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Longitudinal support deformation: importance of support type and support position

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  • Save American Journal of Biomedical En gineer in g 2013, 3(3): 63-69 DOI: 10.5923/j.ajbe.20130303.03 Longitudinal Stent Deformation: Importance of Stent Type and Stent Apposition Mark Pitney1,*, Elizabeth van Niekerk1, Socrates Dokos2, Matthew Pelletier3, William R. Walsh3 1Eastern Heart Clinic, Prince of Wales Hospital, Sydney, 2031, Australia 2Graduate School of Biomedical Engineering, University of NSW, Sydney, 2052, Australia 3Surgical and Orthopaedic Research Laboratories, Prince of Wales Clinical School, University of NSW, Prince of Wales Hospital, Sydney, 2031, Australia Abstract Informat ion on stent longitudinal strength is limited to benchtop models using unconstrained stents, consequently having uncertain clinica l significance. This study investigated the effects of stent apposition and tube (artery) compliance on stent stability. Multilin k Vision, Multilink 8, Pro mus Element, Driver and Integrity were examined. Stent size was 3/18 mm or closest length. Four scenarios were tested: in three the stents were well deployed and in the final test, stents were purposely severely malapposed. In the well-deployed models, three tubes with different co mpliance were tested. Stents were co mpressed by applying a single point force. Malapposed stents compressed extensively at 50 and 100 g m., especially Element, wh ich deformed double the amount of the other stents. At higher forces (150 g m.), the magnitude of difference was reduced. Stent apposition had a profound protective effect and limited differences between stents. Stiffer tubes provided more protection to stent deformation. A ll stents are vulnerable to deformation when severely malapposed but the Element is significantly more so. Well-apposed stents are considerably mo re resistant to longitudinal compression such that differences between stents become minimal. More rigid tubes offered greater protection. Clinical implications regarding stent sizing and postdilation strategy are significant. Keywords Coronary Stents, Longitudinal Co mpression, Apposition, Vessel Co mp liance 1. Introduction To maximize the efficacy and deliverability of coronary stents, their design has traditionally focused on features such as radial strength, recoil, shortening, deliverability, strut thickness and resistance to stent thrombosis. The concept of stent longitudinal strength is relatively new, and though rare, longitudinal stent compression may have devastating effects (1). When stents are init ially deployed, the lu men often remains narrowed at the site of severest stenosis while elsewhere, the stent may not be fully apposed to the vessel wall. Th is increases the feared co mplication o f stent thrombosis and consequently, it is common fo r operators to pass a new balloon through a recently deployed stent to further expand it. Clin ical experience with and consequences of longitudinal deformation have been reviewed elsewhere (2, 3). Recently, case reports have been published of stents that severely deformed longitudinally when operators attempted to pass new balloons through the stents (1, 4, 5). The likely * Corresponding author: (Mark Pitney) Published online at Copyright © 2013 Scientific & Academic Publishing. All Rights Reserved cause was that the balloon catheter tips snagged and exerted a longitudinal force on the stent. Certain stent designs may be mo re vulnerable than others. Although the original clin ical report involved stent testing within tubes (5), recent studies have comprised benchtop models using completely unconstrained stents (6, 7). While elegant and establishing that stents have intrinsic differences, the clinical applicability of these data is arguable. The magnitude of this effect has not been previously studied. This study attempted to mimic clinically relevant scenarios by testing well deployed stents in tubes of various compliances and comparing these to a purposely ma lapposed model using undersized stents. It was expected that the results would be of clinical interest to all operators, especially those treating increasingly co mp lex lesions, where this phenomenon is more likely (5). Arterial co mpliance has the potential to greatly influence the apposition, stability and resistance to deformation of the stents placed within them. This study also attempted to quantify this influence. 2. Materials and Methods The 5 stent platforms most frequently used in our institute over the last 3 years were tested (Table 1). As stent platforms have different names around the world, stents are referred to as listed in Table 1 hereafter. When availab le, the bare metal 64 M ark Pitney et al.: Longitudinal Stent Deformation: Importance of Stent Type and Stent Apposition versions were tested to reduce costs. Stent compression was tested by deploying stents in silastic tubes as detailed below. A single point of fo rce model was emp loyed to mimic the most common clinical scenarios of a postdilation balloon or IVUS catheter catching on to a stent (8). In v ivo, stents can be co mpressed by an anterograde force (e.g. post dilation balloon) or retrograde force (e.g. withdrawal of an IVUS catheter or filter). Th is study used a retrograde force by deploying stents with a 3-0 monofilament nylon 45cm long suture (Ethicon, J&J, New Jersey, U.S.A.) between the stent and the tube. A free end of the suture was brought through the lumen of the stent such that a loop was formed, which subsequently applied force to a single point on the most distal strut. This model was found to be highly reliable in the design phase as opposed to hooks or rods. The connecting suture was attached to the load cell on a MACH 1 M icro mechanical Testing Machine (BIOSYNTECH, Laval, Quebec, Canada). The load cell was configured such that it would move upwards at a constant velocity of 0.1mm/s. Load (grams, gm.) and time (milliseconds) data were recorded at 20 Hz. In the first case, the intrinsic strength of the stents was tested using deployment in an oversized tube such that severe malapposition occurred. The 3.0 mm stents were deployed just under the Rated Burst Pressure of the balloons in an oversized 4.0mm lu men diameter tube. One end of the stent was fixed to prevent mig ration wh ile force was applied to the other end of the stent. Table 1. Tested Stents Pl atfo rm Name in thi s stu dy Manufacture r Size Amount (n = 12) Multilink 8 / Xience Prime ML 8 Abbott Vascular 3x18 mm 3 x 4 Multilink Vision / Xience V ML Vision Abbott Vascular 3x18 mm 3 x 4 Driver / Endeavor Sprint / Endeavor Resolute Driver Medtronic 3x18 mm 3 x 4 Integrity / Integrity Resolute Int egrit y Medtronic 3x18 mm 3 x 4 Omega / Promus Element /T axus Element / Ion Element Boston Scientific 3x16mm 3 x 4 Table 2. Models Tested Te st Name Malapposed Stent Diame ter 3.0 mm Tube diameter /wall thickness 4.0/1.0 mm Pos tdil ati on Nil Apposed Compliant T ube Apposed Normal T ube 3.0 mm 3.0 mm 3.0/0.4mm 3.0/1.0 mm 3.25 @ 18 atm 3.25 @ 18 atm Apposed Rigid T ube 3.0 mm 3.0/1.5mm 3.25 @ 18 atm To model the three other scenarios, stent apposition was ensured by deploying 3.0mm stents in a 3.0mm tube at nominal pressure but then postdilating at 18 at m with a 3.25mm non-co mpliant balloon (3.25mm NC Hiryu, Teru mo, Tokyo, Japan) to standardize apposition. At 18 at m, the postdilation balloon was approximately 0.4 mm larger than the tube lumen. Each stent type was tested three times for each of the four apposition scenarios; severely malapposed, and three scenarios in which the stents were apposed to tubes of three different comp liances. Figure 1 shows an examp le of deformation of well-apposed and severely malapposed stents. Table 2 summarizes the testing methodology. Figure 1. Examples of well-apposed stents prior to (A1) and after (A2) deformation, and severely malapposed stents prior to (B1) and after (B2) deformation. The needle in the B-panels acted as an obstruction to deformation. All stents showed a difference between the apposed version of the stent (panels A1 and A2) and the severely malapposed version of the stent (B1 and B2) American Journal of Biomedical En gineer in g 2013, 3(3): 63-69 65 Silicone tubes of 3mm lu men diameter (GeckoOptical, Joondalup, WA, Australia) were selected to represent a range of possible coronary arterial co mp liances. Concurrently with stent testing, it was confirmed that the tubes had differences in co mpliances by using IVUS and a haemodynamic monitoring systemand measuring tube lu men fro m 80 to 200 mmHg, increasing pressure in increments of 20 mmHg. The 0.4mm thick tube had a d iametric co mp liance of 12.5%/100 mmHg; the 1.0mm th ick tube had compliance of 5.3%/100 mmHg and the 1.5mm thick tube had a compliance of 1.05%/ 100 mmHg . This range in co mpliance attempted to reflect normal, highly co mpliant and stiff ateries respectively. Tubes were compared to each other by measuring compression of all stents at 50, 100 and 150 g m. and calculating a “global average” quantity for each tube. Seria l photographs were taken at 5 second intervals with a Canon EOS 500D with a macro lens (Canon Inc., Ohta-ku, Tokyo, Japan). Image analysis was performed using ImageJ (Rasband, W.S., ImageJ, U. S. National Institutes of Health, Maryland, U.S.A.), to visually determine changes in length. The linear d istance between the upper end of the stent and the point where force was applied was measured by one observer for all image stacks. Care was taken to ensure repeatability by measuring a distance parallel to the sides of the stent. Subsequently, stent compression was calculated as the change in length divided by the original length, and was expressed as a percentage. Image resolution was approximately 25 p ixels per mm. Changes in stent length were co mbined with t ime and force data as produced by the MACH 1 M icro mechanical Testing Machine. As this was a single point of force experiment, the stents did not deform uniformly (as was the case with previous bench top studies which used rim co mpression), but instead, the distance of maximu m deformation was noted. Linear interpolation was used to determine quantities of force and compression as required. Subsequent to testing, stent lumens were photographed with a Dyonics Xenon XL light source and cable receptacle (Smith and Nephew, London, UK) to determine the extent of lu men obstruction after stents were co mp res s ed . One-way ANOVA was performed with M initab® Statistical Software i (M initab Inc., State College, Pennsylvania, U.S.A.) to determine the effects of apposition, tube compliance and stent design, and p < 0.05 was considered significant. One e xpe rimental outlie r was deleted for apposed ML8s at 150 g m. 3. Results 3.1. General Observati ons Well apposed stents displayed characteristic deformat ion of the stent ‘rim’ bending around into the lu men. Malapposed stents displayed significantly more deformation, with examp les of well-apposed and malapposed stents shown in Figure 2. Apposition provided significant protection and tube compliance was an important variable. 3.2. Malapposed Stent Testing This test assessed the intrinsic strength of malapposed stents. All stents suffered considerable deformat ion (Figure 2). At lower forces (50 and 100 g m.), the Element design was significantly more deformab le whereas the remaining stents all behaved similarly. Co mp ression at 50 g m. was 46.7±8.8% (mean ± SD, Element) vs. 18.6±2.8% (mean ± SD of other 4 stents, p = 0.00292) and at 100 g m. was 74.6±1.4% vs. 34.8±6.8% (p = 0.032). At 150 g m., however, the differences between the stents were considerably reduced. Figure 2. Compression for malapposed (left) vs. well-apposed stents in tubes of 1.0mm thickness (right) shown to same scale. Note the profoundprotective effect of apposit ion such that any well deployed stent is much stronger than any poorly deployed stent regardless of st ent type or compressive force 66 M ark Pitney et al.: Longitudinal Stent Deformation: Importance of Stent Type and Stent Apposition 0.4mm thick tube 60 1.0 mm thick tube 50 1.5mm thick tube 40 Severely malapposed 30 Compression (%) 20 10 0 ML 8 ML Vision Element Integrity Driver Figure 3. Compression at 50gm. force applied. Severely malapposed stents are compared with stents apposed to vessel walls of different thicknesses (compliances). Error bars are st andard error 3.3. Apposed Stent Testing Apposition had a dramatic effect on stent stability and longitudinal strength. All well apposed stents deformed by less than 5% at an applied load of 50 g m. regard less of the co mpliance of the tube in wh ich they were deployed (Figure 2). Element (most vulnerable in unconstrained tests), showed an approximate tenfold improvement in longitudinal strength when apposed such that there was considerable equalising in its performance relative to other stents. A comparison of co mpression for different designs at 50 g m. load applied is displayed in Figure 3, showing the dependence of stent compression on stent type (particula rly for ma lapposed stents), as well as vessel compliance in the three apposed models. In the most compliant tube, compression for Element continued to be greater than other stents but only at higher loads. At 150 g m., Element co mpressed by 35.4±10.9%, vs. 23.4±5.3%, for other stents (p =0.16). In the 1.0 and 1.5mm thick tube models, there was no statistical difference between stents at 50 g m. (p = 0.364 in 1.0mm thick tube, p = 0.973 in 1.5mm th ick tube). At h igher fo rces, stents behaved differently but absolute differences were small (2-4% only at 100g m regard less of tube) and there was no stent with clear superior or in ferior performance to other stents (p > 0.09 for all stents in 1.0 or 1.5mm thick tubes at 100 or 150g m.) 3.4. Tube Compli ance Global average stent compression (as defined in methods) was 4% for the tube with usual compliance (1.0 mm wall thickness). Increasing the compliance of the tube (e.g. reduction in stiffness) resulted in greater vulnerability with the 0.4mm wall thickness tube having an average stent compression of 12% (p < 0.0001). Ho wever, making the tube stiffer than normal (less comp liant, wall thickness 1.5mm) offered no additional protection with average stent compression of 5% (p = 0.218). Results are displayed in Figure 4. 16 14 12 10 8 6 4 2 0 0.4mm thick tube 1.0 mm thick tube 1.5mm thick tube Figure 4. Compression averaged over all stents and all loads for well-apposed stents in tubes of different thickness (compliance). Error bars are standard error Average compression (%) American Journal of Biomedical En gineer in g 2013, 3(3): 63-69 67 4. Discussion This study confirmed that intrinsic differences exist between stents with the Element being approximately twice as vulnerable to deformat ion as other stents when ma lapposed. It is noteworthy that all stents are vulnerable to stent deformation when malapposed if sufficient force is applied. A lthough it is unlikely that stents would be entirely malapposed in vivo, long stents sized to the distal vessel in arteries that taper are likely to be poorly apposed. Additionally, somet imes when GTN is ad ministered post stent insertion it becomes apparent that the stent is undersized. Operators should recognise these situations as being of higher risk for deformat ion and should a balloon not pass easily, operators are reco mmended to assume malapposition and change to a lower profile semi co mpliant balloon and ensure apposition at moderate pressure then return with a high pressure non compliant balloon if required. The intrinsic differences between stent types became less important once stents were well deployed as the associated embedding and frictional fo rces dominated stent intrinsic strength in this study. A limitation to this assumption is that this study used silastic tubes which may have different frictional and embedding properties to diseased human arteries. Nevertheless, apposition is clearly protective; although the degree of protection may be over or understated by this study. If vulnerability to deformat ion affects an operator’s stent choice (other factors such as deliverability and conformability may be equally important), then there should be equal focus on ensuring apposition. In terms of tested applied load, the region of interest has been extended fro m the previously deemed ‘clin ically relevant force’ o f 50 gm. to include 100 g m. and 150 g m. Th is was done both to evaluate stents under circu mstances with mo re extreme loads, and to include a marg in of error around the applied load previously estimated at 50 g m., the clinical relevance of which is potentially limited by subjectivity. The potential effect of vessel comp liance was a secondary aim of this study, and was tested by us ing 3 s ilicone tubes of varying compliance (1-12%/ 100mmH). It was hypothesised that stiffer tubes would allow more embedding of stents and thus offer g reater protection. Generally speaking, the results support this hypothesis insomuch as the most compliant tube offered the least protection. However, there was litt le difference between the 1.0 and 1.5mm th ick tubes, suggesting that the effect of arterial co mp liance may plateau. Arteries with coronary artery disease are usually less compliant than healthy arteries. There are many factors affecting arterial co mpliance, including age and sex, applied pressure, diabetes, obesity, plaque type, area and eccentricity, type of coronary artery disease (stable/unstable), and others (9-13). Published studies report compliance for normal human coronary arteries to be between 6 and 9%/100mmHg (6.2%, (14), 8·0% (15), 8.85% (16)). The co mpliances of our tubes thus reflect normal arterial co mpliance, lower than normal and greater than normal. A reasonable assumption fro m this study is that arteries with higher than normal compliance (e.g. soft plaque rupture in young arteries, perhaps myocardial infarction in arteries with positive remodelling) may be more vulnerable to longitudinal stent deformation and may need to be postdilated more carefully. Stent compression was not unifo rm around the ‘rim’ of the stent, because a single point of force was applied. The maximal co mp ression was measured, as only the shortest length of stent visible would provide radial support to an artery. These observations suggest that significant stent deformation may occur before it is fluoroscopically evident. Moreover, the amount of compression likely to be determined fluoroscopically is deemed minimal co mpression, since the longest length of stent visible would be measured, which may even be larger than the orig inal stent length (Figure 5). Figure 5. Eccentric deformation characteristic of compression under a single point of force. There is a clear difference between the maximal and minimal compression Co mparison of this study to previous bench data is complicated. The severely malapposed stent compression results in this study were principally gathered to contrast against the study's apposed models’ results. Outcomes for malapposed stents do not closely mirro r prev iously published results, but it should be noted that the testing methodologies differ. Rather than compressing the stent by pressing on the whole rim of the stent, we applied a single point of force wh ich, we believe, better replicates catheters pushing on a single strut or IVUS catheter catching. Another difference was that stents were supported along the entire bottom strut in previous benchtop s tudies , compared to only two points in the severe malapposition model, where the stent encountered the needle. This prevented it fro m the migrat ion that would otherwise follow for severely malapposed stents. Apposed stents did not require a needle to prevent migrat ion since they were already firmly in contact with the tubes. The issue of whether to express the compression of a stent 68 M ark Pitney et al.: Longitudinal Stent Deformation: Importance of Stent Type and Stent Apposition as a percentage of length or as an absolute displacement (mm), was briefly examined. Different lengths (18mm, 23mm, and 28mm) of M L Vis ion stents, all with the same nominal d iameter of 3mm, were tested. It was found that for the same applied load, longer stents deformed more at the same applied load. Therefore, expressing deformation as a percentage of length gave more reproducible information and thus was employed. This concept of longer stents deforming mo re was not a specific aim of this study and would be an interesting future study with clinical imp licatio n s . 5. Limitations When selecting a stent, in addition to established important features (deliverability, radial strength, visibility and conformability) operators may also need to consider stent intrinsic strength and risk of init ial malapposition. Importantly, appropriate postdilation or other strategies ensuring correct apposition will counteract longitudinal stent deformation. In future, stent longitudinal deformat ion in tortuous vessels may be investigated, since this clinical situation presents an increased opportunity for catheters or other devices snagging on stents. Further elaborations on this work may include modeling the effect of varying other arterial variables such as friction or hardness on stent longitudinal deformation. Sample size was small at 3 stents of each design per apposition group. However, in this study, a total of 60 stents were required in the final testing phase, and a further 10 were used in the init ial design phase. Similarly, one could argue that an important addition to this study would be to apply force to the middle section of a stent and to consider testing in a curved tube in addition to a straight one. This is an area for future work as stent usage becomes a rate limiting factor; especially when all possible investigational variables are included in the study design. For repeatability, arterial co mpliance, surface p roperties, diameter and thickness were assumed uniform, and the material d ifferences between arterial layers (17) have been ignored. The use of a silastic tube is a further simp lification. Silastic tubes may have different frictional and embedding properties to diseased arteries which could change the degree of deformat ion. Nevertheless, the basic premise that apposition is protective and differing co mpliance can alter that protection is reasonably investigated by these testing meth o d s . 6. Conclusions The results of this study have significant clin ical imp licatio n s . All stents are deformable if enough force is applied, particularly when unsupported (severe malapposition). Under these circumstances, the Element is approximately twice as deformab le as the other tested stents. If comp ressed by a single point of force (e.g. post dilation balloon) considerable deformation can occur before it becomes fluoroscopically evident. Apposition makes a dramatic difference and became the ma jor determinant of longitudinal strength with stents being up to 10 t imes more resistant to deformat ion when well apposed. A well apposed Element stent became more resist to deformat ion than any poorly apposed stents, regardless of its intrinsic strength. Arterial co mpliance has an effect, with mo re co mpliant arteries providing less resistive force to counter longitudinal compression. Stents deployed in very compliant arteries may be more vulnerable to this phenomenon. ACKNOWLEDGEMENTS The authors would like to thank the contributions of all the manufacturers in terms of the provision of stents, cooperation and feedback. REFERENCES [1] Williams PD, M amas M A, M organ KP, El-Omar M , Clarke B, Bainbridge A, et al. Longitudinal stent deformation: a retrospective analysis of frequency and mechanisms. EuroIntervention. 2011. Epub Nov. 4. [2] Abdel-Wahab M , Sulimov DS, Kassner G, Geist V, Toelg R, Richardt G. Longitudinal Deformation of Contemporary Coronary Stents: An Integrated Analysis of Clinical Experience and Observations from the Bench. Journal of Interventional Cardiology. 2012;25(6):576-85. [3] Janakiraman E, Subban V, Victor SM , M ullasari AS. Longitudinal deformation – Price we pay for better deliverability of coronary stent platforms. Indian Heart Journal. 2012;64(5):518-20. [4] Hanratty CG, Walsh SJ. Longitudinal compression: a "new" complication with modern coronary stent platforms - time to think beyond deliverability? EuroIntervention. 2011;7(7):872-7. [5] Pitney M, Pitney K, Jepson N, Friedman D, Nguyen-Dang T, M atthews J, et al. M ajor stent deformation / pseudofracture of 7 Crown Endeavor / M icro Driver stent platform: incidence and causative factors. EuroIntervention. 2011;7(2):256-62. [6] Ormiston JA, Webber B, Webster M WI. Stent Longitudinal Integrity: Bench Insights Into a Clinical Problem. JACC. 2011;4:1310-7. [7] Prabhu S, Schikorr T, M ahmoud T, Jacobs J, Potgieter A, Simonton C. Engineering assessment of the longitudinal compression behaviour of contemporary coronary stents. EuroIntervention. 2011. Epub November 2011. [8] M amas MA, Williams PD. Longitudinal stent deformation: insights on mechanisms, treatments and outcomes from the Food and Drug Administration M anufacturer and User Facility Device Experience database. EuroIntervention. American Journal of Biomedical En gineer in g 2013, 3(3): 63-69 69 2012;7:1-9. Epub M arch 2012. 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[14] van Andel CJ, Pistecky PV, Borst C. M echanical properties of porcine and human arteries: implications for coronary anastomotic connectors. Ann Thorac Surg. 2003;76(1):58-64. [15] Tai NR, Salacinski HJ, Edwards A, Hamilton G, Seifalian AM . Compliance properties of conduits used in vascular reconstruction. Br J Surg. 2000;87(11):1516-24. [16] Numao T, Ogawa K, Fujinuma H, Furuya N.[Pulsatile diameter change of coronary artery lumen estimated by intravascular ultrasound]. Journal of cardiology. 1997;30(1):1-8. Epub 1997/07/01. [17] Azarnoush H, Lamouche G, Bisaillon C-E, Degrandpre C, Boulet B. Evaluation of a Technique to Estimate the Compliance of Atherosclerotic Intima In: Bello F, Cotin S, editors.: Springer Berlin / Heidelberg; 2010. p. 20-8. i MINITAB® and all other trademarks and logos for the Company's products and services are the ex clusive property o f Minitab Inc. All other marks referenced remain the property of their respective owners.

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