Self-curing controlled release systems for steroids. Application of prednisolone-based polymeric systems to ear diseases

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Biomaterials 26 (2005) 3311–3318 www.elsevier.com/locate/biomaterials

Self-curing controlled release systems for steroids. Application of prednisolone-based polymeric systems to ear diseases Mar Fernandeza, Juan Parrab, Blanca Vazqueza,, Antonio Lopez-Bravob, Julio San Roma´na a

Instituto de Ciencia y Tecnologı´a de Polı´meros, CSIC. C/Juan de la Cierva, 3, 28006 Madrid, Spain b Hospital Provincial de A´vila. C/Jesu´s del Gran Poder, s/n, A´vila, Spain Received 29 June 2004; accepted 8 September 2004

Abstract An injectable delivery system for prednisolone has been prepared based on a self-curing formulation comprised of poly(methyl methacrylate) particles and hydroxyethyl methacrylate as monomer. The polymerisation reaction was initiated by the redox system 4,40 -bis (dimethylaminobenzydrol)/benzoyl peroxide (BZN/BPO) and followed at 25 1C by measuring the time–temperature profile. A maximum temperature of 53 1C and a setting time of 15 min were obtained, calculated according to standard specifications. The swelling of the cured system was studied in phosphate-buffered saline (PBS) at 37 1C giving a hydration degree at equilibrium of 20%. The swelling kinetics fitted a fickian behaviour at the initial stages of the experiments, with a diffusion coefficient of 0.72  107 cm2/s. The release of the drug was sustained from the beginning without an initial drug burst. The study of the wettability showed a rather hydrophilic character of the surface of the loaded system, and the biocompatibility evaluated through MTT assay revealed the absence of cytotoxicity due to the release of toxic substances. r 2004 Elsevier Ltd. All rights reserved. Keywords: Drug delivery; PMMA/PHEMA; Prednisolone; Ear disease

1. Introduction Steroids are widely used to treat inner ear diseases [1]. Local administration of steroids has been applied in patients for whom systemic steroid treatment has failed or who could not tolerate systemic steroid therapy [2] and also as a way to avoid systemic toxicity. Transtympanic steroids have been applied through a ventilation tube placed with the patient under local anaesthesia [2] or by means of injection or osmotic minipumps [3]. High dose delivery of methyl prednisolone via perfusion at the level of the round window membrane has given good results compared with standard treatment of SSHL [4]. The use of this therapy has also been proposed for the treatment of tinnitus [5] or in cases of intractable Corresponding author. Tel.: 34 91 5622900; fax: 34 91 5644853.

E-mail address: [email protected] (B. Vazquez). 0142-9612/$ - see front matter r 2004 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2004.09.018

Meniere’s disease [6]. However, current methods do not allow for accurate drug delivery to the inner ear. The advantages of using a drug delivery system stem from the best control of the dose and lack of loss of the drug from the middle ear site. Delivery systems based on biodegradable sodium hyaluronate gels [7] or systems based on absorbable gelatin sponges soaked in the drug [6] have been applied in the treatment of vertigo and Meniere’s disease, respectively. The approach presented in this paper is the formulation of a corticoid self-curing delivery system which would be able to provide a sustained release of the drug over long period of time. Self-curing delivery systems are well documented in the literature. Different types of self-curing delivery systems are reported which have been mainly tested for the release of antibiotics. Totally, biodegradable systems of poly(propylene fumarate) [8], or partially biodegradable systems formulated with poly(methyl methacrylate)/

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poly(e-caprolactone) beads have been demonstrated to be effective in the release of vancomycin [9]. Likewise, a system incorporating b-tricalcium phosphate (b-TCP), PEMA and MMA has been developed for the controlled release of analgesic/anti-inflammatory drugs [10]. This paper reports on the preparation and characterisation of an injectable prednisolone delivery system based on a support of poly(methyl methacrylate)/ poly(hydroxyethyl methacrylate). The free radical polymerisation reaction was followed by measuring the exotherm of polymerisation from which the setting parameters were determined. Prednisolone release and swelling kinetics were evaluated in vitro in phosphatebuffered saline (PBS) at the physiological temperature. The wettability of the material was assessed by contact angle measurements and the cytotoxicity was determined using several assays that monitor different aspects of cellular activity.

2. Experimental 2.1. Materials Poly(methyl methacrylate) (PMMA) beads (33 mm of average diameter) were supplied by Industrias Quiru´rgicas de Levante (IQL beads) and have earlier been characterised [11]. 2-Hydroxyethyl methacrylate (HEMA) (Fluka), prednisolone (Sigma) and 4,40 -bis (dimethylaminobenzydrol) (BZN) (Sigma) were used as received. Benzoyl peroxide (BPO) (Fluka) was purified from fractional recrystallization from methanol, mp=104 1C. Thermanox (TMX) control discs were supplied by Labclinics S.L. and plasticware by Sarstedt. Tissue culture media, additives, trypsin, PBS and 3-(4,5dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) were supplied by Sigma. 2.2. Preparation and setting parameters of the self-curing delivery system

statically controlled bath described in a previous paper [11]. Curing parameters were calculated according to the international standard specification ISO 5833 [13]. 2.3. In vitro behaviour 2.3.1. Swelling behaviour Rectangular-shaped samples of 30 mm  10 mm and 1 mm thickness were immersed in PBS at 37 1C. Water uptake was determined gravimetrically at different periods of time. At appropriate times, the samples were removed, blotted quickly with absorbent paper to remove the water attached on its surface and weighed. In all the experiments, a minimum of three samples were measured and averaged. The percentage of hydration degree was calculated according to % hydration degree ¼ ½ðW w  W d Þ=W d   100;

(1)

where Ww is the weight of swollen specimen at time t and Wd is the initial weight of the dry specimen. 2.3.2. Prednisolone release experiments Rectangular-shaped samples similar to those used in swelling experiments were employed. The samples were soaked in 10 ml of PBS and kept at 37 1C. Aliquots were taken at different periods of time and the medium was totally changed by fresh solution. The concentration of prednisolone released was determined by visible–ultraviolet spectroscopy (UV-VIS, Perkin-Elmer Lambda 35) analysing the signal at 248 nm corresponding to prednisolone [14]. A calibrated curve of the glucocorticoid was obtained previously by measuring the absorption of the UV signal of solutions of known concentration in the same medium. A minimum of three samples were measured and averaged. The surfaces of PMMA/PHEMA/PRED samples before and after the release experiment were analysed by ESEM using a Philips XL 30 microscope. 2.4. Surface characteristics

The drug-loaded injectable system PMMA/PHEMA/ PRED was based on a mixture of 68.5 wt% PMMA beads and 30 wt% prednisolone with the inititator BPO (1.5 wt%) as the solid component, and 79.7 wt% 2hydroxyethyl methacrylate, 20 wt% water and 0.3 wt% of BZN as activator of reduced toxicity [12] as the liquid component. A solid:liquid ratio of 1:1 was used. Accordingly, a control formulation (PMMA/PHEMA) was prepared in the absence of the drug for comparison purposes. The exothermic polymerisation temperature profile was registered using a thermocouple connected to a high-sensitivity thermotester positioned within its junction in the centre of the mould at a height of 3 mm in the internal cavity. The mould (10 mm in diameter and 15 mm high) was placed in a thermo-

The contact angle measurements were performed on dry films of material using a Contact Angle Measuring System G10 (Kru¨ss). The surface free energy was calculated by the approach introduced by Fowkes, in which the total surface tension is considered as a sum of independent terms, each representing a particular intermolecular force [15], and by the application of the equation of Owens and Wendt [16], which is an extension to a so-called ‘‘polar’’ component. The liquids used for this purpose were methylene iodide and distilled water. The dispersion force component and the polar force component of the surface energy of water are 21 and 51 mN/m, respectively, and the dispersion force component of the methylene iodide is 50 mN/m.

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2.5. Evaluation of cytotoxicity of leachables from the cured materials 2.5.1. Specimens and in vitro cell culture for biocompatibility experiments Discs of 10 mm diameter and 1 mm thickness of the cured systems and the controls were used for direct and indirect biocompatibility experiments. All specimens were sterilised with poly(ethylene oxide). The negative control was tissue culture plastic, Thermanox, an international standard, and the positive control (toxic agent) was polyvinylchloride (PVC). The cells used in the primary cell culture were cells of African green monkey kidney (VERO) and were cultured at 37 1C. The culture medium was minimal essential medium eagle (MEM), modified with HEPES (Sigma) and supplemented with 10% fetal bovine serum (FBS), 200 mM L-glutamine, 100 U/ml penicillin and 100 mg/ml streptomycin. The culture medium was changed at selected time intervals with care to cause little disturbance to culture conditions. 2.5.2. Microscopic examination. Environmental scanning electron microscopy (ESEM) The materials were placed in a 24-well plate (in duplicate) and seeded with VERO cells at a density of 14  105 cells/ml. These were incubated at 37 1C. The cells were fixed with 1.5% glutaraldehyde buffered in 0.1 M phosphate buffer after a 24 h incubation period. The dried samples were sputter-coated with gold before examination under an ESEM apparatus (Philips XL 30) at an accelerating voltage of 15 KeV. 2.5.3. MTT assay TMX, PVC and films of cured systems were set up in 5 ml of MEM, FCS-free. They were placed on a roller mixer at 37 1C and the medium was removed at different time periods (1, 2 and 7 days) and replaced with other 5 ml of fresh medium. All the extracts were obtained under sterile conditions. VERO cells were seeded at a density of 11  104 cells/ml in complete medium in a sterile 96-well culture plate and incubated to confluency. Then, the medium was replaced with the corresponding eluted extract and incubated at 37 1C for 24 h. A solution of MTT was prepared in warm PBS (0.5 mg/ ml) and the plates were incubated at 37 1C for 4 h. Excess medium and MTT were removed and dimethylsulphoxide (DMSO) was added to all wells in order to solubilise the MTT taken up by the cells. This was mixed for 10 min and the absorbance was measured with a Biotek ELX808IU detector using a test wavelength of 570 nm and a reference wavelength of 630 nm. 2.5.4. Statistical analysis of biocompatibility test The statistical analysis of the extracts of cured systems was made by analysis of variance (ANOVA). In all

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statistical analyses po0:05 was considered as statistically significant.

3. Results and discussion An injectable delivery system for prednisolone, PMMA/PHEMA/PRED, has been developed with potential application in acute acoustic trauma. The system was prepared with PMMA beads in which the prednisolone powder was dispersed. The hydrophilic monomer, HEMA, widely used to obtain hydrogel networks [17], was the monomeric component. The combination of this monomer with the PMMA particles will produce a relatively flexible material which will be able to swell a certain amount of liquid in the physiological medium, contributing to the release of the drug in situ. The chemical structures of the formulation components are shown in Fig. 1. The final material will be a soft and flexible network of PMMA particles embedded into a polymerised matrix of PHEMA. The setting reaction was followed by measuring the evolution of the temperature with time from the onset of the mixture of the components. The mixture was very fluid during the first 10 min, which allowed its injection into the corresponding cavity. The values of curing parameters were determined from the temperature–time profile according to international standard specifications. A maximum temperature of 5370.9 1C and a setting time of 1571.1 min were obtained. The value of maximum temperature is rather lower than 56 1C, which corresponds to the onset of coagulation of albumin [18]. In addition, this low temperature guarantees the activity of the drug after the setting of the cement. The reaction of the control, PMMA/PHEMA, carried out in the same conditions was slower and took place at a temperature not higher than 26 1C, indicating that the prednisolone molecule acts as an accelerator of the polymerisation reaction, although the mechanism is still unclear. The swelling behaviour of the PMMA/PHEMA/ PRED and PMMA/PHEMA systems was studied in PBS at 37 1C by measuring the water uptake with time. Results are shown in Fig. 2. For both loaded and unloaded systems, water uptake increased rapidly with time of immersion during the first 10 h to reach a hydration degree of 6% for the control and 15% for the loaded system. From then on, hydration degree increased at a lower rate to reach a constant value around 10% for the unloaded system and 20% for the loaded system after 100 or 500 h, respectively. The higher hydration degree obtained for PMMA/PHEMA/PRED can be attributed to the presence of prednisolone. Although the chemical structure of this molecule (Fig. 1) is mainly hydrophobic, the hydroxilic groups present may participate in polar interactions with water

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SOLID COMPONENT

CH3 O

HO

CH3 (CH2

OH

CH3

OH

H

O

C)n H

COOCH3

C

H

O

PMMA beads

O O

O C

Benzoyl peroxide (BPO)

Prednisolone

LIQUID COMPONENT CH3 CH2

H3 C

H 2O

C

H3 C

COOCH2CH2OH

2-Hydroxyethyl methacrylate (HEMA)

CH3

CH3 N

C

N

CH3

OH

4,4'-bis(dimethylaminobenzydrol) (BZN)

Fig. 1. Chemical structures of the components of the self-curing delivery system.

21

1.0 0.8

Mt /M inf

15 12 Hydration Degree (%)

Hydration Degree (%)

18

9 6 3

18 16 14 12 10 8 6 4 2 0

0.4 0.2 0

3

6

9

0 0

100

0.6

200

12 15 t (h)

300

18

400

21

24

0.0

27

0 500

600

700

400

600

t

t (h) Fig. 2. Hydration degree versus time of the PMMA/PHEMA/PRED system (’) and the control PMMA/PHEMA (K) in PBS at 37 1C.

molecules increasing the hydrophilicity of the system in comparison to that of the unloaded system. The corresponding reduced sorption curves are plotted in Fig. 3. A straight line was obtained at the beginning of the experiment for both cases showing that, at the first stages, the mechanism of swelling obeys Fick’s second law according to Eq. (2), which can be applied for thin specimens where edge effects can be neglected: M t =M 1 ¼ 4ðDt=pl 2 Þ1=2 ;

200

800

(2)

where Mt is the water uptake at time t, M 1 is the equilibrium water uptake, D is the diffusion coefficient and l is the average thickness of the film [19]. From the corresponding slopes a value of D ¼ 0:32  107 cm2 =s was obtained for the PMMA/PHEMA system and a value slightly higher, D ¼ 0:72  107 cm2 =s; for the

1/2

800

1000

1200

1/2

(s )

Fig. 3. Reduced sorption curve of the prednisolone charged delivery system (+) and the control (  ) in PBS at 37 1C.

prednisolone loaded system. The diffusion coefficient of the control was in the range of those reported by Migliaresi et al. [20] for the water sorption of several HEMA/MMA copolymers. The prednisolone release in PBS at 37 1C is represented in Fig. 4. No burst effect was observed in the initial stages, but the release of the drug was sustained from the beginning, with 5% of prednisolone released in the first 24 h. The totality of the initial amount of drug was delivered in 35 days. A long-lasting sustained release is desirable and difficult to achieve as has been described in other delivery systems proposed in the literature for glucocorticoids [21,22]. The diagram of Fig. 4 shows a rather constant release rate of the drug during the first 15 days considering a pure diffusion mechanism of the drug through the polymeric matrix. It is noteworthy that a slight increase of the release rate

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seems to be observed in the period of 15–35 days that could be related with the migration of the PHEMA component and the consequent increase of porosity of the original matrix. The ESEM photographs of the surface of the PMMA/PHEMA/PRED system before and after the release assay are shown in Fig. 5. The dry surface shows the initial pores which were mainly produced by the entrapment of air due to the stirring of the mass during the mixture of the solid and liquid components. In the photograph of the surface after the release experiment, the PMMA spherical particles of the

Prednisolone release (%)

100

80

60

40

20

0 0

5

10

15

20

25

30

Time (days) Fig. 4. Prednisolone release in PBS at 37 1C.

35

3315

solid component can be clearly observed as a consequence of the migration of the PHEMA component of the matrix and the release of the prednisolone. It is known that the nature of the biomaterial surface, such as wettability or surface free energy, is critical for biocompatibility [23]. The wettability of the surface of the delivery system either in the presence or not of the drug was analysed through contact angle measurements. Table 1 shows the results of the cured systems along with those values corresponding to the homopolymers PMMA [16] and PHEMA [24] separately, for comparison purposes. Water contact angles of the drug-loaded and unloaded systems were in between those measured for PMMA or PHEMA; however, the value obtained for the system containing the drug was lower. Water contact angles measured in this study were comparable to those values reported for a series of HEMA/MMA copolymers with different composition [25,26]. The surface energy of solid (SES) behaved accordingly, and the polar component of SES for the loaded system was closer to that of PHEMA, indicating the more hydrophilic character of this surface. The biological response of the PMMA/PHEMA and PMMA/PHEMA/PRED systems was first analysed by direct microscopic examination at 1 and 2 days after seeding. The results are shown in Fig. 6 and they are compared with those obtained for a PMMA-based system, commonly used in acrylic bone cement formulations [27]. The cells were able to adhere and proliferate

Fig. 5. ESEM photographs of the surface of the PMMA/PHEMA/PRED system before (upper) and after immersion of sample in PBS at 37 1C (down). Left (400  ) and right (800  ).

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Table 1 Wettability properties of the loaded and unloaded delivery system prepared in this work, along with those reported for pure PMMA and pure PHEMA

Water contact angle Diiodomethane contact angle Surface energy solid, gs (mN/m) Polar part, gps (mN/m) Dispersive part, gds (mN/m)

PMMA/PHEMA/PRED

PMMA/PHEMA

PMMA

PHEMA

5974 4571 51 15 37

6772 4772 46 10 36

80 40 40 4 36

5073 3873 58 18 40

Fig. 6. ESEM photographs (900  ) of VERO cells colonization on a PMMA-based system (upper), the PMMA/PHEMA control system (medium) and the PMMA/PHEMA/PRED system (down) at 1 (left) and 2 (right) days.

on the PMMA material showing a normal cellular metabolism, and in 2 days time they formed a layer with numerous cells on the material. However, the cells observed on the PMMA/PHEMA system after 1 day showed spherical morphology indicative of a scarce material recognition and a limited adhesion. After 2

days (on the right hand) only rests of cells on this material were observed. In the PMMA/PHEMA/PRED material, no cells on the surface were observed at all at any time, which can be attributed to the anti-proliferative cellular character of the prednisolone which inhibits cellular adhesion.

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There are many studies dedicated to investigate the effect of surface wettability on the interactions of biological species with solid substrates, but the findings are rather controversial. Some studies report that materials with low surface energies showed low cell attachment [28]. Other studies, however, show that maximum adhesion of fibroblasts occurs on surfaces having moderate water wettability [29]. Our findings are consistent with the latter assumption and support the results reported for a series of HEMA/MMA copolymers, on which optimal adhesion and spreading of human endothelial cells were found to be optimum on the moderately wettable copolymer surfaces [30]. On the other hand, the biological response of the system PMMA/PHEMA/PRED was strongly influenced by the presence of the prednisolone, since glucocorticoids are known to be potent inhibitors of the cellular growth and proliferation [31]. MTT assay was used to measure cell metabolic function. MTT results for the PHEMA/PMMA system formulated in the presence or not of prednisolone are shown in Fig. 7 along with the results obtained for a PMMA cement cured in the same conditions. A significant drop in cell viability in the presence of the eluates of the cements with respect to the negative control TMX was obtained after 1 and 2 days, but it returned to normal in subsequent elutions after 7 days, indicating the absence of mitochondrial damage as a result of ‘‘leachables toxics’’ in any of the experimental formulations, and that all the formulations can be considered as non-toxic. It is noteworthy that any of the PHEMA/PMMA systems behaved similarly to that of PMMA in the MTT assay, in spite of the fact of the differences found among the three systems in the microscopy examination. One can think from these results that the no-adherence of cells to PMMA/ PHEMA surfaces is not due to any toxic reason but due to the difference in hydrophilicity as revealed in

Fig. 7. MTT cytotoxicity results for the control TMX and the test materials. Results are the mean 7SD (n ¼ 8). The extracts were collected over a 7 day period ( po0:05).

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contact angle measurements, and due to the presence of prednisolone for the loaded system.

4. Conclusion A sustained release of prednisolone can be achieved from an injectable delivery system which cures in situ at the physiological temperature to give rise to a rather soft and flexible material. The system absorbed liquids mainly in the first 10 h according to a fickian behaviour; they released the drug without an initial burst, but with a sustained release over 35 days. The analysis of the wettability of the surface showed a rather hydrophilic behaviour attributed to the polar groups of the drug, and the cytotoxic analysis showed no mitochondrial damage due to the release of toxic substances; however, no cell adhesion was observed by microscopic examination which was attributed to the anti-proliferative character of the corticoid.

Acknowledgements Financial support from the Comisio´n Interministerial de Ciencia y Tecnologı´ a, CICYT (MAT2002-04147C02-02) is gratefully acknowledged.

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