Resorbable Technology for Spinal Stabilization

Daniel K. Resnict, Joseph T. Alexander, William C. Welchc a Department of Neurological Surgery, University of Wisconsin, K4/834 Clinical

Science Center, Madison, Wisc., b Department of Neurological Surgery, Wake Forest University, Winston-Salem, N.C., and c Department of Neurological Surgery, University of Pittsburgh, Pittsburgh, Pa., USA

Incredible advances in the surgeon's ability to restore stability to the unstable spine have been made in the last decade. Using metal implants, surgeons can now effectively stabilize any motion segment of the spine. With the concomitant use of autograft or allograft bone, high fusion rates can be achieved in the vast majority of cases. There are, however, drawbacks to the use of such systems.

Autograft bone is not always available for a given application and autograft harvest is associated with nontrivial morbidity in a substantial fraction of patients [1]. The use of allograft bone is associated with lower fusion rates than autograft bone in most cases. The use of allograft is often limited by both cost and availability. Machining of allograft bone to predetermined shapes is a technology- and labor-intensive process, and processing of the bone for infection control and mechanical consistency is expensive. Finally, the use of metallic implants is associated with late complications related to implant failure, erosion into soft tissue structures, and degradation of the bone/implant interface.

In response to these drawbacks, several companies have developed resorbable biomaterials that can serve as temporary fixation devices, structural supports, and osteoconductive or even osteoinductive conduits for new bone growth. These products offer the potential advantages of unlimited supply, significant cost savings, and a reduction in patient morbidity. This chapter is intended to serve as an overview of some of these new biomaterials and their potential applications. This overview will be incomplete, as the pace of new development in this field is staggering. Many of the substances and devices presented are not FDA approved or are FDA approved for uses outside of the spine. The reader is cautioned to review the FDA-approved indications for the use of any of the described products prior to the experimental or therapeutic use of any of the products described.

Resorbable Synthetic Bone Substitutes

Autograft bone remains the gold standard for spinal fusion procedures, especially lumbar posterolateral fusion. Autograft bone is osteogenic, in that it contains the cellular components responsible for new bone formation. Autograft is osteoinductive, in that it contains diffusible proteins which induce neighboring cells to produce bone. Autograft is also osteoconductive, in that it has a structure which allows for the migration of osteogenic cells which can then form new bone [2]. Depending upon the size, shape, and source of the autograft, mechanical strength may be negligible (cancellous) or substantial (tricortical iliac crest, fibula, rib). Drawbacks to the use of autograft include harvest site morbidity as well as limited availability. The use of allograft bone avoids the problem of graft site morbidity, and allograft humeral, femoral, and patellar grafts are available for use as large load-bearing struts. Allograft bone is osteoconductive, and in some preparations may be somewhat osteoinductive (such as in demineralized bone matrix). Unfortunately, allograft bone is not always readily available. Allograft is expensive to preshape into readily used sizes. Finally, some patients may refuse the use of allograft for religious reasons or because of fear of infection.

Although there are many alternatively prepared allograft products (Grafton™, Dynagraft™, Osteofil™) and xenograft products (Collagraft™, True Bone Ceramics™) available for use which may extend the utility of allograft use, the manufacture of these products still requires the use of processed donor (human or animal) bone. In some cases, the extensive preparation of these products has resulted in unexpected toxicity. For example, Botsman et al. [3] and Wang et al. [4] recently presented evidence to suggest that the glycerol carrier in Grafton (Osteotech, Eatontown, N.J., USA) may be toxic in large doses. This discussion will center on synthetic resorbable bone substitutes.

The use of synthetic bone graft material would negate the risks of harvest site morbidity, disease transmission, and limited donor availability. The ability to machine the material would allow for the creation of multiple shapes and sizes that could be used as 'off-the-shelf' bone substitutes. Substantial cost savings may be realized with reasonably priced implants if the costs of autograft harvest (estimated to be USD 1,500-5,000 for iliac crest [5, 6]) and allograft harvest and preparation (machined allograft bone wedge approximately USD 3,000 per level, allograft femoral ring USD 1,200-2,400 per level;

source: University of Wisconsin operating room) are avoided. Ideally, these synthetic bone grafts would be totally absorbed by the body and replaced by native bone.

Calcium sulfate (plaster of Paris) has been used for bone defect repair since 1892 [7, 8]. Peltier and Jones [9] used calcium sulfate pellets to fill unicameral bone cysts in 26 patients with excellent results, and Coetzee [10] used similar pellets to treat osseous defects in the skull and facial bones in 110 patients. Recently, Kelly et al. [7] reported the use of Osteoset™ (Wright Medical Technology, Tenn., USA) a highly processed form of calcium sulfate for repair of bone defects in a series of 109 patients with various bone lesions. Osteoset was used as a bone graft substitute if the surgeon recommended the use of morsellized graft, the bone void was not intrinsic to the stability of the structure, and the void shape could accommodate the calcium sulfate pellets. Overall results were excellent, with 100% pellet resorption and 94% new bone growth (radiographic criteria) noted at 1 year [7]. No spinal fusions were performed during the study, and other materials were used in addition to the pellets in 65% of cases. Hadjipavlou et al. [11] recently reported excellent results with calcium sulfate-filled titanium cages in an interbody fusion model in adult sheep. These authors obtained roughly equivalent results with cages filled with either autograft iliac crest or with calcium sulfate. There are at present, however, no published reports of the use of calcium sulfate for human spinal fusion.

Processed coralline grafts, including hydroxyapatite, have been used extensively for the repair of bone defects. Additionally, these products have been used for spine surgery with a number of different applications. Hydroxyapatite is a poorly crystalline calcium phosphate compound which is similar in mineral composition to bone [6]. Calcium phosphate and hydroxyapatite are osteocon-ductive, if the porosity, pore size, and degree of pore interconnectivity are optimized [6]. Substantial research has established that the ideal pore size is between 100 and 500 |xm [12]. Calcium hydroxyapatite is most commonly formed through processing of coral, which contains calcium phosphate and calcium carbonate in the form of aragonite. Several natural corals, particularly Porites and Goniopora, have pore sizes and pore interconnectivity which allow for osteoconduction. These corals have been harvested, cleansed, and then used as graft material (Biocoral™ Inoteb, France), particularly in dental surgery [13]. These corals may also be chemically treated via the 'replamineform' process. This process results in the conversion of aragonite to hydroxyapatite without changing the three-dimensional structure of the coral [6]. ProOsteon™ 200R, 500R, and Interpore porous hydroxyapatite™ (Interpore Cross International, Calif., USA) are produced in this fashion. The numbers 200R and 500R refer to the nominal pore diameter of the crystalline structure, either 200 or 500 |xm. The mechanical properties of coralline hydroxyapatite grafts are more similar to cancellous rather than cortical bone. In general, increasing porosity increases the osteoconductivity of the graft and lowers the compressive strength.

Coralline hydroxyapatite has been used as a bone substitute for both cervical interbody as well as lumbar posterolateral fusion in animal, and in some cases human trials. In 1996, Zdeblick et al. [14] studied the use of ProOsteon 500 R in a multilevel anterior cervical fusion in a goat model with or without plate fixation. They found that although only 48% of grafts incorporated with the hydroxyapatite alone, 71% incorporated if anterior cervical plate fixation was added. However, none of the implants were completely replaced by host bone at 12 weeks. Biomechanical testing revealed that the hydroxyapatite-fused segments behaved poorly without plate fixation. However the addition of an anterior cervical plate improved performance such that segments fused with hydroxyapatite and a cervical plate behaved similar to segments fused with allograft (but not autograft) bone and cervical plate in torsion. Hydroxyapatite-fused segments were inferior to both autograft and allograft in flexion/extension testing [14]. Boden et al. [15] studied the use of hydroxyapatite in a rabbit model of lumbar posterolateral fusion. They found that the use of coralline hydroxyapatite alone resulted in 0/14 solid fusions. Combination of hydroxy-apatite and autograft bone in a 1:1 ratio resulted in a fusion rate of 50% (7/14). Bozic et al. [16] found that the combination of hydroxyapatite and bone marrow aspirate resulted in 25% fusion rate in a rabbit model. Addition of electrical stimulation (100 |xA implantable direct current stimulator) significantly improved fusion rates. Similarly, Baramki et al. [17] noted a significantly lower fusion rate for hydroxyapatite-grafted animals (50%) than autograft-grafted animals (100%) in a sheep model of posterolateral fusion.

Despite these rather lukewarm animal results, Thalgott et al. [18] used hydroxyapatite (ProOsteon 200R) in combination with an anterior cervical plate for cervical interbody fusions in a group of 26 patients. In this retrospective series of 26 patients, 41 disc spaces were considered 'incorporated' at 24 months. Radiographic fusion was impossible to judge because the graft material was not resorbed at 24 months and remained more radio-dense than native bone. The criteria used for the determination of 'incorporation' was loss of a radiolucent line around the graft. The validity of this radiographic criterion for the determination of bony fusion is questionable. Iseda et al. [19] found that the appearance or disappearance radiolucent stripe around a hydroxyapatite graft did not correlate with changes in the 99mTc-HMDP uptake ratio following surgery in a small group of patients.

Tricalcium phosphate is a synthetically produced ceramic which has a porous structure amenable to osteoconduction. It has been used alone and in combination with hydroxyapatite as a bone graft substitute. It is completely resorbed by the body and is gradually replaced by new bone formation [1].

It is virtually identical to bone in terms of its mineral composition, and no reports of systemic toxicity exist. A comparative study by Emery et al. [20] demonstrated that the use of tricalcium phosphate (pore size 400 |xm) resulted in superior histological results compared to hydroxyapatite or calcium carbonate in a canine interbody fusion model. Toth et al. [21] studied the use of different p-tricalcium phosphate/hydroxyapatite preparations in a goat cervical interbody fusion model and found that p-tricalcium phosphate performed as well or better than autograft at 3 and 6 months following fusion (radiographic, histological, dual-energy x-ray absorptiometry, and biomechanical testing). There was a definite increase in the bony union rate at 3 months with increasing porosity (0% union rate for autograft and 30% porous ceramic, 67% for 50% porous ceramic, and 83% for 70% porous ceramic). Delecrin et al. [1] used a similar compound (Triosite™ Zimmer, France) in a study of 58 patients undergoing surgery for scoliosis. In a prospective randomized study, these authors found that patients who received the ceramic graft had less blood loss and had no graft site complications. The authors noted that half of the autograft group experienced pain at the donor site 6 months following surgery. There were no significant differences between autograft and ceramic groups in terms of radiographic appearance or functional outcome at a minimum follow-up of 24 months (mean follow-up 48 months) [1].

A pure p-tricalcium phosphate synthetic bone product has been recently introduced (Vitoss™, Orthovita, Pa., USA). Vitoss has a unique crystalline structure, imparted by the manufacturing process, which increases its porosity significantly. The porosity of Vitoss is between 88 and 92%, compared to an approximate 55% porosity of hydroxyapatite-calcium carbonate. The increased porosity of this product allows for the rapid penetration of blood and blood products as well as for the migration of osteoblasts. Animal experiments using a humeral defect model have confirmed that the tricalcium phosphate is rapidly and nearly completely replaced by native bone in a matter of weeks (fig. 1, 2). The formation of bone in the humeral model is more rapid and more complete than that seen with coralline hydroxyapatite bone substitutes (fig. 1). In time, the graft is completely resorbed and replaced by new bone. When combined with autogenous bone marrow, the osteoinductive properties of the tricalcium phosphate are superior to demineralized bone matrix in a rat Urist pouch test [22]. A similar p-tricalcium phosphate compound has been developed (but is not yet FDA approved) that has biomechanical properties similar to cortical bone, can be machined and shaped easily, and can even be delivered in a semiliquid or putty form (Cortoss, Orthovita, Pa., USA).

Several other types of compounds have been used as synthetic bone substitutes. Madawi et al. [23] studied a bioactive osteoconductive polymer made of methyl methacrylate, calcium gluconate, and polyamide fibers. These authors

Time (weeks)

Fig. 1. Rapid absorption of Vitoss with regrowth of new bone within the graft.

Time (weeks)

Fig. 1. Rapid absorption of Vitoss with regrowth of new bone within the graft.

Fig. 2. Vitoss tricalcium phosphate bone substitute in a canine humeral defect model. The implant is almost completely resorbed at 6 weeks postimplant (a) and is indistinguishable from normal bone radiographically and histologically at 12 weeks (b).

noted favorable clinical results with this compound in trial of cervical interbody fusion; however, the graft never radiographically incorporated into native bone. Schulte et al. [24] used a bioglass-polyurethane composite to replace vertebral bodies following corpectomy for metastatic lesions in 5 patients. All patients also underwent titanium plate fixation. No instances of implant failure were noted. Roessler et al. [25] reported sobering results in their clinical study of a resorbable compound made up of polyethylene oxide and polybutylene tere-phthalate (Polyactive 70/30™ HC Implants, The Netherlands) for reconstitution of iliac crest bone graft harvest sites. Although this substance had shown very promising results in multiple animal models, human trials failed to show a favorable effect on clinical or radiographic outcome. This study serves as a reminder that there is no substitute for properly designed human clinical trials for the development of bone graft substitutes. Differences in the healing capabilities of the animal species, influences of patient age, and differences in the graft environment all play a role in bone healing [25, 26].

Resorbable Fixation Devices

Implants placed during spinal fusion surgery function in various ways. Screw/rod or screw/plate systems function to provide immediate structural stability, thus allowing early mobilization of the patient, maintaining spinal alignment and promoting the ultimate arthrodesis. Traditionally, these systems were manufactured of stainless steel, which has more recently been largely replaced by titanium alloy. Although both of these metals have performed well in their primary function, there are drawbacks such as stress shielding due to the excessive rigidity and permanence of the constructs that can in turn lead to bone resorption and osteopenia [27, 28]. Corrosion, wear debris and even rare allergic reactions have been seen, more so with stainless steel implants. Metallic artifact on imaging studies can obscure anatomic detail and cause diagnostic dilemmas. Once arthrodesis has occurred, the fixation systems no longer have a purpose [27, 29]. The indications for explantation of spinal fixation hardware are still controversial, but it is commonly performed.

Bioresorbable polymers have been increasingly explored as replacements for metal, bone and nonresorbable synthetic materials in recent years. Obviously, biodegradable implants eliminate the need for explantation. In addition, they can reduce stress shielding, by having a better match of strength and elasticity to bone as well as by gradually reducing load sharing as structural integrity is lost [27, 28, 30]. Lastly, they avoid the problems of metal corrosion, debris, and imaging artifact. This section will explore the properties and potential uses of these polymers in spinal surgery.

Bioresorbable materials and plates have been used to repair orbital and mandibular fractures since 1972 [31, 32]. Polylactic acid (L-isomer) (PLLA) and polyglycolic acid (PGA) and other polyhydroxy acids have been identified in the research to be promising polymers for resorbable implantable devices [33]. PLLA alone is resistant to hydrolytic degradation, which makes implants of this polymer quite strong but takes years for resorption to occur, if it occurs at all. These implants may also result in a foreign body reaction over time and thus require removal [34, 35]. In comparison, pure PGA implants absorb much more quickly when exposed to moisture yet suffer loss of tensile strength and rapid polymer breakdown [36]. Researchers have experimented with combinations of both polymers in order to utilize the favorable characteristics of each. LactoSorb™ (Biomet, Warsaw, Ind., USA) is a resorbable plate system approved for use in

Krebs TCA cycle

Fig. 3. Schematic illustration of the degradation of a prototypical resorbable polymer. (Certain intermediate products of degradation can also be excreted in the urine.)

craniomaxillofacial surgery. It consists of a combination of synthetic PGA and PLLA. This PGA/PLLA copolymer is manufactured in an 82% PLLA and 18% PGA ratio. The increased amount of PLLA in this copolymer gives it degradation characteristics midway between the predominately PGA products and pure PLLA implants [37]. MacroPore™ (MacroPore., San Diego, Calif., USA), another resorbable product with potential spine applications, is produced by the combination of PLLA and a noncrystalline copolymer mixture of poly(L-lactide) and poly(D,L-lactide).

The implant resorption time is most strongly affected by the chemical composition of the polymer, but there are many additional variables in this process [29, 38]. Gogolewski [28] has identified the implant's physical structure and mass, polymeric molecular weight, chain orientation, and presence of additives as some of the many factors influencing the rate of degradation, along with the stress on the implant and the characteristics of the implantation site. For example, he notes that an implant placed under significant load in a highly vascularized site will likely degrade at an accelerated rate [28].

The PGA component of the LactoSorb implant is degraded to glycolic acid by hydrolysis and further degraded to glycine, which may be used for protein synthesis or further converted to serine. All other implant (LactoSorb and MacroPore) components are hydrolyzed to lactic acid, which then enters the tricarboxylic acid cycle (fig. 3) [39]. The use of these plates avoids several potential complications associated with metallic fixation devices, such as migration, stress shielding, and interference with imaging techniques [39]. Currently, the most commonly reported side effect is a sterile inflammatory reaction to the implants [35, 40-43].

Local tissue reaction to polyhydroxy acid implants is governed by the chemical nature of the polymer, the physical characteristics of the implant, and by its degradation rate. Towards the latter part of the degradation process, the implant may rapidly lose structural integrity, and the production rate of polymeric debris may exceed the tissue tolerance and transport potential of the implantation site [27, 38]. This in turn can stimulate what is said to be a 'nonspecific foreign-body reaction' rather than a true inflammatory response [27, 38]. Thus, a large implant made of a fast-degrading polymer is likely to produce a more pronounced inflammation than a small implant composed of a slow-degrading polymer. PGA polymers have been associated with a higher rate of tissue reaction, even including the formation of fibrous capsules, sterile cysts and sinuses. On the other hand, pure polylactic acid (PLA) polymers have a low or nonexistent rate of tissue reaction, while once again, copolymers are intermediate [27, 38, 44].

Histological evaluation of these inflammatory sites has shown typical nonspecific foreign-body reactions. Similar foreign-body reactions were found on histological examination of implant sites of patients without clinical evidence of an inflammatory response [40]. These specimens were obtained upon reoperation for fixation failure. No explanation is available for why some patients have clinical signs of the foreign body reaction while others do not. Most studies have shown the inflammatory response, when clinically present, to occur several months after implantation. Bostman [40] postulates that this time to occurrence of the reaction is reflective of the estimated degradation time of PGA and PLLA/PGA implants which is 90-120 days. Patients who have had plates and screws made solely of PLLA, implanted for fixation of zygomatic fractures, have shown inflammatory reactions from 3.3 to 5.7 years after surgery [35]. These findings seem to correlate with degradation rates of PLLA, stated in some studies to be from 2 to over 4 years [39].

PLA polymers and PLA-PGA copolymers are biocompatible with the dura [45-47]. PLA biocompatibility has also been specifically tested in reference to neural tissue, including brain and spinal cord tissue as well as peripheral nerves. No effect on neuronal cells, nonneuronal cells or axonal growth has been noted [45, 48]. No significant toxicity, carcinogenicity, teratogenicity or mutagenesis has been associated with either PLA or PGA [27, 38].

LactoSorb plates and screws became commercially available for craniofacial applications in 1996. Most of the interest in bioresorbable plates for cranial reconstruction has been in the pediatric patient population because of complications with metallic implants related to bone growth. LactoSorb plates have been used in neurosurgery for fixation of craniotomy flaps, repair of depressed skull fractures and for repair of craniosynostosis [33, 49-53]. Encouragingly, it has been found that the dissolution and gradual loss of tensile strength of the devices minimize growth restrictions as well as the potential for transcranial migration [54]. In addition, good results have been obtained in both the pediatric and adult population when LactoSorb plates and screws were used for cranial fixation and reconstruction.

LactoSorb plates are designed to withstand similar external forces and approximate the strength of traditional titanium plates. In order for a bioresorbable plate to perform similarly to a titanium implant, it must retain similar strength characteristics long enough for osteosynthesis to occur. Because PGA, PLLA, and copolymers are all significantly less rigid than titanium, the design of plate-like implants must be adjusted somewhat. In general, resorbable plate designs incorporate a thicker plate, in order to recapitulate the mechanical properties of a standard titanium implant. Flexural strength is significantly improved in the LactoSorb plate, for example, due to its 'I-beam'-like construction (raised rails along the side of the plate). In vitro studies of flexural strength show the LactoSorb plate has an initial strength comparable to Lorenz titanium plates and retains 70% of this strength for 6-8 weeks [55, 56]. The improved strength does, however, create a much higher profile implant than the traditional metallic plate.

The most common reported clinical difficulties with the LactoSorb products are related to implantation technique. LactoSorb screws are not self-tapping which increases the time necessary to place the screws. Second, unlike metal plates, which can be shaped by hand or with instruments at room temperature, polymer plates must be heated before they can be shaped. Heating is achieved by use of a bag of calcium chloride which, when injected with sterile water, will liberate heat for approximately 20 min. By placing the bag around the LactoSorb plate, the implant can be made malleable. Some authors have stated that the 20-min window may be too small, making a second calcium chloride bag necessary, which involves adding time and expenses. Despite these initial technical difficulties, most authors agree that the LactoSorb system is a safe, effective system for cranial fixation and reconstruction, which, after minimal experience with the system, adds very little time to the operation compared to the traditional titanium plating systems.

Six studies of the use of bioresorbable plates for cranial fixation, either alone or in conjunction with facial reconstruction and orthosis, have been published with a combined total of 115 patients [33, 49-53]. No patients were noted to have significant complications from inflammatory reactions, seizures or failure of the fixation. In fact, clinical studies of patients who have had LactoSorb plates implanted and have been followed for greater than 9 months show no evidence of inflammatory reactions [49, 57]. It is believed that the intermediate rate of degradation for this predominately PLLA copolymer allows for efficient removal of its byproducts without overwhelming the body's mechanisms of clearance [50, 58].

Because of its relatively high profile, LactoSorb can produce a palpable and sometimes visible implant. Studies have noted this factor to occur occasionally; however, the authors state that in all cases the external evidence of the

Fig. 4. Explanted LactoSorb plate. Histological examination of explanted LactoSorb plate demonstates complete resorption of implant material 1 year following surgery (bone on right, normal soft tissues on left) without evidence of persistent inflammatory reaction.

Table 1. Strength of MacroPore implant during degradation

Time after implantation %

Implantation 100

6 months 90

9 months 70

12 months 50

18 months 0

implant disappears within 3-6 months [33, 51, 59]. Most studies of LactoSorb plates show the entire implant to be 95% resorbed by 9 months and completely resorbed by 1 year (fig. 4) [60].

MacroPore is characterized by a degradation time of 18-36 months. The loss of strength during degradation has been well characterized, and is predictable [61] (table1). As noted above, however, the strength and degradation characteristics of any specific MacroPore implant will be influenced by the manufacturing processes, implant size and geometry, and characteristics of the implantation site, among other factors. MacroPore can be formed into a wide variety of shapes, and it can be stored and sterilized utilizing conventional techniques. In appropriate shapes, MacroPore can be heated and shaped for conformation to the actual site of implantation, and it will hold the desired shape once cooled without loss of structural integrity [MacroPore: pers. commun.].

Fig. 5. Absorbable 'spacer'. This resorbable spacer represents one possible avenue of exploration for resorbable technology for spine surgery.

MacroPore devices cannot be readily visualized on routine radiographic studies, although they can be seen distinctly on CT scans prior to significant degradation [Branch: pers. commun; 62]. PLA implants do not degrade MRI images, and MRI scanning has been utilized to evaluate the tissue response of PLA implants [30, 63-65]. PLA implants are visible on MRI images as areas of homogeneous low-signal intensity, which can be distinguished from the high signal intensity of the adjacent bone [63].

MacroPore sheets have been utilized to reconstruct iliac crest donor site defects [66]. Iliac crest reconstruction may diminish pain, prevent bowel herniation through large defects [67], improve cosmesis, and optimize donor site regeneration [66]. The benefits of a protected healing space have been recognized in promoting optimal bone healing [47, 68]. Specifically, if soft tissue is prevented from prolapsing into a bone defect, the regrowth of bone may be better than with graft materials alone. In terms of the iliac crest, once the bone is harvested, the donor site is backfilled with allograft or bone matrix material if desired. A MacroPore sheet is then heated to 70°C, contoured to the defect site and allowed to cool. It is then secured with screws or tacks [66]. This is an example of the potential use of MacroPore as a barrier type of implant.

MacroPore implants are being utilized in pilot studies as load-sharing implants in spine fusion constructs [62]. The versatile nature of this material allows it to be formed into 'cages', 'dowels' and 'interbody spacer' shapes (fig. 5). This, coupled with the desirable strength and degradation characteristics, lack of artifact on imaging, low potential for foreign body reaction and the biocom-patibility with the dura and nervous tissue, makes it a very promising material.

One such study examines the use of MacroPore devices in instrumented posterior lumbar interbody fusion constructs [62]. The devices are placed following complete discectomy, along with morsellized autograft bone, and maintain the disc space height during the early phase of bone healing. Preliminary results show equivalent clinical outcomes to those obtained with the use of allograft bone spacers. The devices produce no artifact on postoperative imaging, allowing better visualization of the maturing bone fusion [62].

Other products are being developed worldwide in order to expand the applications of bioabsorbable products. Bostman et al. [40] used a biodegradable rod for internal fixation of extremity fractures and osteotomies. The composition of the rods was either polyglycolide or a lactide-glycolide copolymer (polyglactin). A small percentage of the 516 patients treated required reoperation for fixation failure (1.2%). The incidence of bacterial wound infections was low (1.7%). The most significant complication was a foreign-body reaction in 7.9% of the patients. The reaction caused a painful, reddened, fluctuant swelling at the operative site 2-4 months after surgery. Godard et al. [69] have investigated a bioresorbable implant for spinal arthrodesis, which is reported to undergo ossification. Brunon et al. [70] have used a bioresorbable plate developed from PLLA for anterior cervical interbody stabilization in 5 patients. The development of bioabsorbable plates, screws, and rods for use in the spine is an area of intense activity at the present time.


Absorbable bone substitutes and fixation devices are being developed for spinal surgery. Advances in material technology may allow for the avoidance of autograft harvest morbidity and allograft preparation costs and toxicities. Resorbable implants may result in decreased morbidity and better radiographic evaluation of spinal fusion procedures. The pace of new development in this field is remarkable. As new products are brought to market, thorough and rigorous testing in well-designed clinical trials will be required to establish their ultimate worth.


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Daniel K. Resnick, MD MS

Assistant Professor, Department of Neurological Surgery University of Wisconsin, K4/834 Clinical Science Center 600 Highland Ave, Madison, WI 53792 (USA)

Tel. +1 608 263 9651, Fax +1 608 263 1728, E-Mail [email protected]

Haid RW Jr, Subach BR, Rodts GE Jr (eds): Advances in Spinal Stabilization. Prog Neurol Surg. Basel, Karger, 2003, vol 16, pp 55-70

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