Permanent implants are necessary for long term, sometimes for the whole life, for many patients; so, the requirements of very good biocompatible materials necessitate especially metals and alloys that do not produce undesirable, local or systemic effects, have most appropriate host response and assure clinical performance of the therapy [1]. The most important requirements of an implant material are: suitability for load-bearing application, low density and Young's modulus (very closed of the natural bone, 1.8-2.1 g/cm2 and respectively, 10-40 GPa), high specific strength, proper surface morphology, topography, energy, and mechanical properties, high corrosion and degradation resistance due to the abrade and toxic effects of the corrosion products [2,3] and very good biocompatibility.
1.1. CoCr alloys as permanent implants Authors: Daniela Ionita and Anca Mazare
Co-Cr-Mo alloys have been successfully used since 1930 as implant materials due to their corrosion resistance in physiological fluids, mechanical properties, wear resistance and good biocompatibility. Such an alloy, named vitallium, with a content of 30% Cr, 7% W and 0.5% C was used in dentistry and was introduced as an alternative to gold based products, which had a very high cost.
As proof that history repeats itself also in the world of materials, depending on the existing technologies, in the last years the use of Co-Cr alloys increased due to the expansion of surface treatment techniques, which improve the quality of materials at a superficial level, thus permitting the replacement of high cost bioimplant as titanium.
From the numerous available Co-Cr alloys the most used as biomaterials are Co-Cr-Mo, standardized ASTM F-75 and F-76, and Co-Ni-Cr-Mo, standardized ASTM F-562. Although alloys with a high content of nickel (25-37%) have an increased corrosion resistance, are problematic due to the toxicity of Ni ions released in the organism. Therefore, the alloy with the most usages is, for now, Co-Cr-Mo with ASTM F-75 composition, also considering that adding of noble metals in the composition didn't improve the electrochemical stability.
In orthopaedy, devices obtained from Co-Cr alloys, as prosthesis, rods, screws, plates, sutures, are used for repairing functions, correcting or fixing the bone tissue or joint after some pathological or traumatic conditions and in dentistry, metals are used for obtaining dental implants, bridges, dental pivots, crowns and other devices which require mechanical and wear resistance.
The corrosion resistance of the CoCrMo alloy is due to the superficial oxide layer formed of Cr2O3 and CrO2. Regarding the CoCr alloy used mainly for hip prosthesis in the case of elderly people, starting from the 1980, it was mentioned in literature that even in the case of patients with properly working prosthesis the presence of an accumulation of Co and Cr ions in the liver, kidneys and spleen 5 or 8 times higher than normal values. It time it was proven that due to the accumulation of Co and Cr ions released from the implant material, the response of the inflamed tissue can lead to implant failure.
Materials used for biomedical applications cover a wide spectrum and must exhibit specific properties. The most important property of materials used for fabricating implants is biocompatibility, followed by corrosion resistance.
Due to their lower corrosion resistance in physiologic media (when compared to other alloys), CoCr alloys devices are used only for temporary implants. However, recently, due to economic reasons, different surface treatments of stainless steel were investigated so that these alloys can be used as permanent implants. In recent decades, a large amount of research has been devoted to surface modification of implant materials. Surface modification technologies can be used to improve biocompatibility, enhance bone bonding, reduce wear and/or corrosion and subsequent release of potentially toxic materials into the body and avoid blood clot formation, etc.
A. Surface nano-functionalization technologies
Recent work has established that key biological processes, including protein adsorption, cell proliferation and gene expression, can be controlled to some extent by using chemical methods to modify the surface properties of biocompatible materials [4].
The most popular and efficient ways to modify surfaces on the nanoscale level involve direct chemical changes with acids and oxidants, deposition of bioactive layer, deposition of antimicrobial layers.
a) Chemical treatments
Chemical treatments using acids, alkalis, and hydrogen peroxide involve chemical reactions at the interface between the materials and solution [4]. Different chemical treatments with acids [5-7] have been used to create micrometer-scale and submicrometer-scale textures on surfaces. Such studies have revealed that chemically treated surfaces can enhance the adhesion and proliferation of osteogenic cells [8-12].
In our case, samples of Co-Cr alloys with dimensions of 10 mm in diameter and 2 mm in height were used. A Co-Cr-Mo alloy with composition very close to that of ASTM F75 alloy was employed. The composition of Co-Cr-Mo alloys, nickel free, was: 28% Cr, 5.2% Mo, 1% Si, 1% Mn, 1% Ni, 0.75% Fe, 0.35% C, 62.7% Co. These samples were ground with silicon carbide papers ranging from 80 to 1200 grit, washed with deionised water and ethanol, dried in air and stored in a desiccator. Finally, before testing, the samples were passivated to simulate clinical conditions in 20% nitric acid for 30 min at room temperature (ASTM F86 standard practice for surface preparation and marking of metallic surgical implants). Nitric acid treatments increased the thickness of surface oxide with higher concentrations of oxygen and chromium that, in turn, appeared to be good for the anti-corrosion properties of the ï¬lm [13].
AFM analysis of chemically treated CoCrMo alloys showed an average roughness of 0.186 ïm ± 0.10 mm.
b) Deposition of bioactive layers like hydroxyapatite or bioactive glass to promote direct bonding with bone
This biomimetic process generally consists of a chemical treatment in an alkaline solution followed by a heat treatment and ending with immersion in a simulated body fluid (SBF). The immersion in SBF can be considered as a first-stage procedure for the bioactivity assessment of a biocompatible material [14].
Following the mechanical processing, the metallic samples were stored in 5M NaOH solution at 60oC for 24 h. After this alkaline treatment, samples were washed with distilled water and subsequently dried for 8 h at 220oC.
The SEM analysis after the alkaline treatment shows a porous surface (Fig. 1) and the EDS spectrum indicated the presence on the surface of the CoCrMo samples of some Na containing compounds, such as chromates which were stabilized during the thermal treatment.
Fig. 1. SEM and EDS image for CoCrMo alloy treated in 5M NaOH
After the thermal treatment, CoCrMo samples were immersed in SBF for 7 days at 37oC and then for 14 days in 1.5 SBF at 37oC.
SEM analysis after immersion showed the deposition on the surface of the CoCrMo samples of a thin layer of amorphous calcium phosphate which will later be used to induce the precipitation of hydroxyapatite crystalline layers. Fig. 2 shows the SEM image and the EDS spectra for the CoCrMo samples covered with hydroxyapatite.
Fig. 2. SEM image and EDS spectrum of the CoCrMo sample covered with HA
After the deposition of hydroxyapatite, the properties of the covered CoCrMo surface were significantly changed. SEM analysis showed a porous layer with pore dimensions of about 5µm and a roughness three times higher than that of grounded surfaces (0.624 nm).
Hydroxyapatite depositions on CoCrMo alloy can be made by both chemical and electrochemical methods and by laser deposition.
Electrochemical deposition has the advantage of creating coatings comprised of wire like nanometric crystals of hydroxyapatite [15] which enhance bone remodelling and maturation [16].
Several authors investigated different conditions of electrolytic process: composition of the electrolyte [17], electrolyte temperature [18], current density [19], composition of the substrate metal, hydrothermal treatment after electrochemical synthesis [20], and hydrothermal-electrochemical conditions, that affect different characteristics of interface as the deposit amounts, size, shape, morphology, microstructure of the HA deposit and the strength bonding of the HA-metallic substrate interface
c) Obtaining of n-Ag coating on CoCrMo substrate
After mechanical processing, the metallic electrodes were silanized for 1.5 hours in APS (3-aminopropyltrietoxisilane), washed again with acetone and distilled water and dried at 25oC. Samples were then subjected to a 3 h incubation of the electrodes in a solution containing: 20 ml of 0.25mM AgNO3, 25 ml of 0.25mM trisodium citrate on which 0.6ml of 10mM NaBH4 were added under vigorous stirring [21].
AFM technique enabled the study of the covered surfaces' topography and the scanned areas measured 20Ã-20 µm, in contact mode. Fig. 3 presents the 3D AFM image of the n-Ag film deposited on the CoCrMo samples. Fig. 3 discloses the dendritic nature of the n-Ag deposition, the value of the average square roughness being high (758 nm).
Fig. 3. AFM image of the n-Ag film deposited on the CoCrMo alloys
The presence of the n-Ag on the surface of the CoCrMo samples led to an increase in the roughness (Rq) from 451 to 545 nm.
A recent review paper by Mendonca et al. [22], as well as some clinical reports [23-25], have indicated that surface change is a valuable strategy to give metallic implants new biological functionalities for in vivo clinical applications.
B. Corrosion in the body
Biocompatibility is the primary requirement for biomaterials. Biocompatibility of implant devices relies on several issues which are presented in Fig. 4.
Fig. 4. Biocompatibility of implant devices
As you can see, corrosion is an important factor that can affect the biocompatibility. The physiological solution is considered a corrosive media to metallic materials. The corrosion of metallic implants due to the body environment can affect the human life in different ways: (i) it may release undesirable metal ions/corrosion products which are non-biocompatible; (ii) it may reduce the life of implant device and therefore, may impose another costly and painful surgery, and (iii) ultimately reduce the human life.
Concerning the electrochemical methods used for characterization of the corrosion processes, they can be divided in two categories. First the polarization methods used to assess the susceptibility to localized corrosion for corrosion resistant materials. Second, measurements performed at the free corrosion potential: open circuit potential (OCP) and electrochemical impedance spectroscopy (EIS) can be used to follow actively corroding systems such as degradable implants.
Corrosion studies on biomaterials have been frequently carried out in simulated body fluid (SBF), Hank's solution, Ringer's solution, and artiï¬cial saliva [26]. The pH of the normal blood and interstitial fluid usually remains between 7.35-7.45, however, it may decrease to 5.2 during implantation in hard tissues, and should return to its normal pH within two weeks of time while the temperature remains about 370C [27].
Several works have been devoted to study surface oxide chemistry of Co-Cr alloy, produced under different conditions [28-31].
The variation of the open circuit potential (OCP) for uncovered samples presents a tendency to shift towards the more active potentials area, until a stability area is achieved. In contrast, the potential for covered samples varies towards more electropositive potentials and achieving, in a relatively short time, an equilibrium state. From the variation of the potential in time, it is observed that the equilibrium state is achieved the fastest for the samples covered with hydroxyapatite. The equilibrium state achieved by the electrodes indicates the presence on the surfaces of CoCrMo alloys of a stable film. The values of the potential after 60 min for the CoCrMo electrode with different surface treatments and different physiological media are presented in Table 1. A positive shift in the OCP was an indication of a stable coating/insulation behaviour [32].
The electrode potential values registered for the SBF solution are more electropositive than those registered for Ringer solution, which is assigned to the composition of SBF, from which, in time, carbonates and phosphates from the solution are deposited on the surface of the metallic electrodes. According to literature date [33], Ca-P compounds precipitate spontaneously from biological fluids when the pH of the solution is 7.4.
Table 1 Potentials after 60 min of immersion in different physiological media for CoCrMo electrodes with different surface treatments
Solution
E (mV) vs. SCE
CoCrMo untreated
E (V) vs. SCE
CoCrMo/HA
E (V) vs. SCE
CoCrMo/n-Ag
SBF
-362
40
120
Hank
-412
-25
20
Ringer
-426
-12
30
Afnor-artificial saliva
-400
10
50
Observations regarding local corrosion were obtained with potentiodynamic polarization and linear polarization methods. Potentiodynamic polarization consisted of plotting the cyclic polarization curves obtained with a Voltalab 40 PGZ 301 potentiostat/ galvanostat and a computer interface - software VoltaMaster 4, by following both the anodic polarization curve and the hysteresis curve. The plotting of the curves began at -800 mV till 1200 mV with a sweep rate of 2mV/s. From these polarization curves the main electrochemical parameters characterizing the corrosion processes and passivation processes were identified: corrosion potential (Ecorr); corrosion critical current (Icorr) and corrosion rate (Vcorr), passivation domain. A platinum electrode was used as an auxiliary electrode, and as reference electrode, a Ag/AgCl electrode immersed in a saturated solution of KCl was used.
The linear polarization method was used to obtain the polarisation resistance (Rp), the corrosion current densities (icorr), corrosion rates (Vcorr) and the ion release rates. The linear polarisation was applied for 50 mV around the open circuit potential with a scan rate of 10 mV/sec, using Voltalab 40 equipment. VoltaMaster 4 program directly supplied the values of Rp, icorr and Vcorr.
From the corrosion point of view, the prevailing redox conditions are important. The oxygen content in the surroundings can vary depending on the specific application. In the case of alloys, the passivity of which is based on the presence of Cr2O3-rich passive film on the surface, highly oxidizing conditions can lead to dissolution of the passive film by formation of soluble Cr(VI) species. Besides dissolved molecular oxygen O2, more active oxygen species such as H2O2 can be formed in biological reactions. The stability of the passive film is dependent on the availability of oxygen.
Fig. 5 presents the Tafel plots of CoCrMo uncovered and covered with HA in SBF solution, from which the electrochemical parameters were determined (Table 2).
Fig. 5. Tafel plots for uncovered CoCrMo and CoCrMo covered with HA in SBF solution
Table 2 Electrochemical parameters for CoCrMo uncovered and covered in SBF solution
Sample
icorr, μA/cm2
Vcorr, μm/year
uncovered CoCrMo
0.414
3.973
CoCrMo covered with HA
0.320
3.071
CoCrMo/nAg
0.108
1.028
Cyclic voltammetry tests (Fig. 6) showed that the presence of coatings on the metallic substrates led to an improvement of the electrochemical stability of the biomaterials in the studied physiological fluids.
There is no active/passive transition and the passive range is established immediately following the Tafel region. The passive range extends from -/0.3 V to approximately 0.4 V for uncoated samples or 0.6 V for CoCrMo/HA at which potential it is interrupted by the current density increase.
At potential higher than +0.8 V (SCE), the potentiodynamic curve shows a peak which, based on [34] can be attributed to the transpassive oxidation of Cr3+ to Cr6+ according to the following reaction:
3Cr3+ + 4H2O → Cr2O3 + 6H+ + 6e−
Fig. 6. Cyclic voltammetry tests
The nanoscale surface modiï¬cation introduces novel bioactive capacity into the arena of metallic biomaterials, and many of the modiï¬cation approaches described in this paper still remain to be tested in vivo.
As a surface modification, several different types of coatings have been employed to reduce the corrosion rate [35]. In addition, ion plating [36] or ion implantation [37] with Ti reduces the corrosion current density.
In conclusion, advances in surface-engineering techniques and nanotechnology promise a new generation of improved prosthetic devices with selective bioactive surfaces, and eventually with ''intelligent surfaces''.
One of the major criticisms of the actual in vitro electrochemical characterization of medical implants is the absence of biological species in the simulated physiological media used. The challenge for corrosion research will be to perform electrochemical tests in media containing living cells and proteins.
1.2. 316L stainless steel as permanent implant Authors: Daniela Ionita and Anca Mazare
Surgical implants are usually made of metallic materials, such as titanium and its alloys, stainless steels and cobalt - chromium alloys. Among all the metallic materials, stainless steel is the most popular because of their relatively low cost, ease of fabrication and reasonable corrosion resistance. However, stainless steel is susceptible to a number of localized corrosion, such as pitting and crevice corrosion, intergranular corrosion and stress corrosion cracking [38].
Stainless steel implants are used as temporary implants to help bone healing, as well as fixed implants such as for artificial joints. Typical temporary applications are plates, medullar nails, screws, pins, sutures and steel threads and networks used in fixing fractures [39]. Although stainless steel is seldom used in developed countries as permanent implants, it is still the most used in emerging countries [40].
The susceptibility of stainless steel to the different types of corrosion, especially pitting corrosion depends primarily on the environmental parameters besides the chemical composition and metallurgical manufacturing condition of the steels. The effects of various anions present in surrounding environment on the pitting of stainless steel have been studied by many authors. Zuo et al [41] reported the inhibition effects of OH-, NO3-, SO42-, ClO4- and acetate ions on pitting of stainless steel in chloride solutions. An increase of Cr content strongly increases the resistance against localized breakdown of passivity.
There are several ways to measure and predict the corrosion behaviour of medical implant alloys that simulate the in vivo situation. For example, Electrochemical Impedance Spectroscopy method can be used to approach the oxide's resistance and capacitance as well as model the oxide with a electronic circuit to further predict the corrosion behaviour in different conditions. Potentiodynamic test scans through a high amplitude DC voltage and the current vs. voltage curves provide information about the corrosion potential, breakdown potential, repassivation potential, and corrosion current density of the test alloys. The area of hysteresis in current vs. voltage curves indicates the susceptibility of metal to pitting corrosion. Linear polarization tests apply potential 10 mV more and less than corrosion potential to predict the corrosion current from the slope of potential vs. current
A stable oxide layer on the passivated metal surface will help to render its corrosion resistance and possess relatively inert in physiological conditions. This passivity could be enhanced by modifying either thickness, morphology, or chemical composition of the surface oxide layer by different treatments [42]. Passivation which improves the surface properties of 316L stainless steel can be performed either thermally, or electrochemically, or passivated in nitric acid [43].
Our research in accordance with other researchers [44] has shown changes in behaviour depending steel surface treatment applied. Thus, in Ringer solution, the nitric acid passivated steel showed corrosion currents babies (icorr = 45 nA/cm2 for nitric passivation) from electrochemically passivated steel (icorr = 76 nA/cm2 for electrochemical treatment) and untreated steel (icorr = 76 nA/cm2). Corrosion potential of surface treated stainless steel is nobler in comparison with the untreated 316L. Surface treatment of 316L stainless steel improves the biocompatibility of the implant because corrosion behaviour is improved as a biocompatibility character.
Many attempts have been made to modify the surface of metal implants with ceramic coatings which intended to minimise the corrosion rate and arrest the release of metallic ions or corrosion products. The literature searches show that the application of ceramic coatings to metal alloys can improve corrosion resistance, wear resistance and bioreactivity that are beneficial for medical devices [45, 46]. It is becoming clear that a continuing need for stringent research and development of new technologies for improving the performance characteristics of current available SS 316L implant is very significant.
1.3. Titanium and titanium alloys as permanent implants Authors: Cora Vasilescu and Silviu Iulian Drob
1.3.1. Titanium as permanent implants
Titanium [47,48] was used as implant material for many years, for its excellent corrosion resistance, due to its native passive film with a thickness of few nanometres. But, titanium has a high Young's modulus of 105 GPa and a low bioactivity that can produce an inefficient implant fixation or implant failure. Many researchers tried to enhance the titanium biocompatibility by different surface treatments as: anodisation [47,49], ion implantation [48], chemical and thermal treatment [50-52], mechanical and chemical treatment [53], sol-gel deposition [54], etc.
The necessity of better implant materials imposed the obtaining of titanium alloys with very good properties.
1.3.2. Titanium alloys as permanent implants
Many titanium alloys were elaborated to satisfy the most important requirements for a good implant material: low density and Young's modulus closed to bone, good resistance to load-bearing, wear, fatigue, tensile stress, high resistance to corrosive attack of the aggressive ions from the human fluid, specially chloride ions, good biocompatibility, bioactivity, ossteointegration, ossteoactivity, etc. Recently, the improvement of the titanium properties as implant material was realized by binary, ternary, quaternary and multi-component alloys.
a) Binary titanium alloys
The alloys containing only non-toxic and non-allergic elements Nb, Ta, Mo, Au, Ag, Pt, Pd will be analysed.
Ti-Nb alloys. Niobium is a β phase stabilizer element with good mechanical properties and high corrosion resistance [55]. Wang et al. [56] obtained Ti-22Nb alloy and Lopes et al. [57] studied Ti-33Nb alloy; both binary alloys have higher corrosion resistance than of Ti or another commercial alloys; their biocompatibility was not yet determined. The Young's modulus of β Ti-xNb (32 to 52 mass %) alloys [58] varied from 72 to 62 GPa with the minimum value of 62 GPa for Ti-42Nb alloy.
Ti-Ta alloys. Tantalum is immune to attack of almost acids, excepting for concentrated hydrofluoric acid [59]. Ti-Ta binary alloys were obtained: Ti-13Ta [60], Ti-xTa (x = 30, 40, 50, 60) [59]; their behaviour in artificial saliva of different pH values, without or with fluoride content was investigated and a better, nobler behaviour than of usually Ti-6Al-4V alloy resulted.
Ti-Mo alloys [61] acquired a typical β equiaxed grain microstructure and after annealing, quenching, cold rolling and recrystallization, an uniform, refined structure with fine grains was obtained. This recrystallization structure rendered to Ti-12Mo alloy a very good corrosion resistance to aggressive ions from physiological fluid, Ringer's solution.
Ti-1M (Ag, Au, Pd, Pt) alloys. Recently, F. Rosalbino et al. [62] obtained Ti-1M alloys with non-toxic elements, better mechanical properties and corrosion resistance than Ti and Ti-6Al-4V alloy. In fluoride containing environment, these alloys revealed an increase in stability of their passive oxide layers and consequently, a decrease of surface activity due to the incorporation of the noble metal cations into TiO2 lattice, thus increasing the dissolution resistance.
b) Ternary titanium alloys
Ternary alloys based on Ti, Nb, Zr, Ta, Sn non-toxic elements will be presented.
Ti-13Nb-13Zr alloy is the most studied ternary alloy [63-68]; this alloy developed a modulus of 79 GPa, exhibited a very high corrosion and wear resistance [64] in human biofluid, and very good biocompatibility [69]. Its ossteointegration was increased by different treatments that deposited on its surface apatite or hydroxyapatite [16,70-73]. However, its high modulus can create long term problems in the cases of the load-bearing solicitations.
Ti-15Zr-4Nb alloy [74-76] was heat treated and the composition of its passive layer was determined; for un-treated alloy, the passive film is formed by Ti2O3 oxide; for heat treated alloy, the oxide layer is composed of TiO2 in the form of rutile. The oxidized Ti-13Nb-13Zr alloy showed a lower corrosion resistance than Ti-6Al-7Nb alloy.
Ti-xNb-13Zr alloys (x = 5, 13, 20) [77] are promising implant materials due to their better biomechanical compatibility and more corrosion resistance than Ti and Ti-6Al-7Nb alloy in Ringer's solution.
Ti-35Nb-xZr alloys (x = 3, 5, 7, 10, 15) [78] with solely β phase microstructure were treated by anodic oxidation to develop nanotubular oxide layers and then, the hydroxyapatite (HA) was deposited by magnetron sputtering method; The HA/Ti thin film-coated nanotubular Ti-35Nb-xZr alloys showed good corrosion resistance in 0.9% NaCl solution.
Ti-(23 to 43 mass %) Nb-(5 to 15 mass %) Sn alloys. Sn addition to binary Ti-Nb alloys decreased the Young's modulus of the ternary Ti-Nb-Sn alloys [58] till a very good value of 40 GPa due to Sn effect to suppression or retardation ï· transformation and to assure β pure phase.
Ti-25Ta-5Zr alloy [79] with α + β structure was obtained and was thermo-mechanical processed in order to optimise the balance strength-elastic modulus; proper mechanical properties and Young's modulus of 55 GPa were realised; treated alloy revealed the improvement of the electrochemical behaviour and mechanical properties as result of the favourable influence of the applied treatment. The alloy surface was processed by anodic oxidation [80] and the obtained oxide layer had a confinable rugosity, more hydrophilic character and very good corrosion resistance.
c) Quaternary titanium alloys
The quaternary alloys that have in their composition only biocompatible metals (Ti, Nb, Zr, Ta) will be examined below.
Ti-15Zr-4Nb-4Ta alloy (with α + β structure) free of cytotoxic elements was obtained and studied by Japanese researchers [81-85]. Having a Young's modulus of 97 GPa, a good balance of strength and ductility, a very good corrosion and biocompatibility, this alloy was considered a good candidate for implant use. This alloy proved a high apatite-forming ability by NaOH solution, CaCl2 solution, heat and water treatment [85,86]; this ability was maintained even in the humid environment; treated alloy has scratch resistance.
Ti-29Nb-13Ta-4.6Zr alloy [87,88] has a good modulus of 60 GPa and a high corrosion resistance in biofluids due to its passive film that contains very protective TiO2, Ta2O5, Nb2O5, ZrO2 oxides but is bioinert and cannot directly bond with living bone. Its bioactivity was enhanced by different surface treatments: heat treatment [89], heat treatment with calcium phosphate invert glass-ceramic [90].
Ti-34Nb-9Zr-8Ta alloy [91] has a β microstructure and relative low elastic modulus of 89 GPa. By aging treatment, the hardness and Young's modulus decreased.
Ti-13Mo-7Zr-3Fe alloy [91]. The composition (free of toxic elements) and heat treatments (homogenization or ageing) assure a homogeneous β structure, a Young's modulus of 89 GPa and good hardness.
Ti-35Nb-7Zr-5Ta alloy [92,93] was obtained by powder metallurgy and after ageing treatment exhibited a β matrix with refined α precipitates homogeneously distributed. The alloy demonstrated a high biocompatibility, a low Young's modulus, being good candidate for implant material.
1.3.3. Necessity of new permanent implant generation
The implant alloys are used for long term (20 years or more) and must satisfy many concerns: low density and Young's modulus does not cause insufficient bone bonding or bone resorption; high mechanical strength, fatigue and wear resistance [2,94] does not release ions, compounds or particles into the surrounding tissues; to contain only non-toxic and non-allergic elements that assure a good biocompatibility [1,95,96]; very good long term corrosion resistance to avoid the accumulation of ions and corrosion products in the adjacent tissues, which, in time, can produce important local changes of the biofluid pH and composition, generating potential gradients, accelerating the corrosion on some zones of the implant [97,98].
Currently available implant alloys do not realise all above mentioned necessities for a permanent implant and a new generation of implant alloys based only on biocompatible elements (Ti, Nb, Zr, Ta, etc.) is under attention of researchers. Taking into account that the quantity of the released metallic ions into physiological fluid is lower 0.3 mg/L [84], especially Ti, Nb, Ta and Zr can accomplish this condition; their biocompatibility decreased in the following order: Nb > Ta > Ti > Zr [95,96]; also, their corrosion resistance is very high because their resultant protective oxides ZrO2, Nb2O5 and Ta2O5 strengthen the TiO2 passive film formed on the alloy surfaces [99-101]. Nb stabilizes the β structure and is high passivating metal [100]. Zr is an izomorphous metal being soluble both in ï¡ and β titanium, limits the corrosion of ï¡ phase, increasing the corrosion resistance [63,102]. Ta has a superior corrosion resistance but limited mechanical resistance [103,104].
Therefore, a new generation of permanent implants is necessary to accomplish the majority of the demands for long term use: very good biocompatibility and corrosion resistance, mechanical properties closed to the bone, non-toxic and non-allergic effects, high fatigue and wear resistance, etc. Only few new alloys realise these conditions and data about their behaviour in physiological fluids are limited.
1.3.4. Anticorrosive performances of some novel titanium based alloys
Three new quaternary Ti-Nb-Zr-Ta alloys were obtained with the aim to satisfy the most requirements of a good permanent implant alloy. Ti-5Nb-10Zr-5Ta, Ti-10Nb-10Zr-5Ta, Ti-20Nb-10Zr-5Ta alloys differ by the Nb content and have a fine, homogeneous ï¡ + β microstructure characterised by a low Young's modulus of 58.24 GPa, 63.47 GPa and 60 GPa respectively. Their anticorrosive performances in Ringer solutions of different pH values, simulating the severe functional conditions of an implant will be presented in this chapter.
a) Alloy composition and sample obtaining
The alloys were obtained by high vacuum, levitation melting and re-melting using pure elements: titanium according to ASTM F 67, niobium 99.81% purity, zirconium 99.6% purity and tantalum 19.59% purity; the alloys compositions are showed in Table 3 that reveals a low level of impurities.
Table 3 Alloy compositions
Alloys
% wt.
Nb
Zr
Ta
Fe
O
N
H
Si
Mg
Al
Ti
Ti-5Nb-10Zr-5Ta
4.09
9.12
4.16
0.036
0.195
0.004
0.0016
0.0011
0.032
0.002
balance
Ti-10Nb-10Zr-5Ta
10.18
9.648
4.466
0.0001
0.15
0.03
0.002
0.0013
0.046
0.002
balance
Ti-20Nb-10Zr-5Ta
20.11
9.54
4.68
0.0001
0.12
0.02
0.011
0.078
0.166
0.192
balance
From ingots were cut cylindrical samples that firstly were grinded and polished to mirror surface; then, the samples were ultrasonically degreased in acetone and bi-distilled water and mounted in a Stern-Makrides mount system.
b) Solutions
The anticorrosive performances of the alloys were studied in Ringer solutions of acid, neutral and alkaline pH, simulating the possible severe functional conditions from human body: acid pH appears after surgery because the hydrogen concentration increases in traumatised tissues and by the in time hydrolysis of the surface oxides; alkaline pH develops in the distress periods of human body [105-107]. Ringer solution composition was (g/L): NaCl - 6.8; KCl - 0.4; CaCl2 - 0.2; MgSO4.7H2O - 0.2048; NaH2PO4.H2O - 0.1438; NaHCO3 - 1; glucose - 1; pH = 7.40; pH = 3.21 was obtained by HCl addition; pH = 9.05 was obtained by KOH addition.
c) Experimental techniques
Four electrochemical methods were used: potentiodynamic and linear polarisation, electrochemical impedance spectroscopy (EIS) and monitoring (for 1500 immersion hours) of the open circuit potentials, Eoc and open circuit potential gradients due to the non-uniformities of the Ringer solutions pH, ΔEoc(pH):
(1)
(2)
(3)
The potentiodynamic polarisation was applied from the cathodic (a potential with about 500 mV more electronegative than Eoc) to the anodic domain (till + 2000 mV vs. SCE) using a scan rate of 1 mV/s; VoltaLab 80 equipment with its VoltaMaster 4 program were used. From the curves, the main electrochemical parameters were determined: Ecorr - corrosion potential, like zero current potential, Ep - passivation potential at which the current density is constant; |Ecorr - Ep| difference represents the tendency to passivation (low values characterise a good, easy passivation); ï„Ep - passive potential range of the constant current; ip - passive current density. If the reverse curve presents lower currents than the direct curve, it results a very stable passive state.
The linear polarization was carried out to obtain Tafel curves for a potential range of  100 mV around Eoc, with a scan rate of 1 mV/sec. The VoltaMaster 4 program directly supplies the corrosion current densities, icorr and rates, Vcorr and polarization resistance, Rp.
The electrochemical impedance spectroscopy was performed at Eoc, using Voltalab 80 equipment; the amplitude of the AC potential was 5 mV and simple sine measurements at frequencies between 0.1 Hz and 102 kHz were applied for each sample. Nyquist and Bode plots were recorded. The electric equivalent circuit was fitted using non-linear, least square program ZVIEW.
d) Results and discussion
Electrochemical behaviour of the alloys
From potentiodynamic polarisation curves (Fig. 7) can be observed a self-passivation behaviour, without active-passive region, with a large passive potential range, ΔEp (more 2000 mV) and low passive current densities, ip for all those three studied alloys. Corrosion potentials, Ecorr (Table 4) exhibited more electropositive values with the increase of the niobium content, due to the favourable influence of this element that, having a nobler corrosion potential acts by its effect of the galvanic couple, ennobling the alloy corrosion potentials. Also, the passivation potential, Ep (Table 4) became more electropositive with the increasing concentration of niobium, as result of the same nobler behaviour of this metal. Tendency to passivation, |Ecorr - Ep| had the best values for the alloy containing the highest concentration of niobium. Passive current densities, ip had low values depicting a high corrosion resistance [77].
Fig. 7. Potentiodynamic polarization curves for Ti-5Nb-10Zr-5Ta, Ti-10Nb-10Zr-5Ta and
Ti-20Nb-10Zr-5Ta alloys in Ringer solutions at 370C
Table 4 Main electrochemical parameters for studied alloys
Alloy
Ecorr (mV)
Ep (mV)
|Ecorr-Ep| (mV)
∆Ep (mV)
ip (µA/cm2)
Ringer solution pH = 3.21
Ti-5Nb-10Zr-5Ta
-450
0
450
>2000
35.0
Ti-10Nb-10Zr-5Ta
-400
-200
200
>2000
4.2
Ti-20Nb-10Zr-5Ta
-300
-100
200
>2000
1.9
Ringer solution pH = 7.40
Ti-5Nb-10Zr-5Ta
-400
0
400
>2000
10.3
Ti-10Nb-10Zr-5Ta
-300
-150
150
>2000
3.9
Ti-20Nb-10Zr-5Ta
-200
-100
100
>2000
0.9
Ringer solution pH = 9.05
Ti-5Nb-10Zr-5Ta
-450
-150
300
>2000
30.1
Ti-10Nb-10Zr-5Ta
-400
-200
200
>2000
10.5
Ti-20Nb-10Zr-5Ta
-250
-150
100
>2000
3.5
Anticorrosive performances of the alloys
The corrosion current densities, icorr and corresponding corrosion rates, Vcorr and ion release rates from Table 5 have low values that characterise a very good corrosion resistance in the "Very Stable" class [97]. High values of the polarisation resistances, Rp show a very resistant passive film. The lowest corrosion rates were obtained in the neutral Ringer solution, the functional normal condition of an implant. Corrosion rates, Vcorr are lower and polarisation resistances, Rp are higher for the alloy with the highest niobium content, i.e. these most protective properties are due to the protective Nb2O5 oxide existing in the alloy passive film and that improves its passive characteristics [82,84].
Table 5 Main corrosion parameters for studied alloys
Alloy
icorr
(µA/cm2)
Vcorr
(µm/yr)
Resistance class
Rp
(kΩ∙cm2)
Ion release (ng/cm2)
Ringer solution pH = 3.21
Ti-5Nb-10Zr-5Ta
0.720
8.496
VS
154.32
863.19
Ti-10Nb-10Zr-5Ta
0.421
4.968
VS
165.13
504.75
Ti-20Nb-10Zr-5Ta
0.135
1.593
VS
180.62
161.85
Ringer solution pH = 7.40
Ti-5Nb-10Zr-5Ta
0.231
2.726
VS
210.24
276.96
Ti-10Nb-10Zr-5Ta
0.107
1.263
VS
230.65
128.32
Ti-20Nb-10Zr-5Ta
0.078
0.921
PS
250.46
93.57
Ringer solution pH = 9.05
Ti-5Nb-10Zr-5Ta
0.535
6.313
VS
163.78
641.40
Ti-10Nb-10Zr-5Ta
0.394
4.649
VS
172.95
472.34
Ti-20Nb-10Zr-5Ta
0.115
1.357
VS
198.54
137.87
VS - Very Stable
Electrochemical impedance spectroscopy studies
Nyquist plots (Fig. 8) exhibited an incomplete, large semicircle showing a capacitive behaviour, a passive insulating film on the surface of all these three alloys. The semicircle diameters and impedance values increased with the increasing niobium content, indicating a more stable, resistant passive film.
Fig. 8. Nyquist plots for Ti-5Nb-10Zr-5Ta, Ti-10Nb-10Zr-5Ta and Ti-20Nb-10Zr-5Ta alloys
in Ringer solutions at 370C
Bode phase plots (Fig. 9) displayed in the low and middle frequency range two phase angles: the first phase angle values varied between -650 and -850, indicating a typical passive films on the alloy surfaces and a near capacitive response of these films [65,87]; the second phase angle have values from -500 to -850, characterising some relaxation processes at the interface with the electrolyte. These angles have lower values in acid and alkaline Ringer solutions, showing a slightly defective passive film, low dissolution processes through film, due to the slightly higher aggressivity of these solutions. Comparing the those three alloys, the best angles of -850 were registered for Ti-20Nb-10Zr-5Ta alloy, demonstrating a better capacitive behaviour, more protective passive film due to the beneficial influence of the niobium.
Fig. 9. Bode plots for Ti-5Nb-10Zr-5Ta, Ti-10Nb-10Zr-5Ta and Ti-20Nb-10Zr-5Ta alloys
in Ringer solutions at 370C
Impedance spectra revealed a passive film with two layers; so, an electric equivalent circuit with two time constants was modelled (Fig. 10) [66,108]: the first time constant is represented by the inner, passive, barrier layer resistance, Rb and capacitance, CPEb; the second time constant is associated with the outer, porous layer resistance, Rp and capacitance, CPEp.
Fig. 10. Electric equivalent circuit
Fitting parameters from Table 6 revealed that the barrier layer resistance, Rb is higher than the porous layer resistance, Rp, i.e. the barrier layer is main responsible for the high corrosion resistance. CPEb component shows the capacitive behaviour of the barrier oxide layer and is responsible for the high phase angle; porous layer resistance, Rp represents the resistance of the porosities from the outer layer, where the species from the biofluid are incorporated; this outer porous layer is related to the alloy biocompatibility; CPEp is associated with the capacitive behaviour of the outer porous layer; the frequency independent parameter n when is equal with 1 indicates an ideal capacitor and when is lower than 0.7 demonstrates some interactions of ions from the electrolyte with the passive film. Capacitances of the porous layer, CPEp are superior to the capacitances of the barrier layer, CPEb showing that the corrosion resistance is due to the barrier layer. Resistances of the porous layer, Rp have lower values in acid and alkaline solutions due to some low corrosion processes; the barrier layer resistances, Rb remained highly (order 105 ï-.cm2) both in neutral, acid and alkaline solutions, proving a resistant, inner barrier layer.
All fitting parameters had the most favourable values for Ti-20Nb-10Zr-5Ta alloy as result of its most suitable composition.
Table 6 Fitting parameters for the electric equivalent circuit with two time constants
Alloy
Rsol
(Ωcm2)
Rb
(Ωcm2)
CPEb
(S sn cm-2)
n1
Rp
(Ωcm2)
CPEp
(S sn cm-2)
n2
Ringer solution pH = 3.21
Ti-5Nb-10Zr-5Ta
11.16
5.9x105
4.7x10-5
0.81
4.8x102
3.9x10-4
0.70
Ti-10Nb-10Zr-5Ta
12.54
6.2x105
4.1x10-5
0.81
5.1x102
3.7x10-4
0.70
Ti-20Nb-10Zr-5Ta
12.78
6.9x105
3.3x10-5
0.82
5.3x102
3.4x10-4
0.70
Ringer solution pH = 7.40
Ti-5Nb-10Zr-5Ta
10.54
7.2x105
3.7x10-5
0.88
4.1x103
1.8x10-4
0.75
Ti-10Nb-10Zr-5Ta
10.83
7.6x105
3.1x10-5
0.89
4.5x103
1.3x10-4
0.75
Ti-20Nb-10Zr-5Ta
10.95
8.1x105
2.6x10-5
0.90
4.9x103
0.9x10-4
0.76
Ringer solution pH = 9.05
Ti-5Nb-10Zr-5Ta
10.03
4.8x105
5.9x10-5
0.80
2.8x103
4.1x10-4
0.73
Ti-10Nb-10Zr-5Ta
10.19
5.3x105
5.2x10-5
0.81
1.3x103
3.6x10-4
0.74
Ti-20Nb-10Zr-5Ta
10.21
5.6x105
4.9x10-5
0.81
9.2x102
2.3x10-4
0.74
Monitoring of open circuit potentials and open circuit potential gradients
Open circuit potentials, Eoc (Fig. 11) became more electropositive in time and after about 700 immersion hours reached a constant level, indicating the thickening of the passive layer [109,110], the increase of its stability. The values of Eoc are placed in the passive potential range of Ti, Nb, Zr, and Ta on the Pourbaix diagrams [111] excepting Zr in acid Ringer solution that is in the active dissolution potential range. In acid and alkaline Ringer solutions can be observed more active Eoc values due to the more corrosive properties of these solutions.
The noblest Eoc values were registered in neutral Ringer solution, showing a very good resistance of the those three alloys in normal functional conditions of an implant. The alloy with the highest Nb content, Ti-20Nb-10Zr-5Ta alloy exhibited the most electropositive Eoc values, i.e. the best passivation, confirming the other electrochemical results.
Fig. 11. Monitoring of open circuit potentials for Ti-5Nb-10Zr-5Ta, Ti-10Nb-10Zr-5Ta and
Ti-20Nb-10Zr-5Ta alloys in Ringer solutions at 370C
Open circuit potential gradients, ΔEoc (Table 7) have low values, which can not generate galvanic or local corrosion because only differences of 600-700 mV can initiate and maintain these types of corrosion [112,113].
Table 7 Open circuit potential gradients developed in Ringer solutions
Alloy
Time (h)
ï„Eoc1(pH) (mV)
ï„Eoc2(pH) (mV)
ï„Eoc3(pH) (mV)
Ti-5Nb-10Zr-5Ta
0
-15
-17
-2
500
-19
-35
-16
1000
-48
-42
-4
1500
-45
-53
-8
Ti-10Nb-10Zr-5Ta
0
-70
20
90
500
-29
-6
23
1000
-63
-63
-0.4
1500
-69
-60
9
Ti-20Nb-10Zr-5Ta
0
12
50
38
500
-99
-27
72
1000
-109
-67
42
1500
-116
-76
40
