Fabrication of Mg alloy tubes for biodegradable stent application

Materials Science and Engineering C 33 (2013) 4746–4750

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Fabrication of Mg alloy tubes for biodegradable stent application Kotaro Hanada a,⁎, Kunio Matsuzaki a, Xinsheng Huang b, Yasumasa Chino b a

Advanced Manufacturing Research Institute, National Institute of Advanced Industrial Science and Technology (AIST), 1-2-1 Namiki, Tsukuba, Ibaraki 305-8564, Japan Materials Research Institute for Sustainable Development, National Institute of Advanced Industrial Science and Technology (AIST), 2266-98 Anagahora, Shimoshidami, Moriyama-ku, Nagoya, Aichi 463-8560, Japan

b

a r t i c l e

i n f o

Article history: Received 16 November 2012 Received in revised form 19 June 2013 Accepted 23 July 2013 Available online 27 July 2013 Keywords: Magnesium alloy Biodegradable stent Tube drawing Microstructure Dimensional accuracy Corrosion

a b s t r a c t Though Mg alloys are promising candidates for biodegradable stents, it is very difficult to fabricate stent tubes with high dimensional accuracy using Mg alloys because of their low deformability. This study aimed to develop thin-walled, high-quality Mg alloy tubes with good performance in stent applications. Cold drawing with a fixed mandrel was carried out for extruded Mg-0.8%Ca and AZ61 alloy tubes using optimized drawing parameters and lubrication, and stent tubes with 1.5–1.8 mm outer diameter and 150 μm thickness were fabricated. A dimensional evaluation showed that the tube dimensional errors were within 0.02–2.5%. Also, an immersion test of pure Mg with different crystal orientations showed that the crystal orientation affected the corrosion properties, results that are the same with other Mg alloys. The crystal orientation of the stent tube could be controlled by changing the deformation amount and direction in the drawing, showing that it is possible to further improve the biodegradability of stents by approaching their fabrication from a processing aspect. © 2013 Elsevier B.V. All rights reserved.

1. Introduction Bare metal stents, which are made of stainless steel or cobalt– chrome alloys, are used worldwide in the treatment of arterial diseases to open stenosed arterial vessels physically. These stents are usually permanently placed in the treated arterial vessel even after the patient has recovered from the disease, and this causes clinical problems such as in-stent restenosis and thrombosis that the current stent technology has been facing [1–8]. A drug-eluting stent introduced recently in the treatment of arterial diseases can provide effective short-term control of in-stent restenosis and thrombosis. However, even the current stent technology cannot remove the long-term risks of developing in-stent restenosis and thrombosis. A biodegradable stent, which temporarily opens the stenosed arterial vessel until vessel remodeling and then is completely absorbed into the body, is expected as a less-invasive stent to solve the above-mentioned problems. The base material of biodegradable stents is required to fulfill the following minimum requirements: 1) biodegradability (device-working duration of 3–6 months and full-degradation time of 1–2 years; 2) biocompatibility (no toxicity, no inflammatory response, and no release of harmful materials); 3) mechanical strength of N300 MPa; and 4) elongation of N 15–18% [9].

⁎ Corresponding author. Tel.: +81 29 861 7181; fax: +81 29 861 7007. E-mail address: [email protected] (K. Hanada). 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.07.033

Because Mg alloys have excellent biocompatibility and appropriate mechanical properties, they have attracted special interest as the most promising candidate material for the biodegradable stents [10,11]. Many studies on the biodegradability and biocompatibility of commercial Mg alloys such as Mg–Al–rare earth (AE) [12], Mg– Al–Mn (AM) [13], Mg–Al–Zn (AZ) [14–21], Mg–Y–rare earth (WE) [22–24], and Mg–Zn–rare earth (ZE) [21,25] series and alloys designed for biological applications [21,25–28], have been conducted; in fact, a few clinical trials in patients with critical limb ischemia and heart disease have already been reported [22,29]. On the other hand, there are a few studies on the mechanical properties and grain structures of Mg alloy tubes for stent applications [9,30]; however, there are very few studies on the dimensional accuracy of stent tubes being an important factor in the stent fabrication. Moreover, Mg alloy tubes with research-based prototype compositions can never be obtained. These are a major reason that the Mg alloy stent development doesn't expand as opposed to the alloy development. Generally, stent tubes require a thin wall of approximately 100–150 μm thickness and sufficient length of N1 m with high dimensional accuracy, which are fabricated by cold drawing extruded tubes [9]. Tube dimensional accuracy is, in particular, an important factor in stent development, because it directly affects the stent quality (dimensional accuracy, with or without defects) and performance (expansibility, radial force). However, Mg alloys that have a highly oriented hexagonal close-packed (hcp) structure show poor deformability at room temperature, and therefore, it is very difficult to cold-draw such alloys into stent tubes with high dimensional accuracy compared with other stent materials.

K. Hanada et al. / Materials Science and Engineering C 33 (2013) 4746–4750

In this study, we propose a fixed mandrel drawing method to enable high-accuracy cold drawing of the Mg alloys. This drawing method that uses a fixed wire mandrel to improve the inner surface accuracy of the drawn tube can obtain high drawing accuracy similar to the plug-drawing method, and postprocessing is not needed to remove the mandrel, unlike the moving mandrel method. Moreover, it is applicable to the drawing of thin-walled fine tubes b1 mm in diameter, which is technically difficult with conventional drawing methods such as free tube drawing, plug drawing, and rod drawing with a moving mandrel. The fabrication of thin-walled Mg-0.8%Ca and Mg-6%Al-1%Zn (AZ61) alloy stent tubes was carried out with the fixed mandrel drawing method at various processing conditions, and the dimensional accuracy and microstructures of the drawn tubes were investigated. Furthermore, control of the oriented hcp structure, which affects the mechanical properties and biodegradability of stents, is carried out by changing deformation concentration in the drawing process. 2. Experimental

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The extruded and drawn thin-walled tubes were cut into samples for three-point bending tests and microstructural observation using a wire electrical discharge machine (Sankyo Engineering DEL-50B). The samples for microstructural observation were first mechanically polished with up to #2000 grit and buffed and were then chemically etched with picric and nitric acids. The grain structure and size of the obtained tubes were observed by an optical microscope (Olympus BX60M). Electron backscatter diffraction (EBSD) analysis was carried out using a JEOL JSM-5910 scanning electron microscope (SEM) to determine the crystal orientation of the obtained tubes. The samples for EBSD analysis were also mechanically polished with up to #2000 grit and buffed and were then further etched with an argon ion beam. In the EBSD analysis, the orientation of the hcp basal face (0 0 0 1) was analyzed in the cross section of the obtained tubes and is shown as pole figures. Three-point bending tests were carried out using a Shimadzu EZ test at a punch speed of 1 mm/min for the as-drawn and as-heat-treated tube samples (1.5 mm outer diameter and 20 mm length) to determine the mechanical strength of the drawn tube.

2.1. Material preparation

2.3. Balloon expansion test

Pure Mg, Al, and Zn of commercial purities (N 99.99 mass%) and Ca (98 mass%) were prepared as raw materials. Mg-0.8%Ca alloy and AZ61 commercial alloy ingots were fabricated by induction melting at N750 °C in He atmosphere followed by solution heat treatment at 420 °C for 24 h. These ingots were cut into cylindrical billets (ϕ 65 mm × 60 mm) for hot extrusion. The billets were extruded at 400 °C with an extrusion ratio of 15:1. The extruded bars were further cut into small billets 10 mm in diameter and 15 mm thick, which were then extruded into thin-walled tubes with outer diameters of 1.9–2.9 mm and approximately 200 μm thick at 450 °C with an extrusion ratio of 68:1. The extruded thin-walled tubes were annealed at 300 °C for 30 min before drawing.

The fabricated thin-walled tubes were processed into stent devices through pulsed laser cutting and electrochemical polishing processes. A balloon expansion test was carried out to evaluate their expansion abilities. The stent device was crimped onto a balloon at the catheter tip, and then was expanded by inflating the balloon with Ar gas at 0.8 MPa. The SEM observation was carried out for the expanded stents.

Cold drawing of the extruded thin-walled tubes was carried out with an area reduction of b14% by the fixed mandrel method. Fig. 1 shows a schematic of the fixed mandrel method. In this drawing process, only the tube is drawn through the gap between the drawing die and the fixed wire mandrel. The drawn tube was annealed at 300 °C for 30 min to remove the work-hardening introduced by drawing. Next, the drawing and annealing of the tube were repeatedly carried out under the same conditions until the outer diameter and thickness were reduced to 1.5–1.8 mm and 150 μm, respectively. The influence of drawing parameters (i.e., area reduction, drawing speed, and lubrication) on the tube formability and accuracy was examined. The dimensional accuracy was determined using a stereomicroscope (Nikon SMZ-U), and a contact surface roughness meter (Mitutoyo SV-600) was used to determine the surface roughness of the drawn tubes.

Drawing die Thin-walled tube Drawing

An immersion test was carried out according to ASTM Standard G3172 to examine the influence of the hcp structure orientation on the biodegradability of Mg. A pure Mg ingot of commercial purity (99.99 mass%) was extruded at 400 °C and worked into square plate samples (10 × 10 × 1 mm3) with a 3-mm diameter hole. Then, the samples, which were cut out along both the radial and axial directions, have prism-face (1 −1 0 0)-rich and basal-face (0 0 0 1)-rich structures respectively. These samples were mechanically polished in dry with up to #4000 grit and washed in acetone, 99.5% absolute ethanol, and distilled water. They were then dried by air blowing. A simulated body fluid (SBF) was prepared [31] and used as the immersion solution. 3

140 120

Drawing stress /MPa

2.2. Cold drawing of thin-walled tubes

2.4. Immersion test

100

10 mm/min

80 60

Breaking

40

5 mm/min 20 0

Fixed mandrel Fig. 1. Schematic of the fixed mandrel drawing method.

0

2

4

6

8

10

12

14

16

Area reduction /% Fig. 2. Drawing limitation of the thin-walled Mg-0.8%Ca alloy tubes.

18

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Drawing stress /MPa

100

3. Results and discussion 3.1. Fabrication and properties of Mg alloy stent tubes

80

BN Mineral oil

60

Water-based lubricant

40

20

Graphite Machine oil

0 0.2

0.3

0.4

MoS2 0.5

0.6

0.7

0.8

0.9

Average surface roughness /µm Fig. 3. Effect of lubricants to the drawing stress and outer surface roughness of the drawn Mg-08%Ca alloy tube.

samples were prepared for the each oriented samples, and those samples were immersed in 50 mL SBF solution maintained at 37 °C in a water bath. After 10 days of immersion, the samples were removed from the SBF solution, gently rinsed with distilled water, and dried at room temperature. In order to evaluate the corrosion properties, the morphology observation and measurement of weight loss to within 0.1 mg were carried out for the immersed samples. Also, to examine the Mg ion elution concentration dissolved in the SBF solution, Xylidyl blue-I (C25H20N3NaO6S) reagent (Wako Pure Chemical Industries, Ltd.) was added to the spent SBF solution, and light absorption measurements at 660 nm wavelength were made using a digital colorimeter (Optima AC-114) for the solution samples (three samples for each material).

Thin-walled Mg-0.8%Ca alloy tubes were cold-drawn at various area reductions to evaluate the limiting area reduction (LAR) of each sample. The applied drawing speeds were 5 and 10 mm/min. As shown in Fig. 2, the drawing stress increased linearly with increasing area reduction; it also increased with increasing drawing speed. LAR tended to increase with increasing drawing speed, which was 14% at 5 mm/min and 15% at 10 mm/min. All the drawn tubes were cracked or fractured when drawn over LAR. The surface roughness of a stent tube must be controlled to involve dimension errors in the tube diameter and wall thickness. Also, a significantly rough surface leads to lowering of the processing accuracy in stent manufacturing and consequent stent performance. The drawing of the Mg-0.8%Ca alloy tubes with a 2.0 mm outer diameter was carried out at an area reduction of 6% with various liquid and solid lubricants to determine an optimal lubricant for minimizing the surface roughness. The drawing stress and surface roughness of the drawn tubes are shown in Fig. 3. Either the drawing stress or the surface roughness was high for the following lubricants: mineral oil, boron nitride (BN), molybdenum disulfide (MoS2), and waterbased lubricant. On the other hand, machine oil or graphite showed low value for both the stress and the surface roughness. In particular, the machine oil showed improved lubrication performance, making it possible to draw tubes with highly smooth surfaces, with an average roughness of 0.24 μm. Fig. 4 shows the longitudinal microstructure of the Mg-0.8%Ca alloy tube. The extruded tube shows a dynamically recrystallized structure with a mean grain size of 39 μm, including a twin crystalline structure at a depth of approximately 45 μm from both the outer and the inner surfaces [Fig. 4(a)]. Twinning deformation oc-

(c)

(b)

(a)

60 µm

60 µm

60 µm

Fig. 4. Longitudinal microstructures of the thin-walled Mg-0.8%Ca alloy tube: (a) as-extruded (ϕ1.9 mm and t170 μm), (b) as-drawn and (c) as-heat treated tubes (ϕ1.5 mm and t150 μm).

(a)

(b)

Ingot

Extruded tube

(c)

Extruded bar

Drawn tube

20 mm

1 mm

1 mm

Fig. 5. Mg alloy stent tubes: (a) from Mg-0.8%Ca alloy ingot to stent tube, (b) cross-section and (c) surface of AZ61 alloy stent tube.

K. Hanada et al. / Materials Science and Engineering C 33 (2013) 4746–4750

the fixed mandrel drawing method that we proposed is very effective for the fabrication of Mg alloy tube. Fig. 6 shows the result of the three-point bending test. The as-drawn Mg-0.8%Ca alloy and AZ61 alloy tubes had bending strengths of 22 N and 27 N, respectively; however, they cracked at the bending point. Upon annealing those tubes, no such cracks were observed at the bending point, although the bending strength decreased to 17 N and 16 N, respectively. The fabricated Mg alloy tubes were processed into stent devices through pulsed laser cutting and electrochemical polishing processes, and their expansion abilities were evaluated with a balloon catheter. Fig. 7 shows the Mg alloy stents after the expansion experiment. It was possible to expand the AZ61 alloy stent to twice its diameter without breaking and cracking at the strut connections. On the other hand, the Mg-0.8%Ca alloy stent was partially broken at the strut connections because of the lack of material ductility and a stent design that easily produces concentrated deformation at the strut connections. Thus, it was found that material ductility and stent design of the Mg-0.8%Ca alloy stent need to be further improved.

Table 1 Dimensional accuracy of the drawn Mg stent tube (ϕ1.8 mm and t150 μm). Dimension error/% b2.5 b0.4 b0.3 0.02

Thickness Outer diameter Inner diameter Straightness

40

As-drawn As-heat treated Bending strength /N

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30

20

10 3.2. Orientation control of hcp structure in Mg alloy stent tubes

0

Mg-0.8%Ca

AZ61

Fig. 6. Bending strength of the Mg alloy stent tube.

curred by cold drawing, and the crystalline structure was statically recrystallized by subsequent heat treatment [Fig. 4(b) and (c)]. As a result, a fine grain structure with a mean size of 14 μm was obtained. Fig. 5 shows the fabricated Mg alloy stent tubes. The tubes show no deformation or decentering, and the thin-wall has highly uniform thickness. Moreover, the surface appearance is very smooth, without shear cracking or tearing. As shown in Table 1, the dimensional accuracy is extremely high in both the Mg-0.8%Ca and AZ61 alloy tubes, which indicates that

(a)

The SBF immersion tests were carried out for pure Mg samples with differently oriented hcp structures, i.e., basal-face (0 0 0 1)-rich and prism− face (1 1 0 0)-rich structures. Fig. 8 shows the appearance of the pure Mg samples after 10 days of immersion in SBF. The basal face-rich sample maintained most of its original shape without local corrosion, and a corrosion product [26], magnesium hydroxide [Mg(OH)2], was uniformly precipitated on the surface; on the other hand, the prism face-rich sample was corroded locally at the edges with Mg(OH)2 precipitation. The weight loss rate was 0.5 mg/cm2/day for the basal-face-rich sample and 0.6 mg/ cm2/day for the prism-face-rich sample. Furthermore, the average Mg ion elution concentration was 397 μg/mL for the basal-facerich sample and 411 μg/mL for the prism-face-rich sample. It is considered that these results indicate that the hcp basal face (0 0 0 1) has much more corrosion resistance than the hcp prism face.

(b)

1 mm

200 µm

(c)

1 mm Fig. 7. Mg alloy stents after expansion: (a) AZ61 alloy and (b) its strut connection, and (c) Mg-0.8%Ca alloy.

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(b)

− Fig. 8. Appearance of (a) hcp basal-face (0 0 0 1)-rich and (b) hcp prism-face (1 1 0 0)-rich pure Mg samples after SBF immersion test.

(a)

(b)

4. Conclusions Characteristics such as the dimensional accuracy and microstructure of stent tubes, as well as the material properties, are very important factors in determining the quality and performance of stent devices; these characteristics are directly affected by the tube processing. The fixed mandrel drawing method was proposed to fabricate thin-walled Mg alloy tubes with high dimensional accuracy. The Mg alloy tubes fabricated by this method had outer diameters of 1.5–1.8 mm and thicknesses of 150 μm, with dimension errors of b 0.4% and b 2.5%, respectively. Moreover, the immersion test of the pure Mg showed that the crystal orientation, as well as the material properties and dimensional accuracy, is very important in determining stent performance. By changing the deformation concentration in the drawing direction, the crystal orientation of the drawn tube could be controlled. The above results make it easy to fabricate high-quality Mg alloy stents with good performance; consequently, acceleration in Mg alloy stent development is expected. References

(c)

(d)

Fig. 9. EBSD analysis result of the thin-walled Mg-0.8%Ca alloy tubes: (a) analyzed area in the tube cross-section, (0 0 0 1) pole figures of (b) as-extruded tube, and as-drawn tubes with (c) 19% and (d) 6% deformation concentrations in the drawing direction.

Moreover, comparable results were reported by Song for electrochemical corrosion of polycrystalline Mg [32]. Although grain refinement has been already revealed to be very effective in improving the corrosion resistance [9,33], the crystal orientation is also one of the important factors that affect biodegradable stent performance. Cold drawing with different deformation concentrations (6% and 19%) in the drawing direction was carried out for Mg-0.8%Ca alloy tubes, and the orientation of the hcp basal face (0 0 0 1) in the tube cross section was analyzed by EBSD [Fig. 9(a)]. The hcp basal face of the extruded tube was oriented horizontally in the circumferential direction [Fig. 9(b)]. On the other hand, by controlling the deformation concentration in the drawing direction, the drawn tube can have a randomly or vertically-oriented hcp basal face in the circumferential direction, as shown in Fig. 9(c) and (d). From the abovementioned results, the orientation of the hcp structure in the Mg alloy stent tube can be easily changed by drawing with controlled deformation. Hence, we should consider the orientation of the hcp structure, along with material design, crystal grain size, and corrosion properties, in order to further advance the development of Mg alloy stents.

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