Hydrolytic Degradation Behavior of 50/50 Poly Lactide-co ... - medIND

Hydrolytic Degradation Behavior Trends Biomater. Artif. Organs, Vol 24(3), pp 131-138 (2010) of 50/50 Poly Lactide-co-Glycolide

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Hydrolytic Degradation Behavior of 50/50 Poly Lactide-co-Glycolide from Drug Eluting Stents Chhaya Engineer1, Jigisha Parikh1* and Ankur Raval2 1

Department of Chemical Engineering, Sardar Vallabhbhai National Institute of Technology, Surat, India Sahajanand Medical Technologies Pvt. Ltd., Surat, India Corresponding author: Jigisha Parikh ([email protected]) 2

Received 8 February 2010; Accepted 27 April 2010 In-vitro degradation of 50/50 PLGA for programmed drug release has been systematically investigated from the drug eluting stents. Mass loss, molecular weight reduction, thermal changes and surface morphology of the drug eluting stents were analyzed as a function of degradation time. It was observed that degradation of 50/50 PLGA occurs in two phases. Quick decrease in molecular weight but little mass loss took place during first phase whereas in second phase, molecular weight reduction is slow with significant mass loss. This phenomenon supports heterogeneous degradation mechanism of PLGA. Drug release was attributed to diffusion rather then polymer degradation.

Introduction Biodegradable polymers have been extensively studied throughout the last few decades. The biodegradable polymer can be broken down and the degradation by-products eliminated from the body [1]. Aliphatic polyesters have found prominence among synthetic biodegradable polymers for implants, the most extensively studied amongst this family of polymers are poly(L-lactide), poly(glycolide), poly(ecaprolactone) and copolymers based on L/DLlactide, glycolide, trimethyl carbonate and ecaprolactone [2]. Polymers with ester linkages in their main chain are ideal candidates for a range of temporary biomedical applications, as the need for surgical removal of the depleted device is eliminated. This has been utilized in the preparation of controlled drug release systems, in sutures and in orthopaedic implants [3]. Controlled release of therapeutic agents remains one of the biggest challenges in drug delivery. Repeated administration of a drug so as to maintain drug concentration within a therapeutic window may cause serious side

effects [4]. With conventional dosage forms, high peak blood concentrations may be reached soon after administration with possible adverse effects related to the transiently high concentration. Suitable drug delivery candidates must therefore not only be biodegradable and biocompatible, but must also exhibit control over the release rate The family of aliphatic polyesters has been by far the dominating choice for materials in degradable drug delivery systems [5]. Many researchers have documented the biodegradation, biocompatibility and tissue reaction associated with poly(lactide) and poly(lactide-co-glycolide) [6, 7]. Despite the growing use of biodegradable polymers, there are still many unsolved problem that hinder to take full advantage of this materials. One example is lack of understanding of the mechanism of polymer degradation which controls the essential processes; like release of drug from the drug delivery system. There are number of factors that affect the biodegradation of polymers such as chemical

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C. Engineer, J. Parikh and A. Raval

structure, molecular weight and its distribution, morphology, processing condition, sterilization, geometry of polymeric device, site of application, etc [8]. One or combination of all these factors play role in the degradation mechanism of polymers.

weighed and dissolved in HPLC grade dichloromethane to prepare coating solution as represented in Table 1. Stents were weighed using analytical balance (Citizen CX-265) having 0.01 mg accuracy. Drug Coating Technique

Various studies are based on the degradation of PLGA microspheres and films but as per our knowledge, however, literature study reveals that degradation behavior of any polymer from coated medical device such as cardiovascular stents is not investigated till date [9]. The objective of this research work was to study the degradation behavior of 50/50 poly lactide-coglycolide incorporated in drug eluting stent. 50/ 50 PLGA in-vitro degradation has been systematically investigated up to 3 months by incubating the drug-polymer coated stents in phosphate buffer saline at 37°C and 120 rpm. Surface morphology, mass loss, molecular weight reduction and thermal changes were determined at various time intervals to study the polymer degradation. Mechanism of degradation for 50/50 PLGA incorporated with Paclitaxel drug and spray-coated on cardiovascular stents was also investigated either it follows homogeneous degradation or heterogeneous degradation. Materials and Methods The coronary stents (Sahajanand Medical Technologies, India) used in the research work were laser-cut from 316L stainless steel tubes. 50/50 Poly lactide-co-glycolide (Mw of 135 KD) was procured from Lakeshore Biomaterials, USA. Poly vinyl pyrrolidone (PVP) (Mw of 1300 KD) was procured from ISP Technologies Inc., USA. Paclitaxel drug was procured from Bioxel Pharma Inc., Canada. The solvent dichloromethane (DCM) and other chemicals used in the current project were of HPLC grade procured from Ranbaxy Fine Chemicals Ltd, India. Nitrogen gas (98% pure) was used as a carrier gas for drug coating. All drug and polymers were stored in a closed air-tight contained under an inert atmosphere before use. Drug Coating Formulation Paclitaxel drug, 50/50 Poly Lactide-co-Glycolide, and Poly vinyl pyrrolidone were accurately

In the current research work, modified air suspension coating technique was utilized for drug coating on coronary stents [10, 11]. Stents were coated with drug-polymer solution and weighed again to analyze the total amount of drug and polymer coated on the stent. Drug Release Drug eluting stents were placed in the phosphate buffered saline (PBS, pH 7.4) at 37°C in a glass vial with constant shaking at 120 rpm. The dissolution medium was replaced daily with fresh PBS solution and the PBS obtained from the glass vials was analyzed for the amount of drug released. Drug was extracted from the release medium using Dichloromethane which was later evaporated using dry nitrogen gas. The residue was dissolved in the mobile phase (Acetonitrile60% v/v, Methanol 5%v/v and Water-35%v/v) and the resultant solution was analyzed for Paclitaxel content by HPLC (HPLC-LC-2010 AHT, Shimadzu0. Polymer Degradation Drug eluting stents were placed in phosphate buffer saline (PBS, pH 7.4) at 37°C in a glass vial with constant shaking at 120 rpm. PBS was replaced daily. Stents were taken out at specific time, washed with HPLC water and dried at 37°C for further analysis to investigate polymer degradation. Gravimetric Analysis: Initial weight (Wi) of coated stents were recorded before placing the

Table 1: Composition of Drug Coating Solution Material

Weight %

Paclitaxel

33 %

Poly Lactide-co-Glycolide

60%

Poly Vinyl Pyrrolidone

7%

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Hydrolytic Degradation Behavior of 50/50 Poly Lactide-co-Glycolide

Mass loss (ML %) was calculated by following equation:

% Paclitaxel Re lease 90 80 % Drug R elease

stent in PBS. At specific interval of incubation in PBS and after drying at 37°C; each stent was weighed gravimetrically for final weight (Wd). Weight of uncoated stent (Ws) was deducted as common factor.

70 60 50 40 30 20 10

ML (%) = [(Wi-Ws) - Wd-Ws)]/(Wi-Ws) x 100

0 0

Differential Scanning Calorimetry (DSC): DSC experiments were performed using a Pyris-1 DSC from Perkin Elmer. Drug and Polymer were extracted from drug eluting stents by dissolving the stents in DCM which was later evaporated using dry nitrogen gas. Samples were placed in a standard aluminum pan. The glass transition was then measured during a heating cycle where the sample was heated from -10° C to 150° C at 10°C/min. Scanning Electron Microscopy (SEM): Scanning electron microscopy (model XL-30 ESEM, Philips, Netherlands) was performed on the drug eluting stents to examine the morphology of drug coated stents before and during degradation. Results and Discussion Figure 1 represents the amount of drug release, as a fraction of the total drug loading against the immersion time in the release medium. Release of drug from the polymeric matrix occurred in two phases. 60-70% of drug was released within 1 week of incubation in PBS. During second phase, remaining 10% drug was released during next 7 weeks. Then after no

15

30

45

60

75

Time (Days)

Figure 1: Drug Release from of 50/50 PLGA at various time interval

% Mass Loss Mass Loss (%)

Gel Permeation Chromatography (GPC): Drug and Polymer were extracted from drug eluting stents by dissolving the stents in DCM which was later evaporated using dry nitrogen gas. Molecular weight of the polymer sample was determined by gel permeation chromatography (GPC) equipped with a differential refractive index detector (RID-10A Shimdzu) and a column (PLgel, 5 µm, Agilent) maintained at 40° C. Degassed THF was used as the mobile phase at a flow rate of 1 ml/min. Sample molecular weight averages were determined relative to polystyrene standards with molecular weights ranging from 162 to 50,00,000 g/mol (Polystyrene Easycal Vial, Aligent).

100 90 80 70 60 50 40 30 20 10 0 0

15

30

45

60

75

90

105

Time (Days)

Figure 2: Mass Loss of 50/50 PLGA at various degradation time

drug release was detected. Studies shows that degradation period of PLGA is 4-6 weeks [5, 7], which indicate that drug release in current research work was diffusion and dissolutioncontrolled rather than degradation controlled. Initial high drug release rate stage was attributed to the release of un-dissolved surfaceconnected drug particles [5], followed by a slow release stage attributed to molecular diffusion through the polymer phase [12]. In-vitro degradation study of the 50/50 PLGA coated on cardiovascular stent showed negligible mass loss during 15 days (5 % and 10 % after 7 and 15 days respectively). After 30 and 90 days of incubation in PBS, 56 % and 85 % mass loss was observed (Figure 2). Figure 3 shows the molecular weight reduction of 50/50 PLGA at different degradation time determined by GPC. Number average molecular weight (Mn) decreased immediately after the


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Roger [14] shown the similar kind of response of PLGA films where significant mass loss was observed after 9 days whereas rapid molecular weight reduction was observed from 0 to 9 days and from 9 to 17 days; molecular weight reduction was slowed down.

Mn Reduction (%)

Mn Reduction (%) 100 90 80 70 60 50 40 30 20 10 0 0

15

30

45

60

75

90

105

Time (Days)

Figure 3: Molecular weight reduction of 50/50 PLGA at various degradation time 0 0

15

30

45

60

75

90

105

ln Mnt/Mn0

-0.5 -1

The polydispersity was found to increase from 1.45 to 1.8 after 15 days. From day 15 onwards, the polydispersity was found to decrease again and after 90 days the polydispersity was found to be 1.4. Similar kind of response for PLGA implants were reported by Chlopek in his work [15]. The narrowing of the molecular weight distribution was due to the simultaneous disappearance of the high molecular weight compound due to chain scission events and of the low molecular weight compounds as they diffuse out of the films into the medium [14].

-1.5 -2 -2.5 Time (Days)

Figure 4: Logarithmic Plot of Mn t/Mn 0 versus time

stents were placed in contact with the media; from 0 to 7 days of degradation Mn decreased from 67,416 g mol-1 to 23,470 g mol-1 (65.52 % reduction). After 15 days, Mn was 13,426 g mol1 (80 % reduction). After 15 days, the decrease in number average molecular weight slowed down significantly from 12,632 g mol-1 (81.26 % reduction) at 30 days to 8,562 g mol-1 (87.3 % reduction) at 90 days. Then after polymer residue was in very less amount which was not able to detect by GPC. The degradation process of the PLGA began with a decrease of molecular weight as a consequence of the hydrolysis of ester bonds of the polymers, which originated smaller polymer chains [13]. When the size of the polymer fragments became small enough and the polymer surface became more porous, degradation products could escape from the matrix and dissolved in the incubation medium. Thus, the mass loss of the 50/50 PLGA from DES began to be observed only after 15 days when a critical value of molecular weight (Mn), reach 13,500. Studies reported by Vey and

Figure 4 shows the logarithmic plots of ratio (molecular weight at time t/molecular weight at initial time) versus time for 50/50 PLGA degradation from stents The phases of the degradation process and their duration were shown, and the degradation rate constants (k) and half-life times (t1/2) were obtained from the prout-tompkins equation and represented in Table 2.

From the above results, it is revealed that 50/50 PLGA degradation from coated stents was preceded in two stages. During stage-I, the molecular weight was decreased rapidly with little mass loss. In stage-II which began after day 15, the decrease in molecular weight was slowed down and severe mass loss was observed. In contrast to the work done by Blanco and Sastre [16] for 50/50 PLGA microspheres, where 1st phase of degradation was observed for 22 days and degradation was slow with a rate constant of 0.014 day-1 compare to the 2nd phase (0.085 day-1); in our case, 1st phase of degradation was observed for 15 days and degradation was fast with rate constant of 0.1064 day-1 compare to 2nd phase (0.0061 day1 ). This difference might be due to various factors such as processing techniques, geometry and physicochemical properties of 50/50 PLGA used in both cases.

Hydrolytic Degradation Behavior of 50/50 Poly Lactide-co-Glycolide Table 2: Degradation Rate Constant and Halflife Data Phase

I

II

Time

0-15

16-90

Degradation Rate Constant (k, day- )

0.1064

0.0061

Half-life (t1/2, days)

6.5

113.6

Regression Co-efficient (r2)

0.9526

0.9955

1

It is well known that the glass transition (Tg) of PLGA is a function of polymer molecular weight. The glass transition temperature is a reversible step change in molecular mobility, from a rigid

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glassy state to a mobile rubbery state. This change is linked to the free volume associated to the polymer chain ends. When the molecular weight decreases, more chain ends are present, therefore, more free volume is generated. The polymer chains have, therefore, more space to move and the rubbery state is reached earlier, i.e. at a lower temperature [17]. The glass transition temperature of the coated stents decreased as soon as the stents were placed in contact with the media. Tg shifted from 45.7 °C to 13.5 °C in the first 15 days. At this early stage of degradation, the decrease in Tg might be due to the decrease in molecular weight of

Figure 5: SEM images of drug coated stents (a) before degradation, (b) after 7 days, (c) after 15 days, (d) after 30 days, (e) and (f) after 90 days

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the polymer. Also, the water diffused into the polymer lead to the decrease of the glass transition of the polymer due to the plasticizing effect of the water molecules [18]. From day 16 onwards, the Tg was found to increase again reaching 32.1°C after 30 days. This increase in Tg was unexpected as the molecular weight of the polymer did not increase but remained constant. Studies reported by Alexis and Hakkarainen suggest that glycolic units degrade preferentially [13, 19]. This effect could be at the origin of the increase in Tg as the glass transition of Poly lactide (55–60°C) is higher than the Tg of Poly glycolide (35–40°C) [20]. Scanning electron micrographs of the surface of drug coated stents taken at different degradation times are shown in Figure 5 (a, b, c, d, e and f). Subsequently after immersion of stents in PBS, small wrinkles appeared on the surface over the first 7 days of degradation. With time, these wrinkles grew in size leading to an increasingly porous surface (Figure 5 (c)) after 15 days. These pores might be generated due to diffusion of drug and hydrophilic polyvinyl pyrrolidone form the coating film and penetration of water in the film. Vey and Roger has also reported similar degradation morphology of PLGA films [14]. After 30 days of degradation, cracks were generated on the coated stent surface providing way for diffusion of oligomers from the bulk of the polymer. Subsequent mass loss after 30 days was observed during gravimetric analysis which supports this phenomenon. After 90 days, significant polymer degradation was observed. The coating film became fragile and polymerfree surface was appeared at most of the locations of the stent. This overall behavior indicates degradation of 50/50 PLGA through chain scission which took place at a faster rate in the bulk of the polymer. Occurrence of such kind of heterogeneous degradation in PLGA has been reported in number of studies [12, 21-25]. The difference between the surface and the bulk of the polymer matrices during degradation is thought to be due to the autocatalysis of the degradation reaction in the bulk of the polymers [12]. At the early stages of the degradation process when

hydrolysis occurs in the bulk of the polymer, the degradation product (lower molecular weight polymer chains with carboxylic end groups) remains entrapped inside the polymer matrices. As a consequence, the concentration of terminal acid groups increases rapidly and the environment becomes highly acidic resulting in the autocatalysis of the PLGA degradation reaction [19-23]. On the other hand when hydrolysis occurs in the surface of the polymer, the degradation products can diffuse more easily out of the surface and this surface layer contains a relatively large amount of buffer solution which results in the neutralisation of terminal acid groups formed [24, 25]. These two combined effects result in a much less acidic environment and therefore significant less autocatalysis reaction at surface. In present study, 50/50 PLGA coated on cardiovascular stents follows bulk erosion. Moreover, degradation period for 50/50 PLGA from stents was observed as 3 months and 70% of drug was released within first week of incubation. This initial burst release of drug might be caused due to combined process of relatively faster drug diffusion than polymer degradation of the matrices [26, 27]. Observations found in the current study suggests that research work for drug release from surface eroding polymers will provide more insights to compare drug release mechanism from surface eroding versus bulk eroding polymers. Conclusion The hydrolytic degradation process of 50/50 PLGA from the drug eluting stent took place in two phases. During the first stage of degradation, significant mass loss was not observed, and the fast decrease in molecular weight took place. During second phase, molecular weight decreased at slower rate and significant mass loss observed as the critical molecular weight was achieved during this stage. The heterogeneous degradation mechanism of PLGA is confirmed from this research work. Drug release from stents was mainly governed by diffusion rather then degradation.

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Acknowledgement The authors express their sincere gratitude to Sahajanand Medical Tech. Pvt. Ltd., India for providing facility to carry out the research work. References 1. M. Borden, Wiley encyclopedia of biomedical engineering, Wiley: New York (2006). 2. R. Lanza, R. Langer and J. Vacanti, Principles of tissue engineering, Elsevier Academic Press: New York, 1265-1270 (2007). 3. P.A. Gunatillake and R. Adhikari, “Biodegradable Synthetic Polymers for Tissue Engineering”, Eur Cells Mater, 5:1-16(2003). 4. L. Geever, D. Devine, M. Nugent, J. Kennedy, J. Lyons and C. Higginbotham, Lower critical solution temperature control and swelling behaviour of physically crosslinked thermosensitive copolymers based on N-isopropylacrylamide”, Eur Polym J, 42, 2540-2548 (2006). 5. D.H. Lewis, In Biodegradable Polymers as Drug Delivery Systems, (Ed.) M. Chasin and R. Langer, Marcel Dekker: New York, 1-43 (1990). 6. U. Edlund and A. Albertsson, “Degradable polymer microspheres for controlled drug delivery”. Adv. Polym. Sci. 157, 67-112 (2002). 7. R.L. Kronenthal, In Polymers in Medicine and Surgery, (Ed.) R.L. Kronenthal, Z. Oser and E. Martin, Plentum Press: New York, 119 (1975). 8. M. Vert, S. Li, and H. Garreau, “More about the degradation of LA/GA-derived matrices in aqueous media”. J. Controlled Release. 16, 15–26 (1991). 9. C. Engineer, C., J. Parikh and A. Raval, “Review on hydrolytic degradation behavior of biodegradable polymers from controlled drug delivery systems”, Trends Biomater. Artif. Organs, (2010) (in press). 10. A. Raval, A. Choubey, C. Engineer, H. Kotadia and D. Kothwala, “Novel Biodegradable polymeric matrix coated cardiovascular stent for controlled drug delivery”, Trends Biomater. Artif. Organs. 20 (2), 101-110 (2007). 11. D. Kothwala, A. Raval, A. Choubey, C. Engineer and H. Kotadia, “Paclitaxel drug delivery from cardiovascular stents”, Trends Biomater. Artif. Organs. 19(2), 88-92 (2006). 12. S. Hurrell and R.E. Cameron, “Polyglycolide: degradation and drug release. Part I. Changes in morphology during degradation”, J. Mater. Sci. Mater. Med. 12(9), 811–816 (2001). 13. F. Alexis, S. Venkatraman, S.K. Rath and L.H. Gan, “Some insight into hydrolytic scission mechanisms in bioerodible polyesters”, J. Appl. Polym. Sci. 102(4), 3111–3117 (2006). 14. E. Vey, C. Roger, L. Meehan, J. Booth, M. Claybourn, A.F. Miller and A. Saiani, “Degradation mechanism of poly(lactic-co-glycolic) acid block copolymer cast films in phosphate buffer solution”, Polym. Degrad. Stab. 93, 1869–1876 (2008). 15. J. Chlopek, A. Morawska-Chochol, C. Paluszkiewicz, J. Jaworska, J. Kasperczyk and P. Dobrzynski, “FTIR and NMR study of poly(lactide-co-glycolide) and hydroxyapatite implant degradation under in vivo conditions”, Polym. Degrad. Stab. 94, 1479–1485 (2009). 16. M.D. Blanco, R.L. Sastre, C. Teijon, R. Olmo and J.M. Teijon, “Degradation behaviour of microspheres prepared by spray-drying poly(d,l-lactide) and poly(d,l-lactide-co-glycolide) polymers”, Int. J. Pharm. 326, 139–147 (2006). 17. R.J. Young and P.A. Lovell, Introduction to Polymers, CRC Press: London, (1991). 18. P. Blasi, S. D’Souza, F. Selmin and P. DeLuca, “Plasticizing effect of water on poly(lactide-co-glycolide)”, J. Controlled Release, 108, 1-9 (2005). 19. M. Hakkarainen, A.C. Albertsson and S. Karlsson, “Weight losses and molecular weight changes correlated with the evolution of hydroxyacids in simulated in vivo degradation of homo- and copolymers of PLA and PGA”, Polym. Degrad. Stab., 52(3), 283–291 (1996). 20. I. Engelberg and J. Kohn, “Physicomechanical properties of degradable polymers used in medical applications - a comparative study”, Biomaterials, 12(3), 292–304 (1991). 21. S.M. Li, H. Garreau and M. Vert, “Structure property relationships in the case of the degradation of massive aliphatic poly(alpha-hydroxy acids) in aqueous-medi; part 1: Poly(DL-lactic acid)”, J. Mater. Sci. Mater. Med., 1(3), 123–130 (1990). 22. F. Alexis, S. Venkatraman, S.K. Rath and L.H. Gan, “Some insight into hydrolytic scission mechanisms in bioerodible polyesters”, J. Appl. Polym. Sci., 102(4), 3111–3117 (2006).

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23. L.K. Chiu, W.J. Chiu and Y.L. Cheng, “Effects of polymer degradation on drug release – a mechanistic study of morphology and transport properties in 50:50 poly (D,L-lactide- co-glycolide)”, Int. J. Pharm., 126(1– 2),169–178 (1995). 24. E.A. Schmitt, D.R. Flanagan and R.J. Linhardt, “Degradation and release properties of pellets fabricated from 3 commercial poly (D,L-lactide-co-glycolide) biodegradable polymers”, J. Pharm. Sci.,. 82(3), 326– 329 (1993). 25. H.T. Wang, H. Palmer, R.J. Linhardt, D.R. Flanagan and E. Schmitt, “Degradation of poly(ester) microspheres”, Biomaterials, 11(9), 679–685 (1990). 26. M.J. Alonso, S. Cohen, T.G. Park, R.K. Gupta, G.R. Siber and R. Langer, “Determinants of release rate of tetanus vaccine from polyester microspheres”, Pharm. Res., 10, 945–953 (1993). 27. S. Cohen, T. Yoshioka, M. Lucarelli, L.H. Hwang, and R. Langer, “Controlled delivery systems for proteins based on poly-(lactic/glycolic acid) microspheres” Pharm. Res., 8, 713–720 (1991).