An updated review on TiNi alloy for biomedical applications

1.1 History and applications

Titanium (Ti) and its alloys are well known for their high corrosion resistance, low elastic modulus, and good biocompatibility, which are much better than those of stainless steel. Ti-based intermetallic alloys exhibit excellent high-temperature mechanical properties, significantly low density, good creep resistance up to 1000°C, and high oxidation resistance up to 900°C. TiNi alloys often play an important role in the biomedical field. Thus, they have been studied as biomaterials since their discovery by Buehler et al. (1963). The porosities of TiNi alloys have been investigated as an implant resource and tissue ingrowth. In fact, TiNi alloys are first used in orthodontic treatments and the bone plate. Gradually, this intermetallic alloy is used more and more widely. The scientific and technological development of TiNi alloy for biomedical applications is shown in Figure 1.

TiNi alloys are naturally studied as a medical and metal material with superelasticity, shape memory, and good biocompatibility. The first medical application was developed in 1971. Andreasen and Hilleman (1971) have implanted a TiNi alloy superelastic orthodontic device for the first time. After that, TiNi alloys have become more and more widely used in medical devices. The current research status of TiNi alloys in medical devices is introduced as follows. The application of TiNi alloys in the treatment of bone fracture was explored by Kang et al. (2002), (2005). They have designed and built bone holders using novel TiNi wires to design two bone holders for 20 patients. Continuous X-ray films, complete blood count, and urine analysis were performed without abnormalities after surgery. This clinical study demonstrated that TiNi alloys are suitable for bone fracture and provided new weapons for surgeons. Then, Szold (2006) pointed out the current status and developments of TiNi alloys in surgical applications. It was shown that TiNi alloy products have been widely used in medicine, such as self-expanding disc spacers made by TiNi alloys, vascular stents, and Nitinol anastomosis clips for left and right lateral anastomosis. It was also pointed out that the mechanical behaviors of TiNi alloys are nonlinear and there is a problem of hysteresis in the finite element modeling process, which will be gradually solved in future research (Tissot et al., 2013; Razali et al., 2016). Poggio et al. (2015) studied the applications of TiNi alloys in the treatment of dental root canal and fully evaluated the advantages and disadvantages of TiNi alloys in the manufacture of the root canal system.

At present, more than 1000 tons of Ti alloys have been implanted in patients worldwide (Batalu et al., 2014). However, the manufacturing cost of TiNi alloys increases gradually due to the increasingly complex part microstructures for biomedical applications. Therefore, the applications and researches of TiNi alloy coatings are more and more extensive in the field of biomedical applications.

In recent years, many researchers have studied a lot of work in the medical applications of TiNi alloy coatings. Thus, it is necessary to summarize the preparation methods, performance advantages, and medical applications to expand the medical applications of TiNi alloy coatings.

1.2 Structure and mechanical properties

The various properties of TiNi alloys, such as shape memory, corrosion resistance, superelasticity, are related to their internal metallographic structure (O’Brien et al., 2011). TiNi alloys produce a memory effect due to the phase transformations between martensite and austenite during stress or temperature changes as shown in Figure 2. TiNi alloy medical products, such as artificial joints, are inevitably subjected to force deformation during use. The temperature of TiNi alloys changes during use. However, the microstructures of TiNi alloys are affected by external forces and temperature. Changes in the microstructure and crystal structures of TiNi alloys result in changes in their properties, such as biocompatibility, shape memory, and other properties. Therefore, it is important to measure and study the microstructures of TiNi alloys.

Figure 1:

Development of the science and technology of TiNi alloys.

Figure 2:

Shape memory of TiNi alloy: (A) thermal shape memory effect and (B) superelastic stress-strain behavior of NiTi (O’Brien et al., 2011). Copyright © Elsevier. License number 4556810296444.

There are generally two kinds of the crystal phases of TiNi alloys (Onda et al., 1992; Yamauchi et al., 1996; Mantovani, 2000): austenite and martensite. There are a large number of researches on the microstructures and types of twinning using transmission electron microscopy and high-resolution electron microscopy to characterize the interface of twinning at the nanometer scale (Nishida et al., 1995). The recoverable strain of TiNi alloy coatings is about 4.5% due to the devitrification and twinning process of martensite.

The initial states and microstructures of the annealed TiNi alloys were first studied in the wide temperature range by Kuntsevich and Pushin (2007) by transmission electron microscopy and X-ray diffraction technologies. The results showed that annealed TiNi alloys are formed to high-strength nanocrystals, and its martensite can undergo two kinds of transformation in B2↔R↔B19′ and B2↔R.

Gunderov et al. (2017) studied the microstructures and deformations of ultrafine grain and nanocrystalline TiNi alloys produced by severe plastic deformation (Gunderov et al., 2013) after high-pressure torsion (HPT) and annealing. HPT-treated samples were annealed at 550°C, and their strengths decreased but ductility increased due to the martensitic transformation of B2 to B19′ occurring in TiNi alloys. The microstructures of TiNi alloys were changed into a ribbon microstructure using the ECAP-Conform technique to improve the tensile strength of TiNi alloys to 1300 MPa and a ductility of 25% (Churakova et al., 2017).

It was also pointed out that the mechanical behavior of TiNi alloys is nonlinear and there is a problem of hysteresis in the finite element modeling process. The CAE method such as finite element is gradually introduced to design medical TiNi alloy products. The finite element simulation of TiNi alloy elastomer composites used as biomaterials was studied by Tissot et al. (2013) to solve the problem of nonlinear behavior and hysteresis of TiNi alloys during finite element modeling. A homogenization method was proposed based on the geometry of TiNi alloy fiber. The superelastic properties, hardening behavior, hysteresis, and anisotropy of TiNi alloys were analyzed by the finite element method. The three-dimensional model of superelastic TiNi alloy archwire in the orthodontic three-stent system was developed by Razali et al. (2016) using the finite element method to achieve an online measurement of the frictional force and the restoring force of the tooth movement in the stent system. The prediction of the bending model is in good agreement with the experimental results.

Good mechanical properties are required to meet the life and reliability requirements of the biomedical products. TiNi alloys are proven to have high wear resistance and low coefficient of friction, and the fatigue resistance of TiNi alloys is much better than traditional bioceramic materials. Therefore, it is necessary to investigate the mechanical properties of TiNi alloy coatings to expand the applications of TiNi alloy coatings in biomedical applications.

2.1 Biomedical properties and surface modification

Determining whether a material is suitable for biomedical applications requires the performances to meet biocompatibility, bioadhesion, biofunctionality, and corrosion resistance. TiNi alloys are materials with great potential for biomedical applications due to their excellent performances (Bombac et al., 2007; Clarke et al., 2010). However, it has a certain resistance in the applications as a biomedical material due to the possible corrosion in physiologic solutions and the dissolution of carcinogenic Ni elements. Therefore, it is necessary to investigate the surface modification of TiNi alloy coatings in depth to improve the biocompatibility of TiNi alloy coatings.

Since the 1990s, many researchers have conducted specific researches about this topic. The TiO₂ oxide layer plays an important role in the surface of TiNi alloys as a bioinert material. The oxide layer prevents Ni atoms from diffusing into the blood and improves the corrosion resistance of TiNi alloys (Firstov et al., 2002; Shabalovskaya, 2002; Tao et al., 2012).

TiNi alloys were investigated for cytotoxicity by Sevcikova and Goldbergova (2017). The investigators used sulforhodamine B assays (Gill et al., 2015), MTT assays (Bernard et al., 2011), expression profiles (Lu et al., 2009), and cell survival assays such as apoptosis assays to assess cytotoxicity (Yeung et al., 2010). Although the biocompatibility of TiNi alloys is still controversial to date, almost all data about the surface is a very good biomedical material.

2.1.1 Corrosion and surface modification

Corrosion is one of the most important factors determining the biocompatibility of TiNi alloys due to a large amount of Cl ions in body fluids. In complex body fluid environments, TiNi alloys are still subjected to wear and corrosion, which leads to the leaching of toxic and carcinogenic Ni elements. Therefore, improving the corrosion resistance of TiNi alloys has been extensively studied.

The surface treatments of TiNi alloys were studied by Villermaux et al. (1996), (1997), and the surface of TiNi alloys was treated by the excimer laser. In Figure 3, the treated surfaces are subjected to much lower corrosion than the untreated surface. The surface of TiNi alloys is homogenized by laser surface melting, and the N element penetrates into the surface of TiNi alloys, causing the hardening of TiNi alloys and the thickening of the oxide layer.

Figure 3:

Scanning electron microscopy of (A) corroded untreated plate and (B) corroded treated pate (Villermaux et al., 1997). Copyright © Elsevier. License number 4554800640225.

The surface homogenization treatment of TiNi alloys was studied by Venugopalan and Trepanier (2000), Shabalovskaya (2001), and Santipach and Honig (2010). The results showed that the passivated and electrolytically polished TiNi alloy disk is more resistant to static corrosion and repassivation (repair ability) than other biomedical materials, and the release of Ni is far below the dietary level.

Shabalovskaya et al. (2004) compared the corrosion properties of medical TiNi alloy wires treated by sand blasting and smooth fine-drawing with alloy wires with black oxide on the surface during the manufacturing process in 0.9% NaCl solution. The studies showed that blasting and smooth fine-drawing are not very effective at the original scale, even exacerbating long-lasting scaling and greatly reducing the corrosion resistance. The reason is that the local defects generated after blasting and smooth fine-drawing lead to the unevenness of the alloy surface, thereby causing local corrosion. Therefore, eliminating the unevenness of the surfaces is an important direction to improve the corrosion resistance of TiNi alloys.

The corrosion resistance of pulsed high-energy density plasma (PHEDP)-modified TiNi alloys was studied by Fu et al. (2000). It was found that TiN nanocrystalline films are formed on the surface of PHEDP-modified TiNi alloys. Due to the plasma films, the size of the TiNi crystal grains was reduced, the adhesion of TiN nanocrystalline films on TiNi alloys was improved, and the corrosion resistance was greatly improved. Tan et al. (2003) used plasma implantation to improve the corrosion resistance of TiNi alloys. TiNi alloys were implanted by oxygen ions at 45 kV, and it was found that the sample had the best corrosion resistance at 21°C at an ion concentration of 1×10¹⁷ ions/cm^-2. Carbon plasma immersion ion implantation (C-PIII) was performed for 2 h to improve the corrosion resistance of TiNi alloys by Shanaghi and Chu (2018). Electrochemical impedance spectroscopy indicated that C-PIII improves the pitting or local corrosion resistance of TiNi alloys.

Corrosion of TiNi alloys in vivo causes the precipitation of Ni, which leads to allergies or carcinogenesis. Corrosion resistance is effectively enhanced by surface homogenization or surface ion implantation, which eliminates the unevenness or generates the reinforcing phase. The precipitation of Ni in vivo is effectively prevented by these methods.

2.2 Superlubrication (friction behaviors)

TiNi alloy biomedical products, such as artificial joints, are subjected to extreme loads during use. For example, a hip prosthesis can withstand 1–2 million cycles per year with articulating forces reaching three to five times the body weight (Brunette, 2001). Artificial joints produced abrasive particles under great friction and fatigue during use, resulting in the failure of the component. TiNi alloys have better mechanical properties such as friction and fatigue properties than traditional ceramic materials. Therefore, it is necessary to study the wear behaviors and other mechanical properties of TiNi alloys in detail.

TiNi alloy coatings have been used in recent years for joint replacement, dental shaping, and skeletal orthodontics due to their good biomedical prospects (Hussein et al., 2015). TiNi alloy coatings are increasingly used in artificial joints and other applications due to their low friction behaviors. Antiwear behavior is the main factor that controls and determines the stability and reliability of biomaterials in long-term clinical applications. In 2003, the abrasive wear of TiNi alloys was studied by Imbeni et al. (2003). As the grain size of the abrasive increases and the roundness factor decreases, the surface roughness of TiNi alloy article increases. The cracks of many medical devices are usually less than 25 μm. Thus, it is necessary to further study the grinding media and the grinding process of less than 20 μm.

It is also necessary to reduce the friction coefficient of TiNi alloy coatings to prevent the generation of abrasive grain particles. As for the tribological properties of TiNi alloys, many researches have been carried out in recent years. The tribological properties of TiNi alloys under liquid lubrications were systematically studied by Zeng (2013), Zeng and Dong (2014), and Zeng et al. (2016), and low coefficient of friction was achieved. In the case of castor oil, the ultra-low friction coefficient of 0.008 was obtained. The ultra-low friction mechanism is attributed to the OH-terminated surface generated by friction-induced castor oil degradation, and the contact surface of the friction pair formed a boundary lubricating film, which lowers the friction coefficient.

The friction and wear behaviors of TiNi alloy coatings in biomedical applications are far less than those of bulk TiNi alloys. Therefore, the friction behaviors of TiNi alloy coatings are studied in depth to improve the reliability and service life of TiNi alloy-coated products.

2.3 TiNi composites

In general, the single TiNi alloy has difficulty to meet the mechanical properties of parts in the biomedical field. Therefore, the mechanical properties of TiNi alloys are often improved by the method of preparing TiNi composites.

The corrosion behavior of TiNi alloys in Cl-rich solution was investigated by Hu et al. (2017). The reinforcing phase of TaC was added to TiB, TiB₂, and TiC reinforced Ti₂Ni/TiNi composite coatings, which greatly improves the corrosion resistance of the composite coatings in Cl-rich solution.

Fatigue refers to an important consideration in the work of TiNi alloy products. During the entire working life of medical products such as prostheses, TiNi alloys are subjected to the repeated cyclic loads. The fatigue performance of Ti coatings on Ti6Al4V alloy substrate under cold gas dynamic spraying (CGDS) was studied by Price et al. (2006). However, it is found that the fatigue properties of the substrate were reduced due to changes in the surface topography of the substrate and the participation in the tensile stress in the substrate caused by the CGDS process. Filip et al. (1999) prepared TiNi/Al₂O₃/Ti alloy composites and studied the structure and properties of TiNi/Al₂O₃/Ti composites to improve the fatigue properties of TiNi alloys. The fatigue life of this new composite material is greater than 10⁷ cycles; thus, the mechanical properties meet the application of bone tissue treatment.

The fatigue performance of TiNi alloys is improved by reducing friction between human tissue and the surface of the part. TiNi endodontic files (EFs) used for root canal treatment were coated by Adini et al. (2011) with cobalt films containing fullerene-like WS₂ nanoparticles, which can improve the fatigue resistance of the TiNi EFs.

TiNi alloys have been widely used in the field of medical devices because of its biocompatibility, corrosion resistance, superlubrication, shape memory effects, and so on. TiNi alloys have been used in the manufacture of minimally invasive implants, orthopedic instruments, dental instruments, and sutures.

Grunert et al. (2016) described the development of orbital floor implants made by TiNi alloys. The alloy can be used for the treatment of orbital fractures as shown in Figure 4, which are inserted in the eye socket through an open approach on the lower eyelid.

Figure 4:

Stages of the planned implant. A surface model (left) is derived from the computed tomography data, picturing the patient-specific geometry. With the help of a three-dimensional model (middle), the implant with the grid structure is made individually (Grunert et al., 2016). Copyright © Elsevier. License number 4555230145843.

3.1 Preparation of coatings

3.1.2 Laser plasma hybrid spray coating (LPHS)/vacuum plasma spraying (VPS)/low-pressure plasma spraying (LPPS)

Hiraga et al. (1999) studied the cavitation mechanism of TiNi alloy coatings by LPHS. The corrosion resistance of TiNi alloy coatings is improved because Ni reduces the phase transition temperature of austenite to martensite with the increase in Ni content. Then, the corrosion resistance of TiNi alloy coatings sprayed by VPS and laser plasma hybrid spraying (LPHS) was studied. The corrosion resistance of two TiNi alloy coatings sprayed by VPS and LPHS methods was 20 and 40 times, respectively, more than ordinary Ti coatings (Hiraga et al., 2001). The preparation of TiNi alloy coatings by VPS was further studied by Bram et al. (2002). TiNi alloy coatings were prepared using 50.8 at% of Ni alloy powder as shown in Figure 5. After heat treatment at 900°C for 2 h, the microstructures of TiNi alloy coatings became uniform, but pure Ni phase still existed. It is well known from biomedical investigations that the existence of pure Ni promotes the appearance of an allergy or leads to allergic reactions, such as the latent inflammation of the surrounding tissue if the allergy has already presented (Bram et al., 2002). Measurements of the wear resistance of TiNi alloy coatings in the as-sprayed state and after annealing are the subjects of further investigations.

Figure 5:

TiNi alloy coatings on a steel substrate produced by VPS (Bram et al., 2002). (A) The coating is porous and inhomogeneous. (B) The coating is less porous and exhibits a higher homogeneity. Copyright © Elsevier. License number 4555250429975.

The applications of LPPS in the preparation of TiNi alloy coatings were also studied (Lin et al., 2014a,b). Ni-coated Ti powders were used as raw materials for preparation, and the results showed that the cavitation resistance of TiNi alloy coatings is superior to 316L stainless steel.

3.1.5 Magnetron sputtering

Magnetron sputtering is commonly used to prepare TiNi alloy coatings. TiNi shape memory films were obtained at low temperature by DC magnetron sputtering by Cicek et al. (2015), as shown in Figure 6. Crystalline TiNi films were obtained by high sputtering current, low working pressure, and high substrate pulse (Figure 6A). TiNi and TiNiAg coatings were obtained through magnetron sputtering technique by Quandt and Zamponi (2008). The utilization of melt-cast sputtering targets followed by annealing in a rapid thermal annealing apparatus led to crystalline films. These TiNi films revealed superelastic properties at a temperature of 36°C.

Figure 6:

Surface and cross-section images of TiNi films: (A) crystalline films and (B) noncrystalline films (Cicek et al., 2015). Copyright © Elsevier. License number 4556390138610.

3.2 Surface modification of TiNi alloy coatings

TiNi alloy coatings directly prepared often have many disadvantages such as insufficient hardness or poor corrosion resistance. Therefore, surface modification is necessary to improve the performance of TiNi alloy coatings.

Richman et al. (1995) and Hiraga et al. (1999) produced TiNi alloy coatings by explosive spray and laser hybrid composite spraying (LPHS), respectively. The results showed that the martensite of TiNi alloys is beneficial to improve the corrosion resistance of TiNi alloy coatings.

The deposition of nanocrystalline TiN protective coatings on the surface of TiNi alloy coatings by DC magnetron sputtering was discussed by Kumar and Kaur (2009). The modified TiNi alloy coatings exhibit high hardness and modulus of elasticity, good wear resistance, and corrosion resistance. At the same time, applications of TiN/TiNi films in the electrochemical sensing of dopamine have been proven, which have important physiologic research prospects in the treatment of Parkinson’s disease.

Tillmann and Momeni (2014) deposited superelastic TiNi coatings on an Si (100) substrate by DC magnetron sputtering and applied the TiCN coatings to the austenitic TiNi alloy coatings to improve the corrosion and wear resistance. The original TiNi alloy coatings were transformed into a new generation of self-healing composite TiNi alloy-based films, which have great reference values for the biomedical applications of TiNi alloy coatings.

High-temperature vacuum annealing and friction stir processing were employed to modify cold-sprayed TiNi coatings obtained from a mechanically blended powder by Huang et al. (2018). The microhardness of the improved layer was increased to 1003.5±65.9 HV_0.1, which is about 4.5 times compared to the TiNi coating prepared by cold spraying (222.5±6.6 HV_0.1).

At present, TiNi alloy coatings have been widely used in the medical field to reduce cost. The microgripper was fabricated by Huang et al. (2003) and Fu et al. (2001) and consisted of two cantilever beams made of Si and coated with TiNi alloy coatings as shown in Figure 7. The manufacturing cost of microgripper is greatly reduced because TiNi alloy coatings are deposited on the top of the Si wafer by sputtering.

Figure 7:

(A) Assembled medium-sized microgripper and (B) cross-section (Fu et al., 2001). Copyright © Elsevier. License number 4555650080131.

Thin-film TiNi shape memory alloy (SMA) microactuator was prepared by Gill et al. (2001) through magnetron sputtering and wet etching. The purpose of the microactuator is to grab microsize objects such as cancerous tumor for removal from the body. As shown in Figure 8, TiNi alloy thin films were deposited by magnetron sputtering. These devices are successfully employed to pattern cylindrical TiNi alloy thin films used for the stent.

Figure 8:

Fabricated microwrapper with TiNi alloys (Gill et al., 2001). Copyright © Elsevier. License number 4556940290099.

5.1 Summary

1. Precipitation of Ni is essential to weaken the biomedical properties of TiNi alloys. Surface modification is an effective method to prohibit the precipitation of Ni and improve the biomedical properties of bulk TiNi alloys. Corrosion resistance is enhanced by eliminating the unevenness or generating the reinforcing phase through surface homogenization or surface ion implantation. The TiO₂ layer is the main factor that affects the biocompatibility of TiNi alloys. Increasing the thickness and density of the TiO₂ layer on the surface of TiNi alloys is an important method to improve the biocompatibility.

2. TiNi alloys are widely used in biomedical applications, but there are difficult-to-process metal materials due to their viscosity and pseudo-elasticity during processing, resulting in an increase in the cost of using. Therefore, the manufacturing process of TiNi alloy coatings has been gradually studied. Thermal spraying technology has shortcomings, but the ductility and memory of TiNi alloys are restored by vacuum annealing treatment. The strength and corrosion resistance of TiNi alloy coatings are enhanced by the decrease in the transformation temperature of austenite to martensite and increase in the proportion of martensite. Currently, TiNi alloy coatings prepared by LPHS have excellent corrosion resistance.

5.2 Remarks

With the development of surface modification technology, the biomedical properties of TiNi alloys are continuously improved. However, the application of bulk TiNi alloys is limited by high processing cost. Researches on the preparations of TiNi alloy coatings not only help to reduce the manufacturing cost of TiNi medical products but also contribute to process parts with more complex shapes. In addition, advanced manufacturing technologies such as additive manufacturing or powder metallurgy are also used to manufacture TiNi alloys. Therefore, further research of the biomedical applications of TiNi alloys is needed.

1. The oxidation mechanism of TiNi alloys is further investigated systematically, and a new process of surface homogenization is proposed to improve the corrosion resistance of TiNi alloys. Improving the thickness of the oxide layer on the surface of TiNi alloys is studied in detail to enhance the lifetime and reliability of TiNi alloy medical products.

2. Reducing the grain size and surface roughness of TiNi alloy coatings is studied to improve blood compatibility and hydrophilicity and reduce thrombus formation.

3. The phase transition mechanism of TiNi alloys is studied in detail to improve the quality of TiNi alloy products made by additive manufacturing technology.