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A new flow diverter stent for direct treatment of intracranial aneurysm
Author’s Accepted Manuscript A New Flow Diverter Stent for Direct Treatment of Intracranial Aneurysm Jiayao Ma, Zhong You, Thomas Peach, James Byrne, Rafik R. Rizkallah www.elsevier.com/locate/jbiomech
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S0021-9290(15)00572-2 http://dx.doi.org/10.1016/j.jbiomech.2015.10.024 BM7384
To appear in: Journal of Biomechanics Received date: 17 June 2015 Revised date: 10 October 2015 Accepted date: 18 October 2015 Cite this article as: Jiayao Ma, Zhong You, Thomas Peach, James Byrne and Rafik R. Rizkallah, A New Flow Diverter Stent for Direct Treatment of Intracranial Aneurysm, Journal of Biomechanics, http://dx.doi.org/10.1016/j.jbiomech.2015.10.024 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting galley proof before it is published in its final citable form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
A New Flow Diverter Stent for Direct Treatment of Intracranial Aneurysm
Jiayao Ma, Zhong You1, Thomas Peach, Ph.D. Department of Engineering Science, University of Oxford, UK
James V. Byrne, Rafik R. Rizkallah, M.D. Nuffield Department of Surgical Sciences, University of Oxford, UK
Article Type: Original Article Word count: 3913 1
Corresponding author: Prof. Zhong YOU, Department of Engineering Science, University of
Oxford, 17 Parks Road, Oxford, OX1 3PJ, Tel.: +44 1865 273137, Fax: +44 1865 283301, Email: [email protected].
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A New Flow Diverter Stent for Direct Treatment of Intracranial Aneurysm
Jiayao Ma, Zhong You1, Thomas Peach, Ph.D. Department of Engineering Science, University of Oxford, UK
James Byrne, Rafik R. Rizkallah, M.D. Department of Neuroradiology, Nuffield Department of Surgery, University of Oxford, UK
Abstract: The use of a stand-alone flow diverter (FD) stent has demonstrated itself as an efficacious endovascular approach to intracranial aneurysm treatment. FD stents that are currently available adopt an interwoven braided design. The relatively low radial stiffness intrinsic to this design could cause difficulty in deployment and poor stent-wall opposition, leading to high complication rates. A new FD stent is proposed to overcome the problems of the interwoven FD stents. The new device is manufactured from a Nitinol tube through a laser-cutting technique, and its unique structure allows for both low porosity and high packaging efficiency. Computational simulation using Abaqus has been conducted to investigate the radial stiffness and longitudinal flexibility of the new device. The new device exhibits high radial stiffness when compared to interwoven FD stents and superior longitudinal flexibility. Results from on-going in-vivo experiments and CFD simulations have also demonstrated the efficacy of the new device as a FD stent.
Key words: intracranial aneurysm; flow diverter stent; radial stiffness; longitudinal flexibility.
1
Introduction
Intracranial aneurysms, a balloon-like localised dilatation on the wall of an artery within the intracranial vasculature, remain an important cause of stroke morbidity and mortality (Bederson et 1
Corresponding author. E-mail: [email protected]. 2
al., 2009). As a safer and more efficacious alternative to conventional treatment methods such as surgical clipping (Dandy, 1938), coil embolization (Guglielmi, 2009), and stent-assisted coiling, the use of a low porosity stent alone as a flow diverter (FD) has gained substantial popularity in the last 10 years or so (Pierot, 2011). The treatment goal of using a FD stent is to correct the hemodynamic disturbance in the aneurysm, and eliminate what is considered to be the main cause of aneurysm progression and rupture (Cebral et al., 2005; Tateshima et al., 2009). A FD stent is placed across the neck of aneurysm sac to reduce blood flow into aneurysm sac, isolating the aneurysm by inducing stable thrombosis. Subsequently biological repair of the aneurysm neck by endothelial overgrowth of the stent mesh occurs to reconstruct the diseased segment of the parent artery (Kadirvel et al., 2014). It has been demonstrated that porosity is crucial for the functionality of FD stents and a porosity threshold of around 70% is the considered the optimum trade-off between device flexibility and efficacy (Liou et al., 2007; Seong et al., 2007). First introduced in 2007, FD stents are considered suitable for treating wide-neck, giant and fusiform or tiny lesions lateral to a large artery, carotid siphon or vertebrobasilar artery (Cirillo et al., 2012). Comprehensive reviews of the current status of FD stents are available in literature (Alderazi et al., 2014; Byrne and Szikora, 2012; Pierot). The FD stents currently available in clinical practice, e.g., the SILK (Balt Extrusion, Montmorency, Frnace) and PED (Coviden, Irvine, California, US), adopt an interwoven structural design, which are manufactured by braiding superelastic wires into a tubular mesh. This design suffers a low radial stiffness that could cause difficulty in endovascular deployment, spontaneous delayed migration or shortening (Chalouhi et al., 2013), and poor stent-wall opposition, risking extra-aneurysmal thrombosis and stroke. Partially due to the low radial stiffness, treatment by FD is associated with relatively high complication rates in comparison with other treatment methods (Byrne et al., 2010; Cirillo et al., 2012). In addition,
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even modest stent oversizing frequently seen in clinical practice could lead to substantial increase in stent porosity, which is also undesirable for a FD stent (Shapiro et al., 2014).
We previously proposed a FD stent design, aimed at providing desirable mechanical properties to enable high device efficacy and deliverability (Ma et al., 2014). The device is manufactured out of a Nitinol tube through a laser-cutting technique. As a further development of the work conducted by the authors, this paper focuses on computational analysis of the key mechanical properties of a new design including the radial stiffness, which ensures the complete opening of the stent in the parent arteries, good stent-wall opposition and secure anchorage, and the longitudinal flexibility, which allows for guidance through tortuous intracranial arteries and good conformability to parent arteries. The layout of the paper is as follows. Section 2 introduces the geometrical design of the new FD stent and its variants. A detailed numerical modelling methodology for the new FD is presented in Section 3. The numerical results are summarized in Section 4, in which the radial stiffness and longitudinal flexibility of the new FD stent are discussed and compared with existing FD stents and intracranial stents. Preliminary results from an on-going in-vivo experiment aare nd the results of CFD simulations are reported in Section 5 to demonstrate the deliverability and efficacy of the new device. Finally, Section 6 concludes the paper by summing up the main findings drawn from the work.
2
The geometrical design
A prototype of the new FD stent, which is manufactured out of a Nitinol tube by laser-cutting technology and then post-processed (Ma et al., 2014), is presented in Fig. 1. The new FD stent has a middle portion of low porosity and two ends of high porosity. The middle portion is to be placed
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across the neck of the aneurysm so as to reduce the volume of blood entering the aneurysm; hence a lower porosity is required in this region. The central section is formed from a series of repetitive modules, which consist of longitudinal waves connected by V-shaped connectors. Compared with the early version of the new FD stent (Ma et al., 2014), the introduction of the V-shaped connectors allows for fine tuning of radial stiffness, which will be demonstrated in the subsequent numerical analysis. Meanwhile, the longitudinal waves adopted in the early version are preserved to maintain the superior longitudinal flexibility and to achieve low porosity. The porosity of the middle portion of the device is tuneable by replacing some straight edges with curved ones, therefore generating two designs, referred to as type I and type II respectively, as presented in Fig. 2(a) and (b). Commercially available interwoven FD stents usually have a nominal porosity around 70% (Gross and Frerichs, 2013), which can be readily replicated by the type II design. The high-porosity ends of the device are designed for convenient packaging of the new FD stent and to accommodate radiopaque markers. In addition, the device ends are less stiff radially than the middle portion so as to achieve a smooth transition of force from the implanted device to the arterial wall. As a result, the strain exerted on the parent artery is reduced.
The middle portion of the new FD stent is critical for the functionality of the device as it determines the porosity, radial stiffness and longitudinal flexibility. The geometry of the middle portion of the new FD stent is shown in Fig. 2(c) and (d), and can be defined by seven parameters: the initial fully-expanded radius of the stent R0, the inner radius of the longitudinal wave r1, the inner radius of the V-shaped connector r2, the ratio of the height of the longitudinal wave to that of the sum of the longitudinal wave and the V-shaped connector a1/(a1+a2), the inner angle of the V-shaped connector α, the strut width w, and the wall thickness t. The design of the middle
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portion of the device can be of either type I or type II, allowing for variation in device porosity ρ.
For a FD stent to be delivered into the small and torturous intracranial arteries, it is essential for the device to be packaged with a very small radial profile. The new FD stent can be radially folded through a two-step process. First, the device is stretched longitudinally from its initial state, Fig. 3(a), to an intermediate state, Fig. 3(b). Only the longitudinal waves are deformed at this step to reduce the device radius. Secondly, the stretched stent is radially crimped to fold the V-shaped connectors, resulting in further reduction in radius at the final state, Fig. 3(c). Upon deployment, the longitudinal waves spring back and the V-shaped connectors are unfolded to restore the initial fully expanded state of the stent.
Fourteen numerical models, including eleven models of type I and three of type II, were built to analyse the effects of stent geometry on its mechanical properties, and specifically the radial stiffness and the longitudinal flexibility. The parameters R0, r1, w and t of all the models were selected as 1.9mm, 0.3mm, 0.05mm, and 0.05mm, respectively, whereas a1/(a1+a2), α, r2, and ρ were systematically varied. The geometries of all the models are listed in Table 1. Note that the porosity of model A4, the design of which is presented in Fig. 2(d), reaches 70.2% matching that of commercially available woven FD stents.
3
Finite element modelling
As previously discussed, radial stiffness and longitudinal flexibility are key mechanical properties for a FD stent, and correspondingly form the basis of discussion in this study. Due to the repetitive nature of the new FD stet design, only a representative section of the middle portion is
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analysed. However, the offset in longitudinal spacing between adjacent V-shaped connector rows in the design makes it difficult to extract a “single module” from the middle portion of the device, and thus, a section with 7 and 8 V-shaped connectors in adjacent rows and a total length l of 6mm, Fig. 2(c), is selected for modelling. The loading scenario is shown in Fig. 4(a): the FD stent is compressed radially by 12 identical rigid cylindrical shells from an initial radius of 1.9mm to 1.4mm, and the radial force per unit length fθ is calculated by the following equation (Zhou, 2009):
fθ
F 2 l
(1)
in which ΣF is the scalar summation of the radial reaction forces of all the cylindrical shells.
Four models, A1, A4, B1, and B3, are bent to investigate the longitudinal flexibility. Both ends of the device, as shown in Fig. 2(a), are attached to each model to form a complete stent. A1 is shown in Fig. 4(b) as an instance. The lengths of A1, A4, B1 and B3 are 15.45mm, 15.45mm, 15.50mm, and 15.65mm, respectively. Each model is subjected to pure bending to reach a curvature of 0.1mm-1, and the bending moment at each end of the device is recorded during the deformation process.
Numerical simulation was conducted using finite element analysis software package ABAQUS/Standard. Superelastic Nitinol, commonly used for intracranial stents and FD stents, was adopted as the material, the properties of which are listed in Table 2 (Rebelo et al., 2006). The details of the numerical modelling are described in the appendix.
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4 4.1
Results Properties of A1
The results for model A1 are presented in detail and can be considered representative of all designs. The von Mises stress contour of A1 subjected to radial compression is shown in Fig. 4(c). It is clear from the contour that the V-shaped connectors are folded substantially to accommodate the reduction in radius of the model. In contrast, the deformation of the longitudinal waves is quite small. Moreover, as would be expected from the deformation mode, high stress concentrations are present at the curved tips of the V-shaped connectors, whereas the stress in the straight sections subject to little bending is small. The location of such stress concentrations suggest that the V-shaped connectors are mainly responsible for the radial stiffness of the FD stent under compression. The radial force per unit length fθ of A1 is plotted against FD stent radius R in Fig. 5(a). It can be seen that due to the introduction of the V-shaped connectors, fθ gradually increases with the reduction in R, a desirable feature for a FD stent.
The von Mises stress contour of A1 subjected to pure bending is presented in Fig. 4(d). A uniform mesh deformation distribution is achieved along the longitudinal length of the model, thus providing homogeneous mesh coverage the aneurysm neck. In addition, large stress only occurs at the curved portion of the longitudinal waves, whereas the rest of the model is subjected to very low stress. This result indicates that the longitudinal flexibility of the FD stent is predominantly 8
determined by the longitudinal waves and the effect of V-shaped connectors is insignificant. The bending moment of A1 is plotted against its curvature in Fig. 5(b). A near-linear relationship between bending moment and curvature is obtained.
4.2
Effects of geometrical parameters on radial stiffness
It is clear from the result of A1 that a significant stress distribution is present throughout the device geometry under a loading scenario of radial compression, hence it is expected that all of the geometric parameters previously discussed will affect the radial stiffness of the new FD stent.
Models A1-A4, all of which have identical a1/(a1+a2), α, and r2 but varying design type (and therefore porosity ρ) are first analysed to investigate the effects of porosity on device stiffness. The fθ vs. R curves of the four models are plotted in Fig. 5(a). It can be seen that the radial force per unit length increases mildly with the reduction in porosity. At a radius of 1.615mm, corresponding to 15% stent oversizing, fθ increases from 0.019N/mm to 0.025N/mm (31.6%) when porosity reduces from 86.3% in A1 to 70.2% in A4. Moreover, fθ converges quickly when the porosity is below 80%: the curves of A3 and A4, which have a difference in porosity of 6.5% are almost identical. This result indicates that within the practical range of porosity, 70%-80%, the radial stiffness of the device varies only marginally with porosity. Hence, porosity does not need to be taken into consideration when tuning the radial stiffness of the new FD stent design. 9
Models B1-B3, differ from each other in terms of the ratio a1/(a1+a2) alone, with all other parameters remaining constant. It should be noted that when a1/(a1+a2) reaches 1, the V-shaped connectors in the design disappear and the device is reduced to the early FD stent design (Ma et al., 2014). When the value is 0, the design becomes a conventional zigzag stent. The fθ vs. R curves of the three models B1-B3 are plotted together with that of A1 in Fig. 5(c). The plot shows that fθ increases significantly with a1/( a1+a2), thus providing an effective way of tuning the radial stiffness of the new FD stent.
The stress contour of A1 in Fig. 4(c) indicates that a high stress concentration occurs at the tip of the V-shaped connectors under deformation, therefore, the α and r2 parameters of the connectors are expected to significantly affect the device radial stiffness. The fθ vs. R curves of models C1-C4 with increasing α are first plotted together with that of A1 in Fig. 5(d). Two observations can be made. First, fθ increases with α, i.e., a more obtuse angle leads to a higher force. Second, fθ is quite sensitive to α, and a wide range of fθ can be achieved by varying α. The fθ vs. R curves of models D1-D3, which have increasing values of r2 are plotted together with A1 in Fig. 5(e). It is clear that fθ also increases with r2. In contrast to α, however, fθ is much less sensitive to r2, and therefore only a small range of fθ is achievable.
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4.3
Effects of geometrical parameters on longitudinal flexibility
The stress contour of A1 subjected to pure bending, shown in Fig. 4(d), indicated that the longitudinal waves of the design undergo significant deformation and play a major role in determining device’s longitudinal flexibility. The contribution to longitudinal stiffness from the V-shaped connectors, which undergo very small deformation under pure bending, is considered insignificant. Therefore only variation in design type and a1/(a1+a2), which are associated with the geometry of longitudinal waves, are analysed to evaluate their respective effects on longitudinal flexibility. The bending moment vs. curvature curves of A1 and A4, which differ only in design type and hence porosity, are plotted in Fig. 5(b). It is clear from the figure that a large reduction in porosity from 86.3% in A1 to 70.2% in A4 only very slightly increases the device bending stiffness. Hence, the effect of design type on longitudinal flexibility can be practically ignored. The bending moment vs. curvature curves of A1, B1 and B3, with varying a1/(a1+a2), are plotted in Fig. 5(f). The figure clearly shows that the larger a1/(a1+a2) ratio, the lower the bending moment, which again demonstrates that longitudinal waves are mainly responsible for the longitudinal flexibility of the new FD stent. In short: the larger the proportion of longitudinal waves in the stent structure, the lower the bending moment required to deform the device, and the more flexible the FD stent.
4.4
Comparison with existing intracranial stents and FD stents
The radial stiffness and longitudinal flexibility can significantly affect the efficacy of a FD stent (Liou et al., 2007; Seong et al., 2007). However, there is no accepted design guidance for either parameter for a successful FD device. Consequently, the approach adopted here is to evaluate the mechanical properties of the new FD stent in comparison to the properties of a number of
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currently available devices. Both intracranial stents that are also laser-cut from Nitinol tubes, including Neuroform3, Enterprise, Wingspan and Solitaire (Krischek et al., 2011), and interwoven FD stents like the PED (Fischer et al., 2012) are considered. Model A4 is selected for the comparison as it has a low porosity, which matches that for commercial interwoven FD stents.
Figure 6(a) presents the value of fθ for each design of intracranial stent with a 15% stent oversizing in all cases. It can be seen that A4 has a relatively high fθ, 11.6% larger than that of Wingspan. This result indicates that the new FD stent is able to provide adequate radial stiffness that exceeds current intracranial stents. Should a lower radial stiffness be required, it can be easily achieved by geometrical modification of the design, such as by reducing α. Note that fθ for PED is about 0.002N/mm (Fischer et al., 2012), less than 10% of that for the new device.
The bending moments of each design of intracranial stent when bent into a 12.7mm arc are plotted in Fig. 6(b). It is clear from the figure that A4 requires the smallest bending moment, only 10.7% of that for Neuroform3. In other words, the new FD stent has the highest longitudinal flexibility, making it ideal for maneuvering through the tortuous arteries of the cerebral vasculature. Data for the longitudinal flexibility of the PED are not available and thus no comparison is made.
5 5.1
In vivo animal experiment and hemodynamic simulation In vivo animal experiment
Rabbit models with elastase-induced arterial saccular aneurysms (Sadasivan et al., 2009) were used to test the efficacy of the new device. A 5F (1.67mm) delivery system was designed to deploy the
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new FD stent conveniently and accurately. A sketch of the design is presented in Fig. 7(a) and a prototype is shown in Fig. 7(b). The new device was loaded at the tip of the system and deployed through an unsheathing process. Stent samples of 1.75mm radius and 14mm length were loaded into the system after two radiopaque markers made from platinum were installed on each end of the stents to improve visibility under X-ray fluoroscopy.
A typical result of one experimental aneurysm before, immediately after, and three months after stenting is shown in Fig. 8. It can be seen from Fig. 8(b) that, first of all, the new FD stent fully opened in the parent artery and conformed very well to the artery wall, and secondly, an immediate isolation of the aneurysm was achieved. The angiography image obtained from the three-month follow-up of the rabbit, which is shown in Fig. 8(c), confirms that no refilling of the aneurysm occurred and that the aneurysm was completely occluded.
5.2
Hemodynamic simulation
A representative aneurysm geometry was extracted from experimental 3D angiographic imaging and used to quantify the flow-diverting effect of the new FD stent. The new device was virtually deployed into the rabbit aneurysm geometry assuming a rigid vessel wall (Peach et al., 2014b). Blood was modelled as an incompressible Newtonian fluid with a density of 1000 kg/m3 and a dynamic viscosity of 0.004 PaS. The 3D unsteady Navier-Stokes equations were solved using the finite volume method using CFD-ACE multiphysics package (ESI Group, Paris, France) with a Central Differencing scheme for spatial variations and a Crank-Nicholson scheme for time marching. The SIMPLE-Consistent (SIMPLEC) pressure correction method and an algebraic multigrid method for convergence acceleration were used (Van Doormaal and Raithby, 1984;
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Webster, 1994).
The aneurysm geometry with and without the device deployed was meshed using a Projected Single Domain mesh, an Omnitree Cartesian tree type, and three near-wall Cartesian layers. Mesh sizes were between 1.67 and 2.81 million elements, assuming a mesh-independent cell density (>4,000 elements/mm3) as discussed in previous works and similar studies (Cebral et al., 2014a; Peach et al., 2014a; Peach et al., 2014b). Time-step independence was also assumed at the chosen time step of 0.01 seconds, in line with similar studies in the literature (Cebral et al., 2014a). The transient inlet boundary condition of volumetric parent vessel flow, in addition to velocity profile and pulse rate, was estimated from the experimental results reported by Cebral et al. (Cebral et al., 2014a).
The results of the CFD simulations indicated that the new FD stent reduces the mean velocity of blood entering the aneurysm dome by a factor of 2.54. Such a reduction is comparable to the mean velocity reduction of 2.64 reported by Cebral et al. (Cebral et al., 2014b) for rabbit model aneurysms treated with the PED, indicating that the flow-diverting effect of the new FD stent is comparable to commercial interwoven FD stents.
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Conclusions
A new FD stent that can be laser-cut from a Nitinol tube, for direct treatment of intracranial aneurysms, has been proposed in this paper. A unique pattern design for the new FD stent, i.e., four longitudinal waves interconnected by V-shaped components, provides high radial stiffness and longitudinal flexibility, as well as convenient packaging and delivery. A low porosity of 70% can be achieved for the new FD stent to effectively divert blood flow from aneurysm sac and
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promote aneurysm thrombosis.
The mechanical properties, specifically the radial stiffness and longitudinal flexibility, of the new FD stent, were numerically analysed and compared with those of several commonly used intracranial stents and commercially available interwoven FD stents. It has been found that the radial stiffness of the new FD stent is an order of magnitude higher than that of the PED and is comparable to that of commonly used intracranial stents. The longitudinal flexibility of the new FD stent, on the other hand, is considerably higher than that of commonly used intracranial stents. In one particular case, model A4 has low porosity of 70.2%, which matches that of commercially available interwoven flow diverter stents; high radial stiffness, which is over 10 times larger than that of commercially available flow-diverter stents and comparable to those of commonly used intracranial stents; and superior longitudinal flexibility, when compared with commonly used intracranial stents. Both in vivo experimental results and CFD simulations show that the new FDS performs well as a flow-diverter and achieves performance comparable to current woven flow-diverters, such as the PED. Such a device therefore serves as an excellent candidate for a new flow-diverter stent.
Study limitations A limitation of the hemodynamic simulations conducted in this study is the use of estimated inflow boundary conditions, although previous studies have shown such estimates to produce reasonable results (Cebral et al., 2014a). Additionally the highly dynamic environment of the rabbit aorta can introduce error in angiographic imaging and subsequent 3D reconstructions. A number of factors including device size, wire diameter, and device pore shape and size are likely
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to influence the flow-diverting effect of a given stent design. Due to obvious constraints, the effects of these parameters have not been fully explored in this study; hence, the sensitivity of the device performance to these parameters is not known. Regarding the elastase induced saccular aneurysms rabbit model, one limitation is that the aneurysm was created in the extracranial carotid artery, and thus was not an intracranial aneurysm. This is attributable to the size constraints of working with rabbits, as well as the challenging nature of controlling the elastase effect on rabbit cerebral circulation. Nearly all previously described animal aneurysm models involved extracranial arteries (Frösen et al., 2006). As the purpose of this model was intended to study the hemodynamic effects of the flow-diverter stents, as well as the cellular interaction of the new device with the parent artery, the extra-cranial carotid elastase induced saccular aneurysm model was appropriate.
Conflicts of interest The authors declare that there are no conflicts of interest.
Acknowledge The authors would like to thank Wellcome Trust and EPSRC for their financial support under grant number WT 088877/Z/09Z, Dr M. Megahed of the ESI Group for allowing the use of the CFD-ACE suite, Dr. Y. Ventikos, and the use of the Oxford Advanced Research Computing (ARC) facility in carrying out this work.
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functional and physical properties of self-expanding intracranial stents [Neuroform3, Wingspan, Solitaire, Leo+, Enterprise]. Minimally Invasive Neurosurgery 54, 21-28. Liou, T.M., Li, Y.C., Juan, W.C., 2007. Numerical and experimental studies on pulsatile flow in aneurysms arising laterally from a curved parent vessel at various angles. Journal of biomechanics 40, 1268-1275. Ma, J., You, Z., Byrne, J., Rizkallah, R., 2014. Design and Mechanical Properties of a Novel Cerebral Flow Diverter Stent. Annals of Biomedical Engineering 42, 960-970. Peach, T., Cornhill, J.F., Nguyen, A., Riina, H., Ventikos, Y., 2014a. The 'Sphere': A Dedicated Bifurcation Aneurysm Flow-Diverter Device. Cardiovascular engineering and technology 5, 334-347. Peach, T.W., Ngoepe, M., Spranger, K., Zajarias-Fainsod, D., Ventikos, Y., 2014b. Personalizing flow-diverter intervention for cerebral aneurysms: from computational hemodynamics to biochemical modeling. International journal for numerical methods in biomedical engineering 30, 1387-1407. Pierot, L., 2011. Flow diverter stents in the treatment of intracranial aneurysms: Where are we? Journal of Neuroradiology 38, 40-46. Rebelo, N., Gong, X.-Y., Hall, A., Pelton, A.R., Duerig, T.W., Year Finite element analysis on the cyclic properties of superelastic nitinol. In Proceedings of the International Conference on Shape Memory and Superelastic Technologies. Sadasivan, C., Cesar, L., Seong, J., Wakhloo, A.K., Lieber, B.B., 2009. Treatment of rabbit elastase-induced aneurysm models by flow diverters: development of quantifiable indexes of device performance using digital subtraction angiography. Medical Imaging, IEEE Transactions 28, 1117-1125.
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Seong, J., Wakhloo, A.K., Lieber, B.B., 2007. In vitro evaluation of flow diverters in an elastase-induced saccular aneurysm model in rabbit. Journal of Biomechanical Engineering 129, 863-872. Shapiro, M., Raz, E., Becske, T., Nelson, P.K., 2014. Variable Porosity of the Pipeline Embolization Device in Straight and Curved Vessels: A Guide for Optimal Deployment Strategy. American Journal of Neuroradiology 35, 727-733. Tateshima, S., Tanishita, K., Hakata, Y., Tanoue, S.-y., Viñuela, F., 2009. Alteration of intraaneurysmal hemodynamics by placement of a self-expandable stent: laboratory investigation. Journal of neurosurgery 111, 22-27. Van Doormaal, J.P., Raithby, G.D., 1984. Enhancements of the SIMPLE method for predicting incompressible fluid flows. Numerical heat transfer 7, 147-163. Webster, R., 1994. An algebraic multigrid solver for Navier-Stokes problems. International Journal for Numerical Methods in Fluids 18, 761-780. Zhou, X., 2009. Design of Stents for the Direct Treatment of Intracranial Aneurysms. University of Oxford, Oxford.
20
Captions of Figures Fig. 1
A physical sample of the new FD stent
Fig. 2
(a) pattern design type I for the new FD stent, (b) pattern design type II for the new FD stent, (c) geometrical parameters definition for the new FD stent, and (d) design of model A4
Fig. 3
Folding process of the new FD stent, (a) initial fully expanded state, (b) intermediate state after longitudinal elongation, and (c) final fully folded state after radial compression
Fig. 4
(a) Radial compression setup, (b) a complete A1 for pure bending, (c) von Mises stress contour of A1 subjected to radial compression, and (d) von Mises stress contour of A1 subjected to pure bending
Fig. 5
(a) fθ vs. R curves of A1-A4, (b) bending moment vs. curvature curves of A1 and A4, (c) fθ vs. R curves of A1 and B1-B3, (d) fθ vs. R curves of A1 and C1-C4, (e) fθ vs. R curves of A1 and D1-C3, and (f) bending moment vs. curvature curves of A1, B1, and B3.
Fig. 6
Comparison of the stents’ mechanical properties, (a) fθ at 15% stent oversizing, and (b) bending moment needed to bend the stents into a 12.7mm arc
Fig. 7
(a) Sketch of the delivery system, and (b) a prototype of the delivery system
Fig. 8
Animal experiment results, (a) before FD stent deployment, (b) immediately after FD stent deployment, and (c) three month follow-up
Captions of Tables Table 1
Geometries of the new FD stent models
Table 2
Material properties of superelastic Nitinol
21
22
23
24
25
Table 1 Model type a1/(a1+a2) α (mm) r2 (mm)
A1
I
0.5
60
0.125
86.3%
A2
II
0.5
60
0.125
81.7%
A3
II
0.5
60
0.125
76.7%
A4
II
0.5
60
0.125
70.2%
B1
I
0.45
60
0.125
86.2%
B2
I
0.55
60
0.125
86.3%
B3
I
0.6
60
0.125
86.3%
C1
I
0.5
40
0.125
84.1%
C2
I
0.5
50
0.125
85.4%
C3
I
0.5
70
0.125
86.8%
C4
I
0.5
80
0.125
87.3%
D1
I
0.5
60
0.0875
86.2%
D2
I
0.5
60
0.1625
86.4%
D3
I
0.5
60
0.2
86.5%
Table 2 Property Value
Definition
EA
50 GPa
Austenite elasticity
EM
37 GPa
Martensite elasticity
Ms
400 MPa Starting transformation stress of loading
Mf
650 MPa End transformation stress of loading
As
350 MPa Starting transformation stress of unloading
Af
80 MPa
End transformation stress of unloading
L
0.055
Maximum residual strain
26
27
PII: DOI: Reference:
S0021-9290(15)00572-2 http://dx.doi.org/10.1016/j.jbiomech.2015.10.024 BM7384
To appear in: Journal of Biomechanics Received date: 17 June 2015 Revised date: 10 October 2015 Accepted date: 18 October 2015 Cite this article as: Jiayao Ma, Zhong You, Thomas Peach, James Byrne and Rafik R. Rizkallah, A New Flow Diverter Stent for Direct Treatment of Intracranial Aneurysm, Journal of Biomechanics, http://dx.doi.org/10.1016/j.jbiomech.2015.10.024 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting galley proof before it is published in its final citable form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
A New Flow Diverter Stent for Direct Treatment of Intracranial Aneurysm
Jiayao Ma, Zhong You1, Thomas Peach, Ph.D. Department of Engineering Science, University of Oxford, UK
James V. Byrne, Rafik R. Rizkallah, M.D. Nuffield Department of Surgical Sciences, University of Oxford, UK
Article Type: Original Article Word count: 3913 1
Corresponding author: Prof. Zhong YOU, Department of Engineering Science, University of
Oxford, 17 Parks Road, Oxford, OX1 3PJ, Tel.: +44 1865 273137, Fax: +44 1865 283301, Email: [email protected].
1
A New Flow Diverter Stent for Direct Treatment of Intracranial Aneurysm
Jiayao Ma, Zhong You1, Thomas Peach, Ph.D. Department of Engineering Science, University of Oxford, UK
James Byrne, Rafik R. Rizkallah, M.D. Department of Neuroradiology, Nuffield Department of Surgery, University of Oxford, UK
Abstract: The use of a stand-alone flow diverter (FD) stent has demonstrated itself as an efficacious endovascular approach to intracranial aneurysm treatment. FD stents that are currently available adopt an interwoven braided design. The relatively low radial stiffness intrinsic to this design could cause difficulty in deployment and poor stent-wall opposition, leading to high complication rates. A new FD stent is proposed to overcome the problems of the interwoven FD stents. The new device is manufactured from a Nitinol tube through a laser-cutting technique, and its unique structure allows for both low porosity and high packaging efficiency. Computational simulation using Abaqus has been conducted to investigate the radial stiffness and longitudinal flexibility of the new device. The new device exhibits high radial stiffness when compared to interwoven FD stents and superior longitudinal flexibility. Results from on-going in-vivo experiments and CFD simulations have also demonstrated the efficacy of the new device as a FD stent.
Key words: intracranial aneurysm; flow diverter stent; radial stiffness; longitudinal flexibility.
1
Introduction
Intracranial aneurysms, a balloon-like localised dilatation on the wall of an artery within the intracranial vasculature, remain an important cause of stroke morbidity and mortality (Bederson et 1
Corresponding author. E-mail: [email protected]. 2
al., 2009). As a safer and more efficacious alternative to conventional treatment methods such as surgical clipping (Dandy, 1938), coil embolization (Guglielmi, 2009), and stent-assisted coiling, the use of a low porosity stent alone as a flow diverter (FD) has gained substantial popularity in the last 10 years or so (Pierot, 2011). The treatment goal of using a FD stent is to correct the hemodynamic disturbance in the aneurysm, and eliminate what is considered to be the main cause of aneurysm progression and rupture (Cebral et al., 2005; Tateshima et al., 2009). A FD stent is placed across the neck of aneurysm sac to reduce blood flow into aneurysm sac, isolating the aneurysm by inducing stable thrombosis. Subsequently biological repair of the aneurysm neck by endothelial overgrowth of the stent mesh occurs to reconstruct the diseased segment of the parent artery (Kadirvel et al., 2014). It has been demonstrated that porosity is crucial for the functionality of FD stents and a porosity threshold of around 70% is the considered the optimum trade-off between device flexibility and efficacy (Liou et al., 2007; Seong et al., 2007). First introduced in 2007, FD stents are considered suitable for treating wide-neck, giant and fusiform or tiny lesions lateral to a large artery, carotid siphon or vertebrobasilar artery (Cirillo et al., 2012). Comprehensive reviews of the current status of FD stents are available in literature (Alderazi et al., 2014; Byrne and Szikora, 2012; Pierot). The FD stents currently available in clinical practice, e.g., the SILK (Balt Extrusion, Montmorency, Frnace) and PED (Coviden, Irvine, California, US), adopt an interwoven structural design, which are manufactured by braiding superelastic wires into a tubular mesh. This design suffers a low radial stiffness that could cause difficulty in endovascular deployment, spontaneous delayed migration or shortening (Chalouhi et al., 2013), and poor stent-wall opposition, risking extra-aneurysmal thrombosis and stroke. Partially due to the low radial stiffness, treatment by FD is associated with relatively high complication rates in comparison with other treatment methods (Byrne et al., 2010; Cirillo et al., 2012). In addition,
3
even modest stent oversizing frequently seen in clinical practice could lead to substantial increase in stent porosity, which is also undesirable for a FD stent (Shapiro et al., 2014).
We previously proposed a FD stent design, aimed at providing desirable mechanical properties to enable high device efficacy and deliverability (Ma et al., 2014). The device is manufactured out of a Nitinol tube through a laser-cutting technique. As a further development of the work conducted by the authors, this paper focuses on computational analysis of the key mechanical properties of a new design including the radial stiffness, which ensures the complete opening of the stent in the parent arteries, good stent-wall opposition and secure anchorage, and the longitudinal flexibility, which allows for guidance through tortuous intracranial arteries and good conformability to parent arteries. The layout of the paper is as follows. Section 2 introduces the geometrical design of the new FD stent and its variants. A detailed numerical modelling methodology for the new FD is presented in Section 3. The numerical results are summarized in Section 4, in which the radial stiffness and longitudinal flexibility of the new FD stent are discussed and compared with existing FD stents and intracranial stents. Preliminary results from an on-going in-vivo experiment aare nd the results of CFD simulations are reported in Section 5 to demonstrate the deliverability and efficacy of the new device. Finally, Section 6 concludes the paper by summing up the main findings drawn from the work.
2
The geometrical design
A prototype of the new FD stent, which is manufactured out of a Nitinol tube by laser-cutting technology and then post-processed (Ma et al., 2014), is presented in Fig. 1. The new FD stent has a middle portion of low porosity and two ends of high porosity. The middle portion is to be placed
4
across the neck of the aneurysm so as to reduce the volume of blood entering the aneurysm; hence a lower porosity is required in this region. The central section is formed from a series of repetitive modules, which consist of longitudinal waves connected by V-shaped connectors. Compared with the early version of the new FD stent (Ma et al., 2014), the introduction of the V-shaped connectors allows for fine tuning of radial stiffness, which will be demonstrated in the subsequent numerical analysis. Meanwhile, the longitudinal waves adopted in the early version are preserved to maintain the superior longitudinal flexibility and to achieve low porosity. The porosity of the middle portion of the device is tuneable by replacing some straight edges with curved ones, therefore generating two designs, referred to as type I and type II respectively, as presented in Fig. 2(a) and (b). Commercially available interwoven FD stents usually have a nominal porosity around 70% (Gross and Frerichs, 2013), which can be readily replicated by the type II design. The high-porosity ends of the device are designed for convenient packaging of the new FD stent and to accommodate radiopaque markers. In addition, the device ends are less stiff radially than the middle portion so as to achieve a smooth transition of force from the implanted device to the arterial wall. As a result, the strain exerted on the parent artery is reduced.
The middle portion of the new FD stent is critical for the functionality of the device as it determines the porosity, radial stiffness and longitudinal flexibility. The geometry of the middle portion of the new FD stent is shown in Fig. 2(c) and (d), and can be defined by seven parameters: the initial fully-expanded radius of the stent R0, the inner radius of the longitudinal wave r1, the inner radius of the V-shaped connector r2, the ratio of the height of the longitudinal wave to that of the sum of the longitudinal wave and the V-shaped connector a1/(a1+a2), the inner angle of the V-shaped connector α, the strut width w, and the wall thickness t. The design of the middle
5
portion of the device can be of either type I or type II, allowing for variation in device porosity ρ.
For a FD stent to be delivered into the small and torturous intracranial arteries, it is essential for the device to be packaged with a very small radial profile. The new FD stent can be radially folded through a two-step process. First, the device is stretched longitudinally from its initial state, Fig. 3(a), to an intermediate state, Fig. 3(b). Only the longitudinal waves are deformed at this step to reduce the device radius. Secondly, the stretched stent is radially crimped to fold the V-shaped connectors, resulting in further reduction in radius at the final state, Fig. 3(c). Upon deployment, the longitudinal waves spring back and the V-shaped connectors are unfolded to restore the initial fully expanded state of the stent.
Fourteen numerical models, including eleven models of type I and three of type II, were built to analyse the effects of stent geometry on its mechanical properties, and specifically the radial stiffness and the longitudinal flexibility. The parameters R0, r1, w and t of all the models were selected as 1.9mm, 0.3mm, 0.05mm, and 0.05mm, respectively, whereas a1/(a1+a2), α, r2, and ρ were systematically varied. The geometries of all the models are listed in Table 1. Note that the porosity of model A4, the design of which is presented in Fig. 2(d), reaches 70.2% matching that of commercially available woven FD stents.
3
Finite element modelling
As previously discussed, radial stiffness and longitudinal flexibility are key mechanical properties for a FD stent, and correspondingly form the basis of discussion in this study. Due to the repetitive nature of the new FD stet design, only a representative section of the middle portion is
6
analysed. However, the offset in longitudinal spacing between adjacent V-shaped connector rows in the design makes it difficult to extract a “single module” from the middle portion of the device, and thus, a section with 7 and 8 V-shaped connectors in adjacent rows and a total length l of 6mm, Fig. 2(c), is selected for modelling. The loading scenario is shown in Fig. 4(a): the FD stent is compressed radially by 12 identical rigid cylindrical shells from an initial radius of 1.9mm to 1.4mm, and the radial force per unit length fθ is calculated by the following equation (Zhou, 2009):
fθ
F 2 l
(1)
in which ΣF is the scalar summation of the radial reaction forces of all the cylindrical shells.
Four models, A1, A4, B1, and B3, are bent to investigate the longitudinal flexibility. Both ends of the device, as shown in Fig. 2(a), are attached to each model to form a complete stent. A1 is shown in Fig. 4(b) as an instance. The lengths of A1, A4, B1 and B3 are 15.45mm, 15.45mm, 15.50mm, and 15.65mm, respectively. Each model is subjected to pure bending to reach a curvature of 0.1mm-1, and the bending moment at each end of the device is recorded during the deformation process.
Numerical simulation was conducted using finite element analysis software package ABAQUS/Standard. Superelastic Nitinol, commonly used for intracranial stents and FD stents, was adopted as the material, the properties of which are listed in Table 2 (Rebelo et al., 2006). The details of the numerical modelling are described in the appendix.
7
4 4.1
Results Properties of A1
The results for model A1 are presented in detail and can be considered representative of all designs. The von Mises stress contour of A1 subjected to radial compression is shown in Fig. 4(c). It is clear from the contour that the V-shaped connectors are folded substantially to accommodate the reduction in radius of the model. In contrast, the deformation of the longitudinal waves is quite small. Moreover, as would be expected from the deformation mode, high stress concentrations are present at the curved tips of the V-shaped connectors, whereas the stress in the straight sections subject to little bending is small. The location of such stress concentrations suggest that the V-shaped connectors are mainly responsible for the radial stiffness of the FD stent under compression. The radial force per unit length fθ of A1 is plotted against FD stent radius R in Fig. 5(a). It can be seen that due to the introduction of the V-shaped connectors, fθ gradually increases with the reduction in R, a desirable feature for a FD stent.
The von Mises stress contour of A1 subjected to pure bending is presented in Fig. 4(d). A uniform mesh deformation distribution is achieved along the longitudinal length of the model, thus providing homogeneous mesh coverage the aneurysm neck. In addition, large stress only occurs at the curved portion of the longitudinal waves, whereas the rest of the model is subjected to very low stress. This result indicates that the longitudinal flexibility of the FD stent is predominantly 8
determined by the longitudinal waves and the effect of V-shaped connectors is insignificant. The bending moment of A1 is plotted against its curvature in Fig. 5(b). A near-linear relationship between bending moment and curvature is obtained.
4.2
Effects of geometrical parameters on radial stiffness
It is clear from the result of A1 that a significant stress distribution is present throughout the device geometry under a loading scenario of radial compression, hence it is expected that all of the geometric parameters previously discussed will affect the radial stiffness of the new FD stent.
Models A1-A4, all of which have identical a1/(a1+a2), α, and r2 but varying design type (and therefore porosity ρ) are first analysed to investigate the effects of porosity on device stiffness. The fθ vs. R curves of the four models are plotted in Fig. 5(a). It can be seen that the radial force per unit length increases mildly with the reduction in porosity. At a radius of 1.615mm, corresponding to 15% stent oversizing, fθ increases from 0.019N/mm to 0.025N/mm (31.6%) when porosity reduces from 86.3% in A1 to 70.2% in A4. Moreover, fθ converges quickly when the porosity is below 80%: the curves of A3 and A4, which have a difference in porosity of 6.5% are almost identical. This result indicates that within the practical range of porosity, 70%-80%, the radial stiffness of the device varies only marginally with porosity. Hence, porosity does not need to be taken into consideration when tuning the radial stiffness of the new FD stent design. 9
Models B1-B3, differ from each other in terms of the ratio a1/(a1+a2) alone, with all other parameters remaining constant. It should be noted that when a1/(a1+a2) reaches 1, the V-shaped connectors in the design disappear and the device is reduced to the early FD stent design (Ma et al., 2014). When the value is 0, the design becomes a conventional zigzag stent. The fθ vs. R curves of the three models B1-B3 are plotted together with that of A1 in Fig. 5(c). The plot shows that fθ increases significantly with a1/( a1+a2), thus providing an effective way of tuning the radial stiffness of the new FD stent.
The stress contour of A1 in Fig. 4(c) indicates that a high stress concentration occurs at the tip of the V-shaped connectors under deformation, therefore, the α and r2 parameters of the connectors are expected to significantly affect the device radial stiffness. The fθ vs. R curves of models C1-C4 with increasing α are first plotted together with that of A1 in Fig. 5(d). Two observations can be made. First, fθ increases with α, i.e., a more obtuse angle leads to a higher force. Second, fθ is quite sensitive to α, and a wide range of fθ can be achieved by varying α. The fθ vs. R curves of models D1-D3, which have increasing values of r2 are plotted together with A1 in Fig. 5(e). It is clear that fθ also increases with r2. In contrast to α, however, fθ is much less sensitive to r2, and therefore only a small range of fθ is achievable.
10
4.3
Effects of geometrical parameters on longitudinal flexibility
The stress contour of A1 subjected to pure bending, shown in Fig. 4(d), indicated that the longitudinal waves of the design undergo significant deformation and play a major role in determining device’s longitudinal flexibility. The contribution to longitudinal stiffness from the V-shaped connectors, which undergo very small deformation under pure bending, is considered insignificant. Therefore only variation in design type and a1/(a1+a2), which are associated with the geometry of longitudinal waves, are analysed to evaluate their respective effects on longitudinal flexibility. The bending moment vs. curvature curves of A1 and A4, which differ only in design type and hence porosity, are plotted in Fig. 5(b). It is clear from the figure that a large reduction in porosity from 86.3% in A1 to 70.2% in A4 only very slightly increases the device bending stiffness. Hence, the effect of design type on longitudinal flexibility can be practically ignored. The bending moment vs. curvature curves of A1, B1 and B3, with varying a1/(a1+a2), are plotted in Fig. 5(f). The figure clearly shows that the larger a1/(a1+a2) ratio, the lower the bending moment, which again demonstrates that longitudinal waves are mainly responsible for the longitudinal flexibility of the new FD stent. In short: the larger the proportion of longitudinal waves in the stent structure, the lower the bending moment required to deform the device, and the more flexible the FD stent.
4.4
Comparison with existing intracranial stents and FD stents
The radial stiffness and longitudinal flexibility can significantly affect the efficacy of a FD stent (Liou et al., 2007; Seong et al., 2007). However, there is no accepted design guidance for either parameter for a successful FD device. Consequently, the approach adopted here is to evaluate the mechanical properties of the new FD stent in comparison to the properties of a number of
11
currently available devices. Both intracranial stents that are also laser-cut from Nitinol tubes, including Neuroform3, Enterprise, Wingspan and Solitaire (Krischek et al., 2011), and interwoven FD stents like the PED (Fischer et al., 2012) are considered. Model A4 is selected for the comparison as it has a low porosity, which matches that for commercial interwoven FD stents.
Figure 6(a) presents the value of fθ for each design of intracranial stent with a 15% stent oversizing in all cases. It can be seen that A4 has a relatively high fθ, 11.6% larger than that of Wingspan. This result indicates that the new FD stent is able to provide adequate radial stiffness that exceeds current intracranial stents. Should a lower radial stiffness be required, it can be easily achieved by geometrical modification of the design, such as by reducing α. Note that fθ for PED is about 0.002N/mm (Fischer et al., 2012), less than 10% of that for the new device.
The bending moments of each design of intracranial stent when bent into a 12.7mm arc are plotted in Fig. 6(b). It is clear from the figure that A4 requires the smallest bending moment, only 10.7% of that for Neuroform3. In other words, the new FD stent has the highest longitudinal flexibility, making it ideal for maneuvering through the tortuous arteries of the cerebral vasculature. Data for the longitudinal flexibility of the PED are not available and thus no comparison is made.
5 5.1
In vivo animal experiment and hemodynamic simulation In vivo animal experiment
Rabbit models with elastase-induced arterial saccular aneurysms (Sadasivan et al., 2009) were used to test the efficacy of the new device. A 5F (1.67mm) delivery system was designed to deploy the
12
new FD stent conveniently and accurately. A sketch of the design is presented in Fig. 7(a) and a prototype is shown in Fig. 7(b). The new device was loaded at the tip of the system and deployed through an unsheathing process. Stent samples of 1.75mm radius and 14mm length were loaded into the system after two radiopaque markers made from platinum were installed on each end of the stents to improve visibility under X-ray fluoroscopy.
A typical result of one experimental aneurysm before, immediately after, and three months after stenting is shown in Fig. 8. It can be seen from Fig. 8(b) that, first of all, the new FD stent fully opened in the parent artery and conformed very well to the artery wall, and secondly, an immediate isolation of the aneurysm was achieved. The angiography image obtained from the three-month follow-up of the rabbit, which is shown in Fig. 8(c), confirms that no refilling of the aneurysm occurred and that the aneurysm was completely occluded.
5.2
Hemodynamic simulation
A representative aneurysm geometry was extracted from experimental 3D angiographic imaging and used to quantify the flow-diverting effect of the new FD stent. The new device was virtually deployed into the rabbit aneurysm geometry assuming a rigid vessel wall (Peach et al., 2014b). Blood was modelled as an incompressible Newtonian fluid with a density of 1000 kg/m3 and a dynamic viscosity of 0.004 PaS. The 3D unsteady Navier-Stokes equations were solved using the finite volume method using CFD-ACE multiphysics package (ESI Group, Paris, France) with a Central Differencing scheme for spatial variations and a Crank-Nicholson scheme for time marching. The SIMPLE-Consistent (SIMPLEC) pressure correction method and an algebraic multigrid method for convergence acceleration were used (Van Doormaal and Raithby, 1984;
13
Webster, 1994).
The aneurysm geometry with and without the device deployed was meshed using a Projected Single Domain mesh, an Omnitree Cartesian tree type, and three near-wall Cartesian layers. Mesh sizes were between 1.67 and 2.81 million elements, assuming a mesh-independent cell density (>4,000 elements/mm3) as discussed in previous works and similar studies (Cebral et al., 2014a; Peach et al., 2014a; Peach et al., 2014b). Time-step independence was also assumed at the chosen time step of 0.01 seconds, in line with similar studies in the literature (Cebral et al., 2014a). The transient inlet boundary condition of volumetric parent vessel flow, in addition to velocity profile and pulse rate, was estimated from the experimental results reported by Cebral et al. (Cebral et al., 2014a).
The results of the CFD simulations indicated that the new FD stent reduces the mean velocity of blood entering the aneurysm dome by a factor of 2.54. Such a reduction is comparable to the mean velocity reduction of 2.64 reported by Cebral et al. (Cebral et al., 2014b) for rabbit model aneurysms treated with the PED, indicating that the flow-diverting effect of the new FD stent is comparable to commercial interwoven FD stents.
6
Conclusions
A new FD stent that can be laser-cut from a Nitinol tube, for direct treatment of intracranial aneurysms, has been proposed in this paper. A unique pattern design for the new FD stent, i.e., four longitudinal waves interconnected by V-shaped components, provides high radial stiffness and longitudinal flexibility, as well as convenient packaging and delivery. A low porosity of 70% can be achieved for the new FD stent to effectively divert blood flow from aneurysm sac and
14
promote aneurysm thrombosis.
The mechanical properties, specifically the radial stiffness and longitudinal flexibility, of the new FD stent, were numerically analysed and compared with those of several commonly used intracranial stents and commercially available interwoven FD stents. It has been found that the radial stiffness of the new FD stent is an order of magnitude higher than that of the PED and is comparable to that of commonly used intracranial stents. The longitudinal flexibility of the new FD stent, on the other hand, is considerably higher than that of commonly used intracranial stents. In one particular case, model A4 has low porosity of 70.2%, which matches that of commercially available interwoven flow diverter stents; high radial stiffness, which is over 10 times larger than that of commercially available flow-diverter stents and comparable to those of commonly used intracranial stents; and superior longitudinal flexibility, when compared with commonly used intracranial stents. Both in vivo experimental results and CFD simulations show that the new FDS performs well as a flow-diverter and achieves performance comparable to current woven flow-diverters, such as the PED. Such a device therefore serves as an excellent candidate for a new flow-diverter stent.
Study limitations A limitation of the hemodynamic simulations conducted in this study is the use of estimated inflow boundary conditions, although previous studies have shown such estimates to produce reasonable results (Cebral et al., 2014a). Additionally the highly dynamic environment of the rabbit aorta can introduce error in angiographic imaging and subsequent 3D reconstructions. A number of factors including device size, wire diameter, and device pore shape and size are likely
15
to influence the flow-diverting effect of a given stent design. Due to obvious constraints, the effects of these parameters have not been fully explored in this study; hence, the sensitivity of the device performance to these parameters is not known. Regarding the elastase induced saccular aneurysms rabbit model, one limitation is that the aneurysm was created in the extracranial carotid artery, and thus was not an intracranial aneurysm. This is attributable to the size constraints of working with rabbits, as well as the challenging nature of controlling the elastase effect on rabbit cerebral circulation. Nearly all previously described animal aneurysm models involved extracranial arteries (Frösen et al., 2006). As the purpose of this model was intended to study the hemodynamic effects of the flow-diverter stents, as well as the cellular interaction of the new device with the parent artery, the extra-cranial carotid elastase induced saccular aneurysm model was appropriate.
Conflicts of interest The authors declare that there are no conflicts of interest.
Acknowledge The authors would like to thank Wellcome Trust and EPSRC for their financial support under grant number WT 088877/Z/09Z, Dr M. Megahed of the ESI Group for allowing the use of the CFD-ACE suite, Dr. Y. Ventikos, and the use of the Oxford Advanced Research Computing (ARC) facility in carrying out this work.
References Abaqus, Version 6.7 ed. SIMULIA Corp., Providence, RI, USA.
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Alderazi, Y.J., Shastri, D., Kass-Hout, T., Prestigiacomo, C.J., Gandhi, C.D., 2014. Flow Diverters for Intracranial Aneurysms. Stroke Research and Treatment 2014. Bederson, J.B., Connolly, E.S., Batjer, H.H., Dacey, R.G., Dion, J.E., Diringer, M.N., Duldner, J.E., Harbaugh, R.E., Patel, A.B., Rosenwasser, R.H., 2009. Guidelines for the management of aneurysmal subarachnoid hemorrhage a statement for healthcare professionals from a special Writing Group of the Stroke Council, American Heart Association. Stroke 40, 994-1025. Byrne, J.V., Beltechi, R., Yarnold, J.A., Birks, J., Kamran, M., 2010. Early experience in the treatment of intra-cranial aneurysms by endovascular flow diversion: a multicentre prospective study. PLoS One 5, e12492. Byrne, J.V., Szikora, I., 2012. Flow diverters in the management of intracranial aneurysms: a review. EJMINT Orig Artic 1225000057, 1-22. Cebral, J.R., Castro, M.A., Appanaboyina, S., Putman, C.M., Millan, D., Frangi, A.F., 2005. Efficient pipeline for image-based patient-specific analysis of cerebral aneurysm hemodynamics: technique and sensitivity. Medical Imaging, IEEE Transactions 24, 457-467. Cebral, J.R., Mut, F., Raschi, M., Ding, Y.-H., Kadirvel, R., Kallmes, D., 2014a. Strategy for analysis of flow diverting devices based on multi-modality image-based modeling. International journal for numerical methods in biomedical engineering 30, 951-968. Cebral, J.R., Mut, F., Raschi, M., Hodis, S., Ding, Y.H., Erickson, B.J., Kadirvel, R., Kallmes, D.F., 2014b. Analysis of hemodynamics and aneurysm occlusion after flow-diverting treatment in rabbit models. American Journal of Neuroradiology 35, 1567-1573. Chalouhi, N., Tjoumakaris, S.I., Gonzalez, L.F., Hasan, D., Pema, P.J., Gould, G., Rosenwasser, R.H., Jabbour, P.M., 2013. Spontaneous delayed migration/shortening of the pipeline embolization device: report of 5 cases. American Journal of Neuroradiology 34, 2326-2330.
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Cirillo, L., Leonardi, M., Dall'Olio, M., Princiotta, C., Stafa, A., Simonetti, L., Toni, F., Agati, R., 2012. Complications in the treatment of intracranial aneurysms with silk stents: an analysis of 30 consecutive patients. Interventional Neuroradiology 18, 413. Dandy, W.E., 1938. Intracranial aneurysm of the internal carotid artery. Annals of Surgery 107, 654-659. Fischer, S., Vajda, Z., Perez, M.A., Schmid, E., Hopf, N., Bäzner, H., Henkes, H., 2012. Pipeline embolization device (PED) for neurovascular reconstruction: initial experience in the treatment of 101 intracranial aneurysms and dissections. Neuroradiology 54, 369-382. Frösen, J., Marjamaa, J., Myllärniemi, M., Abo-Ramadan, U., Tulamo, R., Niemelä, M., Hernesniemi, J., Jääskeläinen, J., 2006. Contribution of mural and bone marrow-derived neointimal cells to thrombus organization and wall remodeling in a microsurgical murine saccular aneurysm model. Neurosurgery 58, 936-944. Gross, B.A., Frerichs, K.U., 2013. Stent usage in the treatment of intracranial aneurysms: past, present and future. Journal of Neurology, Neurosurgery, and Psychiatry 84, 244-253. Guglielmi, G., 2009. The beginning and the evolution of the endovascular treatment of intracranial aneurysms: from the first catheterization of brain arteries to the new stents. Journal of neurointerventional surgery 1, 53-55. Hoh, B.L., Velat, G.J., Wilmer, E.N., Hosaka, K., Fisher, R.C., Scott, E.W., 2010. A Novel Murine Elastase Saccular Aneurysm Model for Studying Bone Marrow Progenitor-Derived Cell-Mediated Processes in Aneurysm Formation. Neurosurgery 66, 544. Kadirvel, R., Ding, Y.-H., Dai, D., Rezek, I., Lewis, D.A., Kallmes, D.F., 2014. Cellular mechanisms of aneurysm occlusion after treatment with a flow diverter. Radiology 270, 394-399. Krischek, O., Miloslavski, E., Fischer, S., Shrivastava, S., Henkes, H., 2011. A comparison of
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functional and physical properties of self-expanding intracranial stents [Neuroform3, Wingspan, Solitaire, Leo+, Enterprise]. Minimally Invasive Neurosurgery 54, 21-28. Liou, T.M., Li, Y.C., Juan, W.C., 2007. Numerical and experimental studies on pulsatile flow in aneurysms arising laterally from a curved parent vessel at various angles. Journal of biomechanics 40, 1268-1275. Ma, J., You, Z., Byrne, J., Rizkallah, R., 2014. Design and Mechanical Properties of a Novel Cerebral Flow Diverter Stent. Annals of Biomedical Engineering 42, 960-970. Peach, T., Cornhill, J.F., Nguyen, A., Riina, H., Ventikos, Y., 2014a. The 'Sphere': A Dedicated Bifurcation Aneurysm Flow-Diverter Device. Cardiovascular engineering and technology 5, 334-347. Peach, T.W., Ngoepe, M., Spranger, K., Zajarias-Fainsod, D., Ventikos, Y., 2014b. Personalizing flow-diverter intervention for cerebral aneurysms: from computational hemodynamics to biochemical modeling. International journal for numerical methods in biomedical engineering 30, 1387-1407. Pierot, L., 2011. Flow diverter stents in the treatment of intracranial aneurysms: Where are we? Journal of Neuroradiology 38, 40-46. Rebelo, N., Gong, X.-Y., Hall, A., Pelton, A.R., Duerig, T.W., Year Finite element analysis on the cyclic properties of superelastic nitinol. In Proceedings of the International Conference on Shape Memory and Superelastic Technologies. Sadasivan, C., Cesar, L., Seong, J., Wakhloo, A.K., Lieber, B.B., 2009. Treatment of rabbit elastase-induced aneurysm models by flow diverters: development of quantifiable indexes of device performance using digital subtraction angiography. Medical Imaging, IEEE Transactions 28, 1117-1125.
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Seong, J., Wakhloo, A.K., Lieber, B.B., 2007. In vitro evaluation of flow diverters in an elastase-induced saccular aneurysm model in rabbit. Journal of Biomechanical Engineering 129, 863-872. Shapiro, M., Raz, E., Becske, T., Nelson, P.K., 2014. Variable Porosity of the Pipeline Embolization Device in Straight and Curved Vessels: A Guide for Optimal Deployment Strategy. American Journal of Neuroradiology 35, 727-733. Tateshima, S., Tanishita, K., Hakata, Y., Tanoue, S.-y., Viñuela, F., 2009. Alteration of intraaneurysmal hemodynamics by placement of a self-expandable stent: laboratory investigation. Journal of neurosurgery 111, 22-27. Van Doormaal, J.P., Raithby, G.D., 1984. Enhancements of the SIMPLE method for predicting incompressible fluid flows. Numerical heat transfer 7, 147-163. Webster, R., 1994. An algebraic multigrid solver for Navier-Stokes problems. International Journal for Numerical Methods in Fluids 18, 761-780. Zhou, X., 2009. Design of Stents for the Direct Treatment of Intracranial Aneurysms. University of Oxford, Oxford.
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Captions of Figures Fig. 1
A physical sample of the new FD stent
Fig. 2
(a) pattern design type I for the new FD stent, (b) pattern design type II for the new FD stent, (c) geometrical parameters definition for the new FD stent, and (d) design of model A4
Fig. 3
Folding process of the new FD stent, (a) initial fully expanded state, (b) intermediate state after longitudinal elongation, and (c) final fully folded state after radial compression
Fig. 4
(a) Radial compression setup, (b) a complete A1 for pure bending, (c) von Mises stress contour of A1 subjected to radial compression, and (d) von Mises stress contour of A1 subjected to pure bending
Fig. 5
(a) fθ vs. R curves of A1-A4, (b) bending moment vs. curvature curves of A1 and A4, (c) fθ vs. R curves of A1 and B1-B3, (d) fθ vs. R curves of A1 and C1-C4, (e) fθ vs. R curves of A1 and D1-C3, and (f) bending moment vs. curvature curves of A1, B1, and B3.
Fig. 6
Comparison of the stents’ mechanical properties, (a) fθ at 15% stent oversizing, and (b) bending moment needed to bend the stents into a 12.7mm arc
Fig. 7
(a) Sketch of the delivery system, and (b) a prototype of the delivery system
Fig. 8
Animal experiment results, (a) before FD stent deployment, (b) immediately after FD stent deployment, and (c) three month follow-up
Captions of Tables Table 1
Geometries of the new FD stent models
Table 2
Material properties of superelastic Nitinol
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Table 1 Model type a1/(a1+a2) α (mm) r2 (mm)
A1
I
0.5
60
0.125
86.3%
A2
II
0.5
60
0.125
81.7%
A3
II
0.5
60
0.125
76.7%
A4
II
0.5
60
0.125
70.2%
B1
I
0.45
60
0.125
86.2%
B2
I
0.55
60
0.125
86.3%
B3
I
0.6
60
0.125
86.3%
C1
I
0.5
40
0.125
84.1%
C2
I
0.5
50
0.125
85.4%
C3
I
0.5
70
0.125
86.8%
C4
I
0.5
80
0.125
87.3%
D1
I
0.5
60
0.0875
86.2%
D2
I
0.5
60
0.1625
86.4%
D3
I
0.5
60
0.2
86.5%
Table 2 Property Value
Definition
EA
50 GPa
Austenite elasticity
EM
37 GPa
Martensite elasticity
Ms
400 MPa Starting transformation stress of loading
Mf
650 MPa End transformation stress of loading
As
350 MPa Starting transformation stress of unloading
Af
80 MPa
End transformation stress of unloading
L
0.055
Maximum residual strain
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