Computational Investigation of the Delamination of Polymer Coatings During Stent Deployment
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- Hopkins, C.G., McHugh, P.E. & McGarry, J.P. Ann Biomed Eng (2010) 38: 2263. doi:10.1007/s10439-010-9972-y
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Recent advances in angioplasty have involved the application of polymer coatings to stent surfaces for purposes of drug delivery. Given the high levels of deformation developed in the plastic hinge of a stent during deployment, the achievement of an intact bond between the coating and the stent presents a significant mechanical challenge. Problems with coating delamination have been reported in recent experimental studies. In this paper, a cohesive zone model of the stent–coating interface is implemented in order to investigate coating debonding during stent deployment. Simulations reveal that coatings debond from the stent surface in tensile regions of the plastic hinge during deployment. The critical parameters governing the initiation of delamination include the coating thickness and stiffness, the interface strength between the coating and stent surface, and the curvature of the plastic hinge. The coating is also computed to debond from the stent surface in compressive regions of the plastic hinge by a buckling mechanism. Computed patterns of coating delamination correlate very closely with experimental images. This study provides insight into the critical factors governing coating delamination during stent deployment and offers a predictive framework that can be used to improve the design of coated stents.
KeywordsPolymer coatingStent deploymentDelaminationCohesive zoneFinite element
Stent coating failure in vivo could lead to several negative clinical complications. Damage to a coating may increase the risk of thrombosis. The dosage of a drug contained in the coating could be reduced or the delivery misguided.14 The breakdown of the coating barrier may allow the release of irritant or toxic ions from the metallic stent, leading to an inflammatory response. Debris from the coating caught in the blood stream could also create micro-embolisms.9,13
The complex geometry of a balloon-deployed stent gives rise to localized regions of high plastic strain which correlate to areas where delamination is predominant as shown in Fig. 1. In order to adequately design coatings to withstand large deformation, knowledge of the mechanisms attributing to delamination is essential.
In this study, a finite-element model is developed whereby a cohesive zone law is applied at the interface of the stent and coating to simulate delamination during the stents crimping and deployment stages. Cohesive zone laws have been established as a means to simulate the decohesion of surfaces.2,17,27,30 A cohesive zone model has been applied in previous work where Abdul-Baqi and Van der Giessen1, 2 simulated the nanoindentation of a coating. Also using this approach, McGarry et al.17 and McGarry and McHugh16 simulated the detachment of cells from an underlying substrate. The delamination of a coating undergoing cyclic heat loads was simulated by Hattiangadi and Siegmund8 using an adapted Xu Needleman cohesive zone model.
In this study, it is hypothesized that debonding of polymer coatings from highly strained regions of the stent surface during deployment can be accurately modeled using a cohesive zone framework at the stent–coating interface. The first objective of this computational study is to investigate the mechanisms of debonding of biomedical stent coatings during stent deployment using a mixed mode cohesive zone model. Secondly, the design parameters which govern the initiation of coating debonding during deployment are analyzed. Finally, a case study is presented in which our modeling framework is applied to two commercially adapted stent designs and three relevant biomedical grade polymer coatings. It is demonstrated that our computational framework is capable of predicting patterns of coating debonding reported in previous experimental studies.
Cohesive Zone Formulation
The cohesive zone model is incorporated into ABAQUS™ Standard (V.6.7-1, ABAQUS Inc., RI, USA) via a user-defined interface subroutine (UINTER). This interface model is applied along the contacting stent and coating surfaces.
Due care is required when selecting the input parameters to the cohesive zone model. The interface bond strength in the normal and tangential directions is assumed equal so τmax = σmax. The normal and tangential characteristic length scales δn and δt are also set equal and given a value of 0.5 μm. This results in the coupling energy parameter are q ≈ 0.43.
After being coated, a stent is crimped onto a balloon and later expanded to a set deployment radius once inside the stenosed artery. To simulate these mechanical processes, boundary conditions on the unit cells are applied primarily in the circumferential direction to insure periodic deformation of the unit. Large deformation kinematics is assumed for all simulations in this study. Preliminary studies were performed using the generic ‘U-shaped’ hinge model to determine appropriate mesh density for the simulations.
Overall, results show close compliance with the predictions of the design curve established in Fig. 8. The dimensionless interface strength values for the SIBS coating on Hinge A of Design 1 and Hinges C and D of Design 2 lie above the design curve; therefore, debonding of the coating is not predicted to occur, even for large εc/L values. From Fig. 12a, the coating remains fully attached to Hinges C and D even when the stent is over-deployed. This can be attributed to the large radius of curvature of these hinges, which leads to an increased dimensionless interface strength, as dictated by Eq. (8).
Figure 11 shows delamination of the coating that occurs on stent Design 1 at a circumferential strain of εc = 1.8. Parameters have been chosen based on delamination observed experimentally (Fig. 9a). As shown in Fig. 10c, tensile delamination initiates at εc/L = 0.4 mm−1 (εc = 0.16). However, from Fig. 11, further deployment results in buckling of the coating from the stent surface in compressive regions of the plastic hinge. This phenomenon first occurs at εc = 0.636 in the vicinity of a geometric discontinuity where a connector strut intersects the plastic hinge. At εc = 1.1, a second occurrence of compressive buckling initiates in a region of high compressive stress on the plastic hinge. From Fig. 11, significant amounts of the coating are predicted to detach from the stent surface following deployment to εc = 1.8.
Figure 12a illustrates the predicted pattern of coating debonding at εc = 1.8 for stent Design 2 with a SIBS coating and an interface strength of σmax = 0.05 MPa. As shown in Fig. 10c, tensile delamination only occurs at hinge B (initiating at εc/L = 4.18, or εc = 1.17), with the coating remaining fully attached to the stent at hinges C and D. Debonding of the coating by buckling on the compressive side of Hinge B is computed to initiate at εc = 1.57 and, as can be seen from Fig. 12a, significant removal of the coating on the compressive side of the hinge is predicted at εc = 1.8. An identical pattern of debonding on the tensile and compressive side of a similar plastic hinge has been reported experimentally by Regar et al.,24 as shown in Fig. 12b.
For the first time a computational modeling framework has been developed to simulate stent coating delamination. We have demonstrated that cohesive zone modeling of a stent–coating interface results in the simulation of debonding patterns of polymer coatings during stent deployment that correlate closely with debonding patterns reported in experimental studies. Furthermore, our simulations have allowed for the identification of critical parameters that govern delamination of stent coatings during deployment.
Simulations of coating delamination on stent geometries reveal that debonding occurs mainly in the high strain plastic hinge regions. This result is supported by previous experimental studies of stent coatings using SEM analysis.14,20,21,24 This study reveals that the initiation of coating debonding in the tensile region of the plastic hinge is dependent on the coating thickness, the coating material, and the curvature of the hinge. Stiff, thick coatings on plastic hinges of high curvature are shown to debond at lower levels of stent deployment. Clearly an increase in interface strength between the stent surface and the coating will delay the on-set of coating debonding. A dimensionless grouping of the critical parameters that govern debonding initiation in tensile regions has been identified. For a required level of stent deployment for any stent design, a critical dimensionless interface strength can be identified, above which tensile delamination will not occur. This should allow for safe design of coating thickness, stiffness, and bond-strength and should provide a significant step toward elimination of coating delamination.
Results demonstrate that debonding initiation at the tensile surface of the plastic hinge of a polymer-coated stent is predominantly Mode I. However, when further stent deployment occurs following debonding initiation, the coating peels away from the curved surface introducing mixed mode fracture which becomes predominantly Mode II along the flat stent strut. Furthermore, Mode II debonding initiation is computed for a hinge with a large radius of curvature in our case study. The conditions that lead to the development of Mode II initiation in a symmetric hinge are currently being investigated in a parallel study. Additionally, mixed mode delamination is also computed in compressive regions of the stent surface, leading to buckling of the coating. Compressive buckling is computed to initiate at a higher deployment level than that that computed for tensile delamination. This may be related to the fact that during stent deployment compressive surfaces of the plastic hinge have a lower curvature (κ) than tensile surfaces, and consequently a higher dimensionless interface strength. However, initiation of buckling does not appear to be governed by our ‘design curve’ for initiation of tensile delamination. The mechanism of compressive delamination is quite different to that computed in tensile regions and further investigation of coating buckling in compressive regions of the plastic hinge is ongoing. The experimental study of Regar et al.24 demonstrates identical patterns of coating buckling in compressive regions as computed by our simulations.
Patterns of coating delamination in the tensile region of the plastic hinge computed in this study are very similar to observed SEM images of coating debonding in the studies of Levy et al.,14 Ormiston et al.,20 Otsuka et al.,14,20,21,24 and Regar et al.24 In this study, an SEM image of tensile delamination in a Cypher stent from Otsuka et al.21 was used to determine the ratio of interface strength to coating stiffness (σmax/E) for this commercially available drug eluting stent. The model was calibrated using reported values of coating thickness,22 and deployment and delamination magnitudes estimated from the SEM image.21 It should be noted that, while no information was available on the coating material or interface strength for this stent design, our methodology allows for the computation of the required (σmax/E), assuming that the coating material exhibits linear behavior in the range of deformation experienced during deployment.
Based on our calibration of the (σmax/E) for reported coating delamination in a Cypher® stent,21 an interface strength in the range ~0.005–0.05 MPa should be expected if coating stiffness lies between that of SIBS (~2 MPa) and Chronoflex (~23 MPa). This range of interface strengths is very similar to a range of ~0.007–0.07 MPa estimated from the investigation of polymer peel tests by Rahulkumar et al.23 The full range of interface strengths used in our parametric studies (0.009–1 MPa) also complies closely with a range between 0.14 and 0.4 MPa estimated from the work of Yan et al.29 for a polymer peel test on a steel substrate. It can therefore be concluded that all results in this study are computed for realistic ranges of interface strengths and for commercially relevant polymer coating materials.3,4,31 We are currently performing a series of mixed mode experimental tests to determine interface properties for polymer coatings on stainless steel substrates, allowing for the precise measurement of τmax and σmax.
One shortcoming of our calibration of coating delamination in stent Design 1 is that we have not included pre-stress in the coating. The experimental image of Otsuka et al.21 would suggest that the coating may contain some pre-stress, as the debonded material appears to be stretched across the plastic hinge. The inclusion of such pre-stress would only serve to accelerate debonding in our models, thus strengthening the motivation for analysis of coating delamination during stent deployment. The effect of pre-stress will be investigated in future simulations.
Results for the multiple hinge Taxus stent geometry (Design 2) reveal very similar patterns of tensile delamination to those computed for a Cypher stent geometry (Design 1). Both designs exhibit a close compliance to our ‘design curve’ with similar debonding initiation being computed for both designs for a given coating material and interface strength. The smallest hinge (B) in Design 2 exhibits the earliest tensile debonding during deployment, whereas the largest hinge (C) exhibits the highest deployment at the onset of tensile debonding, followed closely by hinge A in Cypher stent (Design 1). These results are not unexpected when one considers the non-dimensional geometric parameter (tκ) for each hinge in the two designs. Results suggest that the use of a compliant SIBS coating will result in least debonding for a given interface strength. Regarding the buckling of coatings in compressive regions of the plastic hinge, the earliest buckling is computed in the vicinity of a geometric discontinuity of the Cypher stent. It might therefore be reasonable to suggest that application of a coating to regions of high curvature or geometric discontinuities should be avoided.
It should be noted by the reader that this case study should not be treated as a direct comparison of two commercially available drug eluting stents, as precise information regarding the coating and interface properties are unavailable for both designs. The purpose of our case study is to demonstrate the ability of our modeling framework to predict locations of coating debonding and deployment conditions required to produce debonding for commercially available stent geometries.
In conclusion, our modeling framework allows for the prediction of coating delamination in tensile regions of plastic hinges during stent deployment. Delamination initiation is governed by coating thickness and stiffness, interface strength, and hinge curvature. Simulations also predict coating buckling in compressive regions of the plastic hinge. Such patterns of coating debonding have been reported in experimental studies of commercially available polymer-coated stents, thus providing strong motivation for the predictive modeling framework presented in this study.
C.H. was supported by an Irish Research Council for Science, Engineering and Technology Postgraduate Scholarship.