An assessment of magnesium AZ31 coronary stents manufacture

AZ31 magnesium coronary stents were studied through a manufacturing process chain involving laser cutting, acid pickling, and dip coating. The purpose of this study was to evaluate surface thickness and geometrical dimensions of stents after processing. Stents were dip coated in a solution using PCL with 1% of TiO2. Additionally, AZ31 coronary stents were dynamically tested using a degradation system based on peristaltic pumps. Our results indicate that coated stents degraded slower than AZ31 uncoated control stents. After 4 weeks of dynamic degradation under flowing Hank’s solution, coated stents lost only ∼9% in weight while uncoated stents lost ∼27% in weight. Stents were qualitatively evaluated after four weeks of degradation. Our results demonstrate the formation of micro-pores after one and two weeks of degradation for coated stents. Lamination was observed after three weeks of degradation, meanwhile, uncoated stents resulted with notches and an irregular surface caused by degradation.


Introduction
Biodegradable coronary stents have several specifications (i.e., wider struts, greater stent thickness, lower degradation rate, higher radiolucent properties, and lower corrosion rate) to accomplish stent application requirements. Coronary stents have chronic complications during placement related to restenosis and thrombosis [1]. Therefore, a second intervention must be performed to allow blood flow into the artery [2,3]. Magnesium alloys have been mainly evaluated as a solution for stent replacement due to their corrosion mechanism that allows dissolving the alloy components in blood. Among the variety of magnesium alloys, WE43 and AZ31 alloys have the lowest corrosion rate; 6.82 and 10.08 mm year −1 , respectively [4][5][6]. This is promising when stent durability is a major concern. The corrosion rate of WE43 and AZ31 alloys can vary between 90 d and 4 months depending on the alloying additives and coatings [7]. Magnesium coronary stents are manufactured in a sequence of fabrication processes: laser cutting, acid pickling, sterilization and coating. Laser cutting on magnesium alloys has been studied with considerably good results in reducing surface roughness, dross, slag, and thermal effects [8]. It is of great importance to have a clean surface free of melted particles to accomplish a worthy interface between the coating and the stent's metallic surface. The most common chemical cleaning methods for magnesium alloys are vapor degreasing, solvent cleaning, emulsion cleaning, alkaline cleaning, and acid pickling. Acid pickling baths fulfils important functions to reduce dross particles, slag, or surface contamination in magnesium alloys. These functions are related with removing corrosion, eliminating oxide films and providing a clean surface that is receptive to chemical-conversion coating [9]. Additionally, polymeric coatings have come up as a solution to increase the lifetime of the medical device and to create a protective barrier [10]. Although there is a variety of coating techniques for stent coverings (dip coating, electrotreated coating, plasma coating and spray coating) [11], dip coating is the simplest method to cover medical devices compared to other methods [12]. This technique consists of depositing uniform layers at a controlled speed and quantity of cycles. Some dip coated commercial coronary stents are Cypher TM stent (Cordis, US), BiodivYsio TM (Biocompatibles Ltd UK), ZoMaxx TM (Abbott Vascular, US) and Endeavor (Medtronic Inc., US) Any further distribution of this work must maintain attribution to the author(s) and the title of the work, journal citation and DOI.  [13]. Several studies have worked on controlling the localized corrosion in magnesium alloys through adding rare earth and metallic elements such as lithium, zirconium and calcium in their basic chemical composition [10]. However, polymeric coatings have offered tremendous opportunities in terms of drug loading and control over drug release kinetics [14].Particularly, Polycaprolactone (PCL) polymer has a lower degradation rate than magnesium alloys [15] which preserves the surface's integrity of coated stents. Additionally, the incorporation of nanoparticles with a drug eluting purpose led to create multifunctional coatings [16][17][18]. Tamjid et al [19], performed a drug release study using PCL as polymer matrix and TiO 2 nanoparticles in 2D films. They used diverse concentrations of tetra-cycline hydrochloride (TCH) as an antimicrobial agent. Their study demonstrates that almost all the drug was released within almost 10 h incubation for PCL films. While, the PCL/TiO 2 film released around 50% of the TCH after 10 h of incubation which demonstrated the advantages of adding nanoparticles. Table 1 presents a literature review of acid pickling and coatings in the AZ31 alloy per year. Several works have studied the effect of acid pickling [20][21][22] and coating [15,23] on magnesium materials (i.e. tubes, sheets, rods or disks as raw material). For example, Brusciotti et al [24],evaluated AZ31 disks after being etched and coated with results on surface degradation. In this study, we studied the processing stages (laser cutting, acid pickling, coating and degradation response) to manufacture coronary stents which have geometrical and surface implications due to the micro-size scale of their complex geometry. Additionally, the dynamic degradation response of biodegradable coronary stents due to their corrosion mechanism caused by fluid is of great interest in magnesium alloys. Biodegradable metallic materials have been studied in order to analyze the degradation rate in similar conditions to those in the human body [3,25,26]. For example, Lévesque et al performed a test bench to approximate physiological conditions of coronary arteries in order to study the corrosion mechanisms of magnesium alloys [3]. In this study, AZ31 magnesium coronary stents process chain was investigated. Stents were laser cut and cleaned using the acid pickling method. Additionally, they were dip coated with a PCL and TiO 2 solution and degraded using a dynamic system for four weeks. In the study reported here, the main objective was to evaluate the influence of the process' sequence and their parameters in the stent's quality to simulate the stent performance in similar conditions to those in the human body. Based on the literature review, the contribution of this study is related to the analysis of etching and coating of coronary stents compared with raw geometries such as disks, tubes or sheets.

Magnesium minitubes
Chemical composition of the different raw tube material is shown in table 2. Tubes were acquired with different suppliers. The chemical composition of the tubes with OD of 3 mm and a thickness of 0.25 mm was provided by the supplier (Complex Materials, Eindhoven, Netherlands). While other tubes (OD=3 mm, T=0.22 mm; OD=1.8 mm T=0.16 mm; OD=1.8 mm T=0.11 mm), were bought from Yangzhou Sanming Medical Supply (Yangzhou, Jiangsu, China) and their chemical composition was obtained using atomic absorption spectroscopy (Laboratorios Fairchild, NL,Mexico) based on ASTM E1024-97. Table 2 presents the chemical composition by material used in experimental trials.

Laser cutting
A fiber laser source (YLR/150/1500, IPG photonics, USA) adapted with a fiber core with a diameter of 50 μm, a 120-mm collimator, and a 50-mm focal lens was used for experimental trials. A theoretical spot size of 21 μm was obtained with this configuration. AZ31 minitubes with different outer diameters and thicknesses were laser cut using a high-density polyethylene drape mounted on an acrylic to contain argon and create an inert atmosphere. The chamber set up is explained in our previous work [8]. The tubes were laser cut until an oxygen sensor InPro 6850I (Mettler Toledo, Cd. de Mexico, Mexico) ensuring an oxygen level inside the chamber below 5%. The geometry cut was Palmaz-Schatz stent design [28]. Table 3 presents the applied cutting parameters used for each minitube. Parameters highlighted correspond to our calculations of spot overlap and pulse energy based on our previous work. Intervals were selected based on preliminary tests before pitting corrosion was observed. Samples were washed using the same ethanol/distilled water solution and dried with compressed air.

Dip coating
Eight coronary stents were laser cut, acid pickled then coated. Twenty-four measurements were performed per stent to evaluate thickness. Coronary coated stents were placed in an extraction hood and in a vacuum desiccator to eliminate moisture. Samples were sterilized using a low temperature gas plasma sterilizer (100NX, Sterrad, ASP, CA, USA) for 47 min. The polymer solution was made of PCL (Poly (caprolactone) (MW 80, 000, Aldrich, St. Louis, MO, USA)) with 1% of TiO 2 (99.5%, Sigma-Aldrich, MO, USA) and 90% of CCl 4 as solvent (Chloroform 98%, (Sigma-Aldrich, MO, USA)). Solution was stirred during 4 h at 250 RPM to dissolve the polymer. A dip coated machine was built using a linear guide with a Z resolution of 10 μm and based on a stepper motor actuation. It was programmed using an Arduino card to set the entry and withdrawal speed of 125 mm min −1 . Stents were fully dip coated in PCL with 1% of TiO 2 solution in a vertical position during 5 seconds per cycle and for 10 cycles to allow surface wetting. Samples were suspended with a clip and were dried during 10 min at 70°C per cycle using a 50 w GU10 halogen lamp (Philips, Eindhoven, Netherlands). Specimens were maintained in a desiccator at room temperature. Coated stents were weighed using an Explorer scale (OHAUS, Cd. de Mexico, Mexico). Figure 1 presents the stages (start up, deposition, drainage, and drying) for surface coating. In stent immersion ( figure 1(a)), the stent is placed in solution; in the start-up phase ( figure 1(b)), the substrate is completely wetted; in deposition ( figure 1(c)), a thin film is deposited on the stent and then the excess of solution is eliminated during drainage in the drying stage ( figure 1(d)), to evaporate the solvent with an incandescent lamp ( figure 1(e)).

Coating characterization
Characterization of the PCL solution with TiO 2 was performed using a XRD (x-ray powder diffraction) technique to evaluate the crystallinity of the nanoparticles composite. XRD was assessed between 10°and 85°w ithin a 2θ range, and a step size of 0.026. The equipment used was a Panalytical Empyrean diffractometer (Panalytical, Almelo, The Netherlands), and a CuKα radiation, of which the wavelength was λ=1.5406 Å. The voltage and current applied were 45 kV, and 40 mA, respectively. Infrared studies were developed with Fourier transform infrared (FTIR) (Perkin Elmer, MA, USA) model Spectrum 400 recorded in the wavenumber range of 4000-400 cm −1 at room conditions. 2.6. Dynamic degradation A dynamic degradation system was built using four peristaltic pumps. An Arduino microcontroller and a 4-motor driver were used to control the flow rate. Pumps were mounted on a metal frame and were cleaned using a 70% ethanol and distilled water solution during 15 min before testing. Hank's solution was poured in a recipient by each stent and kept at 37°C using a hot plate while pH was maintained at 7.4, to simulate physiological conditions. The solution was recirculating using a flow rate of 40 ml min −1 to simulate blood flow [29,30]. Stents were maintained inside a hose while the solution was pumped. Figure 2 presents the dynamic degradation system used for experimental trials.  was changed to control the biochemical reactions of degradable magnesium. Dynamic corrosion tests were performed in a total of 12 stents (8 stents were coated with the same conditions and 4 stents were left as control without coating). Coated stents were evaluated per pair to have a replicate. Weight loss (W L ) was calculated using the next formula: where w f and w i are the final and initial weight of the stents, before and after dynamic corrosion tests.

Coronary stent characterization
Surface roughness and tube thickness were measured on the cutting edge and on the coronary stent's surface using an Infinite Focus Variation microscope (Alicona, USA) with a vertical resolution of 0.1 μm, a lateral resolution of 2 μm and a 20× objective lens after acid pickling. Additionally, surface quality was characterized after acid pickling, dip coating, and degradation using a scanning electron microscopy (SEM) model EVO MA25 (Zeiss, Oberkochen, Germany). These images were analyzed using the ImageJ software (NIH, MD USA). Figure 3 presents the surface roughness characterization. For laser cutting and acid pickling the surface roughness was measured on the cutting edge to evaluate the topography. After stents were coated, surface roughness was measured on the stent's surface (lateral roughness). Figure 4 illustrates the coronary stents after laser cutting (a), (c), (e), (g). Hence, some defects are observed after laser cutting, for example dross particles  be performed to minimize surface roughness in tubes with a thickness of 0.25 mm. Additionally, this tube has less content of Zinc compared to other tubes. When the addition of Zinc is greater than 1% it makes the alloy prone to hone cracking and it's corrosion resistance can be improved [9]. Figure 5 illustrates the average surface roughness results after stents were acid pickled. Stents with thicknesses of 0.11 mm and 0.16 mm resulted with a cleaning time of 180 s. For stents with a thickness of 0.16 mm, when time was increased, surface roughness rose. It is attributed to some located pitting corrosion on the cutting edge caused by an overexposure to acid solution. Power models were adjusted in experimental tests and R2 values are illustrated in Figure 5. Minimum values of average surface roughness were in between 0.26 and 0.34 μm. Plasma etching has been presented as a method to improve surface roughness in magnesium, resulting in a 10% reduction (from 0.191 to 0.172 μm) compared with chemical etching (i.e., phosphoric acid etchant)    [27]. Additionally. Zhou et al, presented a study of biological efficiency in vascular implants based on surface texturing gradients [32]. Their results reveled the importance of the role of surface topography in Magnesium materials which have an influence on cell activity. Further experiments can be performed to evaluate cell activity in coronary stents manufactured with different processing parameters. Figure 6 illustrates the thickness evaluation of coronary stents after acid pickling. Thickness reduction in the stent's cutting edge did not follow a trend. However, an experiment was kept until an abrupt thickness reduction change appeared. For stents with a thickness of 0.16 mm, a reduction change from 3.6 μm to 13.2 μm when etching time is in between 150s to 180 s (8% of the thickness) was observed. For stents with a thickness of 0.11 mm, surface roughness (figure 5) was not evaluated after 180 s, due to an increase in thickness reduction from 1.1 μm to 12.5 μm (i.e., 11.5% of the thickness). Stents with thicknesses of 0.22 mm and 0.25 mm required a cleaning time of 240 s to reduce surface roughness on the cutting edge. For stents with thicknesses of 0.22 mm, a maximum reduction of 7.2% of the thickness was obtained after 240 s of acid pickling. Meanwhile, stents with thicknesses of 0.25 mm, a reduction of 13% was found. The etching time differences     titanium. Two shoulders appear at 862 and 642 cm −1 related with the anatase phase presence. Additionally, a Ti-O rutile shoulder is located at 447 cm −1 .

Dip coating
Coronary stents with an outer diameter of 3 mm and a thickness of 0.25 mm were coated. Figure 9 presents the stent thickness evaluation of eight stents after acid pickling and dip coating. The error bars represent the standard error of 24 measurements. The number of measurements were quantified per stent due to the variations presented in each coated sample. The raw material has an inherent thickness variation due to the extrusion process (∼±25 μm). The deviations of coated stents are explained by the stent's proposed geometry. H-strut geometry promotes metal zones to cause conglomerations compared to hollow spaces where less material is deposited per cycle. It causes an irregular cross section of the stent. Lateral surface roughness was quantified and reported in figure 10. Coated stents showed variation among measurements of average surface roughness (R a ). When a stent was coated, small porosities were formed on the surface which increased both surface roughness parameters which, resulted in ∼1.5 μm and ∼10 μm for an average surface roughness (R a ) and a ten point mean roughness (R z ), respectively. Several studies have analyzed the influence of laser and etching parameters on surface roughness [8,22]. However, results are not conclusive for coated stents. According to Ping Li et al, the degree of roughness and the magnesium alloy composition can promote  degradation which leads to osmotic stress for cells [33]. Further experiments can be performed to evaluate the influence of surface roughness in a biological response.

Dynamic degradation
The weight of the stents was monitored after acid pickling, PCL coating and degradation. Figure 11 presents the weight loss percentage per week of degradation with power models adjusted and pH average measurements. AZ31 magnesium alloy without coating was used as control. From results obtained, AZ31 uncoated alloy presents a weight loss percentage of around 27% after four weeks of degradation while stents coated with PCL and TiO 2 presents a weight loss percentage of around 9%. For AZ31 coated with PCL and TiO 2 , pH was effectively monitored and controlled below 7.4. Figure 12 presents a qualitative study of the degraded stents. Coated stents without degradation (figures 12(a), (d), (g), and (j)) were performed with the same dip coated conditions. In deposition stage, the PCL with TiO 2 film was deposited. However, in the drainage stage, a meniscus formation at the end of the stent (red oval) was observed. This meniscus was observed in the H-strut geometry. Although stents were fully coated laterally, it presented a non-uniform thickness. Figures 12(b), (c), (e), (f), (h), (i), (k), (l) present stents after weeks 1, 2, 3, and 4 of degradation, respectively. Additionally, the degraded stent images (figures 12(c), (f), (i), (i)) are a zoom in of degraded stent. After weeks 1 and 2 some pores are observed (figures 12(b), (c), (e), (f)) while after weeks 3 and 4 polymer lamination is observed (figures 12(h), (i), (k), (l)). The usual degradation process in AZ31 alloys is proposed by Hanas et al [34], PCL allows a more controlled degradation, however it degrades over time by hydrolysis of ester bonds and leads to fatal cracks and pitting. The pores observed on the surface after weeks 1 and 2 of degradation have a diameter of ∼3.1 μm while after the second week they presented a pore diameter of ∼9.8 μm. This phenomenon was explained by Chen et al, they found that hydrogen escapes from high purity magnesium surfaces and pushes away the polymer (i.e., PCL and PLA) coatings which causes the polymer's degradation [17].
Additionally, Catauro et al, performed polymeric coatings of 5 and 10 wt% PCL with TiO 2 in titanium disks which presented pores promoting cellular growth [26]. These pores have a diameter of ∼4 μm after 16 weeks which can be compared with our results that show the same results in only 3 weeks. The accelerated degradation in our results can be explained due to the flow rate used in the dynamic degradation system which represents an increase of almost 30.5 times (1.35 ml min −1 versus 40 ml min −1 ). Figures 13(a)-(d) shows the degradation process in uncoated stents. Figure 13(a), (b) presents the stent after one week of degradation and figures 13(c), (d) illustrates the stent's surface after four weeks of degradation. At the end of the corrosion period cracks and notches were developed. According to Ostrowski et al PCL coating showed the least amount of surface corrosion and not visible cracking. [35] Further experiments can be performed to evaluate the incorporation of drugs to coronary stents to promote a slower ion release rate.

Conclusions
This study reports all relevant manufacturing steps of coated AZ31 stents and utilizes realistic stent geometry. Conclusions can be summarized as follows: • A complete process route of AZ31 magnesium coronary stent's fabrication is presented. The stents were laser cut, acid pickled, coated and evaluated through a dynamic degradation process.
• Acid pickling of stents resulted with the best surface quality with 180 and 240 s of etching time for the smallest (0.11 and 0.16 mm) and largest (0.22 and 0.25 mm) thicknesses, respectively.
• Dip coating technique was used to cover coronary stent surface. Coronary stents were coated with a solution of PCL and TiO 2 polymer.  • XRD and FTIR analysis confirmed the presence of titanium dioxide particles embedded in the PCL polymer coating.
• The coated AZ31 magnesium stent with PCL+TiO 2 lost 9% of its weight while the uncoated AZ31 magnesium stent lost 27% of its weight after dynamic degradation. Therefore, the stent degradation was delayed, and the use of polymer functionalized surface could be explored in further research.