International Journal of Pharmaceutics 474 (2014) 134–145

Contents lists available at ScienceDirect

International Journal of Pharmaceutics journal homepage: www.elsevier.com/locate/ijpharm

Pharmaceutical nanotechnology

Amphiphilic chitosan-grafted-functionalized polylactic acid based nanoparticles as a delivery system for doxorubicin and temozolomide co-therapy Antonio Di Martino a,b , Vladimir Sedlarik b, * a b

Polymer Centre, Faculty of Technology, Tomas Bata University in Zlin, nam. T.G. Masaryka 275, 76272 Zlin, Czech Republic Centre of polymer systems, University Institute, Nad Ov9 círnou 3685, 76001 Zlin, Czech Republic

A R T I C L E I N F O

A B S T R A C T

Article history: Received 30 April 2014 Received in revised form 6 August 2014 Accepted 9 August 2014 Available online 12 August 2014

The aim of this work was to investigate the potential of an amphiphilic system comprising chitosan-grafted polylactide and carboxyl-functionalized polylactide acid as a carrier for the controlled release and co-release of two DNA alkylating drugs: doxorubicin and temozolomide. Polylactide and carboxyl-functionalized polylactide acid were obtained through direct melt polycondensation reaction, using methanesulfonic acid as a non-toxic initiator, and subsequently these were grafted to the chitosan backbone through a coupling reaction, utilizing 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide as a condensing agent. ATR-FTIR analysis and conductometric titration confirmed that a reaction between CS and PLA, PLACA2% and PLACA5% occurred. Chitosan-grafted-polylactide and polylactide-citric acid nanoparticles were prepared via the polyelectrolyte complex technique, applying dextran sulphate as a polyanion, and loaded with doxorubicin and temozolomide. The diameter of particles, z-potential and their relationship to temperature and pH were analysed in all formulations. Encapsulation, co-encapsulation efficiency and release studies were conducted in different physiological simulated environments and human serum. Results showed the continuous release of drugs without an initial burst in different physiological media. ã 2014 Elsevier B.V. All rights reserved.

Chemical compounds studied in this article: Chitosan (PubChem CID:71853) 2-hydroxypropanoic acid (PubChem CID:612) Temozolomide (PubChem CID:5394) Doxorubicin (PubChem CID:31703) Keywords: Nanoparticles Drug delivery System Doxorubicin Temozolomide Chitosan Polylactide

1. Introduction Drug delivery is the method or process of administering a pharmaceutical compound to achieve a therapeutic effect in human or animals (Tiwari et al., 2012). A drug delivery system (DDS) is defined as a formulation or device that enables introduction of the bioactive compound in the body and improves its efficacy and safety by controlling the rate, time, and place of release of the drugs in the body (Mainardes et al., 2006). This process includes administration of the therapeutic product, release of the active ingredient and the subsequent transport across the biological membranes to the site of action (Jain, 2008a,b). Several such DDS(s) have been formulated and investigated; these include liposomes, microspheres, gels, nanoparticles and others. Nanoparticles provide massive advantages as regards drug targeting, delivery and release compared to other formulations. The main goals are to improve stability in the biological

* Corresponding author. Tel.: +420 576038013. E-mail address: [email protected] (V. Sedlarik). http://dx.doi.org/10.1016/j.ijpharm.2014.08.014 0378-5173/ ã 2014 Elsevier B.V. All rights reserved.

environment, to mediate bio-distribution of active compounds, and to further drug loading, targeting, transport, release and interaction with biological barriers. The cytotoxicity of nanoparticles or their degradation products remains a major problem, and improvements in biocompatibility are the main concerns of future research. Nanoparticles can be prepared from a variety of materials, such as proteins, polysaccharides and synthetic polymers. Selecting the material is dependent on many factors, including size requirements, the properties of the drug, surface characteristics such as charge and permeability, the degree of biodegradability, biocompatibility and toxicity, the drug release profile desired and so forth (Kommareddy et al., 2005). Polymeric nanoparticles made from natural or artificial polymers have been applied as drug delivery systems with great success and they show potential for many biomedical applications. Drugs may be bound in the form of a solid solution, or through dispersion, or be physically adsorbed or chemically attached (Deo et al., 1997). Polymeric-nanoparticle DDS(s) are chiefly prepared in the following three methods: (i) dispersion of preformed polymers; (ii) polymerization of monomers; and (iii) ionic gelation or coacervation of hydrophilic polymers. Moreover, other methods

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

such as supercritical fluid technology have also been described in the literature (Rabea et al., 2003). Of the several polymers available, nanoparticles made from chitosan (CS) and poly(lactic acid) (PLA) have gained the most attention as a DDS due to the stability, simplicity and uncomplicated preparation method required, and they provide versatile routes of administration (Rabea et al., 2003). CS is a natural biodegradable copolymer consisting of N-acetyl glucosamine and D-glucosamine, obtained by deacetylation of chitin (Pacheco et al., 2011). It demonstrates good biocompatibility, biodegradability and non-toxicity. Nevertheless, application of it has been limited as a result of its poor mechanical properties and water insolubility. PLA linear aliphatic thermoplastic polyester – is a synthetic biodegradable copolymer with favourable mechanical properties. PLA nanoparticles as a carrier for drug delivery have displayed suitable characteristics in the sustained release and protection of drugs from degradation over a relatively long period (Wei et al., 2008). However, a drawback of PLA nanoparticles is the lack of compatibility with cells and blood due to their hydrophobicity (Dong and Feng, 2005; Nobs et al., 2003). CS nanoparticles modified with PLA and its derivatives have been widely reported on, in particular physical modification can improve the properties of nanoparticles in different ways. Firstly, acidic products from PLA degradation can be neutralized by CS due to alkalinity acquiring better biocompatibility (Luckachan and Pillai, 2006). Dev et al. (2010) prepared PLA–CS nanoparticles with no toxicity, highlighting the biocompatibility of the products through degradation of PLA–CS nanoparticles. Secondly, the cationic surface of the modified PLA nanoparticles can promote interaction between nanoparticles and cells. However, there are some drawbacks, in particular related to mechanical properties, which can decrease after physical modification. So as to overcome this and bring about some new properties, such as controlled degradation and biocompatibility, chemical grafting or linking PLA to CS seems to be suitable (Reis et al., 2006). Different anti-cancer drugs classified under various categories such as paclitaxel, tamoxifen, 5 fluorouracil and vincristine have been encapsulated in polymeric nanoparticles, in particular CS, in order to reduce side effects and the frequency of administration (Cho et al., 2008). Temozolomide (TMZ) and doxorubicin (DOX) represent two alkylating agents widely used in cancer therapy. TMZ, an imidazotetrazine derivative of dacarbazine, is used to treat Grade IV astrocytoma and its therapeutic benefit depends on the ability to methylate DNA, often at the N-7 or O-6 positions of guanine residues, causing the death of the tumour cell. DOX is an example of a cytotoxic anthracycline antibiotic and, following FDA approval in the 1970s, it has since been successfully applied to treat a range of cancers including various types of leukaemia, breast cancer, ovarian cancer and lymphomas (Goldstein et al.,1996). However, both drugs possess side effects. TMZ must be administered in high systemic doses to achieve therapeutic levels due to its short plasma half-life of about 1.8 hours (Zhang and Gao, 2007), while DOX creates problems related to developing multi-drug resistance and acute cardiotoxicity (Akiyoshi et al., 1996). In order to overcome these drawbacks, several studies are reported in the literature regarding encapsulation in the polymeric system, especially consisting of polycyanoacrylates for TMZ and solid lipid nanoparticles for DOX. Currently, DOX is actually available in different liposomal formulations as Doxil1 and Myocet1, while for TMZ no nanoformulation is available on the market, even if different studies are being worked on presently (Cheong et al., 2006). The majority of attempts to associate DOX or TMZ to nanoparticulate carriers have used anionic or neutral polymers. Akiyoshi et al. (1996) achieved DOX encapsulation in cholesterol-bearing pullulan hydrogel nanoparticles, though loading levels were very low (98%; doxorubicin hydrochloride 98.0–102.0%, N-(3-dimethylaminopropyl)-N0 -Ethylcarbodiimide hydrochloride, commercial grade, powder; methane sulphonic acid (MSA), N,N-diethylformamide 99% dextran sulphate (DS) (Mw 4  103 g/mol), bromothyol blue and human male serum were supplied by Sigma–Aldrich. L-Lactic acid C3H6O3; 80% water solution (optical rotation 10.6) was purchased from Lachner Neratovice, Czech Republic. The solvents – acetone C3H6O, ethanol C2H6O, sodium hydroxide, sodium phosphate, potassium phosphate, potassium hydroxide and anhydrous citric acid C6H8O7 – were bought from IPL Lukes, Uhresky Brod, Czech Republic. Chloroform CHCl3 (HPLC grade) and acetic acid CH3CO2H (HPLC grade) were purchased from Chromspec, Brno, Czech Republic. 2.1. Synthesis of PLACA Low molecular weight polylactide of up to 103 g/mol was synthesized by following a procedure described by Kucharczyk et al. (2011). In brief, different proportions of L-lactic acid and

136

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

citric acid (0, 2, 5 wt.% regarding lactic acid) were added into a double-necked flask (250 mL) equipped with a Teflon stirrer. The total mass of the mixture at the beginning of the reaction was 50 g (water not included). The flask was placed in an oil bath heated by a magnetic stirrer with heating, and connected to apparatus for distillation under reduced pressure. The dehydration step followed, at 160
C and the reduced pressure of 15 kPa for 4 h. After that the reactor was disconnected from the vacuum pump and the initiator – methanesulfonic acid – was added under stirring. The flask with the initiator of dehydrated lactic acid–citric acid was reconnected to the vacuum source and reaction continued for 24 h at 160
C. The resulting product was cooled down at room temperature and dissolved in acetone. The polymer solution was precipitated into a mixture of chilled methanol and distilled water at the ratio 1:1 (v/v). The product was filtrated, washed with methanol and dried at 45
C for 48 h. This dissolving and precipitation procedure was repeated three times. The pH of the filtrate after polymer separation was checked to ensure that un-reacted citric acid was not present in the polymer. 2.2. Characterization of polylactide-citric acid (PLACA) In order to determine the concentration of terminal carboxyl groups, the acidic number (AN), which represents the amount of KOH (in mg) necessary to neutralize 1 g of substance, was calculated by titration in methanol–dichloromethane (1:1 v/v) with 0.01 M KOH ethanol solution, and bromothyol blue was used as an indicator. The presence of carboxylic groups was proven by UV–vis spectroscopy, carried out by dissolving samples in chloroform (1 mg/mL) in the range 200–400 nm on a Helios Gamma UV–vis spectrometer (room temperature, 0.5 nm resolution in a quartz cuvette, path length 10 mm). 2.3. Synthesis of CS-g-PLA and CS-g-PLACA copolymer CS-g-PLA and CS-g-PLACA were synthesized according to a procedure described by Li et al. (2012). Briefly, 300 mg of chitosan was soaked in 30 mL of N,N-dimethyl formamide (DMF) – for 24 h, and then dissolved by adding 1% (w/v) acetic acid solution. 300 mg of PLA or PLACA, EDC and NHS (molar ratio PLA or PLACA:EDC: NHS = 1:1.5:3) was dissolved in 50 mL of chloroform; the solution was added to CS-DMF solution under vigorous stirring for 48 h at room temperature. The reaction was stopped by adjusting the solution to neutrality with 0.5 M NaOH solution. The final product was precipitated by adding excess ethanol, filtered and freeze-dried. 2.4. FTIR-ATR analysis of CS based polymers Fourier transform infrared spectroscopy-attenuated total reflectance (ATR-FTIR) analysis (Nicolet iS5 FTIR Spectrometer equipped with iD5 ATR accessory, Ge crystal, resolution 4 cm1, 64 scans) was performed on the freeze-dried samples to confirm that the coupling reaction between CS and PLA and PLACA occurred. 2.5. Determination of free amino groups in CS and CS derivatives The percentage of deacetylation degree (N) of CS and CS modified with PLA and PLACA was obtained by conductometric titration, as reported by De Alvarenga et al. (2010). In brief, 200 mg of polymer (CS, CS-g-PLA, CS-g-PLACA2%, CS-g-PLACA5%) was dissolved in HCl 0.050 M solution and titrated by TitraLine Easy SI

Analytics, Germany, using NaOH 0.160 M (0.6 mL of NaOH was added gradually). N (%) was calculated as follows:   C  ðV 2  V 1 Þ  M  100 (1) Nð%Þ ¼ m where C is the concentration of NaOH solution (mol/L); V1 and V2 represent the volumes (mL) of NaOH solution required to reach the 1st and 2nd equivalence point, respectively; M represents the molar mass of the D-glucosamine unit of CS backbone; and m is the amount of polymer (mg) (De Alvarenga et al., 2010). The amount of free amino groups per polymer chain (nNH2) was determined from the N value (Eq. (1)), considering the Mw of the CS equals 100  103 g/mol as indicated by the viscosity parameters obtained from the supplier. Knowing Mw of the D-glucosamine unit (Mw = 179 g/mol) the value of nNH2 per single chain can be approximately calculated. 2.6. Nanoparticle preparation Nanoparticles were prepared following the polyelectrolyte complexation method (PEC). The copolymer (1 mg/mL) was dissolved in the aqueous solution of acetic acid (pH 3.5), and subsequently filtered through a syringe filter (size 0.45 mm) to remove residue and dust. DS was dissolved in deionized water at various concentration (polymer to DS ratio from 0.05 to 1 (w/w)), filtered (filter size 0.45 mm) and 5 mL of the solution was added drop-wise to 5 mL of copolymer solution under stirring and slight heating. An aliquot (2 mL) was withdrawn for subsequent dynamic light scattering analysis, and the rest was freeze-dried. 2.7. Drug loading DOX (1 mg/mL; V = 1 mL) was dissolved under stirring in CH3COOH aqueous solution (pH 3.5). An aqueous solution of dextran sulphate was prepared (1 mg/mL; V = 3 mL) and added drop-wise to the solution containing the drug. The copolymer was dissolved in CH3COOH aq. pH 3.5 to obtain a solution of 1 mg/mL (V = 4 mL). The mixture containing dextran sulphate and DOX (V = 4 mL) was added drop-wise to the copolymer solution under vigorous stirring. Once a colloidal suspension had been obtained, the samples were centrifuged at 10,000  g (Hettich Zentrifuge – Universal 320) for 30 min. and the pellet freeze-dried. In the case of double-loading, DOX (1 mg/mL; V = 1 mL) was dissolved in CH3COOH aqueous solution (pH 3.5), while TMZ (1 mg/mL; V = 1 mL) was dissolved in ethanol, and kept under mild stirring conditions for 30 min. till complete solubilisation had occurred. Both solutions were subsequently mixed. An aqueous solution of dextran sulphate was prepared (1 mg/mL; V = 3 mL) and added drop-wise under vigorous stirring to the mixture containing both drugs. The copolymer was dissolved in CH3COOH aqueous solution (pH 3.5) to obtain a solution of 1 mg/mL (V = 5 mL). The mixture containing dextran sulphate, TMZ and DOX (V = 5 mL) was added drop-wise to the CS derived solution under vigorous stirring and slight heating. Once a colloidal suspension had been obtained, the samples were centrifuged at 10,000  g for 30 min. and the pellet freeze-dried (Cool SafeTM – Scanvan)

Table 1 Acid number (AN), lmax,Mw determined by GPC of PLA, PLACA2% and PLACA5%. Sample PLA PLACA2% PLACA5%

AN (mg KOH/g PLA)

lmax

Mw (g/mol)

Polydispersivity

(nm)

21.4  0.3 32.7  0.2 45.2  0.2

236 239 241

47,000 27,000 5000

2.1 2.5 4.1

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

137

2.8. Dynamic light scattering analysis (DLS) The effect of pH, DS concentration and temperature on the diameter of the nanoparticles (dp, nm) and their z-potential (mV) and electrophoretic mobility (me, mm cm/V s) were analysed through DLS by Nano ZS, Malvern, UK. Analyses were performed in triplicate on suspensions at different pH values (3.5 as prepared, 7.4 and 9). The pH value was increased by adding NaOH solution of 0.1 M. The me of nanoparticles in solution was performed at a different polymer to DS ratio (from 0.05 to 1 w/w) to characterize the surface electrical properties of the nanoparticles. The me values were obtained via the Smoluchowski equation (Eq. (2)), as it is suitable for nanoparticles of any shape dispersed in aqueous solution at any concentration. The Smoluchowski equation (Sze et al., 2003):   ee & me ¼ r 0 (2)

h

where er is the dielectric constant of the media, e0 is the permittivity of free space (C2 N1 m2), h is the dynamic viscosity of the dispersion media (Pa s) and z is the zeta potential (mV).

Fig. 1. ATR-FTIR spectra of (a) CS, (b) PLA, (c) PLACA2% and (d) PLACA5%.

2.9. Drug encapsulation efficiency

2.12. In-vitro drug release studies

Encapsulation efficiency (EE) is defined as the percentage of the mass of drug encapsulated in the polymeric carrier relative to the initial amount of the drug loaded (Yeo and Park, 2004). The encapsulation and co-encapsulation efficiencies of DOX and TMZ in aqueous solutions at different pH levels (3.5, 7.4 and 9) were determined via a UV–vis spectrophotometer from Helios g-Thermo Scientific. The absorbency of the drugs was measured at 480 nm (DOX) and 325 nm (TMZ). The amount of drug was calculated from a standard curve, prepared by measuring the intensity of absorption so as to discern the concentration of the free drugs. Encapsulation efficiency was calculated as follows:   Dt  Df (3)  100 EEð%Þ ¼ Dt

The in-vitro drug release tests were executed on all formulations. 10 mg of each sample was suspended in 10 mL of aqueous acetic acid of pH 3.5 and phosphate buffer of pH 7.4 at 37
C (on a GFL 3033 Incubator, Germany). At predetermined time intervals, an aliquot (1 mL) was withdrawn and the concentration of the drug released was monitored by a UV–vis spectrophotometer (Helios g-Thermo Scientific) at 325 nm for TMZ and 480 nm for DOX. The dissolution medium was replaced with fresh buffer to maintain the total volume. The amount of drug released (DR) was determined by the following equation:   Dt (4)  100 DRð%Þ ¼ D0

where Dt represents the total theoretical amount of the drug added (mg/mL) and Df the concentration of free drug after the encapsulation process (mg/mL). All studies were conducted in triplicate. 2.10. Stability studies

where Dt (mg) represents the amount of drug released at the time t, and D0 (mg) the amount of drug loaded. All studies were conducted in triplicate. Each experiment was performed in triplicate. The calculated concentration (C) data were evaluated by applying a first-order equation (Eq. (4)) and regression processing by the least squares method, using the Solver sub-program of Microsoft1 Office Excel 2003.

Stability studies were carried out by dissolving all formulations (1 mg/mL) in physiological buffer and human serum at 37
C. Variations in the diameters of the nanoparticles over time were followed by light scattering on a Nano ZS, from Malvern Instruments, UK. 2.11. Temperature behaviour Temperature behaviour, in the range 0–60
C, was analysed in order to understand how the system reacted when submitted to gradual warming. Measurements were taken on 2 mL of nanoparticle suspension (concentration 1 mg/mL) on a Nano ZS, from Malvern Instruments, UK, and temperature was increased automatically by 10
C and stabilized for 2 min prior to measurement.

R1 → NH 2 + HOOC – R2 → R1→NH → CO →R 2 + H2O Scheme 1. Amide bond formation between the CS amino group and carboxylic groups of PLA and PLACA.

Fig. 2. FTIR-ATR spectra of (a) CS-g-PLA, (b) CS-g-PLACA2% and (c)CS-g-PLACA5%.

138

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

analysis of the samples can support the idea of COOH attachment on PLA chains. Generally, PLA displays maximum absorbency at the wavelength (lmax) 240 nm (Leu and Chow, 2011). For the authors of the paper presented here, pure PLA showed lmax = 236 nm. The increasing content of citric acid led to slight shift of lmax towards higher values. The sample with 5% of citric acid content showed lmax = 241 nm. It is noticeable that absorption at lmax is heightened with an accordant rise in the content of citric acid. It could be considered that this phenomenon is a consequence of PLA functionalization (Kucharczyk et al., 2011). 3.2. CS-g-PLA and CS-g-PLACA structure characterization 3.2.1. ATR-FTIR analysis ATR-FTIR analysis was performed to prove the occurrence of a reaction between CS amino groups and PLA and PLACA carboxylic groups as indicated in Scheme 1. NH2 + HOOC R2 ! R1 NH CO  R2 + H2O R1 where R1 an R2 represent D-glucosamine unit and the PLA or PLCA chains, respectively.

Fig. 3. FTIR-ATR spectra of CS-g-PLA, DS and CS-g-PLA-DS complex.

C ¼ C max  ð1  ekðttlag Þ Þ

(5)

where C is the cumulative concentration (mg drug/mg polymer) of drug released at given time t (h), Cmax represents the limit value of concentration that can be released from the tested system under given conditions (mg drug/mg polymer), tlag represents the time where no release is observed, and k the kinetic constant (h1), which represents the intensity of the release from the particles at the initial time (t). 3. Results and discussions 3.1. PLACA characterization Table 1 shows the acidic number (AN), lmax and the Mw, obtained through GPC analysis, of PLA and PLA with different amounts of citric acid (PLACA). As can been seen, the highest Mw was achieved for PLA (47,000 g/mol). Increasing the amount of citric acid diminishes Mw values. In addition, the polydispersity (Mw/Mn) of the polycondensates increases from 2.1 (for PLA) to 4.1 (for PLACA5%). The incorporation of citric acid in the structure of a product through a condensation reaction can be proven by determining the carboxyl groups, which should be in excess if a reaction has taken place during product formation. The titration method can provide such information in the form of AN as presented above. The values of AN, shown in Table 1 are in accordance with assuming the occurrence of a condensation reaction between lactic acid and citric acid molecules. Despite the previously mentioned assumption of carboxyl absence in pure PLA, a low AN value was observed. It can be ascribed to the presence of a residual monomer (Kucharczyk et al., 2011). In addition, the indicator – bromothyol blue – has an equivalence point of above pH 7. The rising citric acid content causes an increase in AN. UV–vis Table 2 Percentage of deacetylation degree (N) obtained by conductometric titration (Eq. (1)) and number of free amino groups per polymer chain (nNH2) of unmodified CS, and CS grafted with PLA and PLACA. Polymer

V2–V1 (mL)

N (%)

nNH2

CS CS-g-PLA CS-g-PLACA2% CS-g-PLACA5%

6 5.4 5.4 4.6

84 71 71 61

469 395 395 341

Fig. 4. (a) pH and (b) conductivity (mS/cm) of the solution containing nanoparticles comprising CS, CS-g-PLA and CS-g-PLACA.

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

presence of DS is confirmed by the typical peak located at 842 cm1 (C O S vibration).

Table 3 Particle diameter (dp, nm) and z-potential (mV) of all formulations in an acetic acid and water solution (pH 3.5, 25
C). Sample CS-g-PLA CS-g-PLA DOX CS-g-PLA TMZ CS-g-PLA DOX TMZ CS-g-PLACA (2%) CS-g-PLACA (2%) DOX CS-g-PLACA (2%) TMZ CS-g-PLACA (2%) DOX TMZ CS-g-PLACA (5%) CS-g-PLACA (5%) DOX CS-g-PLACA (5%) TMZ CS-g-PLACA (5%) DOX TMZ

dp (nm)

z-potential

164  14 201  16 188  20 258  28 152  10 172  22 201  30 296  26 283  39 299  48 327  39 298  56

34  0.3 25  0.4 27  0.2 21  0.7 29  0.4 20  0.4 13  0.9 12  0.6 31  0.7 27  0.4 26  0.2 15  0.4

139

3.2.2. Determination of free amino groups Table 2 reports the percentage of deacetylation degree (N) obtained by Eq. (1) and the number of free amino groups per polymer chain (nNH2) per polymer chain presented in CS, CS-g-PLA and CS-g-PLACA. N was determined by conductometric titration. The conductance in (mS/cm) and pH values versus the corresponding volume of titrant is reported in Fig. 4. As can been seen, three phases are observed. The first one (Fig. 4A,B) characterized by a sharp decrease in conductance connected with neutralization of acidic HCl solution. The second one (Fig. 4B,C) refers to the neutralization of the protonated amino groups of CS and the last one corresponds to the excess of the base (Fig. 4C,D). From the intersection of the three lines (Fig. 4AB,BC,CD) it is possible to obtain two volumes (V1 and V2), from which the amount of base necessary for neutralizing the amino groups can be calculated as indicated in Eq. (1) (De Alvarenga et al., 2010). Volume V2–V1 is reduced when CS is grafted with PLA and PLACA (Table 2). The N value of pure CS is in agreement with that reported by the manufacturer (Sigma–Aldrich). The reduced N value could evidence that a reaction between the amino groups of CS and the  COOH groups of PLA and PLACA occurred. While CS-g-PLA and CS-g-PLACA2% have same N value (71%), a lower value is observed in CS-g-PLACA5%, which contains a greater amount of COOH groups (Table 1). The percentage of deacetylation degree (N) and number of free amino groups (nNH2) are reported in Table 2. It can be noticed that both N and nNH2 decrease after conjugation with PLA and PLACA indicating that the reaction occurred. CS-g-PLA and CS-g-PLACA 2% show the same nNH2 value (395) suggesting equal reaction efficiency. On the contrary, CS-g-PLACA 5% have reduced nNH2 (341) indicating enhanced reaction efficiency. The nNH2 value also gives indication of the number of positive charges located along a single polymer chain (Table 2). Both N and nNH2 are in agreement with the z-potential values reported in Table 3 where z-potential is increased with rising nNH2 value.

(mV)

As reported by Li et al. (2012) the reaction between CS amino groups and PLA carboxylic groups, as occurs in the presence of EDC and NHS, leads to the formation of an amide bond between the CS chains and PLA or PLACA. Fig. 1 shows the ATR-FTIR spectra of CS, PLA, PLACA2% and PLACA5%. In accordance with results reported the previous work (Li et al., 2012), the representative peaks of the CS backbone are as follows: 3362 cm1 ( OH stretch), 2979 cm1 ( C H), 1596 cm1 (NH2 deformation) and 1068 cm1 (C O C). The PLA and PLACA representative peaks are at 2994 cm1 (C H stretch of methyl), 1755 cm1 (CQO stretch), 1455 cm1 (CH bend of methyl) and 1184 and 1087 cm1 (C O stretch). Due to the chemical similarity of lactic acid (LA) and citric acid (CA), the differences between PLA and PLACA spectra are difficult to analyse (Kucharczyk et al., 2011). The only difference can be found in the intensity of the peak related to CQO stretching, which is slightly more intense for CS-g-PLACA5% (Fig. 1). Compared to the individual components presented in Fig. 1, the ATR-FTIR spectra of the grafting-reaction products show new peaks observed at 1455 cm1 (methyl asymmetric deformation of PLA and PLACA), 1635 cm1 (the amide bond), and 1755 cm1 (the carbonyl group of the branched PLA and PLACA) (Li et al., 2012). The spectra of CS-g-PLA and CS-g-PLACA (Fig. 2) shows peaks at 3412 cm1 (N H stretching), 3278 cm1 and 3172 cm1 (O H stretching) which are connected through the presence of CS. These results indicate that the reaction is in accordance with Scheme 1. In Fig. 3, the FTIR-ATR spectra of pure DS, CS-g-PLA-DX (1:1 w/w) and CS-g-PLACA2% (1:1 w/w) are shown. DS characteristic peaks related to the symmetrical vibration of C O S (842 cm1) and the pyranose ring (1160 cm1) can be seen (Rosca et al., 2010). CS-g-PLA and CS-g-PLACA spectra were taken on lyophilized material. This treatment of the sample causes particle aggregation followed by subsequent change in the form of polymer chain reorganization. Thus, hydrophobic PLA and PLACA chains occur on the surface of the material, as depicted in Fig. 3, track c (CS-g-PLA). However, the

3.3. Physicochemical characterization of nanoparticles Study approaches include preparing four sets of particles, i.e. without the drug, with DOX or TMZ and with both drugs at different polymer to DS ratio (0.05, 0.1, 0.2, 0.5 and 1). As expected from the literature (Yousefpour et al., 2011), all drug loaded nanoparticles exhibited an increase in diameter (dp) in comparison with the drug free particles. This was probably due to electrostatic interferences of the drugs with the polymer chains, which took place during formation of the nanoparticles as a drug, where loaded at that time. As can been seen in Table 3, the presence of  COOH influenced the z-potential as well dp values. CS-g-PLACA5% showed a favourable trend as only a slight increase in dp occurred when one or both drugs were loaded. This can be

Table 4 Dependence of dp (nm), z -potential (mV) and me (mm cm/V s) of nanoparticles on polymer/DS ratio (w/w). Each experiment was performed in triplicate; average values are reported with SD of up to 10%. Polymer DS/ratio (w/w)

1 0.5 0.2 0.1 0.05

CS

CS-g-PLA

dp (nm)

z-pot

172 164 152 126 98

47 21 10 18 43

me (mm cm/V s)

(mV) 3.62 1.75 1.17 1.42 3.23

CS-g-PLACA2%

dp (nm)

z-pot

me

z-pot (mV)

(mm cm/V s)

dp (nm)

(mV)

164 141 121 106 97

34 15 11 25 49

2.75 1.21 1.34 2.02 3.59

152 115 109 102 111

29 9 19 22 46

CS-g-PLACA5%

me

z-pot (mV)

z-pot

me

(mmcm/V s)

(mV)

(mm cm/V s)

2.34 0.72 1.53 1.77 3.71

283 185 182 167 102

31 25 12 31 49

2.51 1.95 0.97 2.51 3.42

140

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

Table 5 Encapsulation efficiencies (%) of DOX and TMZ at different pH values of release media. pH

CS-g-PLA DOX (%)

CS-g-PLACA2% DOX (%)

CS-g-PLACA5% DOX (%)

CS-g-PLA TMZ (%)

CS-g-PLACA2% TMZ (%)

CS-g-PLACA5% TMZ (%)

3.5 7.4 9

54 48 12

66 52 11

61 46 8

55 41 13

69 52 20

65 56 16

explained by electrostatic interactions occurring in the system. Due to both drugs carrying a positive charge and the greater presence of  COOH groups in ionic form ( COO) in CS-g-PLACA5%, an expansion of the systems is prevented as clearly happens in CS-g-PLA. z-potential values (12–34 mV) indicated that the nanoparticles are relatively stable, and positive values suggest that chitosan chains are located externally (Honary and Zahir, 2013). The presence of drugs in the nanoparticles also influences the z-potential values. In particular, there is a decrease in various extents where observed in accordance with the chemical structure of the polymer used. The results in Table 3 reveal that CS-g-PLACA systems are more sensitive to drug loading in comparison to CS-g-PLA, confirming the influence of COOH groups on the properties of the co-polymer. The relationship between polymer/DS ratio (w/w) versus dp, z-potential and the me of nanoparticles is reported in Table 4. It was found that the lower the polymer/DS ratio, the lower the values of dp, z-potential and me that were obtained (Table 4). When DS concentration increases, its molecules can fill the inter- and intra-molecular spaces between the polymer chain, forming ionic interactions and neutralizing the positive charges located on the CS backbone, thereby causing z-potential and me reduction. When the number of negative charges exceeds the number of positive

charges, a shift from positive to negative z-potential and me values takes place as long as the number of negative charges start to predominate. The pI (isoelectric point) of the systems falls to around 0.3 polymer/DS ratio, except in CS-g-PLACA2% where it is at around 0.4. As the charge on the CS backbone is reduced, the CS chains start to fold, giving a more compact structure characterized by lower dp values (Table 4) (Chen et al., 2007). In the case of CS-g-PLA and CS-g-PLACA, reduction in z-potential and me values occur, thereby increasing DS concentration due to a lack of free amino groups not involved in the amide bond with PLA and PLACA (Table 2). 3.4. Encapsulation efficiency (EE) Drug loading in nanoparticles is achieved in two ways, by incorporating the drug at the time of producing the particles or by adsorbing the drug after forming the particles by incubating them in the drug solution. Thus, it is evident that a large amount of drug can be entrapped by the incorporation method when compared with the adsorption technique (Reis et al., 2006). DOX and TMZ were loaded during the production process of the nanoparticles, leading to higher encapsulation and co-encapsulation efficiencies, as represented in Tables 5 and 6. CS-g-PLACA shows the highest

Table 6 Co-encapsulation efficiencies (%) of DOX and TMZ at different pH values of release media. pH

CS-g-PLA DOX TMZ

CS-g-PLACA2% DOX TMZ

CS-g-PLACA5% DOX TMZ

3.5

20% (DOX + TMZ) 53% (DOX) 47% (TMZ)

32% (DOX + TMZ) 60% (DOX) 40% (TMZ)

36% (DOX + TMZ) 69% (DOX) 31% (TMZ)

7.4

13% (DOX + TMZ) 53% (DOX) 47% (TMZ)

23% (DOX + TMZ) 62% (DOX) 38% (TMZ)

26% (DOX + TMZ) 70% (DOX) 30% (TMZ)

9

4% (DOX + TMZ) 60% (DOX) 40% (TMZ)

6% (DOX + TMZ) 69% (DOX) 31% (TMZ)

4% (DOX + TMZ) 73% (DOX) 27% (TMZ)

Fig. 5. Chemical structures of (a) DOX and (b) TMZ.

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

141

Fig. 6. Stability studies in the physiological environment at 37
C. Dependence of particle diameter (dp) on time of (a) nanoparticles loaded with DOX, (b) loaded with TMZ, (c) loaded with TMZ and DOX. (CS-g-PLA (circle), CS-g-PLACA2% (triangle), CS-g-PLACA5% (square). Each experiment was performed in triplicate; average values are reported with SD of up to 10%.

EE values for both drugs investigated at all the pH values analysed. These results indicate that the presence of  COO helps to keep the drug within the nanoparticles through electrostatic interactions. Generally, it can be observed that increasing pH values cause a reduction in EE of all the samples at a given pH. However, the EE of CS-g-PLACA samples is higher in comparison with systems based on pure PLA. The exception can be observed in DOX containing the CS-g-PLACA5% particle system. It may be explained by the bulky structure of DOX (Fig. 5).

Table 6 reports on the EE of the co-encapsulated system. As can been seen in all cases, the amount of DOX encapsulated is higher than TMZ. The difference is significant especially at neutral and alkaline pH levels. To summarize, the difference in the quantity of DOX and TMZ encapsulated indicates the influence of the different molecular structures and charges of both drugs, as shown in Fig. 5. These results are in agreement with previously published works where drug–polymer electrostatic interactions have been described (Dadsetan et al., 2013). In fact, considering the pKa value

Fig. 7. Stability studies in human serum at 37
C. Dependence of particle diameter (dp) on time of (a) nanoparticles loaded with DOX, (b) loaded with TMZ, (c) loaded with TMZ and DOX. (CS-g-PLA (circle), CS-g-PLACA2% (triangle), CS-g-PLACA5% (square). Each experiment was performed in triplicate; average values are reported with SD of up to 10%.

142

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

Table 7 Temperature–particle diameter (dp) relationship in acetic acid and water solution at pH 3.5 of DOX and TMZ. T (
C)

0 10 20 30 40 50 60

CS-g-PLA DOX

CS-g-PLACA2% DOX

CS-g-PLACA5% DOX

CS-g-PLA TMZ

CS-g-PLACA2% TMZ

CS-g-PLACA5% TMZ

dp (nm)

dp (nm)

dp (nm)

dp (nm)

dp (nm)

dp (nm)

98  8 213  17 241  19 365  29 455  36 274  22 126  10

123  7 277  17 419  25 398  24 305  18 241  14 159  9

276  23 335  28 498  41 336  28 302  25 221  18 128  11

102  7 216  14 225  15 400  26 551  36 325  21 107  7

168  12 277  20 400  29 321  23 297  22 288  21 104  8

315  25 322  25 689  54 427  33 368  29 228  18 235  18

of CS amino groups, they tend to be in neutral form at pH values higher than 6.5, resulting in a lack of polymer and polymer–drug interactions. 3.5. Stability studies One essential aspect to take into account is the colloidal stability of the nanoparticles when immersed in physiological-like media. It is known that CS and many of its derivatives easily tend towards forming aggregates at pH levels higher than 7 (Tsai et al., 2011). This is principally related to the degree of protonation of the amino groups, in particular those located on the top of the nanostructures that determine the surface charge. When the number of protonated groups decreases, due to increasing pH, the surface charge gravitates towards neutrality, causing a reduction in electrostatic repulsion forces amongst the nanoparticles, giving rise to macro structures (Lopez-Leon et al., 2005). For this purpose the authors herein investigated the stability of all formulations in the physiological environment and in human serum at 37
C. Fig. 6 shows the relation between dp and the storage time of particles that were unloaded, single-drug and both-drug-loaded in physiological solution at 37
C. As can been seen, three main phases can be identified. The first one (Phase I, 0–5 days) is characterized by a slight increase in dp (from 5% to 20% according to the system), while in the second phase (Phase II, 6–16 days) a rapid increase in dp takes place (Fig. 6). This could be due to aggregation phenomena caused by a change in the surface electrical properties of the particles. In fact, for the physiological solution, this is related to the screen effect caused by the presence of chlorine ions. In the last phase (Phase III, 17–40 days), dp is stable, showing a value of approximately 1.5 times greater. Stability studies in human serum at 37
C are reported in Fig. 7. In this case, two phases can be detected. The first one (Phase I, up to 14 days) is characterized by a significant increase in dp, related to absorption of the serum component on the surface of the nanoparticles, causing a shift towards the lower

values of z-potential, promoting aggregation phenomena (Saini et al., 2011). In the second phase (Phase II, 15–40 days), dp became stable. The behaviour shown in Figs. 6 and 7 is in agreement with results published in the literature for freestabilizer systems (Leu and Chow, 2011). Nevertheless, when using stabilizers, no aggregation appears and the long-loading stability of the polymeric particle system can be achieved (Jonassen et al., 2012). 3.6. Temperature behaviour In this section, the effect of thermal treatment on dp was analysed. Only a few studies exist on the temperature effect on polymers used as drug delivery systems (Morris et al., 2011; Sershen et al., 2000; Jeong et al., 1997). Indeed, it represents a parameter that should not be underestimated, since a subtle change in temperature can cause various rearrangements of the polymer chains, influencing the entire system including the release kinetic of the present bio-active compound. Temperature behaviour was analysed in the range 0–60
C over six steps. As shown in Tables 7 and 8, increasing the temperature causes an upward trend in dp. However, each composition is characterized by a critical temperature (Tcrit.), where dp reached the maximum value (highlighted in bold in Tables 7 and 8). Reduction in dp was observed above Tcrit., probably due to a reduction in hydrogen bonds amongst CS chains and water molecules, causing a de-swelling in the system. Systems containing CS-g-PLACA showed Tcrit. in the temperature range 20–30
C, while those comprising CS-g-PLA were found at approximately 10
C higher. These results indicate that the presence of  COO influences the thermal properties of the carrier, as they can interact with the free amino groups, avoiding the formation of the hydrogen bond with the water molecules in the release environment. One of the consequences related to this behaviour (swelling and de-swelling) might be the rapid release of the drug in correspondence with the temperature at which the Tcrit. value of dp is located (Martinez-Ruvalcaba et al., 2009; Omidian and Park, 2008).

Table 8 Temperature–particle diameter (dp) relationship in acetic acid and water solution at pH 3.5 of DOX + TMZ. T (
C)

CS-g-PLA TMZ DOX dp (nm)

CS-g-PLACA2% TMZ DOX dp (nm)

CS-g-PLACA5% TMZ DOX dp (nm)

0 10 20 30 40 50 60

119  8 125  9 267  19 301  21 502  36 488  35 300  21

137  8 146  9 331  20 400  25 316  20 308  19 311  19

289  26 346  31 600  53 425  38 321  29 316  28 304  27

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

143

Fig. 8. Release profile in PBS pH 7.4 of (a) DOX and (b) TMZ from CS-g-PLA (circle), CS-g-PLACA2% (triangle) and CS-g-PLACA5% (square). The inner pane shows release in the first 24 h. Line represents fit according with Eq. (5).

3.7. Release kinetic The release profiles of drugs from the nanoparticles depend on several parameters, such as the nature of the delivery system (the chemical structure of polymer), the drug encapsulated and

environmental factors such as pH, temperature, ionic strength and the chemical composition of the release media (Costa and Sousa, 2001). In this work, the in-vitro DOX and TMZ release and co-release profile were studied in phosphate buffer (pH 7.4) at 37
C.

Fig. 9. Cumulative co-release rate in PBS pH 7.4 at 37
C of DOX (circle) and TMZ (triangle), from (a) CS-g-PLA, (b) CS-g-PLACA2% and (c) CS-g-PLACA5%. The inner pane shows release in the first 24 h. Line represents fit according with Eq. (5).

144

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

Table 9 The parameters of release kinetic regression (Eq. (5)) for the single release of DOX and TMZ from CS-g-PLA and CS-g-PLACA systems.

CS-g-PLA DOX CS-g-PLACA2% CS-g-PLACA5% CS-g-PLA TMZ CS-g-PLACA2% CS-g-PLACA5% a

DOX DOX TMZ TMZ

Cmax (mg drug/mg polymer)

k (h1)

tlag (h)

R2

t50 (h)a

0.9 0.9 0.9 0.9 0.9 0.9

0.11 0.035 0.037 0.54 0.055 0.041

0 3 4 0 4 4

0.98 0.86 0.94 0.95 0.96 0.82

5 24 20 2 12 72

t50 represents time where 50% of the drug loaded is released.

The kinetic parameters and coefficient of determination (R2) values obtained by statistical regression analysis (Eq. (5)) of experimentally obtained data (Figs. 8 and 9) are reported in Table 9 (single-loading) and 10 (double-loading). The parameters of Eq. (5) clearly describe the release profile of the systems investigated. Generally, Cmax represents the maximum limit value of the drug content released from the particles. The kinetic parameter k reveals the intensity of the drug release at the initial stage of the process. Systems possessing burst effect show typically high k values. Some of the systems are characteristic of a certain time period between introducing the nanoparticles into the release media and actually delivering the incorporated drug. This phenomenon, known as a lag phase, is represented by the parameter tlag in Eq. (5). In the case of the single-loaded system, the release efficiency, represented by Cmax (Table 8) approaches 0.9 mg drug/mg polymer in all systems. The values of k show the significant effect of both DOX and TMZ release rate over 24 h. While CS-g-PLA based nanoparticles demonstrate a typical burst effect, where the release of 50% of DOX (t50), in Table 9, was observed even after 5 h, the CS-g-PLACA system reaches this level only after 24 h (CS-g-PLACA2%) and 20 hours (CS-g-PLACA5%), as seen in Fig. 8a. This trend was more pronounced for TMZ release (Fig. 8b). In addition, a lag phase period (tlag  3–4 h) was observed for CS-g-PLACA comprising samples. The higher the level of the COOH functionalization, the more obvious the suppression of the burst effect is observable. This is probably due to electrostatic interactions between the COO groups and the positive charged molecules of the drug, slowing diffusion towards the surface of the nanoparticles and, subsequently, in the media (Fu and Kao, 2010). Moreover, the absence of an initial burst indicates that the whole amount of drug was located inside the structure of the nanoparticles, meaning that it was protected from the outer environment. These results demonstrate that CS-g-PLACA systems behave in contrast with most results presented in the literature regarding the release of DOX or TMZ from polymeric nanoparticles, where a burst effect of even higher than 50% of the loaded drug in a few hours is recorded (Zhang and Gao, 2007; Lin et al., 2005; Jiang et al., 2013). Similar behaviour is found for the co-release of DOX and TMZ in PBS (Table 10, Fig. 9). No burst effect and a lag phase of around 2–3 h were observed (Fig. 9a and b). CS-g-PLACA systems Table 10 The parameters of co-release kinetic regression (Eq. (5)) for the release of DOX and TMZ from CS-g-PLA and CS-g-PLACA systems.

CS-g-PLA CS-g-PLACA2% CS-g-PLACA5%

a

DOX TMZ DOX TMZ DOX TMZ

Cmax (mg drug/mg polymer)

k (h1)

tlag (h)

R2

t50 (h)a

0.9 0.9 0.9 0.9 0.85 0.9

0.24 0.21 0.079 0.062 0.061 0.036

0 0 3 3 2 3

0.88 0.86 0.92 0.96 0.91 0.95

3 3 12 12 15 24

t50 represents time when 50% of the drug loaded is released.

confirmed the influence of  COO groups on the release mechanism. The effect of the extent of functionalization on the release kinetic is visible as t50 values (Table 10). However, t50 release was reached after 12 h. A rapid initial release occurred as regards the CS-g-PLA system (k = 0.24 and 0.21 for DOX and TMZ, respectively; (Fig. 8a). In both cases, CS-g-PLA and CS-g-PLACA, DOX and TMZ are released with a well-balanced trend, comparable with that obtained for the single drug (Fig. 8). It indicates that no interference occurred between the drugs during the release. 4. Conclusions Polymeric nanoparticles made from the amphiphilic polymer CS-g-PLA and CS-g-PLACA loaded with DOX and TMZ were obtained through the polyelectrolyte complexation method. Results show that the presence of COOH groups on the PLA side chains influence the dimension and temperature behaviour of the particles as well as the encapsulation efficiency and release rate of the drugs. Low molecular weight polylactide (up to 103 g/mol) and its carboxylated derivatives were obtained via a polycondensation reaction using methanesulfonic acid as the initiator, while UV–vis and acidic number analysis confirmed that a reaction had occurred. PLA and PLACA were grafted to the CS backbone by a condensation reaction in the presence of EDC, and ATR-FTIR and conductometric titration analysis confirmed linking between CS and PLA and/or PLACA. Copolymeric nanoparticles were synthesized via the polyelectrolyte complexation method in aqueous media using dextran sulphate (at different ratio with polymer) as the polyanion. DOX and TMZ were loaded into the system using the same method; nanoparticles diameter (150–300 nm) and z-potential (12–34 mV), related to the presence of one of or both the drugs inside the system, confirmed the nanometrical size of the nanoparticles and their stability in the physiological-like environment. The encapsulation and also co-encapsulation efficiency of both anti-cancer drugs, DOX and TMZ, at neutral and acidic pH (more than 50 mg/g polymer) and a good release rate characterized by the absence of an initial burst, which in CS-g-PLACA release started after 4–5 h, indicate that carboxylic groups play an important role in the release mechanism, as well as that the drugs are located in the core of the system and are well protected from the external medium. Furthermore, the ratio of how DOX and TMZ are released when loaded in the same system is completely balanced, suggesting that CS-g-PLACA is a suitable candidate as a carrier for DOX and TMZ co-administration. Acknowledgements This work was supported by Operational Programme Research and Development for Innovations, co-funded by the European Regional Development Fund and the national budget of the Czech Republic (project CZ.1.05/2.1.00/03.0111), within the framework of a project entitled Centre of Polymer Systems. Antonio Di Martino

A. Di Martino, V. Sedlarik / International Journal of Pharmaceutics 474 (2014) 134–145

would like to acknowledge the Internal Grant Agency of Tomas Bata University in Zlin for co-founding within the project IGA/FT/2014/012. References Akiyoshi, K., Taniguchi, I., Fukui, H., Sunamoto, J., 1996. Hydrogel nanoparticles formed by self-assembly of hydrophobized polysaccharide. Stabilization of Adriamycin by complexation. Eur. J. Pharm. Biopharm. 42, 286–290. Chen, Y., Vellore, J., Wang, F., Benson, H.A.E., 2007. Designing chitosan-dextran sulfate nanoparticles using charge ratios. AAPS PharmSciTech 8, 4. Cheong, I., Huang, X., Bettegowda, C., Zhou, K.W., Vogelstein, S., 2006. A bacterial protein enhances the release and efficacy of liposomal cancer drugs. Science 314 (5803), 1308–1311. Cho, K., Wang, X., Nie, S., Chen, Z., Dong, M.S., 2008. Therapeutic nanoparticles for drug delivery in cancer. Clin. Cancer Res. 14, 1310–1316. Costa, P., Sousa, L.M.J., 2001. Review: modelling and comparison of dissolution profiles. Eur. J. Pharm. Sci. 13, 123–133. Dadsetan, M., Taylor, K.E., Yong, C., Bajzer, Z., Lu, L., Yaszemski, M.J., 2013. Controlled release of doxorubicin from pH responsive microgels. Acta Biomater. 9 (3), 5438–5446. De Alvarenga, E.S., De Oliveira, C.P., Bellato, C.R., 2010. An approach to understanding the deacetylation degree of chitosan. Carbohyd. Polym. 80, 1155–1160. Deo, M., Sant, V.P., Parekh, S.R., Khopade, A.J., Banakar, U.V., 1997. Proliposomebased transdermal delivery of levonorgestrel. J. Biomater. Appl. 12, 77–88. Dev, A., Binulal, N.S., Anitha, A., Nair, S.V., Furuike, T., Tamura, H., 2010. Preparation of poly(lactic acid)/chitosan nanoparticles for anti-HIV drug delivery applications. Carbohyd. Polym. 80, 833–838. Dong, Y., Feng, S.S., 2005. Poly (D,L-lactide-co-glycolide)/montmorillonite nanoparticles for oral delivery of anticancer drugs. Biomaterials 26 (30), 6068–6076. Fu, Y., Kao, W.J., 2010. Drug release kinetics and transport mechanisms of not degradable and degradable polymeric delivery systems. Expert Opin. Drug Del. 7 (4), 429–444. Goldstein, D., Goebel, F.D., Goos, M., Jablonowski, H., Stewart, J., 1996. Efficacy and safety of Stealth liposomal doxorubicin in AIDS-related Kaposi’s sarcoma. The international SL-DOX study group. Br. J. Cancer 73 (8), 989–994. Honary, S., Zahir, F., 2013. Effect of zeta potential on the properties of nano-drug delivery systems – a review (Part 2). Trop. J. Pharm. Res. 12 (2), 265–273. Jain, K.K., 2008a. Drug Delivery in CNS Disorders: Technologies, Markets and Companies. Basel: Jain PharmaBiotech Publications, Switzerland, pp. 1–310. Jain, K.K., 2008b. Transdermal Drug Delivery: Technologies, Markets and Companies. Basel: Jain PharmaBiotech Publications, Switzerland, pp. 1–230. Jeong, B., Bae, Y.H., Lee, D.S., Kim, S.W., 1997. Biodegradable block copolymers as injectable drug-delivery systems. Nature 388, 860–862. Jiang, T., Li, Y.M., Cheng, Y.J., He, F., Zhuo, R.X., 2013. Amphiphilic polycarbonate conjugates of doxorubicin with pH sensitive hydrazine linker for controlled release. Colloids Surf. B 111, 542–548. Jonassen, H., Kjoniksen, A.L., Hiorth, M., 2012. Stability of chitosan nanoparticles cross-linked with tripolyphosphate. Biomacromolecules 13 (11), 3747–3756. Kommareddy, S., Tiwari, S.B., Amiji, M.M., 2005. Long-circulating polymeric nanovectors for tumor-selective gene delivery. Technol. Cancer Res. Treat 4, 615–625. Kucharczyk, P., Poljansek, I., Sedlarik, V., Kasparkova, V., Salakova, A., Drbohlav, J., Cvelbar, U., Saha, P., 2011. Functionalization of polylactic acid through direct melt polycondensation in the presence of carboxylic acid. J. Appl. Polym. Sci. 122, 2. Leu, Y., Chow, W.S., 2011. Kinetic of water absorption and thermal properties of poly (lactic acid)/organomontmorillonite/poly(ethylene glycol) nanocomposites. J. Vinyl Addit. Technol. 17, 1. Li, J., Kong, M., Cheng, X.J., Dang, Q.F., Zhou, X., Wei, Y.N., Chen, X.G., 2012. Preparation of biocompatible chitosan grafted poly(lactic acid) nanoparticles. Int. J. Biol. Macromol. 51, 221–227. Lin, W.C., Yu, D.G., Yang, M.C., 2005. pH sensitive polyelectrolytes complex gel microspheres composed of chitosan/TPP/dextran sulfate swelling kinetics and drug delivery. Colloids Surf. B 44, 143–151.

145

Lopez-Leon, T., Carvalho, E.L.S., Seijo, B., Ortega-Vinuesa, J.L., Bastos-Gonzales, D., 2005. Physicochemical characterization of chitosan nanoparticles. Electrokinetic and stability behaviour. J. Colloid Interface Sci. 283, 344–351. Luckachan, G.E., Pillai, C.K.S., 2006. Random multiblock poly(ester amide)s containing poly(L-lactide) and cycloaliphatic amide segments: synthesis and biodegradation studies. J. Polym. Sci. A 44 (10), 3250–3260. Machinobu, T., Bito, M., Tanimura, M., Katayama, Y., Masai, E., Nakamura, M., Otsuka, Y., Ohara, S., Shigehara, K., 2010. Synthesis and characterization of hybrid biopolymers of L-lactic acid and 2-pyrone-4, 6 dicarboxylic acid. J. Macromol. Sci. A 47, 6. Mainardes, R.M., Urban, M.C., Cinto, P.O., Chaud, M.V., Evangelista, R.C., Gremieo, M. P., 2006. Liposomes and micro/nanoparticles as colloidal carriers for nasal drug delivery. Curr. Drug Deliv. 3, 275–285. Martinez-Ruvalcaba, A., Sanchez-Diaz, J.C., Becerra, F., Cruz-Barba, L.E., GonzalezAlvarez, A., 2009. Swelling characterization and drug delivery kinetics of polyacrylamide-co-itaconic acid/chitosan hydrogels. eXPRESS Polym. Lett. 3 (1), 25–32. Moon, S.-I.I., Deguchi, K., Miyamoto, M., Kimura, Y., 2004. Synthesis of poly lactic by melt/solid polycondensation of glycolic/L-lactics acids. Polym. Int. 53 (3), 254–258. Morris, A.G., Castile, J., Smith, A., Adams, G.G., Harding, S.E., 2011. The effect of prolonged storage at different temperature on the particle size distribution of tripolyphosphate (TPP)-chitosan nanoparticles. Carbohydr. Polym. 84 (4), 1430–1434. Nobs, L., Buchegger, F., Gurny, R., Allémann, E., 2003. Surface modification of poly (lactic acid) nanoparticles by covalent attachment of thiol group by means of three methods. Int. J. Pharm. 250, 327–337. Omidian, H., Park, K., 2008. Swelling agents and devices in oral drug delivery. J. Drug Deliv. Sci. Tech. 18 (2), 83–93. Pacheco, N., Gonzalez, M.G., Gimeno, M., Barzana, E., Trombotta, S., David, L., Shirai, K., 2011. Structural characterization of chitin and chitosan obtained by biological and chemical methods. Biomacromolecules 12 (9), 3285–3290. Rabea, E.I., Badaway, M.E.T., Stevens, C.V., Smagghe, G., Steurbaut, W., 2003. Chitosan as antimicrobial agent: applications and mode of action. Biomacromolecules 4, 1457–1465. Reis, C.P., Neufeld, R.J., Ribeiro, A.J., Velga, F., 2006. Pharmacology: nanoencapsulation I: methods for preparation of drug-loaded polymeric nanoparticles. Nanomedicine 2, 8–21. Rosca, C., Novac, O., Lisa, G., Popa, M.I., 2010. Polyelectrolyte complexes of chitosan with dextran sulphate, synthesis and characterization. Cell. Chem. Technol. . Saini, R.K., Srivastava, A.K., Gupta, P.K., Das, K., 2011. pH dependent reversible aggregation of chitosan and glycol-chitosan stabilized silver nanoparticles. Chem. Phys. Lett. 511 (4–6), 326–333. Sershen, S.R., Westcott, S.L., Halas, N.J., West, J.L., 2000. Temperature-sensitive polymer-nanoshell composites for photothermally modulated drug delivery. J. Biom. Mater. Res. 51 (3), 293–298. Sze, A., Erickson, D., Ren, L., Li, D., 2003. Zeta-potential measurement using Smoluchowski equation and the slope of current–time relationship in electroosmotic flow. J. Colloid Interface Sci. 261 (2), 402–410. Tiwari, G., Tiwari, R., Sriwastawa, B., Bhati, L., Pandey, P., Bannerjee, S.K., 2012. Drug delivery systems: an updated review. Int. J. Pharm. Invest. 2 (1), 2–11. Tsai, M.L., Chen, R.H., Bai, S.W., Chen, W.Y., 2011. The storage stability of chitosan/trypolyphosphate nanoparticles in a phosphate buffer. Carbohydr. Polym. 84 (2), 756–761. Wei, Q., Wei, W., Lai, B., Wang, L.Y., Wang, Y.X., Su, Z.G., Hui, G., 2008. Uniform-sized PLA nanoparticles: preparation by premix membrane emulsification. Int. J. Pharm. 359, 294–297. Yeo, Y., Park, K., 2004. Control of encapsulation efficiency and initial burst in polymeric nanoparticle synthesis. Arch. Pharm. Res. 27 (1), 1–12. Yoo, H.S., Oh, Lee, J.E., Park, K.H., 1999. Biodegradable nanoparticles containing doxorubicin-PLGA conjugate for sustained release. Pharm. Res. 16, 1114–1118. Yousefpour, P., Atyabi, F., Vasheghani-Farahani, E., Movahedi, A.A.M., Dinarvand, R., 2011. Targeted delivery of doxorubicin utilizing chitosan nanoparticle surface functionalized with anti-Her 2 trastuzumab. Int. J. Nanomed. 6, 1977–1990. Zhang, H., Gao, S., 2007. Temozolomide/PLGA microparticles and antitumor activity against glioma C6 cancer cells in vitro. Int. J. Pharm. 329, 122–128.