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• W2297398002 abstract "SUMMARY Drug encapsulation into systems able to deliver at a certain site a given amount of drug over well defined periods is used for controlling the drug release. By encapsulating drugs in biodegradable polymeric microspheres, which in turn are embedded in a hydrogel body, several release mechanisms contribute to the tuning of the release profile of the drug. Mixed hydrogels, consisting of random copolymers of 2-hydroxyethyl methacrylate (HEMA) and methyl methacrylate (MMA) cross-linked by ethyleneglycol dimethacrylate (EGDMA), were chosen to serve as body of the device. Biodegradable poly-e-caprolactone (PCL) microspheres, in which drugs such as levonorgestrel (LNG) were encapsulated, were physically embedded into hydrogels. The hydrogel-microspheres composites has been characterised in terms of water swelling and tensile properties. The release profile of the developed multicomponent drug delivery biomaterial will be discussed. INTRODUCTION Controlled release systems based on microspheres are extensively studied drug delivery systems, due to their ability to maintain optimal drug concentration over a prolonged time, to protect and stabilize the drug and increase the patient compliance by reducing the administration frequency. Poly-e-caprolactone (PCL) is one of the widely used biodegradable and biocompatible aliphatic polyester, having semicrystalline structure and a very low glass transition (-60 °C). Compared to other polymers, degradation of PCL is slow making it suitable for long-term delivery; due to its good drug permeability and biocompatibility, PCL microspheres including drugs dispersed in the polymer matrix have been extensively evaluated for the delivery of active compounds over long periods [1]. Hydrogels are polymers in three-dimensional network arrangements that are insoluble, but able to absorb and retain large amounts of water. In comparison with other synthetic biomaterials, hydrogels closely resemble natural tissues due to their relatively high water content and soft and rubbery consistency, being therefore frequently used for biomedical and pharmaceutical applications. The release of the drug from hydrogel controlled release systems is affected by the rate of water (body fluids) diffusion into the polymer, which in turn depends on the chemical structure of the polymer (polarity of the polymer segments, glass transition temperature, flexibility of the polymer backbone) and on the cross-link density and inter-chain interactions. The physico-chemical properties of the incorporated drug (size, shape, hydrophilicity) and the loading levels have an important contribution to the drug release from these systems [2]. Poly(2-hydroxyethyl methacrylate) (pHEMA) based hydrogels have been investigated for several biomedical applications, such as substrates for cellular and tissue engineering and drug delivery devices, thanks to their ascertained non-toxicity and widespread use as soft contact lenses and intraocular lenses. By embedding microspheres containing drugs into hydrogels, two different release mechanisms can be combined [3]: diffusion through the polymeric matrix for the microcarriers and diffusion through the hydrophilic matrix for the hydrogel. Combining hydrogels and microspheres lead to composites with unique release characteristics, whose mechanical properties could markedly differ from the properties of the initial hydrogel, affecting therefore the ability of the developed device to be used as an implant. For altering the mechanical properties of a hydrogel, the easiest way is to change the relative amounts of physically stronger co-monomer(s) thus modifying the stiffness of the backbone polymer or its hydrophilicity and cross-linking agents – changing the cross-linking density of the polymer network. Changes in the polymer will affect not only the mechanical properties, but also other behaviour of the material as well. The softness, flexibility and mechanical integrity of drug delivery devices based on non-biodegradable materials used as implantable systems, which can be inserted and withdrawn on request, are of major importance with regard to the implantation issue. EXPERIMENTAL METHODS Hydrogels synthesis After removal of dissolved oxygen – by nitrogen bubbling for 5 min from monomer mixtures composed of HEMA and MMA, and containing 0.1 – 0.5% (w/w) EGDMA, into some monomer mixtures either LNG was added, or PCL microspheres (with drug or drug-free) were dispersed. Subsequently, the organic phases were mixed in reagent glasses in a ratio of 3:1 with aqueous solutions of redox initiators (6.4 mg K2S2O8 and 3.2 mg Na2S2O5 / ml water). The nitrogen bubbling continued for another 15 minutes, then the reagent glasses were closed and left under the nitrogen blanket at room temperature until the gelation of the reaction mass was visible. Subsequently, the viscous liquid was stirred to achieve uniform composition and was either aspirated into one-way plastic syringes or moulded between two glass plates spaced at approximately 1 mm distance with a rubber spacer, and let to polymerise overnight at room temperature. The hydrogel rods or membranes respectively obtained were washed for at least five days with distilled water, with daily water changes, to eliminate the unreacted monomer, and were stored in distilled water. Swelling behaviour Freeze-dried rods of pHEMA-based hydrogels were immersed in distilled water at room temperature. At certain time intervals, the hydrogel pieces were extracted from the water, blotted dry with paper towel and weighed. Tensile measurements Specimens were cut from every pHEMA-based hydrogel membrane in a dumbbell shape with a custom-made steel knife. The dimensions of the swollen samples were L = 40 mm and w = 4 mm, with variable thicknesses as a function of the spacer thickness and hydrogel composition. Hydrated grease-covered samples were mounted on the Instron equipped with a 10 N load cell, and a pretension of 0.2 N was applied at a rate of the cross-head displacement of 20 mm / min. During the measurements, a constant displacement of the cross-head of 10 mm / min was used. The rate of displacement produced fracture in the specimen within 1/2 to 5 min of test initiation, as suggested by ASTM D 638-99. RESULTS AND DISCUSSION Different hydrogel formulations manifested different swelling ratios and tensile parameters, as a function of their hydrophobic co-monomer and cross-linker content. The elasticity modulus strongly increased with the increase of the cross-linker content, but this component brought about also the side effect of increased brittleness. 0.0 0.2 0.4 0.6 0.8 1.0 350 400 450 500 550 600 650 700 750 800 850 900 950 Y o u ng 's m o d ul u s, K P a" @default.
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• W2297398002 date "2011-08-01" @default.
• W2297398002 modified "2023-09-27" @default.
• W2297398002 title "Controlled release of drugs from multi-component biomaterials: drug release and mechanical properties studies" @default.
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