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.¹

FIGURE 70-1 Tibial plating with an example of metallic implant.

The use of implantable material is not new in orthopedics. For centuries, 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.

Biocompatibility, or the clinical success of a biomaterial, is directly 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.

This chapter will review the major implantable materials in orthopedics and biologic or tissue responses that may occur. The section is broken down into metals, polymers, hydrogels and biologics.

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).

TABLE 70-1 Common Tissue-Implant Interactions

Implant-Tissue Reaction	Consequence
Toxic	Tissue necrosis
Biologically inert—smooth surface	Implant is encapsulated without bonding
Biologically inert—porous surface	Tissue grows into pores and forms mechanical bonds
Bioactive	Tissue forms interfacial bond with implant (bioactive fixation)
Dissolution of implant	Implant resorption and replacement with soft tissue or bone

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 short-term 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-bearing 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.

TABLE 70-2 Relative Metallic Properties

Metal Types

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).

FIGURE 70-2 Example of pitting corrosion.

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.

FIGURE 70-3 Example of electrochemical cell around metallic implant.

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 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 demonstrated 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.

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 self-lubricating 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 foreign-body 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, 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 water-soluble 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 hydroxyethyl-methacrylate (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 thermoplastic 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. These 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. The 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. Highly-hydrated 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 cost-effective.

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).

TABLE 70-3 Bone Grafting

Bone Graft Family	Description	Classes
Osteoconductive	Provides structure or scaffold to support bone formation	Ca-Phosphate, ceramics, synthetic polymers, bioactive glass, allograft
Osteoinductive	Induces differentiation of cells and new bone growth	BMP, demineralized bone matrix
Osteogenesis	Provides stem cells with osteogenic potential that directly form new bone	Bone marrow aspirate
Combined	Combinations of above	Autograft