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Protective Layer Development for Enhancing Stability and Drug Delivery Capabilities of DES Surface Crystallized Coatings Shady Farah ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b18733 • Publication Date (Web): 13 Feb 2018 Downloaded from http://pubs.acs.org on February 16, 2018

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ACS Applied Materials & Interfaces

Protective Layer Development for Enhancing Stability and Drug Delivery Capabilities of DES Surface Crystallized Coatings Shady Faraha,b a

Institute of Drug Research, School of Pharmacy-Faculty of Medicine, Center for Nanoscience and Nanotechnology and The Alex Grass Center for Drug Design and Synthesis, The Hebrew University of Jerusalem 91120, Israel. b Daniel G. Anderson/Robert S. Langer Labs, Department of Chemical Engineering and The David H. Koch Institute for Integrative Cancer Research, Massachusetts Institute of Technology, 500 Main Street, Cambridge, Massachusetts 02139, USA. Corresponding author, E-mail: [email protected] , Tel.: +1 857-4159686.

ABSTRACT Carrier-free drug eluting stents (DES) based crystalline coatings are gaining prominence due to their function, skipping many limitations and clinical complications of the currently marketed DES. However, their usage has been humbled by crystalline coating inflexibility, limited mechanical and physical properties. This study reports for the first time a protective top coating development for enhancing crystalline coating merits and delivery capabilities. Flexible and water soluble polysaccharide top coating was developed and applied onto Rapamycin (RM) crystalline carpet. Top coating prevented crystalline coating delamination during stents crimping and expansion without affecting its release profile. Crystalline coating strata and its interfaces with the metallic substrate and top coating were fully studied and characterized. The crystalline top coated showed significant physical, mechanical and chemical stability enhancement with ~2% RM degradation after 1 year under different storage conditions. Biocompatibility study of the top coated stents implanted subcutaneously for 1 month into SD rats, did not provoke any safety concerns. Incorporating RM into the top coating to develop a bioactive protective coating for multilayer release purposes was also investigated. The developed protective coating had wide applicability and may be further implemented for various drugs and implantable medical devices.

KEYWORDS: Bioactive Surface Coating; Drug eluting stents; Sequential Release; Rapamycin; Crystallization; Metallic surfaces; Controlled multilayer-release; Hyaluronic acid.

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1. INTRODUCTION Local delivery of drugs using DES provides both biological and mechanical solutions in the management of in-stent restenosis.1 However, DES manufacturing methodologies are based mainly on mechanical processes such as spray and dip coating that tend to generate coatings with poor stability and fast drug release with some long-term safety concerns.2 In marketed DES, a coated therapeutic agent is applied together with a polymer carrier responsible for several functionalities. The carrier holds the drug mechanically, preserves chemical stability of drug, and most importantly regulates the release kinetics of the drug.3-8 Recently, FDA and numerous studies reported safety and clinical issues, i.e., late state thrombosis with regard to available DES. Thus, various developments and clinical studies are ongoing to improve DES performance.5,9-22 Coating durability is one of the stent’s most important characteristics. The stent is expected to go through all sorts of mechanical constrains including crimping, handling deployment in blood vessel and inflation. However, further findings demonstrated DES coating defects, including: cracking, flaking and delaminating in commercially available stents paw path for further investigations.18,23 First, DES must be capable of being stretched without flaking or delaminating.24 Second, the polymer must be able to deliver the drug at a sustained, controlled and predictable rate.25 Surface modification also alters the properties that are exposed to tissues.2 Thus, key requirements for polymers used in the DES matrix carrier systems include suitable physical properties, stability, compatibility with drugs, biocompatibility with vascular tissue, and control of drug release.26,27 It is difficult to have all desired properties together in one polymer, and thus mixtures of polymers are used, but each has its limitations. Research is currently centered on the development and evaluation of new improved DES that maintain impressive clinical benefits while eradicating long-term safety concerns. New technologies developed or under development primarily include use of polymer or carrier-free DES and the use of biodegradable polymers or biodegradable stents.27 Polymer-free stents have a potential long term benefit over traditional polymeric coated DES. Examples are BioFreedom stents that exhibit reduced neointimal proliferation compared to polymer coated Cypher stents28 and the novel VESTAsync-eluting stent effective in reducing lumen loss and neointimal hyperplasia.29 There are ample data showing increased efficacy of the newest stents in interventional cardiology, but more developments are still required, i.e., limited mechanical stability.30,31 We previously reported the development of novel carrier free-DES based crystalline coatings of RM onto a stent's struts surface with a therapeutic dose of ~100µ/per stent, and their in vitro/in vivo behavior.32-34 Here, I have studied the thermal, physical and mechanical stability of the crystalline coating where metallic substrate-crystalline interface, fully studied and characterized. Then, I developed a quickly dissolvable temporary coating that absorbs the mechanical constrains through stent life until deployment, thereafter rapidly dissolves without interfering with the drug release profile. Lastly, I examined the potential of bioactive protective top coating fabrication by incorporation into RM for initial burst release followed by slow profile attributed to crystalline coating, using RM as a role model.

2. EXPERIMENTAL SECTION 2.1. Materials Rapamycin (RM), stents (CoCr smooth and rough stents, 15 mm length, batch M/Z: 450110101203, 4120200275-60) and cylinders tubes (CoCr, 5 cm), were obtained from Alvimedica, Turkey. Macrolane VRF20 (Hyaluronic acid (HA) aqueous solution) was obtained from from Q-Med AB, Sweden. n-Hexane (AR), ethyl acetate (AR), glycerol and hydrochloric acid were purchased from Biolab Ltd. Sodium azide was purchased from MERCK, Darmstadt F. R. Germany. Sodium hydrogen phosphate, sodium phosphate monobasic, sodium carbonate (AR) and methanol (HPLC grade) were purchased from J. T. Baker, Holland. A shrinkable tube (PVLF05J) was obtained from Shrink Sleeve Ltd, UK. Sodium Lauryl Sulfate (SDS), Carboxymethylcellulose (CMC), Alginate, Chitosan, Calcium Chloride, Triethyl citrate and Polyethylene Glycol (PEG, Mw: 1,000-15,000 g/mol) were obtained from Sigma-Aldrich.


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2.2. Methods 2.2.1. Stent/tubes RM surface crystallization or amorphous coating Stents with crystalline coatings of ~100µg/stent were prepared as previously reported.33,34 Shortly, the process was comprised of two steps, seeding and crystallization. For seeding, 50 mg of RM was ground for 3 min, then 1.6 mg of ground RM was transferred into a 5 ml glass vials. 4 ml of nHexane was added to these vials and was sonicated until RM homogeneously dispersed in hexane. Then, stents were mounted onto shrinkable tube placed on needles and located at the center of the vial. These vials were then placed in the ultrasonic bath for 10 min at 30 °C to form seeding layer. Stents were gently removed from the vial dried to yield a thin seeding layer of 15 ± 5 µg/stent. These seeded stents were analyzed by SEM to examine RM seeding nuclei size and distribution; these nuclei serve as starting nucleation points in the next crystallization process.33 Crystallization lunched with 50 mg of RM dissolved in 3 ml of ethyl acetate and 65 ml n-Hexane mixture. Seeded stents were placed in this solution for 5 min at 25 °C to form continues crystalline carpet. They were then dried overnight. Amorphous coatings: were prepared by spraying a drug solution using an ultrasonic spray-coating machine (Medical, Sono-Tek, USA) as previously reported.32,33 An RM solution in ethyl acetate or ethanol (1% w/v) was sprayed at a flow rate of 0.2 mL/min with ultrasonic generator power at 1.2W. Spray-coated samples were examined visually for their homogeneity and then weighed using a microbalance. 2.2.2. Top coating formulations Two coating methods were applied, dipping and spraying. A 0.3% hyaluronic acid (HA) solution was used to make a top coating on crystallized RM, prepared by dissolving 1 ml of Macrolane VRF20 (20 mg/ml HA) in 5.7 ml of DDW and then mixing vigorously for 3 days. Using Sono-Tek system: 0.1 ml solution was sprayed at flow 0.2 ml/min with 10 watts power. Sprayed stents were dried overnight in an active hood. Other polymers (CMC, alginate and chitosan) were applied by a dipping method only. Top coatings releasing RM: were prepared by diluting RM-methanol solution with HA aqueous solution dropwise, followed by vigorous vortex. RM concentration was optimized for 10 mg/ml, and to this solution a HA-F formulation (0.3% HA+10% w/w glycerol) was added to fit a final top coating ratio:drug 1:1, and mixed vigorously. Solution was sprayed using aforementioned system and parameters. 2.2.3. Therapeutic coatings analysis 2.2.3.1. Crystalline coating Coatings were evaluated microscopically for their homogeneity and then weighted using a microbalance. Thermal analysis was determined on a Metler TA 4000-DSC differential scanning calorimeter (DSC) calibrated with zinc and indium standards at a heating rate of 10 °C/min, and on a Stuart Scientific Melting point SMP1 heater (+20-240°C). The crystalline structure was proven by X-ray powder diffraction. Measurements were performed on a D8 advance diffractometer (Bruker, AXS, Germany) with a goniometer radius of 217.5 nm, Gobel mirror parallel beam optics 2° Sollers slits and 0.2 mm receiving slit. The powder samples were placed on low background quartz sample holders. XRD patterns from 20° to 60° 2θ were recorded at room temperature using a CuKα radiation (λ=0.15418nm) with the following measurement conditions: tube voltage of 40 kV, tube current of 40 mA, step mode with size of 0.02° 2θ and a counting time of 1 sec per step. Instrumental broadening was determined using an LaB6 powder (NIST SRM 660). Amorphous coatings were prepared and immediately analyzed to avoid any possibility of selfcrystallization of RM. 20 mg of RM was dissolved in a minimal volume of solvent, 250µl of absolute ethanol or 200µl ethyl acetate (Ar) or 150µl acetone (Ar) followed by filtration with PTFE filters 0.2µm and fast solvent evaporation by N2 flow till complete dryness and analyzed.



2.2.3.2. Top coating 2.2.3.2.1. Profilometry Therapeutic coatings thickness after the top coating was determined by a P-15 profilometer (KLATencor Co., San Jose, CA). Profiles were recorded across a notch in the coating, which was manually scratched by a wooden stick. Also, the roughness of RM seeding and crystalline layers was determined similarly. Samples analysis was performed on cylinders tubes prepared as similar to coated stent. 2.2.3.2.2. Atomic force microscopy (AFM) Stents samples were affixed on glass slides by using melted PEG (Mw 20,000 g/mol). Sample topography was scanned with a Dimension 3100 scanning probe microscope with a Nanoscope-V controller (Bruker, Germany) for these fixed stents by using a Tapping Mode (TM). Stent surface topography was obtained by lightly tapping the surface with an oscillating Si probe (RTESP probe, Bruker, Germany) at f = 300 kHz. Topography changes were studied for bare metal stents, and following seeding, crystallization and HA top coatings (5µm*5µm). Three different points on 3 stents were analyzed. 2.2.3.2.3. Scanning electron microscopy (SEM) Stent samples were placed on conductive carbon paper, coated with gold to a thickness of ~10 nm using a sputtering deposition machine (Polarone E5100). Afterwards, they were imaged using SEM (FEI E-SEM Quanta 2000) at an acceleration voltage of 30 KV. 2.2.4. Stent crimping and inflation Stents with 1.5 mm diameter were mounted on a balloon-expandable catheter system (Bxsonic) and inflated with air. The stents attained a diameter of 3 mm and a final pressure of 10 bar by using a 200 cc inflation device from Biometrics. Samples with or without top coating after inflation were microscopically examined and HPLC quantified. Hot (4.5N) and Cold (6.0N) crimping were also applied on RM crystalline coating and studied. 2.2.5. In vitro RM release from stents in PBS Release study was carried out in phosphate buffered saline (PBS, pH 7.4) with 2 ml medium (with 0.02% of sodium azide) at 37°C. Sampling was carried out by replacement of 1.5 ml of release medium with fresh medium at 6h, 1, 3, 5, 7 days, and then weekly. RM concentration in samples was measured by reverse phase HPLC on C-18 column with a mobile phase consisting of watermethanol (10:90 %v/v). An isocratic mode was set at a flow rate of 1 ml/min and a wave-length of 277 nm. Sample of 20 µl was injected into an HPLC system (Waters, LC-Module-I). Calibration curves were prepared in concentration ranges of 0.05-10 µl/ml. Top coating effect assessment on initial release: A model of a constant reservoir system of the release medium containing 40 ml of buffer was applied. At each sampling time-point a small aliquot of 2 ml was removed and replaced by a fresh buffer to assay drug release kinetics, under physiological conditions (37°C, PBS, pH 7.4, 100 rpm) with 0.02% SDS added to the solution to sustain sink and accelerated conditions. Samples were collected at time points 1, 3, 24h, and 3 days. Mass balance analysis of the remaining RM on stents was quantified by extraction from stents by incubation of the stent in 1 ml methanol and stirred vigorously for 30 min. Then the obtained RM solution was filtered and HPLC analyzed. 2.2.6. Therapeutic coating biocompatibility and histopathology analysis Biocompatibility study was carried out after subcutaneous (SC) implantation of stents in SD rats in a similar way to the early reported study with a slight modifications.34 16 rats were used for this study divided in 3 groups (n=4): Group A - bare metal stent implanted, Group B - stents with the RM crystalline coating, and Group C - RM crystalline coated plus HA top coating. Un-implanted rats (n=4) were used as natural control for drug extraction. At 7 and 28 days post implantation, the


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time point group was sacrificed, and the implantation site was isolated and studied. One isolated stent from each group was visualized by SEM, and the second stent implantation site of each animal was used for histopathological evaluation using the grading method.34 Retrieved stents were imaged using high resolution SEM (30 KV), to determine cell proliferation on bare and RM coated stents compared with RM coated and with protective top HA coated stents. Determination of RM in surrounded SC tissues included determination of concentration of RM in tissues surrounding implanted stent. This involved 4 steps: (1) isolation 2*2 cm of stent surrounding tissue, (2) homogenization of tissues, (3) extraction of the drug from the tissue, and (4) determining drug content by reverse phase HPLC on C-18 column.34-36 An isocratic mode consisting of water-methanol (17:83% v/v) with flow rate of 0.8 ml/min, wavelength of 277 nm and 100 µl of injected samples were applied. 2.2.7. Statistical Analysis All the experiments were repeated at least three times in triplicate. Data are expressed as mean±SD when applicable. Statistical analysis for the drug release profiles was performed on calculated area under curve (AUC) using Prism software (Version 5.00). A two-tailed unpaired Student's t-test, assuming equal variances to compare three replicate means, or One-Way ANOVA followed by Bonferroni post-hoc analysis to compare multiple replicate means. Differences were considered significant when p < 0.05. 3. RESULTS AND DISCUSSION 3.1. RM Crystalline coatings thermal analysis Thermal stability of pharmaceuticals can significantly impact on their use.37-41 To explore the thermal stability of the developed crystalline versus amorphous coatings, a milligrams quantity of RM is needed. Accordingly, a desired coating onto 5 cm long CoCr cylindrical tubes was prepared. The tubes were RM pre-seeded using 200-400 nm grounded crystals and then surface crystallized by dipping into a crystallization chamber. The developed crystalline RM coating was studied and found similar to the crystalline coatings prepared on stents as proved by XRD, a uniform and continuous crystalline carpet with crystals of parallelogram shape (Figure 1a). XRD spectra show that differences of unit cell parameters are not significant and overlap perfectly (theoretical: a = 34.850; b = 13.080; c = 12.250A˚, found: a = 34.883; b = 13.075; c = 12.253A˚). These results indicate the same material phase (Polymorph II), with differences in crystallite sizes.42 The degree of crystallinity was calculated according to the method described by Shujun et al.43 Amorphous coatings were prepared by spray coating of RM solution and immediately analyzed to avoid RM self-crystallization. Amorphous coating development was studied in different solvents: ethanol, ethyl acetate and acetone. Amorphous RM was obtained with no significant peaks shown in XRD spectra. However, in acetone, the amorphous RM contain traces of crystal patterns (Figure 1B), as reported.44 Since acetone is a very good solvent for RM, it can form very small size crystals as was found by XRD of individual peaks (Figure 1B).



Figure 1: (A) XRD analysis of crystalline RM coating collected from CoCr cylinder tubes compared to theoretical values of reference (in red), indicating the same material phase (Polymorph II). (B) XRD spectra for amorphous RM coatings prepared by fast evaporation of RM solutions. Amorphous form with trace of crystal patterns was identified using Acetone. (C) DSC profile of: (i) starting commercial crystalline RM, and peeled off coatings from tubes surface of: (ii) seeding, (iii) crystalline and (iv) amorphous layers (ethanol). Coatings were identified with Tm 190.3°C, 183.7°C and 170.5°C respectively for developed crystalline coating, followed by degradation peak, while amorphous have not showed Tm. (D) Amorphous coating DSC profile (ethanol), identified by Tg 93.94 ± 1.49°C with energy of Delta Cp 0.134-0.192 and degradation peak starting from 170°C.

The thermal stability of RM coatings was studied by DSC. RM samples from seeding, crystalline, amorphous layers and commercial RM were all evaluated, figures 1C and 1D. Two peaks were identified (180-200°C) for commercial RM starting crystals. The first peak was considered a RM melting peak, while the second was assigned for the thermal degradation (Dp). Thermal degradation following crystal melting was also identified using a melting point (Tm) apparatus in which white crystals melt followed by immediate color change to pale yellowish, then turning to brownish-black color with sample bubbling. The difference in Tm, between the starting RM (190.3°C) and the RM seeds (183.7°C) can be explained by the decrease in crystal size from (5-50) µm to (0.3-1.1) µm. A similar phenomenon has been reported for crystalline pharmaceuticals.45,46 For the developed crystalline carpet, Tm was found at 170.5°C. This can be explained because the carpet is comprised of multi drug crystal layers that include a wide range of crystal size, 0.2-5 µm. In comparison, amorphous coatings do not possess a melting peak, alternatively identified by a glass transition temperature (Tg), 93.9 ± 1.5°C, with energy of Delta Cp 0.134-0.192. These results fit the report by Maryanoff et al.47 for an amorphous RM Tg range: 93-95°C. Moreover, a wide degradation peak was identified and proven with a designed heating/cooling/heating cycle. After heating to 240°C with a flow of 10°C/min, the sample experienced fast cooling with a cooling flow of 30°C/min and again heating with flow 10°C/min. No further peak was identified, indicating drug degradation (Figure 1D). 3.2. Physical and Mechanical Stability Challenges of RM Crystalline Coatings RM crystalline-coated stent is made of crystals that have limited mechanical properties. They are fragile and not flexible, and thus it is important to protect crystals during stent handling, crimping and deployment. Figure 2 shows representative SEM images of crystalline coating defects on different stent surfaces following inflation. I have aimed to formulate a temporary fast dissolving thin top coating to absorb mechanical constrains through the stent life until deployment, then it should dissolve or erode right after stent 6 Environment ACS Paragon Plus

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fixation. For that purpose a need for flexible coating must have the capability to: (1) stretch when the stent is expanded to the new dimensions, (2) enhance the stent surface biocompatibility, and (3) not alter the drug release profile significantly. In this case the ratio between the water soluble coating and drug could be minimized, since the coating is required only during deployment and expansion. Several coatings were considered, mainly composed of water soluble polysaccharides, including: carboxymethyl cellulose sodium salt (CMC), sodium alginate, chitosan and hyaluronic acid (HA).48-52 Top coating properties were studied, and potential improvements were examined using plasticizers such as triethyl citrate (TEC), glycerol and PEG Mw 1000-15000 for enhancing the flexibility. Stents after inflation were visualized and evaluated using light microscopy, SEM and HPLC analysis.

A

Rough Surface

B

Smooth Surface

A C

B D

Figure 2: SEM images of metal stents with different surface properties/nature before and after surface crystallization and inflation: (A) rough surface- deviation of hundreds of nanometers to few microns, and (B) smooth surface- deviation of few nanometers, magnification 1000x. (C) And (D) rough and smooth surface, crystalline RM coated after stent inflation, magnifications 1500x and 3491x respectively. In comparison, crystalline carpet exhibited less adhesion to the smooth stent struts.

3.2.1. Stent inflation and protective top coating effect on initial release in vitro Two different methods were tested to apply the top coating: dip and spray coating, were two concentrations tested, 0.5% and 2% w/v. For alginate, an extra step was added to dip the alginate coated stent into 0.1% or 0.5% calcium chloride solution for fixation of the top coating. Top coating weight increase was determined by micro-analytical balance per stent mm length and found: 15-18 µg for CMC, 23-26 µg for alginate, and 50 µg for chitosan 2% aqueous solution, Table S1 (Supporting Information). Table S2 summarizes weight increase as a function of the number of coating layers. It was found that multilayer coating is less effective than a single coating of solution of increased concentration. Top coating characteristics such as flexibility, stretching and strictness were evaluated using a preprepared film from 0.5% and 2% w/v solutions. Table S3 summarizes the difference between the films. The ability to stretch was found to be: Chitosan (Rigid Film) < Alginate < CMC < HA (Flexible Film). TEC and PEG (1000 and 4600) were added to increase flexibility and stretch ability, Table S4. HA and CMC polymers showed the most flexible properties, while alginate and chitosan films were least and were not tested further. CMC was further studied for enhancing performance with five different Mw of PEG at 3 different percentages: 10%, 20% and 30% w/w, Table S5. The formulation of CMC with 10% of PEG 4600 showed the most appropriate film with high capability of stretching and flexibility. Both formulations of HA and CMC (10% of PEG 4600) were applied on both bare and crystalline RM coated stents. Stent coatings were visualized and checked by light microscopy before and after stent inflation, Table S6. The use of HA as an upper coating showed excellent results due to unique flexibility, Figure S1 (Supporting Information). For RM surface coated stents, both leading formulations were used to top coat by a dipping method. Six stents loaded with ~50 µg RM /mm stent length (totally 750 µg RM per stent) were prepared for each type of top coating. The weight increase after top coating process was traced and found to be ~(20-30) µg poly/mm stent length.



Figure 3: (A) Light microscope photo analysis for the stents before (A-1, B-1, C-1) and after stent inflation (A-2, B-2, C-2) for three groups as follow: (A) RM crystalline coated stent without and with top coating of CMC (with 10% w/w PEG 4600) and HA formulations, magnification 4x. Images (B) and (C) cumulative RM release profiles of loading% and absolute microgram released respectively, *: P < 0.05, n.s.: not significant. No significant difference between the releasing profiles of polymer top coated to uncoated, indicating it doesn’t affect the initial release profile.

Representative stents were used to determine drug lost following inflation by HPLC, and the rest used for examine the effect of both top coating and stent inflation on the initial drug release, Figure 3. The groups were as follow: (A) without top coating, (B) with CMC formulation, and (C) HA top coated. Of six stents in each group, 1 stent was utilized for light microscope visualization, 2 for HPLC determination of remaining drug on the stent after inflation, while the remaining 3 stents were used for an initial release study. Representative light microscope photos analysis for the stents before (A-1, B-1, C-1) and after inflation (A-2, B-2, C-2) are given in Figure 3A. Without top coating after inflation, as expected, crystalline coatings were breakable. Figure 3, images (A-1 and A-2), highlight the damage to the crystalline layer specifically in the regions with mechanical stress, i.e.,"Y regions". Drug loss ranged at 20-30% of the total stent drug loading. CMC polymer top coated stents (B-1 and B-2) showed poor efficiency, since protecting coating experienced a 10% drug loss as determined by HPLC. CMC Top coating was not sufficiently flexible, cracked and tore to pieces of clusters. On the other hand, HA showed excellent protective coating properties, and less than 3-4% drug lost was encountered. The polymer coatings exhibit flexible and elastic properties and found to be expansible, without coating breaking. HA showed impressive adaptive ability to the new dimension of the stent without any damage to the crystalline layer. All the stents exhibited a slow releasing profile, were only ~ 6% (30-45 µg) was released along the 3 days. For HA, no significant difference between the release profiles of the top coated and uncoated stents was found, indicating that thin polymer top coating does not affect the release profile, Figures 3B and 3C. 3.2.2. Sono-Tek spray coating system Although among studied polymers, HA exhibited the best protective covering layer according to the previous analysis. However, further improvements are still needed to improve performance, i.e., completely reducing drug lost and preventing webbing-top coating accumulation between stents struts- Figures 4A and 4B. Upon stent inflation, webbing can hinder the expansion process leading to defects in stent structure or in the therapeutic coating, Figure S2. This can cause problems such 8 Environment ACS Paragon Plus

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as release profile disruption.53-55 To avoid this undesired byproduct, I took two major steps: (1) replacing the dipping method with spray coating, and (2) further adjusting the top coating by lowering the concertation and adding glycerol as a plasticizer. Spray coating was performed using the Sono-Tek automated system, Figure S3. A dilute solution of HA and the use of spray coating machine prevented webbing, Figures 4C and 4D. For the spray coating method a study focused on the relationship between the HA concentration,56 number of coating layers, and the amount of coating per stent, Table S7. A 10% w/w of glycerol was added to the sprayed solution to increase flexibility and streamline properties of HA as coating material. Representative light microscope images for top coating coverage development are given in Figure S4. The pictures show that spray coating is more effective than dip coating with no webbing formation. The weight increased significantly by using the spray process, which detected a 12 µg of polymer coating per one layer at 0.1% w/v or 17 µg for 0.1% with 10% glycerol, Table S7.

Figure 4: Representative images for undesired webbing of top coated stent with 1% w/v HA by dipping method: (A) Light microscope photo of stent with crystalline coating top coated, webbing between the stent's struts are marked with red arrows, magnification 4x. (B): SEM image of webbing, magnification 400x. (C) And (D) RM crystalline coating top spray coated with HA formulation (HA-F), before and after stent inflation respectively, magnification 4x. During stents inflation there was no damage observed to the drug crystalline carpet.

Further optimization of the formulation and the coating process caused the 0.3% HA (w/v) solution with 10% glycerol (w/w) to yield the most flexibility and compatibility as the best coating for crystalline coated stents. This formulation was named further as "HA-F" referring to the HA based formulation composed of 0.3% HA + 10% glycerol. The result was a stable coating that covers and protects the drug layer, Figure 4C. No drug lose was observed, and a continuous drug carpet was identified before and after stent inflation. Also, no webbing or damage to the drug crystalline coating was observed. The upper protective layer was found easily to stretch without torn with stent inflation, Figure 4D. 3.2.3. Therapeutic coating topography, top coating thickness and roughness Surface topography analysis is important for implanted devices particularly when inserted into blood vessels.56 Thus, the thickness and surface roughness of the crystalline carpet before and after top spray coating of HA-F was carefully analyzed by AFM, SEM and profilometer analysis. Here, roughness is defined as the standard deviation of Z axis (height). Stents with a clinical therapeutic dose of ~100µg RM crystalline coating per stent were studied. Cross section analysis of the crystalline carpet was performed using a sharp scalpel, Au/Pd coated and analyzed by HR-SEM. Coating thickness and crystal hump height over the upper level of crystals carpet were also evaluated. Adding the top coating was found to smoother the surface as explained in the schematic illustration, Figure 5A. Cross section SEM images revealed that the thickness of crystalline coating is mainly in the interval 3-5µm, while 1 layer of top coating slightly increased the therapeutic coating thickness by 200-500nm. This resulted in a significant change in therapeutic coating roughness, as seen in Figure 5. Stent strut topography for bare metal, seeded, crystalline RM and HA-F top coated stents were examined by AFM. Representative images of the upper crystals layer, crystal hump height, and differences between the lowest-highest crystals edges on the upper top of crystal carpet are summarized in Figures 5C and 5D. The surface roughness of bare metal stent changed from 10 ± 2 nm, to 1115 ± 212 nm after seeding, while RM crystalline coating was identified with the highest diversity of 800-1500 nm. As expected, the HA-F top coating dramatically reduced roughness to 120-150 nm. Roughness changes was also traced by profilometer on cylindrical tubes prepared similarly to stents. A similar predictable roughness change was found, starting from 0 (assuming 9 Environment ACS Paragon Plus


the surface is very smooth) to ~100 nm after seeding, followed by a significant change to almost 920 nm due to crystallization process and finalized to 398 nm as a result of top coating smoothing, Figure 5E.

Figure 5: (A) Schematic illustration of the topography change following spray coating. Top coating smooths the sharp surface of upper layer of crystalline coating. (B) SEM images of: crystalline RM carpet onto stents struts before (i) and after (ii) top coated with 1 layer of HA-F, magnifications 1500x, and its cross section images (iii) and (iv) magnifications 8000x and 2500x, respectively. Highlighting starting seeds (pink arrows), crystalline carpet multilayers-strata (blue arrows), crystals humps in the carpet upper level (red arrows), and top coating (green arrow) while surface roughness dramatically changed. (C) AFM topographical pictures of stent struts: (i) bare stent and after: (ii) seeding, (iii) crystallization, and (iv) top coating. (D) Representative 3D topographic AFM image for top coated stents highlight minimizing the differences between the lowesthighest crystals edges on the upper top of crystal carpet. (E) Profilometer analysis of roughness and coating thickness transformation during therapeutic coating preparation onto cylinders tubes.

3.2.4. Mechanical tests: crimping, balloon inflation and deflation - dislodgement test Stents with therapeutic dose of ~100±10 µg RM crystalline coating per stent, with or without ~50±10 µg HA-F top coating, were prepared and tested for mechanical properties and stent dislodgement. Smooth and rough CoCr stents surface with 50.53 mm2 surface area, were used to study crystalline coating stability and tested for crimping, balloon inflation, deflation and dislodgement (Figure S5). All force values were above 1.2 N to ensure that dislodgement force values be within acceptance criteria.57,58 Hot crimp crystalline coatings showed 4.50 ± 1.3 N and the HA-F top coated 4.08 ± 0.87 N, while for the cold crimp the results were 5.95 ± 0.62 N and 5.80 ± 0.43 N respectively. Expanded stents were dissolved in methanol for drug extraction and then HPLC analyzed. Crimped and nested stents were compared in terms of drug purity. In cold crimp at least 98.4% drug purity was found for top coated, 98.1% in hot crimp, while more than 10% lost for unprotected crystalline coating. Changes in the crystalline coating and adhesion to stent surface with/without top coating over smooth or rough surface were carefully traced with SEM macroscopy, Table 1. For the non-top coated, a significant difference was seen in the adhesion strength of the crystalline carpet to the stent surface. Expansion of smooth surface stents resulted in weak adhesion at the strut bended area, while minimal changes were noticed for the rough surface Figure 2 and Table 1. All top coated stents showed excellent mechanical properties regardless of the native surface roughness level. No coating carpet defects after stents crimping and/or expanding were found. One layer of HA-F as a top coating was sufficient to protect the crystalline carpet from cracking and/or defects. Despite the fact that two layers provide an extra smoother surface (Figure S6),


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multiple layers increase the chance of webbing as discussed previously. This must be avoided (Figures 4 and S2). Table 1: The effect of surface properties and top coating in the inflation process

Magnifications: (A): (i)=1500x, (ii)=700x, (iii)=400x, (iv)=400x; (B): (i)=1500x, (ii)=1500x, (iii)=400x, (iv)=346x; (C): (i)=1500x, (ii)=1500x, (iii)=400x, (iv)=272x; (D): (i)=500x, (ii)=1500x, (iii)=381x, (iv)=602x; (E): (i)=1500x, (ii)=1500x, (iii)=577x, (iv)=400x. Crystalline drug loaded-1 or 2 HA-F coated layers (0.3% HA+10% w/w glycerol).

3.2.5. Chemical stability of RM crystalline coating with HA-F top coating RM is a macrolide identified to have rapid degradation to form newer compounds with opened lactam ring. The degradation rate depends on several factors, mainly temperature, pH and medium concentration.34 A chemical stability study for RM crystalline coating along with HA-F top coating was conducted for 1 year at different temperatures 37ºC, (4-8)ºC and -20ºC. SEM evaluation, drug content analysis and in vitro RM release were carried out for 2 months. Results indicated no change in integrity, morphology and topography of crystalline coating, since both crystalline and top coatings exhibit high stability at different temperatures for the extend time of 1 year, Figure 6A. Release data show identical RM release profiles for 2 months, as can be seen in Figure 6B, and validated with mass balance analysis, Table S8. Nevertheless, no difference between the ratios of degraded/undegraded released forms indicating high stability of crystalline carpet with the top coating, Figure S7. Following the release study, stents SEM visualized for remaining crystalline coatings analysis and found to highly adhere to stent's metal surface, Figure S8. The crystalline coating content was analyzed for 1 and 12 months storage. After 1 month storage no degraded form was detected and negligible difference was found for the stents stored for 1 year were showed 2.76%, 2.46% and 2.33% degraded form for storage at 37ºC, (4-8)ºC and -20ºC respectively. For comparison, amorphous coatings were stored for 1 month at 37ºC and (-20)ºC and analyzed similarly, Figure 6C. The amorphous coating stored at (-20)ºC was identified with partial selfcrystallization where amorphous material shrank in a form of crystals, while at 37ºC coating was soft, fragile and after erosion process, thus indicating a larger degradation process. This observation harmonized with the drug content analysis which showed 20-25% of the drug in degraded form when stents were stored at (-20)ºC while 35-40% with storage at 37ºC, Table S9. Also, stents stored at (-20)ºC exhibited a slightly slow releasing profile (Figure 6D) with a lesser degraded form compared to 37ºC stored (Figure S9). These findings highlight the enhanced chemical stability of 11 Environment ACS Paragon Plus


the active therapeutic agent, RM, due to crystalline coating and preservation of the drug from degradation compared to amorphous coating.

Figure 6: (A) Representative SEM images (magnifications 400x and 1000x, respectively) of crystalline coated stents stored for 1 year at: 37ºC (i) and (ii), (4-8)ºC: (iii) and (iv), (-20)ºC: (v) and (vi). Both crystalline coatings and HA-F top coating exhibit high stability for an extend time of 1 year with no abnormality or cracks identified. These pictures emphasize the coating stability, integrity and topography. (B) Cumulative RM released from crystalline coated stents stored for 1 year. The different storage conditions does not affect the releasing profile and were found identical. n.s.: not significant. (C) Representative SEM images (magnifications 400x and 800x, respectively) of amorphous coated stents prepared by spray coating method: freshly prepared (i) and (ii), and after 1 month stored at: 37ºC (iii) and (iv), or (-20)ºC, (v) and (vi). Coatings stored at (-20)ºC identified with partial self-crystallization, while at 37ºC found soft, fragile with signs of erosion. (D) Cumulative RM released from amorphous coated stents stored for 1 month.

3.3. In vivo examination of protective layer- biocompatibility study For biocompatibility evaluation, 3 groups of stents were implanted and tested. Bare stents (Group A) and both crystalline coating (Group B) and crystalline with protective HA-F top coating (Group C) were implanted SC into SD rats. At 7 and 28 days post implantation rats were scarified, and SEM analysis for implanted stents and histology study of tissues surrounding stents was performed. As similar to the early study34, it was observed that efforts were needed to remove the bare stent from the tissues at implant location, highly adherent to surrounding tissues, Figure 7A. In Group B, stent removal from the implanted site was relatively easier. While Group C found extremely smooth and both were surrounded by a pocket of loose tissue, especially for Group C stents were completely intact. This findings proves that besides the controlled RM release from both stent groups for several weeks reducing growth of surrounding tissue on the stents, the protective layer was effective in protecting the crystalline coating from friction, defects and secured integrity of the coatings. This may have led to symmetrical drug release from the implant as proven by SEM and drug extraction from the nearby tissue. SEM analysis of retrieved stents, showed that stents from Group A identified with massive tissue surrounding the stents from all directions. For Group B, the presence of crystalline carpet of the RM, repeated the early reported significantly reduction of tissue proliferation onto the stent surface while some specific areas tissue covering still be observed or inside the stent, Figure 7B.34 On the other hand, in Group C the therapeutic coating was significantly more successful with complete


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preventing of tissue proliferation onto the stent struts or inside since more drug is localized in the dislodgement site, Figure 7B.

Figure 7: (A) 2 cm2 of isolated tissues containing the stent embedded between the tissue layers (Group A, Day 7). (B) SEM images of implanted stents in SD rats, 28 days post SC implantation for Groups: A (bare stent, i-ii), B (RM crystallized coated, iii-iv), and C and (RM crystallized stent top coated with protective layer, v-vi), upper (120x) and lower row (250x, 500x and 250x respectively) magnifications. (C) RM quantification in surrounding tissues and remaining onto stent struts for Groups B and C. For both time points, 7 and 28 days post implantation, more drug was found for stents top coated in comparison to uncoated, attributed to the top coating protective function during stent insertion to the implantation site. (D) Figures i-vi: Histopathologic evaluation 28 days post implantation for capsule formation surrounding the cavity where the stent was located. Characteristically, Group A: the capsule (arrow) consists mostly of fibroblastic proliferation associated with collagen deposition; there are only minimal sparse mononuclear cells (i.e., lymphocytes) within the capsule (i, 100x and ii, 200x respectively). Group B: the capsule (arrow) consists mostly of fibroblastic proliferation associated with collagen deposition. There are only singular histiocytes within the cavity (iii, 200x and iv, 400x respectively). Group C: relatively minor capsule formation (arrow) consists mostly of fibroblastic proliferation (minimal, grade 1) associated with collagen deposition (slight, grade 2), with no active inflammatory reaction or necrosis is present (v, 100x and vi 400x respectively).

Determination of the RM amount in surrounding tissues and remaining on the stents were done by HPLC. Variable amounts of the drug were present in surrounding tissues and on the stents, in Group B, Figure 7C. The reason for this can be the stent separation process, where some stent struts were partially covered by surrounding tissues. However, in both locations, more drugs were identified in Group C rather than Group B. This can be attributed to the function of the top coating. As expected, results clearly show RM was released in a sustained manner, and the RM amount found in both locations was less on 28 days compared to 7 days post implantation. Histopathologic evaluation of the surrounding tissues from 7 and 28 days implanted samples indicated that for both Groups B and C, the capsule reaction consisted of relatively maturing collagen, generally without any evidence of an inflammatory reaction. In particular, no evidence of the presence of necrosis, foreign body giant cell reaction or any type of increased severity of inflammatory reaction was noted due to HA-F top coating, Figure 7D. Detailed necropsy was performed 7 and 28 days post implantation. Weighing of the heart, spleen, kidneys, lungs and liver was performed using an analytical balance. No difference in organ weight or any gross abnormality was noted between either both RM crystallized without or with protective coating compared to unimplanted suggesting that the developed protective coating onto RM crystallized lack of toxic effects. 3.4. Bioactive top coating preparation-incorporated with RM Drugs applied extra-luminally will diffuse into the media as well as the subintima and can prevent the production of subintimal plaque. However, the concentrations of drug that reach the intima or the endothelial layer are markedly diminished, lessening the chance that drug application will inhibit re-endothelialization of the vessel lumen.59-61 Taking that into account, with the revolution in stenting cardiovascular complications and the growing interest in personalized medicine, I have 13 Environment ACS Paragon Plus


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also examined the potential of bifunctional bioactive top-coating designing. A top coating, protecting the crystalline therapeutic coating during stent inflation and displacement, and can ensure initial burst release of the drug in the final destination site. Such localized saturation of the implantation site with an active drug is anticipated for increasing efficiency of treatment. Accordingly, incorporation of RM into HA-F formulation by mixing HA solution with RMmethanol solution and spraying over bare stents or with active crystalline coatings was examined. Such therapeutic coatings were designed for different releasing profiles, single and multilayer releasing systems, either for fast or fast followed by slow releasing from crystalline coatings. As detailed in the methods section, solutions ratio top coating (HA-F):RM methanol of 1:1 v/v was applied with RM 10 mg/ml using the aforementioned spray coating process. Top coatings incorporated with RM were prepared in duplicate (2 batches of solutions mix) with either 100µl sprayed volume per cycle of coating or overloading with continuous 300µl sprayed volume. Total drug increase per stent was determined by microanalytical balance and HPLC validated for representative stents, Table S10. Immediately following coating preparation, an in vitro study was conducted, Figures 8. A cumulative RM release study from HA coatings in correlation to coating cycles was conducted. The more cycles of coating, the more drug is released, Figure 8A. Nevertheless, continuous RM release from these coating over the study period was found, while 60-90% (25-75 µg) of drug loaded coatings were released in the first 72 hours, Figures 8B and S10. Also the stents coated with minimum coating cycles released the max% of RM loading, Figures 8B.

Figure 8: (A) Cumulative RM release (µg) from HA coatings under physiological conditions. *: P < 0.05. (B) Cumulative RM release% from HA coatings, RM releasing normalized to drug loading focusing on the first 72 hours. *: P < 0.05, n.s.: not significant. Continuous RM releasing from these coating over the study period was found while 60-90% (25-75 µg) of drug loaded coatings are released in the first 72 hours. Stents coated with minimum coatings cycles released the max% of RM loading.

3.5. RM Release using multi-layer coatings and bioactive top coating The effectiveness of the previous process was applied to stents with already therapeutic coating and studied in vitro for 28 days. Different coatings comprised of multiple layers were prepared, Table 2. For crystalline coatings coated with bioactive top layer (protective layer incorporated RM), a burst RM release followed by slow release from crystalline coating was obtained, Figures 9A-D. Blank HA coated on crystallized surface stent did not affect the release profile. However, incorporation RM in HA, it gave initial release more than amorphous coating for the first few hours, followed by slow release profile, Figure 9A-B. Moreover, by normalization of the release profiles with drug loading%, a new releasing profiles lay in between amorphous and crystalline coatings: burst followed by slow release profiles and vice versa, Figures 9C-D. Table 2: HA incorporated RM coatings development Coating Num 1 2 3 4 5 6

Coating Descriptiona Crys-S-1 Crys-S-2 HA-Crys-S-1 HA-Crys-S-2 (HA:RM)-Crys-S-1 (HA:RM)-Crys-S-2

RM Coating (µg)b 173.0 144.0 166.0 (100.0 HA) 155.0 (80.0 HA) 138.0 109.0

Top Coating Processc ----------------2 Cycles100µl 2 Cycles100µl

RM In Top Coating (µg)d ----------------69.0 54.5


HA (µg)e ----------------69.0 54.5

Total Drug Loading (µg)f ----------------207.0 163.5

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7 (HA:RM)-HA-Crys-S-1 114.0 (101.0 HA) 2 Cycles100µl 59.5 59.5 173.5 8 (HA:RM)-HA-Crys-S-2 153.0 (95.0 HA) 2 Cycles100µl 53.0 53.0 206.0 9 Amorph-Crys-S-1 180.0 ----108.0 ----288 10 Amorph-Crys-S-2 188.0 ----69.0 ----257 11 Amorph-S-1 117.0 ----------------12 Amorph-S-2 242.0 ----------------a Abbreviations: S-Stent (metal surface). Crys-Crystalline coated stents. HA-Hyaluronic acid. RM-Rapamycin. (HA:RM)-coating based HA incorporating RM. Amorph-amorphous coated stents. Coating layers order starts from the character S, for example HA-Crys-S it is a mean, stent with first crystalline coated followed by HA coating. b Microanalytical balance and HPLC. c Each cycle- 100µl sprayed volume. d Drug incorporated in the final top coating. e HA weight included glycerol 10% w/w (HA-F). f For both stents strut drug coated and in the HA top coating.

Figure 9: Cumulative RM release from multi-layer coatings under physiological conditions. (A) Cumulative RM release (µg). *: P < 0.05, n.s.: not significant. (B) Cumulative RM release (µg), focusing on the first 24 hours. *: P < 0.05, n.s.: not significant. (C) Cumulative RM release%, RM releasing normalized to drug loading. *: P < 0.05. (D) Cumulative RM release%, focusing on the first 24 hours. *: P < 0.05. Incorporation of RM in protective top coating, it gave initial release more than amorphous coating for the first few hours, followed by slow release profile from crystalline coating. A multilayer release system was found achievable either by incorporation the drug in the protective top coating or by amorphous coating prepared on the top of crystalline coating. *: P < 0.05, n.s.: not significant

4. SUMMARY AND CONCLUSION This study examine crystalline coating thermal, physical and mechanical stability of carrier-free DES, and presents a successful development of novel protective top coating formulation for enhancing crystalline carpet merits, using RM as a model drug. The obtained top coating is durable, flexible, homogenous, uniform and can absorb strains without interfering with the release profile of the crystalline coating. Also, showed enhanced chemical stability with ~2% degraded form after 1 year. Biocompatibility study prove suitability of the developed top coating with no observable toxicity or safety concerns in delivering RM. Bioactive top coatings for multilayer release purposes are stable and achievable by incorporating RM in the developed protective top coating. It can be anticipated top coating can be further implemented for local drug delivery of various other drugs and implanted devices coatings, either single or multiple drugs for sequential release purposes. Supporting Information The Supporting Information is available free of charge on the ACS Publications website at DOI:XXX Details of the polymeric coating weight increase versus: type of polymer, polymer concertation, coating conditions and technique, multi-layer coatings; polymers top coating films preparation and characterization: effect of plasticizer type and percentage; microscope evaluation: top coatings onto bare metal stent, webbing effect on crystalline coated stents, after stent crimping or expanding 15 Environment ACS Paragon Plus


(mechanical test); SonoTek MediCoat spray coating system; top coating SEM images of surface smoothing effect; cumulative RM components degraded/undegraded (µg) and total released in PBS at 37ºC for: crystalline coated stents stored for 1 year at: 37ºC, (4-8)ºC and (-20)ºC, and for amorphous coated stents stored for 1 month at (-20)ºC and 37ºC; SEM images stents after 2 months release and mass balance of stents after release study; drug content component (degraded/undegraded) analysis by HPLC versus storage conditions; HA incorporated RM coatings development and cumulative RM release (µg) from HA coatings, focusing on the first 72 hours in PBS at 37ºC. 5. ACKNOWLEDGMENTS This work was supported in part by a grant from the Israel Science Foundation (ISF) (No. 1051/12). Council for Higher Education-Israel, acknowledged for providing fellowship to Shady Farah. I should like to thank the Unit for Nanoscopic Characterization of the Hebrew University of Jerusalem for performing the XRD (Dr. Vladimir Uvarov) analysis and Dr. Piotr S. Kowalski (Daniel G. Anderson/Robert S. Langer Labs, The David H. Koch Institute for Integrative Cancer Research, Massachusetts Institute of Technology) for his assistance in performing statistical analyses of data sets. 6. REFERENCES (1) Khan, W.; Farah, S.; Domb, A.J. Drug Eluting Stents: Developments and Current Status. J. Control. Release. 2012, 161, 703–712. (2) Levy, Y.; Tal, N.; Tzemach, G.; Weinberger, J.; Domb, A.J.; Mandler, D. Drug-Eluting Stent with Improved Durability and Controllability Properties, Obtained via Electrocoated Adhesive Promotion Layer. J. Biomed. Mater. Res. B Appl. Biomater. 2009, 91, 819–830. (3) Kanjickal, D.G.; Lopina, S.T. Modeling of Drug Release from Polymeric Delivery Systems—a Review. Crit. Rev. Ther. Drug Carrier Syst. 2004, 21, 345–386. (4) M.L. Houchin, E.M. Topp, Chemical Degradation of Peptides and Proteins in PLGA: a Review of Reactions and Mechanisms, J. Pharm. Sci. 2008, 97 (7), 2395–2404. (5) Hu,T.; Yang, J.; Cui, K., Rao, Q.; Yin, T.; Tan, L.; Zhang, Y.; Li, Z.; Wang, G. Controlled Slow-Release DrugEluting Stents for the Prevention of Coronary Restenosis: Recent Progress and Future Prospects. ACS Appl. Mater. Interfaces 2015, 7, 11695−11712. (6) Saleh, Y.E.; Gepreel, M.A.; Allam. N.K. Functional Nanoarchitectures for Enhanced Drug Eluting Stents. Sci. Rep. 2017, 7, 1-12. (7) Han, J.; Farah, S.; Domb, A.J.; Lelkes. P.I. Electrospun Rapamycin-Eluting Polyurethane Fibers for Vascular Grafts. Pharm. Res. 2013, 30, 1735-1748. (8) Farah, S.; Anderson, D.G.; Langer, R. Physical and Mechanical Properties of PLA, and Their Functions in Widespread Applications—a Comprehensive Review. Adv. Drug Deliv. Rev. 2016, 107, 367−392. (9) Spaulding, C. The Question of Drug-Eluting Stent Safety: then and Now. Am. J. Cardiol. 2008, 102, 12J–17J. (10) Slottow, T.L.; Waksman, R. Drug-Eluting Stent Safety. Am. J. Cardiol. 2007, 100, 10M–17M. (11) Nakazawa, G.; Finn, A.V.; Ladich, E.; Ribichini, F.; Coleman, L.; Kolodgie, F.D.; Virmani, R. Drug-Eluting Stent Safety: Findings from Preclinical Studies. Expert. Rev. Cardiovasc. Ther. 2008, 6, 1379–1391. (12) Raja, S.G.; Berg, G.A. Safety of Drug Eluting Stents: Current Concerns and Controversies. Curr. Drug. Saf. 2007, 2, 212–219. (13) Belle, E.V.; Susen, S.; Jude, B.; Bertrand, M.E. Drug Eluting Stents: Trading Restenosis for Thrombosis. J. Thromb. Haemost. 2007, 5, 238–245. (14) Kim, H.L.; Park, K.W.; Kwak, J.J.; Kim, Y.S.; Sir, J.J.; Lee, S.J.; Lee, H.Y.; Chang, H.J.; Kang, H.J.; Cho, Y.S.; Chung, W.Y.; Chae, H.H.; Choi, D.J.; Kim, H.S.; Oh, B.H.; Park, Y.B.; Koo, B.K. Stent-Related Cardiac Events after Non-Cardiac Surgery: Drug-Eluting Stent vs. Bare Metal Stent. Int. J. Cardiol. 2008, 123, 353–354. (15) Sharifkazemi, M.B.; Zamirian, M.; Aslani. A. A Current Problem in Cardiology: Very Late Thrombosis after Implantation of Sirolimus-Eluting Stents. Cardiology. 2007, 108, 273–274. (16) Ruchin, P.E.; Muller, D.W.; Faddy, S.C.; Baron, D.W.; Roy, P.R.; Wilson, S.H. Long-Term Clinical Follow-up of Sirolimus-Eluting (Cyphertm) Coronary Stents in the Treatment of in-Stent Restenosis in an Unselected Population. Heart. Lung. Circ. 2007, 16, 440–446. (17) Unger, F. Drug-Eluting Stents a Nightmare?. Interact. Cardiovasc. Thorac. Surg. 2007, 6, 813–814. (18) Otsuka, Y.; Chronos, N.A.F.; Apkarian, R.P.; Robinson, K.A. Scanning Electron Microscopic Analysis of Defects in Polymer Coatings of Three Commercially Available Stents: Comparison of BiodivYsio, Taxus and Cypher Stents. J. Invasive. Cardiol. 2007, 19, 71–76. (19) Windecker, S.; Mierer, B. Late Coronary Late Thrombosis. Circulation 2007, 116, 1952–1965.


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Table of Contents/Abstract Graphic Protective Layer Development for Enhancing Stability and Drug Delivery Capabilities of DES Surface Crystallized Coatings Shady Faraha,b a

Institute of Drug Research, School of Pharmacy-Faculty of Medicine, Center for Nanoscience and Nanotechnology and The Alex Grass Center for Drug Design and Synthesis, The Hebrew University of Jerusalem 91120, Israel. b Daniel G. Anderson/Robert S. Langer Labs, Department of Chemical Engineering and The David H. Koch Institute for Integrative Cancer Research, Massachusetts Institute of Technology, 500 Main Street, Cambridge, Massachusetts 02139, USA. Corresponding author, E-mail: [email protected] , Tel.: +1 857-4159686.


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