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.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

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.

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.

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