Functional polymeric nanoparticles for dexamethasone loading and release

Colloids and Surfaces B: Biointerfaces 93 (2012) 59–66

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Functional polymeric nanoparticles for dexamethasone loading and release Ilaria Fratoddi a , Iole Venditti a,∗ , Cesare Cametti b , Cleofe Palocci a , Laura Chronopoulou a , Maria Marino c , Filippo Acconcia c , Maria V. Russo a a b c

Department of Chemistry, University of Rome “Sapienza”, P.le A.Moro, 5 – 00185 Rome, Italy Department of Physics, and CNR-INFM-SOFT, Unità di Roma1, University of Rome “Sapienza”, P.le A. Moro 5, 00185 Rome, Italy Department of Biology, University Roma TRE, Viale Guglielmo Marconi, 446 I-00146 Rome, Italy

a r t i c l e

i n f o

Article history: Received 1 July 2011 Received in revised form 6 December 2011 Accepted 8 December 2011 Available online 20 December 2011 Keywords: Functional copolymers Nanostructured polymers Polymeric drug Delivery systems DXM loading and release

a b s t r a c t Poly(phenylacetylene) (PPA) and poly(phenylacetylene-co-acrylic acid) (P(PA-co-AA)), nanoparticles bioconjugated with dexamethasone (DXM) during the synthesis, named PPA@DXM and P(PA-co-AA)@DXM, were prepared by a modified surfactant free emulsion method. The loading was studied as a function of different functionality grades of the copolymer and different amounts of drug, obtaining up to 90% of drug loading for P(PA-co-AA)@DXM with 8/1 PA/AA monomer ratio. The SEM images and DLS measurements showed spheres with average diameters in the range 190–500 nm, depending on the content of acrylic acid monomer units in the copolymer and of DXM loading. -potential and surface charge density of DXM loaded nanoparticles were also investigated and confirm the charge density modulation in the range 0.62–2.68 (␮C/m2 ). The results highlight the enhanced capability of our copolymer of hosting DXM, with the advantage of a control of size, surface functionality, charge and release. Moreover we demonstrate for the first time the ability of P(PA-co-AA) DXM loaded nanoparticles to be used in the apoptosis inhibition of human tumor cells (HeLa). On the basis of the results obtained by comparing the effects elicited in HeLa cells by free DXM versus DXM loaded nanoparticles we confirmed the biological efficacy of our preparation. © 2011 Elsevier B.V. All rights reserved.

1. Introduction In the last decade, significant advances have been made in the development of biocompatible polymers for drug delivery and nanobiomaterials, such as polymer-drug conjugates [1], nanoparticles [2], liposome and polymeric micelles [3] have been employed as potential carriers. This is favored by the accelerating progress of nanotechnology addressed to drug delivery, which is widely expected to change the landscape of pharmaceutical and biotechnology industries for the foreseeable future. In particular, trends of biotechnology and biomedicine trust in the development of polymers with biocompatible and therapeutic properties upon conjugation with proper drugs. So far, a variety of natural, synthetic, biodegradable and non biodegradable polymers have been investigated, [4] and among them, functional polyacetylenes [5], conducting polymers [6,7] and polyacrylates derivatives [8,9] have been proved to be excellent candidates for biomedical applications. Although to our knowledge no clinical protocol has been yet exploited on these materials, much interest has been focused towards this class of polymers. In

∗ Corresponding author. Tel.: +39 06 49913347; fax: +39 06 490324. E-mail address: [email protected] (I. Venditti). 0927-7765/$ – see front matter © 2011 Elsevier B.V. All rights reserved. doi:10.1016/j.colsurfb.2011.12.008

fact, their nanostructures offer the possibility of delivering soluble drugs and targeted substances into cells or tissues for tumor diagnosis and therapy [10]. For example, nanostructured polymer-drug conjugates are used for the intracellular delivery of large macromolecular drugs, in the endosomal trafficking pathways and in visualization of tumor sites, through the complex engineering of composite nanoparticles [11]. Moreover, small molecular weight drugs controlling the proliferation and cells differentiation, can be incorporated into biodegradable scaffolds to induce cellular differentiation and tissue remodeling [12]. The synthetic glucocorticoid dexamethasone (DXM) is a hydrophobic bioactive compound largely used in tissue engineering applications [13] This drug has been used as a potent therapeutic agent for several inflammatory diseases, as it suppresses the expression of inflammatory genes [14] and controls the differentiation of stem cells towards the osteogenic lineage [15] The local treatment of arthropathies is enhanced by the engineering of microparticles of poly(lactic-co-glycolic acid) (PLGA) with superparamagnetic iron oxide, loaded with DXM, which interact with synovial fibroblasts [16]. Besides, implants of poly(␧-caprolactone) loaded with DXM have been successfully exploited with good tolerance for prolonged and controlled intraocular release in rabbit eye [17]. This background supports the perspectives for the investigation of biocompatible polymers of natural and synthetic origin

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as suitable materials for the loading and release of DXM in pathological tissues. Due to the increasing efforts to develop polymers suitable for bio-medical applications, we have chosen to study two synthetic polymers with different properties and functionality. PPA is a biocompatible polymer, since its interaction with cultured cells evidenced the survival and normal proliferation of lymphoma macrophages [18] and polyacrylates are well known materials in biomedical applications [19]. The copolymer, poly(phenylacetylene-co-acrylic acid) (P(PA-co-AA)), contains hydrophobic units of phenylacetylene and hydrophilic units of acrylic acid that enhance the solubility thus combining the physicochemical and biological properties of both components. PPA is a luminescent polymer and could be useful to trace the drug delivery to the target site; on the other hand the presence of acrylic acid units is responsible of controllable functionalization and dimension of polymeric particles [20], allowing to obtain different loading/release capability and grafting to surfaces [21]. In this paper we describe the synthesis, characterization, loading and release of DXM interacting with copolymeric nanospheres bearing acidic functionality, P(PA-co-AA), compared with the behavior of nanostructured poly(phenylacetylene), PPA. Further, preliminary tests have been performed to investigate the biological efficacy of the DXM loaded nanoparticles in a cell system (cervix adenocarcinoma cell line, HeLa cells). 2. Materials and methods 2.1. Materials Acrylic acid (AA) (Aldrich 99% pure), dexamethasone (DXM) (Aldrich 98% pure) and potassium persulfate (KPS) (Aldrich 99.99% pure) were used as received. Phenylacetylene (PA) (Aldrich 99% pure) was distilled under reduced pressure, before use. Other solvents and materials were reagent grade (Aldrich). Deionized water, obtained from Millipore-SIMPAKOR1 (Simplicity 185), was degassed for 30 min with Argon, before use. Buffered solutions at pH 4.0, 7.4 and 10.0 (Aldrich) were used as received.

(1 cm × 1 cm) dried at 25 ◦ C, and subsequently analyzed by SEM, using a SEM LEO1450VP, on metallized samples. The polydispersity index  (PI) was through the relationship PI = Dw /Dn , where obtained di4 / di3 and Dn = di /N are the weight average diameters Dw = and the number average diameters of the particles, respectively, N is the total number of particles and di is the diameter of the ith particle [22]. We analyzed data collected directly on the SEM images from 100 measurements performed on 2 images of non overlapping regions; the uncertainties is valued by standard deviation. The DXM loaded nanoparticles were stored in water suspensions (1 mg/mL) at 4 ◦ C. To investigate the stability and degradation kinetics of polymeric nanospheres, (for example in P(PA-co-AA) with PA/AA = 8/1), the samples were maintained in buffer solution at pH 4.0, 7.4 and 10.0 for different times (2, 6, 24 h, 1 week, 1 month) and related SEM images were acquired. Dynamic light scattering (DLS) measurements were carried out on the nanoparticle aqueous suspensions of PPA, P(PA-coAA), PPA@DXM and P(PA-co-AA)@DXM (0.01–0.2 mg/mL), using a Brookhaven instrument (Brookhaven, NY, USA) equipped with a 10 mW HeNe laser at a 632.8 nm wavelength at a temperature of 25.0 ± 0.2 ◦ C. The correlation functions were collected at  = 90◦ relative to the incident beam, and delay times from 0.8 ␮s to 10 s were explored. Non-negative least-squares (NNLS) [23] or CONTIN [24] algorithms, supplied with the instrument software were used to fit correlation data. The average hydrodynamic radius of the diffusing objects was calculated from the diffusion coefficient D and the Stokes-Einstein relationship, R = (KB T)/(6␲D), where KB T is the thermal energy and  is the solvent viscosity. In this case, the polydispersity was evaluated from the ratio Pdi = ␮/2, where  is the second cumulant of the scattered light correlation function analyzed by the cumulant method and  is the average decay rate. For a monodispersed sample, this index should be zero. The electrophoretic mobility of polymeric particle aggregates was measured by laser microelectrophoresis technique, in a thermostatic cell (T = 25.0 ± 0.2 ◦ C) using a Malvern Zetasizer Mod instrument (Malvern, U.K.). Since the particle size is much larger than the Debye screening length, the -potential was calculated from the measured electrophoretic mobility u, by means of the Smoluchowski Eq. (1). [25] 3u 2ε0 εf (KD R)

2.2. Synthesis and characterization of the nanobioconjugates

ς=

The polymeric nanospheres loaded with DXM, named PPA@DXM and P(PA-co-AA)@DXM, were synthesized by the emulsion technique [20] modified by encapsulating DXM during the synthesis. PPA@DXM was achieved by the following route: 50 mL of deionized water, 1 mL (0.936 g) of PA and different amounts of DXM (10, 20, 30 mg) dissolved in 1 mL of MeOH, were mixed and degassed under Argon for 15 min. In the case of P(PA-co-AA)@DXM, two PA/AA monomer volume ratios were used, i.e. 8/1 and 5/1, by adding to the reaction mixture 0.12 or 0.20 mL of AA, respectively and DXM (10, 20, 30 mg) dissolved in 1 mL of MeOH. A solution of KPS (100 mg in 5 mL) was then added to the mixtures and the temperature was set at 80 ◦ C. After 20 h under vigorous stirring in Ar atmosphere, the reactions were stopped by opening the flask to air and light yellow emulsions were filtered, centrifuged at 5000 rpm and redispersed in deionized water several times, in order to remove all un-reacted chemicals (weight yield: PPA@DXM ∼ 60%; P(PA-co-AA)@DXM ∼ 60%). FTIR spectra have been recorded as nujol mulls or as films deposited by casting from CHCl3 solutions using ZSM5 cells, with a Bruker Vertex 70 spectrophotometer. UV–vis spectra were run with a Varian Cary 100 Scan UV–visible spectrophotometer from CHCl3 or MeOH solutions. Films of self assembled nanoparticles were prepared by casting the water/particles suspension (1 mg/mL) on glass substrates

where  and ε are the viscosity and permittivity of the aqueous phase, ε0 is the dielectric constant of free space, and f(kD R) is the Henry function, which depends upon the inverse of the Debye −1 screening length kD and the particle radius R. For R  kD , valid for relatively large particles in a medium of moderate ionic strength, the Smoluchowski limit applies and f(kD R) = 1.5. The Ohshima relationship [26] was used to calculated the surface charge of the particles from the -potential, Eq. (2):



1 KB T ln ς= ze 6˚ ln(1/˚)

(1)

 ze 2  KB T

Q 4ε0 εR

2 

(2)

where ˚ is the volume fraction of the particles, ze the elementary charge, Q the total electric charge over each particle and R its radius. 2.3. Loading and release measurements The loading analysis of PPA@DXM and P(PA-co-AA)@DXM has been performed on solutions of the nanoparticles (10 mg) dissolved in MeOH recording the variation of the absorbance at = 240 nm (DXM absorption maximum) of the UV–vis spectra. Results were expressed as the mean ± SD of at least three independent measurements. The baseline correction was determined with methanol solutions of P(PA-co-AA) or PPA and the calibration curve was

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processed by using different concentration of DMX in methanol. Drug loading percentage (DL) was calculated from relationship (3): DL =

mgDXM × 100 mgPolymer

(3)

where mgDXM were calculated from the absorbance of PPA@DXM and P(PA-co-AA)@DXM samples. Drug loading efficiency (DLE) as expressed as the percentage of the DXM amount loaded in the nanoparticles (DL) with respect to the amount of the drug added during the nanoparticle preparation (mgDXMadded ), according to the following relationship (4) [27]: DLE (%) =

mgDXMloaded × 100 mgDXMadded

(4)

The DXM release (DR) was assessed by using the UV–vis the baseline correction and the calibration curve were performed by using DXM in phosphate buffer saline (PBS). The dried samples (10 mg), have been placed in vials with PBS (10 mL, pH = 7.4) placed in a thermostatic bath (T = 37.0 ± 0.5 ◦ C), under stirring, to simulate the physiological conditions. At defined time intervals, the supernatant has been withdrawn and submitted to UV–vis analysis. The drug release was calculated from Eq. (5): DR (%) =

mgDXMs × 100 mgDXMtotal

61

In order to verify the stability of the nanomorphology in these conditions, SEM images were acquired after different times: 2 h, 1 week and 1 month, and selected images are reported in Fig. 1. At pH 4.0, P(PA-co-AA) (PA/AA = 8/1) maintained the spherical shape and dimensions, for tests performed up to 1 month (Fig. 1a, c, e). At pH 7.4, the spherical morphology changed from spheres (2 h, Fig. 1b) to fused structures (1 week, Fig. 1d), to enamel-like morphology after 1 month (Fig. 1f). With the increase of pH value up to 10.0, the loss of nanostructured features appeared more quickly, already after 2 h (data not shown). These observations suggest the feasibility of the degradation of P(PA-co-AA) morphologies in relatively short time at physiological pH values, which may be in turn responsible of an enhanced drug release. In fact, pH sensitive nanospheres of a functionalized polymethacrylate that encapsulates rhodamine or paclitaxel have been reported to change the morphology and enhance the drug release upon dissolution in mild acidic conditions, while at pH 7.4 the release is slower, about 1 week [32]. Our results, which differ from the above reported ones in that the morphology modification appears at physiological pH, indicate that P(PA-co-AA) exhibit the desirable property of being pH sensitive. 3.2. Synthesis and characterization of the bioconjugates

(5)

where mgDXMs were quantity of DXM released in the surnatant and calculated from the absorbance spectra. 2.4. Biological tests. Biological efficacy of DXM loaded nanoparticles HeLa cells were grown in DMEM medium containing 10% fetal calf serum as previously reported [28]. Cells were pre-treated with 1 ␮M free or nanoparticle loaded DXM for 24 h before the addition of 1 ␮M staurosporine, a well known pro-apoptotic agent [29] for 24 h. As control, cells were treated for the same times with equal amounts of P(PA-co-AA) nanoparticles alone. After stimulation, cells were harvested and lysed as previously described [30]. Twenty micrograms of cell lysates were loaded on 7% SDS-PAGE and transferred to nitrocellulose filters for Western blot assay [28]. Filters were treated with anti-poly-(ADP-ribose) polymerase (PARP) antibody (Santa-Cruz Biotechnology, Santa Cruz, CA, USA; 1:1000 (v/v)) and with anti-tubulin antibody (Sigma–Aldrich, St. Louis, MO, USA; 1:10,000 (v/v)) to verify the gel loading. Autoradiography and band quantization were performed as described in the literature [30]. A statistical analysis was performed using the ANOVA test with the InStat.3 software system (GraphPad Software Inc., San Diego, CA, USA). In all analyses P values less than 0.01 were considered significant. 3. Results and discussion 3.1. Morphological stability of precursor polymers Preliminary studies on the morphological stability of the precursor polymer and copolymer have been performed. Nanostructured PPA polymer and P(PA-co-AA) copolymers were prepared without the use of any surfactant or emulsifier, following literature reports [31]; this is a modified emulsion polymerization, in which the hydrophobic monomer, in the experimental conditions reported in this work, is able to give rise to the emulsion in water. When maintained in buffered solutions (pH 7.4) PPA retains its pearllike features up to at least 1 month. A different behavior was observed for P(PA-co-AA) nanostructures. When the copolymeric nanospheres were dipped into buffered solutions at room temperature at given pH values (4.0, 7.4, 10.0), some modifications of the morphology were observed.

The synthesis of polymeric nanospheres loaded with DXM, namely PPA@DXM and P(PA-co-AA)@DXM, was achieved through emulsion polymerization modified on purpose, i.e. the bioactive molecule was straightforward added into the emulsion at the beginning of the polymerization reaction. The stability of DXM in emulsion conditions (80 ◦ C, in H2 O) has been previously confirmed by UV–vis measurements. The incorporation of DXM into the polymer/copolymer nanostructures is depicted in Fig. 2. PPA polymer was taken as a reference and its properties upon interaction with DXM were compared with those of P(PA-co-AA). This copolymer was chosen considering the presence of COOH functional groups that allows to achieve nanospheres with acid functionality on the surface [20]. This peculiar feature is expected to enhance the bioconjugation with DXM. Likewise the case of drug-hydrogel interactions [33] the incorporation of DXM may occur by different ways, i.e. through the formation of a covalent bond, leading to ester-like link, or through an electrostatic link, or through hydrogen bonding. The nature of the interaction is not quite assessed yet and needs further studies. However, according to a study that reports the covalent degradable bonding of DXM to a poly(ethylene-glycol) hydrogel network functionalized with a lactide and its hydrolytic driven release [34], we may suggest that a similar type of linkage occurs involving the CH2 OH group of DXM and the COOH functionality of P(PA-co-AA). It has been reported that drug loading has an influence on the morphology and dimensions of amphiphilic copolymeric drug delivery systems [35]. Our experimental results show that the loading of DXM in PPA polymer has a slight influence on the regular spherical morphology, dimensions and polydispersity of the nanospheres (Fig. 3). In fact, the diameter of the PPA nanospheres decreased from 270 nm to 190 nm by increasing the DXM added up to 30 mg. In the case of the bioconjugates P(PA-co-AA)@DXM the dimensions are found in the range 230–500 nm depending on both the monomers ratio (PA/AA = 8/1 or 5/1) and the amount of DXM added during the synthesis (10, 20, 30 mg) (Fig. 3). In comparison with PPA polymer, it may be observed that the dimensions of the copolymeric nanospheres depend on the amount of DXM and the trend from the three materials approximate to nanospheres with diameter in the range 190–290 nm upon the higher DXM amount added (see Supporting Information for SEM images). The shrinkage effect is more evident for the functional copolymer i.e. that having partial hydrophilic character, likewise for other functionalized polymers

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Fig. 1. SEM images of copolymer P(PA-co-AA) nanospheres (PA/AA = 8/1) taken at different times and pH values. (a) t = 2 h, pH = 4.0; (b) t = 2 h, pH = 7.4; (c) t = 1 week, pH = 4.0; (d) t = 1 week, pH = 7.4; (e) t = 1 month, pH = 4.0; (f) t = 1 month, pH = 7.4. (Scale bar 2 ␮m).

[35] through a synergic effect between the acidic monomer and DXM content added. 3.3. DXM loading and release studies In order to assess the interaction of DXM in polymeric and copolymeric nanoparticles, FTIR characterization has been carried out. In particular, peaks in the range 1600–1750 cm−1 , characteristic of DXM functionality and reported in literature for different drug delivery systems [36,37], confirm the loading capacity of our copolymer (see Supporting Information). In order to investigate the loading capacity, increasing amounts of DXM added to the reaction mixture. Drug loading and loading efficiency results were evaluated; all the investigated materials exhibit a loading capability that can be modulated from 28 to 90% and DLE from 2.5 to 8.8% (see Table 1). The most favorable loading was achieved for the samples obtained with PA/AA ratio 8/1 regardless of the amount of DXM added and, more noteworthy, the maximum loading efficiency is found for the sample synthesized with the lowest amount of added DXM. Fig. 4 shows the relationship between the initial amount of DXM into the reaction mixture versus the final effective loading. The PA/AA = 8/1 ratio represents a fair compromise between the hydrophobic component (PA) of the copolymer, that favors the nanomorphology, during the emulsion process [31,38] and the hydrophilic part (AA), that provides the surface charge and functionality for the interaction with bioactive molecules as verified in our previous studies [39,40]. Moreover, the relatively bigger dimensions of this sample, however influenced by drug addition in the synthesis mixture as above discussed in Fig. 3, allow a

better loading with respect to the particles with PA/AA = 5/1 ratio. The reported results highlight the enhanced capability of our copolymer of hosting DXM, with the advantage of a control of size, surface functionality and charge through the different amounts of AA in the composition of P(PA-co-AA), with respect to other systems such as supercritical fluid impregnated chitosan scaffolds [41]. The release of DXM was tested with UV–vis measurements of the bioconjugate PPA@DXM and P(PA-co-AA)@DXM in PBS, performing withdrawals at predetermined times (see Fig. 5), in analogy with different experiments carried out on modified PLG drug delivery systems [42]. These results highlight that PPA polymer has a less efficient daily release, in comparison with the copolymeric materials. The lack of release of DXM from PPA nanoparticles can be due to the strong Table 1 Drug loading (DL) and drug loading efficiency (DLE) for nanoparticle bioconjugates. DL results were expressed as the mean ± SD of at least three independent measurements. PA/AA (v/v)

DXM added (mg)

DL (%)

*

10 20 30 10 20 30 10 20 30

28 49 75 87 89 90 65 77 87

* *

8/1 8/1 8/1 5/1 5/1 5/1 *

PPA homopolymer.

± ± ± ± ± ± ± ± ±

2 3 5 2 4 4 3 5 5

DLE (%) 28 25 25 34 27 25 88 45 30

± ± ± ± ± ± ± ± ±

2 3 5 2 4 4 3 5 5

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Fig. 4. Influence of copolymers composition on the DL in PPA (), in P(PA-coAA) = 8/1 (䊉) and in P(PA-co-AA) = 5/1 (). Error bars quote the standard deviation of at least three independent measurements. Dotted lines guide eye, only.

about 2 h for P(PA-co-AA)@DXM (PA/AA = 5/1). Also in this case, the results can be explained considering the morphological instability of these samples at pH 7.4. SEM images shown in Fig. 6 demonstrate the different behavior of 5/1 sample with respect to 8/1 sample. In fact the copolymer P(PA-co-AA)@DXM (PA/AA = 5/1) appears structured by fused spheres already after 2 h in buffered 7.4 solution (Fig. 6a), and the lack of the nanomorphology is almost complete in less than 24 h (Fig. 6e). In the case of P(PA-co-AA)@DXM (PA/AA = 8/1) the nanomorphology appears maintained until 6 h in buffered solution (Fig. 6d); after that the nanomorphology is disrupted. For copolymeric P(PA-co-AA)@DXM systems, DR = 90% is obtained after 5 days. 3.4. DLS studies

Fig. 2. Schematic representation of bioconjugated nanoparticles: PPA@DXM and P(PA-co-AA)@DXM.

entrapment of the drug into the polymer chains, whose morphology maintains the spherical shape at physiological conditions (SEM results, Supporting Information) The copolymeric samples showed an higher degree of release, after about 10 h for P(PA-co-AA)@DXM (PA/AA = 8/1) and after

Fig. 3. Influence of drug added and composition of the copolymeric nanoparticles on the average size (diameter, D), investigated by SEM. Error bars quote the PI × 10. Dotted lines guide eye, only.

A comparison of the SEM and DLS data was carried out in order to correlate size and polydispersity of the bioconjugate nanospheres by the two techniques. In Fig. 7, we show the typical distribution of nanoparticle diameters obtained by means of SEM and DLS measurements, respectively, for P(PA-co-AA)@DXM sample (PA/AA = 8/1, DXM = 30 mg). As can be seen, despite the differences in the experimental methods (solid state for SEM, water suspension for DLS) the diameters (and the size distribution, as well) evaluated

Fig. 5. Release profile of DXM from PPA@DXM () and bioconjugates P(PA-coAA)@DXM (PA/AA = 5/1 (䊉) and PA/AA = 8/1 (), DXM added = 30 mg) during 24 h. Error bars quote the standard deviation of at least three independent measurements. Dotted lines guide eye, only.

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Fig. 6. SEM images at pH 7.4 and different times of polymeric bioconjugates P(PA-co-AA)@DXM (PA/AA 5/1) with DL 87% or P(PA-co-AA)@DXM (PA/AA 8/1) with DL 90%: (a) P(PA-co-AA)@DXM (PA/AA 5/1), 2 h; (b) P(PA-co-AA)@DXM (PA/AA 8/1), 2 h; (c) P(PA-co-AA)@DXM (PA/AA 5/1), 6 h; (d) P(PA-co-AA)@DXM (PA/AA 8/1), 6 h; (e) P(PA-co-AA)@DXM (PA/AA 5/1), 24 h; (f) P(PA-co-AA)@DXM (PA/AA 8/1), 24 h. (scale bar = 1000 nm).

by the two techniques agree with each other, indicating a linear correlation, thus providing a strong support in favor of the correctness of the size measurement. The charge density of the investigated samples (Table 2) is mainly related to the dimensions of the nanospheres and on the presence/absence of DXM; relatively slight charge variations are found for the samples without DXM in DLS measurements,

Fig. 7. Correlation between the size of P(PA-co-AA) copolymers and P(PA-coAA)@DXM determined by DLS and SEM techniques. The insets show the typical size distribution of P(PA-co-AA)@DXM (sample PA/AA = 8/1, DXM = 30 mg).

(P(PA-co-AA) 8/1, D 490 nm,  1.21 ␮C/m2 and P(PA-co-AA) 5/1 D 400 nm,  0.62 ␮C/m2 ), while the bioconjugates loaded with 20 or 30 mg of DXM show a marked increase of the surface charge (1.51, 2.28 ␮C/m2 respectively). In the case samples without DXM the charge is due only to the functional COOH groups of the copolymer; whereas the higher charge of samples loaded with DXM is attributed to the further functional groups (C O and OH) of DXM which is on the surface of the nanospheres. It is interesting that while the loading is almost the same for the two copolymers, the surface charge is higher for the polymer which should have the smaller number of acidic groups, on the basis of the nominal molar ratio of the co-monomers. From DLE measurements and DLS data it is possible to asses that the DXM loading and surface charge can be tuned by means of the modified emulsion polymerization. The main feature of our material is the easy modulation of nanoparticles functionalization to obtain different release profiles: hydrophobic nanoparticles, as PPA, have a very slow release profile (∼2% in 24 h), whereas P(PA-co-AA), containing different amounts of hydrophilic functionalities, have release profiles depending by the composition (25% in 24 h for PA/AA = 8/1 ratio) as shown above in Fig. 5. These characteristics suggest that the investigated copolymers are suitable for the tailoring of hydrophobic drugs release, which is a critical topic related to polymeric matrix drug entrapment, for example poly(lactic-co-glycolic acid) [43].

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Table 2 Drug loading (DL), diameters D (SEM measurements) and D (DLS), -potential and surface charge  for P(PA-co-AA) and bioconjugate P(PA-co-AA)@DXM nanoparticles. PA/AA (v/v)

DXM added (mg)

DL (%)

D (SEM) (nm)

D (DLS) (nm)

 (mV)

 (␮C/m2 )

5/1 5/1 8/1 8/1

0 30 0 20

– 87 ± 5 – 89 ± 4

400 233 490 290

400 230 500 300

−66.0 −56.0 −71.0 −58.0

0.62 1.51 1.21 2.28

Fig. 8. Biological efficacy of nanoparticles loaded with DXM. HeLa cells were pretreated for 24 h with 1 ␮M of either DXM alone (Panel A) or DXM loaded in P(PA-co-AA) (Panel B) nanoparticles. Staurosporine (St) (1 ␮M) was then added for additional 24 h. The activation of pro-apoptotic cascade was visualized with the appearance of 85 kDa band representing the cleaved form of PARP. Tubulin was used as loading control. The upper panels indicate representative Western blots, lower panel reported the densitometric analysis with SD of three different experiments. *p < 0.01 was calculated with respect to the relative St-treated samples.

3.5. Biological efficacy of DXM loaded nanoparticles As PPA polymer has a less efficient DMX release, the biological efficacy of PPA/AA (8/1) nanoparticles loaded with DXM has been evaluated. DXM modulates diverse and somewhat opposite signaling pathways and physiological processes [44–47]. In order to evaluate the DXM loaded nanoparticles efficacy, here we choose to investigate the cytoprotective mechanism of DXM in HeLa cells. In fact, it is well known that DXM inhibits apoptosis by decreasing the activity of apoptosis effector caspase-3 and, in turn, the cleavage of its substrate PARP which represent the onset of cell apoptosis [35,43,48]. This DXM cytoprotective effect allowed us to evaluate both the efficacy and the biocompatibility of nanoparticles. The results reported in Fig. 8A confirm that the pre-treatment of HeLa cells with DXM alone significantly reduced the amount of PARP cleavage induced by the treatment of staurosporine (St), a well known pro-apoptotic agent. DXM loaded in PPA/AA (8/1) (Fig. 8B) nanoparticles still maintained the same efficacy to reduce the amount of the St-induced cleaved form of PARP. As a whole these data indicate that DXM loaded nanospheres maintain the effect in protecting cells from apoptosis than free DXM. The biological effect has been studied during 48 h, in DMEM medium; these conditions should have some influence in the morphology degradation of P(PA-co-AA) nanoparticles, thus inducing a more efficient DXM release. 4. Conclusions Bioconjugated nanoparticles of poly(phenylacetylene), PPA@DXM, and poly(phenylacetylene-co-acrylic acid), P(PA-coAA)@DXM were prepared by a modified surfactant free emulsion method, loading DXM during the synthesis. The nanoparticles show spherical shapes with diameters in the range 190–500 nm, depending on the presence of hydrophobic substituents of the polymer PPA or on the content of hydrophilic acrylic acid groups

in the copolymers P(PA-co-AA) and more noticeable on the DXM loading. Hydrophilic groups in the copolymers P(PA-co-AA) induce higher loading efficiency, a shrinkage of the average diameter of nanospheres and an enhanced surface charge. The nanostructured bioconjugates P(PA-co-AA)@DXM are pH sensitive, since their spherical morphology is disrupted when maintained in solution at physiological pH after 24 h; PPA@DXM does not show any change in the same conditions. The benefits of using these materials are in the loading capacity of poorly soluble drugs; the loading efficiency can be adapted according to the functionality degree of the copolymer and this also allows to achieve different release time profiles. Moreover, the comparison between the effects elicited in human cell line (HeLa cells) by free DXM versus DXM loaded nanoparticles confirm the biological efficacy and the biocompatibility of our preparation. Cytotoxycity tests will be object of next investigations which can envisage the use of these bioconjugates as promising materials for the delivery of poorly water soluble drugs. Acknowledgments The authors acknowledge for the financial support of this research Progetti di Ricerca, Ateneo Sapienza 2010 (C26A10ZSHC) and FARI 2010 (C26I10ETRK). Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at doi:10.1016/j.colsurfb.2011.12.008. References [1] [2] [3] [4] [5] [6] [7] [8] [9] [10] [11] [12] [13] [14] [15] [16] [17] [18] [19] [20] [21] [22]

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