would imply instability, most burst fractures are not overtly unstable because the dorsal elements are often preserved.

FLEXION–DISTRACTION (CHANCE)

FRACTURES

Flexion–distraction fractures result from a force vector directed both ventrally and rostrally (Fig. 2.6). Such a load is most commonly experienced in deceleration injuries, where the patient is restrained by a single lap belt. This type of injury causes failure of the posterior column with damage to ligamentous components, bony components, or both. The pathophysiology of this injury pattern

Fig. 2.6. There are two fundamental types of Chance (flexion–distraction) fracture. (A) Diastasis fracture through the pedicles and vertebral body. (B) Fracture through the vertebral end plate or disc. (C) The mechanism of injury is depicted. (From Benzel (2001), with permission.)

is dependent on the axis of flexion, with the classic fracture having its axis of flexion anterior to the anterior longitudinal ligament. Several subtypes exist, and each is dependent on the axis of flexion and on the number and degree of column failure. The cervical and upper thoracic spinal columns are the most commonly involved, and the fracture commonly occurs through a bony cleavage plane or the vertebral end plate. The force vector may also be directed dorsally in distraction, as observed in hyperextension-shear injuries. Since this type of fracture involves both the ventral and dorsal columns, they are considered overtly unstable. Injuries that result in facet dislocations are caused by flexion and distraction forces, with or without an element of rotation. The facets may be fractured, subluxed, or dislocated (“locked”), either unilaterally or bilaterally.

DORSAL ELEMENT

Fig. 2.5. The mechanism of injury of a burst fracture: true axial loading without a bending moment. (From Benzel (2001), with permission.)

FRACTURE

This type of a fracture occurs when a load is applied dorsal to the IAR. A common mechanism is hyperextension with axial loading. Dorsal element fractures are most commonly observed in the cervical spine and result in laminar, spinous process, and/or facet fractures.

ANATOMY AND BIOMECHANICS OF THE SPINAL COLUMN AND CORD

ROTATIONAL

FRACTURE-DISLOCATION MECHANISM

This type of injury commonly results in failure of both the middle and posterior columns with varying degrees of anterior column damage. It is caused by a combination of lateral flexion and rotation with or without a component of posterior-anteriorly directed force. The rotational force component results in disruption of the posterior ligaments/articular facet. If the rotational force is sufficient, it can actually rotate the upper vertebral body with the superior portion of the lower vertebral body attached, resulting in a three-column failure. The flexion–rotation injury pattern results in failure of both the middle and posterior columns along with compression of the anterior column.

Regional biomechanics of the spinal column The spinal column can be divided into three mobile (the cervical spine, the thoracic spine, and the lumbar spine) and two fused regions (the sacrum and the coccyx). The mobile segments can undergo axial, lateral, and sagittal rotations and axial, lateral, and anteroposterior translations. Thus, the spine is said to possess six degrees of freedom (DOF). In normal conditions, the spine has evolved to adopt a curvilinear sagittal conformation – with a primary kyphotic thoracic curve, compensated by secondary cervical and lumbar lordotic curves of equal summative magnitude. Any increase in thoracic kyphosis (or loss of lumbar lordosis) results in an increased moment arm (perpendicular distance from the IAR to the gravitational force vector), generating a greater bending moment at each vertebral segment. Since the moment arm (M) is equal to the force (F) multiplied by its perpendicular distance (D) from the IAR, the greater the deformity, the greater the moment arm length. Each vertebral region has unique anatomical and functional features that predispose them to specific injuries. Likewise, the transition areas between the broad regions, such as the cervicothoracic, thoracolumbar, and lumbosacral junctions, are more vulnerable to injury and degenerative changes due to the abrupt change in “stiffness” that occurs at these junctions.

CERVICAL

REGION

The cervical region is the most mobile of the spinal column with a range of motion of approximately 80–90 of flexion, 70 of extension, 20–45 of lateral flexion, and up to 90 of rotation to both sides (Windle, 1980). When the neck undergoes flexion, it is initiated at the lower cervical spine (C4 through C7), followed by motion at C0 (occiput) through C2, C2 through C3, and then C3 through C4. The C6 through C7 segment undergoes a

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brief reversal of motion into extension, followed by a reversal of motion at C0 through C2. The C6 through C7 segment contributes to the end ranges of flexion (Van Mameren et al., 1990). Similarly, extension is also initiated in the lower cervical spine (C4 through C7) and is followed by the beginning of motion at C0 through C2. The middle range consists of varied movement from the mid cervical region, whereas the lower cervical spine is the last to contribute as the column moves into terminal extension (Van Mameren et al., 1990). Axial loading of the cervical spine often results in a transient deformation, or buckling effect, which produces large angulations within the cervical spine as a means of releasing the additional strain energy that has been produced from the vertical loading (Nightingale et al., 1996a). This buckling effect is often a contributor to injury and has been observed experimentally in two distinct orders. First-order buckling results in extension of the upper cervical spine through C5 and flexion through T1. Second-order buckling creates flexion of C1 through C3, extension in C4 and C5, and flexion in C6 through T1 (Nightingale et al., 1996a, b, 2000). The instantaneous axis of rotation (IAR) plays a large role in the reversal of motion observed in the cervical spine (Amevo et al., 1992; Penning, 1995). The center of rotation is located near the superior aspect of the inferior vertebral body. As force is transmitted down the cervical column, the individual vertebrae experience flexion or extension depending on the location of the force vector relative to the IAR. Thus, if the cervical column is moving into flexion, but the force vector passes behind a specific vertebra’s IAR, then that vertebra will extend. The levels at which the spine reverses its motion are where pivot points have been created.

THORACIC

REGION

The thoracic spinal column is a kyphotic segment that is relatively stiff because of costochondral reinforcement. Adding to the stability of the entire thoracic region is the nearly vertically oriented articulating processes as well as the shingle-like oblique arrangement of the spinal processes. This region is not without vulnerabilities, however, since the vertebral bodies are proportionately smaller than those of the lumbar region, making them more susceptible to compressive forces. The narrow canal and relatively poor blood flow in this area also predisposes the spinal cord to injury should fracture occur. The lower thoracic region, T11 and T12, is considered to be the transition zone between the thoracic and lumbar regions. The vertebrae are attached to floating ribs and are less stable. This area resembles the lumbar region in stability and mechanisms of injury.

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LUMBAR

V.J. MIELE ET AL.

Clinical models

AND LOW THORACIC REGIONS

While the lumbar and low thoracic vertebrae are larger (providing more axial strength), the region is significantly more mobile compared to the thoracic spine. This allows for flexion, extension, and rotation of the upper skeleton in relation to the pelvis and lower extremities. Regional stability is decreased by the loss of rib cage stabilization as well as the fact that the spinous processes are more horizontal, which contributes to increased mobility but also provide less mechanical stability. These factors result in an increase in susceptibility to injury.

JUNCTIONAL

AREAS

Any area that has a significant transition in stiffness and mobility is particularly susceptible to injury. The thoracolumbar and lumbosacral junctions are good examples of this. In fact, the thoracolumbar junction is uniquely susceptible to injury and accounts for over half of all spine injuries outside the cervical spine. The transition from the stiff cephalad thoracic spine to the more mobile lumbar segments creates a stress riser at the thoracolumbar junction, which can act as a long lever arm to produce injury at the mobile junctional segments. Additionally, the transition from sagittal kyphosis to lordosis leaves this region susceptible to axial overload. Injuries in these areas are often the result of hyperflexion producing vertebral body failure. The cervicothoracic region is similarly more susceptible to injury. Because the cervical portion has significant mobility versus the relatively immobile thoracic component, it essentially acts as a cantilever beam with the “fixed end” at the cervicothoracic junction, the location of the highest stresses.

Clinically useful biomechanical models of the spinal column have been developed over the past 50 years to predict stability after an injury as well as aid in identifying the mechanism of the injury. Currently, column theory and finite model analysis are the most commonly used methods of modeling (Fig. 2.7).

COLUMN

THEORIES

A two-column model of the spinal column was first developed in the 1960s based mainly on clinical experience with a large number of spinal injuries (Holdsworth, 1970). This model is based on an anterior and posterior column. The anterior column consists of the anterior longitudinal ligament, the vertebral body, intervertebral disc, annulus fibrosus, and posterior longitudinal ligament. The posterior column includes the pedicles, lamina, spinous processes, facet joint complex, ligamentum flavum, and interspinous and supraspinous ligaments. This model proposes that disruption of the posterior column is necessary for spinal instability, suggesting that compression and burst fractures are stable whereas fracturedislocations are unstable. A three-column model was developed in 1983 based on radiographic studies (Denis, 1983). This model suggests that burst fractures, which were considered in the two-column model to be stable, were often actually unstable. The additional middle column consists of the posterior vertebral body, the posterior annulus fibrosis, and the posterior longitudinal ligament. This model also defines the anterior column as containing the anterior longitudinal ligament, the anterior half of the vertebral body, and the related portion of the intervertebral disc and its annulus fibrosus. The posterior column contains

Fig. 2.7. The “column” concepts of spinal stability. The concept described by Louis (left) assigns significance to the vertebral body and the facet joint complexes (lateral masses) on either side of the dorsal spine. Denis’s three-column concept (right) assigns significance to the region of the neutral axis and the integrity of the posterior vertebral body wall (the middle column). The two column construct (left) relies on anatomically defined structures, the vertebral body (anterior column) and the posterior elements (posterior column). Denis’s three-column concept (right) similarly relies on anatomically defined structures. (From Benzel (2001), with permission.)

ANATOMY AND BIOMECHANICS OF THE SPINAL COLUMN AND CORD the bony elements of the posterior neural arch and the ligamental elements, which include the ligamentum flavum, the interspinous ligaments, the supraspinous ligaments and the joint capsule of the intervertebral articulations. In this model, disruption of two columns is required for instability. Thus, the three-column model considered compression fractures to be stable and burst fractures, Chance fractures, and fracture-dislocations unstable (Panjabi et al., 1995).

FINITE

ELEMENT MODELING

The finite element method is a mathematical modeling system that was originally intended for use in structural engineering. It has now been applied to the spine for over 25 years. The spine is separated into a large number of geometric forms, such as cubes or spheres that are called “elements.” These elements interact with each other at junctions called “nodes” (Goel and Gilbertson, 1995). Transforming the spinal column into a finite number of discrete units then allows the stresses, strains, and forces at any given location to be calculated using a computer.

BIOMECHANICS OF SPINAL CORD INJURY Although biomechanical analysis has traditionally been associated with the osteoligamentous spinal column, a better understanding of the biomechanics of spinal cord injury is important. Knowledge of the mechanics of spinal cord injury and the interactions between different anatomical components during trauma provides valuable insight into the pathophysiology of injury and potential management strategies. The spinal cord is a dynamic structure that undergoes significant geometric changes without negative sequelae during normal physiological movements. This ability to change shape without effecting function can be overwhelmed by traumatic forces resulting in dysfunction. The onset, duration, and intensity of these stresses define the magnitude and potential reversibility of the spinal cord dysfunction.

Relevant spinal cord anatomy The spinal cord is a soft, pliable mass of nerve fibers and cells supported by glial tissue. It is elliptically shaped with enlargements in the cervical and lumbar segments and extends from the base of the brain through the central canal to the level of the L1/L2 vertebrae. The spinal cord contains myelinated “long” tracts and interconnecting fibers (white matter) surrounding mostly unmyelinated nerve fibers with supporting glial tissue and a very intense small blood vessel network (central gray matter). Perfusion pressure of the neural tissue from this blood flow increases the structure’s measured stiffness.

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Within the confines of the bony canal, the spinal cord is invested in the soft-tissue structures of the pia, arachnoid, and dura mater as well as cerebrospinal fluid, epidural fat, and veins. These all contribute to the biomechanical and physical behavior of the spinal cord. The dura mater covering the spinal cord is firmly attached to the base of the skull and the second fused segment of the sacrum in the adult (filum terminale). While the cord inside the dura is fairly mobile, it is held more tightly in the flexed position of the spine and in a more relaxed position when the spine is extended. The pia mater invests the spinal cord intimately and on either side of the cord, a pial thickening and extension between the ventral and dorsal nerve roots form the paired dentate ligaments, which tether it to the dura mater laterally (White and Panjabi, 1990). Caudally, another pial thickening forms the filum terminale, which anchors the conus medullaris to the bony sacrum. Lateral attachment of the spinal cord to both sides of the spinal canal via the dentate ligaments (as well as the exiting nerve roots) provides significant fixation and stabilization within the dura in both static and dynamic states. Likewise, direct mechanical stress on the spinal cord can be produced by the dentate ligaments tethering the spinal cord to the central canal. The tensile force of these ligaments on the spinal cord is applied laterally with a slight rostral-to-caudal orientation, due to the caudal inclination of the dentate ligaments with respect to the spinal cord. The application of this force comprises both a transverse vector directed laterally and an axial vector directed caudally (Breig, 1960). The transverse vectors of the paired ligaments oppose one another and thereby help to maintain the cord in a central position within the canal, maximizing the cushioning effect of the surrounding tissues. The axial vectors aid in balancing the tension and reducing axial stresses in the cord (White and Panjabi, 1990). Under conditions of flexion, the tension in these ligaments increases, further stabilizing the spinal cord in the center of the canal (Breig, 1960). These forces, while stabilizing under normal physiological stresses, can contribute to increased tension in pathological conditions (Kahn, 1947; Cusick et al., 1977).

Normal physiological biomechanics of the spinal cord The spinal canal undergoes significant geometric changes with physiological movement that the spinal cord and its surrounding tissues must adapt to. The dural sheath shifts rostrally and stretches with spinal canal lengthening (Adams and Logue, 1971). Significant changes in the size of the ventral and dorsal subarachnoid spaces are also observed with physiological motion of the spinal canal. Studies have demonstrated up to a 43% reduction of the

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ventral subarachnoid space diameter and up to 89% increase in the dorsal subarachnoid space diameter with flexion, as well as up to 9% increase in the ventral subarachnoid space and up to 17% reduction in the dorsal subarachnoid space with extension (Muhle et al., 1998). While biological tissues generally have a nonlinear stress–strain response (van Noort et al., 1981; Bilston and Thibault, 1996; Nightingale et al., 1996a), the spinal cord is viscoelastic and behaves as a linear elastic material under small strains (strain ¼ force/area) (Kakulas, 1984). Rapid viscoelastic relaxation following sustained application of a strain has been observed experimentally with greater than 50% reductions in the stress within the spinal cord at 5 minutes following application of maximum strain (Carlson et al., 1997, 2003a; Ichihara et al., 2003). The distractive load–displacement curve of the spinal cord is biphasic, with an initial phase in which large displacements of more than 10% of its original length occur with minimal force (0.01 N), followed by an abrupt shift to the second phase in which large forces (20–30 N) generate small displacements before mechanical failure (Breig, 1960; White and Panjabi, 1990). These phases are similar to the neutral and elastic zones more commonly discussed relating to osteoligamentous structures and allow the cord to change shape significantly during normal physiological activity. The spinal cord demonstrates similar behavior in compression, with initial large displacements occurring with minimal forces, followed by increasing resistance with smaller associated displacements until the cord buckles (White and Panjabi, 1990). The initial phase of large displacements with small forces has been described as the folding and unfolding of the structure of the cord similar to an accordion, while the subsequent phase of small displacements with large forces represents the cord tissue response to tensile forces (White and Panjabi, 1990). This description has led to the initial phase being named the folding/unfolding zone and the subsequent phase the elastic deformation zone (Breig, 1960). The folding/unfolding zone of the spinal cord accomplishes 70–75% of the length changes between flexion and extension, with the remainder of the length changes occurring in the elastic deformation zone (White and Panjabi, 1990). When a material is stretched in one direction and tends to contract (or occasionally, expand) in the other two directions perpendicular to the direction of stretch it is termed Poisson’s effect. The spinal cord exhibits this phenomenon: an increase in the cross-sectional area of the spinal cord occurs with a decrease in its length and conversely, a decrease in the cross-sectional area of the spinal cord occurs with an increase in length (Panjabi and White, 1988; Muhle et al., 1998). This is primarily the result of the incompressibility of the spinal cord tissue, which plays an important role in pathological states

where ventral or dorsal lesions in the spinal canal compromise the ability of the cord to adapt to such changes and thereby introduce abnormal stresses on the cord. The mechanical differences between the white and gray matter of the spinal cord have been studied with mixed results. It was originally theorized that gray matter was equally or less rigid than white matter (Kahn, 1947; Panjabi and White, 1988; Levine, 1997). However, a study of bovine spinal cords demonstrated significantly higher stress and Young’s moduli in gray matter compared with white matter in the linear portion of the stress–strain curve, which supports the hypothesis that gray matter is more rigid than white matter (Ichihara et al., 2001). The study also found that the gray matter failed at lower strains than the white matter, suggesting that gray matter was also more fragile. These findings were felt to correlate with the sensitivity of the cervical gray matter (i.e., central cord syndrome) following mechanical stress to the spinal cord. Other studies have demonstrated no significant differences in the modulus of elasticity between gray and white matter (Ozawa et al., 2001).

Biomechanics of spinal cord injury A variety of loads and stresses are applied to the spinal cord during trauma, including direct compression load, shear load, tensile load, and bending loads (Panjabi and White, 1988). These loads commonly occur in combination. When a spinal cord injury occurs, the spinal cord undergoes major transient geometric changes that overcome the tissues’ ability to adapt. Direct compression of neural tissue results in a focal application of force. This decreases in magnitude as one travels away from the point of contact (Panjabi and White, 1988). Shear force application results from the displacement of spinal cord segments adjacent to sites of direct compression. It has a zero force value at the site of load application which then increases toward the center of the spinal cord, with a maximal force value in the center of the cord (Raynor and Koplik, 1985; Panjabi and White, 1988). Tensile load is applied uniformly across the cross-section of the segment being stretched. Stretchassociated injury is now widely accepted as the principal etiological factor of myelopathy in experimental models of neural injury, tethered cord syndrome, and diffuse axonal injury. Axonal injury reproducibly occurs at sites of maximal tensile loading in a well-defined sequence of intracellular events: myelin stretch injury, altered axolemmal permeability, calcium entry, cytoskeletal collapse, compaction of neurofilaments and microtubules, disruption of anterograde axonal transport, accumulation of organelles, axon retraction bulb formation, and secondary axotomy (Henderson et al., 2005). Bending loads

ANATOMY AND BIOMECHANICS OF THE SPINAL COLUMN AND CORD introduce differential stresses on the cross-section of the segment being bent, with tensile stress on the convex side of the segment and compressive stress on the concave side of the segment. The maximum stress occurs at the points farthest from the IAR. Injury results from mechanical deformation above the physiological limits of the cord. This can result from compression, torsion, and tension of the spinal cord. As mentioned previously, the distractive load–displacement curve of the spinal cord is biphasic, with an initial phase in which large displacements of more than 10% of its original length occur with minimal force (0.01 N), followed by an abrupt shift to a subsequent phase in which large forces (20–30 N) generate small displacements before mechanical failure (Breig, 1960; White and Panjabi, 1990). In compressive injuries causing less than 1 mm deformation of the cord, it commonly behaves like a spring, with a linear relation between the applied force and the resultant displacement (Somerson and Stokes, 1987). Larger displacements (greater than 1 mm) result in nonlinear characteristics. Various models of spinal cord injury have shown differences in cord damage for various column injury patterns (Fiford et al., 2004; Choo et al., 2007). Three commonly encountered injury patterns/mechanisms include transverse contusion (as would occur in a burst fracture), distraction (as would occur in a column distortion or distraction injury), and dislocation (as would occur in a fracture dislocation). Spinal cord injury models have demonstrated focal strains in contusion and dislocation, while those in distraction were more uniformly distributed throughout the cord (Greaves et al., 2008). The distraction injury mechanism has also been associated with an increased caudal–cranial extent of injury, as compared to contusion (Choo et al., 2007). It also commonly results in the greatest strains in the dorsal column, while the ventral column experienced the least strain. The dislocation mechanism is associated with compressive lateral strains and increased strains in the lateral columns, as compared to contusion (Choo et al., 2007). While damage to the spinal cord primarily occurs from direct injury to the spinal cord neural and supportive glial tissue, it is also the result of alterations in vascular physiology and metabolic derangements (Hung et al., 1981; Torg et al., 1995; Yamada et al., 1995; Carlson et al., 1997, 2003b; Harrison et al., 1999). When the spinal cord undergoes deformation, the axonal membrane is subjected to varying degrees of local stretch damage. Clinical outcome and morphometric characteristics of the spinal cord lesion vary depending on several factors, including force, duration of compression, displacement, impulse, and kinetic energy (Rivlin and Tator, 1978; Dolan et al., 1980; Hung et al., 1981; Guha et al., 1987;

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Panjabi, 1987). The tolerance of neural tissue and the individual cellular components may be dependent on several variables, including spinal region, cellular orientation, and extracellular matrix. Areas with less blood flow can be more prone to injury since they would be more prone to local tissue ischemia. Prolonged compressive stresses on the cord tend to produce more extensive tissue damage and poorer recovery of neurological function compared with shorter periods of compression (Carlson et al., 2003b). The time course of the application of stresses on the spinal cord can also be a determinant of extent of injury. In static or gradually applied loading, much smaller forces are required to generate similar displacements and trauma than in the application of an acute impact force (Hung et al., 1982). Such chronic displacements of the spinal cord appear to be well tolerated, and the function of the cord appears maintained (Ichihara et al., 2003).

Spinal cord injury models Several experimental models have been developed that mimic spinal cord injury. Those that rely on compressive forces to duplicate injuries can be classified as either kinetic or static, according to the biomechanics of the applied forces. Models that involve rapid compression of the cord in less than 1 second are known as kinetic compression models. In practice, most kinetic models compression occurs in less than 100 milliseconds. Static compression models, on the other hand, focus on forces that slowly compress the cord at approximately constant velocity. Kinetic compression models most closely replicate traumatic human spinal cord injuries. They often utilize extradural balloons or clips for compression. Clip application models have the advantage that the force of clip closure can be calibrated precisely and the duration of the compression can be altered over a wide range of times. Static models, which involve a gradual compression of the spinal cord, are useful to model the effect of spinal cord displacement, as well as strain and duration of compression. However, because the load is applied slowly to the cord, these models do not accurately simulate the biomechanical aspects of the majority of spinal cord injuries. It has been demonstrated by using the extradural balloon model that the duration of compression was found to be a significant determinant of neurological recovery (Kobrine et al., 1979). Models using clip compression have shown that clinical recovery varies exponentially according to the force of injury and linearly according to the duration of compression. These studies imply that while the major determinant of recovery is the initial force of injury, the duration of compression is also a

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significant determinant of clinical recovery, even in cases of severe injury forces (Rivlin and Tator, 1977; Dolan et al., 1980; Guha et al., 1987). Since damage to neural tissue is related to peak strain, a major disadvantage of the locally applied force models previously discussed is that they do not measure local strain (Galbraith et al., 1993; Bain and Meaney, 2000). The need to study local strain has been met through the use of the computational modeling techniques of finite element modeling. By virtue of its ability to predict strain distributions through complex structures, this modeling technique may be ideally suited to studying damage of the spinal cord. It has a long track record in engineering applications, has the ability to predict local deformation, and has been used extensively in traumatic brain injury research (King et al., 1995; Ueno et al., 1995; Zhang et al., 2001).

CONCLUSIONS Spine biomechanics is a complex field. Its principals are founded in the study of physics. Once one understands the basics, the application to the clinical arena is straightforward. Both spinal column and spinal cord injury and failure are associated with biomechanical alterations – and hence, can be studied. The strategies for such study have been outlined in this chapter. Much is yet to be learned in this arena. If we, as scientists, maintain a focus on the principles, our knowledge will continue to increase and to enhance our ability to optimally care for patients.

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