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Deployment of a bulk metallic glass based selfexpandable stent in a patient-specific descending aorta Gideon Praveen Kumar, Mehdi Jafary Zadeh, and Fangsen Cui ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.6b00342 • Publication Date (Web): 14 Sep 2016 Downloaded from http://pubs.acs.org on September 20, 2016

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ACS Biomaterials Science & Engineering

Deployment of a bulk metallic glass based selfexpandable stent in a patient-specific descending aorta Gideon Praveen Kumar1, Mehdi Jafary-Zadeh1* and Fangsen Cui1 1

Institute of High Performance Computing, A*STAR (1 Fusionopolis Way #16-16 Connexis

Singapore 138632) E-mail: *[email protected]

KEYWORDS: self-expandable stent; bulk metallic glass; finite element analysis ; crimping; deployment; aorta; nitinol ABSTRACT: The emergence of bulk metallic glasses (BMGs) has been tantalizing in biomedical applications such as development of novel cardiovascular stents. Numerous investigations have confirmed the superior functional properties and biocompatibility of BMGs over conventional crystalline alloys as stent materials. However, a detailed understanding of the mechanical behavior of BMG-based stents during different stages of their application is still scarce. Here, by quantitative finite element analyses (FEA), we explore the deployment process of a BMG-based self-expandable stent in a patient specific descending aorta to evaluate the arterial stresses and the vessel deformation during the stent deployment. We further benchmark the performance of the BMG based stent by comparing the deployment results both with the results of a similar nitinol based stent and the experimental failure strength of the human arteries. Our detailed analyses confirm that the

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proposed BMG stent can be safely deployed in the artery without vessel over-stretching and mechanical failure, preventing unexpected vessel injuries and resultant pathological responses. Our findings would be insightful for further investigations towards realization of novel BMG-based stent applications. INTRODUCTION: As a result of continuously increasing aging population and the need for improvement of living standards, there is a prevalent demand for development of novel biomaterials to enhance the longevity and quality of human life. In the quest for novel biomaterials, both their functional and mechanical properties must be carefully evaluated to guarantee long-term biocompatibility and structural integrity of the medical device. The former (biocompatibility) addresses an important caveat which is the highly corrosive environment in the human body for many conventional materials and alloys, leading to release of toxic ions of some constitutive metallic elements1-3. Hence, biocompatibility is an essential requirement for biomaterials to avoid adverse and toxic effects on the human system. The latter (structural integrity), which poses certain requirements on the mechanical properties of biomaterials, addresses concerns regarding the complex physiological loading by surrounding tissues, blood vessel walls and bones during the operational life of any implantable medical device especially during different disease conditions. Monolithic bulk metallic glasses (BMGs), also known as “liquid-like” metals, have an amorphous microstructure, i.e. lack of atomistic long-range order4, so that they do not contain microscopic structural defects like grain boundaries and dislocations. In the absence of defect driven mechanisms of plasticity, BMGs offer extraordinary mechanical properties such as large elastic strain limit, high strength and resilience (i.e. high elastic energy storage), improved wear resistance, fatigue endurance, and excellent high temperature formability, in comparison to their crystalline counterparts. Besides, studies reporting biocompatibility of BMGs have been on the rise1,

5-8

. Another feature of BMGs which makes it suitable for


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medical device applications is their high imaging compatibility9-10. For example it has been shown that BMGs have better MRI compatibility compared to 316L stainless steel. Owing to the combination of such superior and desirable structural and functional properties, BMGs are in the cutting-edge of materials research for biotechnological applications, including stent devices.11,1, 10, 12-14. Stents are widely employed to restore the patency of narrowed or clogged blood vessel and other human conduits and as device frames in prosthetic heart valve replacement. Despite the success of conventional polycrystalline biometals as stent material, the emergence of BMGs has opened new horizons for development of competitive alternatives for different stent applications9 . The functional properties and biocompatibility of a prototypical Zr-based BMG to be used as stent material have been extensively investigated, and it is confirmed that this class of materials offers better biocompatibility than conventional crystalline alloys in both vitro cell culture and in vivo animal implantations.4, 7, 14-16 The high corrosion resistance of Zr-based BMGs is due to the passive films primarily composed of ZrO2 that are formed on the alloy’s outer surface.10,

16

On the other hand, the high strength of the BMGs greatly

enables the production of stents with thinner struts5, which contributes to its ease of deliverability and reduces the rate of restenosis. For example, in comparison to traditional 316L stainless steel stents, Zr-based BMG stents would require just one-third of the strut cross-section and yet provide in excess of five times the strength.10

However, in contrast

with numerous studies on the aforementioned functional properties of BMGs for stent applications, a complete understanding of the mechanical behavior of BMG-based stents during different stages of their delivery and service is still scarce. A recent study has reported results of finite element based fatigue analysis of a BMG-based vascular stent subjected to pulsatile fatigue loading and showed that the stent can withstand pulsatile pressure loads from blood and the vessel wall9. For the clinical application of BMGs as self-



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expandable stent materials, mechanical assessment of the stent is crucial to assure the integrity of the device during the crimping procedure. Therefore, we have recently conducted in-depth analyses on the crimpability of BMG-based stents for several percutaneous applications17 requiring self-expandable stents. Our analyses demonstrated that BMG-based self-expandable stents containing diamond shaped cells are susceptible to catastrophic failure through shear banding during the crimping stage17. We further deduced that BMGs could be used in non-diamond shaped stents like the Zenith stent graft which is a woven stent where shear banding was not observed during crimping. Figure S1 in the Supplementary Material provides a comparison of strains between our simulated nitinol- and BMG-based Zenith stents in the crimped profile, indicating that the integrity of both stents are preserved during the crimping procedure. Though crimpability is very important, where the stent is progressively compressed to a smaller diameter without failure, its real time deployment in a patient specific blood vessel is equally important which determines the final outcome of the actual surgical procedure. Alarmingly high arterial stresses and vessel stretching, post stent deployment could result in profound long-term histological response including exuberant neointimal proliferation and luminal stenosis18. Hence the aim of this work mainly is to simulate the deployment of a BMG-based Zenith TX2 TAA19 stent in a patient specific descending aorta to quantitatively evaluate the arterial stresses and the vessel deformation due to stent deployment. Since nitinol is currently the most popular self-expandable stent material20-25, we have benchmarked the performance of our BMG stent with that of nitinol. Here, it is noteworthy that computational modeling and simulation (CM&S) has been widely used as a research and development tool to support medical device applications by analysing the integrity, safety and effectiveness of the medical devices. Currently, variety of computational techniques is available such as boundary element method (BEM), finite volume method (FVM), finite difference method (FDM) and finite element method (FEM).


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As the specific application of these techniques, BEM is predominantly used for sound propagation analysis and electromagnetics26-27, FVM is used in computational fluid dynamics28, and FDM is used for certain case-based specific problems29-30. Besides, FEM has proved to be a formidable tool that can be used to analyze ubiquitous biomechanical problems including stent applications and explore the effects of design parameters on the mechanical performance of stents and their resulted stress distributions on the artery wall31-34. Hence, owing to its solid theoretical foundation and computational efficiency, most of the stent related theoretical work available in the literature has used FEM to conduct stent structural analyses which have also been experimentally confirmed in a variety of cases22, 3541

. Our FE analyses are also based on the available documentations provided by The Food

and Drug Administration (FDA or USFDA) on stent analysis42.

MATERIALS & METHODS: We used ABAQUS (v. 6.13-2) (Dassault Systemes, MA) to conduct the finite element analysis and we focused on the stent deployment where the following parts were modeled in the simulations. •

a patient-specific descending aorta;

•

stent;

•

crimper.

In the following subsections, we describe the generation of the patient specific descending aorta model, generation of the stent model, nitinol, BMG and artery material modeling and FEA settings.

Geometry models and meshes The Zenith TX2 TAA Endovascular stent Graft19, 43-45 (Cook Medical Europe, Bjaeverskov, Denmark) as shown in Fig. 1 (a) was studied. This cylindrical graft comprises of a metallic



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stent sutured to a polyester fabric with braided polyester and monofilament polypropylene suture. Since the focal point of this work is the metallic stent, the fabric was not considered in this work. CAD model of the bare metal stent used in this study is shown in Fig. 1(b). It is a woven stent with an outer diameter of 30 mm and a wire diameter is 0.3 mm. The stent was meshed with 6636 linear shear flexible B31 beam elements (Fig. 1(c)), which account for large axial strains and transverse shear strains32-33. Typically, woven stents (e.g. the BMGbased Zenith stent used in this study) are modeled using circular beam elements19, 32-33 as it has been shown that the computation time of beam element is reduced to one third of that of continuum element, where both formulations exhibit comparable results46-47. Moreover, the usage of beam elements in woven stenting simulations has also been successfully validated experimentally48. We processed the DICOM CT images of an aortic model (ascending, arch and the descending) where the geometrical information of the patient specific aorta was extracted as axial images and then imported to Mimics (Materialise, Leuven, Belgium). Segmentation, removal of secondary branches and smoothing of surfaces were done using Mimics before exporting it out as STL file. After some in-house geometric editing and clean-up the STL was then imported into ABAQUS as shown in Fig. 1 (d). 5118 quadrilateral shell elements [(general purpose with reduced integration (S4R)31] were used for modeling the artery since these elements tend to resemble the geometric characteristics of the artery wall wherein the artery dimension in the thickness direction is much smaller than that along the other directions. Moreover, the inherent plane stress assumption for shell elements helps in simplification of the equations and aids in convergence of the solution for this highly nonlinear problem, especially when combining the complex tissue material model with the nonlinear patient specific three dimensional (3D) aorta anatomy49.

Shell elements are

increasingly being used for complicated patient specific artery models to reduce the


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computational time31, 49-50. A Boolean operation was then done to remove the ascending aorta and the arch of aorta to get the patient-specific descending aorta. Since the medical images contain no information on the vessel wall thickness, we assigned the thickness of shell elements to be 2 mm51-53. Fig. 1(e) & (f) shows the final patient-specific descending aorta model and its mesh respectively. The outer diameters of the proximal and the distal ends of the descending aorta were 18 mm and 16 mm respectively. This change in the diameter is because the lumen of the aorta tapers in a way such that the cross-sectional area at the beginning of the descending aorta becomes almost 50% of the area at the aortic root54. This can be expressed by the relationship

‫ܣ‬ሺܼ ሻ = ‫ܣ‬଴ ݁ ି௞௭/௥ (1)

where k is the taper factor, r is the vessel radius, z is the axial distance of the vessel from the aortic root, and A0 is the area at the aortic root. The tapering phenomenon observed in the aorta can be evidently seen in Fig. 1(d)

Material models Nitinol Model

The nonlinear super elasticity of nitinol was modeled as a thermo-mechanical coupled superelastic plastic model implemented by an inbuilt user-defined UMAT for ABAQUS/Standard finite element solver. The model is based on an additive strain decomposition55, in which the total strain is the sum of the elastic strain, the transformation strain, and the plastic strain as follows ߝ = ߝா௅ + ߝ்ோ + ߝ௉௅ (2)



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Even though plastic strains develop as soon as the material is loaded beyond full transformation, exposing it to temperatures above it austenite finish [Af] temperature causes a thermal transformation which makes it recover the strain55-56. The transformation strain is around 6 to 8%, but the elastic strain is much smaller, ranging from 2 to 4 %. Since the transformation strains are large compared to typical elastic strains in nitinol, the material is said to be super elastic57-58. Hence unlike conventional alloys like stainless steel, in nitinol, recovery of up to 10% strains is possible59-61.

The UMAT used in this work is written with several user-defined material constants that define the stress and temperature induced phase transformations occurring between the austenite and martensite phases. Initially, nitinol exhibits austenite crystalline structure with a Young’s modulus of 40000 MPa and Poisson’s ratio of 0.33. When loaded beyond the start of transformation load which typically occurs during stent crimping, nitinol transforms from austenite to martensite. This austenite to twinned martensite transformation is driven by the resolution of shear forces, and happens within a range of stress levels that are characteristic of the material35. The transformation strain can be expresses as,

ߝ்ோ = ߙ∆ߞሺߜ‫ܨ‬/ߜߪሻ, ‫ ܨ‬ௌ ≤ ‫ ܨ ≤ ܨ‬ி

(3)

where ζ is the fraction of the martensite phase, α is a material constant which can be calibrated from uniaxial tensile tests, and F is a transformation potential. The trend is similar for the reverse transformation but at different stress levels. The intensity of the transformation is based on a stress potential law given by ∆ߞ = ݂ሺߪ, ߞሻΔ‫ܨ‬

(4)


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For the manufacturer specific nitinol20 we used in this analysis, the transformation is initiated from a transformation start-load of 449 MPa and ends at a transformation-end load of 540 MPa into a complete martensite phase with a Young’s modulus of 32000 MPa and Poisson’s ratio of 0.33. Martensite begins to transform back to austenite once unloading begins at 250 MPa which happens when the crimped stent is let to self-expand in a blood vessel where the body temperature is typically above the Af. The stent recovers to its original shape with full Austenite reversal which occurs at a transformation unloading stress of 83 MPa. BMG Model

The BMG was modelled as a linear elastic material with Young’s modulus of 80 GPa and Poisson’s ratio of 0.38 obtained from experimental results9. Indeed, using a linear elastic model to represent BMGs is reasonable since they are typically brittle and behave almost as linear elastic materials exhibiting no global plasticity during deformation up to their failure6263

. Furthermore, this material’s elastic strain, yield strength, and resilience (elastic energy

storage per unit volume) is much higher than many other conventional metals and alloys. Especially, the high resilience of BMGs plays an important role in their elastic recovery due to their high effective elastic strain which results in their spring back phenomenon64-65. Artery Model

The mechanical behavior of the descending aorta was modeled using a homogenous, isotropic hyperelastic constitutive model. The constitutive law is based on a reduced polynomial strain energy density function of third order. The 3rd order polynomial form of the strain energy density function U can be written as

ܷ = ‫ܥ‬ଵ଴ ሺ‫ܫ‬ଵ − 3ሻ + ‫ܥ‬ଶ଴ ሺ‫ܫ‬ଵ − 3ሻଶ + ‫ܥ‬ଷ଴ ሺ‫ܫ‬ଵ − 3ሻଷ

(5)



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where I1 is the first invariant of the Cauchy – Green tensor

‫ܫ‬ଵ = ߣଵଶ + ߣଶଶ + ߣଶଷ , ߣ௜ = ‫ܬ‬ଵ/ଷ ߣ௜

(6)

where λi are the principal stretches and J is the total volume ratio. The corresponding material coefficients33 for the polynomial hyperelastic model used in Eq. 5 are C10 = 0.0104673 MPa, C20 = 0.0194098 MPa and C30 = 0.0109830 MPa, which were obtained from experimental results of a realistic aorta which is the blood vessel studied in this work33.

Simulation strategy for stent deployment

The crimper was modeled as a rigid deformable body with 10100 4-mode surface elements (SFM3D4). Axial and circumferential boundary conditions were used to constrain the translational (Uz) and rotational (Uθ) degrees of freedom of the stent and artery models to allow deformation only in the radial direction. Mesh sensitivity studies were performed to ensure that all results were independent of further mesh refinements (For details, see Figure S2 in Supplementary Material and the discussions therein). The numerical analysis is nonlinear and involves large deformations which were considered while defining the step in the ABAQUS setup. Our simulation consists of two steps. •

Stent crimping - here the 30 mm stent was progressively crimped to 8 mm resembling being inserted into a 24F sheath by reducing the diameter of the crimper in a controlled manner. In this step, contact was modeled between the stent and inner surface of the crimper.


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•

Stent deployment - from the crimped configuration, the stent was allowed to reexpand by progressively letting the crimper recoil with the stent against the vessel wall. Here, an additional contact is modeled between the stent and the inner surface of the descending aorta.

We used a surface-to-surface algorithm in order to model the interactions between the different parts and used a frictionless contact between crimper inner surface and the stent and a friction of 0.0566 between the stent and the descending aorta. Fig. 2 shows the different stages of stent deployment in the patient-specific descending aorta. RESULTS & DISCUSSION: Fig. 3 (a) and (b) show the contour plots for the distribution of stresses generated in the arterial walls of descending aorta upon stent deployment made of nitinol and BMG, respectively. It is observed that in both cases there is a base level of stress generated by the overall expansion of the artery. Meanwhile, there are higher stresses of 0.26 MPa and 0.33 MPa due to nitinol and BMG stent deployment respectively, near the contact interfaces where there is stretching of the tapered aorta caused by the expansion of the more uniform circular stent. This is intuitive since there is a large radial deformation against the artery wall. This is because, after crimping, typically the stent is allowed to “self-expand” progressively inside the artery driven by its own stored strain energy. It expands until it reaches an equilibrium state as a function of the vessel’s elasticity. More specifically, upon stent-vessel contact, the vessel begins to expand due to the stent’s outward force till the stent and the vessel reach a stress equilibrium while the stent is in compression and the vessel is in tension25. Once a stress equilibrium state is reached, typically, the stent takes the shape of the vessel as its final form, thereby, getting embedded in the vessel musculature and securing good anchorage and vessel apposition. Fig. 3(c) shows the deformed stent which has taken the shape of the descending aorta which is similar for both nitinol and BMG-based stents. This phenomenon of the stent getting embedded into the vessel musculature is called “stent



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penetration”67. In other words, penetration is the peak radial displacement of the artery on its inner surface due to contact between the stent and the artery after stent deployment. Once the stress equilibrium is reached, penetration can be calculated as the difference between the artery initial inner diameter and the deformed stent outer diameter. Our results showed that penetration induced by the nitinol and BMG stents were 0.32 mm and 0.36 mm, respectively. This difference in penetration can be attributed to the fact that though both nitinol and BMG stents have the same initial (crimped) geometries, their intrinsic mechanical properties in terms of elasticity and deployment mechanisms are different. Nitinol has a unique elastic hysteresis as a result of its martensite to austenite transformation that allows the stent to reexpand upon deployment onto the vessel wall 35, 68-70 . In this case, the stent pushes outwards and positions itself against the artery wall with a low chronic outward force [COF]71. On the other hand, in the case of the BMG stent, the penetration is a result of intrinsic pure elasticity (with no phase transformation) of the stent behaving in analogy with the push-back behavior of a spring released from the compressed state. Here, we recall that BMGs are exceptionally high resilient materials in general, which substantially allows them to store higher amount of elastic energy per unit volume than any other 1500 conventional engineering metals and alloys.65, 72 This implies that to obtain a certain amount of penetration upon deployment, a smaller stents (and hence less material) is needed if the stent is made of metallic glasses rather than other conventional alloys. Extensive work has been done on the deployment of self-expandable33, 68-70 stents in human arteries. However, based on the best of our knowledge, there is a lack of information about direct comparison of the axial and circumferential stresses with experimental results. Researches mostly have just reported the maximum stresses observed in the artery upon stent deployment. In this work, we compare our simulated axial and circumferential stresses in the patient-specific descending aorta post stent implantation directly with existing experimental


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data73 for the ultimate strength of human artery73. This comparison is crucial since during and after stent deployment there could be over stretching and mechanical failure of the arteries which in turn could lead to vessel injuries as well as profound long-term histological response including exuberant neointimal proliferation18. Hence, keeping the vessel stresses below its ultimate strength is imperative73-75. The histogram depicted in Fig. 3(d), compares the axial and circumferential stresses in the aorta due to the deployment of modeled nitinol and BMGbased stents with the ultimate failure strength of the human artery. This figure clearly demonstrates that both the axial and circumferential stresses due to the BMG stent deployment are comparable with those of nitinol stent deployment, and are much lower than the failure limit of human arteries. Hence, it can be concluded that upon careful design, the BMG-based stents can to be used safely during the crimping and deployment procedures without their own mechanical failure and without causing any injury to the blood vessel, implying that BMGs are promising candidates for future stent applications. It is noteworthy that while the material models used in this work to simulate BMG9, nitinol76 and the human artery33 are all obtained from experimental results that are already available in the literature, further experimental studies to evaluate our FE data in real life clinical applications, is our ongoing research as a progression of the current work.

SUMMARY: The objective of this work was to check the feasibility of using BMGs as stent material for self-expandable stent applications. Stent deployment, which is an important process that decides the outcome of the surgical process, was considered here. Finite element modeling and simulation of woven stent deployment in a patient-specific descending aorta were presented. A prototypical Zr-based BMG was used as the stent material and benchmarked against nitinol since currently, nitinol is the most popular material for selfexpandable stent applications. Nonlinear FEA was used to determine the stresses in the



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vessel during and after stent deployment. The obtained simulation results were compared with experimental data for ultimate failure strength of artery vessels. The effect of stent material on the arterial stresses and the stent penetration in the artery were demonstrated. We found that arterial stresses and penetration as a result of a BMG stent deployment was comparable with those of nitinol stent. Additionally, the induced stresses were lower than the failure limit of the arteries which is an encouraging finding of this study. Further to our previous details of crimping analysis, our simulation results of deployment indicate that BMGs can be safely used as a stent material for developing woven stents as they can be crimped and deployed without much difficulty which is very important for the success of any percutaneous procedure. As the progression of the current work, ongoing research on design optimization of BMG-based woven stents is proposed in order to exploit the mechanical properties of BMGs and obtain better performance than nitinol based stents. Our work lays the platform to come up with new BMG based stent designs that could be used for various percutaneous applications requiring self-expanding stents.

ACKNOWLEDGEMENTS: The authors gratefully acknowledge the financial support from the Agency for Science, Technology and Research (A*STAR), Singapore and the use of computing resources at the A*STAR Computational Resource Centre, Singapore.


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Figure 1 CAD geometries and meshes used in this study. (a) Zenith TX2 TAA Endovascular Graft (www.cookmedical.com); (b) Isometric view of the CAD model of the bare metal wire based stent; (c) Stent meshed with beam elements; (d) CAD model of the patient-specific aorta; (e) CAD model of the region of interest, i.e. the patient-specific descending aorta; (f) Patient-specific descending aorta meshed with quadrilateral shell elements.


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Figure 2 Steps involved in the simulation of the deployment process (a) The crimper, stent and the artery are assembled in a cylindrical co-ordinate system; (b) The crimper is used to crimp the stent, where its diameter is reduced from 30 mm to 8 mm; (c) The stent is let to expand into the vessel till a stress equilibrium is achieved between the stent and the vessel.



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Figure 3 Vessel and stent configuration after deployment (a) Contour plot of von Mises stress distribution (S, von Mises) in the post deployment descending aorta when the stent material is nitinol; (b) Contour plot of von Mises stress distribution (S, von Mises) in the post deployment descending aorta when the stent material is BMG; (c) The deformed stent which has taken the shape of the vessel after full deployment; (d) Histogram comparing the axial and circumferential stresses in the artery after stent deployment with nitinol and BMG as the stent material against experimental test results on the failure strength on human arteries.



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For Table of Contents Use Only Deployment of a bulk metallic glass based self-expandable stent in a patientspecific descending aorta Gideon Praveen Kumar1, Mehdi Jafary-Zadeh1* and Fangsen Cui1 1

Institute of High Performance Computing, A*STAR, Singapore 138632


Deployment of a Bulk Metallic Glass-Based Self ... - ACS Publications

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