Biomaterials and Biomechanics of Spinal Arthroplasty

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CHAPTER 266 Biomaterials and Biomechanics of Spinal Arthroplasty

Brian R. Subach, Anne G. Copay, Thomas C. Schuler

The concept of artificial disk replacement is clearly not new. Documented attempts at disk arthroplasty, although unsophisticated, date back nearly 50 years. This chapter attempts to define the biomechanics of the native cervical and lumbar spine, the biomechanics of disk arthroplasty, and the biomaterials available to manufacture arthroplasty devices.

The success associated with replacement of a joint of the appendicular skeleton, such as the hip joint or knee joint, has led to the assumption that joints of the spine may also be amenable to arthroplasty. The functional spinal unit (FSU) is composed of the intervertebral disk and paired facet joints and may be partially augmented (nuclear replacement) or completely replaced (disk arthroplasty, facet joint replacement). Unfortunately, the significant stress placed on the axial skeleton, in combination with the complex biomechanical properties of both the cervical and lumbar spine, such as coupled motion, makes both design of the arthroplasty device and selection of the ideal composite biomaterial a complicated undertaking to say the least. In 1955, Cleveland reported 14 patients in whom he implanted a methyl-acrylic device into the intervertebral space at the time of diskectomy.1 This was followed by Harmon’s use of Vitallium spheres in 13 patients from 1959 to 1961.2 These implants, which were inserted into the lumbar spine through an anterior retroperitoneal approach, led to spontaneous fusion rather than preservation of motion. Another pioneer was Fernstrom, who in 1966 published the outcomes of patients undergoing implantation of a solid stainless steel sphere into the lumbar disk space through a posterior approach at the time of lumbar diskectomy.3 These were important steps that unfortunately met with profound failure. Inadequate surface area resulted in subsidence. Poor design resulted in disk space collapse without restoration of motion. With the advent of more forgiving designs of arthroplasty devices, the challenge has become creation of a device that can withstand hundreds of thousands of wear cycles while allowing ease of insertion and uncomplicated revision.

Historically, anterior cervical decompression with arthrodesis has been a very successful operation for treating both neck pain and cervical spinal cord or root compression syndromes. With the use of interbody graft material and plate fixation, high rates of fusion and high rates of clinical success have been well documented. The surgical goal is solid bony arthrodesis and decompression with a clinical end point of relief of neck pain and neural compression. It has been estimated that failure or pseudarthrosis after attempted anterior cervical fusion may develop in up to 20% of patients.4 Pseudarthrosis may be associated with increasing neck pain, progressive neurological deficit, or the development of spinal deformity. Even in cases of successful fusion, loss of the FSU may transfer biomechanical stress to adjacent levels and presumably lead to more rapid deterioration of affected levels.5,6 Clearly, successful arthroplasty would eliminate the possibility of pseudarthrosis while maintaining neural decompression. This would achieve the ultimate goal of minimizing degeneration of adjacent segments while curing both neck pain and radicular symptoms.

In dealing with the lumbar spine, the biomechanics is clearly different. Not only is the axial load on the lumbar spine greater, but the facet joints and disks are also much larger structures. Perhaps an even more difficult problem is the difficulty in anterior access, particularly after previous anterior lumbar operations have been performed. Unlike the cervical spine, in which anterior access is relatively straightforward, revision surgery in the lumbar spine may present both challenging and possibly life-threatening hurdles. Initially, mobilization of the great vessels can be accomplished carefully by either the spine surgeon or an approach (general or vascular) surgeon. After a surgical intervention, scar tissue often develops around the great vessels and obliterates the normal planes of dissection. An attempt to mobilize vessels in this situation is a significant undertaking. In a perfect world, the implanted arthroplasty device in the lumbar spine would last throughout the entire life of the patient. It would never shed products of motion, known as wear debris, and would significantly eliminate stress on both the facet joints and the adjacent disk levels while relieving both back pain and radicular complaints.

To further define the scope of the problem, more than 80% of Americans will experience a significant episode of neck or back pain at least once during their lifetime. Initially, most back pain was attributed to simple muscular strain. Now, with increased understanding of the progressive injury that affects intervertebral disks, concepts such as diskogenic disease and facet arthropathy predominate. Low back pain has become a problem not only for the individual but also for society as a whole. The significant loss of work days because of back pain, the need for chronic pain medications, and the restriction of activities have had a significant impact on the functioning workforce in this country. As intervertebral disk pathology progresses, not only does loss of hydration of the disk space result in decreased shock-absorbing capacity of the nucleus, but there is also progressive wear. This wear adversely affects the annulus and may result in extremely painful annular tears and often irritation of the traversing or exiting nerve roots associated with the disk spaces. Initial attempts at disk arthroplasty were relatively crude by current standards. At this time, a number of disk arthroplasty devices have been approved by the Food and Drug Administration, and the amount of information available to patients via the Internet is significant.

Anatomy of the Normal Spinal Disk and the Degenerative Process

The intervertebral disk is a weight-bearing and stress-absorbing structure located between the vertebral bodies of the spine. The disk itself is composed of three separate components: the nucleus pulposus, the anulus fibrosus, and the cartilaginous end plates. The central nucleus pulposus is a hydrated gel of proteoglycans with sulfated glycosaminoglycan side chains. The anulus fibrosus has 12 coaxial layers of collagen fibers that insert on the adjacent end plates at an oblique angle. These collagen fibers consist of both type I and type II collagen. The cartilage of the end plates is called hyaline cartilage. It is primarily a proteoglycan gel similar to the nucleus, but it is reinforced by collagen fibers. As wear occurs within the intervertebral disk, there is a progressive loss of cellularity, as well as hydration. With loss of hydration, the proteoglycans begin to lose their function. The amount of type I collagen gradually decreases, whereas the amount of type II collagen gradually increases. Annular fissures or tears begin to develop, loss of mechanical competence occurs, fibrovascular changes begin to develop in the subchondral bone, and as segmental instability develops, pain and radicular irritation may occur. Commonly seen in the later phases of degeneration are sclerotic changes in subchondral bone. End plate sclerosis as a result of aging and degeneration basically blocks the pores that allow diffusion of nutrients across the vertebral body end plate into the disk space. Because the disk space itself has relatively few cells to begin with, it does not require a tremendous amount of blood flow to maintain health and integrity. Conversely, the lack of blood flow to the disk associated with the aging process clearly speeds degenerative destruction of the individual disk components.

Arthroplasty Biomaterials

Spinal disk arthroplasty designs have clearly been influenced by progress made in the field of hip and knee joint arthroplasty. A number of the designs use a combination of metals and polymers. The polymers or plastics essentially provide some measure of shock absorption for the joint while also providing a low-friction surface for joint articulation. Polymers themselves are not strong enough to tolerate the stress of normal joints and are therefore always supported by a metal base. Articular surfaces may involve a metal-on-metal, metal-on-polymer, ceramic-on-polymer, or ceramic-on-ceramic interface. Two polymers that have been used successfully are polyurethane and the derived ultra-high-molecular-weight polyethylene (UHMWPE). Metallic devices have been constructed from solitary metals such as stainless steel, titanium, and cobalt. Newer devices take advantage of metallic alloys. Alloys are homogeneous mixtures or solid solutions of two or more metals. For example, cobalt-chromium alloy, cobalt-chromium-molybdenum alloy, and titanium-aluminum-vanadium7 have characteristics that make them uniquely suited for use in arthroplasty. The alloys seem to wear more slowly than polymers and may resist corrosion better than single-metal implants.

Regardless of whether the prosthesis is metal on metal or metal on polymer, the choice of material used in creating the device must be governed by three principles: articular surface wear, generation of wear debris, and host inflammatory response. The first is the extent of articular surface wear of the device based on repetitive motion cycles. A motion cycle is a typical bending motion of the spine such as flexion-extension or lateral bending. If multiplied by the number of times that the patient bends or twists in a day or a year, the number of motion cycles is considerable. The extent of weight bearing and joint range of motion (ROM) will also determine the amount of friction on the articulating surfaces and therefore the amount of destruction of the device over time. The second factor is generation of wear debris. Because there is no frictionless surface, load on a device combined with motion will generate wear debris. Wear debris, when trapped within the joint, can lead to rapid destruction of the articular surfaces, generation of toxic metal ions, and a profound inflammatory response, depending on the materials involved. The third principle is generation of the host inflammatory response. The inflammatory response can lead to resorption of the bone surrounding the prosthesis, which can result in loosening of the implant and ultimately failure. In general, the greater the degree of wear debris production, the greater the degree of inflammation, and the greater the extent of periprosthetic bone resorption.

To describe the complex interaction of biomaterials in motion preservation devices, a unique discipline known as tribology has developed. Tribology is the science and technology of interacting surfaces in relative motion. It includes the study and application of the principles of friction, lubrication, and wear. Typically, in a normal joint the cartilage in contact with the adjacent articulating surface will wear. That cartilage does cause an inflammatory response in the periarticular tissues, known as arthritis. In the case of disk arthroplasty, wear debris clearly causes an inflammatory reaction. In certain designs, such as those using stainless steel, the extent of the inflammatory response is proportional to the amount of wear debris generated. This inflammatory response is often manifested not only as the cardinal signs of inflammation (pain, heat, redness, and swelling) but also at the microscopic level, where macrophages are actually seen having engulfed metal or plastic fragments. On the molecular level, measurable quantities of potentially toxic metal ions are released into the circulation. The reaction of the body to the inevitable generation of wear debris is the third point to consider when choosing the appropriate biomaterial from which to fabricate an arthroplasty device. It is thought that the inflammatory response associated with wear debris is what leads to osteolysis surrounding the implant. Osteolysis can result in loosening of the implant, abnormal motion of the implant, and eventually misalignment.

Corrosion is an important principle to consider when selecting metallic biomaterials. The chemical reactions that are part of normal human metabolism produce an abundance of oxidizing agents, which creates a destructive environment for metals and alloys. Even the most corrosion-resistant biomaterials will undergo some degree of corrosion. Some metals such as stainless steel decay at a predictable rate, whereas others such as gold and platinum are extremely corrosion resistant. The process entails a coupled oxidation-reduction reaction in which one element gains electrons (oxidizing agent) and the other donates electrons (reducing agent). Most implanted metals, such as titanium, cobalt-chromium, and stainless steel, have a tendency to lose electrons in solution, and as a result, they have a high potential to corrode. The result is dissolution of the metal and the formation of metallic ions. All metals used for human implantation initially corrode and form a thin barrier film. This surface oxidative film offers both a chemical barrier to corrosion and prevents degradation of the deeper metal. If mechanical forces disrupt this layer, the underlying reactive metal atoms become susceptible to corrosion.

Use of a metallic alloy such as cobalt-chromium-molybdenum to minimize the effects of oxidative corrosion unfortunately does not eliminate the concern for wear of the articular surface. Any articulating surface will generate debris secondary to friction. The greater the physiologic load on the device, as well as the greater the ROM, the greater the generation of wear debris. In particular, Hellier and coauthors published the results of determining the absolute and relative wear volume rates of various metal alloys through simulation of an intervertebral disk prosthesis.8 It was found that among all the available alloys, a cobalt-chromium-molybdenum (CoCrMo) alloy generated the least amount of wear debris. It had an average wear volume rate of 0.093 mm3 per million cycles from an arthroplasty device, whereas a titanium alloy containing 6% aluminum and 4% vanadium (Ti6Al4V) had an average wear volume rate of 2.9 mm3 per million cycles. Schmiedberg and colleagues in 1994 used scanning electron microscopy to further define the size and shape of the wear debris fragments generated from an arthroplasty articular surface.7 The fragments from a titanium-aluminum-vanadium surface are extremely rough and irregularly shaped. The size of the fragments ranges from less than 1.0 µm to greater than 30 µm. Fragments from the cobalt-chromium-molybdenum alloy have an irregular polyhedral shape when the alloy was formed from a forged process, but the fragments have a spherical shape ranging between 5.0 and 30 µm when the alloy was formed from a hot isostatically pressed process. Catelas and associates published the results of their study investigating wear debris in metal-on-metal total hip arthroplasty devices.9 The study, published in 2003, demonstrated a significant number of wear debris particles composed predominantly of chromium oxide particles, with estimated loss rates as high as 100 mm3/yr for hip arthroplasty.

Basic Terms for Biomaterials

When selecting materials to use for the manufacture of arthroplasty devices, a number of basic terms should be defined.

Strain is the change in length per unit length when a component is deformed.
Stress is the amount of load per unit area on a component. In general, larger components are used to carry larger loads; stress rather than load is used to normalize the influence of prosthesis size.
Modulus of elasticity is the ratio of stress to strain.
Yield strength is the minimum stress at which a material deforms to the point that if the applied load is released, the material will not return to its original shape.
Malleability is the measure of a material’s ability to be deformed by compression or bending without cracking or rupturing.
Ductility is the measure of a material’s ability to be deformed without breaking. Some materials are relatively brittle, such as ceramics. These materials have low failure strain rates and do not have yield points. Copper has high ductility as evidenced by its ability to be drawn into fine wire.
Toughness is the ability of a material to absorb energy without breaking.
Ultimate strength is the stress at maximum load.
Fatigue strength is the maximum stress that the material can withstand without breaking when subjected to repeat loading.
Corrosion is the gradual deterioration of materials by electrochemical attack. Corrosion not only weakens the implant but also releases metal ions into the body.

Goals for Biomaterials Used for Disk Replacement

1 Preservation of functional kinematics. The device’s design, as well as the biomaterial used, must preserve normal functional joint motion.
2 Preservation of normal biomechanics. The device must not overload adjacent biomechanical structures.
3 Exceed patient life expectancy. The device must outlast the patient.
4 Restoration of normal disk height and foraminal diameter. The device must tension surrounding ligamentous structures for stability while providing indirect foraminal decompression.
5 Integration into surrounding bone. The device must allow complete and permanent bony ingrowth.

Alloys Available for Total Disk Arthroplasty

There are three principal metal alloys used in arthroplasty:

1 Titanium-based alloys
2 Cobalt-based alloys
3 Stainless steel–based alloys

Alloys differ in strength, ductility, and hardness. Each of these factors generally determines the utility of each alloy for different implants. The ability of an alloy to resist corrosion more than any other reason has led to the widespread use of metallic load-bearing implant materials.

Stainless Steel Alloys

All stainless steels are composed of a combination of carbon, iron, chromium, nickel, and molybdenum. The form of stainless steel most commonly used in spinal surgery is designated 316LV based on the American Society for Testing and Materials. The 316 classifies the material as austenitic, the L denotes the low carbon content, and the V is the vacuum under which it was formed. The carbon content must be kept at a low level to prevent accumulation of carbide (chromium carbon) at the grain boundaries. The formation of carbide tends to weaken the material and therefore allows a combination of both corrosion and stress to degrade the material. Molybdenum is added to enhance the corrosion resistance of the metals, whereas chromium allows the formation of chromium oxide on the surface of the metal to resist corrosion. Stainless steel and other implant alloys are surface-treated with substances such as nitric acid to promote the formation of a protective oxide layer. The two aspects of stainless steel that make it most popular are its high ductility and malleability. Additionally, the low cost of stainless steel has been an advantage in manufacturing spinal surgery implants. Unfortunately, the significant amount of metallic artifact generated by stainless steel on magnetic resonance imaging (MRI) and computed tomography (CT) makes this alloy a suboptimal choice.

Cobalt-Chromium Alloys

The two basic constituents of all cobalt-chromium alloys are cobalt, approximately 65%, and chromium, approximately 35%. Molybdenum is often added to improve its mechanical properties. As a result, the dominant implant alloy used for most total joint components is cobalt-chromium-molybdenum (CoCrMo). The corrosion resistance of this alloy is primarily due to the chromium content in the metal, which results in the formation of a protective layer of chromium oxide. These alloys are among the strongest, hardest, and most fatigue resistant of all metals used for replacement.

Titanium Alloys

Titanium alloys were developed in the mid-1940s for the aviation industry and were first used in orthopedics around the same time. Commercially pure titanium and titanium alloy with 6% aluminum and 4% vanadium (Ti6Al4V) remain the two dominant titanium alloys used in total joint arthroplasty. Titanium alloys are particularly good implant material because of their high corrosion resistance in comparison to stainless steel and cobalt-chromium-molybdenum alloys. The stable passive oxide film, primarily titanium oxide (TiO2), protects both titanium alloy and commercially pure titanium components. Titanium-ceramic composites are under development because of their improved metal-on-metal wear resistance in comparison to titanium and its alloys.

Other Materials

Polymers

Polymer materials include polyurethane, polyetheretherketone (PEEK), and polyethylene. Most polyethylene implants are made of UHMWPE, which has a molecular weight of 2 to 10 million g/mol. Typically, total hip arthroplasty devices incorporate a cup made of UHMWPE that articulates with a femoral ball of ceramic or cobalt-chromium alloy. By report, the greatest limitation of this design is wear of the polyethylene, which produces a tremendous amount of wear debris annually. By using chemical and radiation-based techniques to cross-link the polyethylene, it has become possible to improve the resistance of this material to abrasion in joint arthroplasty, thereby increasing its life span.

Surface Coatings

Special surfaces and surface coatings to improve bony ingrowth have become available. These surface enhancements include more traditional roughened titanium, porous coatings made of cobalt-chromium or titanium beads, titanium wire mesh, plasma-sprayed titanium, and newer bioactive nonmetallic materials such as hydroxyapatite or other calcium phosphate compositions. The purpose of surface coatings is to provide another measure of resistance to the osteolytic effect of wear debris.

Imaging Properties of Biomaterials

The imaging characteristics of various materials have been mentioned briefly. Although the specific radiographic characteristics are beyond the scope of this work, it nonetheless remains critically important to visualize both the bony and soft tissue structures of the spine at the level of operated disease, as well as at adjacent levels. CT and MRI are often used to evaluate the adequacy of decompression, determine the completeness of arthrodesis, or investigate the cause of continued radicular complaints. Biomaterials clearly have different imaging properties. For example, polymers such as PEEK are radiolucent and typically contain embedded radiopaque markers. Titanium and ceramic have good MRI qualities. Cobalt-chromium and stainless steel do not image well on MRI because of extensive artifact. As a result, myelography with postmyelography CT is recommended for adequate visualization. A direct comparison of the clarity of cervical arthroplasty devices on MRI concluded that titanium devices, with or without polyethylene, allow satisfactory monitoring of the index and adjacent levels with MRI whereas devices containing metals other than titanium prevent accurate postoperative assessment of the index and adjacent levels by MRI (Tables 266-1 and 266-2).10

TABLE 266-1 Cervical Devices and Materials

image

TABLE 266-2 Lumbar Devices and Materials

Pictures courtesy of Depuy, Synthes, Medtronic Sofamor Danek USA, and Stryker.

DEVICE MATERIAL
Charité

image

Metal-on-polymer

• UHMWPE-cobalt-chromium. A metal ring surrounds the outside of the UHMWPE core so that it can be located on radiographs.
• Spike fixation
• Hydroxyapatite coating

ProDisc-L

image
Metal-on-polymer

• UHMWPE-cobalt-chromium-molybdenum
• Center keel fixation
• Titanium plasma spray coating for bony ingrowth

Maverick*

image
Metal-on-metal

• Cobalt-chromium
• Center keel fixation
• Hydroxyapatite coating

FlexiCore*

image
Metal-on-metal

• Cobalt-chromium-molybdenum
• Spike and central dome fixation
• Titanium plasma spray coating for bony ingrowth

UHMWPE, ultra-high-molecular-weight polyethylene.

* Investigational device limited by federal law to investigational use; not available for sale in the United States.

Normal Biomechanics of the Spine

Rotation and Translation

Rotational movements are movement of the vertebra around an axis. All rotations produce a change in orientation of the vertebra (Fig. 266-1).11

Flexion-extension (bending) is rotation about the x-axis.
Axial rotation (torsion) is rotation about the y-axis.
Lateral bending (side bending) is rotation about the z-axis.
Translational movements are movements of the whole vertebra by the same amount in a given direction. There is no change in orientation of the vertebra. Translation is “gliding” of the vertebra; it rarely occurs by itself but often accompanies rotation and flexion-extension, especially in the cervical spine.

image

FIGURE 266-1 Flexion-extension (A), axial rotation (B), and lateral bending (C).

(Courtesy of Medtronic Sofamor Danek USA.)

The movements of each spinal segment are limited by anatomic structures such as ligaments, intervertebral disks, and facets. Specifically, anatomic structures cause the coupling of motions of the spine; that is, the movements occur simultaneously. Flexion-extension, translation, axial rotation, and lateral bending are physiologically coupled. The exact pattern of coupling depends on the regional variations of anatomic structures.

Center of rotation (COR) is the axis about which the rotation occurs.

The COR for flexion-extension has been found to be located in the posterior third of the inferior vertebral body in both the cervical12 and lumbar13 spine (Fig. 266-2). The COR for axial rotation is in the midline, posterior to the disk but anterior to the facet joints for the lumbar spine14 and aligned with the facet joints in the cervical spine.12 When both translation and rotation occur in the same plane, the COR is variable. Because motions of the spine are typically coupled, they also have a variable COR. In flexion, the COR moves to the anterior side of the spinal column; in extension, to the posterior side; and during lateral bending and axial rotation, to the opposite side of the spinal column.15 A variable COR decreases loads because the working distance of the spinal muscles and ligaments is increased during the motions (Figs 266-3 and 266-4).16,17

image

FIGURE 266-2 Locations of the centers of rotation for the cervical spine as described by Bogduk and Mercer.12 (A) and for the lumbar spine as described by Pearcy and Bogduk (B).13

(A, Reprinted with permission from Elsevier. B, Reprinted with permission from Lippincott Williams & Wilkins.)

image

FIGURE 266-3 Fixed center of rotation (COR). F/E, flexion-extension.

(Courtesy of Medtronic Sofamor Danek USA.)

image

FIGURE 266-4 Variable center of rotation (COR).

(Courtesy of Medtronic Sofamor Danek USA.)

Range of Motion and Its Limits

The FSU is a three-joint complex consisting of the intervertebral disk and the two posterior facet joints. These three joints act together to provide normal ROM and resist supraphysiologic stress. Lumbar facet joints provide resistance to excessive extension and experience high compressive loads in extension. In flexion, they experience minimal load. The facets resist anterior translation because of their orientation, and hence, anterior shear loads increase facet pressure. The posterior ligaments (supraspinous, interspinous, facet capsules) resist posterior translation; in posterior shear loading, the posterior ligaments are placed under tension, whereas the facets are unloaded. Stiffness is the measure of resistance to displacement. In the center of the ROM, the FSU offers little resistance to movement, and large displacement may be caused by small forces. When the FSU gets closer to the limit of its ROM, its stiffness increases and the resistance to further displacement is quite high.18,19 The typical cervical motion segment allows 10 degrees of flexion and extension, 11 degrees of lateral bending, 7 degrees of axial rotation, and 2 mm of translation. A lumbar motion segment typically allows 12 degrees of flexion and extension, 8 degrees of lateral bending, and 7 degrees of axial rotation.

Biomechanics of Total Disk Arthroplasty

Total disk arthroplasty (TDA) refers to devices that replace the majority of the disk while conserving a small portion of the anulus for ligamentous stability.20 Surgical goals are conservation of ROM, restoration of sagittal balance and disk height, and preservation of the facet joints. Because the intervertebral disk consists of concentric layers of interwoven collagen with a proteoglycan gel-filled center, a plausible design for a TDA device would be a thick-walled balloon or a flexible elastomeric device attached to each end plate. A TDA device mimicking the natural disk was indeed designed and underwent clinical trials (Acroflex Lumbar Disc, Depuy Spine, Raynham, MA). However, clinical trials were aborted because of premature mechanical failure of the elastomer and osteolysis caused by debris released from the failed elastomer. In contrast, the majority of currently available TDA devices do not mimic the normal intervertebral viscoelastic disk. They use various forms of low-friction, sliding, gliding, and rotational joints in an attempt to restore normal function to the motion segment.

All TDA devices are capable of flexion-extension, axial rotation, and lateral bending to various extents, but not all TDA devices permit translation. The ability to allow independent translation is the main difference among the TDA devices. Independent translation affects the COR and the location of shear forces. A TDA device with independent translation has a variable COR (as opposed to a fixed COR for a TDA device without translation). As a result, motion of the TDA device follows the natural COR of the FSU and is less likely to overload the facet joints in flexion-extension.

Limits to Range of Motion

A TDA device with independent translation does not resist a shear force. The TDA device translates and the shear force must be resisted by the facet joints. In contrast, a TDA device without independent translation resists a shear force and the shear is instead experienced at the device-bone interface. All current TDA designs lack inherent stiffness. Hence, all TDA devices do not provide an inherent “brake” toward the limits of ROM. However, most TDA devices do have a built-in mechanical stop at the end of their ROM to resist supraphysiologic stress (Tables 266-3 and 266-4).

TABLE 266-3 Biomechanical Properties of Cervical Devices

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TABLE 266-4 Biomechanical Properties of Lumbar Devices

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Consistency between the Natural Center of Rotation and the Center of Rotation of a Total Disk Arthroplasty Device

If the COR of a TDA device does not match the natural COR, soft tissues may be made to lengthen or shorten and facet joints to impinge or distract. The exact location of the COR will depend on both the specific design of the device and positioning of the implant within the disk space during surgery. A variable COR device is less sensitive than a fixed COR device to surgical placement.21 Although spinal movement follows a variable COR, a fixed COR device may be preferred in some situations in which the ability to resist shear is important, such as with a steeply angled disk space or when pars defects are present.20 To this point, spondylolysis has been considered a contraindication to implantation of TDA devices.

The location of the COR is crucial for proper biomechanics, but its measurement remains difficult. In vitro studies are able to measure COR with high precision but they rely on simplified loads and loading influences the location of the COR. In vivo measurements are approximations taken from two-dimensional radiographic studies.22 These difficulties may explain divergent COR localizations.

Adjacent Segments

The conservation of motion allowed by TDA devices is theorized to prevent degeneration of the adjacent spinal segments. The adjacent segments have normal kinematics after TDA, in contrast to the situation after fusion.23 However, there is inconclusive evidence concerning the relative rates of degeneration of adjacent segments in nonoperated and operated patients.24,25

Shock Absorption Properties

A major function of the natural intervertebral disk is shock absorption. The intervertebral disk takes up most of the axial weight of the trunk and upper extremities and distributes it between the vertebral bodies. “Soft” biomaterials (polymers) are believed to have greater shock absorption capacity than “hard” biomaterials (metals). In vitro studies demonstrate some energy absorption ability with a polyurethane sliding core26 but not with UHMWPE ball-and-socket designs.27 The ability of current TDA devices to absorb shock adequately at physiologically significant levels remains a matter of current debate.

Conclusion

Current spinal TDA designs are not meant to mimic the natural intervertebral disk but rather to approximate normal movement patterns and bear spinal loads. Differences in biomaterials will clearly influence both the in vivo behavior and longevity of an arthroplasty device. Current clinical outcomes are probably a result of the interaction between unique patient characteristics, the specific design features of the TDA device, and the implant surgical protocol. Further study is necessary to determine which biomaterials and biomechanical design features (COR, shear resistance, facet loading) are optimal and which surgical protocols (implant positioning, sparing of ligaments, distraction) are predictors of clinical success.

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Sears WR, McCombe PF, Sasso RC. Kinematics of cervical and lumbar total disc replacement. Semin Spine Surg. 2006;18:117-129.

Sekhon LHS, Duggal N, Lynch JJ, et al. Magnetic resonance imaging clarity of the Bryan, Prodisc-C, Prestige LP, and PCM cervical arthroplasty devices. Spine. 2007;32:673-680.

White AA, Panjabi MM. Clinical Biomechanics of the Spine. Philadelphia: JB Lippincott; 1990.

References

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20 Sears WR, McCombe PF, Sasso RC. Kinematics of cervical and lumbar total disc replacement. Semin Spine Surg. 2006;18:117-129.

21 Moumene M, Geisler FH. Comparison of biomechanical function at ideal and varied surgical placement for two lumbar artificial disc implant designs. Mobile-core versus fixed-core. Spine. 2007;12:1840-1851.

22 Crawford NR. Biomechanics of lumbar arthroplasty. Neurosurg Clin N Am. 2005;16:595-602.

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