Basic Biomechanically Relevant Anatomy

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Chapter 136 Basic Biomechanically Relevant Anatomy

Spinal instrumentation has led to many significant advances in how we treat our patients. This treatment may range from the early mobilization after operative treatment of spinal fractures to the prevention of late iatrogenic cervical kyphosis from laminectomy. Selecting the appropriate fixation techniques and instrumentation can be a complex problem. Often, the final surgical solution is one that balances the specific patient requirements, the spinal instrumentation available, and the biomechanics of the spinal pathology. Before we can begin to discuss the biomechanics of instrumentation, we must first understand the biomechanics of the relevant spinal anatomy.

Spine biomechanics is the application of basic engineering and physics principles to a biologic structure, the spinal column. It allows the surgeon to reduce the complex movement and forces on the spine to some basic concepts. The anatomy of the spine must be placed in biomechanical terms, such that the role of the each component is understood in relation to the overall structure and function of the spine. The spine is a complex load-bearing structure made up of bones, joints, ligaments, and muscles arranged to effectively dissipate forces and allow a physiologic range of motion, while at the same time protecting the spinal cord.

Stability may be compromised through trauma, tumor, degenerative disease, or iatrogenic artifacts. Spine surgery itself can often be a destabilizing process, and the surgeon must use instrumentation to restore stability. Spinal implants are designed to impart a force to the spine. To apply instrumentation appropriately, a thorough understanding of the basic biomechanics must be attained. Our goals in this chapter are to introduce the basic concepts of biomechanics and apply them to normal and pathologic spinal anatomy.


Biomechanics involves the application of mechanical engineering principles to biologic problems. It is the science of the action of forces on the human body. It refers to the direction and quality of force as well as to the mechanical properties of the biologic material. It that way, it is possible to describe the force vector of a tendon as well as the plastic deformation of bone. Spine biomechanics is extremely complex, because it must deal with the local as well as global mechanics of the spine. This can be further complicated by any pathologic dysfunction of the spinal column. To appreciate the complexities, it is necessary to first consider some basics about biomechanics, including coordinates, vectors, and moduli.

Vectors, Scalars, Bending Moments, and the Instantaneous Axis of Rotation

A vector is a force oriented in a fixed direction in three-dimensional space (as defined by the right-handed Cartesian coordinate system previously discussed). It possesses both magnitude and direction (i.e., force and velocity), and it may be broken down into its component vectors. A scalar is a quantity that is direction independent and possesses only magnitude (i.e., time, volume). A bending moment occurs when a force acts on a lever or moment arm. A bending moment, when applied to a point in space, results in a tendency to rotate about an axis. When these principles are applied to a vertebral body, the three-dimensional axis about which a bending moment causes the body to rotate; this is known as the instantaneous axis of rotation (IAR). The IAR is the axis about which a vertebral segment rotates at any given instant, and the IAR itself does not move during this instant of rotation or translation. It may be considered a fulcrum about which a segment moves. When a body moves in a plane, there is a point in the body or some extension of it that does not move, and the axis that passes perpendicular to the plane of motion and through this point is the IAR (as defined by White and Panjabi1). The IAR becomes the center of the Cartesian coordinate system used to define the motion. The IAR is dynamic, and as spinal movement occurs, the IAR of each involved segment moves. The location of the IAR depends on the manner in which it is determined and the theoretical foundation on which it is based, and it is therefore subject to error.

Newton’s Laws

Spine surgery involves the application of force to the spine. Fundamental to understanding the stresses withstood by the spine is the concept of action and reaction.2 Newton’s three laws of motion describe how objects respond to external forces.

As the spine is subject to various loads, force couples are created (based on the third law of motion), and the vertebral segments may fail. When applying forces to the spine, these same concepts need to be considered. All instrumentation will fail in time, and therefore the bony alignment must respect the body and gravitational forces or it will fail more quickly. The biomechanical environment must maintain a load balance between the implant and tissue environment to improve implant longevity.

Hooke’s Law and Stress-Strain Curves

When external forces act on a solid at rest, the solid is deformed. According to Hooke’s law, for small displacements, the size of the deformation is proportional to the deforming force. For solids within the elastic zone, the relationship between deformation and force is linear. To make sense of Hooke’s law then as it applies to the spine, the concepts of stress-strain curves and deformation need to be expanded. Deformation of a solid can be either elastic or plastic. Elastic and plastic are defined by recovery once the deforming force is removed, and they exist on a spectrum that can be best illustrated with a deformation curve (Fig. 136-2). When a solid is subject to an external or deforming force, Hooke’s law applies only when the load and deformation occur within the elastic zone of the solid. The elastic zone is defined by elastic deformation, when a solid totally recovers after removal of a stress. Deformation continues to be proportional to the deforming force until the elastic limit of the solid is reached and the linear relationship between force and deformation no longer exists. Beyond the elastic limit of a material is the plastic zone, and this is the zone in which the solid acquires permanent deformation and does not completely recover once the stress is removed. As force continues to be applied, eventually the point of failure is reached. On a stress-strain curve (see Fig. 136-2), the area under the curve represents the amount of energy absorbed and the ultimate strength of a material (amount of energy required to reach failure). The neutral zone is an area of nonengagement where force is applied but deformation is minimal. It precedes the linear area of the curve.

Section Modulus and Moment of Inertia

The section modulus is a concept that is calculated and used to define the strength of an object. It is applicable here with regard to spinal implants and instrumentation as an indicator of the overall strength and, more importantly, potential failure of an implant. The section modulus or strength of a screw is exponentially related to the diameter of the screw. As this diameter increases, the strength increases by the power to the fourth. This concept is also applicable to plate design. The strategy for hole placement for screws must be carefully designed because the holes reduce the section modulus. This then causes failure of the plate at a lower load (Fig. 136-3).

Other important factors relating to implant failure or potential failure that need to be considered are potential force applications (moment arms, loads, bending moments). Stress defines the relationship of section modulus and bending moment. It represents the force per unit area applied to an object and is defined as


where T is stress, M is bending moment, and Z is section modulus. The bending moment (M) is also defined by an equation:


where M is bending moment, F is the force applied, and D is the distance of the moment arm through which the force is applied.

The stiffness of an object (implant) is the resistance to deformability. It is also defined as the slope of the elastic or linear zone of a stress-strain curve. Therefore, the stiffness is affected to a greater extent by the diameter than the strength and increases more rapidly as diameter increases. This applies directly to screws and rods, which are often used in the spine.

Basic Biomechanics of Bone

Bone is composed of osteocytes, which account for 15% of the volume of bone; organic materials such as collagen; and nonorganic materials such as calcium and phosphate. The collagen has high tensile strength, and the nonorganic material resists compression. Because of the lattice orientation of the nonorganic material and the orientation of the cortex, bone is weakest in shear. This difference in strength is an example of an anisotropic material—that is, a material that has different mechanical properties depending on the direction the force is applied. These mechanical properties may also change with age. As people age, the mechanical properties of the bone decline as bone mineral is resorbed. As mentioned previously, this is the primary restraint to compression. This change may predispose older people to mechanical compression fractures.

Bone fractures and the fracture patterns also follow the magnitude and direction of the forces applied. Wolff’s law, a theory developed by the German anatomist/surgeon Julius Wolff (1836–1902), states that bone adapts to imposed stress. If loading on a particular bone increases, the bone remodels itself over time to become stronger to compensate for greater load requirements.3

Rapidly applied loads to the spine may result in failure of the ligaments or bone, or both. The direction of force application, magnitude, and bony integrity all play a role in the fracture pattern and displacement. Several distinct failure patterns can be observed.

Avulsion Fracture

These occur as a result of the strong ligamentous attachments to the spine. The ligaments, which are strongest in tension due to their high collagen content, can avulse bony fragments (Table 136-1). An example of this is the bony avulsion seen with Jefferson burst fractures, where the transverse ligament avulses a portion of the lateral mass of C1.

TABLE 136-1 Spinal Ligament Forces

Ligament Cervical (N) Lumbar (N)
ALL 112 450
PLL 75 324
ISL 240 125
SSL 36 150

ALL, anterior longitudinal ligament; ISL, interspinous ligament; PLL, posterior longitudinal ligament; SSL,supraspinous ligament.

From White AA, Panjabi MM: Clinical biomechanics of the spine, ed 2, Philadelphia, 1990, Lippincott, p 22.

Burst Fracture

The burst fracture typically has both axial load components as well as a flexion moment. Comminution is a function of the magnitude of force that was transmitted to the bone. It can be graded on the amount of the vertebral body that is involved and the amount of displacement. These grades have been used to understand the ability of that bone to support the anterior spine. Gaines described this load-sharing classification and the necessity for long-segment fixation or supplemental anterior support based on the concept of vertebral body comminution and fracture displacement.5 In healthy bone, this energy results in a burst pattern rather than a compression pattern. With failure of the anterior support, the spine falls into flexion with continued energy. The resultant disruption of the dorsal tension band or posterior bony fractures is usually the primary predictor of stability. Most classification systems recognize this, and posterior stability is a primary predictor of necessity of operative stabilization.6

Basic Anatomy of the Spine

The spine is made up of 33 separate vertebral segments that are interconnected by spinal ligaments, facet capsules, and intervertebral discs. Each bony segment is referred to as a vertebra, and each is numbered in a region-specific (i.e., cervical, thoracic) rostral to caudal fashion. Thus, there are 7 cervical vertebrae, 12 thoracic vertebrae, 5 lumbar vertebrae, 5 fused sacral vertebrae, and 3 to 4 fused coccygeal segments. All four normal curves of the spine occur in the sagittal plane. The cervical and lumbar spine is lordotic or concave posteriorly due to the wedge-shaped intervertebral discs. The thoracic and sacral regions are kyphotic or concave ventrally; at least in the thoracic spine where the segments are not fused, this is structural and due to the shape of the thoracic vertebral bodies. The anterior vertebral border height is less than the dorsal vertebral border height, resulting in a ventral or kyphotic curve.

These curves function to increase the flexibility and shock-absorbing capacity of the spine. The overall physiologic alignment must be maintained to prevent construct failure and accelerated degeneration. Sagittal balance is an indicator of physiologic alignment and is particularly important in lumbosacral procedures. In a normal spine, a plumb line called the sagittal vertical axis (SVA) can be dropped from the C7 vertebral body through the lumbosacral junction (Fig. 136-4). If the SVA is ventral to the lumbosacral junction, it is called positive sagittal balance; if the SVA is posterior to this junction, it is negative. Positive sagittal balance has significant implications for body mechanics and muscle work. Many authors describe a positive sagittal balance as an important factor in outcomes and muscle mechanics.8,9 To compensate, the pelvis shifts posteriorly and allows the patient to continue to maintain his or her global balance even with a positive SVA.10 The pelvis may also retrovert and move the center of gravity posterior to compensate for the positive SVA. Additionally, patients maximally extend their hips. Once this fails, they are forced to bend their knees and flex the hips in a biomechanically disadvantageous position.


FIGURE 136-4 Sagittal balance. A plumb line dropped from the C7 vertebral body in the standing position should pass through the lumbosacral junction when normal sagittal balance is present, as depicted.

(From Benzel EC, editor: Biomechanics of spine stabilization, Rolling Meadows, IL, 2001, American Association of Neurological Surgeons.)

Vertebral Body

The vertebral body is the bony cylinder that makes up the ventral aspect of each vertebra, except C1, which has no vertebral body. Each vertebral body consists of an outer rim of cortical bone surrounding a core of softer cancellous bone. The vertebral bodies resist much of the compressive loads placed on the spine in physiologic situations.

In general, the height, width, and depth of the vertebral bodies increase as the spine is descended, which seems to account for this increased strength at lower levels in the spine.1 As previously alluded to, there is a regional variation to the shape of the vertebral body. In the thoracic spine, the anterior border is shorter than the posterior border, creating a structural kyphosis in the thoracic spine. The outer shell of cortical bone is more rigid than the softer cancellous core because it is arranged in vertical lamellae to resist compressive forces. The cancellous bone is made up of trabeculae that are arranged to resist a variety of loads. There is greater compressive deformation of the cancellous bone prior to failure than of the cortical bone because of the increased rigidity of the latter.11,12 Numerous studies regarding the load-sharing properties of the cortical and cancellous bone have been performed showing that the load carried by the trabecular bone varies anywhere from 35% to 90%, depending on age and mineral content of the bone,13,14 and that the strength of the vertebral body decreases with age. Bell et al.15 have shown that there is a direct relationship between the mineral or ash content of the bone and the strength of the vertebral body. A 25% decrease in mineral content resulted in a greater than 50% decrease in vertebral strength due to loss of the trabecular columns that formed the core of the vertebral bodies. In osteoporosis, the mineral content of the bone is affected and results in loss of the horizontal trabeculae within the vertebral body core, thus lengthening the trabecular columns and compromising vertebral body strength.1 In the thoracic spine, the vertebral bodies articulate with the ribs. This articulation with the rib cage significantly augments the strength of the thoracic spine. This articulation with the rib cage and the sternum significantly augments the stiffness of the thoracic spine by as much as 110% in extension and 30% to 45% in the other planes of motion.

The vertebral end plates are formed from the concave, 1- to 2-mm thick cortical bone at the rostral and caudal surface of the vertebral body. These are fused to the cartilaginous end plates of the intervertebral disc by a calcified layer of tissue known as the lamina cribrosa, which permits osmotic diffusion of nutrients for the intervertebral disc. In 1957, Perry16 extensively studied failure of the end plates under compression and noted that one of three mechanisms of failure occurs, depending on the condition of the surrounding discs. Fractures occurred centrally or peripherally or encompassed the entire end plate. In specimens with nondegenerated discs, the fractures typically occurred centrally, and in specimens with degenerated discs, the fractures occurred peripherally. At high loads, the fractures encompassed the entire end plate regardless of degeneration. The strength of the end plate itself has been evaluated as well with regard to resistance to penetration by graft material (subsidence) or disc material (Schmorl node). Kumar et al.17 determined the greatest resistance to penetration to be within the first 4 mm of depth and the end plate to be strongest at the periphery closer to the cortical margin. In fact, the weakest portion of the vertebral end plate is at the center, where it is the thinnest. With a nondegenerated nucleus, axial compression creates a central increase in pressure and deflection of the end plate. This leads to increasing bending stresses in the center of the vertebral end plate and ultimately results in its failure.1 In contrast, in the degenerated disc, the nucleus lacks water content, and there is no build-up of central fluid pressure. The load is transmitted primarily through the anulus at the periphery of the disc and end plate, resulting in fracture in the same location or in the underlying vertebral body.1

Intervertebral Disc

The vertebral bodies are separated at each level by the intervertebral disc. This viscoelastic structure is made up of a central nucleus pulposus, peripheral anulus fibrosus, and cartilaginous end plate of the disc, which separates it from the vertebral bodies above and below. Viscoelasticity is defined as a property of a material that exhibits both viscous and elastic characteristics when undergoing deformation. The material then has different mechanical properties that vary depending on the different loading rates. That is, if a load is applied slowly, the disc can respond with greater deformation, whereas if a load is applied rapidly, less deformation occurs. This property also changes as people age; with increasing age, there is a corresponding decrease in the ability of the disc to deform. There are 23 intervertebral discs starting between C2 and C3 and ending at L5 and S1. They make up 20% to 33% of the total height of the vertebral column, and there are regional differences that seem to parallel those of the vertebrae. The cross-sectional area of the discs increases in the rostrocaudal direction.

The central nucleus pulposus is composed of mucopolysaccharides and mucoprotein, forming a gel with water content ranging from 70% to 90%.1 The water content is highest at birth and decreases with age.18 The nucleus pulposus makes up approximately 30% to 50% of the cross-sectional area of the disc, and it seems to lie more posterior in the lumbar spine at the junction of the middle and posterior thirds of the disc in the sagittal plane.1

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