Corrosion of Ti?6Al?4V alloy medical implants and prosthetics under load




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            Medical implants have had a remarkable
impact in patients suffering from many conditions since its development.
Nonetheless, current advancements in metallic biomaterials used in implants
have a limitation, and it is corrosion. Corrosion is a chemical and physical
degradation of materials due to reaction with environmental factors. Metallic materials
have the tendency to fail by corrosion and metallic biomedical implants are no exception.
While there are have been developments in the increase of corrosion resistance
biomaterials like Ti?6Al?4V alloy,
corrosion continues to be a challenging problem due to multiple biological,
chemical and mechanical factors.                                                                           
                                   Titanium (Ti) alloys, particularly, Ti?6Al?4V, are characterized by high
strength, excellent biocompatibility, corrosion resistance and low density. Ti?6Al?4V is widely used for the
fabrication of medical implants and prosthetics. In addition to biomedical
applications, Ti?6Al?4V have
also been used aircraft and automobiles. Atoms in Ti?6Al?4V alloys include titanium (Ti), aluminum(Al), vanadium (V),
oxygen (O), carbon (C), hydrogen (H), and iron (Fe). The approximate
composition of these atoms are found in Table 1. The micro addition of Al and V
increases grain refinement and strength. Grain refinement is responsible for
the reduction of wear volume loss from a heat treated Ti alloy (Chandramohan
et. al., 2017). In other words, a weaker Ti?6Al?4V alloy produces more wear volume resulting in a rough
surface. Corrosion rate of Ti?6Al?4V increases
as the surface roughens (Dalmau et al., 2013). A roughened implant surface can
be indicative of abrasive wear during loading. The reduction of the alloy
performance by load force, degradation, and corrosion is inevitable.  The types of corrosion mechanisms observed in
loaded Ti?6Al?4V alloy
medical implants and prosthetics are wear-corrosion, fretting crevice corrosion,
selective leaching, and pitting corrosion.



Table 1. Composition of Ti?6Al?4V alloy (Tamilselvi, Raman, and Rahendran, 2006).

Ti?6Al?4V implants are used to assist patients suffering from
osteoarthritis. Osteoarthritis is the inflammation of joints from gradual loss
of cartilage covering the joint bones. Ti?6Al?4V is widely used for the fabrication of acetabular cup in hip
implants (Lomholt, Pantleon and Somers, 2010). The artificial implant is inserted
into the synovial fluid which is found between the acetabular cup and the
femoral head. Synovial fluid is composed of plasma, hyaluronic acid and glycoproteins.
The interaction of the protein and metal ions changes the corrosion rate of the
material (Yan and Neville, 2011).  According
to Hsu et. al. (2004), the corrosion resistance is lower in joint fluid than in
urine or serum because the joint fluid has chloride, protein ions and higher pH
content. The failure of Ti?6Al?4V
implant by corrosion begins with the combination of the mechanical wear by
fretting corrosion and the changes in the synovial fluid composition.

The reaction between oxygen and the
alloy when it is exposed to air or water contributes to formation of oxidation
layer and forms protective layer close to 10nm. (Lomholt, Pantleon and Somers,
2010). The Ti alloy implant forms the layer of oxide which controls the
diffusion rate of ions, and electron. Ti?6Al?4V alloy has a low d-bond character which is reactive with oxygen
and hydrogen (Lomholt, Pantleon and Somers, 2010). One result of the reaction
between the alloy and hydrogen is hydrogen embrittlement. Hydrides are products
formed within Ti?6Al?4V ?-phase when Ti?6Al?4V implant is exposed to H2O (Lomholt, Pantleon and Somers, 2010). The liberated
hydrogen ions lead to the hydrogen embrittlement. In addition to hydrogen
embrittlement, high amounts hydrogen ions can also reduce the pH of fluid
surrounding the implant. According to Yan and Neville (2011), the corrosion
rate of Ti alloy is suppressed in slightly basic joint fluid. So the hydrogen
ions change the acidity of synovial fluid. An acidic environment induces
selective leaching of the beta phase and reduces the corrosion resistance of
the Ti?6Al?4V
implant with or without loading forces (Gilbert et. al., 2012).

Fretting corrosion of the Ti?6Al?4V prosthetics begins with
deterioration of the surface. The shear stresses from normal walking on the hip
joint causes the cartilage to wear leading to the pain in osteoarthritic
patients. A similar mechanism of wear occurs in artificial hip implants. The
surfaces of the artificial acetabular cup and the femoral head which are always
in contact during normal loading activities eventually begins to show signs of
wear loss. The surfaces become rough from the removal of wear debris. Corrosion
rate of Ti?6Al?4V implant
increases as the surface increases in roughness. Frictional and corrosive
behavior of the alloy are influenced by the surface roughness and grain size
(Dalmau et al., 2013). Mechanical wear from a rigorous lifestyle can also increase
the rate of corrosion in the alloy. The debris product from fretting corrosion can
have an effect on the chemical composition of the synovial fluid. Changes in
composition of joint fluid can also lower the corrosion resistance of Ti?6Al?4V implant.

loading on Ti?6Al?4V
implant removes the protective layer which exposes the alloy for further
corrosion. Continuous loading from normal human walking strides can lower the
ability for the alloy to repassivate after passivation and produce wear debris. The non-biocompatible wear debris are toxic
in the body and can cause metallosis. The wear debris is also factor in
depassivation and the reduction in repassivation. The wear products from
continuous loading can get trapped and move between two interfaces causing
abrasion wear. This type of wear where the particles move by rolling and
sliding between two surfaces is known as three-body wear. Figure 1 is an
illustration of how the wear particles remove oxide film by the three-body
abrasion. When the abrasive damage from walking increases, it changes
from abrasive to adhesive wear and leads to the formation of corrosion pits
(Komotori, Hisamori, and Ohmori, 2007). Pitting corrosion is another form of
corrosion penetrating deep from a small surface area of the material. The depth of the pits extends as far as 150mm below a failed
Ti?6Al?4V hip
implant (Gilbert et. al., 2012). Pitting
corrosion begins after the passive layer degrades and growth of the unstable
pits increases after repassivation (Hsu et. al., 2004).

Figure 1. The
removal of passivation layer by three-body rolling abrasion (Wood and Thakare,

            The service
life of the implant can be controlled if the alloy is designed and fabricated for
favorable properties.  Microstructure,
and process treatment influences the longevity of implants. The cryogenic
treatment and long soaking treatment causes the grain size refinement which
improves wear resistance properties of the alloy. The hardness of cryogenic treated
Ti?6Al?4V alloys increased Vickers
hardness values by approximately 8HV from room temperature, and the 17HV when the Ti?6Al?4V alloys were soaked for 72
hours (Gu, Wang and Zhou, 2014). A stronger material with finer grains has more
grain boundary and dislocation sites. The loading forces which may be shear and
frictional forces have to be significantly higher than the forces at these
regions before materials are removed thereby decreasing wear loss. Grain
refinement produced by heat treating Ti alloy is also responsible for the
reduction of wear volume loss (Chandramohan et. al., 2017). Heat treatments
like annealing introduce nuclei that nucleate, recrystallize and grow into new
fine equiaxed grains. The new equiaxed grains increase the ability of the
strengthened alloy to resist deformation, wear and cracks. Both cryogenic
treatment and annealing treatment have the potential of improving the
durability of Ti?6Al?4V hip

The longevity of a Ti?6Al?4V implant can also be increase
by decreasing the rate of corrosion. Oxidation resistance is increase by
reducing the diffusion of metallic ions or the oxides through Ti passive layer.
The thicker the passive layer, the lower the oxidation. To promote the
formation of an anatase/rutile oxide layer on Ti?6Al?4V, the alloy is treated at 600°C and 650°C (Güleryüz and
Çimeno?lu, 2004). However, oxide film is no longer protective and unstable when
there are defects on the surface of the layer. Defects like cracks or vacancies
on an oxide layer increase the diffusion of ions which can affect oxidation
rate.  At high temperatures, the ions are
rapidly diffused through the defective regions in the oxide film. The stability
of the oxide layer is weakened by high rate of oxidation (Güleryüz and
Çimeno?lu, 2004). Heat treatment can have deleterious effects if the rate and
temperature used are not carefully controlled. Heat treatment can cause uneven
oxide film and surface roughening which contribute to corrosion. Other
techniques like excimer laser treatment as well as nano-surface coatings like diamond
film and Al2O3 -TiO2 modified Ti?6Al?4V surface, offer greater
corrosion resistance by multiple folds increasing the implant life service (Manivasagam
et. al., 2010).

Ti?6Al?4V hip implant replacements are invasive and
expensive procedures. Corrosion is a major contributing factor to the decrease
in the service lifetime of the implant, increasing revision surgeries and cost.
There are several factors contributing to the corrosion and failure of Ti?6Al?4V
alloy medical implants and prosthetics under load. The failure of an artificial
hip joint is the result of the combination of wear and corrosion. The failure
mechanism by corrosion starts with fretting corrosion between the modular
interface which produces wear product. Then it is followed by the dissolution
of the beta phase in Ti?6Al?4V implant due to the chemical changes of the
joint fluid after the fretting corrosion. Though Ti alloys have the ability to
repassivate, continuous loading impedes the repassivation process and
contributes to early Ti?6Al?4V implant failure. After the removal of
passivation layer, there are also evidences of pitting corrosion and hydrogen
embrittlement. Cryogenic and heat treatment are a few examples of process
treatments which may improve the corrosion resistance of Ti?6Al?4V
implant. There are limited studies offering scientific explanation for the
preferential dissolution of the beta phase and the impact of process treatments
on corrosion. Further investigation could reveal ways in which corrosion can be
inhibited to increase the longevity of Ti?6Al?4V medical implant for years to come. 

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