Heart Vessels DOI 10.1007/s00380-015-0665-1

ORIGINAL ARTICLE

Repression of wall shear stress inside cerebral aneurysm at bifurcation of anterior cerebral artery by stents Ryuhei Yamaguchi1 · Gaku Tanaka1 · Hao Liu1 · Hiroshi Ujiie2 

Received: 29 August 2014 / Accepted: 20 March 2015 © Springer Japan 2015

Abstract  The effect of a simple bare metal stent on repression of wall shear stress inside a model cerebral aneurysm was experimentally investigated by two-dimensional particle image velocimetry in vitro. The flow model simulated a cerebral aneurysm induced at the apex of bifurcation between the anterior cerebral artery and the anterior communicating artery. Wall shear stress was investigated using both stented and non-stented models to assess the simple stent characteristics. The flow behavior inside the stented aneurysm sac was unusual and wall shear stress was much smaller inside the aneurysm sac. Stent placement effectively repressed the temporal and spatial variations and the magnitude of wall shear stress. Hence, there is an effective possibility that would retard the progress of cerebral aneurysms by even simple stent. Keywords  Cerebral aneurysm · Bifurcation · Wall shear stress · Stent · Hemodynamics

* Ryuhei Yamaguchi [email protected] Gaku Tanaka [email protected]‑u.jp Hao Liu [email protected]‑u.jp Hiroshi Ujiie [email protected] 1

Graduate School of Engineering, Chiba University, 1‑33 Yayoicho, Inage‑ku, Chiba 263‑8522, Japan

2

Department of Neurosurgery, Tokyo Rosai Hospital, 4‑13‑21 Oomori, Ohta‑ku, Tokyo 337‑8570, Japan



Introduction Hemodynamics plays an important role in the progress of cerebral aneurysms. Generally, there is significant evidence to suggest that geometry greatly influences hemodynamics and that hemodynamic forces are important modulators of vascular structure [1–3]. The wall shear stress (WSS) caused by the velocity gradient also affects the platelet aggregation and has an effect on blood viscosity and the intravascular clotting mechanism [4–6]. Temporal and spatial variations and the magnitude of wall shear stress inside aneurysm sacs are thought to correlate with the progress of aneurysms [7]. Recently, cerebral aneurysms have recently been treated with stents accompanying fine platinum wire microcoils inserted into aneurysms [8]. Juszkat et al. [9] inserted stents into cerebral aneurysms with wide necks which is difficult to clip [10]. Follow-up angiography in 1 month showed the decrease of flow rate into the aneurysm and it resulted in the occlusion of aneurysm sac. Benndorf et al. [11] also recognized that aneurysms of the vertebral artery placed two overlapping stents became occluded in 3 months. Many computational and several experimental studies have examined flow structure inside aneurysm sacs along the side wall of straight tube after stent deployment [12– 18]. Although stents seem to have occlusive effects, the relationships between stents and flow behavior, particularly quantitative repression of WSS, have not been adequately clarified experimentally. Low WSS would be closely associated with aneurysm occlusion [7]. Flow diverter (FD) stents have recently been developed and their effects on flow velocity and WSS have been investigated using computational fluid dynamic approach [7, 19–21]. However, instent restenosis is a common phenomenon with bare metal stent in the early stage after deployment [22]. Also, the

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existence of aneurysm might affect to pulsatile energy loss along the major aorta [23]. Furthermore, high WSS might be associated with the intimal tear in patients having overt aortic dissection [24]. The present flow model basically simulated a saccular cerebral aneurysm located at the apex of bifurcation [25] between the anterior cerebral artery ACA (A1–A2) and the anterior communicating artery (ACoA), where aneurysms often develop and rupture. The present model which has an aneurysm at the bifurcation point also simulates such as internal carotid, posterior communicating, basilar and middle cerebral arteries. The effect of stents on the temporal and spatial variations and the magnitude of WSS were experimentally clarified by flow visualization using 2D particle image velocimetry (PIV) in vitro. The Reynolds number and the non-dimensional frequency corresponded to the range of flow parameters in the human cerebral artery. WSS was estimated in stented and non-stented models as an index to assess stent characteristics. Flow structure in stented model is considerably different from that in non-stented model and WSS was noticeably decreased after stent deployment. Furthermore, the minimal temporal and spatial variations in WSS were quantifiably repressed in stented model.

Methods Figure  1 shows a schema of the geometry of the cerebral aneurysm model. The flow model simulates a saccular cerebral aneurysm at the apex of bifurcation [25] between the ACA (A1–A2) and ACoA. The model was connected to an afferent tube (diameter 8 mm) representing the proximal section of the ACA (A1), and to efferent tubes (diameter 6 and 3 mm) representing the distal section of the ACA (A2) and a section of the ACoA, respectively. The scale of the model was approximately fourfold size of the human artery [26]. The aneurysm was an ideal spherical with a neck width (N) and depth (D) of 6 and 12 mm, respectively, and an aspect ratio (AR) of 2.0 [27]. The flow model (R’Tech, Hamamatsu, Japan) is made of silicone resin (Shin-Etsu Chemical, Tokyo, Japan) with a refractive index of 1.41. The material of the simple uncoated bare metal stent (feel stent) (PIOLAX Inc., Yokohama, Japan) is nitinol, the outer diameter and length are 8 and 40 mm, respectively, and the wire diameter, strut width, and porosity are 0.25, 8 mm, and 95 %, respectively. The working fluid was 53 % aqueous glycerin (H2O:C3H8O3; 1:1.128) with a refractive index identical to that of the silicone resin (1.41), a density of 1.13 g/cm3, and a kinematic viscosity ν of 7.00 × 10−6 m2/s at 293 K (20 °C).

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Fig. 1  Schema of aneurysm model with stent (AR = 2.0)

This kinematic viscosity is approximately two times larger than that of blood. To prevent the deflection of laser sheet, it is necessary to employ this working fluid with the refractive index identical to that of the silicon resin. Furthermore, as in this aneurysm model it is difficult to realize this size of 2.2 mm as diameter of ACA (A1) [26], the diameter 8 mm of afferent tube at the proximal section in the ACA (A1) was selected approximately four-fold larger than that of human ACA (A1). Under these conditions, we adjusted to the mean Reynolds number of about 400 in the ACA (A1) [26]. Velocity vector was measured by visualization of the flow field by PIV at the median plane of the spherical aneurysm sac, and signals were processed using DaVis 7 software (LaVision, GmbH, Goettingen, Germany). An imageintense CCD camera (1376 × 1040 pixels) was fitted with an AF NIKKOR lens (Nikon, Tokyo, Japan) and the light source was a 15 Hz double pulse Nd-Yag laser with 30 mJ output. A cylindrical lens produced a laser sheet thickness of 0.3 mm. The scattered material comprised FLUOstar fluorescent polymer microspheres (diameter 14 μm; density 1.1 g/cm3; EBM Co. Ltd., Tokyo, Japan). The spatial resolution of these measurements was 0.2 mm. Wall shear stress was estimated from the velocity in the vicinity of the wall of the aneurysm, based on a linear velocity gradient obtained from the velocity at 0.6 mm from aneurysm wall and that (zero velocity) at aneurysm wall. Although the cerebrovascular flow rate changed in a manner resembling a triangular wave [26], the flow rate was approximated as a sinusoidal waveform defined by the equation Re = Rem(1 + Asinωt), where Rem, A and ω are the mean Reynolds number, flow rate amplitude, and

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pulsatile angular velocity, respectively. The control valve has two slit apertures of 12 mm × 2 mm (width × height). The aperture area sinusoidally changed around a mean area in the direction of width via an eccentric cam. Therefore, the flow rate is expressed as a summation of the steady and pulsatile flow rates [25]. The experiment proceeded under sinusoidal laminar pulsatile flow with a nondimensional   √ pulsatile frequency (Womersley number) α of d0 2 ω/ ν  = 4.0, a mean Reynolds number (Rem) of d0U/ν of 435, one pulsatile cardiac cycle T of 0.9 s, and a flow rate amplitude A of 0.68, where d0 and U are the diameter of the afferent tube ACA (A1) and the mean velocity within ACA (A1), respectively.

Results The effect of the stent on flow structure was examined by PIV at the median plane inside the spherical aneurysm sac. Figure 2 shows the velocity vector inside the spherical aneurysm sac without stent. Pulsatile flow vector is shown as an ensemble averaged from the imaged frame data throughout 50 cardiac cycles at the following four typical phases: early, peak systoles, and mid-, late diastoles. The circle containing an arrow inside the aneurysm sac indicates the direction of vortex rotation and the width of circle line depends on velocity strength. One vortex rotates clockwise inside the aneurysm sac throughout one cardiac cycle. At peak systole ωt  =  π/2, flow velocity became maximum and reached 20 cm/s around the proximal side e of the aneurysm neck. Although the flow rate at early systole ωt  = 0 was the same as that at mid-diastole ωt = π, the velocity was much lower at early systole than at mid-diastole. This difference would relate to the phase lag owing to flow inertia. Furthermore, although the flow rate at late diastole ωt = 3π/2 corresponding to Re = 140 in steady flow was minimal, its clockwise velocity was little larger than that at early systole ωt = 0 corresponding to Re = 435. The impact point indicates in small arrow at the proximal edge e of aneurysm neck at peak systole ωt = π/2 and mid-diastole ωt = π would be related to the progress of cerebral aneurysm. Similarly, Fig. 3 shows the velocity vector inside the spherical aneurysm sac with a deployed stent at early, peak systoles, and at mid-, late diastoles. Flow structure considerably differed between the stented and non-stented models. As indicated in small arrow, the flow impact point moved to the center of aneurysm neck from the proximal edge. In the stented model, the velocity throughout one cardiac cycle was generally much slower. This decrease was generated by a reduction of the flow rate into the aneurysm sac due to the lead and resistance created by the stent. A smooth rotating vortex was not found except for ωt = 0 and

Fig. 2  Velocity vector within aneurysm model in pulsatile flow without stent (Rem = 435, α = 4.0, A = 0.68)

Fig. 3  Velocity vector within aneurysm model in pulsatile flow with stent (Rem = 435, α = 4.0, A = 0.68)

3π/2. At peak systole ωt = π/2, the flow inside the aneurysm sac split into one large and another small vortices. Velocity was comparatively high within the large vortex at the distal wall of the aneurysm and the flow collided at the bottom wall of aneurysm. This situation also arose from the reduced flow rate into the aneurysm sac by stent. However, the flow rate through ACoA was approximately maintained to be equal to that through ACA (A2) because both flow rates were controlled by two apertures at outlet issuing into the atmosphere. Velocity at the proximal side of the aneurysm neck was much smaller at peak systole and mid-diastole in the stented model than in the non-stented model. In the non-stented model, the fully developed flow approaches to the neck of aneurysm sac from ACA (A1). In the stented model, the velocity profile would change owing

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Fig. 4  Distribution of normalized and absolute wall shear stress along aneurysmal wall in pulsatile flow without stent. (Rem = 435, α = 4.0, A = 0.68)

Fig. 5  Distribution of normalized and absolute wall shear stress along aneurysmal wall in pulsatile flow with stent. (Rem = 435, α = 4.0, A = 0.68)

to the lead and resistance of stent. For example, the velocity profile in front of the aneurysm neck in steady flow would be parabolic and flat, in the nonstented and stented models, respectively. Figure 4 shows the distribution of WSS along the aneurysm wall at four typical phases in the non-stented model. The ordinate denotes the absolute non-dimensional WSS normalized by WSS in Poiseuille flow corresponding to a mean flow rate of Re = 435 at the afferent section of ACA (A1). Wall shear stress was small and moderate along θ  = 0–180°, large around the proximal side of the aneurysm neck at θ = 210–285°, and it at θ = 270° reached a maximal value of τw  = 1.9 at peak systole. The proximal side would be generally exposed to high WSS and the averaged WSS with respect to temporal and spatial variations was 0.38. Figure  5 shows the distribution of WSS at four typical phases in the stented model. WSS never varied smoothly, unlike the non-stented model shown in Fig. 4. WSS at ωt = π/2 had two peaks at θ = 180°, 240° and one minimum value at θ  = 210° around the bottom wall of aneurysm that was related to the vortex structure and flow split shown in Fig. 3. WSS at θ = 180° around the bottom wall of aneurysm reached a relatively high τw of 0.82 at peak systole ωt  =  π/2, and the averaged WSS with respect to the temporal and spatial variations at four typical phases was 0.2. A high WSS would damage the arterial lumen and would be involved in the progress of the aneurysm. The WSS magnitude of 0.3 around the proximal side of aneurysm neck in the stented model was much smaller than that of 1.9 in the non-stented model and stagnant flow appeared inside the aneurysm sac. The low WSS would release the load on the arterial lumen.

Discussion

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While the effect of stents on flow structure inside the cerebral aneurysm sac has been recently simulated using a computational fluid dynamic approach [7, 17, 19–21], the experimental studies are few yet. In this study, we tried to experimentally clarify flow characteristics such as WSS and velocity vector around and inside the model of an ideal spherical aneurysm in vitro using PIV. Although there is a limitation in the present study and the flow behavior was only evaluated at the median plane of spherical aneurysm, WSS and velocity vector is strongly emphasized at this median plane which is a characteristic plane. Generally, the averaged WSS (0.20) of the temporal and spatial variations in the stented model was by 47 % decreased from that (0.38) in the non-stented model. Even the present simple stent with 95 % porosity remarkably reduced the temporal and spatial variations as well as the magnitude of WSS. Although Y stent is recently used for the treatment of aneurysm at the apex of bifurcation [28], the present simple bare metal stent effectively reduces WSS. Particularly, at peak systole, the comparison of WSS (0.82) at θ  = 180° around the bottom of the aneurysm in the stented model indicates the decrease by 55 % from WSS (1.8) at θ  = 180° of the proximal edge in the nonstented model. The reduced magnitude and the spatial gradient of WSS might release hemodynamic force against the arterial lumen (endothelial cells) exposed to WSS. Although in these days coil embolization is often used for the treatment of cerebral aneurysm, the present study has suggested that the combination of stent deployment with coil embolization would be more efficient for the treatment of aneurysm [8]. When the coil was applied for

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the treatment of aneurysm, the flow impact point still keeps at the proximal edge of aneurysm [29]. No movement of flow impact point after coil treatment still grow the aneurysm at this same point and the lumen will be exposed high WSS load. However, in the stented model, the flow impact point clearly moves to the center of the aneurysm neck as indicated in small arrow in Fig. 3. It would be related to retard the progress of aneurysm, and the proximal edge was protected. The movement of flow impact point is an important pathological meaning, i.e., arterial lumen would be released from the progress of aneurysm. Stent deployment abruptly changed the distribution of WSS at the proximal side of the aneurysm. Wall shear stress at 270° along the aneurysm wall of the non-stented model reached 1.8, which corresponds to 2.7 Pa as a dimensional value [30]. Without stent, the inflow jet impacted at the proximal edge of the aneurysm and promoted its progress. The WSS at the same location drastically decreases by 1.6 at 270° upon stent deployment. This reduction indicates that reduced WSS around the proximal side of aneurysm would prevent its own progress, which is of important pathological significance. The movement of impact jet point will retard the progress of aneurysm. Surely, in the stented model, there might be a little possibility of initiation bleb around the bottom because WSS reaches relative high 0.8 at 180° at ωt  =  π/2. However, stent deployment globally decreases WSS compared with the non-stented model and the decrease of WSS might be related to repress of aneurysm growth. On the contrary, the repression of WSS at bottom wall might promote the platelet aggregation and thrombosis. Many types of stents and flow diverter stents have been developed. The general objective of stent deployment is to reduce the flow rate into an aneurysm so that flow becomes stagnant inside the aneurysm sac. Although for present model Y stent is much effective for the reduction of WSS and the occlusion of aneurysm, the daughter stent in Y stent needs to insert through narrow mesh of the parent stent and is much expensive. It will be significant valuable that even current simple stent could reduce by at least 55 % of peak WSS. Accordingly, the present experimental results endorse the notion that stent deployment reduces WSS. Stents effectively reduce WSS as described above. In the present experiment, stent was set a little apart from the neck of aneurysm. Ideally, stent should place just at aneurysm neck. However, even if stent might be apart from the neck of aneurysm, the flow rate into the aneurysm was reduced and WSS was drastically decreased at the proximal edge of aneurysm neck. Furthermore, the stent deployment for aneurysm at bifurcation point such as anterior communicating, anterior cerebral, internal carotid, posterior communicating, basilar and middle cerebral arteries is quite effective. Consequently, the reduction of WSS, the

promotion of stagnant flow would retard the progress of cerebral aneurysms.

Conclusions We have carried out an experimental study based on PIV in vitro and confirmed that a stent deployed at an aneurysm induced in a bifurcation can reduce WSSs. This study has suggested that the small variations and low magnitudes of WSS around and inside the aneurysm sac would retard the progress of cerebral aneurysm by even simple bare metal stent. Although the present paper investigates the flow structure within human cerebral aneurysm, this study does not use human subject and it is the research in vitro. With respect to the present study, the authors declare no conflicts of interest. Acknowledgments  This study was supported by a Grant-in-Aid for Science and Culture (#24560208), from the Ministry of Education, Culture, Sports, Science, and Technology, Japan.

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