chemical engineering research and design 8 9 ( 2 0 1 1 ) 328–334

Contents lists available at ScienceDirect

Chemical Engineering Research and Design

journa l homepage: www.e lsev ier .com/ locate /cherd

Effect of copolymer ratio on hydrolytic degradation ofpoly(lactide-co-glycolide) from drug eluting coronary stents

Chhaya Engineera,1, Jigisha Parikha,∗, Ankur Ravalb,2

a Department of Chemical Engineering, Sardar Vallabhbhai National Institute of Technology, Ichchanath, Surat 395007, Gujarat, Indiab Sahajanand Medical Technologies Pvt. Ltd., Wakhariawadi, Nr. Dabholi Char Rasta, Ved Road, Surat, India

a b s t r a c t

The in vitro hydrolytic degradation behavior of poly(d,l-lactide-co-glycolide) (PLGA) has been systematically inves-

tigated from the drug eluting coronary stents with respect to different copolymer compositions. The drug–polymer

coated stents were incubated in phosphate buffer saline (pH 7.4) at 37 ◦C and 120 rpm up to 12 months to facili-

tate hydrolytic degradation. Gel permeable chromatography, differential scanning calorimetry and scanning electron

microscopy were employed to characterize their degradation profiles. The study supports the bulk degradation behav-

ior for PLGA from coated stents. Molecular weight of polymer decreased immediately after immersion in PBS but

mass loss was not observed during first few days. The rate of hydrolytic degradation was influenced by copolymer

ratio, i.e., degradation of 50:50 PLGA was fastest followed by 65:35 PLGA and 75:25 PLGA. The drug release from PLGA

coated stent followed biphasic pattern which was governed by surface dissolution and diffusion of drug rather than

polymer degradation.

© 2010 The Institution of Chemical Engineers. Published by Elsevier B.V. All rights reserved.

Keywords: Poly(d,l-lactide-co-glycolide); Hydrolytic degradation; Restenosis; Drug eluting stent

In recent years, a number of parameters have been iden-

1. Introduction

Coronary stents are a major advance in the treatment ofobstructive cardiovascular disease. However, in the substan-tial number of patients, stent placement can trigger restenosisafter implantation (Costa et al., 2002). The concept of usingdrug eluting stents for prolonged, sufficient, and localizeddrug delivery to address restenosis is an important contem-porary advance in interventional cardiology (Costa et al., 2002;Babapulle and Eisenberg, 2002; Sousa et al., 2003).

Polymers play a critical role in local drug delivery fromthe stent scaffold and to date, attempts to deliver drug with-out polymer have not proven successful (Lansky et al., 2004;Teirstein, 2004; John et al., 2008). The majority of first gen-eration drug eluting coronary stent coatings are based onhydrophobic polymers which retain and release drug in a con-

trolled fashion. Biodegradable polymers are often cited as analternative to biostable polymers for drug eluting stent coat-

∗ Corresponding author. Tel.: +91 261 2201689; fax: +91 261 2201641.E-mail addresses: e.chhaya@ched.svnit.ac.in (C. Engineer), jk parikh

ankur.med@sahmed.com (A. Raval).Received 30 March 2010; Received in revised form 28 June 2010; Accep

1 Tel.: +91 261 2201689.2 Tel.: +91 261 2521251; fax: +91 261 2520252.

0263-8762/$ – see front matter © 2010 The Institution of Chemical Engidoi:10.1016/j.cherd.2010.06.013

ings (Doyle and Holmes, 2009; Hezi-Yamit et al., 2009). Thesepolymers degrade temporally, leaving behind only a bare metalstent.

The family of aliphatic polyesters has been by far thedominating choice for materials in degradable drug deliverysystems. The most extensively investigated and advancedpolymers in regard to available toxicological and chemicaldata are the polylactide (PLA) and poly(lactide-co-glycolide)(PLGA) (Lewis, 1990). The popularity of PLA and PLGA is fur-ther explained by the fact that FDA has approved them for thenumber of clinical applications (Edlund and Albertsson, 2002).However the biocompatibility of these polymers, specificallyin a vascular setting, depends to a large extent on degradationkinetics. Moreover, drug elution kinetics from the stent alsogreatly depends upon polymer degradation.

@yahoo.co.in, jkp@ched.svnit.ac.in (J. Parikh),

ted 29 June 2010

tified that influence the polymer degradation. Among themare the copolymer composition (Vert et al., 1984), morphol-

neers. Published by Elsevier B.V. All rights reserved.

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gy (Vert et al., 1992), autocatalysis by acidic degradationroducts inside a matrix (Li et al., 1990a,b), presence ofrugs (Li et al., 1996) or other excipients (Heller, 1986) andreparation technique (Mathiowitz et al., 1990). It was alsostablished that the hydrolytic degradation rate is higher incidic pH than in neutral media (Kiss and Vargha-Butler, 1999).he hydrolytic polymer degradation was found to be depen-ent on the degradation temperature, namely, whether itas above or below the glass transition temperature (Tg) of

he selected polymer sample (Kiss and Vargha-Butler, 1999).olymers in amorphous state degrade faster than crystallinetate (Vert et al., 1992). Preparation technique is also criti-al because spray-dried particles degrade faster than particlesrepared by solvent evaporation (Giunchedi et al., 1998). Sizend geometry of polymeric device also influence degradationehavior. Large size plates degrade faster and heterogeneoushan thinner films (Dunne et al., 2000; Grizzi et al., 1995).he physico-chemical properties of the incorporated drug asell interaction between polymeric matrix and drug is criticalarameter which has a strong effect on polymer degrada-ion and the drug release (Chen et al., 2000). For example,ydrophilic drugs or excipient facilitate water penetration inhe system and lead to the creation of highly porous poly-

er networks upon drug leaching; thus accelerate polymeregradation. In contrast, lipophilic drugs hinder water diffu-ion into the matrix and retard polymer degradation (Kloset al., 2008). For acidic drugs, faster hydrolysis of ester bondsecause of acid catalysis can be observed which acceleratesolymer degradation (Li et al., 1996; Frank et al., 2005). In con-rast, in the case of basic drugs two effects can be observed:ase catalysis of ester bond cleavage and neutralization ofarboxyl end groups of polymer chains which minimizes orliminates the autocatalytic effect of acidic chain ends. Thushe degradation can be accelerated or slowed down depend-ng on the relative importance of the two effects (Li et al.,996; Tarvainen et al., 2002; Maulding et al., 1986; Bodmeiernd Chen, 1989). One or combinations of all above mentionedactors play a role in the degradation mechanism of polymers.n current research work, we have focused on effect of copoly-

er composition of PLGA on degradation behavior keeping allther factors constant.

Many researchers have documented degradation studiesor PLGA in the form of microspheres and films but as perhe literature study, degradation behavior of any polymer fromoated medical device such as coronary stents is not investi-ated till date (Engineer et al., in press). The objective of thisesearch work was to study the effect of copolymer ratio onydrolytic degradation behavior of poly(lactide-co-glycolide)

ncorporated in drug eluting stent. Effect of copolymer com-osition on in vitro hydrolytic degradation of PLGA has beenystematically investigated up to 12 months by incubatinghe drug–polymer coated stents in phosphate buffer saline at7 ◦C and 120 rpm. Surface morphology, mass loss, molecu-ar weight reduction and thermal changes were analyzed atarious time-intervals as an indicator of polymer degradation.

. Materials and methods

.1. Materials

he coronary stents (Sahajanand Medical Technologies, India)sed in the research work were laser-cut from 316L stainlessteel tubes. Poly(lactide-co-glycolide) (PLGA) having different

sign 8 9 ( 2 0 1 1 ) 328–334 329

lactide/glycolide ratios (50:50 PLGA with Mw 135 kDa, 65:35PLGA with Mw 118 kDa and 75:25 PLGA with Mw 120 kDa)were procured from Lakeshore Biomaterials, USA. Poly vinylpyrrolidone (PVP) (Mw of 1300 kDa) was procured from ISPTechnologies Inc., USA. Paclitaxel drug was procured fromBioxel Pharma Inc., Canada. The solvent dichloromethane(DCM) and other chemicals used in the current research workwere of HPLC grade procured from Ranbaxy Fine ChemicalsLtd., India. Nitrogen gas (98% pure) was used as a carrier gasfor drug coating. All drug and polymers were stored in a closedair-tight container under an inert atmosphere before use.

2.2. Methods

2.2.1. Drug coatingPaclitaxel drug, poly(lactide-co-glycolide), and poly vinylpyrrolidone were accurately weighed and dissolved in HPLCgrade dichloromethane to prepare drug coating solution asrepresented in Table 1. Three different solutions were pre-pared for Group-I (50:50 PLGA), Group-II (65:35 PLGA) andGroup-III (75:25 PLGA). Stents were weighed using analyticalbalance (Citizen CX-265) having 0.01 mg accuracy.

In current research work, modified air suspension coatingtechnique was utilized for drug coating on coronary stents(Raval et al., 2007; Kothwala et al., 2006). Stents of all threestudy groups were spray-coated with drug–polymer solutionand vacuum dried for 1 h for solvent evaporation from coat-ing film. The coated stents were weighed again to analyze thetotal amount of drug and polymer on the stent.

2.2.2. Drug releaseThree stents of each study group were placed in 10 ml ofphosphate buffer saline (PBS, pH 7.4) at 37 ◦C in standardmeasuring flasks with constant shaking at 120 rpm. The dis-solution medium was replaced daily with fresh PBS solutionand PBS obtained from the flask was analyzed for the amountof drug released from the stent. Paclitaxel drug was extractedfrom the release medium using dichloromethane which waslater evaporated using dry nitrogen gas. The residue was dis-solved in the mobile phase (acetonitrile 60%, v/v, methanol5%, v/v and water 35%, v/v) and the resultant solution wasanalyzed for Paclitaxel content by High Performance LiquidChromatography (HPLC-LC-2010 AHT, Shimadzu).

2.2.3. Polymer degradationDrug eluting stents from all three study groups were placedin individual flasks having 10 ml of PBS (pH 7.4) at 37 ◦C withconstant shaking at 120 rpm to facilitate hydrolytic degrada-tion. PBS was replaced daily. Stents were taken out at specifictime, washed with HPLC water and dried at 37 ◦C in oven for12 h for further analysis to investigate polymer degradation.

2.2.3.1. Gravimetric analysis. Initial weight (Wi) of each drugeluting stents was recorded before placing the stent in PBS.After each time-interval of incubation in PBS; each stentwas weighed gravimetrically for final weight (Wd). Weight ofuncoated stent (Ws) was deducted as common factor.

Mass loss (%) was calculated by following equation:

Mass loss (%) = (Wi − Ws) − (Wd − Ws)Wi − Ws

× 100

2.2.3.2. Gel permeation chromatography (GPC). Because theamount of drug and polymer present on single stent is very

330 chemical engineering research and design 8 9 ( 2 0 1 1 ) 328–334

Table 1 – Drug–polymer formulations for stent coating.

Parameters Group-1 Group-2 Group-3

Drug Paclitaxel (33%, w/w) Paclitaxel (33%,w/w)

Paclitaxel (33%,w/w)

Polymer 50:50 PLGA (60%, w/w) and PVP (7%, w/w) 65:35 PLGA (60%,w/w) and PVP(7%, w/w)

75:25 PLGA (60%,w/w) and PVP(7%, w/w)

Solvent Dichloromethane Dichloromethane DichloromethaneDrug dose 1.4 �g/mm2 1.4 �g/mm2 1.4 �g/mm2

33:67 33:67

Fig. 1 – Drug release from all groups of stents at varioustime-intervals.

Drug:polymer Ratio 33:67

less for detection by GPC, drug and polymer was extractedfrom 10 stents at each time-interval and extract of 10 stentswas analyzed for GPC. Drug and polymer were extracted bydissolving the stents in DCM which was later evaporatedusing dry nitrogen gas. Molecular weight of the test sam-ple was determined by gel permeation chromatography (GPC)equipped with a differential refractive index detector (RID-10AShimadzu) and a column (PLgel, 5 �m, Agilent) maintained at40 ◦C. Degassed tetrahydrofuran (THF) was used as the mobilephase at a flow rate of 1 ml/min. Molecular weight averages oftest sample were determined relative to polystyrene standardswith molecular weights ranging from 162 to 5,000,000 g/mol(Polystyrene Easycal Vial, Agilent).

2.2.3.3. Differential scanning calorimetry (DSC). Because theamount of drug and polymer present on single stent is veryless to detect by DSC instrument, drug and polymer residuewas extracted from 10 stents at each time-interval and theextract was analyzed for measurement of glass transition tem-perature (Tg). Drug and polymer were extracted from drugeluting stents by dissolving the stents in DCM which was laterevaporated using dry nitrogen gas. DSC experiments were per-formed using a Pyris-1 DSC from Perkin Elmer. Test sampleswere placed in a standard aluminum pan. The glass transitiontemperature was then measured during a heating cycle wherethe sample was heated from −10 ◦C to 150 ◦C with a heatingrate of 10 ◦C/min.

2.2.3.4. Scanning electron microscopy (SEM). Scanning elec-tron microscopy (model XL-30 ESEM, Philips, Netherlands) wasperformed to examine the surface morphology of drug elutingstents after each time-interval during degradation. No gold-sputtering was performed on these stents.

3. Results and discussion

3.1. Drug release

Fig. 1 represents the amount of Paclitaxel release, as a fractionof total drug content against the immersion time in phosphatebuffer saline. The in vitro drug release follows biphasic releasepattern, i.e., a burst phase and a diffusive phase. In all threegroups of drug eluting stents; average 59% (n = 3 samples pergroup, SD = 4.6) of Paclitaxel was released cumulatively withinfirst 4 days of incubation in PBS. It is believed that the highinitial burst was due to surface dissolution of drug Paclitaxel(Chen and Ooi, 2008). Upon immersion into the PBS, the drugwhich was loosely bounded on the surface or embedded into the coating layer on the stents came in the contact with

water and was released into the surrounding medium. It wasobserved that rate of Paclitaxel release in the burst phasedepends on the rate of water uptake, and was affected by

the composition of polymer (Chen and Ooi, 2008). 50:50 PLGAwhich is more amorphous; provides high burst release com-pared to other two groups. In the second phase, remaining10% drug was released after 60 days of immersion in PBS.Then after no drug release was detected. During this diffusionphase, the drug was migrated from the core to the surfaceof the polymer due to the concentration gradient and wasslowly released from the polymer (Chen and Ooi, 2008; Hurrelland Cameron, 2001). No remarkable difference in drug releasebetween three study groups was observed in this phase prob-ably because of less amount of drug present after burst phase.

Due to the short-term nature of the drug release observedin current study, the principal mechanism of drug release isconsidered as drug dissolution and diffusion rather than poly-mer degradation. The high level of drug release during theburst period was observed as the polymer was in the formof thin film having large surface area which may account fora high proportion of the un-dissolved surface-connected drugparticles (Lewis, 1990).

Fig. 2 – Mass loss from all groups of stents as a function ofdegradation time.

chemical engineering research and design 8 9 ( 2 0 1 1 ) 328–334 331

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ig. 3 – Molecular weight reduction of PLGA at variousegradation times.

.2. Mass loss

ig. 2 plots the changes in mass loss of the drug elutingtents as a function of hydrolytic degradation time. In an aver-ge 37% mass loss (n = 3 samples per group, SD = 1.13) wasbserved among all three study groups after 1 month of incu-ation in PBS. This mass loss in this period may occurred dueo release of drug Paclitaxel and PVP in surrounding mediaecause composition of drug coating comprised 33% (w/w)f Paclitaxel drug and 7% (w/w) of hydrophilic PVP (Table 1)nd it was observed that major amount of drug and PVPere released within this time period. As degradation pro-

eeded, rapid mass loss was observed in all three groups;robably due to release of water-soluble oligomeric fragmentsnd monomer products into the degrading medium (Pitt etl., 1981). The results show that mass loss was fastest in 50:50LGA and slowest in 75:25 PLGA, although the difference isot much significant. Degradation studies of PLGA in the formf microspheres also show the similar kind of effect of poly-er composition on mass loss (Pitt et al., 1981; Blanco et al.,

006).

.3. Molecular weight reduction

hen the degradation of the polymers was examined, theolecular weight profile of PLGA was found to decrease with

ime (Fig. 3). It can be seen that 50:50 PLGA degrade much fasterhan 65:35 PLGA and 75:25 PLGA. As more glycolic acid unitsre incorporated into the polymer, the chain scission reactionan occur more readily at the abundant G–G and L–G linkagesn the polymer backbone (Park, 1995).

In first month of degradation, even though mass loss wasegligible, the Mn of the polymers decreased rapidly withegradation time. This corresponded to the bulk degradationechanism. As water absorbed into the polymer, it promote

ydrolytic chain scission on the linkage of ester bonds in theolymer backbone and the molecular weight decreased withegradation time. When the size of the polymer fragments wasmall enough and the polymer surface became more porous,egradation products could escape out from the matrix andissolved in the incubation medium (Pitt et al., 1981; Alexis etl., 2006). Thus, the mass loss of the PLGA from DES began to bebserved only after 1 month of degradation. Studies reportedy Vey et al. (2008) show the similar kind of response for PLGAlms where rapid molecular weight reduction was observedrom 0 to 9 days; from 9 to 17 days, molecular weight reductionas slowed down and significant mass loss was observed.

The polydispersity was found to decrease from 1.51 to 1.25fter 3 months for 50:50 PLGA. For 65:35 PLGA, polydispersityas decreased from 1.49 to 1.20 and for 75:25 PLGA, it reaches

Fig. 4 – Change in Tg of PLGA at various degradation times.

from 1.47 to 1.27 after 3 months. After 3 months, no signif-icant decrease in polydispersity was observed with respectto copolymer ratio (average reduction among groups = 16.8%,SD = 2.96). Similar kind of response for PLGA implants wasreported by Chlopek et al. (2009) in his work. The narrowingof the molecular weight distribution was due to the simulta-neous disappearance of the high molecular weight compounddue to chain scission events and of the low molecular weightcompounds as they diffuse out of the films into the medium(Vey et al., 2008).

3.4. Thermal changes

Fig. 4 plots the changes in glass transition temperature withrespect to the degradation time. The Tg of PLGA decreasedwith degradation time for all the groups. The decrease in Tgwith degradation time was occurred probably due to the loss ofamorphous material arising from hydrolytic scission of poly-mer chains (Chen and Ooi, 2008). Since the polymer used wereamorphous in nature, water was able to penetrate easily intothe polymer matrix because of the open nature and looselypacked chains resulting in decrease in Tg due to plasticizationeffect of the amorphous polymer. Faster decrease in Tg after 3months in 50:50 PLGA (% reduction = 52%) compared to 75:25PLGA (% reduction = 30%) is probably due to more amorphousand hydrophilic nature of 50:50 PLGA, which may increasewater uptake.

3.5. Surface characterization

Scanning electron micrographs of the surface of drug elutingstents coated with 50:50 PLGA, 65:35 PLGA and 75:25 PLGAtaken at different degradation times are shown in Figs. 5–7respectively. Initially, the surface of drug eluting stent wassmooth without any cracks or delamination of coating layer(Figs. 5–7(a)). Subsequently after immersion of stents in phos-phate buffer saline, small wrinkles were appeared on thesurface (Figs. 5–7(b)). With time, these wrinkles grew in sizeleading to an increasingly porous surface (Figs. 6 and 7(c)).These pores might be generated due to diffusion of drug andhydrophilic poly vinyl pyrrolidone form the coating film andpenetration of water in the film. Study reported by Vey et al.(2008) has also shown such degradation morphology of PLGAfilms.

After certain time of degradation, cracks were generatedon the coated stent surface providing way for diffusion ofoligomers from the bulk of the polymer (Figs. 5(c), 6(d) and 7(d))(Pitt et al., 1981). Subsequent mass loss was observed in gravi-

metric analysis has supported such phenomena. After 90 days,significant polymer degradation was observed. The coatingfilm became fragile and polymer-free surface was appeared

332 chemical engineering research and design 8 9 ( 2 0 1 1 ) 328–334

Fig. 5 – SEM images of Group-I DES (a) before degradation, (b) after 1 month, (c) after 3 months, (d) after 6 months, (e) after 9months, and (f) after 12 months.

b) af

Fig. 6 – SEM images of Group-II DES (a) before degradation, (months, and (f) after 12 months.

at most of the locations of the stent. No trace of polymer wasobserved on stent surface of first groups at 9 months (Fig. 5(e))and in other two groups at 12 months (Figs. 6(f) and 7(f)). Thissurface characterization supports the overall observation of

faster rate of degradation for 50:50 PLGA compared to 65:35PLGA and 75:25 PLGA.

Fig. 7 – SEM images of Group-III DES (a) before degradation, (b) a9 months, and (f) after 12 months.

ter 1 month, (c) after 3 months, (d) after 6 months, (e) after 9

The observations from current study show that althoughthe molecular weight of PLGA decreased immediately uponcontact with water, the mass loss did not start until a criticalmolecular weight of the polymer was reached. This finding

supports the bulk degradation mechanism of PLGA from thedrug eluting stent. The data of GPC profiles suggesting the

fter 1 month, (c) after 3 months, (d) after 6 months, (e) after

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apid decrease in molecular weight of PLGA does not nec-ssarily mean that their degradation products immediatelyiffuse out into the aqueous medium. This may be due to the

ow-molecular-weight degradation products which were gen-rated because of hydrolysis of polymers were retained withinhe polymer bulk, and are then released out at a slow rate.his observation supports the existence of the diffusion bar-ier layer at the surface, which controls the release rate of theegradation products (Park, 1995).

The degradation process of PLGA coated on drug elutingtents of all three groups occurred in two stages. During firsttage, molecular weight reduction of polymer was rapid dueo hydrolysis of polymer backbone. Mass loss was very lessuring this stage. Glass transition temperature was reducedith degradation time due to penetration of water in poly-er bulk creating plasticization effect. Polymer degradationas not apparent on coated stent surface during first stage ofegradation as only small pores were developed on the sur-ace. During second stage, molecular weight reduction waslow and rapid mass loss was observed due to diffusion ofegradation products in the media. Reduction of glass transi-ion temperature was not significant in second stage. Surfaceharacteristics changed drastically during this stage of degra-ation. Small cracks were developed on the surface allowingligomers and small fragments to come out from the polymerulk (Pitt et al., 1981). Coating film became fragile and polymeras disappeared at the end of the second stage.

The degradation of PLGA involves chain scissions of esterond linkages in the polymer backbone by hydrolytic attackf water molecules. The difference in degradation behavioretween different copolymer compositions of PLGA can bexplained as follows. Poly(d,l-lactide-co-glycolide) is an amor-hous polymer. Changes in the hydrophilic/lipophilic balances well as steric effects are responsible for the changes inhe degradation rate of PLGA. This polymer built from two

onomers GA and LA and contains four types of bonds:A–GA, LA–LA, GA–LA and LA–GA. These bonds have differ-nt hydrolysis rates. The GA units are more hydrophilic thanhe LA units. The more GA content in 50:50 PLGA compared to5:35 and 75:25 PLGA provide more GA–GA units in the poly-er backbone, thus making it more hydrophilic. This higher

ydrophilicity results in a higher water absorbing capacity ofpolymer which in turn will increase the hydrolysis rate of

he ester bonds. PLGA polymers containing 50:50 ratio of lac-ic and glycolic acids have been reported to hydrolyze muchaster than those containing higher proportion of either of thewo monomers (Heller and Hoffman, 2004; Brannon-Peppasnd Vert, 2000). Similar results were observed in present study,here 50:50 PLGA degraded fastest, followed by 65:35 PLGA

nd lastly, 75:25 PLGA.Additionally, the degradation experiment in this study was

erformed in phosphate buffer saline having 7.4 pH. To mimiche actual clinical situation, the stents were incubated at 37 ◦Cnd 120 rpm. It has been demonstrated that in vitro and inivo degradation profiles of PLGA were similar in terms ofolecular weight decrease (Kenley et al., 1987). However, in

ivo degradation is difficult to predict due to the complicatedature, which involves not only tissue interaction, but alsoerhaps enzyme participation in degradation (Therin et al.,992). Menei et al. (1993) had observed faster degradation ofLGA microspheres in vivo versus in vitro system and attributed

his to lipids or other biological compounds present in vivocting as plasticizers favoring the uptake of water into theolymer. In addition, the faster degradation in vivo may be due

sign 8 9 ( 2 0 1 1 ) 328–334 333

in part to the foreign body response (Ali et al., 1994; Williamsand Mort, 1977; Williams, 1979; Tokiwa and Suzuki, 1977).This response results in the accumulation of cells such asmacrophages around the foreign body leading to a walling-offthe region. Free radicals, acidic products, or enzymes pro-duced by these cells during the foreign body response mayaccelerate polymer degradation. Study of in vivo degradationbehavior of PLGA from drug eluting stents and effect of copoly-mer composition on in vivo degradation is the future scope ofcurrent research work.

4. Conclusion

The degradation process of PLGA coated on drug eluting stentsof all three groups occurred in two stages. During first stagemolecular weight reduction was rapid with little mass losswhile in second stage molecular weight reduction was sloweddown and significant mass loss was observed due to the diffu-sion of degradation products in to the surrounding media. Asthe glycolic acid ratio in the copolymer increases, the rate ofdegradation increases. This is due to the glycolic acid inher-ent high reactivity with water and its greater hydrophilicitycompared with lactic acid, which will increase water uptakeand, thus, the rate of degradation. The drug release fromPLGA coated stent followed biphasic pattern which was gov-erned by drug dissolution and diffusion rather than polymerdegradation.

Acknowledgement

The authors express their sincere gratitude to SahajanandMedical Tech. Pvt. Ltd., India for providing facility to carry outthe research work.

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