Novel Hybrid Flow Diverter for Intracranial Aneurysm

An intracranial aneurysm is a weakened area in the wall of a cerebral artery which causes abnormal localized ballooning of the blood vessel. As an aneurysm grows, it puts pressure on adjacent structures and may eventually rupture, leading to severe complications or even sudden death. The standard treatments for intracranial aneurysms include traditional craniotomy and endovascular coiling. The purpose of these treatments is to stop the blood flow to an aneurysm to reduce the risk of rupture. I n recent years, another new device, “flow diverter”, has gained popularity. It is placed in the parent artery to divert the blood flow away from the weakened area, isolating aneurysms from normal circulation. Although flow diverter stents have great potential, there remain clinical issues to be resolved. This paper proposes a unique hybrid flow diverter, the first of its kind in the world, for treatment of the intracranial aneurysm. The hybrid flow diverter is designed to have variable mesh densities, with the denser side facing an aneurysm to block the blood flow and the lighter side facing the artery to prevent stenosis. It is deployed in the main cerebral artery next to an aneurysm to divert the blood flow away from the weakened aneurysm. Simulation results showed that the hybrid flow diverter reduced the blood flow into an aneurysm by a whopping 75 – 95%. The residence time of the blood flow inside an aneurysm was 12.47 times longer with the hybrid flow diverter, which may trigger thrombogenic reaction to fill an aneurysm and thus reduce the risk of rupture.


Introduction
Cardiovascular disease (CVD), cancer, and stroke are the leading causes of death worldwide. Stroke can be divided into two main types: ischemic and hemorrhagic stroke. Hemorrhagic stroke may occur due to the spontaneous bursting of an intracranial aneurysm [1]. An intracranial aneurysm is a cerebrovascular disorder in which structural weakening of the wall media results in localized pathological dilation of the blood vessel. About 6% of the general population develop one or multiple intracranial aneurysms, which are relatively common in adults [2]. Patients harboring aneurysms are often unaware of their presence because most of these lesions are asymptomatic and small. The growth of aneurysms appears to be unpredictable and highly variable. As an aneurysm grows, it puts pressure on surrounding brain tissues and may eventually rupture, leading to fatal subarachnoid hemorrhage (SAH). SAH has high mortality, ranging from 25 to 50%, and causes severe disability in about 30% of survivors. Moreover, the rebleeding process may cause approximately 30% of untreated patients with intracranial aneurysms to die in the subsequent decade [3][4].
Treatment options for intracranial aneurysms generally include traditional craniotomy and endovascular coiling. In traditional craniotomy, an opening is made in the skull under general anesthesia, and a tiny clip is placed across the neck of an aneurysm to stop the blood flow into it. The purpose of surgical clipping is to isolate the aneurysm from the original circulation and thereby reduce the risk of rupture.
Craniotomy has many limitations, such as the location, size, and geometry of the aneurysm, as well as the high possibility of complications induced by nerve damage and large aneurysms (larger than 25 mm) [5]. In addition, if the patient is old or in poor physical condition, the craniotomy is not recommended due to the high risks of invasion, the large wound, and the lengthy operation time.
Currently, a safer alternative to traditional craniotomy is a minimally invasive technique known as endovascular coiling. In this technique, an aneurysm is filled with coils to block it off from the main artery. The occlusive agent is a detachable platinum coil delivered through a micro-catheter positioned within an aneurysm. Although the success rate is high, limitations to the application of endovascular coiling treatment remain. One is the risk that the platinum coil may fall into the blood vessel and cause stroke when the aneurysmal neck diameter is greater than 4 mm or the aneurysmal dome to neck ratio is less than 2. In addition, although post-embolization angiography can show the degree of occlusion, the density of the coil packing cannot be adjusted, which may result in unstable and non-uniform strength. Also, if an aneurysm is not under equilibrium stress when the coils are inserted, it may cause the micro-catheter, guidewire, or coil to penetrate the aneurysm, leading to rupture and hemorrhage [6][7][8][9].
In recent years, another new device, the flow diverter, has been proposed to solve the problems mentioned above. The flow diverter is placed in the parent artery to divert blood flow away from the weakened area, isolating the aneurysm from the original circulation. Studies have shown that even for giant aneurysms with a high morbidity rate after endovascular coiling, deployment of a flow diverter has a high success rate for aneurysm sac embolization [10][11]. Although flow diverter stents have great potential, there remain clinical issues to be addressed. For example, excessively high metal coverage of the flow diverter may induce in-stent restenosis and block the parent artery, while low metal coverage may reduce the effect of flow diversion [12]. Also, the flow diverter should have advanced longitudinal flexibility to enter the brain, and its visibility also needs to be improved [13].
In this paper, a novel stent for interventional management of intracranial aneurysms, the hybrid flow diverter, is proposed. The hybrid flow diverter concept is the first of its kind in the world. It was designed with variable metal densities such that the denser side faces an aneurysm to block the blood supply and the lighter side faces the parent artery to prevent in-stent restenosis. This provides a novel design to maintain the flow diverter performance of blocking blood flow and simultaneously reduce the clinical problems with comprehensive simulation verification.

Methodology
The development of the hybrid flow diverter is revealed, including the optimal design, finite element analysis of the manufacturing and deployment processes, and hemodynamic analysis before and after deployment. The parametric design methodology achieves the optimal design. Several finite element models were developed to analyze the structural integrity of the hybrid flow diverter during the manufacturing and deployment processes. The fatigue resistance of the device was evaluated by Goodman life analysis. Computational fluid dynamics models were established to assess the hemodynamic performance of the device. No human samples were involved in this study.

Flow Diverter Basic Design
Flow diverters can be divided into two major types: laser-cut and braided metallic structures. The basic design of the laser-cut flow diverter was used as the standard stent in this study. A stent is an assembly of a series of cuddle rings interconnected with bridging links. The unit cell of a ring consists of a crown and a strut. A crown is a curved structure composed of two concentric arcs with different radii, and a strut is a straight portion connecting two adjacent crowns. A simple flow diverter, as 6 illustrated in Fig. 1, comprises several components: the strut width, ring spacing, crown radius, thickness, crown number, and ring number. These geometric parameters in different design patterns have critical impacts on the mechanical properties of flow diverters [14][15]. In this study, the mathematical relations between the different geometric parameters were established. Defining the relations between these parameters allowed quick modification of the parameters as needed and the automatic generation of new designs, which greatly improved the efficiency of design modification.

Hybrid Flow Diverter Design
The design pattern of a hybrid flow diverter is typically laser-cut directly onto a nitinol micro-hypotube. To avoid neointimal hyperplasia by high-profile metal, we propose a flow diverter with variable metal densities. The denser side faces an aneurysm to inhibit the blood flow, and the lighter side faces the parent artery to prevent in-stent restenosis. To make a flow diverter having a variety of metal densities, we defined the mathematical relations between different geometric parameters. For example, once the mesh density of the denser side (strut width, crown number and crown radius) was set, the crown radius of the lighter side could be automatically calculated for a fixed crown number and tube diameter.
In this study, a flow diverter was designed with different ring numbers, which strongly affect the strength of the denser and lighter sides. Due to the higher mesh density and greater ring number on the denser side, the crown expansion degree is relatively smaller, since the dense side is far stronger than the lighter side. On the other hand, the crown of the lighter side stretches to a great extent when expanding due to its low strength. Therefore, the design has to balance the strengths of the mesh densities of the two sides.
Through adjustment of several designs and optimal parameters, a more balanced design of the mesh densities was found. Based on this design, the geometric model of the hybrid flow diverter proposed in this paper was constructed. Figs. 2a and 2b show the geometric structure of the hybrid flow diverter, which is mainly divided into four rings. The distal and proximal rings in the axial direction have low density, and the middle rings in the circumferential direction are divided into low and high density.
The ring number ratio of the high vs. low density in the middle section is 2 to 1.
Therefore, the denser side can block the blood flow to an aneurysm, and the lighter side, deployed in the parent artery, can avoid inducing neo-intimal hyperplasia at the same time.

Finite Element Analysis
In this paper, finite element analysis was conducted using the ABAQUS/Standard FEA solver (Dassault Systems Simulia Corp., Providence, RI, USA) with the user-defined UMAT subroutine [16]. The effects on the hybrid flow diverter in the manufacturing process were analyzed to confirm that the hybrid flow diverter could expand to the desired shape without material damage during manufacturing. After heat treatment and finalization, the hybrid flow diverter was further deployed in a model of an intracranial aneurysm to investigate its clinical behaviors. By obtaining the geometrical shape of the hybrid flow diverter deployed in the parent artery, we could ensure that no material damage would occur during the deployment process.

Material Properties
To make the hybrid flow diverter self-expand to its target shape and dimensions in the parent artery, we used nitinol (NiTi) as the material due to its unique superelasticity and shape-memory properties. It is noted that, when force is applied, the austenite transforms into stress-induced martensite. After the load is removed, nitinol returns to its original shape due to the transformation between austenite and martensite. Since nitinol has more complex mechanical properties than ordinary metals, the hyperelastic properties of the ABAQUS finite element model needed to be set with the super-elastic user-defined material subroutine UMAT. The material used in this study was 2.0 mm nitinol hypotubes manufactured by the Minitube company.
To make the material properties of the UMAT simulation closer to the actual material properties, we used the ABAQUS user plugin to correct the data repeatedly. It was found that the nitinol hypotube had about 12% elastic strain. When the strain caused by the load is less than 12%, then upon unloading, the material can recoil to its original shape. If the strain is more than 12% after the load is applied, the material is unable to revert to its original shape due to plastic deformation. Therefore, in this study, 12% strain was set as the safe upper limit to avoid the risk of material damage in the clinic.

Manufacturing Simulation
Several finite element models were developed to evaluate the structural integrity and pulsatile fatigue life of a hybrid flow diverter subjected to loading conditions consistent with current manufacturing and clinical practice [16][17]. These included hybrid flow diverter expansion and its corresponding heat treatment during manufacturing, crimping of the hybrid flow diverter inside a catheter for delivery, its release into a blood vessel, and pulsatile fatigue life under systolic/diastolic cycles relative to heart beats. The procedures were simulated in the following five major steps: Step 1: Expand the hybrid flow diverter to 4.0 mm ID.
Step 2: Heat treat the hybrid flow diverter after expansion.
Step 3: Crimp the hybrid flow diverter inside the 1.5 mm ID catheter for delivery.
Step 4: Release the hybrid flow diverter into the 4.0 mm ID blood vessel.
To simulate the hybrid flow diverter expansion during manufacturing and the crimping inside a catheter, two rigid cylindrical sleeves of different diameters were added to the model, with one inside the hybrid flow diverter and the other one outside the hybrid flow diverter [18]. During the expansion and subsequent heat treatment, the inner sleeve was expanded to the target size and the resulting stress/strain values in all elements of the FEA model were re-set to zero to simulate the stress relief process during heat treatment. After completion of the final expansion, the outer sleeve was applied to crimp the hybrid flow diverter inside the catheter for delivery. The third sleeve of 4 mm was then created to imitate the blood vessel. When the catheter was removed, the hybrid flow diverter sprang back until it contacted this 4 mm cylindrical surface.
Since the crowns of the hybrid flow diverter are subjected to high stresses/strains during expansion, the hybrid flow diverter was meshed with the incompatible mode 8-node brick element (C3D8I), which provides accurate predictions of maximum stress/strain with its integration points. On the other hand, two cylindrical sleeves were meshed with the three-dimensional, 4-node surface element with reduced integration (SFM3D4R). Because of the numerous contact interactions between surfaces in the simulation process, the function of the ABAQUS contact pair was adopted to set the contacts, which were defined to prevent surface penetration [17]. In this study, both the contact pairs and the self-contact tangential friction coefficients were set to 0.1.

Fatigue Life Analysis
It is necessary for the hybrid flow diverter to withstand many types of stress and deformation when it is deployed in an artery, and the rhythmic pulsation of blood will present a major challenge to its fatigue resistance. If the structure cannot withstand extended periods of pulsation, the material may suffer fatigue damage and eventually fracture. The Food and Drug Administration (FDA) provides non-clinical engineering tests of intravascular stents and their associated delivery systems in the guidance document. FDA recommends that Goodman fatigue life analysis should be used to determine the fatigue safety factor under physiologic loading that simulates blood pressure conditions in the human body [19]. Following the simulations of stent manufacturing, a ± 3% stent diameter oscillation was applied to simulate pulsatile fatigue loading. The strain values of the integration points on the principal axis of the hybrid flow diverter obtained by finite element analysis were reprocessed to calculate the fatigue safety factor of the hybrid flow diverter, and then the fatigue resistance of the hybrid flow diverter was obtained. The results showed that fatigue failure would occur if the strain state in the device satisfied the following relation [16][17]: where a  is the strain amplitude applied to the device, e  is the material endurance limit, m  is the mean strain applied to the device, and u  is the material ultimate strain. The Goodman diagram is a plot of the normalized strain amplitude u a /   (on the y-axis) versus the normalized mean strain u m /   (on the x-axis). The equation represents the failure line.
The fatigue safety factor (FSF) is defined as the ratio of the strain amplitude to the modified endurance limit. If the FSF is less than 1.0, it indicates that stent fatigue failure may take place due to pulsatile loading. If the FSF is greater than 1, it represents that the structure of the medical device is not prone to fatigue damage, and its fatigue resistance increases with increases in the FSF.
The position with the minimum FSF is the place where fatigue failure is most likely to occur. Therefore, the minimum FSF was used to evaluate the fatigue resistance of the hybrid flow diverter.

Stent-to-artery Ratio
One of the complications of treatment with flow diverters is the occurrence of in-stent stenosis due to its high metal coverage. Therefore, the hybrid flow diverter was designed with a higher area of coverage on the side facing the aneurysm and a lower area of coverage on the side facing the parent artery. The stent-to-artery ratio is considered an important clinical attribute, and it can be calculated using the following equation: where stent A is the contact area between the outer surface of the hybrid flow diverter and the artery, and artery A is the inner surface area of the artery.

Deployment Simulation in Intracranial Aneurysm Model
The heat-treated hybrid flow diverter was further deployed in the intracranial aneurysm model to obtain the steady-state results of the interaction between the hybrid flow diverter and the arterial wall. These results would indicate whether the hybrid flow diverter could be bent and deployed in the parent artery of an aneurysm.
They could also be used to analyze whether material damage would occur during the deployment of the hybrid flow diverter. The procedures were simulated in the following three major steps: Step 1: Crimp the hybrid flow diverter.
Step 2: Bend the hybrid flow diverter into the 4.0 mm ID parent artery.
Step 3: Release the hybrid flow diverter, and the deployment is completed.
The deployment models included the vascular geometry model of an intracranial aneurysm, the hybrid flow diverter after heat treatment, and the cylindrical sleeve to an aneurysm into hard, medium, and soft, and the three levels of the corresponding risk of aneurysm rupture as low, medium and high [22][23]. We used the hierarchy of soft as the material property of an aneurysm. The Mooney-Rivlin model was adopted to construct the material of the aneurysm, while the parent artery was constructed with reference to the second order hyperelastic constitutive model proposed by Creane et al. [24]. The materials of the arterial wall were assumed to be hyperelastic, homogeneous, isotropic and incompressible. The strain energy potential of the two materials can be defined as follows: 1  01  2  11  1  2  22  20  1  02  2   3  3  3  3  3 3 where s W is the strain energy potential of the aneurysm, p W is the strain energy potential of the parent artery, 1 I is first deviatoric strain invariants, 2 I is second deviatoric strain invariants, and 10 C , 01 C , 11 C , 20 C and 02 C are material parameters, which are shown in Table 1.
For the intracranial aneurysm model, the sweep method can be used to generate hexahedron meshes after proper geometric partitioning, so it was meshed with an 8-node hexagonal element with reduced integration (C3D8RH). The 4.0 mm ID hybrid flow diverter after heat treatment was adopted in the deployment simulation.
The element types of the hybrid flow diverter and the cylindrical sleeve were meshed with the same method in the manufacturing simulation.

Computational Fluid Dynamics
In this study, the blood flow model of an intracranial aneurysm was used for hemodynamic analysis to evaluate the capability of the hybrid flow diverter to block the blood flow and to analyse the wall shear stress before and after the deployment.
After the manufacturing and deployment simulation, the hybrid flow diverter was further used to establish a hemodynamic model to analyse the clinical behaviour of the hybrid flow diverter.

Material Properties
Blood usually behaves like a Newtonian fluid when the shear rate is greater than 100 s -1 [25]. However, due to flow disturbances, the actual shear rates in stented arteries could become lower than 100 s -1 . Therefore, in this paper, blood flow was assumed to be steady, non-Newtonian, incompressible, and laminar. The effects of other components in the blood, such as cells, proteins, ions, nutrients, and excreta, were not considered. Since the range of human body temperature changes is not obvious, the influence of temperature on blood was not taken into account.
Due to the low Reynolds number of blood, it was input as a laminar fluid with a density of 1060 kg/m 3 [26]. The Carreau model was used to describe the non-Newtonian characteristics of the blood and is given in Eq. (6): where   and 0  are the viscosities when the shear rate approaches infinity and zero, respectively,  and n are the material coefficients, and  is the shear rate.  [27][28]. Fig. 3b plots the relation between dynamic viscosity and shear strain rate based on the above parameters for the Carreau-Yasuda Model.

Governing Equations
In this study, the blood flow was assumed to be an incompressible, single phase and laminar flow with no impact of temperature. Therefore, only fluid mass and momentum conversation were considered, and the governing equations are the conservation laws of mass and momentum, given in Eqs. (7) and (8), respectively: where  is velocity,  is density, p is pressure, and ij  is shear stress, which is viscosity  times shear strain rate  : Simulations were conducted by using the CFD solver FLUENT (ANSYS, Inc., Canonsburg, PA, USA) with a finite-volume method. The flow equations were discretized for each cell in the model. The pressure-velocity coupled algorithm, capable of solving velocity and pressure fields simultaneously, was used to provide robust solutions.

Womersley Flow
Blood flow is a pulsatile flow with periodic variations, so we adopted the analytic  In this paper, the arterial wall was assumed to be a fixed, no-slip surface, and the hybrid flow diverter was also assumed to be a rigid body fixed on the arterial wall.

Observation Indicators
To predict the clinical behaviours of the hybrid flow diverter, the blood inflow and inflow velocity of an aneurysm, residence time, and wall shear stress were used as observation indicators in this study. The stability of the flow within an aneurysm, such as the blood flow rate or flow into an aneurysm, has been used in many studies to assess the risk of aneurysm rupture [30][31]. In this study, the blood flow velocity inside the aneurysm was analysed and the blood flow into the aneurysm was measured to assess the stability of the flow inside the aneurysm. In the flow field model of the intracranial aneurysm, the plane with a normal vector toward the inside of the aneurysm was set to monitor blood flow (Fig. 5a). The net blood flow through the plane was the sum of inflow and outflow: where net Q is the net blood flow, in Q is the net blood inflow, and out Q is the net blood outflow through the plane. The blood flow into an aneurysm can be obtained by calculating the dot product of the blood inflow velocity through the plane with the normal vector of the plane: (12) where n is the unit normal vector of the plane, A is the area of the plane that intersects an aneurysm, and in V is the blood inflow velocity through the plane, which can be obtained by the Heaviside step function.
Janiga et al. proposed the parameter of residence time [32], which can evaluate the influences of different aneurysm geometries, as defined below: where r t is residence time, in Q is the inflow of the aneurysm, and AS V is the volume of the aneurysm. The residence time represents the stability of blood flow within an aneurysm. The greater the residence time, the more stable the blood is, and thus the lower the risk of aneurysm rupture. Since the residence time considers the factor of aneurysm volume, differences in residence time can be compared even for aneurysms of different volumes.
After stenting, neo-intimal hyperplasia and thus restenosis occur in some patients.
Clinical results have shown that low wall shear stress or oscillating wall shear stress corresponds to the occurrence of the greatest neo-intimal hyperplasia in stented arteries [33][34]. Moreover, excessively high or low wall shear can induce neo-intimal hyperplasia, so wall shear stress has been seen as an important indictor to predict the potential risk of restenosis. In FLUENT, the definition of vascular wall shear stress is as follows: where v is the blood flow velocity, and n is the normal vector to the arterial wall. In this paper, the absolute value of shear stress is used as the observation indicator.
In addition, an intracranial aneurysm in an area of low flow velocity and low wall shear stress can easily lead to platelet thrombosis [35]. Some studies have also shown that high wall shear stress increases the risk of aneurysm rupture, so blood flow in an aneurysm with low wall shear stress is relatively favourable [36][37][38]. It has been suggested that shear stress of less than 5 dynes/cm 2 leads to endothelial proliferation of smooth muscle cells [39], so the size of the low wall shear area (that is, the area of wall shear stress less than 5 dynes/cm 2 ) on the arterial wall was also an important observation indicator.

Manufacturing Simulation
Figs. 6a and 6b show the strain distribution of the hybrid flow diverter when it was expanded to 4.0 mm ID and crimped to 1.5 mm OD. After the expansion, the balance between metal densities and feasibility was successfully achieved by the parametric design. The hybrid flow diverter could both prevent blood from flowing into an aneurysm and avoid stimulating the growth of the endothelial cells of the parent artery at the same time. Icon SDV24 (solution-dependent state variable 24) is the variable name in ABAQUS UMAT. In the hyperelastic material of UMAT, SDV24 represents variables for the equivalent elastic strain. As shown in Fig. 6a and 6b, in the process of either expansion or crimping, the regions near the crown indicated higher strain energy, while it was relatively low in the parts near the struts. The strain levels were 1.37% and 1.70% when the hybrid flow diverter was expanded to 4.0 mm ID and crimped to 1.5 mm OD, respectively. In addition, in both steps, the elastic strain of the nitinol material was kept within 12%. Thus, it was below the upper limit of safe working strain, indicating that the material of the hybrid flow diverter would not be damaged during the manufacturing process.

Fatigue Life Analysis
According to the simulation results, Goodman fatigue life analysis defined fatigue safety factors of the hybrid flow diverter in terms of normalized mean strain and strain amplitude (Fig. 6c).

Stent-to-Artery Ratio
A hybrid flow diverter with various metal densities is proposed in this paper to achieve different stent-to-artery ratios. Fig. 6d compares the metal densities from the lighter side to the denser side of the hybrid flow diverter after expansion. The different metal densities entailed dissimilar artery area coverages, and the stent-to-artery ratios were 19.58% and 10.42% for the sides facing the aneurysm and the parent artery, respectively. Therefore, the hybrid flow diverter indeed blocked the blood flow into an aneurysm with relatively high metal coverage and avoided excessive contact with the parent artery on the lighter side.

Deployment Simulation
Fig. 5c presents the simulation results of the hybrid flow diverter deployment.
When the hybrid flow diverter moved with the bending cylindrical sleeve and the centerline overlapped with the parent artery, the maximum strain value was 1.19%, which was still within the safe strain range of the nitinol material. The strain gradually decreased to 0.84% as the hybrid flow diverter was released in the parent artery.
Therefore, no material damage would take place during the deployment process. Fig.   5c also shows the final result of the hybrid flow diverter deployment to the parent artery.

Inflow Rate
Fig. 5b plots the inflow ratio of the surface monitor and inlet boundary before deployment of the hybrid blood diverter. The calculated results showed that the blood volume in the aneurysm accounted for about 3-6% of the blood volume at the inlet boundary. Fig. 7a plots the inflow rate measured by the surface monitor before and after deployment of the hybrid flow diverter, respectively. After deployment, the blood flow into the aneurysm was significantly reduced, by 75-95%. As shown in Fig.   7b, the hybrid flow diverter prevented most of the blood from flowing into the aneurysm.

Residence Time
The results on the residence time showed that the hybrid flow diverter significantly increased it (Fig. 8a).

Velocity Distribution
Since the maximum increase in the residence time was 0.79 seconds, that time was used to present the streamline and the contour of flow velocity for an intracranial aneurysm. Fig. 9a presents the contour plots of comparison for the streamline before and after the deployment of the hybrid flow diverter. The streamline indicated that the hybrid flow diverter successfully blocked the blood flow into the aneurysm, and only a small amount of blood entered the aneurysm through the hollows. As shown in the partial enlargement of the streamline in Fig. 9b, the blood flow into the aneurysm was substantially reduced. Fig. 10a shows the central section of the flow velocity contour before and after the deployment of the hybrid flow diverter. From the partial enlargement of the flow velocity shown in Fig. 10b, it is clear that the blood flow rate inside the aneurysm decreased sharply after the deployment, and the aneurysm also maintained a stable blood flow inside it. Therefore, the hybrid flow diverter effectively blocked the blood flow into the aneurysm.

Wall Shear Stress distribution
Since the maximum increase in residence time was 0.79 seconds, that time was also used to present the contour of wall shear stress distribution for an intracranial aneurysm. Fig. 11a shows the wall shear stress distribution before and after the deployment of the hybrid flow diverter. Due to the faster flow velocity in the aneurysm before the deployment of the hybrid flow diverter, only a few areas in the aneurysm had low shear stress (< 5 dynes/cm 2 ). After the deployment of the hybrid flow diverter, the reduced blood flow into the aneurysm was accompanied by a decrease in the blood flow rate, so the wall shear stress of the aneurysm also fell sharply, to mostly less than 5 dynes/cm 2 . As an area of low flow rate and low wall shear stress is prone to thrombosis due to platelet aggregation, an aneurysm would eventually heal as the thrombus filled up the entire lumen over time. According to some studies, higher wall shear stress will increase the risk of aneurysm rupture.
Therefore, the hybrid flow diverter not only reduced the wall shear stress of the aneurysm but also alleviated the risk of aneurysm rupture. Fig. 11b shows a partial enlargement of the wall shear stress distribution contour after the deployment of the hybrid flow diverter. In the parent artery, the region near the crowns of the hybrid flow diverter had a low wall shear stress distribution. This phenomenon was the same as those in previous studies, which indicated that the reduced wall shear stress promoted endothelial cell proliferation, thus resulting in stenosis. Therefore, to avoid neo-intimal hyperplasia of the parent artery due to high metal coverage, the hybrid flow diverter was designed with a lower crown number on the side facing the parent artery so that the area of low wall shear stress could be reduced and excessive endothelial cell proliferation could be prevented.

Conclusions
A