Concepts and Mechanisms of Spinal Biomechanics

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CHAPTER 265 Concepts and Mechanisms of Spinal Biomechanics

Vincent J. Miele, Tarun Bhalla, G. Alexander Jones, Edward C. Benzel

Over the past decade, increased understanding of the spinal column has rewarded the neurosurgeon with a plethora of new techniques and instrumentation. Along with these advances, however, comes the need to understand the principles of the various methods of spine stabilization available so that decision making is appropriate in choosing which of these tools to use. Essential to this process is a solid understanding of the concepts and mechanisms of spinal biomechanics. This chapter attempts to provide a foundation for this subject and is presented in three sections. The first reviews the relevant anatomy and basic biomechanical principles. The effects of the aging process and spinal pathophysiology on spinal biomechanics are the focus of the second section. Section three focuses on surgical planning and includes a discussion of newer technologies such as motion preservation devices.

Basic Biomechanical Principles

Biomechanically Relevant Anatomy

Spinal stability is maintained by a variety of anatomic structures that have evolved to provide resistance against deforming forces. These structures include both osseous and ligamentous components.

The vertebral body (VB) is the main axial load–bearing structure of the spine. Its cylindric shape, bounded peripherally by cortical bone and rostrocaudally by end plates, confers it with superior biomechanical properties. The width and depth of VBs increase as one descends in the spine to accommodate increased axial load (Figs. 265-1 and 265-2). The relative weakness of the L5 vertebra can be explained by the asymmetry in height between the ventral and dorsal cortical walls.

FIGURE 265-1 Vertebral body diameter versus spinal level. The width (blue line) and depth (red line) of the vertebral bodies are depicted separately.

(From Benzel EC. Biomechanically relevant anatomy and material properties of the spine and associated elements. In: Benzel EC, ed. Biomechanics of Spine Stabilization. Rolling Meadows, IL: AANS Publications; and New York: Thieme; 2001:1. Reprinted by permission.)

FIGURE 265-2 Vertebral compression strength versus spinal level.

The intervertebral disk serves as a shock absorber and is the primary stabilizing structure of the motion segment. It is composed of the nucleus pulposus (a hydrated core of proteoglycans suspended in a loose collagen network) located in the posterocentral area of the disk and the annulus fibrosus (a fibrocartilaginous ring designed to provide structural support). The loads that these structures must tolerate cyclically for decades is impressive. Compressive loads on the lumbar intervertebral disks are 1.0 to 2.5 times body weight during normal walking. During the lifting of 14- to 27-kg objects, axial compressive loads in the lumbar spine increase up to nearly 10 times body weight, with anteroposterior shear loads approaching double body weight.^1,² Concentric axial loads cause equally distributed forces within the disk, whereas eccentrically placed loads result in bulging of the annulus on the side of the applied force along with associated displacement of the nucleus to the opposite side (Fig. 265-3). Shearing and rotational forces are resisted by the annular fibers, which lie at a 30-degree angle with respect to each other. As the disk deteriorates, its isotropic load transfer properties are lost and load transfer becomes concentrated at the periphery (annular insertion) of the vertebral end plates.^3,⁴

FIGURE 265-3 The intervertebral disk. The fibers of the annulus fibrosus are oriented radially and in opposite directions throughout several layers. A, The nucleus pulposus (dashed outline) is contained (in nonpathologic situations) by the annulus fibers. B, Bearing of an axial load results in an even distribution of force. C, An eccentrically borne load results in bulging of the annulus fibrosus on the concave side of the resultant spinal curve and tension on annulus fibrosus on the convex side of the curve. The nucleus pulposus, however, tends to move in the opposite direction as the annulus fibrosus bulge when an eccentric load is borne (solid to dashed outline) (D).

In conjunction with the intervertebral disk, the facet joints provide additional load-bearing and stabilizing functions between segmental levels. Their orientation (Fig. 265-4) serves to facilitate or limit degrees of motion (Fig. 265-5) and therefore plays an important role in spinal stability. The cervical facets are coronally oriented and resist translation while facilitating flexion, extension, and rotation. Conversely, the lumbar facets are sagittally oriented (with the exception of L5-S1) and resist rotation while allowing significant flexion and extension. The thoracic facets are intermediately oriented and thus provide an “intermediate” restriction of translation and rotation. Both extension and ventral translation tend to load facets, whereas flexion and dorsal translation unload them. Degeneration of the intervertebral disk, loss of disk height, and alterations in sagittal alignment result in greater load transfer to the facet joints.

FIGURE 265-4 Facet joint orientation: relative coronal plane orientation in the cervical region (A), intermediate orientation in the thoracic region (B), and relative sagittal orientation in the lumbar region (C). The facet joint orientation changes substantially in the lumbar region; here, the facet joint angle (with respect to the midline) is depicted versus the spinal level (D).

FIGURE 265-5 Segmental motions allowed at the various spinal levels (combined flexion and extension, pink line; unilateral lateral bending, red line; and unilateral axial rotation, gray line).

The spinal ligaments provide passive stabilization of the vertebral column. Their bone-to-bone interface and elastic properties provide both tension band and translational support. The tension band contribution to spinal stability is related to both the ligament’s tensile strength (Fig. 265-6) and the moment arm through which it acts. As discussed later, the moment arm is the perpendicular distance from the instantaneous axis of rotation (IAR) to the applied force vector. The amount of resistance (counterbending moment) that a ligament provides is proportional to its distance from the IAR (Fig. 265-7).

FIGURE 265-6 Failure strength of spinal ligaments versus the spinal region. ALL, anterior longitudinal ligament; CL, capsular ligament; ISL, interspinous ligament; LF, ligamentum flavum; PLL, posterior longitudinal ligament.

FIGURE 265-7 A, Relative lever arm (moment arm) length of ligaments causing flexion (or resisting extension). B, The ligaments and their effective moment arms. Note that this length depends on the location of the instantaneous axis of rotation (IAR). An average” location is used in this illustration. ALL, anterior longitudinal ligament; black dot, IAR; CL, capsular ligament; ISL, interspinous ligament; LF, ligamentum flavum; PLL, posterior longitudinal ligament.

As opposed to the skeletal muscles, which generate long bone motion and span one or two articulations, the paraspinous musculature (and associated abdominal musculature) spans multiple segments (Fig. 265-8). The primary function of the paraspinous musculature is to stabilize the spinal column rather than produce motion. An exception is the action of the erector spinae muscles when arising from a forward flexed position. In general, any imbalance in muscular forces causes movement about an axis. Conversely, a balancing of muscle and other intrinsic forces about an axis results in no net movement. The ventral abdominal musculature is critical in counterbalancing the erector spinae muscles to provide stability.

FIGURE 265-8 Effects of muscles on stability. A, Muscles provide stability by virtue of the orientation of their attachments to the spine. B, In some situations, as with the rectus abdominis muscle, the muscle may influence spinal movement indirectly (i.e., without direct attachment to the spine). Similarly, this muscle (as well as others) may stabilize the spine by balancing opposing muscle function, thereby resulting in no movement. A significant degree of stability is thus provided. Lateral bending is achieved by the contraction of muscles attached to the lateral aspect of the spine; for example, the quadratus lumborum, attached to the right and left muscle actions, likewise results in no movement.

(From Benzel EC. Stability and instability of the spine. In: Benzel EC, ed. Biomechanics of Spine Stabilization. Rolling Meadows, IL: AANS Publications; and New York: Thieme; 2001:40. Reprinted by permission.)

The rib cage, acting as a barrel attached to the spine, adds significant stability to the upper and middle thoracic segments. Both the costovertebral and costosternal joints are essential to this contribution (Fig. 265-9).

FIGURE 265-9 Illustration of the stability provided to the spine by the rib cage. A, A spine without a rib cage can bend excessively. B, Addition of the rib cage moderately increases stability. C, Sternal attachments are required for achievement of the full stabilization potential of the rib cage. Removal of the effects of either the sternum or the ribs causes a significant diminution in stability.

Basic Biomechanical Principles

Biomechanical analysis assesses the effects of energy and forces on biologic systems by using physics and physical principles that have well-accepted definitions. A reference for many of these definitions can be found in Table 265-1.^5,⁶

TABLE 265-1 Basic Biomechanics Definitions

Kinematics	Study of the motion of objects without considering the factors that cause or affect the motion. The latter is the subject of dynamics.
Momentum	Product of mass and velocity
Moment	Circular force creating a rotational vector around an axis
Torque/bending moment	Product of the force applied to the lever arm multiplied by the perpendicular distance from the axis
Coupling	When more than one non-collinear force acts about the same axis and the resultant force moment is the sum of the individual forces
Stress	Force/load applied to an object divided by its cross-sectional area
Strain	Change in length of an object secondary to a deforming force
Stress/strain	Aids in defining an object’s intrinsic material properties
Stiffness	Relationship of stress or force and strain or deformation
Deformation	Change in shape or size secondary to stress and strain on an object from applied forces and moments. It is a structural property of a material that depends on the shape, size, and intrinsic material properties.
Elastic deformation	Occurs when strain on a material is totally recovered once the stress is removed
Plastic deformation	Occurs at the point where stress is no longer proportional to strain
Yield point	Point at which elastic deformation becomes plastic deformation
Ultimate tensile strength/breaking point	Point at which an object fails
Strength	Maximum stress that a material can sustain—coincides with the area under the stress-strain curve to the point of its ultimate tensile strength
Intrinsic material properties	Independent of an object’s shape and size—thus, its study requires that the effect of the object’s shape and size (geometry) be 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 the deformability of a solid object.
Isotropic objects	Intrinsic material properties independent of the direction of loading and with a randomly dispersed internal structure (metal, glass, plastic)
Anisotropic objects	Intrinsic material properties dependent on the direction of loading and with an orderly internal structural arrangement (bone, intervertebral disks, ligaments/tendons)

Measurements of Force

The concepts of moment arm, bending moment, and IAR are critical to the understanding of spinal biomechanics. Forces applied to the spine can be simplified into component vectors (forces that are oriented in a fixed and well-defined three-dimensional direction). These vectors may then act on a lever (moment arm) to create a bending moment. This force is relative to a fixed point in space, the IAR, and results in rotation—or the tendency to rotate—about this axis. With respect to the spinal column, the IAR is the axis about which each motion segment rotates at any particular point in time and is the center of the x-, y-, and z-axes of the Cartesian coordinate system. The IAR can also be thought of as a dynamic fulcrum of movement that changes in response to applied forces.

The IAR also changes with the spinal level and migrates during intervertebral motion. The flexion-extension IAR of a cervical or lumbar motion segment is slightly dorsal and caudal to the center of the caudal vertebral end plate. An exception to this position is found at L5-S1. At this level the IAR lies within the disk space instead of below the caudal end plate. The IAR also moves proximally when going from C2 to C7. Changes in the IAR are important because they can result in excessive facet joint or posterior ligamentous loading.⁴

Deformation and Elasticity

The shape of all solids can be altered by external forces. For small displacements, the size of the deformation is proportional to the deforming force. This is known as Hooke’s law. Although valuable for analyzing the effects of force on various spinal instrumentation/constructs, it becomes unreliable with larger displacements because the neutral zone (component of the physiologic range of motion associated with significant flexibility and minimal stiffness at small loads) is exceeded and the elastic limit of the material is reached. The elastic limit is the point at which the force’s linear relationship departs from being linear between the size of deformation and the deforming force and results in permanent distortion. At its extreme, the solid may break; this is known as the point of failure (Fig. 265-10).

FIGURE 265-10 Typical stress-strain curve for a biologic 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. The entire colored area under the curve represents strength, whereas the tan area represents resilience.

(From Benzel EC. Physical principles and kinematics. In: Benzel EC, ed. Biomechanics of Spine Stabilization. Rolling Meadows, IL: AANS Publications; and New York: Thieme; 2001:24. Reprinted by permission.)

Stability versus Instability

The goal of all spinal stabilization techniques is to establish and maintain a nonpathologic relationship between the neural elements and the surrounding osseous and extrinsic soft tissues in a biomechanically favorable environment. Normal physiologic loads vary significantly. Therefore, stability should be interpreted as circumstance-dependent rather than an all-or-none phenomenon. For example, spinal stability should be provided when landing from a jump; conversely, a certain degree of spinal laxity is required to stoop for tying the laces of one’s shoes.

The inherent structure of the spine provides a physiologic and functional degree of freedom of motion. Normal range of motion includes translation and rotation about the three anatomic axes (Fig. 265-11) to provide six potential movements referred to as degrees of freedom. Segmental motions at the various spinal levels (see Figs. 265-4 and 265-5) are generally determined by facet orientation, bony anatomy, associated ligaments, and supporting structures.

FIGURE 265-11 Cartesian coordinate system with the instantaneous axis of rotation as the center. Translation and rotation can occur in both their respective directions about each axis.

Instability is the inability to limit excessive or abnormal spinal displacement in any plane. The term excessive underlies the difficulty in quantifying abnormal displacement. The extensive literature on acute spinal instability clearly illustrates the difficulty associated with such a definition process.

Many classification schemes for instability have been introduced over the last few decades. In general, instability is referred to as either acute or chronic. As discussed previously, instability is not an all-or-none phenomenon but rather an array of increments ranging from stable to grossly unstable. Moreover, acute instability can be described as overt or limited. Both these types (more commonly overt) may progress to a more chronic state if left unattended. Acute instability is most frequently encountered with traumatic, infectious, and neoplastic conditions. Chronic instability may, as previously suggested, be a sequela of an acute process, but it may also result from degenerative changes. Chronic instability may be subdivided into glacial instability (in which the deformity progresses slowly, like the motion of a glacier) or dysfunctional segment motion. In the latter there is no progression of the deformity, but rather a pain syndrome generated by dysfunctional motion. This is synonymous with mechanical instability.

Overt instability is defined as an inability of the spine to support the torso during normal activity. For such instability to occur, loss of VB or disk integrity must be combined with loss of integrity of the dorsal elements. This results in a circumferential loss of spinal integrity. Loss of ventral column integrity, as observed with wedge compression or burst fractures, can be readily seen on lateral plain radiographs and sagittal computed tomography or magnetic resonance imaging (MRI) scans.

The loss of dorsal integrity may be more challenging to assess. Plain films or sagittal tomograms may show splaying of the spinous processes or frank fractures of the dorsal elements. Dorsal pathology may also be noted on clinical examination by pain on palpation or loss of midline soft tissue definition. MRI, especially sagittal T2-weighted images, may be most useful in assessing dorsal ligamentous integrity, and the addition of fat suppression or short tau inversion recovery (STIR) sequences may offer even superior quality images. Overt instability is synonymous with gross instability and should be treated surgically in nearly all cases.

Limited instability is defined as the loss of either ventral or dorsal spinal integrity with preservation of the other. This is sufficient to support most normal activities. Isolated laminar fractures or ligamentous disruption (as illustrated on fat suppression T2-weighted MRI) with intact ventral elements is an example. Conversely, isolated wedge or burst vertebral fractures, with preserved integrity of the dorsal elements, are considered to constitute limited instability. Occasionally, underestimation of dorsal ligamentous injury may lead to overt instability being mistaken for limited instability. This is less likely if MRI is used liberally. Dynamic flexion-extension radiographs may also be useful in the context of limited instability. Dynamic radiographs may be misleading when guarding is present and may even be dangerous when underlying overt instability is present. Clinical judgment must guide the use of such imaging. Limited instability is usually managed nonoperatively with bracing. Surgery may be indicated if there is a significant risk for chronic instability (Table 265-2).

TABLE 265-2 Classification of Instability

TYPE	EXAMPLES
Overt	Inability of the spine to support the torso during normal activity
Limited	Loss of either ventral or dorsal spine integrity, with the preservation of the other
Glacial	Instability that is not overt and does not demonstrate a significant chance of rapid development of progression of the deformity but, like a glacier, progresses gradually with time. Substantial external forces do not cause movement or progression of the deformity.

Implant Properties

In addition to the deformation and elasticity properties discussed previously, understanding of spinal implants is not complete without a perception of section modulus and moment of inertia. Section modulus is an indication of the strength of an implant, which is defined as its resistance to failure in flexion. It is a function of cross-sectional area, as well as geometry. The moment of inertia portrays stiffness and is a measure of the implant’s distribution around its center. The diameter of a rod (or the core diameter of a screw) significantly affects strength (to the third power), as well as stiffness (to the fourth power).

Many constructs depend on screws to maintain attachment to the spine. The important anatomic aspects of a screw include the head, core, thread, and tip. Inner diameter can remain constant along the length of the screw or taper. Section modulus remains unchanged along the length of the screw in the former and rises exponentially in the latter. Pullout resistance is proportional not only to the volume of bone between screw threads but also to the triangular area defined by the screw: the perpendicular and dorsal VB surface. Alterations in thread pitch and distance between threads affect the interthread bone volume. Although screw length does not contribute significantly to pullout resistance, rigidly triangulated screws significantly increase resistance. Thus, increasing the screw angle (toe-in) increases the triangular area and thus pullout resistance (Fig. 265-12).

FIGURE 265-12 Screw core (minor) and outside (major) diameters, thread depth, and screw pitch.

(From Benzel EC. Implant-bone interfaces. In: Benzel EC, ed. Biomechanics of Spine Stabilization. Rolling Meadows, IL: AANS Publications; and New York: Thieme; 2001:159. Reprinted by permission.)

The Aging Spine

Cervical Spondylosis and the Aging Spine

The degeneration associated with aging has a significant impact on the spinal column’s fundamental biomechanics. As would be expected, younger spinal columns are more flexible and exhibit a significantly greater range of motion than do their older counterparts ⁷^–¹⁰ as a result of an accumulation of changes, including facet joint osteoarthritis, dehydration of the intervertebral disk, and loss of normal spinal alignment.¹⁰ It seems that the cervical region is most susceptible to these changes, in part because of its high degree of mobility, as well as the more complex geometry of the uncovertebral joints.¹⁰

Spondylosis

Many of the changes associated with age start at the microscopic level. The functional proteoglycans in cartilage decrease significantly with age, thereby resulting in a gradual reduction in the tissue’s water-binding properties and an associated decrease in its shock-absorbing ability, as well as increasing friability.¹¹ Stiffness of the cartilage also increases, partially as a result of increased cross-linking between fibrous proteins, especially the collagens.¹² Both these changes result in increased susceptibility of the cartilage and tendinous attachments to injury. The intervertebral disks also lose much of their functional ability with age. Although this is probably the result of millions of load cycles throughout an individual motion segment’s functional life span, it is accelerated by an age-related decline in transport of metabolites within the avascular matrix.¹³ As one would surmise, this also increases the disk’s vulnerability to injury.

Osteoporosis

The aging process is commonly associated with a gradual decrease in bone mineral density that can lead to osteoporosis and fracture. Because these changes are generally initiated in cancellous bone, the VBs of the spine are often involved.¹⁴ The compressive strength of cancellous bone is related to the square of the apparent density, and the elastic modulus is also related to the apparent density.¹⁵ Thus, the strength of the VB and its resistance to axial loading are immediately compromised. However, the process is more complex. As an adaptive response to decreased bone mineral density, the trabecula of the bone remodel. Specifically, the ratio of vertical to horizontal trabecular orientation increases, especially in the anterior third of the VB. As this ratio increases, the ability of the VB to resist axial loads becomes greater, and so does the directional dependency of load resistance. The decrease in horizontal trabecula results in a decrease in elastic modulus and strength in the transverse direction and an increased vulnerability to forces other than pure axial. This could partially explain the higher incidence of VB fractures associated with osteoporosis in the anterior part of the VB (wedge fracture).¹⁴ Thus, although an osteoporotic VB adapts to resist axial loads, it sacrifices its ability to withstand loading in directions different from the direction of the principal trabecular orientation or the shear load.

Spinal Alignment

Age-related postural changes affect the relative orientation of adjacent vertebrae and profoundly alter stress distributions within the apophyseal joints and intervertebral disks. A perfectly straight spine would theoretically be an ideal axial-loading spinal configuration, but it would tolerate eccentric loads poorly and provide limited flexibility. The spine has therefore 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. This results in a balanced configuration that is necessary for a bipedal upright posture (Fig. 265-13). Any increase in thoracic kyphosis (or loss of lumbar lordosis) leads to an increased moment arm (i.e., perpendicular distance from the IAR to the gravitational force vector), which generates a greater bending moment at each vertebral segment (Fig. 265-14). The moment arm (M) is equal to the force (F) multiplied by its perpendicular distance (D) from the IAR (M = F × D). The greater the deformity, the greater the length of the moment arm; hence, “deformity begets deformity.” The same mechanism applies to deformities in the coronal plane (scoliosis) (Fig. 265-15).

FIGURE 265-13 A kyphotic posture (as is present in the thoracic spine) increases the length of the natural moment arm (D) and thus the magnitude of the bending moment resulting from an eccentrically placed (with respect to the instantaneous axis of rotation) axial load (arrows).

(From Benzel EC. Trauma, tumor, and infection. In: Benzel EC, ed. Biomechanics of Spine Stabilization. Rolling Meadows, IL: AANS Publications; and New York: Thieme; 2001:75. Reprinted by permission.)

FIGURE 265-14 Depiction of the injury force vector causing a ventral wedge compression fracture. D, length of the moment arm (from the instantaneous axis of rotation to the plane of F); F, applied force vector; IAR, instantaneous axis of rotation; M, bending moment.

FIGURE 265-15 “Deformity begets deformity.” A kyphotic deformity, as depicted, is associated with a moment arm of length d. If an axial load (straight arrows) is applied, a bending moment (M) is created (curved arrow) (A). Application of this load, including the bending moment, results in the application of an eccentric load to the spine (greater ventrally than dorsally). This begets further deformation, in this case kyphosis (B).

Biomechanics of Intervention

Iatrogenic Spinal Destabilization

A surgeon’s appreciation of the biomechanical properties of the ligamentous and osseous structures of the spine is vital to avoid destabilization during ventral and dorsal spinal exposures.

During dorsal decompression, preservation of the facet joints, interspinous ligaments, and uncovertebral joints, when possible, will minimize the risk for iatrogenic destabilization. Approximately one third to half of the facet joint may be resected without causing destabilization. In the lumbar spine, the instability associated with lumbar facet disruption is typically glacial and does not promote a significant risk for the rapid development or progression of translational deformities. Lumbar facet integrity may be minimally disrupted during laminectomy if an optimal trajectory is used and the pars interarticularis remains intact. The uncovertebral joints regulate extension and lateral bending motion and resistance to torsion. Damage (especially to the posterior uncovertebral joints) can result in loss of these resistive forces. Although the interspinous ligament is relatively weak, it has biomechanical advantages related to its long moment arm. This ligament is deficient at L4-5 and usually absent at L5-S1.

During ventral decompression, the anterior longitudinal ligament (ALL), posterior longitudinal ligament (PLL), annulus fibrosus, and VB, all of which contribute to stability, are often sacrificed to some extent. The ALL is not usually significantly disrupted because of the extensive width of the ligament. It provides a tension band–like effect, which is an especially important contributor to postoperative spinal stability secondary to its position ventral to the IAR. Ventral annulotomy during anterior lumbar interbody fusion may significantly disrupt the ALL, thus requiring the PLL and dorsal elements to provide adequate tension banding. The PLL resists flexion/distraction and has far less biomechanical strength because of its weak intrinsic mechanical properties and short moment arm in relationship to the IAR. Obviously, a decompression that exposes the thecal sac disrupts the PLL completely. It is also significantly impaired after many ventral surgical approaches in which the intention is to decompress the spinal cord. The degree of ventral spinal integrity that remains after corpectomy is determined by the proportion of bone remaining in the ventral component of the VB. Although resection of the ventral portion may result in loss of stability, loss of the middle and dorsal thirds may not result in loss of spinal integrity if the ventral section is intact, the ALL remains intact, and dorsal column osseous and ligamentous integrity remains intact.

Pathologic or iatrogenic reduction in stability, if biomechanically significant, must be compensated for by one or more of the following three therapeutic maneuvers: postural correction, other nonoperative management (including spinal splinting) that provides time for osseous and ligamentous healing to offset the acute disruption of spinal integrity, or placement of a ventral spinal bone strut or instrumentation or dorsal instrumentation.

Construct Design Principles

Planning for instrumentation is based on numerous biomechanical principles. Depending on the situation, the surgeon needs to choose which principle needs to be applied to obtain lasting stability.

Construct Failure

A construct needs to survive 3 to 5 million cycles of loading after insertion to provide support for 1 year. The vast majority of spinal implant failures are secondary to surgeon-related underestimation of these stresses, poor construct design, and improper patient selection. Construct failure occurs when the implant, implant-bone interface, or component-component juncture becomes incompetent.

Implant failure may result from either instantaneous or cyclic overload. Fatigue failure occurs as a result of the accumulation of microinjuries or damage to the instrumentation. It is dependent on the intrinsic material properties of the implant, as well as its exposure to repetitive stress. Instrumentation (plates, rods, and screws) breaks at the point at which maximum stress is applied. This is the point at which the ratio of the applied bending moment and the section modulus is greatest. All constructs have stress concentration points, which are often iatrogenic. These points can be structural imperfections or surface irregularities made on a rod or plate during contouring/bending. They can also occur at areas with sudden changes in cross section and drill holes.

Degradation of the screw-bone interface results in toggling of the screw (moving in a windshield wiper motion). Sufficient force can also result in the screw cutting or pulling out of the bone. Long rigid (fixed moment arm) constructs tend to load the more caudal screws far more than the rostral screws and are associated with a high failure rate (Fig. 265-16).

FIGURE 265-16 Screw-bone interface failure can occur via toggle (A), cutout (B), and pullout (C). Plate fracture can occur at the point of maximum stress application (i.e., maximum bending moment application at a weak point in the plate [low section modulus]). The latter may occur at a stress riser caused by plate contouring. This is illustrated by radiographs D and E.

(From Benzel EC. Deformity prevention and correction: Component strategies. In: Benzel EC, ed. Biomechanics of Spine Stabilization. Rolling Meadows, IL: AANS Publications; and New York: Thieme; 2001:371. Reprinted by permission.)

Bone-Bone Interface Failure

Factors affecting the incidence and extent of nonunion/subsidence include the closeness of fit of the bone graft in the VB mortise, the surface area of contact between the bone graft and VB, and the character or quality of the contact surfaces. The extent of end plate preservation and proximity of the contact point to the edge of the VB affect the quality of the contact surfaces. Greater construct strength is achieved when the cortical portion of a graft is positioned in line with the cortical surface of the VB. Some implants take advantage of the boundary effect by providing support at the edge of the VB, thus maximizing resistance to axial loading. The greatest biomechanical advantage with respect to interbody axial load-bearing ability is achieved when a strut graft is nearly the same size as the VB in terms of contact surface area. Creation of a lateral buttressing effect can help minimize bone-bone interface failure, and lateral fit of the bone graft against the wall of the corpectomy trough further optimizes the interbody bone-bone interface relationships. Increasing bone-bone contact surface area will decrease the chance of graft pistoning (Fig. 265-17).

FIGURE 265-17 An interbody strut’s penetration is inversely proportional to its cross-sectional area of contact with the vertebral body. Note that the thinner (less substantial) strut pistons are further than the more substantial strut pistons (A). A mortise that is poorly fitted (matched) to the bone graft (B) increases the chance that a space will persist between the bone graft and the mortise (shaded area) (C) or that pistoning (D) will occur.

(From Benzel EC. Spinal fusion. In: Benzel EC, ed. Biomechanics of Spine Stabilization. Rolling Meadows, IL: AANS Publications; and New York: Thieme; 2001:123. Reprinted by permission.)

Biomechanics of Nonfusion Implants

As nonfusion devices gain popularity, an understanding of their biomechanical effect on the implanted and adjacent motion segments will be required. Ideally, they should mimic the IAR and range of motion of the normal functional unit, which varies by level and individual. Currently, these devises fall into three categories: nuclear implants, total disk replacement (TDR), and posterior stabilization devices.

Biomechanics of Nuclear Implants

Nucleus pulposus replacement implants are designed to restore disk turgor and tension in the annulus fibrosus, as well as reestablish the disk’s ability to transfer loads uniformly across the disk space.³ Remodeling of the VB end plates has been observed after nuclear replacement. Such remodeling is probably the result of changes in load concentration on the end plates after placement of the device and is being addressed by increasing the area of the device in contact with the end plate, as well as by minimizing implant stiffness.

Biomechanics of Total Disk Replacement

To avoid wear-related failure and progressive facet joint arthrosis and minimize the effect on adjacent motion segments, TDR biomechanics needs to be as close to the original disk as possible. Biomechanical constraint of the TDR, defined as the limitation of pure anteroposterior or lateral translational intervertebral motion, will affect the stress on the remaining facet joints. The ball-and-socket articulation TDRs have much higher constraint and secondarily less shear translation. Because facet cartilage tends to be loaded by anterior shear loads, the more constrained implants are able to reduce loading of the facets at the expense of greater loads on the implant and the implant-bone interface.³ Similarly, less constrained TDRs may allow increased facet loading. Contrarily, a constrained TDR dictates the arc or rotation. This can result in increased loading of the facets, especially at the extremes of motion.³

In constrained TDRs, the location of the IAR and the radius of curvature of the articulation are important predictors of its biomechanical characteristics. A relatively posterior IAR better reflects the physiologic IAR and results in greater range of motion.¹⁶ The radius of curvature is the distance from the IAR to the implant’s surface, and this determines the relative proportion of rotation and apparent translation observed in a ball-and-socket joint. With a small (round) radius of curvature, rotation dominates. With a large (flat) radius of curvature, translation dominates. Although the interplay of flexion-extension and translational motion affects facet loading, its preservation is important in reducing stress on adjacent levels, one of the primary goals of TDRs.

One of the leading causes of TDR failure in the lumbar spine is subsidence, which is largely a result of the compliance mismatch between a metallic implant and the adjacent vertebral end plate. Because the outer rim of the vertebral end plate is thicker than the central area, seating the TDR on this thickened peripheral ring of bone could minimize this risk.