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Vol. 8. Issue 4.
Pages 3399-3414 (July - August 2019)
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Vol. 8. Issue 4.
Pages 3399-3414 (July - August 2019)
Original Article
DOI: 10.1016/j.jmrt.2019.06.006
Open Access
Investigation of cast and annealed Ti25Nb10Zr alloy as material for orthopedic devices
Raúl Bolmaroa, Anca C. Paraub, Vasile Prunac, Maria A. Surmenevad, Lidia R. Constantinb, Martina Avalosa, Cosmin M. Cotrutd,e, Raluca Tutuianuc, Mariana Braicb, Danut V. Cojocarud, Ioan Danf, Sorin Croitorug, Roman A. Surmenevd, Alina Vladescub,d,
Corresponding author

Corresponding author.
a Instituto de Física Rosario, Laboratorio de Microscopía Electr´nica de Barrido, Centro Científico Tecnológico, Ocampo y Esmeralda, 2000 Rosario, Argentina
b National Institute of Research and Development for Optoelectronics - INOE 2000, 409 Atomistilor St., 077125, Magurele, Romania
c Institute of Cellular Biology and Pathology Nicolae Simionescu of the Romanian Academy, 8 B.P. Hasdeu, Bucharest, Romania
d Physical Materials Science and Composite Materials Centre, National Research Tomsk Polytechnic University, Lenin Avenue 43, Tomsk, 634050, Russia
e University Politehnica of Bucharest, 313 Spl. Independentei, 060042, Bucharest, Romania
f SC R&D Consulting and Services SRL, 45 M. Ghiculeasa St., 023761, Bucharest, Romania
g Tehnomed Impex Co SRL, 1 Sos. Pantelimon, 021591, Bucharest, Romania
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Tables (3)
Table 1. Mechanical characteristics of investigated alloy.
Table 2. Surface potential measurements of the studied samples.
Table 3. Electrochemical parameters of the investigated alloys in Ringer solution with different pH values (corrosion potential — Ecorr, corrosion current density — icorr, polarization resistance — Rp).
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In the present work, we report the preparation of a novel titanium-based alloy, namely Ti25Nb10Zr, by cold crucible levitation melting technique. The cast alloy consists of a complex microstructure with large Beta phase grains (54%, ˜50–150 µm) with a regularly connected net of Alpha′ (orthorhombic, 46%) phase running along boundaries and across the grains and keeping a regular misorientation with respect to the Beta phase. An intermeshed 51% Alpha and 49% Beta phases with lamellar microstructure were found by annealing. The electrochemical tests showed that both alloys were affected by the corrosion process. A good corrosion resistance in SBF at 37 °C was found for the cast form. The cast alloy is more resistant when immersed into solutions with pH2 and pH7, while the annealed one is resistant in pH5 solution. Surface potential of both alloys is negative, with the annealing process leading to a slight decrease of that property. Collectively, the biological results indicate a more favorable viability on cast form as compared to annealed one, suggesting that the cast alloy is promising for biomedical applications.

Corrosion resistance
Surface potential
Full Text

Titanium-based alloys offer a wide variety of properties, adequate for materials used for medical applications, such as good corrosion resistance and biocompatibility inside the human body and a good ductility for processing the complex geometries such as those of medical devices [1–5]. In biomedical applications, 316L stainless steel and Ti-based alloys are well used since many years [1–5]. Unfortunately, each type of material exhibited advantages and disadvantages. For example, 316L steel is well accepted in biomedical field due to its greater anticorrosive properties in a wide range of corrosive environments, smoothness, biocompatibility and cleanability [6]. Firstly, the 316L steel was introduced in the medical field to minimize the apparition of pitting corrosion of the metallic implants after insertion inside the human body. Unfortunately, this material proved in time to be highly susceptible to a localized corrosion, to stress corrosion, cracking and crevice corrosion inside of human body [7]. Despite of this disadvantage, 316L steel demonstrated to be fine for load bearing and fixation devices. Moreover, this alloy derived from the 18-8 stainless steel alloy with high Cr and Ni contents, causing a toxic effect in time. In many cases, 316L is used often for locations in which strength is not required for an extended duration, due to improper mechanical properties. Nevertheless, reported studies estimated that 316L steel is used for production of approximately 60% of the surgical implants used in the United States [8]. For example, Ti alloy is well used for producing of orthopedic hip stems due to its excellent fatigue strength, good density, and its ability to diminish stress shielding. Moreover, the Ti alloys formed an external oxide layer with high corrosion resistance in corrosive media. The main disadvantage is based exactly on one of its advantages – external oxide layer – which breaks down quickly and the formed debris act as a third-body, which accelerates the wear process, leading to the rejection of the implant. Other disadvantage is related to a low elastic modulus when compared with 316L steel, low shear strength and high cost for manufacturing. Despite all these, in orthopedic applications, Ti6Al4V is commonly used [9,10]. In recent years, it was reported that this alloy produced cytotoxic reactions due to the Al and V elements [11] and its replacement with other biocompatible alloys is currently under way. Other problem of Ti6Al4V alloy is its “stress shielding”, because of the difference between its elastic modules (˜114 GPa) and that of the human bone (max 30 GPa). Many efforts were directed to replace Ti6Al4V alloy by others with improved mechanical, anticorrosive, tribological, and biological performances, such as Ti–Nb [4,12], Ti–Zr [13–15], Ti–Ta [16], Ti–Zr–Ta [17], Ti–Zr–Mo [18], Ti–Nb–Zr [19,20], Ti–Zr–Pd–Si [21], Ti–Zr–Mo–Sn [22], Ti–Zr–Ta–Ag [23], Ti–Zr–Cu–Fe [24] etc. The challenge is huge, because the main concern is to develop a novel biocompatible alloy with low elastic modulus (E) close to that of the human bone (E = 2–30 GPa). A survey of the literature showed that alloys of the system Ti–Nb–Zr are the proper choice due to their E value, which can range from 70 to 14 GPa (i.e. 14 GPa for Ti19NbZr; 54 GPa for Ti10Nb9Zr; 62 GPa for Ti13Nb13Zr or Ti34Nb25Zr or Ti5Nb9Zr; 50 GPa for Ti(18-20)Nb(5-6)Zr; 70 GPa for Ti16Nb10Zr). The E value can be reduced by alloying with additional elements. In this manner, the system of Ti–Nb–Zr–X, where X could be Ta [25,26], Fe [27], Mo [28–30], Sn [31], Hf [32], Ta–Si [33], Ta–Si–Fe [34] has been investigated. It was revealed that the reduction of E could be achieved by suppressing the transformation at high temperatures of β phase into α′ [35]. However, additional elements will decrease the E value, but the production cost will increase, what will be reflected in the final cost of the implants. Thus, the best solution is to keep the number of alloying elements at the minimum and to find the best composition for the proposed applications at a decent price. For each alloy, the conditions of heat treatment, such as type, aging temperatures and cooling rates, play an important role in the final properties. Thus, our study aims to investigate the Ti25Nb10Zr alloy as a possible candidate for orthopedic implants. Moreover, the effect of thermal treatment in vacuum on the chemical, structural, morphological, anticorrosive, tribological and biological properties is analyzed.

2Materials and methods2.1Alloys’ preparations

The Ti25Nb10Zr alloy was obtained by cold crucible levitation melting technique (CCLM), using a FIVES CELES — CELLES MP 25 furnace with 25 kW nominal power and a melting capacity of 30 cm3. The thermal annealing treatment of Ti25Nb10Zr was performed in vacuum using the following steps: 3 h with heating rate of about 5 °C/min, 7 h holding time at 930 °C, and further cooling for 2.5 h with 3 °C/min until 300 °C. Ti6Al4V-ELI alloy was used as control material specimen (Bibus Metals AG, Germany). According to manufacturer quality certificate, the alloy was annealed at 730 °C for 1 h, then hot rolled and machined. The samples under investigation in this paper are labeled as follows: Ti6Al4V as control material, Ti25Nb10Zr cast and Ti25Nb10Zr annealed.

2.2Alloy characterization

The elemental composition was examined by energy dispersive X-ray spectroscopy (EDS) using a scanning electron microscope (SEM, TableTop 3030PLUS, Hitachi) equipped with EDX system (Quantax70, Bruker). Electron Back Scattering Diffraction was performed using a FEG FEI Quanta 200 equipment, furnished with an EDAX-OIM-TSL system, a Digiview camera and TSL 7.3 post-processing software.

The mechanical properties of the alloys were evaluated by uniaxial tensile tests using a GATAN MicroTest 2000N module on a Tescan Vega II microscope. The ASTM E8 standard was used to determine the ultimate tensile strength, 0.2% proof stress in elongation, and the ASTM E111 standard for the calculation of initial elastic modulus. Furthermore, the stress-strain relationship is studied based on the Ramberg–Osgood equation, in the elastic region, and their parameters used to determine the tangent and secant modulus. The hardness was measured using a model FM-700 microhardness tester, by applying a force of 50 gf.

In vitro corrosion resistance was studied in Ringer solution with three different pH values (2, 5 and 7) at human body temperature (37 ± 0.5 °C) using a Potentiostat/Galvanostat VersaSTAT3 in the following steps:

  • Open circuit potential for 48 h.

  • Linear polarization resistance (LPR) from −0.01 to +0.01 V with scan rate of 1 mV/s.

  • Tafel curves from −0.25 V to +0.25 V with scan rate of 1 mV/s.

  • Potentiodynamic curve from −1 V to 2 V with scan rate of 1 mV/s.

A saturated calomel electrode (SCE) was used to determine all potentials. The Ringer solution (NaCl = 8.69 g L−1; KCl = 0.30 g L−1; CaCl2 = 0.48 g L−1) was prepared in laboratory using the reagents from Sigma Aldrich (Germany). The values of pH were achieved by adding a solution of HCl 1 Mmol.

Investigation of the surface potential distribution on the surface was performed by Kelvin probe force microscopy (KPFM). The measurements were carried out by means of a NanoScanTech (Dolgoprudny, Russia) Certus Standard V atomic-force microscope (AFM) under normal conditions. Single crystal silicon probes doped antimony, namely Tips Nano NSG30/Au probes, were employed. The KPFM is the two-pass technique used to measure contact potential variations, with an atomic force microscope (AFM), on a nanometer scale. Preparation and examination of the surfaces is done in air. During the first pass (or scanning), the surface profile is recorded under semi-contact conditions. Then, in the second pass, the probe is lifted above the surface at the height dZ. An AC voltage U·sin(ωt) is applied to the probe. This voltage induces oscillations of the probe at its resonance frequency ω. The probe oscillation amplitude is maintained equal to zero through the feedback loop by altering the dc component of the bias voltage U0(x,y).

The contact angle was determined using a KSV-Instruments Attention TL101 system (Biolin Scientific, Stockholm, Sweden) at room temperature (23 ± 1 °C). For measurements, a micro-syringe (Hamilton Company, Reno, NV, USA) was used to weep the liquid (Ringer solution) on the surfaces of the Ti-based alloy.

For cell culture investigation, each sample was first sterilized by keeping for 24 h in 70% v/v ethanol, then washed for several times in free-endotoxin sterile water and kept for other 24 h in culture medium with 10% fetal bovine serum. The biocompatibility assay was carried out with osteoblast-like cells of the human osteosarcoma cell line MG63, which were incubated at 37 ºC in a minimum 95% humidified air atmosphere containing 5% CO2. The cells were seeded at a density of 5000 cells/cm2, being cultured in Dulbecco's Modified Eagle's Medium (DMEM) with 1‰ glucose, 10% heat-inactivated fetal bovine serum, and antibiotics: penicillin (100 U/mL), streptomycin (72 U/mL), and neomycin (50 U/mL). Before sterilization, each sample was precisely measured and the correct volume of cell suspension was calculated and added over the sample, in order to obtain the needed cell density on sample cultured surface [36]. For biocompatibility evaluation, a XTT viability assay (Cell Proliferation Kit II) was performed [37], after a culture period of six days on sample surface. This assay is used to assess cell viability as a function of redox potential. Viable respiring cells convert the XTT (2,3-bis-(2-methoxy-4-nitro-5-sulfophenyl)-2H-tetrazolium5-carboxanilide) to a water-soluble orange formazan. The reduction of XTT is facilitated by incorporation of an electron mediator that scavenges available electrons on the plasma membrane, leading to the formation of a reactive intermediate that reduces XTT to an intensely colored formazan whose optical density was read at 470 nm with a reference at 630 nm. The results were normalized in order to control sample result Ti6Al4V.

MG63 were cultured for 3 and 6 days on the metallic probes as described in the previous section. At the selected time points, a Live/Dead assay was performed using a dedicated kit from Molecular Probes according to the manufacturer’s protocol. Briefly, cells were stained with a solution containing 2 µM calcein AM and 4 µM EthD-1 for 30 min, followed by washing with PBS and mounting with Prolong Gold Antifade Mountant (Invitrogen). For vimentin and type I collagen staining on MG63 cells cultivated for 6 days on the metallic probes, the following protocol was performed: cells were washed with PBS, fixed with 4% PFA and permeabilized with 0.1% Triton X for 15 min at room temperature. After another PBS wash, the samples were incubated with blocking buffer consisting of 1% BSA, for 30 min, followed by anti-vimentin (Sigma-Aldrich) diluted 1:50 and anti-collagen I ( Thermo Fisher Scientific) diluted 1:200, for 1 h. The samples were washed 3 times in PBS, followed by goat anti-mouse conjugated with Alexa Fluor 568 (Thermo Fisher) and goat anti-rabbit conjugated with Alexa Fluor 488 (Thermo Fisher) diluted 1:2000 for 1 h. The samples were washed thrice with PBS and mounted with Fluoroshield with DAPI (Sigma-Aldrich). The images were acquired using Axio Observer Z1 Carl Zeiss microscope.

For each test, it was used more replicates for each structural state of Ti25Nb10Zr alloy and also for reference alloy (Ti6Al4V) for validation of the results, as following:

  • -

    2 replicates for EBSD analysis,

  • -

    3 replicates for mechanical, corrosion and KPFM tests,

  • -

    5 replicates for biomedical assay and contact angle.

3Results3.1Composition and microstructure

The elemental composition of the casted and annealed Ti25Nb10Zr alloy is presented in Fig. 1. It can be observed that both compositions are very close, meaning that the composition is not affected by annealing treatment. In Fig. 1, it can also be seen the distribution of each constituent element of alloy, being obvious that all elements are homogenously distributed on the surface. SEM images on as-cast Ti25Nb10Zr sample showed typical biphasic microstructures apparently consisting of α (light) and β (dark) phases, where polyhedral grains of α and β phases, at the boundaries, can be distinguished (Fig.1). The annealed alloy shows a microstructure with the apparent presence of α-needles mono-phase, with Widmanstätten patterns. However, the analysis by EBSD shows a much richer microstructure and phase composition in both pre- and post-annealing behavior.

Fig. 1.

(a) Ti25Nb10Zr casted and (b) Ti25Nb10Zr annealed.


The samples were studied by EBSD in a FEG-SEM FEI Quanta 200 equipped with EDAX TSL-OIM, DigiView EBSD camera and TSL 7.3 analysis software. The results for regular Ti6Al4V samples are presented as representative for the power of the technique and the paramount differences with the microstructure developed by Ti25Nb10Zr alloy.

A cleanup subroutine accepting 3° misorientation and a minimum of 3 pixels was applied to each scan. Less than 10% of the points were re-indexed. Ti6Al4V sample shows a slightly elongated microstructure of Alpha (93%, 9 µm2 area size) and Beta phases (7%, 1.6 µm2 area size). That is only self-evident by the distribution of the Beta phase (Fig. 2). Textures are typical of round swaged Ti samples (Fig. 2). Scans of 120 µm × 100 µm consist of 1.368.578 points separated by a step size of 100 nm in a hexagonal arrangement. Average confidence index CI = 0.26 and average image quality IQ = 3575 for Alpha phase and CI = 0.24 and IQ = 2902 for Beta phase were obtained. Texture is typical for swaged bars.

Fig. 2.

Ti6Al4V. Inverse Pole figure maps and Pole figures for (a) Alpha phase and (b) Beta phase. Scans of 120 µm × 100 µm comprise 1.368.578 points separated by a step size of 100 nm in hexagonal arrays. Average confidence index CI = 0.26 and average image quality IQ = 3575 for Alpha phase and CI = 0.24 and IQ = 2902 for Beta phase were obtained. Textures are typical of swaged bars.


The cast Ti25Nb10Zr alloy shows a very complex microstructure with large Beta phase grains (54%, ˜50–150 µm) with a regularly connected net of Alpha′ (orthorhombic, 46%) phase running along boundaries and across the grains and keeping a regular misorientation with respect to the Beta phase (Fig. 3). Scans of 120 µm × 150 µm consist of 2.079.693 points separated by a step size of 100 nm in a hexagonal arrangement. Average confidence index CI = 0.15 and average image quality IQ = 2791 for Alpha′ phase (orthorhombic) and CI = 0.41 and IQ = 3783 for Beta phase were obtained. The pole figures show the few large Beta phase grains present in the scan and the many small orthorhombic crystals keeping a crystallographic relationship with the Beta phase, with the Beta <001> and <011> directions with which the <011> and <001> directions of the Orthorhombic Alpha′ phase stay aligned, respectively.

Fig. 3.

Casted Ti25Nb10Zr alloy. Inverse pole figure maps and pole figures for (a) Alpha′ (orthorhombic) phase and (b) Beta phase. Scans of 120 µm × 150 µm comprise 2.079.693 points separated by a step size of 100 nm in hexagonal arrays. Average confidence index CI = 0.15 and average image quality IQ = 2791 for Alpha′ phase (orthorhombic) and CI = 0.41 and IQ = 3783 for Beta phase were obtained. Pole figures are not representative of the texture but only presented to show the orientation correlation between the crystal orientations of both phases.


The annealed Ti25Nb10Zr sample shows an intermeshed 51% Alpha and 49% Beta phases in an acicular microstructure (Fig. 4). Scans of 120 µm × 100 µm consist of 1.368.578 points separated by a step size of 100 nm in a hexagonal arrangement. Average confidence index CI = 0.18 and average image quality IQ = 3159 for Alpha phase and CI = 0.23 and IQ = 2883 for Beta phase were obtained. Alpha phase was previously absent in the cast sample and the whole sample completely changed the microstructure with the Alpha and Beta needles growing probably from an unstable microstructure composed of Alpha′ and Beta phases that were present in approximately equal amounts.

Fig. 4.

Annealed Ti25Nb10Zr alloy. Inverse Pole figure maps and Pole figures for (a) Alpha phase (hexagonal) and (b) Beta phase (BCC). Scans of 120 µm × 100 µm comprise 1.368.578 points separated by a step size of 100 nm in hexagonal arrays. Average confidence index CI = 0.18 and average image quality IQ = 3159 for Alpha phase and CI = 0.23 and IQ = 2883 for Beta phase were obtained. Pole figures are only presented to show the orientation correlation between the hexagonal and cubic phases.


Pole figures for cast and annealed Ti25Nb10Zr were not representative of textures, they were only meant to show the correlation between phases’ crystalline orientations. Grain sizes in Fig. 3 were close to 100 µm for BCC phase, which are so large that only neutron diffraction could be used. Despite the Orthorhombic phase shows very small crystals they are firmly correlated with the underlying BCC crystals. Fig. 4 shows many rather small needles. However, BCC phase shows a single orientation for all needles comprising more than 100 µm across the scan. The orthorhombic phase shows more orientations but all of them correlated through the characteristic martensitic BCC → Alpha′ phase transformation. In this case, also the grain size is beyond the capabilities of X-ray diffraction texture determination.

3.2Mechanical properties

The mechanical properties (Young modulus, ultimate tensile strength, yield strength — 0.2% proof stress, elongation to fracture) of the cast and annealed alloy are presented in Table 1.

Table 1.

Mechanical characteristics of investigated alloy.

Mechanical properties  Ti25Nb10Zr casted  Ti25Nb10Zr annealed  Ti6Al4V[Refs.] 
Ultimate tensile strength σUTS (MPa)  786 ± 26.8  1171 ± 39.3  895–1300 [1,3,5,9–11,15] 
0.2 yield strength σ0.2 (MPa)  542 ± 18.4  541 ± 18.8  830–1100 [1,3,5,9–11,15] 
Elongation to fracture εf (%)  36 ± 1.4  10 ± 0.3  10 [1,3,5,9–11,15] 
Elastic modulus E (GPa)  65 ± 1.8  60 ± 2.1  100–120 [1,3,5,9–11,15] 

The experimental stress-strain curves of both cast and annealed alloy are illustrated in Fig.5. In the literature, Ti-based alloys used for biomedical applications exhibited elastic moduli and ultimate tensile strength ranged from 55 MPa to 85 GPa and from 596 MPa to 1080 MPa, respectively. For example, the Ti35.3Nb5.1Ta7.1Zr presented low elastic modulus (55 GPa) and low ultimate tensile strength (596.7 MPa) compared to other types of Ti-based alloys [12]. In our case, cast alloy exhibited an elastic modulus of 65 ± 1.8 GPa and 60 ± 2.1 GPa after annealing, values rather between the spread limits one can expect for the measurement method. Despite the completely different microstructures and phase compositions of as cast and annealed samples, the elastic properties are slightly changed. Between cast and annealed states, the largest modification is the substitution of Alpha′ phase by Alpha phase, keeping more or less the same proportion of Beta phase. In first approximation elastic properties are quite microstructure insensitive. The main effect contributing to elastic properties is the strength of the interatomic bond between atoms. If the composition and compactness does not significantly change between phases, the expected variation for average elastic constants is rather low. A very low variation is confirmed indeed by our results [38]. Elastic modulus of human bones is ranged up to 20 GPa, much below the measured values, but the currently studied alloy is closer to the needed values.

Fig. 5.

Stress–strain curves of casted and annealed alloy (the presented results are only for one of replicate).


Ultimate tensile strength of 786 MPa increased after annealing to 1171 MPa. This property is appreciated in many applications, but in biomedical field it is not necessary to be so high. A human tibia or a femur of a 20–39 years old person has an ultimate tensile strength of 174 MPa and 124 MPa, respectively [39]. It is true that a bone is approximately 20% stronger for brisk walking than for slow walking, but in our opinion, a high ultimate tensile strength is not essential. Before failure, the cortical bone sustains greater stress, but less strain. For example, before failing in clinical practice, the cancellous bone can tolerate strains of 75%, while cortical bone will fracture when the strain exceeds 2% [40]. Probably this effect can be related to its structure, because the cancellous bone is porous and filled with blood, marrow, and body fluid. The elongation of the alloy before breaking is lower for annealed alloy, which is a positive result. For example, elongation of human tibia or femur for a 20–39 years old person is 1.50 and 1.41 %, respectively [39].

To explain this behaviour of both as-cast and annealed samples, one must consider the following two mechanisms. First mechanism refers to the influence of constituent phase fractions, a higher phase fraction of a certain phase having a higher influence upon properties. One must consider this aspect and the mechanical resistance properties (ultimate tensile strength and yield strength) of all involved phases: α′-Ti phase exhibits much lower mechanical properties in comparison with β-Ti phase, and also, both α′-Ti and β-Ti phases exhibits much lower mechanical properties in comparison with α-Ti phase. In the case of ductility (elongation to fracture) an inverse behaviour was reported, α′-Ti phase being more ductile in comparison with both β-Ti and α-Ti phases [41–44]. In this study, for as-cast microstructural state a balance of 46% α′-Ti / 54% β-Ti phase fraction was found, while in the case of annealed microstructural state a balance of 51% α-Ti / 49% β-Ti phase fraction. The balance of constituent phase fraction must be taken into account, playing an important role in final exhibited mechanical properties. The second mechanism refers to the influence of grain-size of constituent phases, being known that the size of crystalline grains influences mechanical properties (expressed by the Hall–Petch correlation [45,46]; a smaller grain size leads to increasing mechanical resistance properties). In our case, if one considers the influence of grain-size, for all α′-Ti, β-Ti and α-Ti phases, it can be expected that higher mechanical resistance properties (ultimate tensile strength and yield strength) will be obtained in as-cast state in comparison with annealed state.

The shape of the strain–stress curve of both microstructural states revealed significant differences in terms of mechanical resistance and ductility properties. In addition, one can observe a linear elastic behaviour, with a constant elastic modulus, for both as-cast and annealed samples.

3.3Surface potential

The typical images of the surface potential distribution over the surface obtained via KPFM are presented in Fig. 6. The results of KFPM, measured on different areas, are presented in Table 2. It can be observed that the surface potential of all the samples is negative. It is also revealed that annealing results are in the slight decrease of the surface potential, besides the case of Ti6Al4V alloy, where the surface potential remains almost unchanged. In the case of Ti25Nb10Zr alloy, the cast alloy has a lower surface potential average value than Ti6Al4V alloy. After annealing, the surface potential decreased significantly, regardless of the investigated area.

Fig. 6.

Surface potential images of the developed samples (30 × 30µm2): (a) Ti6Al4V, (b) Ti25Nb10Zr cast, (c) Ti25Nb10Zr annealed.

Table 2.

Surface potential measurements of the studied samples.

SampleSurface potential (V)
30 × 30µm2  15 × 15µm2  5 × 5µm2 
Ti6Al4V  −1.467 ± 0.124  −1.206 ± 0.109  −0.752 ± 0.047 
Ti25Nb10Zr casted  −1.996 ± 0.165  −1.419 ± 0.117  −0.887 ± 0.52 
Ti25Nb10Zr annealed  −2.625 ± 0.258  −2.190 ± 0.184  −1.190 ± 0.83 
3.4Contact angle

For the present paper, the contact angle investigation was performed using the Ringer solution with different pH values of 2, 5, and 7. Fig. 7 shows the evolution of contact angle with the pH values of Ringer solution. It can be seen that the contact angle increases for both cast and annealed Ti25Nb10Zr alloy. It can be observed that after annealing, the contact angle decreased significantly, tending to values of a more hydrophilic surface. At the acidic pH, the surface has more hydrophilic character. During the years, it was demonstrated that the implants are rapidly accepted by the human body if its surface is hydrophilic [47]. For all the investigated alloys, an increase in contact angle was found for solution with pH5 and pH7. More evident, it is observed in the case of annealed Ti25Nb10Zr alloy.

Fig. 7.

Contact angle vs pH values of Ringer solution (results are mean of five replicates).

3.5In vitro corrosion behaviour

Corrosion resistance is an important property for materials used in biomedical applications. This is a process, which can be used to predict the degradation of a surface imitating the conditions of the human body. In this study, the corrosion tests were performed in Ringer solution with different pH values of 2, 5 and 7, in order to establish the influence on corrosion behavior. It was reported that immediately after surgery, an implant is surrounded by fluids containing fibrin and chloride ions of relative low pH (5 or lower till 2), because hydrogen concentration increases in the traumatized tissues [48]. Then, an inflammatory reaction will take place and the pH will increase till 7 or 9 [49].

The potentiodynamic curves performed at three pH values are presented in Fig. 8. Based on steps of corrosion tests, the main electrochemical parameters, which are calculated using LPR and Tafel extrapolation [50–53], are presented in Table 3. The corrosion behavior of the investigated alloys can be explained by the following criteria:

  • 1

    Corrosion potential (Ecorr): surface resistant to corrosion means more electropositive potential. According to our results, it can be seen that the cast alloy is more resistant in solution with pH2. In solution with pH5, the Ecor values were not so different, meaning that both alloys behave more or less the same. In solution with pH7, the annealed alloy has a more electropositive Ecor, indicating the presence of more resistant, stable passive films probably due to the passivation effects of the Zr and Nb elements. A related point to be taken into consideration is that the Ecor values followed a similar trend to that of EOC. The values of corrosion potential of both alloys are placed in the passive potential range of Ti, Zr and Nb of the Pourbaix diagrams, indicating that the surface of the alloy is stable and resistant to corrosive solution.

  • 2

    Corrosion current density (icor): surface resistant to corrosion has low icor value. In a solution with pH2 and pH7, the cast alloy exhibited low icor value, while in pH5 both alloys have an appropriate value.

  • 3

    Polarization resistance (Rp): high Rp means high corrosion resistance. Also according to this criterion, the cast alloy is more resistant into solution with pH2, followed by annealed alloy in solution with pH5 and cast alloy in solution with pH7. The obvious differences of corrosion results in solution with pH7 can be related to the roughness (discussed below).

Fig. 8.

Potentiodynamic curves of investigated alloy in Ringer solution with pH2, 5 and 7 (the presented graph are only for one of replicate).

Table 3.

Electrochemical parameters of the investigated alloys in Ringer solution with different pH values (corrosion potential — Ecorr, corrosion current density — icorr, polarization resistance — Rp).

Sample  pH of Ringer solution  Ecorr (mV)  icorr (μA/cm2Rp (kΩ × cm2
Casted  2−28.5 ± 1.4  0.286 ± 0.038  304.57 ± 14.24 
Annealed  −66.8 ± 2.8  0.304 ± 0.043  267.61 ± 10.26 
Casted  5−114.0 ± 2.7  0.634 ± 0.052  235.69 ± 9.57 
Annealed  −103.6 ± 2.9  0.565 ± 0.048  346.41 ± 21.89 
Casted  7−246.7 ± 5.7  0.113 ± 0.015  764.23 ± 19.97 
Annealed  −197.7 ± 3.2  0.131 ± 0.018  209.16 ± 9.54 

According to the potentiodynamic curves, above the Ecorr, the current density values for both alloys, investigated in solution with pH2 and pH5, continued to be constant until 2 V from ≈0.7 V. In general, the constant current region is associated to the passive region.

During the real life inside of the human body, a pH value of an implant surface can be different from one zone to another, leading to potential gradients, which will produce galvanic corrosion and accelerate the corrosion process. This phenomenon takes place because the fluids wets the implant surface non-uniformly. In order to evaluate this phenomenon, we calculated the open potential gradients of the investigated alloys:

ΔEOC1= EocpH=2 − EocpH=5: casted = 0.046 V; annealed = 0.026 V;
ΔEOC2= EocpH=2 − EocpH=7: casted = 0.260 V; annealed = 0.014 V;
ΔEOC3= EocpH=5 − EocpH=7: casted = 0.213 V; annealed=-0.012 V;

For annealed alloy, the differences are small, while for cast alloy were found to be of maximum 0.26 V, but there was still a small value according to the literature, in which it was reported that the value higher than 0.7 V can induce and sustain the galvanic cell [54,55]. To summarize the corrosion test results, we can say that the cast alloy can be more appropriate for the use in biomedical field than annealed alloy and it cannot cause galvanic or local corrosion reaction, even if the implant surface exhibited a large pH difference in various zones.

In order to evaluate the damage of the surface after the electrochemical tests, a comparison between the roughness, before and after corrosion tests, was performed (Fig. 9). Before corrosion tests, each sample was polished to the same roughness (around 100 nm), in order to remove this parameter from the list of factors, which can influence the corrosion process. It can be observed that both alloys were affected by the corrosion process. The solution with pH2 does not affect the roughness of the surface too much, indicating that the alloys are resistant in acidic medium. In the case of the surface immersed in solution with pH5, the roughness of both surfaces increased compared with the initial values, indicating that the surface is affected by the corrosive attack. After the test in solution with pH7, the roughness of both alloys were significantly modified, with values for cast alloy that significantly increased, meaning that the surface has many peaks and for the annealed alloy the values decreased, indicating that the surface has many valleys, which were formed during the tests and therefore the corrosion process on the surface was nonhomogeneous. The surface of both alloys is more deteriorated by the solution with pH7. These results are in good agreement with the data obtained during the electrochemical measurements (Fig. 10 and Table 3).

Fig. 9.

Roughness (Ra) values of investigated alloys before and after electrochemical tests.

Fig. 10.

Viability results of the investigated alloys. The results of casted and annealed alloys were normalized to control sample result (Ti6Al4V).

3.6In vitro biological assay

In vitro biological investigations are the first step for the prediction of the possibility that a material can or cannot be used as a biomaterial. It is true that in vitro results can differ from in vivo material behavior, but they can help us know if we can perform further in vivo tests, which are more expensive than those in vitro.

According to XTT results, it can be seen that the samples promote the cell culture grow at a good level. Further tests are needed in order to evaluate other biological properties of these materials. To conclude the biological results, we can say that the developed cast and annealed alloys can be good candidates for medical applications.

MG63 cells adhered and proliferate on the casted Ti25Nb10Zr and annealed Ti25Nb10Zr samples with no obvious difference compared to Ti6Al4V as reflected by the images at 3 and 6 days (Fig. 11). Furthermore, red colored spots corresponding to dead cells were insignificant on all samples, indicating that the alloys do not interfere with the MG63 cells survival.

Fig. 11.

Fluorescence images of MG63 cells grown on TiAlV, TNZ and TNZtt coated sampled stained using the Live/Dead kit assay, showing viable-green and dead-red cells.


MG63 cellular morphology as reflected by the intermediate filament protein vimentin was evidenced by immunoflourescence at 6 days post-seeding (Fig. 12), showing that the cells on both alloys maintained a similar vimentin distribution compared to Ti6Al4V. Considering that type I collagen is an important component of the bone tissue and impairment of its secretion is often associated with pathological manifestations, we evaluated the collagen I expression by immunofluorescence. MG63 cells grown on all coated samples stained positively for type I collagen (Fig. 12). These results suggest that both casted and annealed Ti25Nb10Zr alloys support the survival of MG63 cells without affecting cellular morphology or type I collagen expression.

Fig. 12.

Fluorescence images of MG63 cells grown on TiAlV, TNZ and TNZ tt for 6 days and double immunolabeled for vimentin (red) and type I collagen (green). Cells nuclei were stained with DAPI (blue).


Over the past years, the ageing of populations increased and the demand for implant has intensely grown, because people want to keep their life at the same level of activity as ever. Therefore, the demand for implants with high-performance characteristics has been amplified in many fields of medicine such as cardiology, orthopedics, trauma, neurosurgery, urology, plastic surgery, spine, dental, and wound care. In many of such areas, the metallic biomaterials are more adequate than other types of biomaterials due to their characteristics. In spite of their advantages, the clinical use of metallic biomaterials leads to different complications such as poor osseo-integration, high inflammation responds, mechanical instability, necrosis, and infections, leading to prolonged patient care and pain. For this reason, the biomaterials science is still a domain under development with huge possibility to offer a diversity of biomaterials processed by various methods. Currently on the market, there is a wide variety of biomaterials, which is an advantage for the patients who can select the material that meets the precise purposes of their treatment.

For fracture fixation, angioplasty and bone remodeling, the metallic alloys are commonly used such as stainless steel (316L), cobalt–chromium (CoCr) alloys, and titanium alloys, due to their long-term stability under corrosive in vivo conditions and good mechanical properties. Ti6Al4V alloy is the most used alloy form Ti alloy group. In the last years, it was demonstrated that this alloy has some toxic reaction and many companies try to replace it with other alloys, which are more adequate for these applications. In this paper, we prepare a novel Ti alloy, alloyed with biocompatible metals, such as Ti and Zr for orthopedic devices manufacturing. We chose a composition with 25% Nb, 10% Zr and Ti as balance, based on the previous experience of the group in biomaterials design. The alloy was prepared by the cold crucible levitation melting technique. We also modify the microstructure of the alloy by the thermal annealing treatment in vacuum at 930 °C, in order to achieve the requirements imposed by the medical applications (such as density, thermal conductivity, thermal expansion, radiolucency, malleability and corrosion, as well as biocompatibility and toxicity are thoroughly considered). A stable crystalline structure reduces the risk of crack propagation and failure in the implant. For example, a solid solution in face-centered crystal structure confers easy shaping and bending and higher ductility relative to the body centered cubic crystal structure. Moreover, in the case of 316L steel it was reported that an optimum “fiber texture” with fibers elongated into fibrous or spindle shapes and parallel to the long axis of the implant accepts high deforming forces imposed on implant due to various anatomical place of its [56]. Furthermore, implants made of annealed alloys retain better ductility than casted implants, which typically possess lower strengths. Manufacture prefers casted alloys just for significant economical savings. The structural characteristics of metals and alloys make them suitable candidates for many biomedical applications such as load-bearing orthopedic, but all metallic materials are susceptible to chemical attack or corrosion from reacting with human physiologic media surrounding an implant. Usually, the degradation of implant is not only related to its structural integrity, may also outcome in a general in vivo reaction. Thus, biomaterials used for biomedical applications should have a good resistance to oxidation and high chemical stress resistance and high protection to corrosive attack.

As it was expected, the microstructure is different: the cast alloy showed a very complex microstructure with large Beta phase grains (54%, ˜50–150 µm) with a regularly connected net of Alpha´ (orthorhombic, 46%) phase running along boundaries and across the grains and keeping a regular misorientation with respect to the Beta phase. The annealed alloy exhibited an intermeshed 51% Alpha and 49% Beta phases in an acicular microstructure.

The mechanical properties are not so different: small differences were found for elastic and σ0.2 yield strength, while significant differences were seen for ultimate tensile strength and elongation to fracture, but the values still remain close to other biocompatible alloys. Compared with biocompatible Ti6Al4V alloy, the advantage of this alloy consists in its constitutive elements, which are biocompatible and will not produce toxic or allergic effects along with sufficient mechanical properties.

The experimental part showed large differences in the corrosion behavior of these two forms of Ti25Nb10Zr alloy. The cast alloy proved to be more resistant to corrosive solution with pH2 and pH5, while in pH7 it suffered a degradation process, while annealed alloy was affected by all solutions regardless of the pH value. These results cannot be related to the roughness effect because both alloys were polished to the same values of roughness. In this case, we believe that the good corrosion resistance of cast alloy is due to the wettability properties; the cast alloy tends to be more hydrophobic than the annealed one. In the literature, it was reported that the hydrophobic surfaces are more resistant to corrosion than the hydrophilic ones, due to their low surface energy [57–64]. He et al. reported that the probability of adsorption of the Cl aggressive species on the solid surface is reduced by increasing the water angle [65]. Thus, we can say that the cast alloy, which has higher contact angle, has a high anticorrosive protection, due to the reduced Cl absorption.

We believe that the good corrosion resistance of cast alloy can also be due to the beta phase, which is a bit higher than that of the annealed alloy. There is an old study of Nishimura and Hiramatsu who reported that the alloy with beta phase has a slightly better corrosion resistance than those with alfa phase, which depend on the nature and concentration of the alloying elements [66]. Moreover, Davis investigated the relation between the corrosion of different alloys and their various microstructures and he found that the alfa phase is more active with less corrosion resistance in corrosive environments compared with beta phase, which is noble [67]. Based on these published results, we consider that the good corrosion performance of cast alloy can be related to the amount of beta phase from this alloy. According to the emf series (electromotive force series), the Ti–Nb–Zr system forms a strong cell with potential differences (4.26 V vs SHE), greater than that found for Ti6Al4V (3.545 V vs SHE), possible origin of its better corrosion resistance. In a galvanic couple, the metal with more negative value in emf series, considered as anode, will be firstly corroded. The metals with more electropositive potential in the emf series have the highest corrosion resistance and the corrosive solution should have high oxidizing power in order to corrode them.

As biological tests are the most important evaluation for the development of new biomaterials, we can say, according to our results, that both cast and annealed alloys are suitable for further investigation in order to develop new orthopedic biomaterials. We did not notice significant changes in the viability results between cast and annealed alloys.

Tehnomed S.R.L. company manufactured few medical devices prototypes made of cast Ti25Nb10Zr (Fig. 13) such as plate for osteosynthesis, screw and screwdriver. This alloy was easily formable to produce some implantable devices with different shapes and geometry. The screwdriver was designed just to fasten the screws without deteriorating their surface during the fastening process. In case of release of some particles from the screw or screwdriver, the intent is to have the same material presents as debris.

Fig. 13.

Prototypes made of casted Ti25Nb10Zr: plate for osteosynthesis, screw and screwdriver.


A novel titanium based alloy with the composition of 25% Nb, 10% Zr and Ti as balanced (Ti25Nb10Zr) was prepared by cold crucible levitation melting technique and annealed in vacuum at 930 °C, in order to be used in biomedical applications for orthopedic devices (implants, screws, plates, etc.). The experimental results showed:

  • The cast alloy shows a complex microstructure with large Beta phase grains, (54%, ˜50–150 µm) with a regularly connected net of Alpha′ (orthorhombic, 46%) phase running along boundaries and across the grains and keeping a regular misorientation with respect to the Beta phase. By annealing, an intermeshed 53% Alpha and 47% Beta phases with acicular microstructure were found.

  • The cast alloy, after polarization studies in SBF with pH2, exhibited an un-attacked surface, indicating a high corrosion resistance of the alloy in corrosive solutions. After the polarization studies in pH5, a slightly attacked morphology was detected. The annealed one is resistant in solution with pH5.

  • Surface potential on both alloys is negative, but the annealing process leads to a slight decrease of the surface potential.

  • After annealing, the contact angle significantly decreased, tending to a more hydrophilic surface. In fact, the hydrophilic increment for all used experimental pH seems to be the only variable definitely favouring the annealed alloy instead of the cast one. However, this high hydrophilic behaviour can also favour corrosion, although marginally.

  • The elastoplastic behaviour of the tensile curve of cast alloy presents a linear elastic behaviour with a constant elastic modulus. For the annealed alloy, one can see a strain-hardening elastoplastic behaviour, indicating a high plastic deformation in the alloy.

  • Cast and annealed alloys differ only in the microstructure. Despite the fact that those differences are quite noticeable, the final influence on many of the biocompatibility assays seems to be marginal. Biological assay showed that both cast and annealed alloys are suitable for the development of new orthopaedic devices.

Conflicts of interest

The authors declare no conflicts of interest.


The present work was supported under two grants of the Romanian National Authority for Scientific Research, CNCSUEFISCDI:

  • Project number PN-II-PT-PCCA-2014-212 (OSSEOPROMOTE), within PNCDI III;

  • Project number COFUND-ERANET-RUS-PLUS-CoatDegraBac (no. 68/2018), within PNCDI III.

The EDS, SEM and XRD results were acquired using the systems purchased by the Sectorial Operational Programme “Increase of economic Competitiveness”, ID 1799/SMIS 48589/2015. A part of work is also suported by PROINSTITUTIO Project – contract no. 19PFE/17.10.2018.

EBSD results were obtained at the Laboratory of Scanning Microscopy in Centro Científico y Tecnológico Rosario, Argentina. The authors acknowledge the collaboration of Eng. Pablo Risso and Biol. Vanina Tartalini for their contribution on sample preparation.

The surface potential measurements (KPFM) were financially supported by the Russian Science Foundation (15-13-00043).

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