Medical titanium is one of the most widely used functional implant materials. Low density, high specific strength and a stable passive film make titanium and its alloys suitable for aerospace, energy equipment and surgical devices alike. For medical use the decisive properties are biocompatibility, corrosion resistance in body fluid, an elastic modulus as close as possible to that of bone, and the ability to be processed into complex shapes without losing mechanical integrity.
Development Generations of Medical Titanium Alloys
The history of medical titanium falls into three periods. The first generation was represented by unalloyed titanium and Ti-6Al-4V. The second generation moved to alpha-plus-beta alloys such as Ti-5Al-2.5Fe and Ti-6Al-7Nb, which reduced or eliminated vanadium for better biological tolerance. The third generation is defined by metastable beta alloys developed for better biological performance and a lower elastic modulus, and this approach now dominates mainstream medical device development.
Domestic research began in the 1970s, when a nonferrous metals research institute developed the Ti-2.5Al-2.5Mo-2.5Zr alloy. In the 1990s, independently developed Ti-6Al-4V, Ti-5Al-2.5Fe and Ti-6Al-7Nb materials became available, and later work produced the metastable beta alloy Ti-24Nb-4Zr-7.6Sn, whose low modulus and niobium-rich composition avoid both vanadium and aluminium. Breaking through into new alloy systems while broadening the clinical application of existing ones remain the two main directions of development.
Corrosion Behaviour of Titanium in the Human Body
Titanium is thermodynamically unstable, with a standard electrode potential of about -1.63 V, yet the metal passivates easily. In air and in aqueous solution it forms an oxide layer only a few nanometres to several tens of nanometres thick, and that layer delivers the corrosion resistance the metal is known for.
The corrosion resistance of a medical material matters for two reasons. First, metal ions and corrosion products released from an implant penetrate surrounding tissue and can trigger physiological reactions of varying severity. Second, body fluid can severely reduce the performance of a material and lead to rapid damage or even failure. The body environment is chemically complex, and it promotes dissolution of trace elements and changes in the stability of the oxide layer. Even slight friction damages the passive film; in an oxygen-poor region the film is less stable and cannot be repaired immediately once damaged, which makes localised corrosion more likely. Repeated movement of the body and the use of surgical instruments make that situation almost unavoidable.
Plastic deformation matters as well. Deformation alters the microstructure and therefore the corrosion performance of the material, and the effect varies greatly with the degree of deformation. Internal stress concentrations create interfacial and grain boundary defects, so increasing plastic deformation weakens corrosion resistance.
Corrosion Mechanisms and Stress Corrosion Cracking
The passivating film reduces the area available for active dissolution and slows the dissolution rate, and it can also repair itself after damage. Corrosion of titanium implanted in living tissue can nevertheless take several forms: pitting corrosion, stress corrosion cracking, crevice corrosion, galvanic coupling corrosion and wear corrosion.
Stress corrosion cracking is the rupture of a metal under the simultaneous action of tensile stress and a corrosive medium. Tensile stress first ruptures the protective film, producing a pitting or crevice site that acts as a crack origin; the crack then grows into the depth of the metal, and continuing stress repeatedly breaks the film until a crack perpendicular to the stress direction forms and the component fails.
SCC of titanium alloys is the result of three factors acting together: environment, stress and material. The phenomenon is highly selective, and changing any one of the three is usually enough to prevent it.
Medium. Titanium alloys can suffer SCC in many aqueous solutions, in distilled water, in organic solutions and in hot salts, and the mechanism differs between media.
pH value. Views on the effect of pH differ, but sensitivity generally falls as pH rises; at pH 13 to 14 SCC is often inhibited, while the local chemistry inside a crack can reach pH 2 to 3 even when the bulk solution is neutral.
Potential. The effect of potential is critical. Beta alloys in halide solutions show aggravated SCC near -600 mV, while no cracking occurs below about -1000 mV. Ti-8Al-1Mo-1V is sensitive between -500 mV and -600 mV in chloride and bromide solutions, and in iodide-bearing solutions the sensitive range moves above 0 mV.
Temperature. Sensitivity increases with temperature in general. In hot salt and air environments, near-alpha alloys become SCC sensitive above roughly 450 degC. Ti-6Al-4V containing palladium or molybdenum is less sensitive in a mixed H2S, CO2, NaCl and sulphur solution at 200 degC than at 250 degC. For implants the temperature range is narrow, so this factor is of limited practical importance.
Chloride concentration. The higher the chloride concentration in solution, the greater the SCC sensitivity.
Stress level. Residual stresses from cold working, forging, welding, heat treatment or assembly account for a large share of SCC failures, reported at about 40 %. External service stresses, and stresses generated by the volume change of corrosion products, act in the same way: the higher the stress level, the shorter the time to cracking.
Material condition. Chemical composition, segregation, microstructure, grain size, crystal defects, heat treatment and surface condition all change susceptibility. Small additions of palladium, molybdenum or ruthenium reduce it. Peak-aged Ti-6Al-4V and Ti-15V-3Cr-3Al-3Sn are more sensitive than the annealed condition, and holding the oxygen content of Ti-6Al-4V below 0.13 % greatly reduces SCC sensitivity.
Three practical countermeasures follow. Residual stress can be removed by full or local annealing after manufacture, provided the effect on strength, ductility and toughness is acceptable. Alloying with palladium, molybdenum or ruthenium raises resistance in the specific medium that causes the problem. Surface treatment improves surface quality, biocompatibility and wear resistance, and therefore delays crack initiation.
Crevice Corrosion and Wear Corrosion
Crevice corrosion occurs when a medium becomes trapped in a gap between a metal part and another metal or non-metallic component. Oxygen inside the gap is consumed faster than it can be replaced, so the crevice becomes anodic and the passive film is destroyed. The process normally passes through three stages: consumption of oxygen inside the crevice, formation of a macro-cell with a fall in pH, and activation and dissolution of the passive film until it is completely destroyed. In Hank's solution at 37 degC the measured crevice corrosion severity of implant materials decreases in the order NiTi, NiTiCu, 316L stainless steel, then Ti-6Al-4V and pure titanium together, so unalloyed titanium and Ti-6Al-4V have strong crevice corrosion resistance in simulated body fluid.
Wear corrosion is the accelerated attack caused by relative motion between a loaded metal surface and the surrounding medium at high velocity. When an implant is inserted it is abraded to some degree by the instruments used, and that abrasion damages the existing oxide film. If the film is not repaired in time, the implanted metal corrodes further and can eventually fail. Polishing the surface, avoiding mixed-metal junctions and minimising instrument contact all reduce the risk.
Grade Selection and Property Comparison
| Material | Nominal composition | Elastic modulus | Typical use |
|---|---|---|---|
| Pure titanium Gr2 | Unalloyed, controlled oxygen | About 105 GPa | Dental components, non-load-bearing devices |
| Pure titanium Gr4 | Unalloyed, higher oxygen | About 105 GPa | Dental implants, fixation hardware |
| Ti-6Al-4V (Gr5) | Al 5.5-6.75 %, V 3.5-4.5 % | About 110 GPa | Bone plates, screws, femoral stems |
| Ti-6Al-4V ELI (Gr23) | Same base, oxygen 0.13 % max | About 110 GPa | Load-bearing implants, fracture fixation |
| Ti-6Al-7Nb | Al 6 %, Nb 7 %, vanadium-free | About 105 GPa | Cementless stems, vanadium-free devices |
| Ti-13Nb-13Zr | Near-beta, niobium and zirconium stabilised | About 79 GPa | Hip stems, low-modulus applications |
| Ti-24Nb-4Zr-7.6Sn | Metastable beta | Below about 60 GPa | Low-modulus research and next-generation implants |
The relevant product standards are ASTM F67 and ISO 5832-2 for unalloyed titanium, ASTM F136 and ISO 5832-3 for Ti-6Al-4V ELI, ISO 5832-11 for Ti-6Al-7Nb, ASTM F1713 and ASTM F1813 for the low-modulus beta alloys, and GB/T 13810 for wrought titanium materials for surgical implants. Cortical bone has an elastic modulus of roughly 10 GPa to 30 GPa, so the mismatch between implant and bone remains the central design constraint, and stress shielding is the reason beta alloys with a modulus below 80 GPa attract so much interest.
Frequently Asked Questions
Q: Why does a low elastic modulus matter for a titanium implant?
A modulus close to that of cortical bone reduces stress shielding, so the bone around the implant keeps carrying load instead of resorbing.
Q: Is Ti-6Al-4V still used in implants?
Yes, but the extra-low interstitial grade Ti-6Al-4V ELI is preferred for load-bearing devices because lower oxygen content improves toughness and SCC resistance.
Q: What causes stress corrosion cracking in titanium implants?
It needs tensile stress and a specific corrosive medium together. Residual stress from machining or welding is the usual stress source and can be reduced by annealing.
Q: How does the passive film on titanium repair itself?
The film reforms spontaneously in any oxygen-containing medium, but inside an oxygen-depleted crevice or under continuous abrasion, repair cannot keep pace with damage.
Q: Does plastic deformation affect implant corrosion performance?
Yes. Deformation concentrates internal stress at interfaces and grain boundaries and weakens corrosion resistance, and the effect grows with the amount of deformation.
Q: Which alloying elements improve SCC resistance?
Small additions of palladium, molybdenum or ruthenium are the standard approach, combined with a vanadium-free or low-oxygen base composition.
Q: How is crevice corrosion avoided in implant design?
By eliminating tight gaps, avoiding mixed-metal junctions and selecting a grade whose film can repassivate where oxygen is scarce.
