One of the most widely used polymers for nanoparticles is the biodegradable and biocompatible poly (d,l-lactide-co-glycolide) (PLGA), which has been approved by the FDA for certain human clinical uses. Poly (lactide-co-glycolide) polymers undergo hydrolysis upon administration to the body, forming biologically compatible and metabolizable moieties (lactic acid (L) and glycolic acid (G)) that are eventually removed by the citric acid cycle [1].

The unique structure of PLGA nanoparticles, composed of a hydrophilic surface and a hydrophobic core, provides a drug-carrying reservoir and also enables them to dissolve in aqueous solutions. Many approaches are proposed for the preparation of PLGA nanoparticles. The emulsification-evaporation method, spontaneous emulsificationsolvent diffusion method (SESD) [2], and nanoprecipitation method are all widely used for preparing various diameters of PLGA nanoparticles. During nanoparticles formation by emulsification evaporation and SESD approaches, toxic organic solvents such as CH₂Cl₂ and CHCl₃ are usually used. To meet the requirement for clinical use, the residual solvents should be completely removed from the PLGA particles [3].

Interfacial deposition is a process used for the production of nanocapsules; however, this is not a polymerization technique but an emulsification/solidification technique. In interfacial deposition, a fifth compound is introduced, of oil nature, miscible with the solvent of the polymer but immiscible with the mixture. The polymer deposits on the interface between the finely dispersed internal phase droplets and the aqueous phase, forming nanocapsules. An aqueous solution is used as the dispersing medium. The main difference is that polymers such as PLA are dissolved together with the drug in a solvent mixture (eg, benzyl benzoate, acetone, and phospholipids). This mixture is injected slowly into a stirred aqueous medium, resulting in the deposition of the polymer in the form of nanoparticles of about 230 nm in size [4].

The interfacial deposition method is recommended for the incorporation of hydrophobic drugs into polymeric nanoparticles [5,6]. As described by several authors and also as demonstrated in the work of Fonseca et al. [7] the establishment of a protocol that allows nanoparticles precipitation, while avoiding extensive diffusion of the drug along with the solvent aiming at obtaining high values of drug encapsulation is a challenging issue [8]. Also, the aggregation of PLGA NPs during solvent-evaporation process is a notable problem regardless different preparation method. In order to prevent PLGA nanoparticle aggregation, polymer stabilizers are often used. Many stabilizers such as poly (vinyl alcohol) (PVA), poly(vinyl pyrrolidone) (PVP), Tween 80, Fluonic 127 (poloxamer 407), Fluonic 68 (poloxamer 188), didodecyl dimethyl ammonium bromide (DMAB) are also excellent stabilizer candidates [3]. These stabilizers are coated on the surface of PLGA nanoparticles and can affect the zeta potential, particle size and particle surface properties. PLGA nanospheres have been suggested to be a good carrier because of the safety and achieving sustained release. The degradation time of PLGA can be altered from weeks to months by varying the molecular weight of the copolymer or the lactic acid to glycolic acid ratio in copolymer [9].

The objective of this work was to obtain an optimum formulation of PLGA nanoparticles by utilizing factors which are likely to influence their mean size, surface morphology, encapsulation efficiency, and drug release. Considering lomustine as a lipophilic model drug, biodegradable lomustine loaded nanoparticles were prepared by interfacial deposition method. This is based on the interfacial deposition of a polymer following displacement of a semipolar solvent miscible with water from a lipophilic solution. The solubility and compatibility of the drug was tested with materials used for preparing the drug delivery system. The influence of various formulation components such as PLGA concentration, grade of PLGA: PURASORB PDLG 5002 and PDLG 7502, organic phase volume, lomustine content, PVA and Pluronic F68 concentrations as 1% and 3%w/v, PVA and pluronic F68 solution content (aqueous phase volume) 1:10 as that of organic solution on the characteristics of nanoparticles was investigated. The cytotoxicity of the selected PLGA-lomustine nanoparticles was evaluated on the L132 Human Lung cancer cells. Also stability and biodistribution of selected nanoparticles was observed with toxicological/ histopathological studies.

Materials

Polymer PLGA was a gift sample from Purac Biomaterials, Gorinchem, The Netherlands. Lomustine pharmaceutical grade was obtained from Fujian Provincial Medicines and Health Products, Xiamen Import and Export Corporation (China). Acetone used of Qualigens, India; Pluronic F68 was used of Hi-Media, India. All other chemicals were of analytical grade.

Methods

Preformulation studies

Saturation Solubility of lomustine in PVA solution, Pluronic F68 solution, in polymers PDLG 5002, PDLG 7502 and in phosphate buffer saline pH 7.4: Lomustine solubility in oversaturation conditions (C > 10 x Cs) was obtained in PVA solution (1%w/v), PluronicF68 solution (1%w/v), PURASORB PDLG5002 solution (1%w/v), PURASORB PDLG7502 solution (1%w/v), Phosphate buffer saline (PBS) pH 7.4 by dispersing 300 mg of drug in 100 ml of respective solutions for water solvent and 3g or more until saturation of drug in 5ml of respective solutions for acetone solvent. The suspensions were stirred under constant magnetic stirring at 37 ± 0.5°C for 24 h (adequate time for equilibration), filtered through a syringe filter (0.45 μm) and then assayed spectrophotometrically at 230 nm by dilution with ethanol 95%v/v for acetone solvent, and at 230.4 nm by dilution with PBS pH 7.4 for water solvent.

Differential scanning calorimetry (DSC) of lomustine, 1:1 mixtures of drug:polymer (PDLG 5002 and PDLG 7502): The DSC runs were carried out on the DSC (DSC-60, Shimadzu Co., Kyoto, Japan) instrument. Five to ten milligrams of drug and polymer was placed in aluminium pans and hermatically sealed. The heating rate was 10°C/min; nitrogen served as purge gas and the system was cooled by liquid nitrogen.

Fourier-transform infrared spectroscopy (FTIR) of pure drug, polymer and 1:1 mixture of drug: polymer (PDLG 5002 and PDLG7502): Infrared spectra were recorded on Shimadzu 8400S FTIR spectrophotometer using KBr pellet method.

In-vitro lomustine release: The in vitro drug release of lomustine nanoparticles was determined on a Franz diffusion cell. As regards to sink condition, 2 mg of lomustine in 2ml PBS pH 7.4 was placed in donor site and 50 ml PBS in receptor chamber with membrane (cellulose membrane, mw cut off 12,000 Da) placed in between, incubated at 37°C under magnetic stirring (200rpm). At specific time intervals, 1ml of medium was removed and replaced with the same volume of fresh PBS. The samples were analyzed spectrophotometrically at 230.4 nm.

Preparation of nanoparticles: Nanoparticles were prepared by the modified interfacial deposition method [7]. Components are listed in Table 1. An organic solution was prepared by dispersing PLGA and lomustine in 10ml acetone one by one. The organic phase was added drop wise into 1%w/v polyvinyl alcohol (PVA) aqueous solution 1:5 under magnetic stirring at room temperature. Stirring was continued until complete evaporation of organic solvent. Subsequently, nanoparticles were separated by centrifugation in Remi C-24 cooling centrifuge at 19,000 rpm at 4°C for 40 min. The drug-loaded nanoparticles (DNPs) were diluted with 10ml of distilled water, the dispersion was centrifuged and supernatant was discarded. This was repeated three times to remove traces of acetone from the DNPs. The final volume was adjusted to 10ml with water. The residue was used for further characterization by evaporating to dryness to constant weight at reduced pressure at 35°C. Blank nanoparticles were prepared as per above method omitting the drug. All these nanoparticles and their preparations were fabricated in laminar air flow workstation.

Table 1: Components of various batches of fabricated lomustine nanoparticles, their particle size, polydispersity index, zeta potential, drug content and entrapment efficiency.

For study of effect of different formulation components, the following variables were studied. a) PLGA concentration, b) PURASORB PDLG 5002 and PDLG 7502, c) Organic phase volume, d) Lomustine content, e) PVA and Pluronic F68 concentrations as 1% and 3%w/v, f) PVA and Pluronic F68 solution content (aqueous phase volume) 1:10 as that of organic solution. Levels of all variables were chosen on their usual concentrations reported in the literature.

Nanoparticle characterization

Particle size, polydispersity index (PDI), zeta potential and scanning electron microscopy (SEM): The method used by Mehrotra and Pandit [10] was used. The Zetasizer used was of Malvern instruments, UK DTS Ver 5.10 and SEM was performed on Hitachi Japan S-3400N.

Drug content and encapsulation efficiency: The method published by Mehrotra and Pandit [10] was used.

In- vitro drug release study: The method published by Mehrotra and Pandit [10] was used. Here all method is similar only lomustine was determined by dialysis method using Franz diffusion cell. Lomustine nanoparticles in 2ml PBS pH 7.4 was placed in donor site and 50 ml PBS in receptor chamber with membrane (cellulose membrane, mw cut off 12,000 Da) placed in between which was washed with double distilled water previously and incubated at 37°C. For evaluation of release kinetics, the obtained release data was fitted into spherical matrix model also along-with first order, zero order and Higuchi equations.

In- vitro cytotoxic activity: The method used by Mehrotra and Pandit [10] was used here. The nanoparticle formulations used were of batches PD1, PD4 and PD13 based on particle size, physical stability and drug release data.

Stability studies at room temperature (25°C) and at 45°C: The stability study was carried out of freshly prepared and dried under reduced pressure nanoparticles were conducted in sealed vials placed in stability chamber maintained at 25°C, 60% RH and at 45°C, 70% RH. The nanoparticles subjected to stability tests were analyzed over three month’s period for drug content with a frequency of one month sampling. Three formulations were used for study namely formulation PD1 which was basic formulation, formulation PD4 with different polymer i.e. PDLG 7502 and formulation PD13 with different stabilizer i.e. 3%w/v Pluronic F68 based on particle size, physical stability and drug release data.

In- vivo biodistribution study in albino mice: Healthy albino mice (20 ± 5 g) were used for study. The animals were divided into groups of six mice each. Then 0.1 ml of the suspension in buffer solution (PBS pH 7.4), containing lomustine, lomustine-loaded nanoparticles of different batches: PD1, PD4, PD13 (40 mg/kg lomustine equivalent) and buffer solution were intravenously administered into the mice through tail vein, separately in different groups. At predetermined time intervals (12, 24, 36 and 48 h) blood samples (1ml in duplicate) were collected from the carotid artery, placed into heparinized test tubes and centrifuged to get corresponding plasma samples. Afterward, the animals were dissected and each tested organ (liver, lung, kidney, spleen heart) was excised. Blood was centrifuged in eppendorf tubes at 4°C (15000 rpm for 10 min), and plasma was collected. Organs of interest (lung, kidney) were washed with 0.9% w/v saline, blotted dry and were stored at –80°C until analysis for lomustine. Aliquots of harvested organs (80-150 mg) were homogenized in saline and acidified to pH 3.0 with acetic acid. Both the plasma samples and tissue homogenates were stored at 4°C for 12 h. Then tissue homogenates and plasma samples were treated with two volumes of cold mixture of acetonitrile/ methanol (1/1 v/v) to precipitate proteins and extract lomustine. The obtained suspensions were centrifuged at 15,000 rpm for 10 min, and 20 μl of the clear supernatant was injected into the HPLC system to determine the concentration of lomustine. The HPLC analysis method used by Al-Shammary (1990) was used. The HPLC condition was modified for better analysis of drug in bio-samples. Mobile phase was a mixture of acetonitrile/water/methanol, 48:41:11v/v/v. Detection of drug was by UV adsorption measurement at 234 nm (flow rate 1ml/ min). The concentration of drug was determined based on the peak area at the retention time of 2-9 min by reference to calibration curve. The assay was linear over the tested concentration range 8 ng to 340 ng. For histological tissue analysis, kidney and lungs were fixed in neutral buffered 10% formalin, processed by standard methods in paraffin, sectioned at 5–10 μm, and stained with hematoxylin and eosin (H & E).

Statistical analysis: It’s done a same analysis as published by Mehrotra and Pandit [10]. Mean values of nanoparticle size, polydispersity index, encapsulation efficiency and cytotoxic activity were compared.

Saturation solubility of drug in stabilizer, surfactant, polymer and phosphate buffer

Observed saturation solubility of lomustine in different solutions is shown in Figure 1. The experimental values are the average of three replicates and standard deviations did not exceed 3% of the mean value. The higher solubility of lomustine in phosphate buffer saline (PBS) pH 7.4 could be attributed to its reported solubility in saline solution [11]. The saturation solubility of lomustine was significantly higher in PURASORB PDLG5002 solution (1%w/v), PURASORB PDLG7502 solution (1%w/v) which was a result of its reported solubility in acetone.

Figure 1: SEM image of Nanoparticles: a)Basic formualtion PD1 Magnification 5.00k SE, b) Batch PD4 Mgnification 10.0k SE, c) Batch PD11 Magnification 5.00k SE, d) Batch PD12 Magnification 3.00k SE, e) Batch PD13 Magnification 5.00k SE, f) Batch PD14 Magnification 20.0k SE, g) Batch PD15 Magnification 3.00k SE.

Differential scanning calorimetry (DSC)

The compatibility of the materials used for preparing the drug delivery system was characterized before analyzing the effect of the formulation process on the stability of different materials involved. In this study, the physical interaction between lomustine and PLGA present in the 1:1 physical mixtures was studied. Thermogram of pure lomustine showed an endotherm with sharp melting at 91°C. Pure PLGA polymer PDLG 5002 and PDLG 7502 exhibited relatively distorted thermal transition, confirming the amorphous nature of PLGA. From DSC tracings of the corresponding 1:1 physical mixtures of lomustine: PLGA a sharp endotherm at 90.06°C and 91°C was observed. This endotherm can be easily attributed to the melting of crystalline lomustine that persisted in the physical mixture without any interaction. There was no interaction between polymer and drug.

Infrared spectroscopic analysis

The infrared spectroscopic analysis was performed to observe possible interactions between drug and polymer and to complement the results from DSC. The IR spectra for both drug and polymer and physical mixtures of drug: polymer was observed.

The characteristic bands observed from the IR data of lomustine included characteristic ring absorption of cyclohexyl ring of lomustine at 1533 cm^-1 and 1490 cm^-1, C=O str. at 1703 cm^-1. In the FTIR spectra of PLGA the characteristic absorption bands at 1733 cm^-1 and ester group, 846 cm^-1 C-C str. vibrations were identified.

In the FTIR spectra of 1:1 physical mixture of lomustine: PLGA (PDLG 5002 and 7502) the characteristic IR spectra were very similar, showing all the bands of the functional groups of lomustine and PLGA identified in the isolated compounds. The maintenance of these characteristic bands of both the drug and PLGA polymer, as well as the absence of new IR bands, indicates that there was no chemical interaction between the lomustine and the PLGA, demonstrating that lomustine does not react with the polymer and is only dissolved in the PLGA polymeric matrix. Any interaction present in physical mixture was tested in the thermal analysis experiments also as given above. In fact, since both the drug and the polymer have similar lipophilicity character a homogeneous distribution of the molecular drug into the matrix structure of the polymeric nanoparticles could be expected.

Particle size and zeta potential

The nanoparticles were of nanometric size with homogeneous and a narrow particle size distribution and negative zeta potential (Table 1). Smaller particles were obtained when organic phase volume was increased to 20ml and 30ml and with 50ml of 3%w/v pluronic F68 surfactant content.

This preparation method is applied to types of PLGA polymers, which varied in copolymer ratio lactic acid: glycolic acid (L/G). The difference in L/G ratio did not have a significant influence on the average size, only slight increase was observed in size. This was in agreement with findings of Jiang et al. [12]. This could be regarded as a technical advantage of the interfacial deposition method, since the formation of nanoparticles would not be influenced by such factors as the L/G ratio and the polymer species.

Pluronic F68 incorporated was oriented at organic solvent/water interface to reduce efficiently the interfacial tension, which resulted in significant increase in the net shear stress at a constant energy density during phase mixing [13] and resultant emulsification and promoted the formation of smaller emulsion droplets. Thus, the mean diameter of nanoparticles decreased with the presence and increase of Pluronic F68 concentration.

Acetone is a freely water-miscible organic solvent. When organic phase volume increased, the rapid dispersion of considerable amount of acetone into the external aqueous phase contributed to a significant decrease of the interfacial tension, thereby decreasing the particle size.

The negative charge on the PLGA nanoparticles was attributed to the presence of uncapped end ionized carboxyl groups of the polymer at the particle surface [14]. PVA has been extensively used as a promising stabilizer for PLGA nanoparticles. The mechanism of PVA binding with PLGA has been proposed to be due to the interpenetration of PVA and PLGA molecules during nanoparticle formation [15]. The presence of PVA formed a stable coating network on the polymer surface. This network shielded the surface charge and moved the shear plane outward from the particle surface, which resulted consequently in a slightly negative zeta potential. Despite this comparatively weak zeta potential, the nanoparticles stabilized by the layers of PVA surrounding the nanoparticles by steric hindrance.

The reduction seen in zeta potentials by using pluronic F68 surfactant was due to the fact that the coating layers shield the surface charge and move the plane of shear outwards from the particle surface [16,17]. Similarly the results obtained by Zou et al. [18] also correlate the zeta potential with shielding of the surface negative charge of PLGA.

Scanning electron microscopy (SEM) of some critical batches

The SEM analysis of nanoparticles of batches PD1, PD4, PD11, PD12, PD13, PD14, and PD15 was performed which shown smooth spherical particles as shown in Figure 2a-g respectively.

Figure 2: SEM image of Nanoparticles: a)Basic formualtion PD1 Magnification 5.00k SE, b) Batch PD4 Mgnification 10.0k SE, c) Batch PD11 Magnification 5.00k SE, d) Batch PD12 Magnification 3.00k SE, e) Batch PD13 Magnification 5.00k SE, f) Batch PD14 Magnification 20.0k SE, g) Batch PD15 Magnification 3.00k SE.

All the nanoparticles were spherical with smooth surface. No aggregation, separated particles with smooth surfaces confirmed suitability of different formulation parameters selected for preparation of nanoparticles. From the observed formulation parameters PVA was a critical factor affecting significantly a preparation process especially a purification and isolation step as there was observed some residual matter in SEM photomicrographs of formulations containing more amount of PVA than 50ml 1%w/v PVA (Figure 2c and 2f). This might be attributed to increased viscosity of aqueous solution in presence of increased amount of PVA, which require extra washings of residue after centrifugation. Surfactant pluronic F68 in all concentrations used produced nanoparticles with smooth surfaces with method of interfacial deposition used.

Encapsulation efficiency

The effect of variable process parameters on entrapment efficiency is given in Table 1. Entrapment efficiency was increased significantly with increasing polymer concentration. It was increased slightly with using copolymer PDLG 7502 and increased significantly with increasing amount of lomustine (p<0.05). It was decreased with increasing amount of surfactant pluronic F68 from 1% to 3%w/v.

Lomustine demonstrated the highest encapsulation efficiency in this process used. More lipophilic drugs do not suffer from the problems of leakage of drug to the external medium, resulting in improved drug content in the nanoparticles. The entrapment efficiency of drug in nanoparticles correlated well with its solubility in different stabilizer solutions [19]. Due to greater solubility of lomustine in pluronic F68 solution than in PVA solution, on mixing of two phase’s lomustine migrates rapidly to the external phase resulting in low entrapment, while in PVA solution, the hydrophobic interaction of the lomustine and polymer supersedes due to the lower solubility of lomustine in PVA solution, resulting in higher entrapment efficiency. In these lomustine nanoparticles, hydrophobicity of drug and relative proportions of drug/polymer/solvent might be one of the causes of high encapsulation efficiency [20].

Danhier et al. [21] observed that nanoparticles prepared with interfacial deposition/nanoprecipitation method could achieve higher encapsulation efficiency than with the simple emulsion technique. Thus the process itself is beneficial for encapsulation efficiency of hydrophobic drugs. Addition of droplet stabilizer less polar for drug further showed synergistic effect.

As the amount of lactide of PLGA increased, encapsulation efficiency of lomustine increased slightly. With the increase of L: G ratio of PLGA, the interaction or affinity of drug probably increased [13,22].

The encapsulation efficiency of lomustine increased significantly (p < 0.05) with the increase of PLGA concentration. This phenomenon probably resulted from the increase of viscosity. Increasing viscosity could increase the drugs resistance diffusional into the aqueous phase and thus enhance the drugs incorporation into nanoparticles. Additionally, larger nanoparticles had higher drug entrapment efficiencies.

The slightly increased encapsulation efficiency with increase of PVA concentration was probably caused by the increase in particle size. Moreover, with the increase of pluronic F68 concentration, more molecules of lomustine might be partitioned out rapidly into the aqueous phase during phase mixing procedure and less drug molecules remained in emulsion droplets to interact with PLGA molecules, hence decreasing the encapsulation efficiencies.

The increase of the lomustine content resulted in a significant increase in encapsulation efficiency (p < 0.05) as the lomustine concentration in the organic phase increased and then more drug molecules could interact with PLGA molecules, resulting in the increased encapsulation of lomustine. As the initial lomustine amount increased, the amount of lomustine partitioned into the aqueous phase probably reduced during phase mixing procedure. Thus more lomustine molecules could remain in emulsion droplets to interact with PLGA molecules and the encapsulation of lomustine could be further enhanced.

When aqueous phase volume increased, the amount of lomustine precipitated in the aqueous phase resulted in less lomustine retention in the internal phase to interact with PLGA molecules and then shown lower encapsulation efficiency.

The encapsulation efficiency of lomustine increased first with the increase of organic phase volume (p < 0.05) then it decreased. This occurred possibly because the change of acetone volume affected the partition of drug in the organic phase. The hydrophobic molecule lomustine was easy to dissolve in acetone. When acetone volume increased, more lomustine molecules were carried into the aqueous phase or at organic solvent/water interface by considerable amount of acetone; and thus less lomustine molecules remained in the internal phase to interact with PLGA molecules and lower entrapment efficiency was obtained. Afterwards, further increase in acetone volume increased the partition of lomustine in the organic phase. Hence, more lomustine molecules interacted with PLGA molecules in the internal phase and enhanced the encapsulation.

In- vitro drug release

The percentage drug release of lomustine in Franz diffusion cell was complete. The in vitro release behaviour of lomustine from PLGA nanoparticles; from the polymer matrix exhibited a fast initial release during the first 24h followed by a slower and continuous sustained release over 48 hr.

The release rate of the lomustine from the nanoparticles and its appearance in the dissolution medium was governed by the partition coefficient of the drug between the polymeric phase and the aqueous environment in the dialysis bag and by the diffusion of the drug across the membrane as well. The dialysis bag retained the nanoparticles and allowed the diffusion of the drug immediately into the receiver compartment [23]. The biphasic release profile, with an initial burst of drug release attributed to surface associated drug, followed by a phase of slower release as drug entrapped in the nanoparticles diffuses out into the release medium.

Also, the effect of different formulation variables on drug release was observed. Effect of polymer amount on drug release profile of lomustine (Figure 3a) with respect to basic formulation showed that the drug release was slow as the amount of polymer was increased. This was because of increased viscosity and diffusion path. Molecular weight and crystallinity of the polymer influence drug release and degradation of the nanoparticles [2]. Also this drug release mirrors polymer mass loss and was therefore dependent on polymer composition and molecular weight, the higher the molecular weight and lactide content of the polymer the longer the lag time (reflected in a large t_max) before the commencement of the polymer degradation controlled phase [24].

Figure 3: Drug release profile of lomustine from nanoparticles: a) as a function of polymer amount b) as a function of organic phase volume c) as a function of lomustine content d) as a function of PVA and surfactant Pluronic F68 concentration.

When organic phase volume was increased drug release rate was increased (Figure 3b). This was due to increased solubility of lomustine in increased volume of organic solvent resulting increased diffusivity and hydrodynamics at the interface.

When drug loading was increased drug released rate was increased (Figure 3c). The reason proposed by Corrigan and Li [24] for higher drug loading giving more rapid release, was that a high density of interconnecting channels (‘active sites’) increase polymer permeability in the presence of drug and that result in increased PLGA degradation and erosion. In addition the polymer degradation rate is influenced by the presence of drug which may interact [24] and cause plasticization of the polymer, a lowering of the Tg and more rapid polymer degradation and hence drug release.

A slight increase in drug release was observed when concentration of surfactant Pluronic F68 was increased (Figure 3d). This was attributed to the high concentration of drugs dispersed, which effectively increases the proportion of drug linked to the particle surface–liquid interface.

The nanoparticle size was also associated with changes in drug release kinetics. The smaller sized nanoparticles prepared with lower amounts of PLGA showed higher drug release rates. This release behaviour may be explained by a corresponding increase in the total nanoparticles surface, resulting in a larger drug fraction exposed to the leaching medium. Smaller nanoparticles size lead to shorter average diffusion path of the matrix-entrapped drug molecules [26]. The diffusion distances encountered in the particles were small which allowed drug trapped in the core to rapidly diffuse out and also for the release medium to diffuse in.

Mechanism of drug release

The correlation coefficients (r2) values for a plot of Higuchi’s equation were more linear than for other plots. The values were in the range of 0.9758 to 0.9976, which was always higher than other equations indicating square-root time dependant release kinetics of lomustine from the lomustine loaded nanoparticles. The release exponents (n) were <1.0 indicating non-fickian release mechanism.

Stability study

Nanoparticles were studied for stability at room temperature i.e. 25°C/60%RH and at 45°C/70%RH for three months and drug content was found as shown in Table 2.

Table 2: Drug content of nanoparticles stored at different temperatures for 3 months.

After storage for 3 months at 25°C/60%RH, nanoparticles displayed no significant change (p>0.05) of drug content for all the three tested nanoparticle batches. At 45°C/70%RH the drug content was maintained and not significantly (p>0.05) affected for one month, after 3 months nanoparticles drug content was slightly affected during storage. This result showed that PLGA nanoparticles of lomustine prepared possessed good storage stability at 25°C/60% RH, due to an adequate proportion of different constituents in the preparations. At 45°C/70% RH nanoparticles with higher concentration of pluronic F68 preserved the content of a thermolabile drug lomustine, proving it a suitable stabilizer.

In- vitro anticancer activity

The percent viable cells were reduced with nanoparticles formulation as shown in Table 3. In-vitro cytotoxic activity tested showed that cell viability was decreased significantly by lomustine nanoparticles and percent cell viability and IC₅₀ values at 24h incubation are given in Table 3.

Table 3: Cytotoxic activity of lomustine and nanoparticles of batch PD1, PD4 and PD13 against L132 cellline at 24h.

Lomustine nanoparticles showed significant anticancer activity at concentration of 100 μg/ml. The observed cytotoxic effect on cells probably arose from the degradation products of polymer or adhesion of solid nanoparticles to the cell surface [27-29]. Surface charge of nanoparticles determines the performance of the nanoparticle system in the body, e.g. interactions with cell membranes [2]. Degradation products may change the pH of the environment or interact with the cells due to their exposed functional groups or surface charge. It has been shown that higher toxicity is associated with more rapidly degraded polymers [28,29].

Biodistribution profile of lomustine and lomustine nanoparticles in mice

There was fast decay of lomustine in plasma when lomustine solution was administered to mice (Figure 4) and at 24 h post administration, the lomustine concentration in plasma could not be detected.

Figure 4: Lomustine concentration in plasma in albino mice at various time intervals after administration of lomustine solution and nanoparticles of batch PD1, PD4, PD13.

The plasma concentration of lomustine after administration of nanoparticles was higher (Figure 4) and still showed lomustine concentration in plasma at 48 h over 14 μg/ml. The lomustine concentration profiles in plasma shown here indicate that lomustine level in blood could be maintained for an extended time period for nanoparticle sample due to the long circulating property of the polymeric nanoparticles. It was found that circulation of PLGA nanoparticles loaded with lomustine could be improved by stabilization with a stabilizer. Lomustine was still detected in the serum 48 hours after initial administration. As a result of the rapid clearance of free drug over 24 hours, drug found in the serum is believed to be encapsulated in nanoparticles. The higher plasma concentration of PD13 nanoparticles was due to the presence of Pluronic F68 layer on the surface, which shifted the shear plane of the diffusive layer to a larger distance. Surface modification or coating by biocompatible (hydrophilic) polymers improve uptake of nanoparticles and enhance stability. PEG, poloxamers and poloxamines are examples [2].

The distribution profiles of lomustine in lungs and kidney showed that compared with lomustine solution (Figure 5a), the biodistribution of lomustine was changed in the group treated with lomustine-loaded nanoparticles of batch PD1, PD4 and PD13 (Figure 5b-5d). Lomustine concentration in lung was elevated with the nanoparticles than with lomustine solution and it was in decreasing order with nanoparticles of batch PD13>PD4>PD1. Lomustine elimination was slow when nanoparticles were administered to mice as it was detected in kidney even after 48 h post administration as compared to lomustine solution.

Figure 5: Biodistribution of lomustine in albino mice after administration of a) lomustine solution, b) lomustine nanoparticles batch PD1, c) lomustine nanoparticles batch PD4, d) lomustine nanoparticles batch PD13.

Here, For nanoparticles taken up by opsonisation of the mononuclear phagocytic system (MPS) present in the liver, spleen, bone-marrow and kidney, instead of particle size, deformability of nanoparticles might be considered to understand their distribution in tissues and opsonization by MPS, because the real particle size can be changed against the deformability of particles (i.e. the mechanical strength, swelling behavior, capacity to undergo hydrolysis and subsequent biodegradation) in the blood stream.

Histopathological examination of the lung and kidney of albino mice was carried out to study the presence of any toxicity or inflammatory response after administration of lomustine nanoparticles as compared to vehicle and lomustine solution at 24 h post administration.

In Group 1 (Figure 6a-b), Group 3 (Figure 6e-f), Group 4 and Group 5 (Figure 6g-h) no microscopic change was observed in lungs and kidney. In Group 2 the microscopic alterations were observed in lungs and kidney. Vacuoles in alveoli were seen in lungs (Figure 6c). Fatty changes were observed in kidney (Figure 6d).

Figure 6: Photomicrograph of H & E stained histological section (magnification 10 x 10) showing the effect 24 h after administration of a) 0.1ml/kg of vehicle PBS pH 7.4 as a control sample (group 1) on the albino mice kidney, b) 0.1ml/ kg of vehicle PBS pH 7.4 as control sample (group 1) on the albino mice lung, c) 0.1ml of 40mg/kg lomustine solution (group 2) on the albino mice kidney, d) 0.1ml of 40mg/kg lomustine solution (group 2) on the albino mice lung, e) 0.1ml of 40mg/kg lomustine equivalent nanoparticles batch PD1 (group 3) on the albino mice kidney, f) 0.1ml of 40mg/kg lomustine equivalent nanoparticles batch PD1 (group 3) on the albino mice lung, g) 0.1ml of 40 mg/kg lomustine equivalent nanoparticles batch PD13 (group 5) on the albino mice kidney, h) 0.1ml of 40mg/kg lomustine equivalent nanoparticles batch PD13 (group 5) on the albino mice lung.

The most affected organ by drug toxicity in the histopathological study is generally kidney where degenerative and fatty changes of nephrotoxicity can be detected. Generally the other most significant organ which shows histological changes is lung showing lesions. As reported in literature in mice lung lesions like dark pigmented alveolar macrophages, eosinophilic foreign bodies and alveolar hyalinosis same alveoli vacuoles are marked by only a free drug administration and absence of these changes by nanoparticles administration protects histology of the organs. It shows nanoparticles can significantly change the biodistribution of the loaded drug. An optimized biodistribution may lead to improved drug efficacy and, at the same time, reduced side effects and act as a reservoir to gradually release the encapsulated drug.

More encouragingly, drug concentration in lung was also markedly elevated several folds with the nanoparticles, possibly due to the filtration effect of the lung capillary bed. Since lung has commonly been an organ that antitumor drugs could not efficiently reach through i.v. administration, an increased dose level or the use of aerosol is often needed to achieve efficacious lung concentration [30]. Therefore, our results suggest that lung targeting could be easily realized with i.v. injection of the drug-loaded nanoparticles, which makes these nanoparticles valuable for lung cancer treatment. Drug concentration in kidney show significant difference between groups treated with drug-loaded nanoparticles and with free drug injection with time observed between those groups.

It was concluded that formulation variables could be exploited in order to enhance the incorporation of lomustine into PLGA nanoparticles by interfacial deposition method. Based on the optimal parameters, it was found that lomustine-PLGA nanoparticles with expectable properties could be obtained. A technical advantage of the interfacial displacement method is that the formation of nanoparticles is not much influenced by the L/G ratio and the polymer species. This method is simple, mild and practically easy because lomustine-anticancer drug, demonstrated the highest encapsulation efficiency. Here it is notable that, mechanical strength and its ability to be formulated as a drug delivery device are controlled by molecular weight and its intrinsic viscosity. The intrinsic viscosity increase with the increase in PLGA concentration could increase the drug resistance diffusion into the aqueous phase and thus enhance the drug’s incorporation into nanoparticles. The lomustine loaded PLGA nanoparticles stabilized by 3%w/v pluronic F68 in the ratio of 1:5 organic phase: aqueous phase volume (Batch PD13) formulated is optimized as best formulation with stability of 96.5%w/w drug content and highest controlled drug release time profile with IC₅₀ value 13μg/ml and in-vivo release with increased lung concentration of 35.2 μg/ml of drug even at 48 hrs, with slowed elimination 16 μg/ml and no histopathological toxicity.