Abstract
The effect of inflow waveform on the hemodynamics of a real-size idealized sidewall intracranial aneurysm (IA) model was investigated using particle imaging velocimetry (PIV). For this purpose, we implemented an error analysis based on several PIV measurements with different time lags to ensure high precision of velocity fields measured in both the IA and the parent artery. The relative error measured in the main part of the circulating volume was <1 % despite the three orders of magnitude difference of parent artery and IA dome velocities. Moreover, important features involved in IA evolution were potentially emphasized from the qualitative and quantitative flow pattern comparison resulting from steady and unsteady inflows. In particular, the flow transfer in IA and the vortical structure were significantly modified when increasing the number of harmonics for a typical physiological flow, in comparison with quasi-steady conditions.
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Notes
In this adaptive correlation algorithm, the interrogation windows are refined twice from \(128 \times 128\) pixels to \(32 \times 32\) pixels i.e \(\approx 1.3\times 1.3\)mm to \(\approx 0.33 \times 0.33\,\hbox {mm}\).
Adaptive PIV correlation algorithm (Dynamic Studio, Dantec Dynamics) was used to compute the velocity fields in IA model. This algorithm optimizes the size of the interrogation windows (\(32 \times 32\) pixels to \(64 \times 64\) pixels i.e. \(\approx 0.28 \times 0.28\,\hbox {mm}\) to \(\approx 0.55 \times 0.55\,\hbox {mm}\)) according to the local density of seeding particles and the velocity gradient.
Or unsteadiness when considering a steady inflow.
For a 2D velocity field \(\sigma _{v}\approx [(\sigma _{v_x}v_x)^2+(\sigma _{v_y}v_y)^2]^{1/2}/v\).
\({\mathfrak {I}}\) is the imaginary part of component in parenthesis.
Note that this effect becomes negligible in larger r = 3 mm tubes (not shown here).
The corresponding Reynolds numbers are \(Re = 68, 136, 273, 545, 1,091\).
The quasi-steady approximation consists in approximating the time dependence of pulsatile flow at time t by the steady velocity field with inlet flows \(q(t)\). The time dependence of each component, \(v_{x,y}\), of the quasi-steady velocity field was interpolated locally using a linear interpolation between the five steady flow measurements at \(q = 1, 2, 4, 8, 16\,\hbox {ml/s}\) shown in Fig. 4b.3–4 and Fig. 5.
We define the velocity exchange at the neck surface as \(v^{\mathrm{ex}}_{S_{\mathrm{N}}} = 1/|S_{\mathrm{N}}|\int _{S_{\mathrm{N}}}|\mathbf{v}\cdot \text {d}{\varvec{\sigma }}|\). This quantity measures the flow transfer between the parent artery and IA, normalized to neck area.
\(q_{\mathrm{max}}\) and \(q_{\mathrm{min}}\) are the maximum and minimum inflow rates during the cardiac cycle.
rect is the rectangular function.
Close to \(y_j\approx 0\) (in the center of the cylinder) the linear system becomes unstable. It is therefore recommended to use a polynomial development of \(v_x(y_j)\) to solve Eq. (10).
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Acknowledgments
We would like to thank the technical support of the LMH for the realization of the experimental setup. P. Bouillot thanks the Vasco Sanz Foundation for its support.
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Appendix: Laser thickness correction
Appendix: Laser thickness correction
Due to its finite thickness, h, the laser sheet enlighten particles located in different out-of-planes distances from the center of the measurement plane. The computation of the velocity field from the recorded pictures by means of cross-correlations is therefore influenced by all the velocities in the illuminated sheet. Assuming a cylindrical symmetry of the actual velocity profile \(v_x(y,z) = v_x(s)\) with \(s = \sqrt{y^2+z^2}\), the measured velocity profile in the symmetry plane at the \(j^{\text {th}}\) radial measured location, \(v^{\text {mes}}_z(y_j)\), can be written as a weighted sum
w(z) weights the effect of each illuminated particle on the computed velocity field according to their out-of-plane distance, z. The second equality comes from a Taylor expansion of \(v_x\) at fixed \(y = y_j\) and for \(z = 0\). \(w_n = \int z^nw(z)\text {d}z\) is the nth momentum of w(z).
For a uniform weighted functionFootnote 11 \(w(z) = h^{-1}\text {rect}\left( z/h\right)\) (i.e. assuming that all the enlighten particles have the same effect on the computed velocity fields) and keeping only the first \(n\leqslant 2\) components of the Taylor expansion in (10), we can write
Using the finite difference approximation \(\text {d} v_x /\text {d} s|_{s = y_j}\approx [v_x(y_{j+1})-v_x(y_{j-1})] /(y_{j+1}-y_{j-1})\), Eq. (11) becomes a linear system which can be solved forFootnote 12 \(v_x(y_j)\). The results shown in Figs. 3 and 8a assume a measured laser thickness \(h = 1\,\hbox {mm}\).
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Bouillot, P., Brina, O., Ouared, R. et al. Multi-time-lag PIV analysis of steady and pulsatile flows in a sidewall aneurysm. Exp Fluids 55, 1746 (2014). https://doi.org/10.1007/s00348-014-1746-0
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DOI: https://doi.org/10.1007/s00348-014-1746-0