The unique combination of material properties found within orthopedic alloys is the result of the metallic bonds formed between metal atoms. The positively charged nuclei of metal atoms form a crystal lattice structure and are held together in a sea of valence electrons (electrons of the outer orbital shells). The resulting balance between positive and negative charges results in overall neutrality. This freely flowing sea of valence electrons within any metal solid is responsible for the related material properties of high thermal and electrical conductivity. Although the nuclei of metal atoms are held within a sea of electrons, these nuclei are fixed in closely packed crystalline arrays, forming a distinct arrangement. The more closely packed these metal nuclei, the stronger the resultant bonding. The structure of the metal crystals can be broken down to a three-dimensional repeating unit (a unit cell) that generally takes one of three configurations (Fig. 11-1): face-centered cubic (FCC), body-centered cubic (BCC), and hexagonal close packed (HCP). The nondirectionality of the metallic bond allows these unit structures to be stretched, deformed, broken, and reformed, permitting defects (dislocations) to pass through the structure. Alloying elements can be used to fill the spaces between metal atoms—also known as interstitial elements—or replace the base alloy element within the lattice—also known as substitutional elements. Either way, alloying elements change or distort the basic crystal structure of a metal to alter or enhance the material properties by hindering dislocation movement throughout the lattice (e.g., the addition of 6% aluminum [Al] and 4% vanadium [V] are used to enhance the mechanical properties of Ti).

As metals cool from a liquid state, crystals begin to form at nucleation sites within the liquid. These crystals grow, forming a granular structure (or a polycrystalline array). Such crystal grains can be observed microscopically on the surface of a polished metal specimen (Fig. 11-2). The microstructural morphology (grain size, shape, etc.) predominantly determines the mechanical properties of the metal. Finer grain sizes generally result in increased yield strength, fatigue strength, and fracture toughness, thereby decreasing the chances for implant fracture. Insufficient quality control in the manufacture of earlier total hip arthroplasty (THA) stem designs demonstrated that large grain sizes within the stems resulted in decreased fatigue strength, which resulted in fracture of the implants in vivo. All the grains within a pure metal have the same crystal structure. This “single phase” of a chemically homogeneous metal can be maintained, in some cases, even after the addition of other metals with similar atomic size. However, if added elements are not of similar size or atomic structure, then additional phases develop. Therefore, in a multiphase metal, there will be two or more distinct types of grains or crystal structures.

Although metal alloys do contain a variety of phases and crystal structures in practice, there are usually only a few dominant phases present. The three aforementioned alloys used in TJA usually exhibit the following phases: Ti-based alloys are a mixture of an alpha-phase (HCP) and a beta-phase (BCC) (Fig. 11-3). Co-based alloys are generally single phase (FCC) after casting, but an HCP structure can be introduced by cold working of the alloy. Additional carbidic or intermetallic hard phases may occur depending on the carbon content and heat treatment. Implant-grade stainless steels are so-called austenitic steels comprised of the FCC phase. The amount of
carbon in the steel is intentionally kept to low concentrations, because it can form additional phases, which may settle at grain boundaries and lower the overall strength of the alloy.¹,²,³,⁴,⁵,⁶,⁷,⁸

All metal alloy implants corrode in vivo. When severe, the degradative process may reduce structural integrity of the implant, and the release of corrosion products is potentially toxic to the host or can cause adverse local tissue reaction and subsequent implant failure.⁹,¹⁰ Electrochemical corrosion of implants includes generalized forms of corrosion uniformly affecting an entire surface and localized forms of corrosion affecting areas of a device relatively shielded from the environment (crevice corrosion) or at seemingly random sites on the surface (pitting corrosion).

Metal corrosion is governed by the thermodynamic driving forces that cause corrosion (oxidation-reduction) reactions and the kinetic barriers that limit the rate of these reactions. The chemical driving force (ΔG) determines whether corrosion will take place under the conditions of interest. If the free energy for oxidation is less than zero, then oxidation is energetically favorable and will take place spontaneously. During corrosion, positive and negative charges (metal ions and electrons, respectively) leave one another for more chemically stable partners. The metal ions generally leave to form an oxide or another more stable ionic compound (or are released into solution), and the electrons are left behind in the metal
and undergo other electrochemical reactions on the surface such as the reduction of oxygen or hydrolysis of water. A charge separation across the metal-solution interface contributes to what is known as the electrical double layer and creates an electrical potential (much like a capacitor).

The corrosion resistance of implant alloys to this electrochemical driving force is primarily due to the formation of surface barriers that limit implant corrosion. Were it not for the protective barriers that form on the surfaces of implant alloys, vigorous corrosion would take place. Kinetic barriers prevent corrosion by physically limiting the rate at which oxidation and reduction processes can take place. Alloys used in orthopedic implants rely on the formation of passive films to prevent significant electrochemical dissolution from taking place. These films consist of metal oxides that form spontaneously on the surface of metals in such a way that they prevent further transport of metallic ions or electrons, or both, across the film. Passive films must have certain characteristics to limit further oxidation:

The oxide film will change crystal structure, size, and thickness depending on the conditions of the electrolytic solution (e.g., Co-Cr alloy in serum will form a thin oxide layer (2 to 5 nm) versus the thick oxide layer (10 to 20 nm) formed in nitric acid). Thus, various surface treatments, known as passivation, have been used to improve the barrier effect of the oxide film. Typically, such treatments have included a hot 35% nitric acid bath, boiling in distilled water, and anodization. Implant metals are generally passivated using a series of steps that include (1) thorough cleaning using detergent, ultrasound, and heat; (2) cleaning again with ethanol; (3) rinsing in deionized water; (4) treating with a dilute acid solution (e.g., 35% nitric acid); and (5) rinsing and sterilization. However, the development of treatments to optimize the shape, integrity, and protective ability of oxide films of the various implant alloys remains an ongoing process.

One of the current issues associated with implant alloys is the corrosion observed around the modular junction connections of retrieved joint replacement components, which use metal-on-metal conical taper connections (Fig. 11-4).¹⁰,¹¹,¹² Dual-modular stems contain additional potential corrosion sites due to neck-stem modularity.⁹ Accelerated corrosion can take place in the crevices
formed by these tapers in vivo due to (1) the occurrence of micromotion, fretting, and subsequent abrasion of the passive film and (2) the depletion of oxygen. Corrosion has been observed in the Co-based alloy systems used in these modular taper connections through such mechanisms as intergranular corrosion, fretting corrosion, phase boundary corrosion, etching, and selective dissolution of Co.¹¹,¹² Oxide-induced stress corrosion cracking and selective corrosion of the beta-phase have also been observed in Ti alloy stems.¹⁰,¹³ However, stainless steel alloys generally corrode to a greater extent than Co or Ti alloys (Table 11-2).⁴,⁵,¹⁴,¹⁵,¹⁶

Stainless steels were the first metals to be widely used for orthopedic applications in 1926. However, it was 1943 before American Society for Testing and Materials (ASTM) 304 was developed as a standard implant material. All steels are comprised of Fe and carbon and may contain Cr, nickel (Ni), and molybdenum (Mo). Trace elements such as manganese (Mn), phosphorous, sulfur, and silicon are also present. Carbon and the other alloying elements affect the mechanical properties of steel through alteration of its microstructure.