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Research Papers

Fabrication and Testing of Planar Stent Mesh Designs Using Carbon-Infiltrated Carbon Nanotubes OPEN ACCESS

[+] Author and Article Information
Anton Bowden

Department of Mechanical Engineering,
Brigham Young University,
Provo, UT 84602

Manuscript received February 15, 2013; final manuscript received September 23, 2013; published online October 17, 2013. Assoc. Editor: Shaurya Prakash.

J. Nanotechnol. Eng. Med 4(2), 020903 (Oct 17, 2013) (7 pages) Paper No: NANO-13-1007; doi: 10.1115/1.4025598 History: Received February 15, 2013; Revised September 23, 2013

This paper explores and demonstrates the potential of using pyrolytic carbon as a material for coronary stents. Stents are commonly fabricated from metal, which has worse biocompatibilty than many polymers and ceramics. Pyrolytic carbon, a ceramic, is currently used in medical implant devices due to its preferable biocompatibility properties. Micropatterned pyrolytic carbon implants can be created by growing carbon nanotubes (CNTs), and then filling the space between with amorphous carbon via chemical vapor deposition (CVD). We prepared multiple samples of two different stent-like flexible mesh designs and smaller cubic structures out of carbon-infiltrated carbon nanotubes (CI-CNT). Tension loads were applied to expand the mesh samples and we recorded the forces at brittle failure. The cubic structures were used for separate compression tests. These data were then used in conjunction with a nonlinear finite element analysis (FEA) model of the stent geometry to determine Young's modulus and maximum fracture strain in tension and compression for each sample. Additionally, images were recorded of the mesh samples before, during, and at failure. These images were used to measure an overall percent elongation for each sample. The highest fracture strain observed was 1.4% and Young's modulus values confirmed that the material was similar to that used in previous carbon-infiltrated carbon nanotube work. The average percent elongation was 86% with a maximum of 145%. This exceeds a typical target of 66%. The material properties found from compression testing show less stiffness than the mesh samples; however, specimen evaluation reveals poorly infiltrated samples.

FIGURES IN THIS ARTICLE
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A stent is a flexible, tubular mesh which physicians surgically insert into the passages of the human body to reduce localized flow constrictions and support a collapsed passageway. Stents are used in multiple applications throughout the body, and for each application, various stent designs have been implemented [1]. A stent is a compliant mechanism, which depends on flexible members rather than rigid-body joints to achieve motion [2]. Specifically, stents are designed to be relatively small for insertion, but capable of expanding once inside the body. When using compliant elements, material selection is crucial, since the device must be designed around the material's maximum stress and strain.

Stents were initially made from stainless steel, a material known to resist corrosion and also capable of tolerating high stress, even plastic deformation, without fracture. For many engineering applications, stainless steel is an appropriate choice because of its strength, well-understood and repeatable properties, and reasonable cost. However, in biomedical applications, it may not react as well with the body as other materials, especially ceramics and polymers [3]. To improve biocompatibility, metallic, ceramic, and polymeric coatings are used with some success [3,4]. To the body, stents are foreign objects and are targeted by the body's defense mechanisms, causing known stent problems, including restenosis and thrombosis. In stent implantation, restenosis is reclosing of the stented passageway, due to tissue growing from the passageway wall onto and over the stent as an immune system response to protect the body. Over time, this tissue begins to hinder blood flow [5]. Another common problem is thrombosis, which is the formation of a blood clot that further restricts blood flow. While restenosis can increase the risk of thrombosis, it is also possible for a stent to cause thrombosis without first experiencing restenosis. Thrombotic events remain the primary cause of death after surgical complications [6]. Fortunately, materials with improved biocompatibilty, including many ceramics, have been discovered and engineered [7]. Pyrolytic carbon (also known as pyrolytic graphite) is one such engineered ceramic material which the body typically accepts [8]. Carbon and carbonaceous materials are generally well tolerated by animal cells and are therefore commonly used in medical implants [7,9,10]. Several forms of carbon-based material have been demonstrated, including graphene [11] and pyrolyzed polymer films [12]. Pyrolytic carbon has been used in applications such as finger joint replacement implants and heart valve replacements [13,14]. It is not uncommon to design a biomedical device constructed from steel or titanium because of their superior mechanical behavior, but coat the implant in pyrolytic carbon to keep the body from adversely reacting to the bare metal [10,15]. Pyrolytic carbon is a manmade material that can be created using CVD, a process commonly used for MEMS fabrication.

Unfortunately, using a pyrolytic carbon coating on a metal stent would not work well. As the stent expands, it undergoes large, plastic deformations. Pyrolytic carbon coatings are brittle, and would crack and flake during stent expansion. Therefore, a different approach for creating carbon stents is needed.

Research at Brigham Young University has led to a new method for fabricating systems on the micro and mesoscales, including compliant MEMS, using CNTs [16,17]. Carbon nanotube-templated microfabrication (CNT-M) is achieved by patterning a growth catalyst on a substrate and growing a forest of vertically aligned CNTs on top of that pattern. The voids between the CNTs are then filled, or infiltrated, with a material to form a solid structure using CVD. While much research has focused on the remarkable material properties of CNTs themselves, we use them as a physical framework for the infiltration material. Because the volume of CNTs is orders of magnitude below that of the filler material, the resulting properties of the infiltrated structure are primarily determined by the infiltration material. These voids can be filled with amorphous carbon, creating a CI-CNT material. Because the infiltrated carbon is pyrolytic carbon, the structure as a whole becomes a pyrolytic carbon structure. As shown in Ref. [18], initial experiments with carbon deposition have shown that carbon-infiltrated structures exhibit a remarkable degree of compliance and strain in bending, where tension is the typical failure mode. The high strain and superior biocompatibility marks CI-CNT as a potentially appropriate and even superior material to typical steels and alloys for stent fabrication.

The purpose of this paper is to test stent patterns made from the CI-CNT material to verify that they can endure the deflections needed in stent deployment. In order to better understand the ability of pyrolytic carbon to function in stent applications, flexible planar stent meshes were designed, as well as simple cubic structures. These designs were fabricated using the CI-CNT process, largely as outlined in Ref. [18], with modifications unique to the samples in this study. The mesh and cube samples were then tested to demonstrate flexibility characteristics, and a modulus of elasticity was calculated to validate each sample's material properties.

Test Pattern Design.

The purpose of the sample meshes was to test the ability of the CI-CNT material to flex and deflect in the way that a stent would, with the same scale of force and displacement. As discussed above, stents are tubular devices made to be inserted into body passageways to hold a collapsed passageway open. The processing required to fabricate cylindrical CI-CNT structures is still under development. However, much can be learned and demonstrated with stent designs fabricated in a planar configuration. The designs in this study were created as planar versions of flexible meshes that could also be implemented successfully in a tubular configuration.

As noted in Refs. [19] and [20], coronary arteries affected by plaque can have diameters reduced to 1 mm, a 66% decrease from a healthy artery diameter of about 3 mm. Typical metal stents are designed and fabricated such that they can be inserted into a small passageway and, once in position, are then expanded by a balloon, plastically deforming to the open or expanded configuration. Because the CI-CNT material is similar to a ceramic, it exhibits a low ductility [18]. We can, however, take advantage of the elastic properties and relatively large strain capability of the material by designing stent meshes in the open size as fabricated, and elastically compress them to a small size for insertion into the body. Once in the body, the stents could be released to their initial, as-fabricated size. Using compliant mechanism theory, two designs were created to achieve the 66% change in size. These designs used long, thin segments, and rounded corners to minimize stress concentrations and to distribute the stress along larger segments of the material, rather than focusing the stress on a single location and creating a weak point incapable of handling large deflections. Figure 1 illustrates the sample designs that were fabricated out of CI-CNT material, called the curved and rectangular designs. The two designs represent two potential shapes for stent meshes, one containing curved and angled segments and the other containing only straight line segments.

The mesh pattern designs were incorporated into a full-size, dark field photolithography mask for use with a 4-in. wafer and positive photoresist. Multiple instances of the patterns were arranged to make effective use of the entire mask and wafer area.

In addition to the mesh designs, small cubes with sides of 1 mm each were fabricated alongside for additional testing. In Ref. [18], the author performed much work to quantify material properties as tested in tension, with the CNTs aligned in one direction. As ceramics commonly perform better in compression than in tension, the cubes provided a means to test the compressive strength of the CI-CNT material. Furthermore, the cubes were tested by applying compression in one of two directions: either aligned with the nanotubes or transverse to the nanotubes. This provided a way to measure property differences in each direction.

Fabrication.

Fabrication followed the same general CNT-M process as used in Ref. [16], with the appropriate modifications made to the procedure to infiltrate the sample with carbon instead of silicon. The process is described below and illustrated in Fig. 2.

Using standard 4-in. silicon wafers, a 30-nm layer of alumina (Al2O3) was deposited using an e-beam evaporator (a). The alumina serves as a buffer layer to prevent diffusion of the iron layer into the silicon at the elevated CNT growth and infiltration temperatures. Following the alumina deposition, standard photolithography procedures were used to pattern AZ-3312 positive photoresist on a silicon substrate using a single full-size photolithography mask (b). Then, a 7-nm layer of iron was deposited using a thermal evaporator (c). The iron layer served as the catalyst for CNT growth and we chose to deposit 7 nm based earlier research [18]. The final catalyst pattern was obtained by sonicating the patterned wafer in liftoff agent to remove any remaining photoresist (d). The wafers were then diced into four individual pieces appropriately sized for the furnace.

The CNTs were grown using a CVD process (e). Samples were placed in a 1-in. quartz tube furnace and heated from room temperature to 750  °C in about 15 min, while flowing H2 at 218 sccm. Once the temperature reached equilibrium, 275 sccm of C2H4 was added for 30 min to grow the carbon nanotubes. After 30 min, the C2H4 was switched off, stopping the nanotube growth. Following growth, carbon infiltration was performed, also using CVD (f). The furnace was allowed to ramp up to the infiltration temperature of 900  °C. Once the furnace had reached the new equilibrium temperature, 327 sccm of C2H4 was again added, and infiltration via CVD commenced. Infiltration lasted for 30 min. When infiltration was complete, the furnace was shut off, but not opened, thus allowing the samples to cool slowly to around 650  °C. Once this temperature was reached, the furnace was opened and samples were allowed to cool much faster to a temperature appropriate for handling. Samples were then removed from the furnace. Slight intrinsic stresses on the large end “pads” of the samples resulted in the pads self-separating from the silicon substrate, likely during cooling.

To effectively separate the delicate samples from the substrate, additional etching processes were required. The CVD process creates a floor layer of carbon on the substrate surface where CNTs were not grown, which must be removed. In addition, to remove the sample from the silicon substrate, the surface layer of silicon would need to be wet etched away, separating the substrate and the infiltrated structure. The carbon floor layer was removed using a planar dry etching machine flowing O2 and setting the generator power to 200 W for a duration of 5 min (g). To fully remove the carbon floor layer, this dry etch process was repeated 3–5 times until the shiny surface of the silicon substrate was visible through the sample's infiltrated CNT structure. Then, KOH was used to wet etch away the alumina and silicon until the sample was released from the substrate. The sample was placed in a bath of KOH solution heated to 100  °C and etched for approximately 30–45 min, or until the infiltrated CNT sample had visibly detached from the substrate (h). Once the sample was released, it was removed from the KOH bath, rinsed with room temperature distilled water, and dried accordingly. Note that once the sample had released from the substrate, extreme care was taken in handling the sample to minimize any sample breaks or fractures before testing. An example of a completed sample of the curved design is shown in Fig. 3 and more detail in Fig. 4.

Testing Procedure.

Each mesh sample was tested in an Instron tabletop tensile testing machine, using specially machined fixtures attached to a force transducer to gradually expand the stent mesh and record the resulting force. The large pad and hole at the ends of each sample were used to secure the sample in the special fixtures while minimizing any stress concentration or effect of the clamping system on the delicate samples. The hole and shaft configuration also allowed the sample to slightly rotate, if needed, during expansion, to find the natural position of lowest energy, rather than attempting to manually set this position. Using a high resolution camera fitted with a zoom lens, we observed and visually recorded each test. Figure 5 shows how the testing equipment and samples were arranged.

Each mesh sample was tested individually by placing the unstretched mesh structure in the fixtures and expanding the sample at a constant rate of 0.5 mm/min until fracture occurred. Deflection and force values were continuously recorded as well as picture capture at a rate of 5 frames per second during each test. Following each expansion test, the data were copied from the testing program into a spreadsheet. Initially, unstretched frames and frame-before-fracture images from the camera were saved for expansion measurements and test videos of each sample were compiled.

Cubic structures were tested with a similar tensile testing machine, but in a compression environment. The millimeter-sized cubic samples were placed in between two flat-ended fixtures which were themselves clamped in the tensile machine's hydraulic gripping jaws. Each flat-ended fixture was machined out of hardened steel to minimize deflection resulting from the fixtures themselves. Cubic samples were compressed at a rate of 3 μm/min with CNTs oriented both in line with the compression load, and transversely oriented. Following each compression test, the data were compiled from the testing program into a spreadsheet.

Data Analysis.

As stated above, the purpose of this effort was to demonstrate the ability of CI-CNTs to flex and deflect as would be necessary for a stent of similar shape and size and to validate material properties as noted in Ref. [18] using stent-like geometry. The cubic compression samples were fabricated and tested to obtain raw material properties. These material properties could be easily taken from the raw compression data due to simple cubic geometry. Because the complex mesh geometry being tested was not a simple, constant cross-section beam, we could not use the Instron force data to directly find a modulus of elasticity, the strain, or the ultimate strength. Instead, a model was created in ansys to simulate each test independently of one another for both stent designs.

A single “cell” from each flexible design's array of repeated cells was constructed in the ansys simulation environment. Using PLANE183 elements, planar geometries with a constant thickness were created. The thickness of each sample was measured using a microscope to provide the thickness value for each simulation. One end of the simulation cell was fixed, while the other was allowed to move according to the images captured during testing. The images recorded during testing were used to find the deflection values for each test. For each sample, the unstretched and frame-before-fracture images were loaded into a cad software drawing environment. Once loaded, measurements of the initial, unstretched single cell, and the same deflected cell were taken on the pictures using the dimensioning tools; the deflection value was found by finding the difference between the two measurements. Examples of these images are shown in Fig. 6. To obtain the true dimensioned values, rather than a value skewed or scaled by the camera or dimensioning software, a reference dimension was also taken in the cad drawing environment whose actual dimensions were measured using a digital microscope. Together with the reference dimension, the actual cell displacement was calculated and applied to the model in ansys.

In addition to the known displacement, the force at the maximum displacement was continuously recorded by the Instron machine. Because the sample mesh designs were configured in a “serial” pattern, the same force is applied in each individual section of the design. Within each individual section undergoing the same force, a number of cells were configured in a “parallel” pattern. Each cell and section could be analyzed similar to a system of serial and parallel springs. By taking the measured maximum force from each test, we could then adjust it according to the specific geometry being in parallel or series that was simulated in ansys. This adjusted force then became the target for the ansys simulation. For each test, an initial guess for Young's modulus was given and ansys was allowed to iteratively solve the simulation, changing the Young's modulus until the resulting reaction force matched that which was measured with the Instron. In addition to Young's modulus, we were also able to obtain maximum values for stress and strain from the ansys simulations.

Figure 7 shows a typical force–deflection curve for the CI-CNT mesh designs. As mentioned previously, the plot reveals that the curve is nearly linear and additionally, samples do not exhibit plastic deformation before failure. This translates to a linear stress–strain relation for the material up to failure. The sudden drop of force back to zero indicates that failure is characteristic of an instantaneous, brittle failure.

Figures 8 and 9 show the simulation-calculated values of maximum strain and Young's modulus for each sample tested and Fig. 10 shows the percent elongation for analyzed cells. All three plots show error bars based on the 95% confidence intervals calculated from the data, and the design type (curved or rectangular) for each of the 11 samples is indicated. For stents, which rely heavily on the compliance of the material to function properly, a high maximum strain is desirable. High strains translate to greater deflection before failure. The highest average strain measured in this stent pattern experiment was greater than 1.4%, while the lowest average strain value measured was 0.4%. Modulus calculations had a similar range of values. The highest average modulus calculated was close to 15 GPa, while the lowest was close to 5 GPa. Percent elongation to failure values were consistently above 45%, with only two samples failing to reach the 66% elongation target and one sample reaching approximately 145%.

One of the most striking features of the data is that, out of the 11 samples, each of the four highest maximum strains was achieved by a rectangular design. No corresponding difference is seen in Young's modulus. This suggests that the rectangular design is capable of achieving higher overall elongations and strains compared with the curved design. This increase may indicate that the CI-CNT material performs better under pure bending stress, since the stresses in the curved design include both bending and radial stresses in the curved segments. It may also simply result from random variation, since sample 7, which is also a rectangular sample, performed no better than most of the curved designs.

In the figures showing values for average modulus and strain, there are also data representing a maximum and minimum. These values came from the variability in the beam's physical dimensions. The critical flexible segments were designed to be a certain size, but upon investigation and measurement with a digital microscope, it was found that the beam widths varied. In some cases, the difference was as large as 30% from the designed value, though the average variation was just 2.4%. These variations probably arise due to random variations in the exposure of the photoresist during lithography. Additionally, it was found that actual beam width varied on each sample. Approximately, 30 measurements were taken for each sample and from them, an average, maximum, and minimum beam width was calculated. These values were then used to reanalyze the data, resulting in a modulus and strain range for each sample.

Even though each mesh sample performed largely as anticipated, we can see from the results that a large amount of variability is present in the data. After conducting the experiments, we feel there were some major factors contributing to the variation. First, these tests were completed over a relatively long period of time so that they are affected by process variation over time. More specifically, each machine or piece of equipment used was likely affected by other experiments or work using the same machine. Second, the photomask itself, being made as an emulsion-based transparency mask, had line edges that degrade over time. These degraded edges lead to rough surfaces on the samples and stress concentrations where the sample would fail at lower stresses. Third, the samples were designed to have a large number of cells throughout the entire mesh. Only one or two of the cells could be captured in the view area of the camera and analyzed. It is possible that sample failure on some samples occurred somewhere on the mesh outside of the images being viewed. While the same force was felt throughout the sample, some sections or cells, due to geometric inconsistencies, may have undergone larger deflections. If this is the case, these sections would fail before the imaged section, and affect the results as simulated in ansys.

Figure 11 shows a summary of all cubic samples' tests in compression. The light lines indicate samples whose CNTs were oriented transversely to the load, while dark lines indicate CNTs aligned the same direction as the load. As loading began, both the transverse and aligned samples had the same Young's modulus. At some point, however, the aligned samples underwent a change to a reduced modulus, as shown by the dark lines suddenly continuing along a different slope. The highest ultimate strength was about 190 MPa, while initial Young's modulus was about 1.4 GPa. For samples that underwent a change in stiffness, the second, less stiff part of the curve gave a modulus of about 290 MPa. It was also interesting to note that the orientation seemed to affect the way each sample failed in compression. If the sample was loaded transversely, fracture occurred so quickly and completely that remnants of the samples could hardly be recovered. If the sample was loaded aligned with the CNTs, it went through a slow material separation, effectively “smashing” into a powder rather than fracturing.

The Young's modulus measured from the cubic compression samples is consistently lower than that measured from the mesh samples. Upon further investigation of the cubic compression samples, it was found that infiltration likely caused the reductions in Young's modulus. Figures 12 and 13 show the same compression sample after fracture. Due to the relatively large (1 mm cube) nature of these cubes, the amorphous carbon did not penetrate fully throughout the forest of nanotubes under the infiltration parameters used for the flexible meshes. While no exact density measurement was taken, the images in Figs. 12 and 13 show voids in the structures' composition which are not present on SEM images of the much smaller structures. This suggests that sample size may affect infiltration density, changing effective material properties.

The details shown in Fig. 13 may also explain the force–deflection behavior of the compression samples with CNT's aligned with the load. Internally, the sample appears to consist of bars of carbon deposited around individual nanotubes. When the load is transverse to the nanotubes, these bars would be in bending, leading to a smooth deflection during loading. However, when the load is aligned with the nanotubes, the individual bars could buckle at sufficient load, which would lead to the “knee” in the curve of Fig. 11.

Examination of Fig. 13 also suggests that two failure modes interact together during sample fracture. Several of the carbon sheaths surrounding a nanotube are fractured, suggesting that the carbon matrix itself can crack during failure. However, there are also several bare nanotubes protruding from the fractured surface. This suggests that these nanotubes first slid out of their carbon sheath as it fractured, before the nanotubes themselves failed. Hence, failure occurs with a combination of fracture of some carbon sheaths, combined with the sliding of nanotubes out of some sheaths.

An experiment was conducted to validate the potential capabilities of the CNT-M process to create stent-like geometry of the same shape and size as stents currently on the market. Once the structures were fabricated, they were tested in tension to demonstrate the material's flexibility and measure strain values. In addition to the measurement of strain values, Young's modulus was also measured/calculated to confirm the material as pyrolytic carbon. Next, the percent elongation of individual cells was measured. We found that stent-like structures composed of pyrolytic carbon fabricated using the CNT-M process will deflect sufficiently, even in excess of the target 66% elongation, without fracturing, and could therefore potentially be used in such applications. Lastly, even though the Young's modulus of the cubic compression samples (1.4 GPa) was much lower than the average Young's modulus of the mesh samples (about 10 GPa), an ultimate strength of 190 MPa was still achieved, which is slightly higher than the maximum ultimate strength of 181 MPa found in mesh sample tension tests. This suggests that the material is potentially stronger in compression.

While this research performs a similar exploration into the material properties of the CI-CNT material as discussed in Ref. [18], the load orientation on the material was different. Due to the volumetric dominance of infiltrated carbon in these structures, we expect loading in every direction, regardless of CNT orientation, to produce similar results. However, it is possible that the presence of CNTs may cause slight inconsistencies between values reported in this work and that of Ref. [18] as introduced by the compression testing in this work. A useful study would be the investigation and quantification of the CI-CNT material's anisotropy. It would also be beneficial to perform a study on the effects of infiltration density and sample size on material properties. As processing techniques for cylindrical fabrication become available, the CI-CNT material could be a potentially interesting choice for arterial stents.

We acknowledge Clarke Capital Partners for their support and funding in part of this work, as well as the BYU Integrated Microfabrication Lab. We also thank Darrell Skousen for his collaboration in this work.

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References

Schmidt, W., Lanzer, P., Behrens, P., Topoleski, L., and Schmitz, K.-P., 2009, “A Comparison of the Mechanical Performance Characteristics of Seven Drug-Eluting Stent Systems,” Cathet. Cardiovasc. Interv., 73, pp. 350–360. [CrossRef]
Howell, L. L., 2001, Compliant Mechanisms, Wiley Interscience, New York.
Hansi, C., Arab, A., Rzany, A., Ahrens, I., Bode, C., and Hehrlein, C., 2009, “Differences of Platelet Adhesion and Thrombus Activation on Amorphous Silicon Carbide, Magnesium Alloy, Stainless Steel, and Cobalt Chromium Stent Surfaces,” Cathet. Cardiovasc. Inter., 73, pp. 488–496. [CrossRef]
Serruys, P., Kutrykm, M., and Ong, A., 2006, “Coronary-Artery Stents,” New Engl. J. Med., 354, pp. 483–495. [CrossRef]
Serruys, P., Luijten, H., Beatt, K., Geuskens, R., de Feyter, P., van den Brand, M., Reiber, J., ten Katen, H., van Es, G., and Hugenholtz, P., 1988, “Incidence of Restenosis After Successful Coronary Angioplasty: A Time-Related Phenomenon. A Quantitative Angiographic Study in 342 Consecutive Patients at 1, 2, 3, and 4 Months,” Circulation, 77, pp. 361–371. Available at: http://circ.ahajournals.org/content/77/2/361 [CrossRef] [PubMed]
Iakovou, I., Schmidt, T., Bonizzoni, E., Ge, L., Sangiorgi, G. M., Stankovic, G., Airoldi, F., Chieffo, A., Montorfano, M., Carlino, M., Michev, I., Corvaja, N., Briguori, C., Gerckens, U., Grube, E., and Colombo, A., 2005, “Incidence, Predictors, and Outcome of Thrombosis After Successful Implantation of Drug-Eluting Stents,” J. Am. Med. Assoc., 293(17), pp. 2126–2130. [CrossRef]
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Figures

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Fig. 1

Sample mesh designs configured to undergo large deflections. On the left is the curved design, and the rectangular design is on the right.

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Fig. 2

CNT-M process with carbon infiltration

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Fig. 3

Example of sample size comparison to a United States penny

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Fig. 4

Sample mesh example after KOH release and rinse

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Fig. 5

Planar mesh test setup with Instron and gripping fixtures

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Fig. 6

Camera view sample of images where measurements of deflection were taken

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Fig. 7

Typical force–deflection curve for the stent mesh tensile samples

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Fig. 8

Strain values as calculated from the ansys analyses. C stands for the curved design, while R stands for the rectangular design.

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Fig. 9

Modulus values as calculated from the ansys analyses. Labels are as in Fig. 8.

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Fig. 10

Percent elongation of each analyzed test cell. Labels are as in Fig. 8.

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Fig. 11

Plot showing compression samples in both directions

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Fig. 12

SEM image of broken transverse compression sample

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Fig. 13

SEM image of broken transverse compression sample with detail on infiltration quality, revealing multiple voids in the material

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