Fabrication of the locally reinforced open-cell nitinol stents for intracranial vascular aneurysm

ABSTRACT

Intracranial vascular stent-assisted coil embolization has been widely used for intracranial vascular aneurysm. However, there are serious complication risks, such as thromboembolism, stent migration and spring coil prolapse, due to the inappropriate radial force, poor wall apposition and oversized mesh. Therefore, a locally-reinforced intracranial vascular stent was designed to improve the radial force and wall apposition performance. Based on finite element analysis and substantial measurements, the local radial force has improved 38.9% and thestability of the coil embolism has been enhanced as well. The releasing test in vitro suggests the stent can adhere to the vascular wall very well without the ‘Gator backing’ phenomenon. Furthermore, the implanting test in vivo shows complete endothelialization without restenosis and hyperplasia in 4 weeks. The outcomes of this study provide optimal design of the locally reinforced nitinol stent used for intracranial aneurysm therapy with reduced risks of vascular embolism and haemorrhagic apoplexy after implantation.

KEYWORDS:

• Nitinol stents
• wall apposition
• radial force
• open-cell
• closed-cell

Introduction

Intracranial aneurysm is a common cause of haemorrhagic cerebrovascular disease, accounting for 13% of cerebrovascular disease, which mostly occurs in complex and tortuous parts of the blood vessels like the junction of the cerebral arteries with vascular branches, such as the main artery branch point of the circle of Willis [1]. Intracranial vascular stent-assisted coil embolization has become the most effective solution to achieve aneurysm occlusion, especially for intracranial wide-necked, fusiform complex aneurysm lesions (aneurysm neck >4 mm or the ratio of aneurysm sac to aneurysm neck <1.5). Different intravascular stents based on metal materials have been developed to date, such as nanostructured austenitic alloys [2,3], copper-containing austenitic antibacterial stainless steel [4] and nitinols. Among them, nitinols are intensively applied in the manufacturing of vascular stents since they combine high hyperelasticity, good biocompatibility and corrosion resistance. Two kinds of nitinol stents have been in clinical use for spring coil embolism: open-cell stents and closed-cell stents. However, clinical experiences with these stents have confirmed the aneurysm recurrence rate of 12.7% and thrombotic complications rate of 4.5% after stent-assisted coiling treatment [5], which are mainly attributed to the poor administration of the radial force and wall apposition abilities.

Typical closed-cell stents are prone to be oval or crescent, and kinking at the tortuous vessel [6], which results in poor apposition to the vessel wall [7,8] and higher thromboembolic complication rates [9]. Dashti et al. [10] reported that the Enterprise stent with the closed cell structure was displaced after implantation for 4 months, which may also be related to stent malapposition. On the other hand, the open-cell stents have unconnected free ends, which renders better wall apposition to blood vessels; however, the lack of radial force is a critical aspect. In addition, when the stent is located in a tortuous vessel, from a structural point of view, the free ends may cause the mesh to protrude.

For stent-assisted coil embolization in the treatment of intracranial aneurysms, clinical results show that the incidence of in-stent restenosis complications is only 1.28% [11], while the risk of thromboembolism is as high as 4.2% to 17.1% [12], which is attributed to the vascular injury caused by the excessive radial force of the stent. In addition, due to the large mesh or low support force, there is no sufficient supporting to ensure the stents anchoring, resulting in the phenomenon of spring coil prolapse. Therefore, for the treatment of haemorrhagic diseases, the appropriate administration of the radial force from the perspective of structural designing is crucially important.

In order to obtain a proper radial force, Shobayashi et al. designed a closed-cell endovascular stent, which is dominated by a twined strut geometry to reduce the stiffness of the stent, but the wall apposition when contacted with the adjacent vessel wall was not assessed [13]. Li et al. introduced a regular arc structure into the middle of the traditional sinusoidal open-cell stent ring, and the radial strength for the optimised stent increases by approximately 40% [14]. However, the open-cell stent design will inevitably produce problems such as the free ends prolapsing.

As a consequence, a middle-enhanced intracranial vascular stent was designed for assisting coil embolization in the treatment of intracranial aneurysms in this paper. The main body of the stent is composed of an open cell structure to improve the flexibility and wall apposition of the stent. In addition, closed cell structures with different lengths are arranged in the middle part to reduce size of meshes, achieving mechanical enhancement and improving the stability of coil embolism. The local radial force of the stent and its wall apposition ability were evaluated by the finite element analysis (FEA). A series of mechanical properties of the stent were also tested, including three-point bending test, flat plate test, stent crimping test and conveying and releasing test.

Materials and methods

Design of stents

A centrally reinforced hybrid-cell nitinol stent was designed, containing closed cells and open cells (Figure 1a), the former is mainly used to provide better radial load capacity to the stent, and the later is employed to increase the conformability and wall apposition.

Figure 1. (a) Schematic diagram of stent structure, (b) open cell structure, (c) closed cell structure, and (d) the whole stent view.

The strut widths of the open cell and closed cell are 0.03 mm and 0.05 mm (Figure 1b,c), respectively. The wall thickness of the stent is 0.05 mm. Three closed cell structures were added in different parts of the stent: closed cell 1 (CC1, length-to-width ratio of 1.5) and closed cell 2 (CC2, double diamond structure) were designed to improve the stability of the stent, while closed cell 3 (CC3, length-to-width ratio of 1) was used to reduce the mesh size, achieving mechanical enhancement and improving the stability of spring coil. The open cells account for 30% of the total number of cells, and the distance of the locally reinforced closed cells is 2.25 mm corresponding to the aneurysm length. The stents were fabricated with nitinol tubes (out diameter = 2.8 mm and wall thickness = 100 µm, EUROFLEX GmbH, Pforzheim, Germany) by laser cutting, followed by cleaning, expansion and heat treatment, the stents with the diameter of 4.5 mm and length of 25 mm were fabricated, as shown in Figure 1d.

Material properties for finite element analysis

FEA was used to evaluate the local force in the middle position and wall apposition of intracranial vascular stents, since the built-in user material subroutine is more suitable for modelling of nitinol materials and has good capability to cope with problems involving large nonlinear analysis. In the numerical models, the nitinol stent was assumed to be homogeneous and isotropic. The material properties were obtained from previous work [15] (Table 1).

Table 1. Material properties of the nitinol stent (Auricchio model).

ρ	$E_A$	${\nu }_A$	$E_M$	${\nu }_M$	${\epsilon }^L$	$(\frac{\delta \sigma }{\delta T}{)}_L$	${\sigma }_L^S$	${\sigma }_L^E$	$T_0$	$(\frac{\delta \sigma }{\delta T}{)}_U$	${\sigma }_U^S$	${\sigma }_U^E$	${\sigma }_{CL}^S$	${\epsilon }_v^L$
6450 kg/m^3	68000 MPa	0.3	36000 MPa	0.3	0.04	2.1 MPa/K	520 MPa	635 MPa	37°C	2.1 MPa/K	180 MPa	26 MPa	610 MPa	0

Local radial force evaluation

Three stents with different lengths of CC3 structure were designed, abbreviated as CC3A (0 mm), CC3B (2.25 mm) and CC3C (4.75 mm), respectively, as shown in Figure 2. To verify the mechanical enhancement of different CC3 structure, the three-dimensional models of 4.5 mm diameter stents with different CC3 structures were established by wrapping a 2D sketch of the design around a cylinder. The 3D models of the stents were imported in FEA software. The meshes on all stents are built with C3D8R elements (an 8-node linear brick, reduced integration, hourglass control) of 0.01 mm size.

Figure 2. Schematic diagrams of (a) CC3A, (b) CC3B and (c) CC3C structures.

The local radial force in the middle of the stent was assessed to confirm the effectiveness of the middle reinforcement. In detail, the stent was placed on a flat plate and then the local radial force was measured during compression course. The indenter was withdrawn back to allow the stent recovery after a displacement of 0.7 mm. During the expansion simulation, the indenter was discretized using R3D4 elements (a 4-node 3D bilinear rigid quadrilateral) and the flat plate was discretized by SFM3D4R elements (4-node quadrilateral surface element, reduced integration).

Wall apposition evaluation

Wall apposition describes the ability of implanting stents to remain when contacting closely with the adjacent blood vessel wall. For comparison, a fully closed-cell stent was designed (named Stent-CC) was deployed in a vessel with different diameters, 3.5 mm at one end and 2 mm at the other end. A displacement load was applied to the stent by a rigid clamped shell, pressing the outer diameter to 1.8 mm. The clamped shell was then withdrawn to allow the stent to expand freely and adhere to the wall of the blood vessel. The cross section in the transition segment was obtained to visually observe the apposition of each stent with the vessel wall.

In this model, the mechanical properties of natural blood vessels were simplified to homogeneous and isotropic model, assigning the vessels’ density of 1200 kg/m³, Poisson’s ratio of 0.45 and Young’s modulus of 4 MPa. The meshes on the vessel are built with C3D8R elements. Nlgeom is open during the wall apposition test. A static, large deformation analysis was performed using the default contact algorithm. A ‘surface-to-surface contact’ algorithm was defined between the vessel inter surfaces and the outer surfaces of the stent. The friction in the tangent direction was defined as a penalty function assuming the friction coefficient of 0.5 for all contract surfaces, and ‘Hard contact’ was defined in the normal direction.

Mechanical properties of as-fabricated nitinol stents

Stents fabricated by nitinol tubes were expanded to 4.5 mm. The morphology was observed using scanning electron microscope (SEM, Inspect F50, FEI, Hillsboro, U.S.A), and the chemical compositions were characterized by energy dispersive spectroscopy (EDS).

The radial force of the fabricated stent was analysed using the ‘Flat Plate Test’. The stent was placed between two flat plates, and the radial force was then measured during compression of the two plates. The flat plate was withdrawn back to allow the stent recovery from 50% compression (2.25 mm displacement). The radial force was collected in the period. To establish the relationship between the bending moment and displacement, two lower static supports were displaced in a selected span 20 mm. The upper load applicator sat in the middle point of the span. In detail, the upper load applicator moved down at a speed of 0.05 mm/s then withdrew back to allow recovery of the stent due to the spring-back effect after a displacement of 4.5 mm. The force value was collected in the period, which was multiplied by the length to obtain the bending moment. In addition, in order to further characterize the radial force of the stent, the TTR2 radial force testing machine (Blockwise Engineering LLC, AZ, U.S.A) was used for stent crimping test. During the whole process, the stent was compressed or released along the radial direction, and the radial force of the stent was recorded by the force sensor.

Finally, an in vitro silicone hose model for the evaluation of delivery performance of stents was built, which was composed of hoses with the same inner diameter (3 mm) and different curvature radius (4 mm, 5 mm, 6 mm, respectively). Subsequently, the stent was released into the model through the 2.3 F Prowler® Select Plus microcatheter (0.021 inch inside diameter) and transmission-related properties were investigated, such as delivery, releasing and apposition of the stent.

In vivo stent implantation

Three stents were implanted into the abdominal aorta of three adult male Japanese white rabbits weighing 3–3.5 kg through the delivery device. After 28 days, the test samples were taken and fixed with 10% paraformaldehyde solution at 4°C for 5 days. Then the samples were cut in half. One half was cut longitudinally along the artery direction. After gradient dehydration, the endothelialization degree of the stent was observed by SEM. The other half of stents were embedded in epoxy resin and sliced into 100 μm thin pieces. Haematoxylin-eosin (H&E) staining (Solarbio Science and Technology Co., Ltd., Beijing, China) was performed to observe the new endothelial cell layer. Animal welfare and the experimental design were approved by the Ethics Committee of the Shenyang Research Institute of Chemical Industry, China.

Results

Local radial force of the stent

The local radial force in the middle of the stent is shown in Figure 3. The reaction force of the indenter subjected to the stent was 0.0162 N (CC3A) < 0.0225 N (CC3B) < 0.0337 N (CC3C), respectively. With the increase of the middle reinforcement length, the force value in the middle of the stent increases, and the enhancement effect is more significant.

Figure 3. Local compression test simulation of stents with different lengths of CC3 structure: A CC3A, B CC3B, CCC3C, D-F enlarged views of the black dashed boxes.

In addition, there are structural differences between the closed cell design and the open cell design. The mesh wires of the closed cell structure are connected, and the force transferring can be realized when compressed, inducing higher resistance to external forces and uniform deformation (Figure 3e,f). Open cell structures with unconnected mesh, forming units with open ends, provide better flexibility during bending process, but supporting performance is relatively poor. Most importantly, free ends are easy to extend under pressure (red circles in Figure 3d), thus increasing the risk of thrombosis.

Wall apposition of stents

The closed cell structure Stent-CC and the open cell structure CC3B were simulated to be released into vessels with varying inner diameters, and the FEA results are shown in Figure 4.

Figure 4. Deployment simulation of stents into vessels with varying inner diameters: (a, b) wall apposition of the stent, (c, d) S11 stress distribution of the stent on the vessel wall, (e, f) S33 stress distribution of the stent on the vessel wall. (a, c, e) Stent-CC, (b, d, f) CC3B stent.

The cross-sectional micrographs of the stent in the transition region of the inner side of the vessel are shown in Figure 4(a,b). It is obvious that the Stent-CC mesh of the closed-cell stent is connected to each other, and the force exerted on the retracted end of the stent is transmitted to the whole stent, and the wall apposition is poor in the transition area of the inner side (Figure 4a, red circle area), forming a horn mouth shape, which may lead to an increased risk of migration. As for the open-cell stent CC3B, typical free ends are attached to the wall at both ends of the transition region because part of the mesh is not connected (indicated by the arrows in Figure 4b). The stress distribution of different stents acting on the vessel wall in different directions are shown in Figure 4c–f. It can be seen that the stress on the vessel wall in the S11 direction (radial) of Stent-CC (Figure 4c) is significantly higher than that of the open-cell stent CC3B (Figure 4d), mainly concentrated at one end of the small-diameter vessel, indicating that the closed-cell stent at one end of the large-diameter vessel has a weaker interaction with the vessel wall, thus is more likely to migrate. In addition, the stress of the open-cell stent in the S11 direction of the vessel wall is evenly distributed at both ends of the vessel, achieving a good anchoring effect. On the other hand, for the S33 direction (axial), the shear stress on the vessel wall of Stent-CC (Figure 4e) is also higher than that of the open-cell stent CC3B (Figure 4f). As such, the difference of the diameter at the position of distal and proximal vessels may exert a compressive force on the stent, driving it to move to a more stable position, resulting in the migration of the stent into the vessel with larger diameter.

Characterization of the as-fabricated stents

Based on the simulation results, the middle-enhanced stent with the length of 25.5 mm was processed, and the size of the CC3 cell is about 2.25 mm (Figure 5a), consists of both open cell and closed cell structures (Figure 5d,e), which are separated from each other by rounded transitions (Figure 5b,c), thus reduces stress concentration and avoids damage to vessels due to the sharpness of the stent. Fifteen positions of the stent were selected randomly for measurement and verification, the width of the 12 periodic V-shaped unit mesh was 0.037 ± 0.003 mm, the width of the 8 periodic V-shaped unit mesh was 0.050 ± 0.003 mm, and the wall thickness was 0.054 ± 0.003 mm. EDS analysis confirmed that nitinol alloy scaffold consisted of 50.81 atomic percentages of Ti and 49.19 of Ni (Figure 5f).

Figure 5. Dimension verification and composition of laser-cut stents: (a) the whole stent image, (b, c) structural connecting rod, (d) open cell structure of the stent, (e) closed cell structure of the stent, (f) the chemical composition of the stent measured by EDS.

Finally, the as-fabricated stent was tested to evaluate mechanical properties, as shown in Figure 6. The local radial force of the stent at a displacement of 0.7 mm is 0.0225 N (Figure 6a), which corresponds well to the simulation results. When compressed to 50% of the stent’s nominal diameter (2.25 mm), this value increase to 0.086 N, much higher than the radial force value (a range of 0.008–0.026 N) of other aneurysm stents as reported [16–18]. Through the three-point bending test (Figure 6b), the final bending moment of the bracket was calculated to be 0.148 N·mm. Meanwhile, the radial force was 0.0161 N/mm by flat plate test (Figure 6c). In addition, the chronic abduction force of the stent obtained from stent crimping test was about 1.03 N (the position of the red dot in Figure 6d).

Figure 6. Mechanical properties of the as-fabricated stent: (a) the relationship between the local radial force and displacement, (b) the relationship between the bending moment and displacement, (c) the relationship between the radial force per unit length and displacement, (d) the relationship between the chronic outward force and diameters.

The results of stent releasing into different models in vitro are shown in Figure 7. The stents were released into silicone hoses with a centerline radius of 4 mm, 5 mm, and 6 mm, respectively (Figure 7a–c). It is obvious that during the releasing processes with different centerline radii, the stents could be successfully delivered and released through the microcatheter, showing good wall apposition and excellent compliance. After that, the stent was further released into a silicone aneurysm vascular model with an inner diameter of 4 mm (Figure 7d,e). For the tortuous vessel with aneurysm (Figure 7e, the aneurysm sac is about 9.1 mm wide, the aneurysm neck is about 7.5 mm, the ratio of above two is about 1.21, and the diameter of the aneurysm vessel curvature circle is 5 mm), good wall apposition for the stent can be observed, most important, at the neck of the tumour, no ‘Gator backing’ phenomenon at the free end of the open cell appeared.

Figure 7. Stent deployment process: 3 mm inner diameter bent silicone tube with a centerline radius of (a) 4 mm, (b) 5 mm, (c) 6 mm and (d-f) aneurysm silicone vessel model with inner diameter of 4 mm.

In vivo animal implantation experiment

The rabbit abdominal artery model was used to evaluate the stent implantation in vivo for 4 weeks. The H&E staining showed that a thin layer of endothelial cells formed and wrapped the whole stent post surgery for 4 weeks, and there was no restenosis and hyperplasia (Figure 8a–c). Blood vessel tissue covered on the stent was further examined by SEM. It can be seen that the stent struts were wrapped in the middle layer of blood vessels, which has been almost completely healed, as shown in Figure 8d–f [19–21].

Figure 8. (a-c) H&E staining of stent samples after implanting in rabbit abdominal artery for 4 weeks, (d-f) stent strut encased in vascular tissue and the enlarged views (SEM).

Discussion

Intracranial aneurysms are generally found in complex and tortuous parts of the blood vessels, often accompanied by branching, such as the junction of the cerebral arteries. Many clinical studies have shown that closed-cell stents have a higher risk of spontaneous delayed migration [22–25], especially for the treatment of basilar artery apical aneurysms, where the stent extends from the basilar artery to the posterior cerebral artery (PCA). When stents implanting, the difference in the inner diameter of the vessel is about 1–1.5 mm, which is considered to be the main reason for the spontaneous delayed migration of the stent. Compared with the closed-cell stents, the open-cell stents can provide better adherence when implanted in blood vessels with varying inner diameters owing to the unconnected free ends. Therefore, the open cell structures are preferred for the auxiliary coil embolization stents for intracranial aneurysm.

However, when the open-cell stent is located in a tortuous vessel, its free ends may cause the mesh to protrude, forming a ‘Gator backing’ at the neck of the aneurysm [16], which might cause coil prolapse. Yang et al. established linear theoretical model to estimate mechanical properties of stents with rhombus cells under specific loading conditions [15]. According to their algorithm, the open cell structure resembles a spring with less stability, while the closed cell structure is closer to a truss, where the loading can be distributed evenly and thus providing the higher radial force. Additionally, most intracranial aneurysms range from 2.1 to 9.3 mm in width, and thus the appropriate enhanced length for stents are required [26]. Therefore, setting a 2.25 mm length single diamond cell CC3 structure with a length-to-width ratio of 1 in the middle part of the stent can be expected to improve the local radial force, and with the mesh size decreasing, the stability of the coil embolism is enhanced. In addition, although the open-cell stents showed good adherence and anchoring with little migration, Broadbent et al. [27] found that the open-cell stents may migrate due to less radial force. Therefore, closed cell structures CC1 and CC2 were added to the stents in this study, which can further improve the stability of the stents and avoid migrating risk.

Previous studies have shown that the wire width, wall thickness and mesh size of the stent have great impact on the radial force [14,28–30]. It is well known that too thick the stent wire will lead to restenosis, while too thin the stent wire cannot provide appropriate radial force. Therefore, based on our previous work [31], the strut widths of the open cells and closed cells are set as 0.03 mm and 0.05 mm, respectively. In addition, the cell shape also influences the radial force. Karnessis et al. designed different cellular structures, including inverted hexagonal honeycomb, V-type and chiral auxetic structures, the results demonstrated that the adoption of auxetic tubular structures, through appropriate designing, can contribute to extending considerably the curvature that the tube can reach before undergoing kinking [32]. The open-cell stent with rhombus/diamond structure has good structural stability, but the wall-apposition and bending performance are slightly poor. Therefore, an uneven design with different functions for the stent is adopted in this work, that is to say the main body is composed of an open cell structure, and a closed cell structure with different lengths arranged in the middle parts, as such, a middle-enhanced intracranial vascular stent for assisting coil embolization in the treatment of intracranial aneurysms is obtained.

The local radial force test shows that the closed cell structure indeed improves the local compression resistance of the stent, and increases the stability of the coil. According to the feedback force value of the three-point bending test, the final bending moment of the stent after processing is 0.148 N mm, which is superior to other commercial intracranial vascular stents. The radial force obtained from the flat plate test was 0.0161 N/mm, which was much higher than the performance of commercially available stents (0.0065–0.0116 N/mm [16]). Therefore, the newly designed stent has good local compression resistance at the neck of the aneurysm and appropriate wall apposition as well.

Conclusion

Based on the FEA, a locally reinforced intracranial vascular stent for assisted coil embolization in the treatment of intracranial aneurysms was designed with the appropriate radial force and good wall apposition. The main body of the stent was composed of an open cell structure, and CC3 closed cell structures with different lengths were set in the middle part to reduce the mesh size. The local radial force increased from 0.0162 N (CC3A) to 0.0225 N (CC3B), and the stability of the coil embolism was enhanced as well. The releasing test in vitro suggested that the stent can adhere to the vascular wall very well, without the ‘Gator backing’ phenomenon at the free end of the open cell. Complete endothelialization of the stent was observed after in vivo implantation of the stent for 4 weeks.

Acknowledgments

This work was supported by the National Key Research and Development Program (2018YFC1105503). We also thank Chen Liu and Nina An in Shenyang Research Institute of Chemical Industry for helping complete animal experiments. The FEA using Abaqus (SIMULIA Inc., U.S.A.) were provided by National Supercomputing Center in Shen Zhen, China.

Disclosure statement

No potential conflict of interest was reported by the authors.

Additional information

Funding

The work was supported by the National Key Research and Development Program of China [2018YFC1105503].