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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_6011_Библиотеки_им_академика_М_И_Перельмана.pdf
C H A P T E R 6 8 Guided Lumbar Interbody Fusion
455
stabilizing a ship’s mast, during activities such as lifting, by using compressive loading and bilateral activation.
8
Thus, preserving the psoas muscle
function can be related to better spinal stability and protection.
Regarding direction of approach, Bergey reports that a “left-sided
approach to the surgery is preferred to a right-sided approach because it
is easier to dissect the aorta off the spine than to dissect around the more
friable inferior vena cava.” Bergey concludes in his study that the advantage
of a lateral transpsoas approach over a standard anterior transperitoneal
approach is in the ease of access to the upper lumbar spine (L1-L4), a place
where anterior techniques are frequently complicated by the location of the
great vessels. As the indications of minimally invasive lateral approaches
expand to treatment of coronal plane deformities, other factors play a larger
role. The concave side of such curves usually allows convergence to one point
and minimizes the number of incisions needed to approach more than one
level. However, the disc is being approached from the collapsed side, and it
will take more work to access the disc. Once the annulotomy is made from
the side of the collapse, theoretically it should allow the maximal release and
optimize the correction.
Bergey expressed concern about a 30% incidence of transient groin/
thigh pain, these symptoms being consistent with the cutaneous innervation of the genitofemoral nerve. As a result, Bergey recommends staying
in the anterior one third of the psoas muscle to avoid nerve root injury.
Visualization and protection of the genitofemoral nerve should avoid
permanent paresthesias in the anterior thigh.
2
The genitofemoral nerve
mainly branches from the L1 and L2 nerve root, travels through the psoas
major muscle anteriorly, and descends along with the abdominal surface of
the psoas major. Moro et al. report that the level where the genitofemoral
nerve passes through the psoas major muscle ranges from the cranial third
of the L3 vertebral body to the caudal third of the L4 vertebral body.
It then descends on the surface of the psoas muscle, normally under the
cover of the peritoneum, and divides into the genital and femoral branches.
The genital branch passes outward on the psoas major and pierces the
fascia transversalis or passes through the internal abdominal ring. It then
descends along the back part of the spermatic cord to the scrotum, and
supplies, in the male, the cremaster muscle. In the female, it accompanies
and ends in the round ligament. The femoral branch of the genitofemoral
nerve descends on the external iliac artery, sending a few branches to it
and, after passing beneath the Poupart ligament to the thigh, supplies the
skin of the anterior aspect of the thigh about midway between the pelvis
and knee.
2
Moro states that the risk of injuring the genitofemoral nerve
increases when splitting the psoas major muscle at lower levels. However,
there are reports of having succeeded in remitting the symptoms without
a serious problem.
9
Moro’s anatomic study of the lumbar plexus with respect to retroperitoneal endoscopic surgery clarifies the safety zone of the psoas major muscle
during retroperitoneal endoscopic surgery using cadavers. Lumbar spines
were removed from embalmed cadavers, and from L1-L5, each specimen
was cut in parallel with the lumbar disc space and the lumbar vertebra at the
cranial third and caudal third of each lumbar vertebral body. The distribution and relationship of the lumbar plexus and nerve roots was analyzed
using computer images. Moro reports, from the results of the study, that the
safety zone may be at L2-L3 and above, due to the presence of the genitofemoral nerve between the cranial third of the L3 vertebral body and L4-L5.
If the possibility of damaging the genitofemoral nerve is not considered, the
safety zone should be at L4-L5 and above. Moro recommends starting from
the abdominal edge of the vertebra when spreading the psoas major muscle
at L2-L3 and below, because nerves are not located in the abdominal surface
of the vertebra. Because the lumbar plexus and nerve roots were wholly contained within the psoas major muscle, one can safely split between the psoas
major muscle and vertebral body to protect the nerves (retract posteriorly).
Moro continues, stating that the method for retracting the psoas muscle
anteriorly and reaching to the lateral surface of the vertebral body may be
useful; however, according to the present study, it is the danger zone where
the lumbar plexus and nerve roots are located in the center of the vertebral
body and dorsally. Additionally, at the L5-S1 level, there are the L4 nerve
root, L5 nerve root, femoral nerve, and obturator nerve between the psoas
major muscle and the lumbar quadratus muscle. Therefore, those nerve tissues must be checked and protected with the endoscope or the alternative
transperitoneal approach should be considered.
9
ADVANTAGES/DISADVANTAGES OF GLIF
Advantages
1. Access allows addition of posterior fixation and posterior releases without rotating patient, to reduce procedural time, anesthesia, and blood
loss, which can also help in driving down costs
2. Delivers an implant directly across anterior column, engaging the
apophyseal ring without requiring drastic repositioning
3. Allows for a larger implant to be delivered with a larger amount of bone
graft
4. Protects / avoids mobilization of the great vessels
5. Retroperitoneal approach decreases incidence of injuring the peritoneum
6. Preserves the spine’s natural stabilization elements:
○ Anterior longitudinal ligament
○ Posterior longitudinal ligament
○ Posterior elements (facets / lamina)
7. Allows direct visualization, via hinged portal, allowing the surgeon to
easily confirm anatomy and ensure soft tissue and nerves are protected
8. Provides alternative approach for revision procedures
Disadvantages
1. Initial trajectory is defined using blind dilator techniques, although
neurostimulation can be used
2. A learning curve is associated with new procedures
Intraoperative neurologic surveillance may also provide added benefit in
avoiding the exiting nerve roots, especially at L4-L5, where the L3 nerve root
9
can cross the disc space and may be at risk if the approach is in the anterior
one half of the psoas muscle.
2
Peloza validates the use of an electrically elicited
electromyography (EMG) monitoring system for nerve avoidance during a
posterolateral approach to the spine. Electrically elicited EMG monitoring
works by initiating an electrical impulse that causes nearby nerves to depolarize, which generates a muscle contraction in the corresponding myotome(s).
These impulses can be detected using peripheral EMG electrodes. Peloza
uses adhesive EMG surface electrodes applied to the patient’s legs, providing
EMG monitoring of the myotomes associated with the spinal levels of interest: vastus medialis for L2-L4, tibialis anterior for L4-L5; biceps femoris for
L5-S2; medial gastrocnemius for S1-S2. Peloza concludes that such a system
could assist the spine surgeon in safely accessing the intervertebral disc space
for minimally invasive lumbar interbody procedures.
10
CONCLUSION/DISCUSSION
Published literature indicates that the lateral transpsoas retroperitoneal
approach to the spine can be a safe, minimally invasive method for accessing the anterior lumbar spine. Early results show low rates of morbidity, few
serious complications, successful rates of fusion, minimal narcotic requirements, and no difference in surgical outcomes comparing obese to nonobese
cases. Results from studies of XLIF, which is most analogous to the GLIF
technique, state that the surgical corridor for obese patients and nonobese
patients is essentially the same, thus making this technique easier on obese
patients in whom anterior or posterior approaches would be more difficult.
The data suggest that a lateral access to the lumbar spine is most preferable at
the lower levels between L2 and L5; above L2 presents challenges associated
with the ureter and renal artery and below L5 can be more easily approached
with a transperitoneal approach because those levels do not require mobilization of the great vessels, as they are below the vessel bifurcation.
The most significant complications associated with transpsoas
approaches are groin/thigh pain related to disruption of the genitofemoral
nerve or peritoneal perforation while establishing exposure to the spinal column and damaging the exiting nerve roots / lumbar plexus. However, symptoms related to interference of the genitofemoral nerve seem to be transient
and are reported to remit in 6 weeks. As with any surgical procedure there
should be a “bailout” plan; in this case it would be to convert to mini-open
in the case of encountering scar tissue from previous surgeries or any other
unforeseen complication. These less conventional lateral approaches to the
spine have a steep learning curve and hands-on training in a laboratory is
recommended.
456
P A R T V I I I The Future of the Aging Spine
Published articles indicate a left-sided approach to the surgery is preferred because it is easier to dissect the aorta off the spine than to dissect
around the more fragile vena cava. The literature reinforcs the necessity of
correct patient positioning and accurately finding the location of the incision. Diligently positioning the patient and locating the incision will aid in a
smoother operation and procedure, as reorienting surgical portals can often
be cumbersome. The access instrumentation and technique should be constructed to promote ideal “muscle-splitting” techniques compared to musclesacrificing techniques. Handling of the psoas muscle should be done with
care, as the area in and around the muscle indicates the most significant area
for injury. The literature shows that it is optimal for the anterior one third of
the psoas muscle to be carefully dissected and, optionally, aided by the use of
intraoperative neurologic surveillance. Postoperative prescription of bracing
for up to 3 months may also be considered.
References
1. M.P. Steinmetz, D.K. Resnick, Use of a ventral cervical retractor system for minimal access
transforaminal lumbar interbody fusion: technical case report, Operative Neurosurgery 60
(2) (2007) E175–E176.
2. D.L. Bergey, A.T. Villavicencio, T. Goldstein, J.J. Regan, Endoscopic lateral transpsoas
approach to the lumbar spine, Spine 29 (15) (2004) 1681–1688.
3. A. Olinger, U. Hildebrandt, W. Mutschler, M.D. Menger, First clinical experience with an
endoscopic retroperitoneal approach for anterior fusion of lumbar spine fractures from levels
T12 to L5, Surg. Endosc. 13 (1999) 1215–1219.
4. D. Bergey, J. Regan, Lateral endoscopic transpsoas spinal fusion: review of technique and
clinical outcomes in a consecutive series, Spine J. 3 (5) (2003) S166.
5. L. Pimenta, R. Diaz, F. Phillips, F. Bellera, F. Vigna, M. Da Silva, XLIF: 90 degrees, minimally invasive surgical technique to treatment lumbar degenerative scoliosis in adults: clinical
and radiological results in a 15 months follow-up study, Minimally Invasive and Reconstructive Spine Department at Santa Rita Hospital World Spine III Interdisciplinay Congress in
Spine Care Meeting, Rio Janeiro, Brazil, 2005.
6. N.M. Wright, XLIF: the first 10 patients at Washington University, Washington University
School of Medicine. St. Louis, Missouri World Spine III, Rio de Jinero, Brazil, (Sep 2005).
7. M.H. Mayer, A new microsurgical technique for minmally invasive anterior lumbar interbody fusion, Spine 22 (6) (1997) 691–699.
8. P.L. Santaguida, S.M. McGill, The psoas major muscle: a three dimensional geometric study,
J. Biomech 28 (3) (1995) 339–345.
9. T. Moro, S.I. Kikuchi, S.I. Konno, H. Yaginuma, An anatomic study of the lumbar plexus
with respect to retroperitoneal endoscopic surgery, Spine 28 (5) (2003) 423–428.
10. J. Peloza, Validation of neurophysiological monitoring of posterolateral approach to the spine
via discogram procedure, 9th International Meeting on Advanced Spine Techniques, Center
for Spine Care, Dallas, Texas, 2002.
Laser and Ozone Spinal Decompression
James J. Yue and David A. Essig
69
k e y p o i n t s
Define the postulated etiologies of mechanical and radicular pain.
Define the theory of laser disc decompression.
Define the role for laser disc decompression.
Define the postulated mechanism of action of ozone-oxygen chemodiscolysis.
Define the role of ozone-oxygen chemodiscolysis.
INTRODUCTION
Low back pain is one of the most frequent chief complaints in medical
practices. Approximately 80% of the populations of Western countries will
suffer from at least one episode of low back pain in their life. Pain is often
characterized as either radicular or postural. While radicular pain is often
due to an offending disc herniation, the etiology of low back pain is poorly
understood. Proposed pathogenesis includes both mechanical and inflammatory mechanisms, including deformation of the annulus, stimulation
of the nociceptive components of the spinal root, ischemia, venous stasis,
prostaglandins, and cell-mediated immune response.
cedures have been utilized to address the treatment of low back pain. These
include both motion-sparing and fusion interventions. While these procedures have excellent short-term effects, they have been linked to longer-term
complications including recurrent disc herniation, postoperative scarring,
and adjacent motion segment disease. As a result, less invasive procedures
have been developed. Two such procedures, laser spinal decompression and
ozone chemodiscolysis, have shown significant promise.
1
Various surgical pro-
LASER DECOMPRESSION
Peter Choy and David Asher were the first to use laser energy to evaporate
disc material in 1986. Their initial results were poor, but subsequent studies have had good to excellent outcomes in upto 80% of patients.
various different lasers have been described, most of them use approximately 1200 joules per disc in a pulsatile manner. The principles of treatment are based on the hypothesis that the intervertebral disc functions as
a closed hydraulic system. Thus, an increase in water content within the
disc increases the pressure, as a result of the inelastic annulus fibrosus. The
energy from the laser seeks to evaporate intradiscal material to decrease
intradiscal pressure. Furthermore, the energy is hypothesized to denature
and renature proteins, causing irreversible changes to the structure of the
disc and its ability to rehydrate.
As a result of the laser’s energy, the biomechanical properties of the disc
are also affected. Experiments have shown that there is a negative correlation
between laser energy and disc stiffness. While there is a decrease in intradiscal pressure, there is an overall increase in disc circumference and height as a
result of the decrease in disc stiffness. The duration of these biomechanical
changes has also been shown to be a function of laser energy. High-energy
lasers have been shown to maintain disc height reductions on radiographs
and MRI at 12 weeks of follow-up in animal studies.
3
3
2
While
There is significant debate as to the type of laser most suitable for percutaneous laser decompression. Optical analysis of the properties of degenerative discs and the lasers used in clinical practice has revealed that the
wavelength of the Ho:YAG laser provides the highest absorption rate (83%
at 2060 nm wavelength). CO
effective at ablating disc material in vitro.
is chosen, a temperature of 100° C must be attained within the nucleus
pulposus for the treatment to be affective. It is important that this heat be
confined to the nucleus so as to avoid the potential for destruction of the vertebral endplates, possibly causing a sterile spondylodiscitis, which has been
reported in animal models. Clinical complications that have been reported
include discitis, vertebral osteomyelitis, worsening of low back pain, and failure of the percutaneous probe requiring surgical decompression. Overall,
the complication rate is believed to be around 0.5%.
While there is significant debate with regard to the optimum amount
of energy that should be delivered to ablate the nucleus pulposus to reduce
intradiscal pressure, indications for the procedure are confined to contained
disc herniations. Patients with extruded disc herniations, sequestered herniations, narrowed intervertebral spaces, vertebral abnormalities, or those
patients experiencing severe neurologic symptoms should be excluded from
receiving treatment.
als regarding percutaneous laser disc decompression, observational studies
have shown positive evidence in support of the technique. These studies
have shown an average relief of 72% at 1 year, with sample sizes of at least 50
patients. A systematic review of the literature revealed Level II-2 evidence
for short- and long-term relief with a Grade 1C strong recommendation.
4
Although there is a lack of randomized controlled tri-
lasers have also been found to be the most
2
3
No matter what type of laser
2
4
OZONE CHEMODISCOLYSIS
The concept of chemodiscolysis or chemonucleolysis first gained popularity with chymopapain approximately 40 years ago. However, it fell out of
favor as a result of reports of serious neurologic complications in addition to
reports of anaphylaxis. Recently, there has been resurgence in the popularity
of chemodiscolysis using an oxygen-ozone mixture delivered percutaneously.
This technique is based on the principle that pain is generated by mechanical pressure as well as by radicular and periganglionic inflammation.
niated discs are believed to cause pain through an autoimmune reaction as
well as through the triggering of cytokine release. An oxygen-ozone mixture
is believed to achieve its effect not only through chemodiscolysis, but also
through an antiinflammatory effect similar to a corticosteroid. Postulated
mechanisms of action of O
tissue hypoxia, inhibition of proteases, release of immunosuppressive cytokines, and disc dehydration through rupture of water molecules.
The approach is similar to other percutaneous disc techniques. Using
CT guidance, a needle is inserted via a paravertebral approach into the center
of the disc. A mixture of oxygen and ozone gas is injected into the disc and
the foraminal spaces. In a recent study of 2900 patients who were treated
with this method, good results with regard to the VAS score were achieved
in 85% of patients, with no neurologic or infectious complications reported.
Furthermore, a recent randomized study contrasting intraforaminal steroid
and local anesthetic injections with oxygen-ozone injection demonstrated
that oxygen-ozone was more effective at 6 months than steroid injections.
gas include improved oxygenation to reduce
2-O3
1
Her-
6
457
5
458
P A R T V I I I The Future of the Aging Spine
CONCLUSION
Although percutaneous laser disc decompression and oxygen-ozone chemo-
discolysis are exciting new technologies, there is a lack of prospective data
concerning their effectiveness compared to other interventional procedures.
To date, studies have been unable to show an advantage over standard lumbar discectomy.
ing sequestered or extruded disc herniations as well as severe neurologic
dysfunction, the specific patient that could best benefit is not well defined.
It is unclear whether these treatments offer a temporizing solution or a
cure for discogenic and radicular pain. Also, will these minimally invasive
techniques bias future operative procedures? Further studies will hopefully
delineate the appropriate clinical scenarios for these techniques as well as
their long-term results.
7
Furthermore, while specific exclusion criteria exist, includ-
References
1. C. Andreula, M. Muto, M. Leonardi, Interventional spinal procedures, Eur. J. Radiol. 50 (2)
(2004) 112–119.
2. V. Singh, R. Derby, Percutaneous lumbar disc decompression, Pain Physician 9 (2) (2006)
139–146.
3. B. Schenk, P.A. Brouwer, M.A. van Buchem, Experimental basis of percutaneous laser disc
decompression (PLDD): a review of literature, Lasers Med. Sci. 21 (4) (2006) 245–249.
4. V. Singh, et al., Percutaneous lumbar laser disc decompression: a systematic review of current
evidence, Pain Physician 12 (3) (2009) 573–588.
5. M. Muto, et al., Low back pain and sciatica: treatment with intradiscal-intraforaminal O(2)-O
(3) injection: our experience, Radiol. Med. 113 (5) (2008) 695–706.
6. M. Gallucci, et al., Sciatica: treatment with intradiscal and intraforaminal injections of steroid
and oxygen-ozone versus steroid only, Radiology 242 (3) (2007) 907–913.
7. J.N. Gibson, G. Waddell, Surgical interventions for lumbar disc prolapse: updated Cochrane
Review, Spine 32 (16) (2007) 1735–1747.
The Biochemistry of Spinal Implants:
Short- and Long-Term Considerations
Shawn Hermenau, Anne Prewett, and Ravi Ramachandran
70
k e y p o i n t s
Biocompatibility of a material is directly related to the host’s response to
an implanted object. It is multifactorial in nature and may include implant
material, size, shape, location of implant, and duration of implantation.
Release of biomaterial from an implant may have local and systemic effects.
Local or systemic toxicity, mutagenesis, carcinogenesis, and hypersensitivity
reactions are all examples of biologic response to implanted materials.
Major metals used today all have relatively safe biologic profiles and are
relatively inert, but biologic responses have been documented and are well
described.
Polymers and hydrogels are new and emerging fields of implants with
minimal biologic effects.
Bone grafting material can be classified into several groups based on desired
effects. ese materials are typically inert and can be used safely.
HISTORICAL BACKGROUND
In 1892, Sir William Aruthnot Lane began to fix tibia fractures with ordinary steel (Figure 70-1). He was successful in treating a large number of
patients, but noted that the steel plates he used became corroded after time.
Fortunately, and unbeknown to him, the rust that formed acted as a pseudoinsulator (oxide layer), and prevented further degradation and, likely, failure
of the plate. If he had used a dissimilar metal, this layer would not have
formed and a severe electrolyte reaction would have ensued, leading to the
destruction of the metal and inflammation of the tissues. Though metals
had previously been implanted in patients, it was with this advancement
that the use of metal implants for fracture stabilization became a practical
procedure.
the use of biomaterials has enabled us to heal patients as well as to learn
more about material properties and the body’s response to these materials.
This learning process has produced the arsenal of safe materials used today.
The “safety” of a material is in part determined by its biocompatibility.
dependent upon the response of the host tissue to perturbation brought
about by the foreign material. Biocompatibility is very dependent on the
site of implantation, the function and size of the implant, and the duration
of implantation. An unintentional consequence of implanting objects into
a host is the solubility of implanted material and its dissemination into
bodily tissue. This dissemination may be local or throughout the body at
distant sites, with little or no effect or with potentially life-threatening
effects.
and biologic or tissue responses that may occur. The section is broken down
into metals, polymers, hydrogels and biologics.
1
The use of implantable material is not new in orthopedics. For centuries,
Biocompatibility, or the clinical success of a biomaterial, is directly
This chapter will review the major implantable materials in orthopedics
Tissue Response to Biomaterials
The biocompatibility of a material is directly related to the tissue response
generated by the material. These are time-dependent processes and can be
viewed in two different but interconnected ways: first, the bulk properties of
a material, and, second, the physiochemical surface properties of the material, both of which contribute to the initial incorporation and long-term
survival of biologic prostheses (Table 70-1).
The bulk properties of a material can mimic those that they are intended
to replace or augment. Material designs are targeted for the optimization of
function with specific prostheses — wear, strength, and modulus of elasticity. Typically, the bulk materials may have low and unintended systemic distribution in the body over time and may be responsible for potential negative
effects such as hypersensitivity or carcinogenicity.
The surface physiochemical or biochemical properties of a material
directly relate to incorporation of implants and are more crucial to the shortterm success or biocompatibility of a material or implant. The effects of
material surface biochemistry are seen in protein adsorption and mediation
of cell attachment in the implant assimilation.
Tissue response to implanted biomaterials typically follows a predictable
pattern. First, tissue injury and blood-material interaction occurs in the
wound bed. During this phase, a hydration shell is formed around the
implant. This stage is crucial to determining which proteins and molecules
and, hence, cells will adhere to the prosthesis during later stages of incorporation. Hours after implantation, the material becomes covered with proteins from the extracellular matrix, marking the second stage of implantation.
The third stage may occur from minutes to days after implantation and is
marked by the arrival of cells that adhere to the material surface. Cell adherence through integrins is mediated by earlier protein precursors and adsorption. Intercellular protein adsorption occurs, and further cell-mediated
changes are seen on the material surface. Enrichment of surface proteins
(Vroman effect) may mediate cell adherence and subsequent incorporation
of the device into a specific biologic tissue. This final stage may take days
(biodegradable suture), months (bioabsorbable implants), or years (total
disc replacement), depending upon the implanted material and clinical
goals. Adverse responses can occur throughout the assimilation process.
Blood clots, fibrous capsule formation, or foreign body giant cell formation
may result as a consequence of exaggerated or prolonged stimulation of the
immune system.
METALS
Current implantable metal alloys with wide use in orthopedics are 316L
stainless steel, cobalt-chromium alloys, titanium alloys, and tantalum (Table
70-2). In general, metals are used routinely for weight-bearing or load-bear-
ing implants such as plates, nails, stems, and screws. Though biocompatibility is good with metals, there are issues of concern. Corrosion, metallic
toxicity, hypersensitivity, genotoxicity, and carcinogenesis all have been
described in the literature with the use of metallic implants.
459
460
F IG UR E 7 0- 1 Tibial plating with an example of metallic implant.
P A R T V I I I The Future of The Aging Spine
TA BL E 70 1 Commo n Tis sue-Impla nt Interaction s
Implant-Tissue Reaction Consequence
Toxic Tissue necrosis
Biologically inert—smooth surface Implant is encapsulated without
Biologically inert—porous surface Tissue grows into pores and forms
Bioactive Tissue forms interfacial bond with
Dissolution of implant Implant resorption and replace-
bonding
mechanical bonds
implant (bioactive fixation)
ment with soft tissue or bone
Metal Types
Titanium
Although titanium has excellent heat and corrosion resistance capabilities, it
is difficult to form and machine into desired shapes. Also, its extreme chemical reactivity with air, combined with other factors, has caused the cost of
titanium components to be very high. It is used in aerospace applications
where weight and temperature resistance are very important, and in military
applications, where it provides extreme corrosion resistance and durability.
Titanium is also used in biomedical applications such as prosthetics and
implants, due to its biologic inertness.
Pure titanium and titanium alloys are used in the making of orthopedic implants such as total disc replacements, stems, nails, and plates.
There are several titanium alloys that have been developed. The most
commonly used alloy is Ti-6Al-4V. Ti-6Al-4V is composed of titanium, aluminum (6%), and vanadium (4%). These alloys have high corrosion resistance compared to stainless steel and Co-Cr. A passive oxide
coat (TiO
) forms on titanium and its alloys, which protects the metal
2
further from corrosion and enhances the metal’s biocompatibility profile.
These materials are classified as biologically inert biomaterials or bioinert. As such, they remain essentially unchanged when implanted into
patients. The human body is able to recognize these materials as foreign, and
tries to isolate them by encasing them in fibrous tissues. However, they do
not elicit any adverse reactions and are generally well tolerated. Furthermore,
they do not induce allergic reactions such as those observed with stainless
steel and cobalt-chrome implants, which have some nickel in their composition and may elicit a nickel hypersensitivity reaction in surrounding tissues.
Titanium and its alloys possess suitable mechanical properties to be used
in orthopedics, such as strength, bending strength, and fatigue resistance.
Other specific properties that make it a desirable biomaterial are density and
elastic modulus. In terms of density, it has a significantly lower density than
other metallic biomaterials, implying that these implants will be lighter than
similar items fabricated out of stainless steel or cobalt-chrome alloys. Having a lower elastic modulus compared to the other metals is desirable, as the
metal tends to behave more like bone itself, which is desirable from a biomechanical perspective. This implies that the bone hosting the biomaterial is
less likely to atrophy and resorb.
As a clinical benefit, the scatter associated with titanium is far less than
with other metals and makes future imaging studies better. These are not
ferromagnetic metals and are safe to use in MRI magnets.
Cobalt-Chrome
The main components of cobalt-chrome alloys are cobalt, chromium,
molybdenum, and some nickel. Cobalt alloys are combined with chromium
and molybdenum to increase the metal’s corrosion resistance. Cobaltchrome alloy was the first alloy that was introduced in dentistry in the
1930s, and since then, has proved its clinical effectiveness as a biomaterial.
The components of Co-Cr are elemental, as noted above, and therefore
must be classified as more biologically favorable in principle, as elements
that have no function in the human body. For essential elements, the body
has diverse ways of decomposition and utilization. There appears to be certain threshold values, below which no interaction takes place.
Ores of cobalt are accompanied by nickel. Complete separation of the
elements is never possible. The relevant standards stipulate a maximum
nickel content of 0.1%. Concentrations of greater than 0.1% have to be
declared. Alloys with less than 0.1% of nickel can be designated as nickelfree. In standard cobalt-chrome implants, the release of nickel will amount
to approximately 0.00003 mg/cm
2
(0.03 μg/cm2) in the first week and constantly decline thereafter. If one compares this to the daily uptake in food,
i.e., approximately 0.19 to 0.90 mg, (190 to 900 μg), toxicological or allergic
stress appears very improbable.
Stainless Steel (316L)
The composition of stainless steel has varying percentages of iron, chro-
mium, nickel, molybdenum, and carbon. The most common stainless steel
alloy in orthopedic implants is SS 316L. The designation 316L by the
ATSG is broken down as follows: the 300 series represents the austenitic
family (crystalline structure) of steel, and the L means that the carbon content of the stainless steel is below 0.03%; this will reduce the sensitization
effect, precipitation of chromium carbides at grain boundaries, due to the
high temperatures produced by welding. The effect of these precipitates can
weaken the material by increased corrosion at the grain boundaries. Stainless steel is chemically treated with nitric oxide to form a passive oxide layer,
to further increase its corrosion resistance.
Stainless steel is a strong material, with better ductile properties than all
the other implantable materials. It has fallen out of favor in the United
States, but worldwide is still the most commonly used metal implant. The
decrease in U.S. utilization is due to the superior strength, corrosion resistance, and mechanical properties of titanium and cobalt-chrome. Additionally, the biocompatibility profile of stainless steel is less favorable than the
other metals, with more reports of hypersensitivity reactions because of its
higher nickel content.
Tantalum
Tantalum is a gray, heavy, and very hard metal. When pure, it is ductile and
can be drawn into fine wire, which is used as a filament for evaporating metals
such as aluminum. Tantalum is almost completely immune to chemical
attack at temperatures below 150
°
C, and is attacked only by hydrofluoric
TA BL E 70 2 Relat ive Meta llic Proper ties
The hydroxide
C H A P T E R 7 0 e Biochemistry of Spinal Implants
461
Characteristic
Stiffness Medium High Low Low
Strength Medium Medium High High
Corrosion resistance Low Medium High High
Biocompatibility Low Medium High High
316L Stainless Steel Cobalt Chrome Titanium Tantalum
Electrochemical cell action
driven by the energy of
oxidation continues the
corrosion process
Fe
2+
e
-
OH
O
2
-
e
-
Iron hydroxide forms
and precipitates
Cathode action
reduces oxygen
from air, forming
hydroxide ions
Electron flow
Iron
quickly iodizes
to form rust
Water droplet
2+
Fe
-
OH
O
2
-
e
-
e
Anode action causes
pitting of the iron
F IG UR E 7 0- 2 Example of pitting corrosion.
acid, acidic solutions containing the fluoride ion, and free sulfur trioxide. At
high temperatures, tantalum becomes much more reactive. Tantalum is used
to make a variety of alloys with desirable properties such as high melting
point, high strength, and good ductility. Tantalum readily forms oxides and is
most stable as +5 tantalum pentoxide. Elemental tantalum unites strength
and corrosion resistance with excellent biocompatibility. Tantalum is the
metal used in the construction of Trabecular Metal (Zimmer). The cellular
structure of Trabecular Metal resembles bone and approximates its physical
and mechanical properties more closely than any other prosthetic material. Its
unique, highly porous, trabecular configuration is conducive to bone formation, enabling rapid and extensive tissue infiltration and strong attachment.
Corrosion
Most fluids in the human body are of similar chloride content and pH to sea
water (20 g/L and 7.4); therefore many metals used in orthopedic implants
have been those most resistant to corrosion in sea water. Corrosion is, simply, the dissolution of metallic ions in aqueous solution. Electrochemical
cells are produced in the body when these metallic implants are used and
equilibria of metallic ions in solution are achieved within body fluids over
time (Figure 70-2).
Generally three types of corrosion exist with the use of metallic implants
and include (1) galvanic, (2) crevice or pitting, and (3) fretting corrosion.
Galvanic corrosion is corrosion due to the use of dissimilar metals in contact with one another or electrochemical dissolution. Pitting corrosion is a
form of localized corrosion that leads to the creation of small holes or
defects in the metal (Figure 70-3). The driving power for pitting corrosion
is the lack of oxygen around a small area. This area becomes anodic while
the area with excess of oxygen becomes cathodic, leading to very localized
galvanic corrosion. The corrosion penetrates the mass of the metal, with
limited diffusion of ions, further increasing the localized lack of oxygen.
The mechanism of pitting corrosion is probably the same as crevice corrosion. Finally, fretting corrosion, as defined by the ASM Handbook on
Fatigue and Fracture, is: “A special wear process that occurs at the contact
area between two materials under load and subject to minute relative
motion by vibration or some other force.” The relative small motion causes
mechanical wear and material transfer at the surface of the metals, followed
by oxidation of that debris and the freshly exposed surface. This debris then
acts as an additional abrasive product that is often harder than the original
metal and perpetuates the process.
Distribution of Metal in Body Fluids
A prosthetic device constitutes a pool of trace elements or alloy in the
body, which, when mobilized by corrosion, dissolution, and wear, are distributed in local tissue or potentially at sites distant to the original site of
implantation. Metallic particles can be found in local tissues such as articular joint capsules, muscle, and regional lymph nodes, or at distant tissues
such as abdominal paraaortic lymph nodes, liver, spleen, and pancreas.
Studies have looked at distribution of metal ions in body fluids following
prosthetic joint implantation. Only slight increased levels of Co and Cr in
serum and urine have been noted in patients 2.5 years after implantation
of the prosthesis.
1
Another paper reported increased deposition of metallic particles in the liver, spleen, and abdominal paraaortic lymph nodes, in
a postmortem study. Larger metallic burdens were seen in patients with
failed total joint replacements. In most of the patients evaluated in the postmortem study, the concentration of metallic particles in the liver and spleen
was low, and no toxic effects were apparent on histological exam of the surrounding tissue.
2
Animal studies have shown that nickel, cobalt, or molybdenum introduced into tissue is quickly transported and eliminated in the urine within a
relatively short time. Chromium is not eliminated as quickly and can accumulate in tissues and red blood cells. The hexavalent Cr often will be reduced to
trivalent Cr and become cell-associated, therefore accumulating in the body.
Mutagenesis
Metallic particles disseminated throughout the body are feared to have
potentially deleterious effects. Some early studies were published raising
the question of biomaterials being responsible for mutagenesis at distant
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P A R T V I I I The Future of The Aging Spine
Liquid film on
surface receives
ions from metal
2+
Fe
Oxide layer
protects surface
but allows
electrons to
flow to oxygen
Oxygen-deficient
anodic region
near bottom of
corrosion pit
Metal
F IG UR E 7 0- 3 Example of electrochemical cell around metallic implant.
-
e
O
2
sites in the body. Mutagenesis, or genotoxicity, is the disruption of DNA
resulting in the production of aberrant proteins that lead to cellular or
tissue dysfunction. Genotoxicity or mutagenesis can serve as an indicator
for the potential carcinogenicity of a material. A paper in 2003 looked at
potential mutagenesis of cobalt-chrome and titanium implants. The conclusion was that neither material showed evidence of mutagenesis in bacterial assay and mammalian cell assays. Although, by itself, this is not
enough evidence to be able to state that these implants are not mutagenic
or genotoxic, in combination with the reports of cancer (next section) in
patients with Co-Cr or titanium implants, it appears to be supported
clinically.
Carcinogenicity
Although local and systemic deposits of metallic ions have been demon-
strated in patients with implanted metal, the associative relationship of
these toxic effects has yet to be established. Concentration-related connections between orthopedic implants and malignant degeneration have been
questioned and potential case reports have been published. The International Agency for Research on Cancer concluded that implanted foreign
bodies of metallic titanium, cobalt-chromium, and stainless steel appear not
to be directly carcinogenic in humans.
There is sufficient evidence in experimental animals for the carcinogenicity of implants of cobalt, nickel, and nickel alloy powder containing
approximately 66% to 67% nickel, 13% to 16% chromium and 7% iron. This
noted, there is inadequate evidence in experimental animals to establish the
carcinogenicity of orthopedic-type implant materials of chromium metal,
stainless steel, titanium metal, or titanium-based alloys.
There is inadequate evidence in humans for the carcinogenicity of
metallic implants and metallic foreign bodies, although numerous case
reports and small power studies have been published and controversy has
been generated. Out of the large number of patients with orthopedic
implants, a total of 35 cases have been reported of malignant neoplasms
arising from the bone or the soft tissue in the region of an implant. Fourteen cohort studies of patients following total knee or total hip replacement from six countries were performed to investigate cancer incidence in
these populations. One study showed a small increase in overall cancer
incidence, while the remaining studies showed overall decreases. Four of
these studies suggested a possible increased risk for specific cancers,
including Hodgkin disease, non-Hodgkin lymphoma, leukemia, and kidney cancer. However, results of several other studies were not consistent
with this observation. Additionally, two case-control studies, one including cases with soft-tissue sarcoma and the other including lymphoma and
leukemia, were carried out in the United States. These studies failed to
establish a causal effect. Most of the studies did not have information on
possible confounding variables such as immunosuppressive therapy or
rheumatoid arthritis for the lymphomas and analgesic drugs for kidney
cancer. The follow-up in most of the studies may have been too short to
evaluate cancer occurring many years after exposure. A total of 23 cases of
sarcomas, 23 cases of carcinomas, and 7 cases of brain tumors have been
reported at the site of metallic foreign bodies, mainly bullets and shrapnel
fragments.
Hypersensitivity
The first report of hypersensitivity with a metallic orthopedic implant was in
1966 by Foussereau and Laugier. They reported on a patient with an eczematous dermatitis and associated this hypersensitivity reaction with nickel. Since
then, a growing body of literature has described metal hypersensitivity reactions to stainless steel, cobalt-chromium, and, to a lesser degree, titanium
implants. Though well documented, these metal hypersensitivity reactions
remain unpredictable and poorly understood events relative to orthopedic
implants.
The prevalence of metal hypersensitivity in the general population is
approximately 10% to 15%. Metals known to cause reactions are nickel,
beryllium, cobalt, chromium, and to a far lesser extent, titanium and tantalum. Nickel is the most common sensitizer, with a prevalence of
approximately 14%. Cross-reactivity between nickel and cobalt exists. In
patients with metal prostheses, the prevalence of metal hypersensitivity is
approximately 25%, and in patients with failed prostheses, the prevalence
reaches 60%. It is unclear whether the failure is a result of hypersensitivity
or whether increased degradation products in the body due to the implant
failure result in increased hypersensitivity reactions.
Dermal contact and ingestion of metals is known to result in an immune
response causing hives, eczema, redness, and itching. Resultant metallic
degradation products may sensitize the body and generate similar effects.
A temporal association between implantation and clinical manifestations
of these symptoms has been shown. Implant-related hypersensitivity reactions are typically cell-mediated reactions (type IV delayed-type hypersensitivity).
Implant degradative products from corrosion or mechanical wear will
react and bind to proteins in tissue and form organometallic complexes. It is
these complexes that become antigens, sensitize T cells, and eventually
result in a T-cell–mediated immune response. T cell release of cytokines,
including IL-3, granulocyte-macrophage colony stimulating factor, INF-α,
and TNF-β, then leads to the activation and infiltration of macrophages
responsible for the immune response seen in these delayed-type hypersensitivity reactions.
Clinically, the immunologic response within the periprosthetic area may
include vasculitis, fibrosis, muscle necrosis, osteolysis, and metallosis. This
cascade of events may result in mechanical failure of the device or inability
of the implant to be integrated into the biologic system, and may mandate
removal of the biomaterial. Removal of a device that has served its function
and can be safely removed should be considered, as this may alleviate some
of the symptoms for the patients.
Though hypersensitivity reactions to orthopedic implants are not common, more frustrating is the lack of predictability for avoiding this complication. No evidence exists to support the use of routine allergy testing or skin
testing of metals in patients undergoing implantation of a metallic device. In
the event of temporally related skin symptoms and metallic implantation,
skin sensitivity testing should be considered. Until more studies are conducted to better define the role of delayed-type and humoral immune
hypersensitivity reactions in patients with metallic implants, the risk to
patients should be considered minimal.
POLYMERS
Introduction
Synthetic polymers are occupying a growing role in implant construction.
They accord numerous advantages including radiolucency and elasticity.
While the vast majority are biologically inert, their wear and degradation
processes and properties reflect on their suitability as implants. The following sections will outline the principal polymers used in disc arthroplasty, in
fusion, and as bioabsorbable interbody spacers.
C H A P T E R 7 0 e Biochemistry of Spinal Implants
463
UHMWPE
The polymer with which spine surgeons have had the most experience and
the longest-reaching data is ultra high molecular weight polyethylene
(UHMWPE). The material is composed of extremely long chains, with
molecular weights numbering in the millions. The longer chain serves to
transfer load effectively to the polymer backbone, resulting in a very tough
material. UHMWPE is highly resistant to corrosive chemicals, exhibits low
moisture absorption, has a low coefficient of friction, and is both selflubricating and resistant to abrasion.
Total disc arthroplasty’s CHARITÉ Artificial Disc has been implanted
since the 1980s. The implant consists of two metallic endplates that articulate with a central UHMWPE disc. Modern manifestations of this design,
still incorporating the UHMWPE articulation, include the Synthes ProDisc-C, the Cervitech PCM, the LDR Spine Mobi-C, the Aesculap AG
Activ-C, and the DePuy Spine Discover.
PEEK
Polyether ether ketone (PEEK) is an organic polymer thermoplastic.
Molecularly, it consists of phenylene rings that are linked via oxygen bridges.
There have been no reports of biologic adverse response/reaction with this
material. The plastic is used either alone or with a carbon fiber reinforcement. It is a member of the polyaryletherketone family, which includes several other polymers with applications in spine surgery.
The first utilization of PEEK was in spinal cages in the 1990s, by
AcroMed. An advantage of the polymer in this function is its radiolucency,
which facilitates radiographic assessment of fusion in vivo. The majority of
current implants employing PEEK are cervical and lumbar spinal cages.
Examples include Zimmer Spine’s BAK Vista radiolucent Interbody Fusion
System, Surgicraft’s STALIF anterior lumbar fusion cage, Scient’x’s CC
interbody fusion cage, Depuy Spine’s OCELOT Stackable Cage System,
and the Nubac Disc Arthroplasty Device.
Recent studies with PEEK cages attempt to accelerate fusion by incorporating hydroxyapatite, 40% β-tricalcium phosphate/60% hydroxyapatite,
or rhBMP-2. Possible applications for the material in posterior dynamic stabilization, interspinous process decompression systems, posterior rods, and
total disc replacement are also being explored.
PLA and PGA
Bioabsorbable devices in the spine are composed of polymers known as
alpha-polyesters or poly-(alpha-hydroxy) acids. These include polylactic acid
(PLA) and polyglycolic acid (PGA). PLA is based on a lactic acid monomer,
while PGA is based on a glycolic acid monomer. Both substances safely
degrade completely in vivo and are used in bioabsorbable interbody spacers.
Advantages of bioabsorbable polymers over metals in spine surgery include
the avoidance of imaging artifact on postoperative radiographs and a modulus
of elasticity closer to that of native bone, lessening stress-shielding. As the
implant is ultimately resorbed, complications such as implant erosion and
migration may be avoided. Disadvantages of the use of bioabsorbable polymers are an initial strength that cannot match their metal counterparts and the
possibility of generating an inflammatory response with breakdown products.
Implant Performance and Failure
The reaction of the implants and the materials that compose them with the
body is largely a result of the processes of wear, degradation, and oxidation.
In the following section, we will discuss the theoretical and observed complications specific to polymer implants.
UHMWPE
Given our relatively short experience with the application of UHMWPE in
spine surgery, most relevant in vivo clinical data come from retrieval studies
of the CHARITÉ Artificial Disc. Some data have also been published on
the ProDisc-L. Both implants are constructed with a UHMWPE insert.
Relevant wear data concerning other materials are largely a product of lab
studies. Decades of experience from total hip and knee arthroplasty have
demonstrated that UHMWPE wear particles have the ability to cause
implant failure through macrophage-mediated aseptic osteolysis. Similar
complications have been observed in spine surgery. Osteolysis has been
observed around certain total disc replacement designs, including the
CHARITÉ implant. The particle load and resulting inflammatory response
in the periprosthetic area is reported to be proportional to that observed in
total hip arthroplasty.
CHARITÉ components, retrieved for intractable pain and/or facet
degeneration, frequently displayed one-sided wear patterns. The dome of
the components typically exhibited burnishing, and the rim showed evidence of plastic deformation, burnishing, and fracture, thought to be produced by impingement. Similar patterns are described in the ProDisc-L and
Prodisc-C.
In addition to impingement, rim damage observed in polyethylene total
disc replacement retrievals has also been associated with postirradiation oxidation. Analysis of explanted CHARITÉ cores has shown that the exposed
rim experiences severe oxidation after 10 or more years. The central dome is
protected from in vivo oxidation due to contact with the metallic endplates.
The end product of UHMWPE wear is the creation of wear debris and
ensuing aseptic loosening. The biology of aseptic loosening has been extensively studied and described in the hip and knee total joint arthroplasty literature. The cellular response to UHMWPE consists primarily of giant
cells and macrophages. The magnitude of the response is directly related to
the volume of debris. The role of these cells is to detect, phagocytose, and
degrade any foreign material. In the process, these cells release chemical
messengers, including cytokines and other mediators of the inflammatory
process. As a result, a foreign-body granulomatous response is initiated.
Macrophages fuse, forming giant cells to wall off the foreign material.
Osteoclasts are activated by the cytokines IL-1b, IL-6, IL-8, PGE2, and
TNF-α. Osteolysis is postulated to be the product of both osteoclastic and
macrophage– and giant-cell–mediated bone resorption.
PEEK
Theoretical implant complications of PEEK cages are the same as those
observed with metal devices. These include subsidence, wear, debris production, and fracture. Wear debris has been identified in periprosthetic biopsies,
but no evidence of an inflammatory reaction to the particles has been
described. In vitro and in vivo studies suggest that PEEK particles appear to
be harmless to the spinal cord.
Examination of systemic, intramuscular, and intracutaneous toxicity of
PEEK has revealed no adverse side effects. There do not appear to be issues
with sensitization or gene toxicity. Extensive in vitro testing with fibroblast,
macrophage, and osteoblast cell lines shows no cytotoxicity, immunogenesis,
or genotoxicity of PEEK.
PLA and PGA
Alpha polyesters are degraded by hydrolysis. The process releases their
respective monomers, which are then incorporated into normal cellular
physiologic processes. Lactic acid is produced from PLA, and glycolic acid
from PGA. Lactic acid eventually ends up in the citric acid cycle, while glycolic acid can be excreted in urine. The rate of degradation is based both on
factors inherent to the implant and polymer, including molecular weight,
crystallinity, and porosity, and other local factors including vascularity and
loading conditions.
As the implant begins to degrade and fragment, particle removal via a
foreign-body reaction begins. The rate of degradation is associated with the
degree of inflammatory response, synovitis, and even activation of the complement cascade. PLA has the slower rate of degradation, while PGA has
the faster. As a result, PLA degradation has been associated with a foreignbody reaction as late as 143 weeks after implantation, while a foreign-body
response reaction to PGA has been seen as early as 3 to 6 weeks after
implantation. As a result of this inflammatory response, complications such
as sterile sinus tract formation, synovitis, hypertrophic fibrous encapsulation, and osteolysis have been described.
HYDROGELS
Synthetic Hydrogels
Synthetic polymers exhibit low toxicity, and have been used in medical
applications for a period of more than 60 years. Many polymer systems have
been employed, including polyacrylonitrile, polyamides, polyethylene,
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P A R T V I I I The Future of The Aging Spine
polymethylmethacrylate, polytetrafluoroethylene, polyurethanes, and silicones. Products made from these polymers have been used as bone and tissue replacements, as drug delivery devices, and have been variously employed
in nearly all medical disciplines, including heart surgery, orthopedic surgery,
ophthalmology, gynecology, and plastic surgery with remarkable success.
Synthetic polymers are also utilized in topical applications and as coatings
for stents and other implants.
As a result of their synthesis, hydrophobic polymers generally contain
minute amounts of residual impurities, such as monomers, degradation products, stabilizers, catalysts, and solvents. Regardless of the purity of the polymer,
there is the potential for small quantities of these impurities to migrate into
the recipient of products devised from these polymer systems. These impurities are very difficult to remove completely from the polymers, and they can
migrate over long time periods from the polymers into the surrounding tissue.
Water-insoluble, hydrophilic polymers that absorb large quantities of
water relative to their initial weight are called hydrogels. Through the
absorption of water or surrounding media, hydrogels expand in both weight
and volume and can be viewed to some extent as “solidified water.” Synthetic
hydrogels may have several important advantages as biomaterials, when
compared with the classic hydrophobic polymers.
Hydrogels are typically permeable to aqueous solutes, thereby permitting the removal of water-soluble impurities by simple aqueous extraction.
As hydrogels are principally composed of water, they are highly biocompatible and exhibit reduced potential to invoke an inflammatory process. The
potential for fibrosis and encapsulation is diminished with hydrogel
implants relative to traditional hydrophobic synthetic polymer implants. As
a result, hydrogels show low adherence to tissues, making them excellent
candidate materials for adhesion barriers. Hydrogels also exhibit low friction relative to surrounding tissue. The higher the water content of the
hydrogel, the lower the friction.
Protein and lipids can be deposited on the surface of hydrophobic polymers due to denaturation. Cell adhesion proteins are frequently denatured
in this fashion and can lead to cellular attachment and fibrosis. Hydrogels,
due to their high water content, are resistant to lipid and cell attachment and
spreading. As a result, hydrogels often exhibit low adhesion of platelets and
other thrombotic cellular elements.
Hydrogels are permeable to water and to small molecular weight watersoluble substances. Part of the water in the swollen hydrogel is available as
free water, which provides a diffusion path through the polymer’s structure
for molecules up to a certain size. At the same time, the polymer network
acts as a barrier for larger molecules and for cells, bacteria, and viruses.
One of the first hydrogels used widely was Ivalon. Composed of polyvinylalcohol cross-linked with formaldehyde or glutaraldehyde, the material
was hard in the dry state, and soft and pliable in the swollen state. The material found uses as an implantable device and was used variously as a repair
material for anorectal reconstruction, breast augmentation, middle ear tympanoplasty, and orthopedic surgery. Complications related to loss of tensile
strength and a tendency to become brittle over long-term surgical implantation limited its use.
The potential of synthetic hydrogels as biomaterials was first recognized
by Wichterle and Lim in the 1950s. Hydrogels based on hydroxyethylmethacrylate (HEMA), sparingly cross-linked by diesters of diglycols
(mono-, di-, tri-, and tetra-) and methacrylic acid were tested on animals
mid-century and later developed for soft contact lenses. They later were
tested for use as an implant material for reconstructive, plastic, ophthalmic,
thoracic, orthopedic, and general surgery; and for drug delivery. Covalently
cross-linked polyHEMA is very stable chemically and thermally, and is
resistant to enzymatic degradation. This polymer is extensively used in the
soft contact lens industry, either as pure poly(HEMA) or in various copolymers, such as PolyHEMA and Polyvinylalcohol. In addition, both polyHEMA and PVA are resistant to degradation due to the carbon-carbon
backbone, which is chemically very stable. Polymers such as polyamides,
polyesters, and polyurethanes lack the C-C backbone yet have found a wide
application as medical hydrogels. Although their in vivo stability cannot
match the stability of polymers with the C-C backbone, they have been
found to be stable in tissue for over 1 year, with no loss of mass or mechanical properties. Their widest application in medicine today is found in hydrophilic coatings on stents and catheters, wound and burn dressings, and
controlled drug release formulations.
Hydrolyzed Pan Hydrogels – Development and History
The last group of synthetic hydrogels is the HPANs. It is a family of ther-
moplastic hydrogels, based on acrylic multiblock copolymers. HPAN copolymers form hydrogels using phase separation and formation of crystalline
clusters, which cause physical cross-linking. HPAN copolymers are formed
by a partial controlled hydrolysis of polyacrylonitrile (PAN). Their formation requires just a simple chemical reaction (hydrolysis) and they contain
no monomers, cross-linkers, catalysts, or other toxic residuals.
HPAN hydrogels belong to a family of hydrogels based on partial
hydrolysis of PAN, generally called HPANs (hydrolysed PANs). First generation of HPANs was developed in the Institute of Macromolecular
Chemistry of the Czechoslovak Academy of Sciences in the Czech Republic, and its synthesis, composition, and properties were described in a number of papers. These materials were found to be highly biocompatible and
were used in contact lenses and orthopedic implants.
Additional HPAN advantages as compared to other hydrogels are:
(1) Mechanical strength even at high water content: HPAN hydrogels are
comparable in elasticity and tensile strength to tissues with a similar
water content, such as cornea, vitreous body, cartilage, and nucleus
pulposus of the intervertebral disc. ese materials are particularly
resistant to tear propagation. HPAN is probably stronger and more
resistant to mechanical damage than any other current synthetic hydrogel of the same liquid content.
(2) As for other hydrogels, HPAN hydrogels are permeable to water-
soluble compounds. e maximum size of these molecules (permeation
limit) can be controlled by the water content of the hydrogel. Molecules
smaller than the permeation limit (drugs, nutrients, metabolites, salts,
gases) can be transported through the hydrogel using either diffusion or
hydraulic flow mechanisms. At its maximum water content, HPAN
can pass solutes of molecular weight up to 100,000 Daltons. Highlyhydrated HPAN has also very high hydraulic permeability similar to
cartilages and comparable tissues.
BIOLOGICS
Bone Graft
Bone graft is required to fill voids to achieve fusion of motion segments and
to unite fractured bones. The ideal bone graft, considered the gold standard
to which all others are compared, is autogenous bone graft or autograft. The
ideal bone graft substitute should be osteogenic, biocompatible, bioabsorbable, able to provide structural support, easy to use clinically, and costeffective.
The normal host response to autograft is divided into several phases. As
mentioned earlier, with any “foreign” object implanted into the body there is
hemorrhage and inflammation, and next, invasion by vascular elements
from the periphery that bring in precursor cells to osteoblasts and osteoblasts themselves. The rim of osteoblasts deposits new bone on the outer
edges of the graft and remodeling begins. This stage may take weeks to
months, depending on the type of graft, and is completed after the graft is
fully incorporated into the host tissue in a seamless fashion.
The bone grafts and their substitutes are divided according to their
properties of osteoconduction, osteoinduction, osteogenesis, or a combination of these. Osteogenic refers to a material that produces bone-forming
cells that directly lay down new bone in an area. Osteoinductive refers to a
material that can stimulate the differentiation of stem cells into osteogenic
cells. Osteoconductive materials are those that provide a porous scaffold to
support the formation of new bone. In addition, there are materials that
provide more than one of the above characteristics and are considered combination materials (Table 70-3).
Osteoconduction refers to the process in which the three-dimensional
structure of a substance is conducive to the ongrowth and ingrowth of new
bone. Osteoconductive bone graft substitutes are commercially available and
vary in chemical composition, structure, and resorption rates. Understanding the basics of each type and the reason to use a specific one of them will
assist with surgical success. The Table 70-3 groups these materials into
classes and describes some of their basic properties.