surface nano-architectures and their effects on the

Surface nano-architectures and their effects on the mechanical properties andcorrosion behavior of Ti-based orthopedic implants

Shuilin Wu a,b, Xiangmei Liu b, Kelvin W.K. Yeung c, Huan Guo b, Penghui Li a, Tao Hu a,C.Y. Chung a, Paul K. Chu a,⁎a Department of Physics & Materials Science, City University of Hong Kong, Tat Chee Avenue, Kowloon, Hong Kong, Chinab Ministry-of-Education Key Laboratory for the Green Preparation and Application of Functional Materials, Faculty of Materials Science and Engineering, Hubei University, Wuhan 430062,PR Chinac Division of Spine Surgery, Department of Orthopaedics & Traumatology, Li Ka Shing Faculty of Medicine, The University of Hong Kong, Pokfulam, Hong Kong, China

a b s t r a c ta r t i c l e i n f o

Available online 22 October 2012

Keywords:Ti-based alloysSurface modificationMechanical propertiesCorrosion resistanceOrthopedic implantsNano surface architecture

Ti-based alloys are widely used in orthopedic implants because of their excellent mechanical properties andbiocompatibility. However, in the complex physiological environment and influenced by other factors likecyclic loading, these implants can lose their natural properties and become incompatible or cyto-toxic tothe body tissues and cells. Two of the causes are the deteriorated mechanical properties in this dynamicenvironment and the effects of by-products stemming from corrosion. In serious cases, implant failure re-sults. Nano architectures on the surface of these alloys can improve the in vivo and in vitro biocompatibilityof these alloys. This paper reviews recent progress pertaining to the design and construction of nano surfacearchitectures on Ti-based alloys and their effects on the mechanical properties and corrosion resistance in thesimulated physiological environment.

© 2012 Elsevier B.V. All rights reserved.

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 132. Design and preparation of nano architectures on Ti-based alloys . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 14

2.1. Surface nano films/layers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 152.1.1. Plasma immersion ion implantation and deposition (PIII&D) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 152.1.2. Oxidation. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 172.1.3. Hydrothermal processes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 192.1.4. Plasma spraying . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 192.1.5. Electropolishing. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 202.1.6. Magnetron sputtering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 212.1.7. Sol–gel . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 21

2.2. Surface nanocrystallization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 212.3. One-dimensional nanostructures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22

2.3.1. Nanotubes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 222.3.2. Nanowires or nanofibers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22

3. Surface nanostructures on porous Ti-based alloys. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 234. Conclusion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 24Acknowledgments. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 24References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 24

1. Introduction

Among common biomaterials which include bioceramics, biopoly-mers, and biometals, metals are most widely used in hard tissue re-pair or reconstruction implants due to their excellent mechanical

Surface & Coatings Technology 233 (2013) 13–26

⁎ Corresponding author. Tel.: +852 34427724; fax: +852 34420542.E-mail address: [email protected] (P.K. Chu).

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properties [1,2]. For instance, biomedical stainless steels (316 and316L), Co-based alloys, Ti-based alloys, and tantalum have foundcommercial success in orthopedics and dentistry [3]. In addition totheir typically better biocompatibility in comparison with commonceramics and polymers, the high strength and excellent ductility ofmetals ensure high resistance to fracture thus favoring long-termload bearing. The easy processability of metals by casting and conven-tional machining promotes easy production of biomedical implantswith a complex shape [3–5].

Potential bone substitutes should possess good biocompatibilityand matching mechanical properties with natural bones. In the caseof biometal implants, some by-products arising from corrosion andwear can degrade the biocompatibility of implants in the complexphysiological environment. For example, harmful metal ions leachedfrom the implants and wear debris may induce inflammation, cellapoptosis, and other detrimental tissue reactions. For example, triva-lent Cr ions are released from CoCr alloys [6]. 316L stainless steel iscommonly used in temporary osteosynthetic implant devices suchas screws and hip nails but suffers from crevice corrosion and wearcaused by the high stress and oxygen-depletion under load-bearing[7–9]. It is well accepted that that materials' stiffness close to that ofnatural bones can provide good load transfer so that new bone forma-tion can be accomplished effectively. Unfortunately, the typicallylarger Young's moduli of common metals give rise to stress shieldingand impede the growth of new bone tissues. As shown in Table 1[3,4,10–12], among the common biometals, Ti-based alloys haveYoung's moduli most similar to those of nature bone. In combinationwith their higher corrosion resistance and other excellent mechanicalproperties, Ti-based alloys are generally regarded to be the mostpromising implant materials at present.

It is well known that Ti-based alloys are bioinert and their biocom-patibility is related to the positive response of cells and tissues to thematerials' surface. Therefore, the surface biological characteristics ofTi-based implants play an essential role in bone tissue repair andcan be selectively and optimally altered by surface modification tech-niques. However, while some biological properties can be enhanced,improper surface treatment may degrade the corrosion resistanceand mechanical properties from the perspective of bone implants.Release of toxic metal ions like Ni, V, Nb, and Al from some Ti-basedalloys may cause the concentrations of these elements to exceed thesafe limitations in adjacent body tissues and fluids [13,14]. In this re-spect, surface nano-structuring has been demonstrated to be effective

in not only enhancing the materials' biocompatibility but also im-proving the corrosion resistance and mechanical properties [15–22].In this review, we aim at summarizing recent progress pertaining tothe development of surface nano-architectures and discussing theireffects on the mechanical properties and corrosion resistance ofTi-based alloys.

2. Design and preparation of nano architectures on Ti-based alloys

Many surface modification techniques have been developed to ob-tain various nano structures on the surface of biomedical materials.

Table 1Mechanical properties of metallic biomaterials, some polymers and ceramics, as well ascortical bone properties for comparison [3,4,10–12].

Materials Young's modulus(GPa)

Yieldstrength(MPa)

Ultimatestrength(MPa)

Bone 1–2 (trabecular)10–27(compact)

10–90(trabecular)

120–200

Ti-basedalloys

CP-Ti 100–115 170–480 240–550Ti6Al4V 110 860 930Ti6Al7Nb 105 795 860Ti5Al2.5Fe 110 820 900Ti13Nb13Zr 79–84 863–908 973–1037Ti12Mo6Zr2Fe 74–85 1000–1060 1060–1100Ti35.5Nb7.3Zr5.7Ta 55–66 793 827Ni–Ti 80 70–140 700–1100

Fe-based alloys 200–205 170–690 540–1000Co-based alloys 220–230 450–1500 655–1900316L stainless steel 193 205–380 515–585304L stainless steel 193–200 215 505Ta 188 140–345 205–480UHMWPE 0.5 – ~3Al2O3 350–380 – 400Plasma-sprayed-ZrO2 200 – 100–300

Fig. 1. SEM micrographs of titanium: (a) before implantation; (b) after Ca ion implan-tation at 10 kV; and (c) SIMS depth profiles of sample in (b) [39].

14 S. Wu et al. / Surface & Coatings Technology 233 (2013) 13–26

The typical methods include lithographic techniques [23,24], imprint-ing [25], femtosecond laser-texturing [26], shot peening [27], laserdeposition [28], plasma spraying [29], gaseous plasma treatment[30], ion bombardment [31], chemical etching [32], electrochemical

process [33], physical vapor deposition [34], self-assembly [35], andso on. Different nano architectures can be produced to cater to partic-ular applications. With regard to Ti-based alloys, in order to accom-plish bone integration, it is necessary to have good bone inductioncapability. Meanwhile, good mechanical properties and higher corro-sion resistance are also required. Therefore, a proper surface modifi-cation technique should not only enhance specific surface propertiesbut also preserve the desirable bulk attributes. The nano architectureson Ti-based alloys can be classified into the following categories.

2.1. Surface nano films/layers

Owing to the large amount of Ti in Ti-based alloys, a thin nano filmor layer composed of titanium oxides, nitrides or titanates is oftenprepared on the surface of these alloys.

2.1.1. Plasma immersion ion implantation and deposition (PIII&D)Plasma immersion ion implantation and deposition (PIII&D) was

first introduced by Conrad et al. and Tendys et al. [36,37]. This tech-nique offers many advantages over conventional surface modifica-tion methods because of its non-line-of-sight nature. It is oftenperformed to modify the surface of metals such as stainless steels,tool steels, pure aluminum, and titanium-based alloys to improvemechanical properties such as hardness, friction coefficients, wearresistance, and corrosion resistance. Several ion sources are usedin this course.

2.1.1.1. Metal PIII&D. Calcium PIII&D has been shown to enhance thebiocompatibility of titanium alloys [38–40] both in vitro and in vivo.Liu et al. [39] have used Ca-PIII to modify the surface of titanium.After exposure to air, a thin layer composed of calcium hydroxide

Fig. 2. Potentiodynamic polarization curves of Na-PIII, Na and O PIII, and untreated NiTi[42].

Fig. 3. SEM images after electrochemical tests (25×): (a) control and (b) 20 kV O PIIIand annealed at 600 °C [17].

Fig. 4. Nano-hardness and Young's modulus profiles acquired from the NiTi alloys:(a) untreated NiTi and (b) C2H2 PIII&D NiTi [16].

15S. Wu et al. / Surface & Coatings Technology 233 (2013) 13–26

and calcium carbonate is formed on the surface (Fig. 1). The depthprofile in Fig. 1c indicates that this layer is composed of a 500 nmthick deposited layer and an implanted layer 700 nm in thickness. So-dium PIII&D has also been performed by Maitz et al. [41] to enhancethe biocompatibility of titanium. In comparison with the untreatedsample, Na PIII&D increases the breakdown potential (Eb) of theNiTi alloy from 300 mV to 463 mV (shown in Fig. 2) [42], indicatingthat metal ion PIII&D is capable of improving the corrosion resistanceof NiTi alloys.

2.1.1.2. Gas PIII&D. Many gaseous ions with the exception of noble gas-ses react with titanium to form stable titanium compounds during PIIIand nitrogen, hydrogen, oxygen, water vapor, and so on have beenused in PIII&D to modify Ti-based alloys. Poon et al. [17] have shownthat oxygen PIII followed by annealing significantly improves the cor-rosion resistance of NiTi alloys. As shown in Fig. 3, there are many cor-roded holes on the untreated sample but only a few small pits can befound on the plasma-modified sample. It can be ascribed to the

formation of the 150 nm thick titanium oxide surface barrier layer.This film created by PIII&D is more compact and uniform than the na-tive oxide layer thereby offering better protection [17]. Using C2H2

PIII&D, a carbon-containing coating about 160 nm thick is producedon the surface of NiTi substrate. The hard titanium carbide formed atthe interface between the carbon coating and NiTi substrate enhancesthe surface mechanical properties such as Young's modulus and hard-ness (Fig. 4) [16]. The same effects can be achieved by nitrogen PIII[19,22,43,44]. Besides the enhanced Young's modulus and hardness[19], the 100 nm thick TiN film fabricated by N PIII improves thewear resistance of the NiTi alloys (Fig. 5) [22]. In addition, N PIII doesnot compromise the super-elasticity of NiTi as long as the sample tem-perature is not excessively high by adjusting the implantation parame-ters during the experiments [22]. Similar to N2, C2H2 and O2, H2O canalso lead to the formation of a thin titanium oxide layer with anano-structured surface as shown by the AFM image (Fig. 6). ThisH2O PIII&D surface improves the corrosion resistance of the NiTi alloyappreciably [45].

2.1.1.3. Dual PIII&D. Double-element plasma immersion ion implanta-tion and deposition can endow Ti-based alloys with excellent com-prehensive performance that cannot be achieved by single elementPIII&D [15,42,46–48]. Wen et al. [15] prepared titanium oxide and ti-tanium nitride gradient films on Ti6Al4V by PIII&D. The Ti ions creat-ed by a Ti cathodic arc source diffuse into the vacuum chamber andare subsequently implanted into the negatively biased sample. Atthe same time, a nitrogen or oxygen gas plasma is sustained in thevacuum chamber by an external radio frequency plasma source torealize nitrogen or oxygen ion implantation [15]. The depth profile re-veals that the thickness of films is in the range of 510–940 nm. Thehardness of the gradient film is better than that of a single-elementfilm and the nano-hardness of the gradient film can be as high as19.5 GPa, resulting in much improved wear resistance and adhesion[15]. As shown in Fig. 2, PIII&D of sodium and oxygen works betterthan sodium PIII&D alone in improving the Eb of NiTi alloys in

Fig. 5. Variations in the wear rate on the NiTi samples under different loads [22].

Fig. 6. Surface morphologies of NiTi by AFM: (a) untreated and (b) H2O PIII [45].


simulated body fluids (SBF). Specifically, the Eb value of the alloy in-creases to 850 mV after the dual sodium and oxygen plasma treat-ment [42]. The recent study by Li et al. [46] shows that sequentialAg and N PIII&D into titanium produces nanoparticles with a size of~10 nm on the surface and X-ray photoelectron spectroscopy (XPS)depth profiling reveals the formation of a 100 nm thick Ag film onthe surface as shown in Fig. 7a. On account of the large amount ofAg nanoparticles and thicker TiN layer, potentiodynamic polarizationand nanoindentation tests reveal that the Ag and N co-implanted tita-nium sample exhibits the highest corrosion resistance and hardness,as illustrated in Fig. 7b and c, respectively [46]. Hence, dual PIII&D en-hances the corrosion resistance and mechanical properties ofTi-based alloys.

2.1.2. OxidationOxidation of Ti-based alloys produces a titanium oxide surface

film, but oxidation at a high temperature in air frequently yieldsworse corrosion resistance [17]. Hence, alternative oxidation

methods have been proposed to improve the biocompatibility andcorrosion resistance of titanium alloys [21,49,50].

2.1.2.1. Hydrogen peroxide oxidation. Since the surface of titanium issensitive to H2O2 [51,52], H2O2 is often used to pretreat Ti-based al-loys to enhance the bioactivity [53,54]. The study by Chu et al. [21]shows that low-temperature oxidation of NiTi shape memory alloy(SMA) in H2O2 forms a titania scale enriched with Ti–OH groups.Wang et al. [55] have found that a subsequent heat treatment pro-duces a micro/nano structured surface on the titania film, as shownin Fig. 8, and it influences the surface hydrophobic properties whichare different from those observed from that of the NiTi sample oxi-dized by H2O2 but not heat treated [21]. The electrochemical testsalso reveal that the oxidized and annealed film possesses improvedcorrosion resistance [55].

2.1.2.2. Fenton's oxidation. Fenton's oxidation which is used industrial-ly to remove inorganic and organic pollutants from waste water isconducted by catalytically decomposing H2O2 into hydroxyl radicals(OH−) according to the following reactions [56]:

Fe2þþH2O2→Fe3þþOH• þ OH− andH2O2þFe3þ→HOO• þHþ þ Fe2þ:

The ferrous ion (Fe2+) initiates and catalyzes the decompositionof H2O2, resulting in the generation of hydroxyl radicals. The hydroxylradicals which have one unpaired electron are strong and highly reac-tive oxidants and have a higher oxidation potential (2.8 V) than H2O2.

Chu et al. [49] compared the results obtained by chemicallypolishing of NiTi (Fig. 9a) to those acquired by Fenton's oxidation.A more uniform and denser nanostructured titania film with depletedNi in the surface region and a graded interface with the NiTi sub-strate are observed from the NiTi sample after Fenton's oxidation(Fig. 9b). As shown in Fig. 9c, this film withstands large plastic defor-mation similar to NiTi alloys in the full martensite state (−196 °C)and can almost recover its original shape similar to NiTi SMA in thefull austenite state, i.e. when the temperature is higher than Af (aus-tenite finish temperature). The results disclose that the surface filmhas very good mechanical stability. In addition, electrochemicaltests show that the Fenton oxidized NiTi has a higher Eb and lowercorrosion current than the chemically polished sample (Fig. 9d). Thefavorable results are attributed to the good mechanical propertiesand enhanced corrosion resistance offered by the graded interfacestructure between the titania film and NiTi substrate. In addition,the nanostructured titania film can better accommodate elongationand compression [49].

2.1.2.3. Micro-arc oxidation. Micro-arc oxidation (MAO) is performedin an aqueous electrolyte by applying a pulsed DC current to thespecimen. It is a room-temperature electrochemical process and suit-able for the formation of thick and uniform coatings on componentswith complex geometries. This technique has been utilized to fabri-cate oxide ceramic coatings on Ti-based alloys, and the MAO coatingsenhance not only the corrosion and wear resistance, but also the bi-ological performance [50,57–59]. Wang et al. [50] have prepared tita-nium oxide ceramic coatings by MAO in the galvanostatic regime onbiomedical NiTi alloys in a H3PO4 electrolyte using a DC power sup-ply. The layer produced exhibits varied micro/nano surface morphol-ogies (Fig. 10a) depending on the MAO time. The potentiodynamicpolarization tests disclose enhanced corrosion resistance on titaniumafter MAO (Fig. 10b). By changing the chemical composition of theelectrolyte, different MAO coatings can be obtained. For instance,Liu et al. [57] have fabricated a protective layer composed of Al2O3

on the surface of NiTi alloy to improve the wear resistance, asshown in Fig. 11.

Fig. 7. Effects of Ag/N dual implantation on pure titanium: (a) Ag XPS depth profile;(b) potentiodynamic polarization behavior; and (c) surface hardness [46].


Fig. 8. SEM images of TiO2 film prepared at different temperatures: (a) 350 °C, (b) 400 °C, (c) 500 °C, and (d) 600 °C [55].

Fig. 9. Effects of Fenton's oxidation on NiTi SMAs: Surface morphology of (a) chemically polished sample and (b) sample after Fenton's oxidation for 24 h. (c) Shape recovery be-havior of NiTi strips. Left: chemically polished NiTi; right: after Fenton's oxidation for 24 h. The samples are subjected to large plastic deformation at−196 °C and then warmed upto Af (left to right). (d) Electrochemical behavior of Fenton oxidized NiTi (curve 1) and chemically polished NiTi (curve 2) [49].


2.1.3. Hydrothermal processesAlkaline hydrothermal treatment is effective in enhancing the bio-

compatibility of biometals [60–62]. After H2O2 treatment, Chu et al.[54] immersed NiTi (50.8 at.% Ni) SMA into a 10 M NaOH aqueous so-lution at 60 °C for 24 h. A titania (TiO2) layer was first found on thesurface of the NiTi substrate, and then a porous sodium titanate

(Na2TiO3)/titania film with many TiOOH groups and a trace of Ni2O3

was produced in situ by the reaction between the partial TiO2 phasein the titania layer and NaOH solution. After surface modification,the most satisfactory bioactivity was observed from samplessubjected to both oxidation and hydrothermal treatment, and thehighest corrosion resistance was displayed by the samples after un-dergoing alkaline treatment followed by heating at 700 °C [61].

2.1.4. Plasma sprayingPlasma spraying is a subset of thermal spraying which uses an

electrical arc to generate a high-temperature plasma designed tomelt and spray materials onto a surface [63]. It is very cost-effectiveand a flexible means to deposit coatings with various thicknessesfrom the micro scale to the macro scale [63]. Liu et al. [18] have pro-duced a bioactive nanostructured TiO2 surface on Ti6Al4V with agrain size smaller than 50 nm using nanoparticle plasma sprayingfollowed by H PIII (Fig. 12). In this course, some kinds of Magneliphases are formed as usually found in such titania layer. This has im-portant repercussions on the biological behavior of such coating. Thetreated materials favor deposition of hydroxyapatite (HA). Xie et al.[20] have found that the dicalcium silicate (C2S)/yttria stabilized zir-conia (YSZ) composite coatings prepared by atmospheric-pressureplasma spraying have better durability and more superior mechanicalproperties than pure C2S coatings deposited on Ti6Al4V.

Fig. 10. MAO modified titanium: (a) Surface morphology and (b) Potentiodynamicpolarization curves [50].

Fig. 11.Weight loss from coated and uncoated NiTi against GCr15 steel ball after slidingfor 30 min [57].

Fig. 12. Surface SEM views of as-sprayed TiO2 surfaces: (a) Nano-TiO2 coating and(b) higher magnification picture of (a) [18].


2.1.5. ElectropolishingElectropolishing (EP) is conducted to improve the surface finish of

a metal by making it anodic in a suitable solution such as concentrat-ed phosphoric acid, sulfuric acid, their mixtures, or perchloric acid/acetic acid solutions [64]. Traditionally, electropolishing is employedto minimize the surface roughness of anodic metals but recently,this technique has been extended to produce a nanostructuredsurface in order to enhance the biocompatibility and corrosion resis-tance of Ti-based alloys [65–68]. Chu et al. [66] have fabricated a10 nm thick smooth titanium oxide film which suppresses nickelrelease from the NiTi substrate. On the basis of this work, a newphotoelectrocatalytic oxidation (PEO) process has been developed

to improve the surface properties of biomedical NiTi [67]. The NiTi al-loys are first electropolished before undergoing photoelectrocatalyticoxidation. A nano-textured and graded titanium oxide surface layerabout 250 nm thick is formed on the surface as shown in Fig. 13bthat inhibits nickel release (Fig. 13c). In comparison with EP,(EP+EPO) does not degrade the mechanical properties of the NiTi al-loys [67].

Fig. 13. Electropolished NiTi, surface morphology of NiTi (a) EP sample and(b) EP+PEO sample, and (c) the nickel release concentration in SBF after immersiontests at 37±0.5 °C for 2, 5 and 10 weeks [67].

Fig. 14. Mechanical properties and corrosion behavior of different NiTi samples modi-fied by MS: (a) Young's modulus [71] and (b) nano-hardness [71], and(c) potentiodynamic polarization curves in different solutions. Curve 1: EP NiTisubstrate, curve 2: ZrN film, and curve 3 ZrCN film [72].


2.1.6. Magnetron sputteringAs one of the most widely used physical vapor deposition (PVD)

techniques, magnetron sputtering (MS) is often employed to depositdiamond-like carbon (DLC) coatings on Ti-based alloys. It is aplasma-based technique in which inert gas atoms are ionized andaccelerated by the potential difference between the negatively biasedtarget (cathode) and anode to bombard the target surface resultingin sputtering of atoms which condense on a substrate to form afilm [69]. Using this technique, Chu et al. [70,71] have fabricatedZrN and ZrCN films on NiTi to improve the surface load bearing abilitydue to the enhanced Young's modulus and surface hardness, asillustrated in Fig. 14a and b. As shown in Fig. 14c, both the Eb (break-down potential) and Icorr (corrosion current) values measured bypotentiodynamic polarization confirm that the amorphous DLC filmproduced by MS improves the corrosion resistance of NiTi significant-ly [72].

2.1.7. Sol–gelA sol is a colloid suspension of solid particles in a liquid and a gel is

a substance that contains a continuous solid skeleton enclosing a con-tinuous liquid phase. The sol–gel process generally includes five mainsteps: (1) hydrolysis and polycondensation, (2) gelation, (3) aging,(4) drying, and (5) densification and crystallization. Traditionally,the sol–gel technique is used to deposit a thin ceramic coating lessthan 10 μm thick [63]. This process offers advantages such as coatinghomogeneity, stoichiometry control, purity, ease of processing, con-trolled composition, and ability to coat large and complex substrates[63,73,74]. By controlling the withdrawal speed during thedip-coating, an anatase film 40–50 nm thick can be produced andCheng et al. [75] have fabricated a film as thick as 400 nm on NiTi al-loys using the sol–gel technique in conjunction with ensuing hydro-thermal treatment. Using the sol–gel method and subsequent steamcrystallization, a crystallized titania film with a thickness of 750 nmhas been prepared on NiTi alloys (Fig. 15), the surface of which iscomposed of particles of about 100 nm [76]. The coated NiTi hasmuch better corrosion resistance than the untreated sample in thesimulated physiological solution [75,76]. Besides titania, other ceram-ic nano coatings such as HA and ZrO2 have been produced on Ti-basedalloys to enhance the mechanical properties and corrosion resistance[73,74].

2.2. Surface nanocrystallization

Nanocrystalline materials are structurally described to havenanometer-sized grains with a large number of grain boundaries, and

Fig. 15. SEM micrograph of titania coating on NiTi produced by sol–gel and streamprocess, showing a tilted view with a part of the coating removed [76].

Fig. 16. The cross-sectional microstructure of (a) coarse grain (CG) in the as-receivedNiTi; (b) shot peened NiTi specimens; and (c) bright field TEM image at a depth ofabout 13 μm of the shot peened NiTi, the inset is the corresponding selected area elec-tron diffraction (SAED) pattern [81].


possessing many novel properties not observed in their coarse-grainedcounterparts [77,78]. Generally, surface nanocrystallization of metals isachieved by surface mechanical attrition treatment (SMAT) that isoften performed using ultrasonic shot peening or sand blasting[78–80]. In the former process, small metallic balls with diameters of1–10 mm are usually used while the latter makes use of a flow of silicaparticles 200 to 300 μm size to blast the targets. The SMAT technique iseffective in enhancing the corrosion resistance and surface mechanicalproperties of biomaterials such as Ti-based alloys and stainless steels[79,81–84]. Hu et al. [81] have observed that shot peening induces var-ious degrees of grain refinement in NiTi along different depths. In

comparison with the original coarse grain (CG) size of 40 to 80 μm,the nanocrystallites have dimensions of 10 to 30 nm and the averagevalue is about 20 nm, as shown in Fig. 16. Fig. 17 reveals that boththe surface hardness and wear resistance are significantly enhanced.This arises from strengthening and work hardening of the NiTi inducedby grain size refinement during shot peening [81,82]. The surface nano-crystalline structure also improves the corrosion resistance of the NiTialloy, as indicated by the increased Eb and decreased corrosion currentdensity (Icorr), shown in Fig. 17c [82]. In comparison, sand blasting in-duces the formation of a nano-crystalline surface layer on pure titani-um with an average grain size of about 50 nm which produces about10% higher cycle fatigue strength than the untreated titanium. Sandblasting and sand blasting plus annealing improve the corrosion prop-erties of the alloy and among the two samples, the nano-crystallinesurface layer generated by the latter process yields better corrosionresistance [84].

2.3. One-dimensional nanostructures

One-dimensional nanostructures fabricated on Ti-based alloyshave been shown to enhance cell adhesion, proliferation, and spread-ing [85–88]. Nanotubes and nanowires, two most common one-dimensional nanostructures, are often produced on Ti-based materialsby anodic oxidation, hydrothermal process, or combined processes.

2.3.1. NanotubesA titania nanotube layer can be produced on Ti-based alloys by anod-

ic oxidation. The nanotube size and length can be readily controlled byadjusting the parameters in anodizing oxidation. The synthesized one-dimensional nanotubes play an important role in osseo-integration be-cause they accelerate the adhesion of bone cells onto the TiO2 [89,90].Cell adhesion can be improved by up to 400% due to mechanicalinterlocking between the osteoblasts and nanotubular titanium oxidelayer, that is, an interlocked cell structure with the topography ofnanotubes produced by the filopodia of the growing cells into the nano-tube pores [90]. Liu et al. [91] have fabricated a series of nanotubes withdifferent diameters on titanium by changing the anodic voltages(Fig. 18) and anodic polarization and electrochemical impedance spec-tra (EIS) were used to determine the corrosion resistance. Whereasthese nanotubes exhibit higher open-circuit potentials and a smallercorrosion current density at 0.4 VSCE than the mechanically polishedTi in artificial saliva, its size also can significantly influence the electro-chemically stability. Accordingly, nanotubes with a size of 15 nm havethe highest corrosion resistance [91]. Titania nanotubes with diametersbetween 30 and 110 nm have been synthesized on Ti–13Nb–13Zr alloyby anodic oxidation. They degrade the corrosion resistance but follow-ing a heat treatment at 150 °C, positive results are observed from thecorrosion current density and potentiostatic polarization plots in simu-lated physiological solutions [92]. Readers are referred to Ref. [93] forthe detailed characteristics and properties of the titania nanotubesprepared on Ti-based alloys [93].

2.3.2. Nanowires or nanofibersOne-dimensional nanowire or nanofiber scaffolds provide a larger

surface area to absorb proteins and present more binding sites to cellmembrane receptors thus favoring protein absorption [94]. Thesenanowires and nanofibers also boast the ability to attract mineral de-position in vitro [85]. Generally, peptide nanofibers self-assemble andcovalently adhere onto the Ti-based alloy by pre-tailoring the surfacechemistry without changing the mechanical properties of the sub-strate. Stupp et al. [86–88] have fabricated one-dimensional peptidenanofibers on Ti6Al4V and NiTi alloys to enhance the surface biocom-patibility, as shown in Fig. 19a, but there has been no study on thesubsequent effect on the corrosion behavior of these alloys. Titaniaand titanate nanowires can be easily fabricated on Ti-based alloyshydrothermally and the representative SEM image is depicted in

Fig. 17. Mechanical properties and corrosion resistance of NiTi after shot peening(SMAT): (a) hardness vs. depth [81], (b) variation of wear volume with the appliedload during SMAT [81], and (c) Potential dynamic curves in the 0.9% NaCl physiologicalsolution. CG refers to coarse grain [82].


Fig. 19b [95,96]. The current research focus is on the interfacial be-havior between cells and nanowires/fibers as extracellular matrix(ECM) but there have been a few reports on the associated mechani-cal properties and corrosion behavior. Titanate nanowire films havebeen prepared on titanium using alkali treatment (AT) and ATfollowed by heat treatment (ATH). These titanate nanowire filmsshow worse nano mechanical properties compared to the substratesuch as hardness of only 0.01 to 0.26 GPa and elastic moduli between1.7 and 2.8 [97], possibly ascribed to the loose structure of thenanowire films.

3. Surface nanostructures on porous Ti-based alloys

Besides excellent mechanical properties, a porous structure simi-lar to that of natural bone allows tissues ingrowth and mass transferof body fluid and nutrition. Hence, porous Ti-based alloys arepromising substitutes for hard tissue repair or reconstruction[14,86,96,98–100] and can be fabricated by powder metallurgy[99–101]. Similar to other artificial biomaterials, the surface of theseporous alloys has to be modified to satisfy clinical requirementssuch as biocompatibility, corrosion resistance, surface mechanicalproperties, and biomimetic performance. However, the 3D porousstructure renders traditional line-of-sight techniques such as laser ni-triding, PVD and CVD difficult and ineffective. Some non-line-of sighttechniques can treat all the exposed area of porous Ti-based alloy, butonly induce the formation of bioactive titanium oxide nano films[14,102,103]. Our recent work reveals that a low-temperature hydro-thermal treatment can induce the formation of titanate nanowires onthe entire exposed areas of 3D porous Ti-based metals such as Ti andNiTi. The process is schematically illustrated in Fig. 20 [96]. The elec-trochemical stability of the surface-modified porous NiTi is assessedin simulated body fluid (SBF) by electrochemical impedance spec-troscopy (EIS) [104] and the raw and modeled data are shown asthe Nyquist plots in Fig. 21. The EIS results reveal that the surfacestructures produced by different modification techniques influencethe electrochemical behavior of the porous surface, and O PIII and

Fig. 18. Surface morphologies of TiO2 nanotubular layer formed on Ti by anodization in 0.5 wt.% HF at: (a) 5 V, (b) 10 V, (c) 15 V, and (d) 20 V [91].

Fig. 19. Morphologies of one-dimensional nano phases: (a) peptide nanofibers [88]and (b) titanate nanowires on Ti [95].


chemical treatment are useful techniques to produce protective nanofilms on the surface of porous NiTi alloys [104].

4. Conclusion

Recent progress related to the design and fabrication of surfacenano-architectures on orthopedic Ti-based alloys is reviewed. Thesenano-architectures can be divided into three categories: micro/nanofilms or layers, nanocrystalline surfaces, and on-dimensional nanostruc-tures.Micro/nanofilms and layers are typically produced on ceramicma-terials to improve the biocompatibility in vitro and in vivo by techniquessuch as plasma immersion ion implantation and deposition, oxidation,hydrothermal processes, plasma spraying, electropolishing, magnetronsputtering, and sol–gel process. The effects of nano films on themechan-ical properties and corrosion behavior are discussed. In general, thesefilms enhance the corrosion resistance and mechanical properties suchas surface hardness, elasticmodulus, andwear performance. Sometimes,multiple techniques ormethods yield better results than a single process.Surface nanocrystallization triggered by shot peening or sand blasting,namely SMAT (surface mechanical attrition treatment), can effectivelyenhance surface mechanical properties such as elastic modulus andwear resistance due to work hardening. The refined grains also play arole in the corrosion resistance improvement. One-dimensional nano-structures like nanowires, nanofibers, and nanotubes are attractingmuch attention because they can create a biomimetic structure or inter-face that can actively adjust protein adsorption as well as cell and tissuebehavior on the biomedical implants. However, their effects on the sur-face corrosion resistance andmechanical properties tend to vary. For ex-ample, titania nanotubes and the annealed ones exhibit differentcorrosion properties and the titanate nanowire films in fact have inferiormechanical performance. Recent research suggests that non-line-

of-sight surface modification techniques can effectively produce nano-structures on all the exposed surfaces on porous Ti-based alloys thusboding well for the anti-corrosion capability.

Acknowledgments

This work was jointly supported by the City University of HongKong Applied Research Grant (ARG) No. 9667066, the Hong Kong Re-search Grants Council (RGC) General Research Funds (GRF) No. CityU112212, NSFC Nos. 50901032, 51101053, and 81271715, the Ministryof Education Specialized Research Foundation for Doctoral Program ofUniversities No. 20094208120003, as well as the Wuhan ChenGuangResearch Program Grant No. 201150431134.

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Fig. 21. Nyquist diagrams of porous NiTi alloys measured in SBF at 37 °C and EIS is modeled by R(QR)(Q(R(QR))): (a) sample MP, (b) oxidation sample A, (c) O-PIII sample B,(d) O-PIII plus subsequent 450 °C air oxidation sample C, and (e) H2O2 pre-oxidation at 80 °C for 4 h plus subsequent NaOH treatment at 60 °C for 24 h sample D [104].


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