journal of the mechanical behavior of biomedical materials 32 (2014) 166 –176

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Research Paper

MRI compatible Nb–Ta–Zr alloys used for vascular stents: Optimization for mechanical properties Hui-Zhe Li, Jian Xun Shenyang National Laboratory for Materials Science, Institute of Metal Research, Chinese Academy of Sciences, 72 Wenhua Road, Shenyang 110016, China

ar t ic l e in f o

abs tra ct

Article history:

With the increased usage of magnetic resonance imaging (MRI) as a diagnostic tool in

Received 7 September 2013

clinic, the currently-used metals for vascular stents, such as 316L stainless steel (SS), Co–Cr

Received in revised form

alloys and Ni–Ti alloys, are challenged by their unsatisfactory MRI compatibility, due to

5 December 2013

their constituents containing ferromagnetic elements. To provide more MRI compatible

Accepted 14 December 2013

vascular stents, the Nb–xTa–2Zr (30 rx r70) series alloys were selected in the current

Available online 27 December 2013

work. Several key properties of these alloys were optimized in terms of stent requirements,

Keywords:

including magnetic susceptibility, elastic modulus and tensile properties. In the as-cast

Vascular stent

state, a single-phase solid solution with bcc structure was formed in the alloys. The

MRI compatibility

volume magnetic susceptibility (χv) and Young0 s modulus (E) of the alloys scaled linearly

Tensile properties

with the Ta content. Increasing the Ta content gave rise to the decreased χv and the

Niobium

increased E, together with the elevated yield strength but less-changed elongation. From

Tantalum

multiple requirements for the stents, the Nb–60Ta–2Zr alloy exhibits an optimal properties, including the χv of about 3% of the 316L SS, the E of 142 GPa superior to pure niobium, high mass density of 12.03 g/cm3 favored to the X-ray visibility, yield strength of 330 MPa comparable to the 316L SS and a elongation of 24%. These remarkable advantages make it quite promising as a new candidate of stent metals. & 2013 Elsevier Ltd. All rights reserved.

1.

Introduction

The vascular stent is a tiny, expandable metallic mesh tube. It is implanted in the coronary or peripheral arteries of humans to reduce blockages and keep blood flowing smoothly. Percutaneous coronary intervention (PCI) treatment with implanted stents has been a conventional approach in clinic. Reliability of stents is directly determined by the material properties from several aspects (Hanawa, 2009; Mani et al., 2007; O0 Brien and Carroll, 2009). The materials used for stents should possess excellent corrosion resistance and n

Corresponding author. Tel.: þ86 24 23971950; fax: þ86 24 23971215. E-mail address: [email protected] (J. Xu).

1751-6161/$ - see front matter & 2013 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.jmbbm.2013.12.015

mechanical properties, such as high radial hoop strength, sufficient ductility, low yield ratio and good fatigue resistance. The currently-used materials for vascular stents fabrication typically include the 316L stainless steel (SS), cobalt– chromium alloys, pure tantalum and nickel–titanium shape memory alloys. Nickel, chromium or molybdenum ions can be eluted from the stainless steel and Co–Cr alloy stents. The action of blood, saline, proteins, and mechanical stress promotes the release of these ions, thus, which causes allergic reactions and also induces intimal hyperplasia and in-stent restenosis (Koster et al., 2000).

journal of the mechanical behavior of biomedical materials 32 (2014) 166 –176

The magnetic resonance imaging (MRI) compatibility of stent materials has become more and more important due to the increased usage of the MRI as a diagnostic tool in clinic. This imaging technique possesses remarkable advantages that show high contrast of human tissues, and it is noninvasive. Unlike other X-ray imaging that exposes the patient to ionizing radiation, the MRI does not bring about adverse effects caused by the iodine-based contrast agents. Thus, the MRI examination is more desirable in imaging diagnosis, especially for diagnosing the brain, spine, joints and ligaments. However, the currently-used stent metals significantly lack MRI compatibility (Hug et al., 2000; Lenhart et al., 2000; Shellock and Shellock, 1999; Teitelbaum et al., 1988). Most of them can be magnetized in the intense magnetic field, such as 1.5 T, due to their ferromagnetic or paramagnetic elements, such as Fe, Ni, Co and Cr. Under this intense magnetic field, the MRI artifacts inevitably appear in some extent, which is induced by localized distortion of the applied magnetic field, arising from the difference in the magnetic susceptibility between the stent metal and the examined tissues, characterized by volume magnetic susceptibility (Δχ V ¼ χ V  χ Vtissue ). Also, the artifact effect is dependent upon the shape and size of stents as well as of the nature of applied magnetic field (Schenck, 1996). The magnetically-induced displacement force, torque and generated heating on the stents under the magnetic resonance environment are not critical to damage the blood vessels. However, the artifacts usually distort and enlarge the outline of the stents in MRI. This generally results in inaccuracies during image reading, and even causes a misleading diagnosis, particularly for the postoperative assessment of arterystenosis. Therefore, lower magnetic susceptibility could ensure better MRI compatibility and also reduce the artifact size, which is a pressing need for future stent metals. In the current work, an effort was made to design a new alloy based on the metals of Nb, Ta and Zr. These metals were selected because of their low magnetic susceptibility (Schenck, 1996), suitable mechanical properties, excellent corrosion resistance and biocompatibility (Branzoi et al., 2008; Eisenbarth et al., 2004; Matsuno et al., 2001; MetikosHuković et al., 2003). As indicated by Vaishnav et al. (1994), pure Ta has been used to make the Wiktor tantalum wire coronary stent, which has been proven to be a valid means of treating acute complications during angioplasty in clinic. On the other hand, the Ta with a large mass density (ρ¼ 16.6 g/cm3) is supposed to have a high radiopacity, thus providing a better X-ray visibility for locating and deploying the stents under Xray fluoroscopy. The highly refractive nature of tantalum, with a high melting temperature, would have created significant fabrication challenges from alloy development to stent electropolishing (O0 Brien et al., 2008a). In addition, the Nb stents were also used in the coronary arteries of patients (Beier et al., 2006). These stents showed no significant difference in contrast to the stainless steel stents, in terms of the minimal lumen diameter, percent diameter stenosis, late lumen loss and the rate of major cardiac adverse events at 6 months. However, there was a propensity towards more neointimal in-growth in the Nb stent group. This may be attributed to the larger strut

167

thickness (100 μm) of these stents due to their relatively lower strength in contrast to the stainless steel stents with a strut thickness of 85 μm. Besides biocompatibility and image-resolvability, the mechanical properties of an alloy are vitally important. High strength is necessary for the stent to resist the force imposed on it by the vascular walls. This high strength allows the stents0 struts thinner to improve flexibility and to reduce cross-sectional profiles so that the stent can easily travel through the narrow vessels (O0 Brien and Carroll, 2009). Based on clinical follow-ups, implantation of a stent with thinner struts may significantly reduce the probability of postoperative restenosis, as the thinner struts produce weaker vascular trauma, accompanied by a significant reduction in neointimal hyperplasia (Kastrati et al., 2001; Schwartz et al., 1992). The stent made of 316L SS keeps the strut thickness at a level of 120–140 μm, while the strut thickness of Co–Cr alloy is reduced down to 80–90 μm, owing to its high elastic modulus and yield strength. Thus, the stent made of 316L SS yields an increased rate of restenosis in contrast to the Co–Cr alloy. Several studies have concluded that (Briguori et al., 2002; Rittersma et al., 2004) a thinner strut significantly promotes the reduction of restenosis events. With all the above considerations, zirconium was selected as an alloying element to play a role of solid-solution strengthening in the binary Nb– Ta alloys, because of its solid solubility of  9 at% in Ta and a large difference in atomic radius (  8%) between Zr and Ta. High elastic modulus and good ductility are also necessary for the alloy, to ensure minimal elastic recoil and reliable expandability. The Young0 s modulus of Nb and Ta is about 103 GPa and 185 GPa, respectively. Nb and Ta form a complete solid solution with a body-centered cubic structure. Finally, combining the above-mentioned multiple factors, the Nb–Ta– Zr ternary alloys were chosen to address their feasibility and advantage as a new alloy for the fabrication of MRI compatible stent. To this end, the microstructure and composition of Nb–xTa–2Zr (30r xr 70) alloys are optimized for magnetic susceptibility, elastic modulus and tensile properties.

2.

Materials and methods

2.1.

Alloy preparation

A series of alloys with nominal composition of Nb–xTa–2Zr (x¼ 30, 40, 50, 60, 70 in wt%) were prepared using element pieces of Nb, Ta and Zr with purity better than 99.9 wt% as starting materials. The alloy ingots weighing about 75 g were prepared by arc-melting under a Ti-gettered argon atmosphere in a water-cooled copper hearth. Each ingot was remelted and flipped several times to ensure compositional homogeneity. A additional group of Nb–50Ta–yZr (y ¼2, 4, 6 wt %) alloy were also fabricated with the same approach. The Nb–60Ta–2Zr alloy was chosen to prepare the largesize ingot with a diameter of 90 mm and a height of 240 mm using electron beam melting (EBM). Before the subsequent run of melting, the surface oxide layer of the last-run melted ingot was removed. After coating with glassy protective lubricant for forge processing, the ingot was heated at 1100 1C for 1 h, then forged into a green body of 117 mm in

168

journal of the mechanical behavior of biomedical materials 32 (2014) 166 –176

width, 40 mm in height and 270 mm in length. Subsequently, the alloy piece cut from the forged bulk was annealed at 1300 1C for 1 h in a furnace under flowing argon atmosphere. Final chemical composition of the bulk alloy was examined using the chips taken from several representative locations in the bulk. It was determined to be Nb–61.6Ta–1.6Zr. The gaseous impurities of hydrogen and oxygen were determined to be less than 1 ppm and 30 ppm, respectively.

2.2.

Microstructure characterization

Microstructures of the alloy ingots and forged bulk were examined with a Quanta 600 scanning electron microscope (SEM) (FEI, Eindhoven, The Netherlands). Specimens for SEM observation were cut from the ingots, mounted in epoxy resin, mechanically ground with emery paper to 2000 grits and then polished with diamond paste. The polished surfaces of the specimens were etched in a solution consisting of 40% HCl, 40% HF and 20% HNO3 (in vol%). Grain size of the alloys was roughly evaluated with the linear intersection method for SEM images. Crystalline phases in the as-cast ingots were analyzed by X-ray diffraction (XRD) using a Rigaku D/max 2500 diffractometer (Rigaku, Tokyo, Japan) with monochromatic Cu Kα radiation. Lattice parameter of the crystalline phase was measured using (321) diffraction line. Step-scanning mode was used within this range to ensure the accurate measurement of the diffraction peak position. A silicon wafer was used to calibrate the error induced by instrument drift. Even though, in principle, multiple diffraction lines in XRD spectra can be used for the calculation of lattice parameter, however, the linear fitting with the extrapolating function of ð1=2Þðð cos 2 θ= sin θÞ þ ð cos 2 θ=θÞÞ does not work well in the current cases (not shown here). The data scatter is attributed to the selection of θ values lower than 601 for calculation, which yields bigger errors. It is a well-known fact that the change of Δθ at higher θ value corresponds to much lower Δ sin θ, then more accurate value of calculation. Thus, in the case of Nb–xTa–2Zr alloy, only one peak with a relatively high intensity is available at the high-angle (θ4601), i.e. the (321) diffraction line. In other word, the method of extrapolating function in terms of multiple diffraction lines in XRD spectra is not valid under the present situation.

2.3.

Magnetic susceptibility

Magnetic susceptibilities of the as-cast alloys were measured using a MPMS-7S superconducting quantum interference device (SQUID) (Quantum Design, San Diego, US). Cylindrical specimens of about 0.6 g in weight were taken from the ingot, then polished and weighed. Commercial pure Ta (99.98 wt%) rod with the equivalent weight was used as a reference. External magnetic field was incrementally applied to the sample from zero up to 1.5 T at room temperature. The corresponding magnetic moment was recorded simultaneously. Using the data of the magnetic moment against applied magnetic field strength, the slope of the tangent line at 1.5 T was calculated for each specimen as the magnetic susceptibility. One sample was tested for each alloy.

2.4.

Mechanical properties

Elastic constants of as-cast Nb–xTa–2Zr alloys, including the Young0 s modulus (E), shear modulus (G), bulk modulus (B) and Poisson0 s ratio (ν), were measured by using resonant ultrasound spectroscopy (Quasar, Albuquerque, NM). Cylindrical specimens of 4 mm in diameter with known volume and mass were placed between the piezoelectric transducers. Four specimens were prepared for each composition to give the average value. Two independent elastic constants C11 and C44 from each specimen were obtained for calculating the elastic modulus. Relative deviation of the experimental data for E, G, B and ν is 0.6%, 2%, 4%, and 0.8%, respectively. Mass density (ρ) was measured using the Archimedes method in ultra-pure water at room temperature, with more than three samples tested for each composition. Vickers hardness test of the specimens was carried out on a Mitutoyo MVK-H3 hardness testing machine using a load of 200 g and a dwell time of 10 s. At least twenty individual measurements were performed for each alloy to ensure the reproducibility. Uniaxial tensile tests were conducted at room temperature on a 2 kN Instron 5848 micromechanical testing machine (Instron, Norwood, MA) using a laser extensometer. A constant strain rate of 8  10  4 s  1 was used, and at least four specimens were measured for each alloy composition. The tensile specimens taken from the alloy ingot or bulk were machined with a nominal gauge dimension of 8 mm in length, 2.5 mm in width and 1 mm in thickness.

3.

Results

3.1. Microstructure characterization of as-cast Nb–xTa–2Zr alloys Fig. 1(a)–(f) displays a set of SEM secondary electron images of Nb–xTa–2Zr (x¼ 30, 50, 70) alloys as the representative, indicating that the microstructure changes at different locations of the ingots as the Ta content in the alloys increases. Microstructures of Nb–40Ta–2Zr and Nb–60Ta–2Zr are substantially similar to the case of Nb–50Ta–2Zr and Nb–70Ta– 2Zr, respectively, which are not shown here. As shown in Fig. 1(a)–(c), the bottom portion of alloy ingots contacted with copper hearth under melting exhibits coarse equiaxed grains of a single-phase solid solution. After etching, grain boundaries are clearly visible in the alloys at x¼ 30, 50, 70 with the average grain size of 550 μm, 470 μm and 390 μm, respectively. As the Ta content increases, the grain size is gradually reduced. The microstructure of Nb–xTa–2Zr alloys near the free surface of ingots is shown in Fig. 1(d)–(f). Significant grain refinement appears in Nb–xTa–2Zr (x¼ 30, 50) alloys. The grain boundaries become slightly vague in observation. Grain size was in a range of 170–220 μm, due to rapid solidification. However, as shown in Fig. 1(f), the grain boundaries are undistinguished in the Nb–70Ta–2Zr alloys because the dendritic feature of grown grain is prevalent. The secondary dendrite arm spacing of the dendritic crystals was about 25 μm. Fig. 2 shows the XRD patterns of the as-cast Nb–xTa–2Zr alloys. As seen in Fig. 2, the Nb and Ta phases were identified

journal of the mechanical behavior of biomedical materials 32 (2014) 166 –176

169

Fig. 1 – SEM images of as-cast Nb–xTa–2Zr (x ¼30, 50, 70) alloys. (a)–(c) at a location near ingots bottom contacted to copper hearth. (d)–(f) at a location of ingots top near free surface.

parameter of the alloys. It is consistent with a fact that the XRD peak positions shift towards the side of low angle. Quantitatively, the lattice parameter, a, was calculated by pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi combining dhkl ¼ a= h2 þ k2 þ l2 with the Bragg equation of 2dhkl sin θ ¼ λ, then, qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi a ¼ λ h2 þ k2 þ l2 =2 sin θ ð1Þ

Fig. 2 – XRD patterns of the as-cast Nb–xTa–2Zr series alloys.

as the bcc structure. The diffraction peaks of Nb and Ta were substantially overlapped, owing to a tiny difference in interplanar spacing. Peak positions of all Nb–xTa–2Zr alloys shifted to the side of low diffraction angle, with respect to the diffraction peak position of pure Nb and Ta. This suggests that the lattice of the alloy expanded, as signaled by the increased lattice parameter. In terms of the phase diagram of the binary Nb–Ta system, a single-phase solid solution could be formed within the entire compositional range. However, the atomic radius of Zr (0.160 nm) is  8% larger than that of the Nb (0.147 nm) and Ta (0.149 nm). Thus, the Zr which dissolved in the lattice of the Nb–Ta matrix is responsible for the increase of the lattice

where dhkl is the interplanar spacing of the crystal planes of Miller indices (hkl), θ is the diffraction angle of (hkl) planes, and λ is the wavelength of the X-ray radiation (0.1541 nm). In this regard, Miller indices of the (321) plane were adopted for the calculation, together with the accurately measured 2θ values around 120.71. The calculated a values of the Nb–xTa– 2Zr alloys are listed in Table 1, together with the values of pure Nb and Ta for comparison from Ref. Winter (2010). It is interesting to note that the a values of the Nb–xTa–2Zr alloys remain almost unchanged at 0.3316 nm, without significant dependency on the Ta content due to nearly identical a values of Nb (a ¼0.3303 nm) and Ta (a¼ 0.3306 nm). However, compared with pure Nb and Ta, the a values of the alloys were increased only by  0.4%, this is mainly attributed to the presence of 2.9 at% Zr as the alloying element. As a result, solid-solution strengthening is expected to be the predominant mechanism of the investigated alloys.

3.2.

Magnetic susceptibility of as-cast Nb–xTa–2Zr alloys

The measured data of volume magnetic susceptibility (χv) for Nb–xTa–2Zr alloys and pure Ta are tabulated in Table 1 as well. It shows that the χv of the Nb–xTa–2Zr alloys is

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Table 1 – Lattice parameters, volume magnetic susceptibilities (χv), mass densities (ρ) and elastic constants of the investigated Nb–xTa–2Zr alloys. For comparison, the corresponding parameters of pure Nb and Ta are also listed from Refs. Schenck (1996), Winter (2010). Material

Lattice parameter (nm)

χv (  10  6)

ρ (g/cm3)

E (GPa)

G (GPa)

B (GPa)

ν

Nb Ta x ¼ 30 x ¼ 40 x ¼ 50 x ¼ 60 x ¼ 70

0.3303 0.3306 0.3315 0.3316 0.3316 0.3316 0.3316

237 179 234 229 220 214 205

8.57 16.64 9.96 10.53 11.26 12.03 12.74

103 185 119 126 134 142 149

38 69 43 46 49 52 55

170 200 163 168 170 177 177

0.400 0.340 0.379 0.375 0.369 0.366 0.360

3.3. Elastic and mechanical properties of as-cast Nb–Ta–Zr alloys

Fig. 3 – Plot of volume magnetic susceptibility (χv) against Ta content for as-cast Nb–xTa–2Zr alloys. The upper and lower lines represent linear fitting with the measured data and the calculated with a rule of mixture, respectively.

decreased by 12% with the increase of Ta content, from 234  10  6 to 205  10  6. In contrast, the χv of pure Ta is determined to be 179  10  6, which is nearly the same as previously reported by Schenck (1996). This also means that the susceptibility of our ternary alloys is comparative to that of the Nb, Ta and Zr metals. Fig. 3 displays a plot of volume magnetic susceptibility of the Nb–xTa–2Zr alloys against the Ta content in the alloys. The measured data present a good linear relationship (R ¼ 0.99) with the Ta content. The susceptibility of the alloys is also estimated with the rule of mixture, χ ¼ x1 χ 1 þ x2 χ 2 þ ⋯ þ xi χ i

ð2Þ

where the xi and χi are the concentration in atomic percent and volume susceptibility of pure elements as the component, respectively. The volume susceptibility calculated with Eq. (2) manifests an underestimation with respect to the measured data, as shown in Fig. 3. The maximum difference between the two cases is less than 6%. It suggests that in this single-phase solid solution alloy, the susceptibility is fundamentally dependent upon the susceptibility of each element and their concentration. The dependency roughly follows the rule of mixture.

The measured mass density (ρ) and elastic constants of the as-cast Nb–xTa–2Zr alloys are summarized in Table 1, including the E, G, B and ν. For comparison, the values of pure Nb and Ta are also listed from Ref. Winter (2010). As the Ta content increases from 30% to 70%, the ρ of Nb–xTa–2Zr alloys increases by  28%. The density of Nb–70Ta–2Zr is 40% and is 60% larger than that of the L605 Co–Cr alloy and 316L SS, respectively. This implies that the Nb–xTa–2Zr alloys possess a higher X-ray linear attenuation coefficient, which corresponds to a higher radiopacity and better X-ray visibility. As the Ta content increases, the E and G are increased by 28%, whereas the B and ν of the alloys vary in a range of 163– 177 GPa and of 0.36–0.38, respectively, as presented in Table 1. Among the investigated alloys, the Nb–70Ta–2Zr alloy possesses the highest E value, which situates between Nb and Ta (see Table 1). This means that the E value is significantly dependent on the chemical composition of the alloys. In the case of a stent, sufficient radial hoop strength and negligible recoil are important issues, which are partially dominated by Young0 s modulus of the material. Thus, from the view of the radial hoop strength and recoil, high E value of the alloy with high Ta content exhibits an accessible advantage. Fig. 4 shows a plot of Young0 s modulus against the Ta content for the Nb–xTa–2Zr alloys. As indicated, the E exhibits a good linear scale with the Ta concentration. The E also can be estimated in terms of the rule of mixture, E ¼ x1 E1 þ x2 E2 þ ⋯ þ xi Ei

ð3Þ

where the xi and Ei are the atomic percent and Young0 s modulus of the pure element as a component, respectively. The difference between the calculated and the measured data is less than 3%. Thus, it is reasonable to estimate the E of the Nb–xTa–2Zr alloys by using the rule of mixture. Fig. 5 (a) and (b) illustrate tensile engineering stress–strain curves of the as-cast Nb–xTa–2Zr and Nb–50Ta–yZr (y¼ 2, 4, 6) series alloys, respectively. The obtained mechanical properties are summarized in Table 2. The alloys exhibit the typical uniform deformation prior to failure. As the Ta content increases in a range of 30–70%, the yield strength, s0.2, increases by  32% from 262 MPa to 347 MPa, while the ultimate tensile strength, sb, increases by 25% from 354 MPa to 443 MPa, which is caused by the solid solution-strengthening effect. In contrast, no distinct change for the elongation, δ, of the alloys is

journal of the mechanical behavior of biomedical materials 32 (2014) 166 –176

171

Fig. 4 – Plot of Young0 s modulus (E) against Ta content for ascast Nb–xTa–2Zr alloys. The upper and lower lines represent linear fitting with measured data and calculated with a rule of mixture, respectively.

presented, with a variation of 17–20%. The yield ratio, s0.2/sb, and specific strength, sb/ρ, is determined to be about 0.8 and 35  103 N m/kg, respectively. Moreover, the Vickers hardness (HV) values of the Nb–xTa–2Zr alloys are listed also in Table 2. The HV values increases by 18% as the Ta content increases from 1431769 MPa to 1686778 MPa. The trend of increased HV is consistent with the dependency of Ta content on the strength of the alloys. For the as-cast Nb–50Ta–yZr alloys, the s0.2 increases by  24% from 303 MPa to 376 MPa as the Zr content increases from 2% to 4%, while the change of sb is insignificant, as shown in Fig. 5(b). However, the reduction of elongation is remarkable, down to  2%. The increasing of Zr content up to 6% results in the degradation not only for strength but also for ductility. Thus, the alloy with 2% Zr possesses an optimal combination of strength and ductility. Fig. 6 shows a plot of yield strength, ultimate tensile strength and elongation against the Ta content for the Nb–xTa–2Zr alloys. As indicated, both of the s0.2 and sb significantly increase as the Ta content increases. Within a compositional range of 60–70% Ta, the increase of s0.2 and sb becomes steady, with an increment less than 5%. Nevertheless, the elongation drops down slightly from 20% to 17%. In light of these properties, the Nb–60Ta–2Zr alloy has an optimal combination in strength and ductility. Its s0.2 is comparable to that of 316L SS, while the ductility indicated by the elongation of 20% needs further improvement as performed in subsequent processing.

3.4. Microstructure and properties of forged and annealed Nb–60Ta–2Zr alloy The EBM melted Nb–60Ta–2Zr ingot was forged at 1100 1C to reduce casting defects and to refine grain size. Subsequently, the ingot was annealed at 1300 1C for 1 h to refine the grains through recrystallization. Fig. 7(a) displays a SEM image of the annealed Nb–60Ta–2Zr alloy. A single-phase microstructure

Fig. 5 – Tensile engineering stress–strain curves for as-cast (a) Nb–xTa–2Zr and (b) Nb–50Ta–yZr alloys, tested at a strain rate of 8  10  4 s  1. The curves in (a) are shifted relative to each other for clarity.

with the uniformly equiaxed grains is presented. The average grain size is approximately 30 μm, which is one order of magnitude smaller than that of as-cast state. Fig. 7(b) shows tensile engineering stress–strain curves of the annealed Nb– 60Ta–2Zr alloy, with 6 tested specimens. In contrast to the ascast state, as seen in Table 2, the s0.2 of 33279 MPa remains unchanged, while the sb of 474710 MPa is increased by 10%, then accompanied by an increase of sb/ρ up to 39  103 N m/kg. Moreover, the s0.2/sb is reduced to 0.7, reflecting an improved work-hardening ability during plastic deformation. As an indicator of work-hardening ability, strain-hardening coefficient, n, is determined according to a relation of sT ¼ KðεT Þn

ð4Þ

where the sT and εT are true stress and true strain, respectively, and the K is the strength coefficient. There is no physical significance for the K. It is simply treated as the true stress required to cause a true strain of unity. The K is about 757 MPa and the n value of the alloy is obtained as 0.17. It is noteworthy that for the forged and annealed alloy, the elongation reaches to 24%, which is almost 20% larger than that of as-cast state. Such improvement in ductility is more favorable to the expansion reliability during stent

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Table 2 – Measured mechanical properties of as-cast Nb–xTa–2Zr and Nb–50Ta–yZr alloys. Material

HV (MPa)

s0.2 (MPa)

sb (MPa)

s0.2/sb

δ (%)

sb/ρ (103 N m/kg)

x ¼ 30 x ¼ 40 x ¼ 50 x ¼ 60 x ¼ 70 x ¼ 60 (forgedþannealed) y¼4 y¼6

1431769 1441778 1539778 1705749 1686778 1607769 – –

262711 28677 30378 33279 347710 33279 376720 310724

354712 386731 394712 432713 443712 474710 409717 316724

0.7 0.8 0.8 0.8 0.8 0.7 0.9 1.0

2073 1774 1771 2072 1772 2473 2.370.5 0.470.1

35.5 36.7 35.0 35.9 34.8 39.4 – –

Fig. 6 – Yield strength (r0.2), ultimate tensile strength (rb) and elongation (δ) against Ta content for as-cast Nb–xTa–2Zr alloys.

deployment, since upon deployment the balloon expandable stent undergoes as much as 20%–30% plastic strain (Murphy et al., 2003).

4.

Discussion

4.1.

Magnetic susceptibility

MRI compatibility has become a key issue for metallic devices as the extensive application of MRI diagnosis in clinical practice. Artifacts size in the image is influenced by several factors, including magnetic susceptibility of the materials, shape and size of the stent, strength of the external magnetic field, echo time, etc. (Schenck, 1996; Starcukova et al., 2008). Magnetic susceptibility has been used as a measure to characterize the MRI compatibility of a metal. Lower magnetic susceptibility is associated with smaller size of the artifacts. For the currently-used stent materials, the artifact effect for the Ta and Ni–Ti stents is much less than for the 316L SS and Co–Cr alloy stents, owing to their lower magnetic susceptibility (Hug et al., 2000; Lenhart et al., 2000). To this end, development of new alloys with excellent MRI compatibility is considerably significant. Such effort has been made to account for some noble metal elements owing to their paramagnetism nature. A Pt–Ir alloy was employed to fabricate bare stents (Bhargava et al., 2000; Trost et al., 2004). However, its mechanical properties remain unsatisfactory. Severe stent recoil happens in this alloy in contrast to the

Fig. 7 – (a) SEM image and (b) tensile engineering stress– strain curves from six tensile specimens for Nb–60Ta–2Zr alloy bulk fabricated by EBM, with subsequent forging plus annealing treatment.

316L SS, as indicated by the recoil ratio of 16% for Pt–Ir and 5% for 316L SS. In addition, a Pd–Ag alloy with low magnetic susceptibility manifests weaker artifacts even in comparison with the Ta and Ni–Ti, as proven by in vitro tests. But, its mechanical properties and biocompatibility have not been fully understood yet (Hoare et al., 1953; van Dijk et al., 2001). Recently, a new Nb-based alloy, Nb–28Ta–3.8W–1.3Zr, was designed for MRI compatible stents (O0 Brien et al., 2008a, 2008c). The stent made from this alloy provides suitable compressive strength and elastic recoil, good corrosion resistance and low magnetic susceptibility, which is significantly

journal of the mechanical behavior of biomedical materials 32 (2014) 166 –176

superior to the 316L SS. Its elongation of 16.7% is slightly lower as the stent expansion required. As is well known, to minimize the imaging artifacts, it is preferred that the magnetic susceptibility of metallic devices is close to the level of human tissues, in a range of (–11.0)– (–7.0)  10  6. In this current work, we put forward a Nb–60Ta– 2Zr alloy with a low magnetic susceptibility of 214  10  6. The difference in the Δχv between Nb–60Ta–2Zr alloy and human tissue reaches to a level of 225  10  6. For comparison, Fig. 8(a) shows the χv of conventional stent metals and several new MRI compatible alloys. As shown in Fig. 8(a), the χv of Nb–60Ta–2Zr alloy is between the value of pure Ta and Nb, and is approximately 60% larger than that of pure Zr and its alloys (Nomura et al., 2009; Suyalatu et al., 2010). In addition, a Zr61Ti2Cu25Al12 amorphous alloy (denoted as ZT1) is included as well (He et al., 2011). Its χv was determined to be 108  10  6, nearly the same as the pure Zr. In contrast to conventional 316L SS, Co–Cr alloys and Ni–Ti alloys, the χv of our Nb–60Ta–2Zr alloy is remarkably reduced by a factor of  20, 3 and  0.13, respectively. Furthermore, in vivo

Fig. 8 – Comparison of (a) volume magnetic susceptibility, χv, and (b) Young0 s modulus, E, for annealed Nb–60Ta–2Zr alloy, with the currently-used stent metals and several Zr-based alloys. ZT1 represents Zr61Ti2Cu25Al12 bulk amorphous alloy (He et al., 2011). The χv data for Zr–3Mo is taken from Ref. Suyalatu et al. (2010). The χv data for remainder materials are taken from Ref. Schenck (1996). The E data are compiled from Refs. Mani et al. (2007), O0 Brien et al. (2008a), Brunski (2004), Cardarelli (2000), Davis (2003) and Park and Kim (2003).

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assessment of MRI compatibility for the current Nb–60Ta–2Zr alloy is an on-going work.

4.2.

X-ray visibility

Apart from the magnetic-resonance examination, X-ray fluoroscopy is also a routine tool to monitor the location and deployment of stents during implantation. Thus, in some sense, the X-ray visibility of a material is also a critical issue to determine the success of stent implantation. Due to the effects of scattering and absorption, the X-ray linear attenuation coefficient (μ) of a material is expressed as the following exponential law, I ¼ I0 e  μd

ð5Þ

where the I0 and I are the intensity of incident and output X-rays, respectively. And d is the thickness of material. The μ of a given alloy (μalloy) is attained as μalloy ¼ μmalloy ρalloy

ð6Þ

μmalloy ¼ ω1 μm1 þ ω2 μm2 þ U U U þ ωi μmi

ð7Þ

μmi ¼ Kλ3 Zi 3

ð8Þ

where the μm-alloy and ρalloy are the mass attenuation coefficient and the density of the alloy, respectively; the i stands for the component in the alloy. The μmi, ωi and Zi are the mass attenuation coefficient, weight percent and atomic number for the alloying elements, respectively. The λ is the wavelength of the X-ray, and the K is a constant. Therefore, based on Eqs. (6)–(8), the μalloy is dependent on the alloying elements (ωi and Zi) and the density of the alloy (ρalloy) under a given wavelength of X-ray. Radiopacity is supposed to be directly proportional to the μalloy. The larger the μalloy value, then the higher the radiopacity. In this sense, the heavyweight Ta and Ir (ρTa ¼ 16.6 g/cm3, ρIr ¼22.5 g/cm3) were incorporated as alloying elements in many alloys to enhance the μalloy, such as Nb–28Ta–3.8W–1.3Zr, Ti–17Ir, Ti–45Ta–5Ir (O0 Brien et al., 2008a, 2008b). Coating the pure Au (ρAu ¼19.3 g/cm3) or Ta on the surface of the bare 316L SS is also an approach to improve X-ray visibility. Unfortunately, the goldcoated stents give rise to severe artery restenosis (Kastrati et al., 2000; Macionczyk et al., 2001). On the other hand, the excessive radiopacity is undesirable, because it yields numerous artifacts which interfere with the lumen patency such as the scenario of pure Ta, when used for computed tomography (CT) imaging (Maintz et al., 2003, 2006). Therefore, it is necessary to balance the conflict between the radiopacity requirements in various imaging techniques. In this work, the Nb–xTa–2Zr alloys with densities in a range of 9.96–12.74 g/cm3 exhibit more suitable radiopacity, in comparison with the 316L SS, Co–Cr alloy and Ni–Ti alloy. Using Eqs. (6) and (7), the X-ray linear attenuation coefficient of Nb–60Ta–2Zr (μNb–60Ta–2Zr) can be estimated by   ð9Þ μNb–60Ta–2Zr ¼ ωNb μmNb þ ωTa μmTa þ ωZr μmZr ρNb–60Ta–2Zr The values of the μm-Nb, μm-Ta and μm-Zr at 100 keV are 1.04 cm2/g, 4.30 cm2/g and 0.97 cm2/g (Hubbell and Seltzer, 2004), respectively. Then, the value of μNb–60Ta–2Zr is obtained to be 36.02 cm  1, which is 12-folds larger than μFe of 2.92 cm  1. This implies that the X-ray radiopacity of

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Nb–60Ta–2Zr alloy is significantly superior to 316L SS, even for the coronary stents with the strut thinner than 100 μm.

4.3.

Elastic modulus associated with stent recoil

Stent recoil in the artery is of much concern for the application of balloon expandable stents, which is a potential source responsible for the in-stent restenosis (Barragan et al., 2000). In terms of clinical practices, acute recoil values of the stents varies in a range of 3.5%–8.2%, as the 316L SS tubular stent (Palmaz–Schatz) and the Ta coil stent (Wiktor) (Dejaegere et al., 1994; Haude et al., 1993). For the highly-recoiled stent, a higher balloon pressure is required in order to compensate the loss of lumen patency. However, the over-high inflation pressure and over-expansion may lead to vessel injury. The resultant inflammation reaction, endothelial cell growth and intimal proliferation have an effect to promote the in-stent restenosis (Rogers and Edelman, 1995; Schwartz et al., 1992). Thus, as an ideal stent material, the minimized recoil is expected. The factors influencing on the stent recoil are complex, including the material properties, geometric design and pressure imposed on the stent. From the material perspective, high elastic modulus, high work-hardening rate, low yield strength as well as strong radial strength are expected to prevent the effect of elastic recoil. As presented in Table 1, the measured Young0 s modulus (E) data of the Nb–xTa–2Zr alloys partially reflects the recoil performance. The E is a major factor which affects radial strength according to the product of E and t, where the t is the stent wall thickness. This product was used to assess the ability of a stent to support the arterial wall. As shown in Fig. 8(b), taking the Nb–60Ta–2Zr alloy as an example, the E of 142 GPa is nearly double that of the Ni–Ti alloy (  83 GPa for austenite phase), and comparable to that of pure Ta and 316L SS, with a reduction of 24% only. Although the E of Nb– 60Ta–2Zr alloys is 40% lower than that of the Co–Cr alloy, the stent recoil can be further resisted by geometric design of the stent. As indicated by the Nb–28Ta–3.5W–1.3Zr alloy with the E of 129 GPa, the recoil level of its stent is comparable to the L605 Co–Cr alloy (O0 Brien et al., 2008a). Furthermore, a small elastic strain (ε ¼s0.2/E) of a material reflects a recoverable strain, thus related to a lower elastic recoil. The ε of the Nb–60Ta–2Zr alloys is about 0.23%, which is comparable to that of 316L SS of 0.17% and Co–Cr alloy of 0.16–0.32%. Consequently, the recoil property of Nb–60Ta–2Zr alloys is qualified for stent application.

4.4.

Trade-off between strength and ductility

Mechanical properties are key factors for stent material screening. High strength of a material is required to fabricate the thin-strut stent with a smaller profile and enhanced flexibility, which is very important to reduce the restenosis rate. Sufficient ductility and small yield ratio play a role to ensure the right expansion and deployment of the stent without failure. In many scenarios, the materials designed for MRI-compatible stent is mainly challenged by the inadequate mechanical properties. In the current work, to provide a good combination of high strength and reasonable ductility, the solid-solution and grain refinement are used as the major

Fig. 9 – Plot of yield strength (r0.2) against elongation (δ) for the currently-used stent metals. Data are compiled from Refs. Mani et al. (2007), O0 Brien et al. (2008a). L605 alloy (ASTM F90) is selected as the representative of Co–Cr alloy. For Ni–Ti alloy, the yield strength of austenite phase and maximum strain recovery ( 8.5% after plastic deformation) are chosen.

strengthening approaches. The alloying with element Zr with large atomic radius remains forming a single-phase Nb–Ta solid solution. For these solid-solution alloys, the finallyoptimized Nb–60Ta–2Zr alloy presents a good combination in strength, ductility and Young0 s modulus. The grains in the alloy were refined by forging and subsequent annealing from 500 down to 30 μm. Such microstructure provided not only the enhanced tensile strength and lower yield ratio but also better work-hardening ability. It is worthy emphasizing that the elongation of the annealed alloy (  24%) is increased to be comparable to the pure Ta, enough to reach the required ductility for stent expansion. For a comparison, Fig. 9 presents a plot of yield strength against elongation for currently-used stent materials. Note that even though it is inferior to the Co–Cr alloys, the yield strength of annealed Nb–60Ta–2Zr alloy is comparable to the 316L SS already, and is significantly superior to pure Ta and Nb, let alone biodegradable stent metals like the Mg or Fe alloys. On the other hand, elongation of the current Nb–60Ta– 2Zr alloy is comparable to that of the pure Ta, in spite of its inferiority to the Co–Cr alloy and 316L SS. In contrast to the Nb–28Ta–3.5W–1.3Zr alloy developed by O0 Brien et al. (2008a), the current Nb–60Ta–2Zr alloy manifests superior ductility, with the elongation of 24% for Nb–60Ta–2Zr and of 17% for Nb–28Ta–3.5W–1.3Zr.

5.

Conclusions and outlook

To develop new alloys for MRI compatible vascular stents, the Nb–xTa–2Zr (30r xr70) series alloys were selected to optimize several key properties including the magnetic susceptibility, elastic modulus and tensile properties. In the as-cast state, a single-phase solid solution with bcc structure was formed in the alloys. The volume magnetic susceptibility (χv) and Young0 s modulus (E) of the alloys scaled linearly with the

journal of the mechanical behavior of biomedical materials 32 (2014) 166 –176

Ta content, substantially followed by the rule of mixture. Increasing the Ta content gives rise to a reduced magnetic susceptibility and higher Young0 s modulus, together with the elevated yield strength but less-changed elongation. For the electron-beam melted and subsequently-forged alloy, the ductility can be further improved by grain refinement and microstructure homogenization. Considering the multifactorial requirements of the stent, the Nb–60Ta–2Zr alloy provides optimal properties, including the volume magnetic susceptibility of about 3 % of 316L SS, Young0 s modulus of 142 GPa superior to pure niobium, high mass density favored to the X-ray visibility, yield strength of  330 MPa comparable to the 316L SS and elongation of  24%. These remarkable advantages make it quite promising as a new candidate of stent metals. Further work is in progress, including the electrochemical corrosion behavior under simulated plasma solution, fatigue property, hemocompatibility and in vitro cellular responses.

Acknowledgments The authors gratefully acknowledge the assistance in magnetic susceptibility tests from Dr. Xin-guo Zhao.

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