INTRODUCTION:

Kranz and Bodmeier developed in 1997 and 1998 respectively a novel polymeric based in situ forming micro particle (ISM) system as an attempt to control the release of drug from such systems. The ISM system consists of an inner drug polymer-solvent phase which is emulsified into an outer phase usually oil.

This emulsion when injected into buffer or gets in contact with physiological fluids, the internal polymeric phase solidifies and micro particles are formed spontaneously. The ISM offers may advantages over its corresponding polymeric in situ forming implant such as; little myotoxicity, better syringability and inject ability (since viscosity is highly dependent on the outer oil phase and not on the polymeric phase) and lower initial burst effect. The process of ISM system preparation is quite simple compared to the conventional methods for the preparation of micro particles (Kranz and Bodmeier, 2007).^[1]

Kranz and Bodmeier studied the release of diltiazem hydrochloride and buserelin acetate from two different in situ forming systems namely; in situ implants (ISI) and in situ micro particles. Either poly(d,l-lactide) (PLA) or poly(d,l-lactide-co-glycolide) (PLGA) in DMSO, NMP or 2-pyrrolidone was used to form polymeric solutions that were used as in situ implants. The ISM systems were prepared by utilizing the previously described polymeric solutions that were emulsifying into peanut oil at different polymeric solution to oil phase ratios. The release of both drugs from the in situ implant systems showed an initial high burst release compared to the release from the ISM system. They concluded that the ISM system significantly reduced both drug initial burst effect when compared to the in situ implant systems (polymer solutions) and attributed that effect to the presence of an outer oil phase which had made a partial barrier between the inner polymer solution and the outer aqueous medium. Another possible mechanism for the lower initial drugs burst was the less porous surface of the ISM compared to the ISI system.^[1] Another comparative study between both systems (ISI and ISM) was conducted on bupivacaine hydrochloride utilizing poly(d,l-lactide) (PLA) as a biocompatible polymer which was dissolved in various organic solvents to prepare ISI while the ISM was prepared using peanut oil as external phase at different polymer phase to oil phase ratios as previously described. A reduced initial bupivacaine hydrochloride release was also exhibited from the ISM compared to the ISI and they also attributed this behavior to presence of external oil phase and the less porous surface of the ISM. Ahmed et al. also reported the same results for haloperidol in vitro and in vivo release from ISM and ISI systems.^[8]

Jain et al. was developing a novel in situ method for the preparation of inject able stable dispersion of PLGA micro globules (premicrospheres or embryonic microspheres). The preparation made up of two oil phases. The oil phase I consists of a mixture of PLGA/ triacetin/drug/PEG 400/tween 80. This mixture is added drop wise to oil phase II which composed of miglyol 812 and span 80 and homogenized to produce rubbery inject able dispersion of PLGA micro globules. The produced embryonic or pre-microspheres harden, shrink, were able to entrapping the drug and form true microspheres in situ within 17 minutes. One major advantage of this system is its ability to control the release of cytochrome c from few days to weeks. The burst of the drug was less than 30% (within the first 24 hrs) of the total drug load and they attributed the major amount to the unencapsulated drug. We expected that the formulation and processing factors of this method could be optimized to give lower initial drug release.^[5] Another factor that could play a role in the release of drugs from in situ implant system is the drug lipophilicity. Deadman et al., studied the effect of drug lipophilicity on its release profile from different controlled release vehicles such as, PLGA micro particles and in situ forming depots. They reported that, although there was minor effect of drug lipophilicity on the in vitro studies that effect was obvious in vivo which attributed to the interactions between the formulation and the biological tissue. Moreover, recent developments in this field such as the first Phase I trial of Risperidone-ISM™ suggest there is a possibility for prolonged release of bioactive molecules applied as intramuscular injection.^[4] ISM formulations have shown advantages in comparison to the higher viscous ISI solutions. Easier inject ability through smaller needle sizes, improved muscle compatibility observed in Sprague Dawley rat^[12] and decreased initial burst release of low molecular drugs and peptides in vitro and in vivo were related to the additional lower viscous continuous phase.

MATERIALS AND METHODS:

Materials:

Biocompatible Polymer:

US Pat. No. 2003/0049320 A1^[12] describes the polymer is a long chain polymer, amorphous, semi crystalline or crystalline in nature. Preferably, the long chain polymer is one with a molecular Weight in the range of 500 to 100,000 daltons as measured by gel permeation chromatography against polystyrene standards. The chosen polymer could be biodegradable or non-biodegradable. For parenteral applications, a biodegradable polymer with a degradation profile occurring Within 1 Week to 1 year, is desirable. Examples of such biodegradable polymers useful in this invention include but are not limited to poly-L-lactic acids, poly-DL-lactic acids, poly-L-lactides, poly-DL-lactides, poly(L-lactic acid-co-glycolic acids), poly(DL-lactic acid-co-glycolic acids), poly(L-lactide-co-glycolides), poly(DL-lactide-co-glycolide), polyglycolides, polycaprolactones, polycarbonates, polyorthoesters, polyaminoacids, polyethylene glycols, polyethylene oxides, polyvinyl alcohol, polyvinyl pyrrolidone, polyoxyethylene-polypropylene block copolymers, polyethers, polyphosphazenes, polydioxanones, polyacetals, polyhydroxy butyrates, polyhydroxy valerates, polyhydroxycelluloses, chitin, chitosan, polyanhydrides, polyalkylene oxalates, polyurethanes, polyesteramides, polyamides, polyorthocarbonates, polyphosphoesters, star-branched polymers and copolymers, betacyclodextrin, polysaccharides, gelatin, collagen, albumin, fibrin, fibrinogen, polyketals, polyalkylene succinates, poly(malic acid), polypropylene oxides and other biodegradable polymers, known to a person skilled in the art of drug delivery and their copolymers, terpolymers, combinations and mixtures thereof. These polymers can either be used alone or as copolymers created from the different monomers in different ratios or mixtures of two or more different polymers or copolymers to achieve a variety of release profiles and degradation rates. The copolymers could either be random copolymers in a variety of comonomer ratios or block copolymers. Such polymers could be end-blocked or free carboxylic acid end group polymers or mixtures of these or polymers with other end groups. Preferred polymers are those with a lower degree of crystallinity and a higher degree of hydrophobicity. Such polymers include but are not limited to poly-L-lactic acids, poly-DL-lactic acids, poly-L-lactides, poly-DL-lactides, poly(L-lactic acid-co-glycolic acids), poly(DL-lactic-acid co-glycolic acids), poly(L-lactide-co-glycolides), poly(DLlactide-co-glycolide) polyglycolides, polyanhydrides, poly orthoesters, polycaprolactones and their combinations and copolymers. These polymers also include those created from interlinked segments of D- and L-lactide, or combinations of these with DL-lactide.

Biocompatible organic Solvents:

The solvents of this invention should be completely water-soluble and miscible with aqueous media in all proportions and include without limitation N,N‘-dimethyl acetamide (DMA), glycofural, dimethyl sulfoxide (DMSO), N-methyl-2-pyrrolidone (NMP), water, 2-pyrrolidone, ethanol, propylene glycol, polyethylene glycol, glycerol, sorbitol, dimethyl formamide (DMF), dialkylamides, caprolactam, glycerol formal, decylmethyl sulfoxide and other polar solvents, because of their exceptional solvating capability for the polymers described above, their non-volatility and their complete miscibility with water and with each other. The viscosity of the polymer solution is governed by the type of polymer, concentration of the polymer and molecular weight of the polymer. A particular solvent or solvent composition should be chosen for each polymer to provide a polymer solution of optimum solubility and of optimum viscosity. When a drug will be incorporated into the polymer solution, the solvent used in the invention must provide a polymer solution with a high enough viscosity to carry a fairly high drug load but should not be too viscous for processing for the purposes of the invention. This is also true when a bioinactive agent is used. The choice of solvents and solvent systems for different polymers is within the scope of understanding for a person skilled in the art of making polymer based drug delivery systems.^[12]

Biocompatible oils:

US Pat. No. 2003/0049320 A1^[12] describes biocompatible, nontoxic, nonirritant, and a non-solvent for the polymer. The oil is chosen from classes of oils which are allowed for pharmaceutical parenteral use. Such oils include without limitation various grades of animal oils such as whale oil or shark liver oil, or vegetable oils such as sesame seed oil, cottonseed oil, poppy seed oil, castor oil, coconut oil, canola oil, sunflower seed oil, peanut oil, corn oil, soyabean oil, or their fractionated counterparts such as capric-caprylic triglycerides and their salts with other acids. Preferably, the oil is chosen from super refined fixed vegetable oils such as sesame seed oil, soyabean oil, castor oil, fractionated coconut oil, poppy seed oil and such other pharmaceutically acceptable vegetable oils and their derivatives. Isopropyl myristate can also be used. Other classes of oils and their derivatives or mixtures of different oils in different proportions are known to those skilled in the art and also fall within the scope of this invention. There is no limitation to the kind of biocompatible oil chosen as long as it is gelled by the emulsifiers of the invention. The oils used in this invention are biocompatible, nontoxic, nonirritant, and a non-solvent for the polymer. The oil is chosen from classes of oils which are allowed for pharmaceutical parenteral use. Such oils include without limitation various grades of animal oils such as whale oil or shark liver oil, or vegetable oils such as sesame seed oil, cottonseed oil, poppy seed oil, castor oil, coconut oil, canola oil, sunflower seed oil, peanut oil, corn oil, soyabean oil, or their fractionated counterparts such as capric-caprylic triglycerides and their salts with other acids. Preferably, the oil is chosen from super refined fixed vegetable oils such as sesame seed oil, soyabean oil, castor oil, fractionated coco nut oil, poppy seed oil and such other pharmaceutically acceptable vegetable oils and their derivatives. Isopropyl myristate can also be used. Other classes of oils and their derivatives or mixtures of different oils in different proportions are known to those skilled in the art and also fall within the scope of this invention. There is no limitation to the kind of biocompatible oil chosen as long as it is gelled by the emulsifiers of the invention.^[12]

Polymer Concentration:

It is preferred to use polymer concentrations of between 1-90% w/w with respect to the solvent in the polymer phase. Even more preferably the polymer concentrations are in the range of 5-70% w/w. An even more optimum concentration is that between 10-60% w/w with respect to the solvent. The molecular weight of the polymer, copolymer or mixtures of polymers and their crystallinity will determine the solution viscosity. Thus a high molecular weight polymer will provide a solution of higher viscosity at a lower concentration when compared with a lower molecular weight polymer from the same class. Polymer solutions of concentrations of up to 60% w/w can be processed by raising the temperature of the polymer solution up to 25-75° C. Such concentrated polymer solutions of 10-60% w/w allow the delivery of higher loads of biologically active substances in smaller volumes of the final delivery system in contrast to the prior art compositions. Polymers concentrations up to 70% w/w can be processed by raising the temperature up to 95° C. Polymer concentrations of greater than 60% w/w to 90% w/w can be prepared by heating to 75-95° C. If a low melting polymer is used then the polymer solutions of greater than 60% w/w to 90% w/w can be processed at temperatures below 75° C. The polymer solution will generally comprise 0.01- 60%w/w of the total composition. More preferably the polymer solution will comprise 5-50%w/w and even more preferably 10-40%w/w of the total composition.^[12]

Biocompatible emulsifier:

The continuous oil phase contains from 5-70% w/w of the non-ionic emulsifiers sorbitan monostearate (span 60), sorbitan monopalmitate (span 40)and sorbitan monooleate (span 80) or a mixture thereof. The percentage of these non-ionic emulsifiers added to the oil phase will depend upon the amount of the emulsifier required to gel the continuous oil phase in the presence of the polymer solution. The higher the amount of the polymer solution that is to be emulsified the greater the amount of emulsifier is required. Also, a higher percentage of the emulsifier would impart additional stability to the gelled polymeric dispersion through an increase in the droplet stabilization. The determination of the percentage of the emulsifier required to form the gelled polymeric dispersion can be determined by a person skilled in the art of forming disperse systems. Other emulsifiers that can be used in the polymer solution may be chosen from but are not limited to polysorbates, lecithins, other sorbitan esters of fatty acids, or other emulsifiers used in the formulation of disperse systems. These emulsifiers are used in concentrations of 01-60% w/w with respect to the polymer solution. More preferably the Weight percentage of the emulsifier with respect to the polymer solution is between 5 and 50%.^[12]

Preparation of in situ forming micro particles drug delivery system:

Sonication method:

In situ implants (polymer solutions) are prepares by mixing PLA or PLGA with the solvents (2-pyrrolidone, NMP or DMSO) in glass vials until the formation of a clear solution. For the in situ implants the polymer concentration is kept constant at 40% (w/w, based on the amount of solvent and polymer). Goserelin acetate then dissolved in the polymer solution (10% (w/w), based on the polymer). The ISM systems are prepare by emulsifying the drug containing polymer solutions (PLA or PLGA in 2-pyrrolidone, NMP or DMSO) (polymer phase) into a peanut oil phase (oil phase) at a polymer to oil phase ratio of 1:1, 0.5:1, 0.25:1 and 0.1:1 by probe sonication for 30s under ice cooling. The polymer concentration varies between 0% and 40% PLA or PLGA (w/w, based on amount of solvent and polymer). Pluronic F 68 (1% (w/w) based on the amount of the total formulation). Than dissolved in the polymer phase and aluminum monostearate (2% (w/w), based on peanut oil) in the oil phase to increase the stability of the emulsions. The active agent (10% (w/w), based on the weight of the polymer) was dissolved in the polymer phase for the preparation of the Goserelin acetate-containing ISM systems.^[1]

Homogenization method:

Jain et al.^[5] have described a novel method for in situ preparation of inject able biodegradable PLGA microspheres which did not involve the use of any unacceptable organic solvents. The delivery system is a dispersion of PLGA microglobules (premicrospheres' or &embryonic microspheres') in an acceptable vehicle mixture (continuous phase) and whose integrity is maintained by the use of appropriate stabilizers. A solution of PLGA, triacetin, a model protein (cytochrome c), PEG 400, and Tween 80 (Oil Phase 1) is added drop wise with continuous homogenization to Miglyol 812-Span 80 solution (Oil Phase 2), thereby inducing phase separation (coacervation) of PLGA and forming PLGA microglobules (containing cytochrome c) dispersed in the continuous phase. This novel drug delivery system (NDDS) is dispersion and has a viscous consistency, but is sufficiently syringeable. When injected, it comes in contact with water from aqueous buffer or physiological fluid and as a result, the microglobules harden to form solid matrix type microparticles entrapping cytochrome c (in situ formed microspheres). Cytochrome c was then released from these microspheres in a controlled fashion.

Two syringe/connector system:

The O/O–ISM system comprises of polymer solution phase and external oil phase. Polymer solution phase can be prepared by dissolving a biodegradable polymer such as PLA or PLGA in a water-miscible, biocompatible solvent (this may also act as a plasticizer for the polymer) such as NMP, 2- pyrrolidone, dimethyl sulphoxide (DMSO), triacetin and/or low molecular weight grade of polyethylene glycol {such as polyethylene glycol (PEG) 200 or 400}, which are able to form highly concentrated polymer solutions in combination with surfactants such as polyethylene glycol sorbitan monooleate (Tween 80) or polyoxyethylene polyoxypropylene copolymer (Pluronic F 68). Peanut oil and sesame oil (oil for injection) can be used as a biocompatible external oil phase with surfactants such as sorbitan monooleate (Span 80) or triglyceride (Miglyol 812) with/without aluminum stearate or aluminum monostearate. Subsequently, accurately weighed internal and external phases (in different ratios) are loaded preferably into polypropylene syringes A and B, respectively. The two syringes are coupled with a Connector and emulsification can be achieved by pushing the content of syringe A into syringe B to and fro for 50 mixing cycles. At the end of this, the contents are pushed into one syringe, the connector is removed and a needle is attached, and is ready for use.^[8]

Figure 1. Preparation of an injectable ISM-emulsion prior administration in a two syringe/connector system.

However, there is still the need for improvement. One challenge, especially associated with non-aqueous ISM systems, is the relatively low ISM emulsion stability. Koerber observed a beginning phase separation within minutes despite the utilization of various potential emulsion stabilizers (Koerber 2007)^[9]. Coalescence of unstable viscous polymer solution droplets (lumps formation) prior administration may complicate an easy injection through thin needles. Furthermore, emulsion instability will result into the in situ formation of implant-like polymer lumps rather than microparticles. Differently shaped polymer matrices exhibit different surface areas which will likely affect the drug release pattern.

Figure 3: SEM of ISM of PLGA (RG 503 H)in 2-Pyrollidone after (A) 24 h, (B) 14 days and (C) 28 days stored in buffer medium pH 7.4. (D) ISM system of PLA (R 203) in 2-pyrollidone after 28 days in buffer 7.4 pH.

Characterization of ISM:

Syringe ability and Micro carrier Formation:

Syringe ability and micro carrier formation from the novel systems was determined by filling the formulations into glass syringes fitted with needles of various gauges, ranging from 14-23 gauge, and injecting the formulation into glass vials containing pH 7.0 phosphate buffer containing 0.02% Tween 80 and 0.02% sodium azide at 37°C., hereinafter stated to be the “aqueous medium”. The tubes were then capped and placed in an orbital shaker at 37° C. and mixed at 100 oscillations per minute. Syringe ability is described as the smallest bore needle through which the formulations can be delivered with ease. Micro carrier formation is defined as the formation of a uniform dispersion within a maximum time period of 24 hours with the absence of any lumps or aggregates when observed visually.^[12]

Particle Size Measurement:

The gelled compositions were filled into glass syringes fitted With 18 gauge needles and approximately 0.5-1.0 g of the gelled compositions were injected into glass tubes containing 10 ml of pH 7.0 phosphate buffer containing 0.02% Tween 80 and 0.02% sodium azide. The tubes were capped and placed in an orbital shaker at 37°C. and mixed at 100 oscillations per minute for 24 hours. The sizes of the formed micro carrier dispersions were measured using a Malvern particle size analyzer by laser light scattering.

Drug Release from the Novel Systems:

The novel gelled polymeric dispersions (0.5 g) were injected using syringes attached With 18 gauge needles followed by the addition of 5 ml of the release medium into pieces of dialysis tubing tied at one end. The other end of the sacs were tied with threads and the sacs were placed into screw-capped glass tubes containing 15 ml of pH 7.0 phosphate buffer containing 0.02% w/v Tween 80 and 0.02% w/v sodium azide. The tubes were placed in a reciprocating incubator shaker maintained at 37° C. with an oscillation speed of 100 oscillations per minute.

At different sampling points post initiation of the study, the release medium was removed from the tube and replaced with fresh medium. The amount of drug released into the medium was assayed by HPLC. The actual amount of the polymer-drug solution incorporated in the final formulation was taken as the basis for the calculation of drug release and encapsulation efficiencies. The amount of biologically active agent entrapped within the particles was determined by the difference in the actual amount of drug incorporated in the final formulations during processing and the amount released in one day.^[12]

Stability of ISM:

Some aspects of the stability of ISM systems have been investigated by researchers. Jain et al investigated the effect of myoglobin (macromolecules) on the release characteristics of ISMs and reported that the physical stability of myoglobin (helical structure) was unaffected by formulation, process, and storage conditions. Kranz et al⁽⁶⁾ also investigated the stability of ISM systems and showed that the stability of the ISM increased by dissolving Pluronic F68 in the polymer phase while placing aluminum monostearate in the oil phase of the system. The stability of PLGA and leuprolide acetate in in-situ forming systems, either ISMs or polymer solutions and lyophilized sponges, was investigated by Dong et al. It was found that degradation of PLGA increased with increasing storage temperature and water content in biocompatible solvents. Faster degradation occurred in polar protic solvents (2-pyrrolidone, PEG 400, triethyl citrate) than in polar aprotic solvents (NMP, DMSO, triacetin, ethyl acetate). The presence of leuprolide acetate significantly accelerated the PLGA degradation, especially in solution state. PLGA was stable in oily (sesame oil, soybean oil, medium chain triglyceride) suspensions at 4° C and degraded only slightly faster than solid powder at 25° C. No interaction between oils and PLGA was observed as indicated by an unchanged Tg of approximately 47° C. Finally, leuprolide acetate was chemically stable in sponges, oils and polymer solutions in suspension state, but unstable (aggregation) when dissolved in the polymer solutions and stored at 25 and 40°C. Glycerol monostearate (GMS) showed superior stabilizing potential prolonging the emulsion stability from a few minutes to more than 12 h. Flow behavior analysis, differential scanning calorimetry, polarized light- and Cryo-electron microscopy revealed, that the stabilization was caused by an immediate, more than 5-fold viscosity increase in the continuous phase after emulsification and by a stabilized interface through a liquid crystalline GMS layer around the polymer solution droplets.^[20]

Creatine kinase activity in vitro interference assay:

All of the tested solvents were evaluated to determine whether they stimulated or inhibited CK activity. Briefly, as described in ^[^10,11^], rabbit muscle CK Type I was prepared by dissolving approximately 1 mg of the enzyme in 10 ml of balanced salt solution (BSS) at pH 7.4. A given aliquot of this solution was spiked into incubation vessels containing the BSS at 37º C and bubbled with 95% O₂/5% CO₂. The tested solution was added to the test incubation vessel, while the same volume of 0.9% sodium chloride injection served as control. The amount of CK, approximately 500 U/L, was the same in both the test and control incubation vessels. All studies were conducted for 30 min. CK was determined spectrophotometrically at 340 nm using a commercially available kit, which is based upon the change in the absorbance caused by a reduction of NAD to NADH.

In vitro myotoxicity studies:

Extensor digitorum longus (EDL) muscles (approximately 150 mg) were isolated from male Sprague–Dawley rats as previously described^[^10,11^]. Briefly, rodents were administered an anesthetic dose of sodium pentobarbital and sacrificed via cervical dislocation. The EDL muscles were injected with the test solution or emulsion formulation (15 ml) using a 100 ml Hamilton syringe equipped with a needle guard to control the depth and angle of injection. The injected muscles were placed into a Teflon coated plastic basket and immersed in 9 ml of a carbogenated (95% O₂/ 5% CO₂) BSS. The solutions were drained and fresh BSS was added at 30-min intervals. The drained solutions at 30, 60, 90 and 120 min were analyzed for CK using a commercially available Spectrophotometric kinetic assay. Myotoxicity was calculated from the cumulative sum of the CK values (U/L) over a 120-min period. Phenytoin (50 mg/ml in normal saline) and 0.9% normal saline served as positive and negative controls, respectively.

In vivo myotoxicity studies:

Studies were conducted using male rats as described previously^[^10,11^]. Briefly, rats were catheterized and allowed to recover for 3 days prior to the study to allow CK-levels to stabilize at baseline. Following intramuscular injection (0.3 mL) in the right thigh muscle, blood samples (0.5 ml) were collected via the carotid artery at 0, 0.5, 1, 2, 4, 6, 8 and 12 h. The blood samples were centrifuged immediately and plasma was stored at 20ºC for analysis of CK level, while blood cells were reconstituted in heparinized (40 U/mL) normal saline solution (0.25 mL) and reinjected into the rat following the next sample to maintain blood volume. Myotoxicity was assessed by the area under the plasma CK curve.

DISCUSSION:

Micro particles prepared by the classical solvent evaporation method, the use of the lower molecular weight PLGA resulted in ISM with a lower initial release than ISM prepared with the higher molecular weight PLGA. A slower solvent diffusion from the low molecular-weight PLGA solution droplets into the release medium led to a less porous structure of the resulting micro particles, thus explaining the lower initial release. PLGA with free carboxylic acid end groups led to a lower drug release compared to PLGA with esterified end groups. 6-month controlled release leuprolide ISM could be obtained by blending poly(lactides) (PLA) with different molecular weights. The polymer concentration plays an important role in the drug release from in situ forming systems. A decrease in the drug release from in situ forming implant systems with the increasing polymer concentration was already reported. A higher polymer concentration led to a more viscous solution, which delayed the polymer precipitation and resulted in a less porous polymer matrix with a slower drug release. In ISM-systems, the initial release decreased dramatically from 62.7 to 43.7and 11.7% with an increasing polymer solution concentration of 20, 30 and 40% (w/w), respectively. The effect of polymer concentration on the second release phase (after initial release) was marginal. ISM-systems prepared with 40% RG 503H formed lumps during the emulsification into the external oil phase due to the high viscosity of the inner polymer solution and fast diffusion of NMP into the oil phase. A decrease in the internal polymer to the external oil phase ratio (1:1 to 1:2.5, w/w) led to a decreased initial release (41.6–27.0%). More oil decreased the direct contact area between the inner leuprolide-polymer phase and the release medium and increased the diffusion pathway of the drug/droplets to the oil/release medium interface, thus resulting in a lower initial release. The initial release increased with increasing surfactant (Span 80) concentration in the oil phase (w/w), which could possibly be explained with the smaller particle size at the higher surfactant concentration because of a reduced interfacial tension between the polymer solution and the oil.

The initial drug release from ISM systems (40% (w/w) PLAG based on the solvent and polymer, polymer: oil phase ratio of 0.25:1) prepared with different solvents decreased in the rank order of DMSO>NMP> 2-pyrrolidone. After 20 h, 70.8% drug release from ISM systems prepared with DMSO. This initial burst decreased to 39.9% and 30.2% for the ISM systems prepared with NMP and 2-pyrrolidone, respectively. The particle size of the ISM decreased in the rank order of 2-pyrrolidone >NMP>DMSO. Ternary solvent blends of dimethyl sulfoxide (DMSO), ethyl acetate and water are used to adjust the protein solubility in order to facilitate the incorporation of either dispersed or dissolved protein into the polymer solution. The pharmaceutically acceptable solvent DMSO is use because of its ability to dissolve both the model protein and the biodegradable polymer (PLGA). The GRAS-listed biocompatible ethyl acetate dissolves the polymer but is a non-solvent for the protein. Ethyl acetate is used in order to allow adjustments of the protein solubility. Additionally to DMSO and ethyl acetate, water has been introduced into the solvent system since preliminary investigations showed decreased dissolution times for protein in DMSO in the presence of small amounts of water (from about 2h to less than 0.5 h). This would be desirable for formulations, which require reconstitution prior to administration, e.g. where protein and polymer have to be stored separately from the solvent systems due to storage instability. On the other hand, water could alter the protein release patterns of in-situ forming drug delivery systems through an accelerated phase inversion of the PLGA solutions. The in situ forming formulations were found to be much less myotoxic than the polymeric liquid implant because the solvent has been diluted with the external aqueous phase 1–10 times. The low myotoxicity of the ISM systems with the dilution of external phase was confirmed by the in vivo myotoxicity data. It suggested that the ISM formulations were interesting drug delivery systems which were well tolerated in muscle tissues.

CONCLUSION:

In situ forming micro particle (ISM) systems offer a new encapsulation technique that provides prolonged release of drug along with much greater ease of preparation and administration than conventional micro particles and surgically implanted systems. ISMs are an attractive alternative to parenteral drug delivery, especially those prepared by existing complicated microencapsulation methods.

ACKNOWLEDGEMENT:

The authors would like to thank the Gujarat Council of Science and Technology (GUJCOST) for research grant GUJCOST/MRP/2014-15/386 and Shree Dhanvantary Pharmacy Analysis and Research Center (SDPARC) for support research work via using the lab and instrumental analysis for the study in Institute.

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Received on 29.01.2016 Modified on 17.02.2016

Res. J. Pharm. Dosage Form. and Tech. 2016; 8(2):127-134.

DOI: 10.5958/0975-4377.2016.00017.3