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CHAPTER 38/DYNAMIC STABILIZATION / 383
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CHAPTER 39
Lumbar Artificial Disc Replacement: Rationale and Biomechanics
Geoffrey M. McCullen and Hansen A. Yuan
Stability of the vertebral motion segment is dependent
upon the structural integrity of the disc, facet joints,
and the musculoligamentous complex. Intervertebral
discs transmit 85% of the axial load and provide rotational/translational support and shock absorption (1,2). In
a standing position, the intradiscal pressure at L3-4 is
1,000 N, increasing to 3,000 N when sitting, leaning forward, or carr ying 20 kg (3,4). The estimated number of
walking cycles a person performs is approximatel y 2 million per year (5). Each walking cycle loads the disc 150
to 1,250 N in compression and −40 N to +450 N in shear
(5). The estimated number of lifting c ycles is 125,000 per
year, each producing an estimated compressive load
between 200 and 2,250 N (5).
The magnitude and frequency of such imposed loads
on the spinal motion segment will inevitably lead to wear
and degenerative changes over time. Three sequential
clinical and biomechanical stages of spinal degeneration
have been proposed: dysfunction, instability, and stabilization (6,7). Segmental motion increases with the severity of disc degeneration (8) and facet degeneration most
typically follows (9,10). Segmental axial rotation motion
increases with the degree of cartilage degeneration within
the facet joints (8). Stabilization occurs with advanced
disc collapse associated with marked end-plate sclerosis
and osteophytosis.
RATIONALE FOR THE ARTIFICIAL DISC
Spinal arthrodesis signif icantly alters the biomechanics of the spine and results in an increased compensatory
motion and mechanical loading in the adjacent free
segments, particularly within the facet joints (11–15).
Changes in the adjacent segments become more pronounced as the fixation range extends and as the rigidity
of the construct increases (13,16,17). A 45% increase in
intradiscal pressure occurs adjacent to a fusion (17,18).
Elevated intradiscal pressure acts to alter the diffusion
characteristics of nutrients into the disc (19,20) and leads
to biochemical changes including an elevation of type 1
collagen and a decrease in proteoglycan chondroitin sulfate, type 2 collagen content, and water content (2,21).
Single photon emission computed tomography imaging
4 years after lumbar fusion demonstrates 62% with
increased uptake within the vertebral bodies and facet
joints in the free motion segment adjacent to the fused
segment (22). Symptomatic juxtafusion degeneration is
estimated to occur in 35% of postfusion patients (23).
Discectomy adversely affects the mechanical properties of the segment and alters coupled motions within
adjacent segments (24,25). The level above a discectomy
experiences an increased anterior-posterior translation in
flexion and increased lateral translation with lateral bending irrespective of the amount of disc removed (24).
Removal of the nucleus from the disc causes the inner
margins of the annulus to bulge inward rather than outward during loading causing delamination of the fibers
and further degeneration (26–28). Subtotal discectomy
induces significantly less abnormal motion than total discectomy in all loading modes (24,25).
The aim of an artif icial disc is to decrease pain, provide stability while maintaining near-ph ysiologic motion,
and diminish the likelihood of adjacent-level degenerative cascade. Two artificial disc interventions have
evolved: (a) a total disc replacement (nucleus and annulus) and (b) a nucleus substitute. Many designs hav e been
suggested. Few devices have been analyzed in biomechanical or clinical studies.
TOTAL DISC REPLACEMENTS
Total disc replacements are used to treat degenerative
disc disease when removal of all possible sources of
discogenic pain, including the annulus, is desired. The
384
CHAPTER 39/LUMBAR ARTIFICIAL DISC REPLACEMENT / 385
elective indications for a total disc replacement are single- or double-level degenerative disc disease, juxtafusion degeneration, and postdiscectomy axial pain that
have failed a significant trial of nonoperative care.
Absolute contraindications include: local or systemic
presence of tumor or infection; osteoporosis (dual energy
X-ray absorptiometry T-score less than −2.5); obesity
(body mass index: weight (kg)/height (m) greater than
40); overlying thoracolumbar kyphosis; spondylolisthesis
(greater than grade 1); disabled posterior elements unable
to contribute in load-sharing with the prosthesis; and
facet osteoarthritis and facet-mediated pain patterns
(Table 39-1).
The natural spine triple joint complex provides six
degrees of freedom: (a) compression, (b) distraction, (c)
flexion, (d) extension, (e) lateral bending, and (f) rotation.
For the total disc replacement, the degree of “constraint”
refers to the relative range of motion of the prosthetic joint
compared with the healthy intact joint within each of these
degrees of freedom (5,29). An “unconstrained” device
provides no mechanical restraint or limit. An “under-constrained” device imposes limits to motion outside of the
naturally occurring constraint of motion. With a “critically” constrained implant, motion is allowed within the
physiologic range but blocked beyond. The “over-constrained” prosthesis prohibits natural motion by imposing
limits within the normal range of motion.
Many current total disc devices are rotationally unconstrained and require intact posterior elements (facets and
musculoligamentous complex) for load sharing and biomechanical stability. Over time, such a device risks overloading, injury and degeneration of the posterior elements, particularly the facets. Alternatively, designs with
rotational over-constraint would off-load the facets at the
expense of increasing stress at the implant-host interface.
In addition, the majority of current devices have high
compressive stiffness (over-constraint in axial compression) resulting in decreased shock-absorbing capacity
and increased risk of vertebral body fracture with significant loading.
All total disc replacements are implanted through anterior open approach (transperitoneal or retroperitoneal).
During implantation, the anterior longitudinal ligament
TABLE 39-1. Biomechanical contraindications for the
artificial disc
Osteoporosis
Obesity
Significant deformity
Spondylolisthesis (> grade 1)
Disc height <5 mm
Facet incompetence, facet mediated pain
Posterior ligament compromise
Additional contraindications for nucleus replacement:
Incompetent anulus
End-plate sclerosis, Schmorl nodes
and the anterior annulus are excised. These two anterior
ligamentous structures provide a balanced tensile resistance to rotation on the opposite side of the center of rotation from the facets (30). Methods of retaining or reconstructing the anterior longitudinal ligament may lead to
restoration of normal load sharing and segmental stiffness (30). Ligament tensioning is the key to restoring
segmental sagittal balance. Restoring the “normal” disc
height can cause over-stretching of the spinal ligaments
that would limit ultimate segmental motion. To avoid
over-distraction, it is advisable not to increase the height
of a degenerated disc by more than 3 mm (31).
The total disc implant requires support from the
peripheral cortical shell of the vertebral body. The endplate periphery is stronger while the center is weaker
(32,33). The device end-plate “foot print” should cover
the largest possible area of the vertebral end plate for
optimal load distribution and to decrease the formation of
heterotopic ossification occur ring at the end-plate margins.
Improvements in material and design specifications
are leading slowly toward the ideal artificial disc (Table
39-2). Material properties to be considered are cytotoxicity, fatigue strength, and the modulus of elasticity. The
design must attempt to approximate the natural disc
dynamics, plan for compatibility between materials, use
safe insertion technique, and have a reliable fixation
between host and implant.
Metal Devices
The primary advantage for using an all-metal total disc
replacement is the optimal fatigue strength. Biocompatibility has been shown in other orthopedic applications
including joint arthroplasty. Designs include the springs
with hinge and the ball-in-socket.
Kostuik has developed a device that uses two springs
coiled between plates with a posterior hinge allowing
flexion and extension (34) (Fig. 39-1).This device allows
15° to 20° of axial rotation in the sagittal plane, 3° of lateral bending, and less than 1° of axial rotation (5). The
hinge pin has a decreasing diameter as the ends of the pin
are approached. This feature, combined with elongated
holes in which the pin rotates, is intended to allow a small
amount or rolling lateral rotation (3° to 6°) (5). The
springs sit in pockets that are designed to minimize offaxis loading and provide for a mechanical stop to prevent
the springs from being loaded to the point that the coils
touch (5). The springs are made of titanium alloy and are
designed to provide adequate stif fness in both fle xion and
extension with a combined resistive torque in sagittal
plane rotations of 2.24 N per degree (5). The springs have
been successfully tested to more than 100 million full
deflection cycles without failure (5). Hot isostatically
pressed cobalt chrome alloy, with high carbon content to
improve wear characteristics, is employed in the remain-
386 /SECTION V/SPECIFIC CLINICAL ENTITIES
TABLE 39-2. Total disc device comparison
Kostuik Salib Lee nonmetal Mar nay metal, Buettner-Janz metal,
metal metal elastomer polyethylene polytheylene two
spring-hinge ball-socket composite cup-cap moving articulations
Shock absorption Yes No Yes No No
Axial rotation Over constrained Unconstrained Critically Unconstrained Unconstrained
Translation Over constrained Over constrained Unknown Over constrained Critically constrained
Flexion-extension Under constrained Under constrained Critically Critically Under constrained
Center of rotation Fixed Fixed Moves Fixed Moves, replicates
Potential Galvanic Metal-on-metal, Material Polyethylene, Polyethylene, wear
problems corrosion, wear delamination, wear
bulky, difficult, Device-host
insertion fixation
constrained
constrained constrained
instantaneous
axis of rotation
der of the device. Vertically projecting tabs are placed at
the front and side of the plate members through which
screws are placed for fixation. Using different metals,
galvanic corrosion is a concern but minimal corrosive
issues have been seen in short duration animal models
(5). In a sheep model, with histologic analysis at 6
months, fibrous tissue does not g row between hinges or
around coils (35). If such an in-growth were to occur, it
would be expected to significantly interfere with the disc
mechanics.
Reminiscent of hip arthroplasty, a prosthetic device
shaped as a ball and socket has been proposed (36) (Fig.
39-2) and has subsequently been redesigned. The Maverick (Medronic, Sofamor Danek, Memphis, TN) is made
entirely of cobalt chrome and is axially rigid. There is no
inherent rotational stiffness (unconstrained). Finite element analysis (FEA) has predicted that surgical variables
will modulate the loads transmitted through the posterior
elements after implantation (30). When the Maverick is
placed anteriorly, the FEA predicts that facet loads will
increase 2.5 times. A Maverick implanted posteriorly
within the intervertebral space will successfully unload
the facets in axial compression.
Nonmetal Devices
Whereas the principal advantage of metals is fatigue
strength, the primary benef it of using nonmetals such as
rubber and the elastomers (silicone, polyurethane, polyethylene) is the biomechanical similarity to the natural
disc (37). By using a material with a lower modulus of
elasticity, it is easier to replicate disc dynamics, attenuate
shocks, and distribute loads evenly across the end plates.
Difficulties, however, arise when attempting to develop a
long-lasting nonmetal component that has a stable hostdevice interface. In addition, a polymer disc prosthesis
that has optimal rigidity in axial compression may possess insufficient torsional rigidity. Alternatively, a polymer disc that is suitably rigid in torsion frequently
becomes too rigid axially (38). P olymers do not allow the
defining of a stable center of rotation for a disc implant.
Lastly, delamination commonly occurs between polymer
materials.
In 1975, Stubstab and Urbaniak introduced a total disc
constructed entirely of synthetic nonmetal materials (39).
Silicone or other elastomers such as polyurethane were
formed into the general kidney shape with a fluid-f illed
FIG. 39-1. Kostuik’s all-metal device with two springs coiled
between plates with a posterior hinge allowing flexion and
extension (34).
FIG. 39-2. All-metal ball-and-socket design by Salib (36).
CHAPTER 39/LUMBAR ARTIFICIAL DISC REPLACEMENT / 387
core between flat superior and inferior elastomeric plates.
Simulating the annulus, the device was covered by a
weave of Dacron fibers creating pores for tissue
ingrowth. This device has been implanted into eight
chimpanzees followed for one year (40). Erosion and
reactive end-plate bone formation have been attributed to
an unstable host-implant interface.
Lee and Parsons (41,42) have designed what has been
referred to as the “New Jersey disc” (Fig. 39-3). Two
material designs have been described. In the first, a soft
elastomer central core is surrounded by a fiber-reinforced
polyurethane (12 layers of Dacron with alternating multidirectional pattern: −45, 0, −45). In the second, the core
is surrounded by a nonf iber reinforced polymer, C-flex
(polysiloxane-styrene-ethylene-styrene-butylene) (Concept Polymer Technologies, Clearwater, FL) (38,43). Two
stiff end plates using elastomer or hydroxyapatite frame
the disc. This disc has been studied both in vitro and in
finite element models (43,44). Questions remaining for
the New Jersey disc include: material biocompatibility,
fatigue, and wear resistance; and the adequacy of shortand long-term implant f ixation.
The “3-DF” disc is a three-dimensional fabric woven with an ultra–high-molecular-weight polyethylene
(UHMWPE) fiber and a surface spray-coated with bioactive ceramics (45,46). Made of one material, the device
avoids a composite interface. No changes in biomechanical parameters have occurred after 2 million loading
cycles with compressive loading of 200 N (45). Creeprelaxation testing has demonstrated a viscoelastic straintime curve almost identical to the normal intervertebral
disc (45). The “3-DF” de vice has been studied using an in
vivo sheep model (46). Highlighting the problems inher-
FIG. 39-3. The New Jersey disc designed by Lee et al (41).
A soft, central elastomer core is surrounded by either a fiberreinforced polyurethane or a nonfiber- reinforced polymer.
Two stiff end plates using elastomer or hydroxyapatite frame
the disc.
ent in animal models, complete removal of the end plates
was required in order to fit the 10 mm artificial disc. This
created a different implant-host interface than is clinically intended. The segmental biomechanics and interface histology were evaluated at 4 and 6 months. Implant
displacement without complete dislodgement was noted.
Those devices implanted with temporary internal fixation
(Kanada one-rod Smooth Rod system) were firmly in
place and demonstrated a successful bony bonding to the
host. Work is now being directed toward the creation of a
supplemental bioabsorbable spinal fixation device.
Combination Metal and Nonmetal Devices
To overcome the shortcomings found when using metals or nonmetals alone, designs have combined the materials. Most commonly, this has taken the form of a metalpolymer-metal sandwich disc. The metal tray is emplo y ed
to improve fixation with spikes, tabs with screws, or
porous coating for bony ingrowth. With the component
thus stable and fixed, the central polymer provides the
needed flexibility.
In 1991, Steffee developed the Acroflex disc (DePuy,
Acromed Corporation, Cleveland, OH) (47). This disc
replacement consists of a hexane-based, carbon black–
filled, polyolefin rubber core vulcanized to the two titanium plates. Fixation is accomplished with a porous coating promoting ingrowth and four 7-mm cone-shaped
posts that extend into the vertebral body. A significant
subsurface shear is experienced at the rubber-metal junction producing failures with fracture of the rubber core
(48).
The ProDisc (Spine Solutions, Inc., New York, NY) is
a cap-cup matching articulation designed by Marnay (49)
(Fig. 39-4). The end plates are made of a cobalt chrome
molybdenum alloy with a central fixation keel projecting
through the vertebral end plates. The convex bearing surface is made of UHMWPE that snaps into the inferior end
plate. The de vice has a fixed center of rotation and is critically constrained in flexion-extension, unconstrained in
rotation, and over-constrained in translation.
Buettner-Janz and Schellnack developed the Link SB
Charité disc (Waldemar LINK GmbH and Co., Hamburg,
Germany) (Fig. 39-5) (50). Since the first design in 1984,
three revisions have been made. The current device (Link
SB Charité III) includes two symmetric end plates made
of cobalt chromium alloy. Immediate fixation is by
anchoring teeth projecting from the outer end plate. For
bony ingrowth, the end plates are porous coated with two
layers of thin sintered titanium beads (pore size ranging
from 75 to 300 microns) covered by electrochemically
bonded hydroxyapatite. A primate model has demonstrated ingrowth over 48% of the end-plate surface (51).
The inner end-plate surface contour is an oval-shaped
concave surface that articulates with a central high-density polyeth ylene core [modulus of elasticity 420 to 1,200
388 /SECTION V/SPECIFIC CLINICAL ENTITIES
A
A
B
C
FIG. 39-4. The Marnay device (49). A: The polyethylene
cups snaps into the inferior end plate. The device is unconstrained in axial rotation. B, C: The device in flexion and
extension with movement occurring around a fixed center of
rotation.
mega Pascals (MPa)] (52). The device allows 20° of
motion in flexion-extension and lateral bending (underconstrained). There is no limit to motion in axial rotation
(unconstrained) (53). Hysteresis (the conversion of strain
energy into heat formation during cyclic loading) of the
polyethylene occurs at loads less than 4.2 kN with incipient irreversible polyethylene deformation occurring with
loads between 6 and 8 kN (53).
With movement in flexion and extension, the polyethylene core moves: in flexion the polyethylene moves posteriorly and in extension it moves anteriorly. This changing of the instant axis of rotation (IAR) is meant to
replicate the natural disc (52). Because of the changing
IAR, the Link SB Charité disc needs to be placed poste-
B
C
FIG. 39-5. The SB Charité disc (50). A: Both top and bottom
inner end-plate surfaces articulate with the central polyethylene. With flexion, the polyethylene moves posteriorly. B: In
extension, the polyethylene moves anteriorly, changing the
instantaneous axis of rotation. C: The device allows 20° of
motion in flexion-extension and lateral rotation and is unconstrained in axial rotation.
riorly within the intervertebral disc space. Anterior positioning of the device will significantly decrease the available motion (52,54).
In a human cadaveric biomechanical study, the Link
SB Charité prosthesis showed an average percentage
increase in segmental axial rotational range of motion of
44% when compared to the intact control (55). The
increase in axial rotation reflects the excision of the anterior longitudinal ligament required during insertion and
the unconstrained nature of the disc in axial rotation.
Flexion and extension range-of-motion between the in-
CHAPTER 39/LUMBAR ARTIFICIAL DISC REPLACEMENT / 389
TABLE 39-3. Factors affecting in situ total disc device
motion
Preoperative segmental stiffness
Facet and ligament degeneration, end-plate osteophytes
Device positioning
Anterior positioning decreases motion
Device sizing
Over-stretching ligaments decreases motion
Under-sized “foot print”
Device subsidence
Heterotopic ossification formation
Bracing
Best to avoid
tact specimen and a segment implanted with the Link SB
Charité disc were not significantly different. The disc
preserved the normal mapping of segmental motion at the
operative level and at the adjacent levels.
Clinically, postoperative segmental motion is equal to
preoperative motion (54,56). The motion permitted by the
implanted prosthesis appears to be determined by the
extent of the degenerative changes in the adjacent structures such as the facet joints and neighboring spinal ligaments. The a v erage postoperati v e range-of-motion for the
SB Charité III is 9° of flexion and extension (52,54,56).
At the L4-5 level 9.6° flexion, 3.4° extension, and 4° lateral bending motion has been reported (52). At L5-S1
these values are decreased with 6° flexion, 3.5° extension, and 3.2° of lateral bending (52). Clinical studies
confir m that anterior device positioning decreases postoperative motion (54). Over-distraction causes excessive
tightening of the posterior ligaments and decreases segmental motion. To obtain and maintain available motion,
early postoperative motion is encouraged (54) (Table 39-
3).
In published series, further surgery has been required
in up to 24%, the majority for “pain control” as a result
of ongoing or developing facet-mediated pain (56).
Device dislocation/migration occurs in 2% to 6.5%, usually secondary to poor size selection (54,57). Eight percent of the devices have subsided most commonly
because of an undersized device “footprint” (54).
NUCLEUS REPLACEMENTS
There are two potential indications for nucleus replacement: (a) as a adjunct to discectomy (open and percutaneous nucleotomy), addressing the biomechanical alterations of discectomy; and (b) to relieve back pain
resulting from early stage degenerative disc disease that
has failed to improve with nonoperative means.
Nucleus substitutes should be made of biocompatible
materials; restore intervertebral height and tension the
annulus with normal end-plate load distribution; recreate
the intradiscal hydraulic pumping mechanism for pressure modulation and nutrient delivery to the remaining
nucleus and the inner annulus; and demonstrate safe, possibly minimally invasive implantation with a revision
strategy in cases of failure.
Nucleus replacement requires the satisfactory condition of the annulus and the vertebral end plates. Sufficient annular containment is essential to stabilize the
device in the disc space and to avoid excessive motion
that could lead to migration and an elevated rate of wear
on end-plate or implant surface. When end plates are
weakened by osteoporosis or Schmorl nodes or challenged by obesity (weight greater than 90 kg or body
mass index greater than 30) subsidence of the implant
into the vertebral body can occur.
Successful tensioning of the annulus will depend upon
the implant modulus of elasticity and the implant/cavity
conformity. The inward bulging of the annulus after
nucleotomy is re v ersed by a nuclear implant with Young’s
modulus of elasticity of 3 MPa (range 0.2 to 40 MPa)
(27). Restoration of intervertebral disc height with a
nucleus replacement can best be accomplished if the
starting disc height is greater than or equal to 5 mm. In
late stage disc disruption, with failure of the annulus, end
plate and facets, a nucleus replacement would certainly
be ineffective. Annulus repair and methods to achieve
annulotomy healing remain elusive.
Fernstrom developed the first nucleus replacement in
1966 (58). These devices were spheric, solid, stainless
steel balls, up to 16 mm in diameter that were meant to
serve as intervertebral spacers while allowing movement.
Contact area with the vertebral bodies was small. No
pressure modulation was possible with position change.
The device did not restore the normal load distribution
across the end plate and was ultimately abandoned
because of implant migration or subsidence.
In the 1960s, work began with the injection of self-curing silicone and polyurethane into the disc space (3). The
potential advantages of this technique were the minimall y
invasive, percutaneous route of insertion resulting in
reduced annular injury and the in situ polymer curing
that offered the possibility of obtaining good end
plate–implant conformity. The problems encountered
included: long polymerization reaction times; incomplete
polymerization; exothermic reaction temperatures; high
injection pressures; polymer containment; and the production of particulate wear debris (59). Subsequent
efforts in nucleus replacement have been directed toward
developing contained systems that allow a more reproducible fabrication and accurate prosthetic placement.
In 1988, Ray introduced the Prosthetic Disc Nucleus,
or “PDN”, (Raymedica, Minneapolis, MN) with dual disc
cylinders resting side-by-side (60). In the original procedure, a trephine was used, from the posterior approach,
to bore a passage into the center of the disc. After
nucleotomy, the cylinders were slid into position and
filled with a water-absorbing gel held within a semipermeable membrane. Prefilled devices have subsequently
390 /SECTION V/SPECIFIC CLINICAL ENTITIES
replaced in situ fill of the cylinders. The outer device
layer of the PDN is constructed with a flexib le but inelastic polyethylene f iber jacket that constrains the ultimate
swelling of the hydrogel.
The PDN implants have passed U.S. Food and Drug
Administration guidelines for biologic safety, cytotoxicity, carcinogenicity and long-term animal implantation.
Fatigue studies to greater than 50 million cycles under
simulated physiologic loads (200 to 800 N) have been
successful (61). After nucleotomy in vitro insertion of the
PDN has restored the segmental mobility in all movement/directions (62).
The first PDN clinical trials began in 1996. Initially,
the PDN used a hydrogel polymer, HYPAN 68 (polyacrylonitrile-polyacrylamide), which absorbs 68% of its
weight in water. Early analysis indicated that this device
was too rigid, promoting pressure concentrations on the
cartilaginous end plate (61). In 1998, a change was made
to HYPAN 80, which absorbs more water and thus is
softer. A higher rate of implant migration with a 38%
revision rate was recognized (61). Further changes in
design and surgical protocol followed. PDN devices are
now two differently shaped units: a tapered anterior unit
and a rectangular posterior unit placed transversely and
connected by tethering sutures (Fig. 39-6). To accommodate the two units the end-plate anteroposterior diameter
must be 37 mm or greater (61).
For device containment, the annulotomy must be kept
as small as possible with a sequential annular dilating
technique rather than a cut entry. Using lateral decubitus
positioning and a lateral retroperitoneal exposure,
A
Bertagnoli has developed the anterolateral transpsoas
approach (ALPA). The annulus is cut to create a hinged
flap. The overlying psoas muscle covers the lateral annulotomy site, diminishing the risk of implant extrusion
(63). A restricted postoperative protocol with the use of a
brace for 6 weeks is recommended (63).
Moderate to severe end-plate remodeling, likely due to
changes in load distribution, has been noted in some
patients implanted with the PDN (63). Sixty patients with
PDN have been followed with postoperative MRI (64).
Six weeks postoperatively there is an increase in degenerative end-plate changes in 27% of the implanted levels
(64,65). At one year, the incidence of such changes
increases further to 70%. Advanced de generati ve changes
within the end plate at 1 year are seen in 7%. The clinical
implications of these radiographic changes have yet to be
determined.
Other nucleus replacements are emerging in various
stages of early development. Bao has designed a polyvinyl alcohol (PVA) hydrogel nucleus replacement in the
form of beads held within a retaining membrane (66)
(Aquarelle Hydrogel Nucleus, Stryker/Howmedica, Mahwah, NJ). It is inserted in the dehydrated form, through a
5 mm cannula. Water, drawn from the surrounding tissues, fills the membrane to a 70% water content and the
prosthesis nearly doubles in size. This in situ expansion
creates an interference fit. Like the natural nucleus, the
hydrogel replacement is able to absorb and relinquish
water with changes in applied load. Biomechanical studies have confirmed restoration of disc function after
hydrogel implantation (67,68). Pre v enting implant migration and extrusion remains a problem. The Newcleus
(Sulzer SpineTech, Edina, MN) is an elongated, elastic
memory-coiling, spiral-shaped polycarbonate urethane
(69,70). It is inserted into the disc through a cannula
device to create a spiral coil of the material. The Prosthetic Intervertebral Nucleus, or “PIN” (Disc Dynamics,
Minnetonka, MN) is a polyurethane balloon inserted in
deflated form into the disc and then inflated with an
incompressible liquid (71). The size and shape of the
nucleus prosthesis can be significantly altered during
implantation and regained or restored after implantation.
B
FIG. 39-6. The PDN devices (62). A: Sagittal view. B: Axial
view demonstrating a tapered anterior unit and a rectangular
posterior unit placed transversely and connected by tethering sutures.The annulus and end plates must be in satisfactory condition.
DEVICE LONGEVITY
The generation of particulate debris can occur as a
result of wear and corrosion and should be expected any
time an artif icial disc implant is used. The accumulated
magnitude of the debris over time and the resultant clinical sequelae after disc replacement has yet to be determined. In long-term patient outcomes for total knee and
hip replacements, microscopic wear fragments have disseminated to the liver, spleen, and paraaortic lymph
nodes (72). Local wear debris can incite a local cytokinemediated [tumor necrosis factor-α, prostaglandin-E2,
interleukin (IL)-1, IL-2 or IL-6] immunologic reaction
CHAPTER 39/LUMBAR ARTIFICIAL DISC REPLACEMENT / 391
that causes destruction of bone at the prosthetic-bone
interface (73). The optimal particle size to disseminate
and to incite macrophages to release cytokines is 1 to 5
microns (72,73).
Among the metals implanted in the spine, titanium
debris fragments appear to be the most reactive in eliciting a cytokine-mediated response (73,74). Polyolefin
rubber compound particles have been generated in vitro
and placed in subcutaneous pouches in rats and adjacent
to the lumbosacral dura in sheep (75). A local foreign
body reaction was created but no particle migration from
the site of implantation was identified and no apparent
local or systemic toxic effects resulted.
In a nonhuman primate study with the Link SB Charité III, histochemical assays showed no accumulation of
particulate debris (no titanium, UHMWPE, or cobalt
chrome) and no cytokines 6 months after implantation
(51). The lack of long-term evaluation or anatomic differences including the absence of synovial fluid in the
disc may be the reason why particulates have not been
generated.
The SB Charité disc has more than 10 years of clinical
use but no long-term clinical outcome studies. Neither
osteolysis nor late loosening has been reported in clinical
studies. It is a possibility that some cases will be identified with longer duration follow-up. Repeat anterior
retroperitoneal approaches are fraught with potential hazard because of the development of a thick scar enveloping the spine and vascular structures. Therefore, when
total disc devices fail, the safest revision strategy is a posterior spinal instrumented fusion.
CONCLUSIONS
With innovations in materials and design, the artif icial
disc is becoming a reality. Optimal short-term clinical
results will depend upon patient selection.
Poor clinical results are expected when disc replacement is applied to advanced degeneration (76). Longterm outcome will depend upon the ability of the device
to replicate natural kinematics and to withstand the repetitive loads over time. Long-term clinical evaluations after
disc replacement will be essential to identify the rate and
extent of progressive degenerative change within the
facet joints and the adjacent segments.
For accurate comparability between developing devices, biomechanical terminology and methodology for
the artificial disc must become standardized. The optimal
device constraint for particular clinical situations must be
determined.
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