¤ ( A ) A high-grade stenotic lesion with 

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... of stent struts around the bridge following expansion. The authors found that stents with V- or S-shaped links were easier to inflate than the Palmaz-Schatz stent, but tended to fore- shorten more. S-shaped links are more flexible than V-shaped links. The expanded diameters of the stent only, balloon only, and stent and balloon were mapped as a function of internal pressure. The stent size increased nonlinearly from the original 3-mm diameter to , 5.5 mm as the internal pressure increased from 0 to , 0.4 MPa. This study demonstrated that modeling a quarter of the stent can reduce computational time while keeping the same measurements of displacement and stress (measurement differences of , 1.0%). Ju and others 18 expanded upon Xia’s work by modeling computationally the closed-cell stents with RUC and with a free end (RUC + ). The study compared 3 models (Panel, RUC, RUC + ) applied to 2 stents (Palmaz-Schatz and a sinusoidal stent); they found good agreement with previous studies without losing computational accuracy by way of RUC + , yielding the most accurate results for the inner surface and ends of the stent. Computational models of stents in the coronary artery were created by Lim et al. 19 with open or closed cells linked by bend- shaped structures. A realistic, transient, non- uniform balloon expansion process was ana- lyzed using FEM. Seven commercially available stents (Palmaz-Schatz PS153, TenaxTM, MAC Standard, MAC Q23, MAC Plus, Coro- flex, and RX Ultra Multi-link) were compared to recommend design parameters that could reduce the risk of restenosis due to foreshortening or dog-boning (flaring of stent ends due to excessive balloon expansion). Although dog-boning is an issue related to balloon-expandable stents, Lim’s modeling can be applied to self-expanding stents to suggest design parameters. Foreshortening ranged from 0.0% to 12.9% and dog-boning values were all positive, indicating overex- pansion of the distal portion of the stent. Hence stents with closed cells tend to have greater foreshortening and dog-boning. Bedoya et al. 20 investigated stent design parameters based on vessel stress distribu- tion and radial displacement using a generic stent model with varying connector bar lengths (strut spacing), radius of curvature at the crown junctions, and axial amplitude. Strut spacing changed the results most drastically in that small strut spacing (1.2 mm versus 2.4 mm) could induce higher stresses over a larger region. The authors recommended that future stents be designed with large strut spacing, blunted corners at bends, and high axial amplitude. Timmins et al. 21 expanded upon Bedoya’s work by using FEM to numerically evaluate arterial wall stress, lumen gain, and cyclic deflection (change in radial position from diastolic to systolic pressure) of a stent. An algorithm was developed to optimize the stent design using single values representing these 3 variables in the computational model. The goal of the algorithm was to maximize lumen gain and cyclic deflection while minimizing wall stress, using different weighting factors to emphasize the importance of various parameters. When circumferential stress re- duction or a cyclic deflection increase was emphasized, strut spacing and axial amplitude increased while radius of curvature decreased. Increasing the emphasis on lumen gain resulted in a decrease in strut spacing and axial amplitude and an increase in radius of curvature (Fig. 11). With the recent introduction of intravascular ultrasound (IVUS) with virtual histology (VH, Fig. 12A), it may be possible to better understand the influence of lesion characteristics (e.g., plaque thrombogenicity) on the significance of cell size, stent flexibility and conformability, and plaque prolapse potential, all of which are predictors of stroke. The angiographic anatomy may provide a better understanding of lesions more demanding of flexibility or scaffolding, while IVUS-VH might suggest lesions that are thrombot- ically active and not suited for stenting regardless of cell design (Fig. 12B,C). It is reasonable that the cell structure itself can be designed to act as its own embolic protection filter, i.e., an increase in surface area coverage may prevent plaque prolapse and em- bolization, effectively trapping plaque frag- ments between the stent and the vessel wall. With VH and an improved understanding of the lesion characteristics, we are better equipped today to choose the most appropriate stent. Early clinical experience with helically designed stents has demonstrated excellent flexibility and kink resistance. The helical design allows uniform cell size at flexion points without scaling. In terms of clinical applications, the calcified and tortuous lesion may be best suited to the more conformable open or helically designed stents. Greater scaffolding may be more appropriate for vulnerable plaque or ulcerative lesions with type C characteristics (i.e., with high anatomical risk, Fig. 13). It is evident that there remain unresolved clinical issues and the need for improvement in stent designs. While we look forward to robust datasets and device evolution, it is important that the interventional community remains mindful of the lessons of recent clinical trials in both the US and Europe. An interesting observation in the CAPTURE study was the existence of a difference (although not statistically different) in outcomes among inexperienced versus experienced operators. 22 Major stroke and death occurred at a 3.4% event rate in the lesser experienced operators compared to 1.1% in the operators with significant experience. Furthermore, an all stroke and death rate comparison of inexperienced and experienced operators also showed a difference of 6.9% versus 4.6%, respectively. The lesson to be learned is that interventionists can and do achieve excellent stenting results when they are well trained in the use of the devices, the stenting protocol, and appropriate patient selection. Current stent geometries can be evaluated using computational modeling tools, which also aid in the design and optimization of new stents that improve upon the shortcomings of the existing devices in the market. Computational studies can parametrically vary stent geometries, which give insight into which configurations have the most favorable effects on the vessel wall and local hemodynamics. Virtual design of medical devices using FEM reduces the prototype building and evaluation periods. The desired features of a carotid artery stent include (1) scaffolding adequate to control plaque prolapse but with acceptable flexibility, (2) conformability, and (3) radial strength to track the lesion, appose the vessel wall, and control recoil. Evidently, no matter how tight the scaffolding, there is still no stent available on the market to control emboli # 100 m m in diameter, which comprise most of the emboli produced during CAS. When used with an efficient distal protection filter and combined with a flow reversal system, we may be able to reduce most of the embolic events occurring in the periprocedural period. Nonetheless, pre- dicting and controlling delayed post-procedure events remains a challenge. By means of FEM and design optimization algorithms, new stent configurations can be evaluated in a simulated environment prior to prototype building. The modeling issues relevant to stent design include (1) curvature changes, (2) compliance mismatches, and (3) low wall shear stress. Previous studies have found that stents with shorter and fewer struts, as well as reduced thickness, yield better results. Compliance-matching stents aid blood flow and reduce stress concentrations. Careful consideration must be taken in the design of the stent edges, as the ends have a tendency to induce elevated mechan- ical stresses, cause large curvature changes that lead to low wall shear stress, and damage the arterial wall after ...