17.1 Introduction

Titanium and its alloys are biologically inert materials widely used in the production of biomedical implants due to their high tensile and fatigue strength, low Young modulus, and superior corrosion and wear resistance.¹ Moreover, they are biocompatible, providing direct bonding with bone surface in orthopedic and dental surgery, and are inert in physiological environments.²

Titanium is a material with a high superficial energy and after implantation it provides a favorable body reaction that leads to direct deposition of minerals on the bone-titanium interface and titanium osseointegration.³ This is due mainly to the thin TiO₂ film naturally formed on its surface, which is very stable and chemically inert. Nevertheless, bulk titanium used in implants presents three main problems⁴:

– high processing energy due to melting and casting difficulties

– slow reaction with the human tissues resulting in difficulties in bone attachment

– higher elastic modulus compared to bone

One of the ways to reduce the problems associated with stress generated by the incompatibility between the elastic modulus of Ti implant and that of the bone is to fabricate titanium implants with controlled porosity.⁵ Recently, porous Ti scaffolds with good mechanical properties (particularly, elastic modulus and stiffness) that provide a favorable environment for bone growth and a significant enhancement in the osteoblastic activity were produced.⁶

Although “commercially pure” titanium has acceptable mechanical properties and has been used for different implants, presently, for most biomedical applications, titanium is alloyed with nickel or with small amounts of aluminum, vanadium, or niobium. These elements improve some of the Ti characteristics such as strength, mechanical properties, corrosion resistance, and so on. Titanium alloys most frequently used in dental and orthopedic implantations are NiTi (Nitinol), Ti-6Al-4V, and Ti-6Al-7Nb.

NiTi is a material with excellent shape memory effect and super elasticity and damping capacity suitable for use in orthopedic implants. In dentistry, the material is used in orthodontics for brackets and wires connecting the teeth. However, it should be mentioned that its high Ni content constitutes a potential threat to its safe use in vivo.⁷

Ti-6Al-4V is the material used for ~ 20-30% of commercially available medical devices. It contains 6% Al and 4% V by weight and is by far the most common Ti alloy. As in the case of NiTi, during the Ti-6Al-4V use for implants the release of metallic ions in the physiological environment could be harmful to the health. Thus, Al and V may cause neurological disorders and bone diseases when they are released in the human body.⁸

Ti-6Al-7Nb is another titanium alloy with excellent strength and biocompatibility dedicated for surgical implants, used especially for replacement hip joints. It is a typical α + β alloy, with Al as stabilizer for the α phase and Nb as stabilizer for the β phase.² It is usually stable in neutral halide solutions, but in acidic media is prone to accelerated dissolution.⁹

Irrespective of the material nature, the most important requirement for the long-term success of implants is the stable interface between the biomaterial and the surrounding tissue.¹⁰ Hence, an improvement of all Ti-based materials corrosion resistance is of particular importance. In this regard, TiO₂ naturally formed on the surface is one of the main candidates for protective coatings on Ti and Ti alloys substrates due to its thermodynamic stability, chemical inertia, and low solubility in body fluids.¹¹ Titania with specific structures of anatase and rutile was found to stimulate apatite formation¹² and were proved to be much more beneficial for bone growth than the amorphous TiO₂.¹³ However, the native TiO₂ layer is very thin, irregular, and porous and there is also the risk of viruses being carried by this layer. Moreover, the TiO₂ naturally formed on Ti surface is not sufficient to ensure a strong chemical bond with bone tissue and a safe, long lifetime of the implants. Consequently efforts are made for thickening the oxide layer along with an improvement of its surface ability to promote osteosynthesis and to increase its corrosion and wear resistance.

The development of nanostructures has allowed unprecedented access to an intimate interaction with the biological environment. For bone regeneration and repair, re-creating substrates on a nanoscale allows scientists to probe the first level of the bone structural hierarchy.¹⁴ Studies have shown that TiO₂ films consisting of nanostructures (nanotubes, nanorods, nanosheets, etc.) exhibit high hydrophilicity, improving thus the bioactivity and bone-bonding behavior of materials for implants.^15,16 However, there is still a lot to do in the direction of improving the behavior of implant biomaterials. Some of the most widely used methods to modify the characteristics of materials surface are presented below.

17.2 Surface Modification Methods

As already mentioned, optimization of the biomaterial surface is necessary in order to trigger osseointegration, to minimize metallic ion release, and to ensure a longer lifetime of an implant. There are several possible ways to form a surface layer on the implant materials¹⁷:

(i) materials deposition on the substrate (e.g., by sol-gel method, spray method, or electrolytic deposition)

(ii) material removal from the substrate (e.g., by etching)

(iii) substrate surface transformation (e.g., by anodic oxidation)

Some of the most used methods to modify the surface of Ti and Ti alloys are presented below.

17.3 Sol-Gel Method

Sol-gel process is a method for producing solid materials from small molecules that is suitable for preparing different coatings (e.g., silicium and titanium oxides) on the surface of Ti-based materials. It involves conversion of small molecules (precursors) into a colloidal solution (sol) and then into an integrated network (gel) consisting of either discrete particles or network polymers.

The main advantages of the sol-gel method are¹⁸: (1) low processing temperature (avoiding volatilization of entrapped species), (2) the possibility to cast coatings in complex shapes, and (3) the use of compounds that do not introduce impurities into the end product. Among the disadvantages of this method one can count the relatively long period of time for processing flow and the difficulties related to phase separation occurring especially in hybrid coating synthesis.

The sol-gel process involves four stages: (1) hydrolysis, (2) condensation/polymerization of monomers, (3) growth of particles, and (4) gel formation.¹⁸ These processes are influenced by several experimental parameters such as pH, temperature, concentration of the reactants, and presence of additives.

Traditional precursors for sol-gel coatings are alkoxysilanes such as tetraethyl orthosilicate (TEOS) and tetramethyl orthosilicate (TMOS), but efforts are made in order to find less toxic and more environmental friendly precursors.

TiO₂ can be prepared, for example, starting from precursor solutions containing an alkoxide (e.g., tetrabutylorthotitanate) and diethanolamine dissolved in ethanol¹⁶ and mixing them with water (with or without additives) in a certain ratio.

SiO₂ coatings can be prepared starting from a precursor solution consisting of tetraethylorthosilicate Si(OC₂H₅)₄, H₂O, C₂H₅OH, and HCl, mixed at a certain molar ratio.¹⁹ Different compounds (inhibitors, pigments, etc.) can be added in order to improve the physicochemical properties of the coating.

Irrespective of the nature of the film, the two main techniques used to apply a sol-gel coating on the surface of a metallic substrate are dip-coating and spin-coating.

17.3.1 Dip-coating

By this technique, the material from which the film is produced is put into solution, and then the substrate is progressively dipped into and is extracted from the solution at a controlled rate (Figure 17.1). After the solvent evaporates, a thin and homogeneous film is produced. The thickness of deposited liquid film coatings depends on the coating solution properties such as density, viscosity, and surface tension, as well as surface withdrawal speed from the coating solution. The thickness of the film is generally bigger than that prepared by spin-coating with the same solutions.

Dip-coating was successfully used, for example, to prepare sol-gel-derived Al₂O₃ films on γ-TiAl-based alloys,²¹ porous TiO₂ films,¹⁶ hydroxyapatite (HA) coatings,²² and SrO-SiO₂-TiO₂ on NiTi,²³ and so on.

17.3.2 Spin-coating

In the case of spin-coating, an amount of solution is placed on the substrate that is rotated at high speed in order to spread the fluid by centrifugal force (Figure 17.2). After the evaporation of the solvent, a thin, homogeneous film is formed. As in the case of dip-coating, final film thickness and other properties will depend on the nature of the sol-gel coating (viscosity, drying rate, surface tension, etc.) and on the parameters chosen for the spin process. Higher spin speeds and longer spin times give birth to thinner films.²⁵ Generally, a moderate spinning speed is recommended. The drying rate should also be slow in order to ensure film uniformity. A supplementary drying step is sometimes necessary after the high-speed spin step to further dry the film without substantially thinning it.

By using spin-coating, SiO₂ films were deposited on Ti-48Al alloy,¹⁹ sphene (CaTiSiO₃) ceramics on Ti-6Al-4V²⁶ alkoxide-based HA nanocoatings on Ti and Ti-6Al-4V substrates,²⁷ HA²⁸ and HA/polymer coatings on Ti,²⁹ TiO₂ on Ti-6Al-4V,³⁰ and many others.

17.4 Laser Oxidation

Laser surface treatment is a thermal process that has an advantage over conventional furnace treatment. It is based on the heating caused by the light adsorption of the surface layer and the cooling ensured by the high conductivity of the material.³¹ The adsorbed laser energy results in a thin surface layer with desired properties while the bulk of the material is unaffected.

By using laser treatments followed by AFM investigations, it was proved that the morphological modifications of Ti-based alloys play a key role in cell adherence and proliferation by influencing cell-material interactions.³²

Porous Ti samples were fabricated by using laser power in an attempt to improve the osteoconductivity and reduce the problems associated with stress shielding.⁵

NiTi were treated with a pulsed Nd:YAG laser in air, by using selected parameters aiming to obtain uniform oxide films of controlled thickness.⁷ The corrosion resistance of the so-treated NiTi samples increased about 15 times, while the surface Ni/Ti ratio was reduced. The last feature is important as the high Ni content in NiTi constitutes a potential threat to its safe use in vivo.

17.5 Anodic Oxidation

Anodic oxidation is a traditional surface modification method of Ti and its alloys that allows the obtaining of the desired properties of oxide layer by controlling the electrochemical parameters, such as electrolyte composition and concentration, applied potential or current, temperature, and so on.

Anodic oxidation can be performed potentiostatically³³ or galvanostatically³⁴ in different electrolytes. The most common electrolyte solutions used for anodization are H₂SO₄,^12,33 Na₂SO₄,³³ CH₃COOH,^33,35,36 H₃PO₄,³³ and HF.^13,14 Mixture of inorganic substances or hybrid inorganic-organic solutions could be also used.^2,37,38 It was reported that porous anodic films consisting of and/or anatase are produced by anodizing in H₂SO₄ and Na₂SO₄ solutions, while amorphous TiO₂ films are produced in CH₃COOH and H₃PO₄.³³ Self-organized TiO₂ nanotubes can be generated from viscous electrolytes consisting of glycerol or ethylene glycol with addition of ammonium fluoride.³⁸ Viscosity of the electrolyte plays a key role in the nanotubes growth process.

The oxidation potentials are usually determined before the anodic oxidation step from potentiodynamic polarization experiments.³⁵

Oxide layers formed by anodization differ according to the substrate, for example, oxidation in chromic acid produced a TiO₂ layer on titanium substrate, but layers consisting of TiO₂ and Al₂O₃ oxides on a Ti-6Al-4V substrate.³⁹ The minor oxides formed by anodic oxidation in the TiO₂ layer on the Ti-6Al-7Nb alloy are Al₂O₃ and Nb₂O₅.³⁵

In order to prepare layers with high apatite-forming ability, anodic oxidation can be carried out in complex solutions containing calcium and/or phosphate ions.⁴⁰

17.6 Plasma Electrolytic Oxidation

Plasma electrolytic oxidation is a relatively inexpensive and simple method used for increasing surface roughness of the materials, allowing also the synthesis of ceramiclike oxide films on some metals.⁴¹ It is similar to anodizing, but it employs higher potentials so that the resulting plasma modifies the structure of the oxide layer. Thus, the method combines electrochemical oxidation with a spark treatment in the plating bath. If this contains aluminate, phosphate, silicate, and sulfate anions or some of their combinations, bioactive layers enriched in these elements are formed.⁴² The incorporation of Ca and P into an oxide film formed via anodic oxidation on Ti-15Mo alloy⁴¹ or Ti-6Al-7Nb,⁴³ as well as the formation of a bioactive film on Ti-6Al-4V⁸ after the application of a voltage higher than the breakdown voltage of the oxide layer, were also reported. Plasma electrolytic oxide coatings are generally recognized for high hardness, wear, and corrosion resistance.

17.7 Electrolytic Deposition

A promising route for deposition of thin oxide films from aqueous solutions is the electrolytic deposition method using TiCl₄ or TiOSO₄ as starting materials in the presence of hydrogen peroxide.⁴⁴ The procedure consists in galvanostatic electrodeposition at 0 °C from a solution obtained by slow addition of TiCl₄ and H₂O₂ and subsequent adjustment of the solution volume. The resulting films are uniform, of high purity, present excellent adhesion to the substrate, and have good mechanical properties. The electrolytic deposition can be followed by annealing.

A possible mechanism for TiO₂ electrodeposition is based on hydrolysis of the peroxo-complex of Ti in the presence of OH⁻ ions generated at the cathode.⁴⁵

${\text{TiCl}}_4\to {\text{Ti}}^{4+}+4{\text{Cl}}^{-}$

${\text{Ti}}^{4+}+{\text{H}}_2{\text{O}}_2+\left(n-2\right){\text{H}}_2\text{O}\to {\left[\text{Ti}\left({\text{O}}_2\right){\left(\text{OH}\right)}_{n-2}\right]}^{\left(4-n\right)+}+n{\text{H}}^{+}$

${\left[\text{Ti}\left({\text{O}}_2\right){\left(\text{OH}\right)}_{n-2}\right]}^{\left(4-n\right)+}+m{\text{OH}}^{-}+k{\text{H}}_2\text{O}\to {\text{TiO}}_3{\left({\text{H}}_2\text{O}\right)}_x$

$2{\text{TiO}}_3{\left({\text{H}}_2\text{O}\right)}_x\to 2{\text{TiO}}_2+{\text{O}}_2+2x{\text{H}}_2\text{O}$

A fluorine-doped HA/ZrO₂ two-layer coating was prepared by electrodeposition on titanium in ZrO(NO₃)₂ aqueous solution and subsequently in a mixed solution of Ca(NO₃)₂, NH₄H₂PO₄, and NaF. The adherence of the two-layer coating was found to be significantly higher than that of the pure HA coating.⁴⁶

17.8 Combined Methods

Sometimes, the combination of two or more methods is used in order to improve surface layer properties and to attain the desired characteristics. Thus, the combined laser/sol-gel synthesis of calcium silicate coating was produced on Ti-6Al-4V substrates for improved cell integration.⁴⁷

Anodic oxidation treatment applied on laser-processed porous Ti samples was used to produce bioactive TiO₂ nanotube arrays that significantly changed the surface properties and enhanced the apatite-forming ability of porous TiO₂ samples in simulating body fluid.⁵ By creating the nanotube arrays, the healing time due to increased biocompatibility can be significantly reduced.

The polymeric sponge replication method followed by microarc oxidation (MAO) was used to prepare HA/TiO₂ hybrid coatings on highly porous Ti scaffolds.⁶ For this purpose, a polyurethane sponge was covered with titanium hydride slurry (consisting of TiH₂ dispersed in ethanol containing triethyl phosphate and polyvinylbutyl), and then heated at 800 °C to convert TiH₂ to Ti. The surface was then coated with a hybrid, bioactive, microporous HA/TiO₂ layer using MAO.

17.9 Protective Films

The protective layers produced on the surface of Ti and Ti-based alloys for enhancing their corrosion resistance, as well as their biocompatibility, are of different natures: oxides, composites, HA, hybrid organic-anorganic, ceramic, and so on.

17.9.1 Oxides

Oxide-based films are the most frequently produced on the Ti substrates. The most important factors deciding their corrosion resistance are the oxide thickness, and its evenness and compactness. The main oxide formed on the surface of Ti alloys is TiO₂, but minor oxides of the alloying elements (Al₂O₃, Nb₂O₅, NiO, etc.) are also produced.

TiO₂ can be produced in different crystallographic forms (anatase, rutile, brookite; see Figure 17.3), depending on the substrate nature and the preparation conditions.

From the point of view of osseointegration abilities, crystalline forms are superior to amorphous ones and nanostructured layers are superior to microstructured ones. For example, rutile can induce apatite formation, especially when it has a crystal orientation in the (101) plane, because the matching structure with apatite (002) promotes the nuclei formation for epitaxial crystal growth. On the contrary, amorphous titania layers are not able to induce apatite formation.³³

Osteoblasts are very sensitive to the different materials’ topographies and respond differently on submicron- and nanometer-sized surfaces. This is why nanostructured TiO₂ are preferred and are often generated on the top of the thin TiO₂ naturally formed on the Ti surface, aiming to accelerate osteogenesis. Besides, the systems based on nanotubes are very hydrophilic and promote greater cell adhesion and increased channeling for fluid exchange that contributes to bone remodeling.¹⁴

The adhesion/propagation of the osteoblasts is substantially improved on vertically aligned and laterally spaced TiO₂ nanotubes.¹⁵ This type of nanotube can be obtained by anodic oxidation of Ti in diluted HF acid^13,14 and confer to the substrate increased surface area, porous structure, and superior osteoblast cell growth.

Anodization of the same material in different electrolytes leads to the formation of different oxidic nanostructures. For example, oxide layers obtained by anodic oxidation of Ti₆Al₇Nb alloy in (NH₄)₂SO₄ + NH₄F solution are more homogeneous than those obtained in a solution containing glycerol and NH₄F and water.² The concentration of F⁻ in the electrolytes ions is also important: the larger its concentration, the higher the diameters of pores and the surface roughness. It should be also noticed that, as the electrolytes used for anodization are fluoride-based, the nanotubes always contain a certain amount of incorporated fluorine,⁴⁸ which could be useful for the enhancement of antibacterial properties of the surface.

17.10 Corrosion Studies

The most important requirement for a biomaterial is its compatibility with the human body. The bioimplants should not cause adverse effects such as allergies or inflammations and should not be toxic. Other important requirements for a successful implant material are very high mechanical, corrosion, and wear resistance in order to ensure a long service period (over 15 years).

In the physiological environment, biomedical implants are subjected to corrosion, especially to mechanically accelerated electrochemical processes such as stress corrosion, corrosion fatigue, and fretting corrosion. Fretting corrosion, alone or in combination with crevice corrosion, has been identified as one of the most important modes of corrosion of Ti-based implants such as hip, knee, and shoulder replacements.^9,68

Due to corrosion processes, even the most resistant materials undergo electrochemical exchange and release metal ions in the physiological environment, producing different inconvenient or even implant failure.⁶⁹ For Ti-based alloys, mostly Ti is released, but also Al and V were detected in the case of Ti-Al-V alloy.⁹ It is accepted that the tolerable corrosion rate for metallic implants should be about 2.5 × 10^− 4 mm/year.⁷⁰ In this context, the failure due to corrosion remains one of the challenging clinical problems and all the implantable materials should be tested from a corrosion behavior point of view, apart from other properties. Moreover, different approaches should be used in order to enhance the corrosion resistance.

A way to stabilize the passive film formed on Ti-based surfaces is Ti alloying with selected elements that stabilize either the α-phase (Al) or the β-phase (Mo, V, Fe) of the alloy and enlarge its passive range.^71,72 Other methods consist, as already mentioned, in modifying the surface of the implant materials in a convenient way.

In order to investigate the corrosion behavior of Ti-based biomaterials, different artificial physiological solutions can be used simulating plasmatic serum (Ringer’s and Hank’s solution), saliva, or sweat. The composition of some of these artificial physiological solutions is presented in Tables 17.2–17.4.

The in vitro evaluation of the corrosion behavior is an important step in the elaboration of new biomaterials. The corrosion resistance of biomaterials used for implants is usually tested by electrochemical methods because they are accelerated tests that allow evaluating rapidly the performance of materials and predicting their long-time behavior. The most common electrochemical methods are open circuit potential measurements, cyclic voltammetry, potentiodynamic polarization measurements, and electrochemical impedance spectroscopy.

By comparing three titanium-based materials (Ti, NiTi, and Ti-6Al-7Nb), it was reported that Ti-6Al-7Nb alloy possesses the greatest corrosion resistance in simulated Hanks physiological solution at pH 7.4 and 37 °C, this behavior being valid for the untreated samples as well as for the samples anodically oxidized at 3.0V in acetic acid.³⁵ The polarization resistance values, determined from EIS measurements for the oxidized Ti-6Al-7Nb alloy during the later stages of immersion in Hank’s solution, are one order of magnitude higher than those determined for the untreated alloy. Thus, the anodized samples possess a much higher corrosion resistance than the untreated ones.

A two-layer model was proposed to accurately describe the corrosion behavior of the oxidized Ti alloys. Thus, it is generally accepted that the surface film consists of a two-layer oxide, composed of a dense inner layer and a porous outer layer⁷⁶ (Figure 17.4).

The high corrosion resistance is due to the thin barrier-type inner layer, while its osseointegration ability is attributed to the porous outer layer. The significant enhancement of the polarization resistance values at longer immersion time in the physiological solution could be due to the stabilization of the barrier layer (e.g., by some sealing processes inside pores of the oxide film⁷⁷), which might diminish the tendency of the alloy to corrosion. The nature of the porous layer depends on the nature of the alloy and the anions present in the solution.⁷⁸ The two-layer model accurately describes also the electrochemical behavior of HA coatings on Ti-6Al-4V in Ringer’s physiological solution.⁷⁹

17.11 Conclusions

Surface modification is one of the ways to significantly improve the corrosion and wear resistance of Ti-based materials along with their surface texture and biocompatibility. Protective layers of different natures (oxides, HA, composites, hybrid layers, etc.) can be generated on the materials’ surface by various methods such as electrooxidation, laser oxidation, and sol-gel spin coating technique.

The performance of biomaterials used for implants should be evaluated by determining their physical, chemical, and biological properties. As the materials behavior in vivo is a complex one, the understanding of the implant materials behavior requires joint efforts of materials science, medicine, biology, and engineering. The in vitro experiments should be considered only as screening tests.