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Loading and Protection of Hydrophilic Molecules into LiposomeTemplated Polyelectrolyte Nanocapsules Francesca Cuomo,*,† Andrea Ceglie,† Marco Piludu,‡ Maria G. Miguel,§ Björn Lindman,§,∥ and Francesco Lopez*,† †

Dipartimento di Agricoltura, Ambiente Alimenti (DIAAA) and CSGI, Università degli Studi del Molise,Via De Sanctis, I-86100 Campobasso, Italy ‡ Dipartimento di Citomorfologia, Università di Cagliari, Cittadella Universitaria, S.S. 554 bivio Sestu, 09042-Monserrato (CA), Italy § Chemistry Department, Coimbra University, 3004-535 Coimbra, Portugal ∥ Physical Chemistry, Lund University, P.O. Box 124, 22100 Lund, Sweden ABSTRACT: Compartmentalized systems produced via the layer-bylayer (LbL) self-assembly method have been produced by alternatively depositing alginate and chitosan layers onto cores of liposomes. The combination of dynamic light scattering (DLS), ζ potential, and transmission electron microscopy (TEM) techniques provides detailed information on the stability, dimensions, charge, and wall thickness of these polyelectrolyte globules. TEM microphotographs demonstrate the presence of nanocapsules with an average diameter of below 300 nm and with a polyelectrolyte wall thickness of about 20 nm. The possibility of encapsulating and releasing molecules from this type of nanocapsule was demonstrated by loading FITC-dextrans of different molecular weights in the liposome system. The release of the loaded molecules from the nanocapsule was demonstrated after liposome core dissolution. Even at low molecular weight (20 kDa), the nanocapsules appear to be appropriate for prolonged molecule compartmentalization and protection. By means of the Ritger−Peppas model, non-Fickian transport behavior was detected for the diffusion of dextran through the polyelectrolyte wall. Values of the diffusion coefficient were calculated and yield useful information regarding chitosan/alginate hollow nanocapsules as drug-delivery systems. The influence of the pH on the release properties was also considered. The results indicate that vesicletemplated hollow polyelectrolyte nanocapsules show great potential as novel controllable drug-delivery devices for biomedical and biotechnological applications. and compositions.10−13 Recently we proposed a new method of hollow nanocapsule fabrication by combining liposomes and biocompatible chitosan/alginate multilayers via the layer-bylayer (LbL) self-assembly method.6 The self-assembly of polyelectrolytes onto nano- or microsized templates thus generates containers stabilized in solution by polyelectrolyte shells whose composition and assembly conditions influence the multishell properties.14 The complexes formed by the assembly of different polyelectrolytes behave like a new material having characteristics different from those offered by the individual polymers.15 The multilayer, indeed, is a network that controls the permeability of the capsule wall. This is a feature that makes capsules assembled via the LbL method good candidates for drug-delivery systems. Permeability, in fact, depends on the number of interactions between the polymers and on the number of assembled layers. For these reasons, the knowledge of the nature of the interactions involved in such processes becomes crucial.16−19 Materials assembled following

1. INTRODUCTION The development of material design has led to the rapid growth of novel functional hybrid materials for a variety of different applications with increasing complexity. Significant advances have been made in the development of stable multicompartmental tools with well-defined shape and size, giving responsive features relevant especially in the fields of medicine and biology.1,2 Among these, self-assembly processes provide a large number of different delivery nanodevices.3−8 For example, the synthesis of hollow polymer capsules has attracted enormous attention because of their unique structural properties and because they offer the opportunity to design multicomponent devices. Polymer capsules are basically structures with the shape of a globule produced by coating a spherical template made of different materials, such as colloidal inorganic or organic particles, latex particles, and emulsion droplets with one or more polymer layers. The use of the layer-by-layer (LbL) technique for the accomplishment of compartmentalized systems is widely recognized.9 This strategy allows the assembly of oppositely charged polyelectrolytes, on the basis of electrostatic interactions, onto several templates with different sizes, shapes, © 2014 American Chemical Society

Received: May 21, 2014 Revised: June 11, 2014 Published: June 19, 2014 7993

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performed with FITC-dextran molecules of different molecular weights, we provide new insights into the transport properties of the nanocapsules through the polyelectrolyte wall.

this approach are substantially made of an internal core, generally aqueous after the template removal, protected by polyelectrolyte layers. Depending on the strategies adopted for drug loading, a drug can be encapsulated inside the capsules before the layer deposition or after the core removal. In the first case, the drug is entrapped in the template or the template can be the drug itself.20,21 Structural and functional properties and studies on the molecular release kinetics of these devices are relevant aspects influencing the utility of such systems as drug-delivery devices. Accordingly, drug encapsulation is attracting a dramatic interest in biomedical applications22−27 since this kind of approach offers the opportunity to reduce the dose of drug administered and, consequently, the corresponding side effects of drug therapy, enhancing at the same time the therapeutic efficacy. The production of containers loaded with drugs is mainly directed toward the realization of drug-delivery systems having dimensions in the nanometer range because nanosized systems have the potential to remain in circulation for longer times compared to the microsized particles, which are subject to clearance via the lymphatic system.28 Therefore, the biological activity of nanosystems would be greater.29 Zhao et al.30 proposed a strategy to incorporate and release anticancer drugs (danorubicin and doxorubicin) in preformed microcapsules. The drugs were introduced into the device on the basis of a charge interaction mechanism. Oppositely charged poly(allylamine hydrochloride) (PAH) and poly(styrenesulfonate) (PSS) were assembled onto PSS-doped CaCO3 colloidal particles via LbL to build core−shell particles. After the removal of the carbonate cores, hollow microcapsules with entrapped PSS were obtained, which showed the spontaneous loading ability of positively charged drugs. Yan et al.31 demonstrated the loading and sustained release of 5fluorouracil, a hydrophobic molecule, in microcapsules made of poly(L-glutamic acid)/chitosan assembled onto melamine formaldehyde (MF) templates. The drug was encapsulated into the capsule core after template dissolution. Thomas et al.32 investigated the release of doxorubicin encapsulated in hollow biocompatible nanocapsules made of chitosan/heparin assembled onto SiO2 nanoparticles. The encapsulated doxorubicin was successfully internalized in MCF-7 cell lines, and in vivo experiments on BALB/c mice revealed that the drug circulation time was increased compared to that of free doxorubicin. Several models of transport mechanisms are presented in the literature, and in all of them the strong relationship between the structure and the function of the device is emphasized.33,34 Generally, the diffusion process is affected by the structural properties of the material, the physicochemical properties of the solutes, the release environment, and the possible interactions between the wall of the device and the drug molecules.35 The Higuchi model, the Ritger−Peppas model, the Weibull function, and the Hixson−Crowell model are among the most used models for quantitative studies of drug-release profiles from particles.36 These models describing the transport behavior are strictly related to Fick’s law of diffusion that offers the basis for the description of solute transport from polymeric matrices and focus mainly on two categories: Fickian and nonFickian behavior.37 In the present study, we provide further details on the structure of hollow nanocapsules and show that the vesicular template can be exploited as a drug reservoir for model watersoluble molecules. Additionally, by means of a release study

2. EXPERIMENTAL SECTION 2.1. Materials. L-α-Phosphatidylcholine (egg yolk lecithin) was purchased from Avanti Polar Lipids. Didodecyldimethylammonium bromide (DDAB), sodium chloride, low-molecular-weight sodium alginate, low-molecular-weight chitosan, Triton X-100, and FITCdextran 20, 40, and 70 kDa as well as dialysis tubing cellulose membranes, with molecular weight cutoffs of 3.5−5 and at 100 kDa, were purchased from Sigma-Aldrich. 2.2. Preparation of Liposome-Templated Chitosan/Alginate Nanocapsules. Unilamellar liposomes (80 nm) were prepared by reverse-phase evaporation according to the method described by Szoka38 with a 6.5:3.5 molar mixture of phosphatidylcholine and DDAB. Lipid and surfactant were dispersed in 3 mL of diethyl ether, and to this 1 mL of Hepes buffer (20 and 150 mM NaCl, pH 7.4) containing 10 mg/mL of FITC-dextran was added, forming a twophase system which was mixed by means of a sonicator tip to become a dispersion of inverted micelles. The majority of the organic solvent was removed by a rotary evaporator; meanwhile, the inverted micelles became an aqueous suspension of liposomes. Finally, an additional 2 mL of buffer was added and the suspension was left for an additional 45 min on the rotary evaporator to remove any trace of ether. The final lipid concentration was 16 mg mL−1 (i.e., 20 mM). The liposomes were then sequentially extruded through 0.1 and 0.05 μm polycarbonate membranes before use. Excess fluorescent dye was removed through a dialysis membrane (12 h) with a cutoff of 100 kDa against 500 mL of buffer solution (Hepes pH 7.4). The procedures of chitosan/alginate deposition onto the surface of colloidal particles and lipid core removal were reported elsewhere.6 2.3. Nanocapsule Characterization. The average hydrodynamic diameter and the ζ-potential values of the aggregates were determined by means of DLS measurements, and the electrophoretic mobilities, by laser Doppler velocimetry using a Malvern UK commercial instrument Zetasizer-Nano ZS90 operating with a 4 mW He−Ne laser (633 nm wavelength). The average aggregate size was estimated with a fixed detector angle of 90° by a cumulant analysis of the autocorrelation function using the software provided by the manufacturer. The working temperature was kept constant at 25 °C with a Peltier element integrated into the apparatus. DLS autocorrelation functions of the scattered light intensity were carried out with DTS 5.0 software provided by the manufacturer, which allowed the measurement of the distribution of the scattering intensity versus the hydrodynamic diameter. For the measurement of the ζ potential, the electrophoretic mobility of the aggregates was determined by laser Doppler velocimetry. The samples were placed in dedicated disposable capillary cells. The cells were calibrated before each set of measurements with a latex standard solution (−50 ± 5 mV). The ζ-potential values were calculated by the Smoluchowski approximation of Henry’s equation.39 2.4. Transmission Electron Microscopy. TEM high-resolution characterization of the samples was carried out placing appropriately diluted drops of each suspension on Formvar-coated 100 mesh grids for 5 min. Excess suspension was adsorbed by touching the edge of the grids with filter paper. Finally, the grids were air dried, observed, and photographed with a JEOL 100S transmission electron microscope. 2.5. FITC-Dextran Encapsulation Efficiency. The encapsulation efficiency (EE%) of FITC-dextran into liposomes and into core−shell capsules with five shells of polyelectrolytes was calculated as follows:

EE%(liposomes) amount of dextran entrapped in liposomes = × 100 amount of dextran loaded

EE%(core−shell) actual amount of dextran core−shell capsules = × 100 theoretical amount of dextran in core−shell capsules 7994

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The theoretical concentration of dextran in core−shell capsules was calculated as the expected concentration of the dye after the dilution of dye entrapped in liposomes after polyelectrolyte deposition. 2.6. FITC-Dextran Release. The samples were dialyzed through a membrane with a cutoff of 100 kDa to follow the release of FITCdextran. The bulk dialysis was prepared according to the ionic strength of the sample and with Hepes 20 mM at pH 7.4 or acetate buffer 20 mM at pH 4.5. The samples were left in the bulk to equilibrate for half an hour, and in this time window we did not observe any change in dextran concentration. Subsequently, 100 μL of Triton X-100 11 mM was added to the sample for liposome rupture, and after 20 min the sample was placed again in the dialysis tube. Twenty minutes was long enough for Triton X-100 to interact with liposomes, inducing a vesicleto-mixed micelle transition.6 At specific intervals, samples were withdrawn for spectrophotometric analysis of FITC-dextran and Triton X-100 using a Cary 100 Bio UV/vis spectrophotometer from Varian. Spectra were collected in the 800−200 nm range, at 25 °C, using quartz cells with a path length of 1 cm. Absorption spectra were corrected by fitting the spectral region between 300 and 500 nm with a power law: A = K0λ−k, where λ is the wavelength in the range defined above and k (for Rayleigh scatterers, k ≈ 4) and K0 are adjustable parameters. Then, the calculated scattering curve has been extrapolated up to 220 nm and finally subtracted from the measured absorption spectrum. The absorbance values collected were converted into concentration values according to the calibration curve previously established.

3. RESULTS AND DISCUSSION 3.1. Characterization of the Nanocapsules. Core−shell nanoparticles were assembled via LbL deposition onto

Figure 2. TEM photograph of core−shell particles (capsules with liposomes inside) with three layers (panel A) and five layers (panel B).

Figure 1. Average particle size (black dots) increasing with layer deposition and ζ-potential values (red dots) alternating with polyelectrolyte adsorption. The first values for both y axes refer to the bare liposomes.

Figure 3. TEM photograph of hollow capsules (after liposome removal) with five layers of polyelectrolytes.

liposome templates loaded or not with the FITC-dextran molecules. Liposomes were made of phosphatidylcholine, a zwitterionic lipid without a net charge, and didodecyldimethylammonium bromide (DDAB), a positively charged surfactant bearing a quaternary ammonium moiety in its headgroup. The presence of DDAB gives the liposome surface a net positive charge, which is essential for the further polyelectrolyte deposition. Negatively charged alginate and positively charged chitosan were alternatively assembled onto the templates, and the variations in size and surface charge were monitored by dynamic light scattering (DLS) and the ζ potential.6,14 Figure 1 shows the values of the diameters with the increase of layer number as well as the alternation of the ζ potential with the addition of negatively or positively charged polyelectrolytes.

After the deposition of five layers, the average diameter measured by DLS was 300 nm and the surface charge was negative since alginate was deposited as the outer layer. The growth of the capsule wall with the polyelectrolyte layer deposition and the shell thickness was investigated by Transmission Electron microscopy. In Figure 2, core−shell assemblies with three (Figure 2A) and five (Figure 2B) layers are displayed. From the photographs, it was found that core− shell particles with five layers have a shell thickness of about 20 nm while in the presence of just three layers the thickness is ca. 15 nm, thus proving the thickening of the polyelectrolyte multishell with polymer addition. For all samples, the diameters of the aggregates are in full agreement with the DLS data (Figure 1). 7995

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Figure 4. Release of 20 kDa FITC-dextran from nanocapsules with five layers of polyelectrolytes. The sketch represents the release process triggered by the addition of Triton X-100. Figure 6. Release profiles of FITC-dextran (20, 40, and 70 kDa) at pH 7.4 and 4.5 fitted to the Ritger−Peppas equation. The inset on the left shows the k values calculated from the equation. The inset on the right shows the values of n. The line drawn on the graph indicates the boundary between the values of n typical of Fickian (0.43).

investigated. After the layer deposition, the amount of dextran encapsulated into the core−shell structures with five layers was about 50% of the dextran encapsulated into the templates. This decrease can be attributed to the procedure for the preparation of core−shell particles. The samples are indeed centrifuged after the deposition of each layer to eliminate the possibility of macroaggregate formation after polyelectrolyte addition. The release of FITC-dextran from the nanocapsules was studied by dialysis experiments. Samples were placed in dialysis tubes having a cutoff of 100 kDa in order to allow the passage of dextran molecules. Dialysis was carried out versus 500 mL of buffer solution at pH 7.4 and 4.5, which correspond to the pH conditions of blood circulation and tumor tissue, respectively. The release of FITC-dextran was triggered by the addition of the nonionic surfactant Triton X-100, which causes liposome dissolution.6 This made the loaded polymer, FITC-dextran, free to diffuse from the aqueous core of the nanocapsules through the polyelectrolyte multishell (cf. sketch in Figure 4). In Figure. 4, the time-dependent normalized release of 20 kDa FITC-dextran through the nanocapsule wall is shown. From this data, two main points can be highlighted, i.e., (a) after core dissolution the multishell is still able to retain molecules at low molecular weight and (b) the release ability of these devices was clearly evident just after liposome core dissolution. The release profiles of FITC-dextran (20, 40, and 70 kDa) from nanocapsules prepared with five alternating layers of alginate and chitosan at pH 7.4 and 4.5 are shown in Figure 5A,B, respectively. The nanocapsules displayed different release rates depending on the FITC-dextran molecular weight. At first glace, a higher release rate for 20 kDa FITC-dextran is evident. At the same time, nanocapsules loaded with 40 and 70 kDa dextran show largely overlapping release profiles. One of the parameters used to compare different profiles of release is the estimation of the time necessary to release a determined

Figure 5. Time-dependent normalized release of FITC-dextran (20, 40, and 70 kDa) at pH 7.4 (A) and pH 4.5 (B). Lines corresponding to a 0.2 value of normalized release are drawn on both the graphs to underline the time necessary to release 20% of the loaded molecules.

Direct proof of the existence of hollow structures after the core dissolution from core−shell particles with five layers is reported in Figure 3, where the presence of the intact structures is clearly visible. All of the structures in the present figure, even if polydisperse in size, have a constant shell thickness of about 20 nm, which is the same thickness as deduced from Figure 2B. 3.2. FITC-Dextran Release. FITC-dextran is a polyglucose functionalized with fluorescein isothiocyanate residues. FITCdextran molecules with different molecular weights (20, 40, and 70 kDa) were separately loaded into the liposomes during their preparation. The percentage of molecule retention in the liposome template was higher than 55% for all of the molecular weights 7996

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Table 1. Best Fit Values Calculated from the Fitting of the Data to the Ritger−Peppas Equation for Dextran Molecules of Different Molecular Weight 20 20 40 40 70 70

kDa kDa kDa kDa kDa kDa

pH

k

7.4 4.5 7.4 4.5 7.4 4.5

0.58 (±0. 15) 0.98 (±0.11) 0.088 (±0.015) 0.058 (±0.021) 0.13 (±0.034) 0.061 (±0.017)

n 0.728 0.588 0.805 0.914 0.718 0.863

20 20 40 40 70 70

kDa kDa kDa kDa kDa kDa

dextran dextran dextran dextran dextran dextran

7.4 4.5 7.4 4.5 7.4 4.5

Dapp (× 10−16) (cm2/s) 6.117 5.426 0.352 0.703 0.226 0.301

(±0.1706) (± 0.1042) (± 0.0094) (± 0.0306) (± 0.0750) (± 0.0087)

R2 0.9491 0.9862 0.9615 0.8941 0.9346 0.9578

percentage of drug.36 Our experimental data show that this type of nanocapsule released 20% of its 20 kDa FITC-dextran in about 2 h; on the other hand, those loaded with 40 and 70 kDa FITC-dextran are released slowly; in fact, after 8 h of dialysis, the amount of drug released is still lower than 20%. As can be seen from panel B of Figure 5, under acidic conditions (pH 4.5) the release profiles are quite similar to those detected in a neutral environment. This absence of substantial differences in the release profiles for pH 4.5 and 7.4 is advantageous if one considers that the cell uptake of particles would occur by endocytosis. This means that if the medium of delivery has slightly acidic or neutral pH conditions, then the cargo will not be lost during the administration route, particularly for highmolecular-weight drugs. After all, it has also been reported by Yan et al.31 that the release of 5-fluorouracil from poly(Lglutamic acid)/chitosan microcapsules shows similar drug release behavior in pH 1.4 and 7.4 buffers. From these results, it can be inferred that, even at low molecular weight (20 kDa), the nanocapsules appear to be appropriate for the prolonged protection of loaded drugs. In a first rough analysis, it would seem rather obvious that molecules with higher molecular weight would find a larger obstruction in passing through the polyelectrolyte network. On the other hand, a possible mechanism of pore formation based on the hydrostatic pressure difference arising as a result of the water flux expanding existing pores or creating new ones may also be taken into account.40 As mentioned above, the diffusion process is also strongly affected by the interactions between the wall of the device and the drug molecules.35 It is well known that for transport behavior Fick’s law of diffusion based on the unraveling of Fickian and non-Fickian behavior remains the main reference. In this study, an exponential approximation model was applied to fit the obtained experimental data. This exponential relation is the Ritger−Peppas model, which is a simple power law used to describe the general solute release behavior from different polymeric devices:41−44 C = kt n C0

transport behavior

0.99052 0.99436 0.99223 0.98472 0.98194 0.98512

non-Fickian non-Fickian non-Fickian non-Fickian non-Fickian non-Fickian

In this relation, C/C0 is the fractional amount of the drug released at time t, n is a diffusion exponent which indicates the release mechanism, and k is a characteristic constant of the system. For spherical geometry, n lower than 0.43 accounts for a Fickian diffusion release mechanism; values of n in the range between 0.43 and 1 indicate anomalous non-Fickian transport. The experimental data fitted to the Ritger−Peppas equation are reported in Figure 6, and the diffusion exponent n and mechanism of the release were analyzed as summarized in Table 1. Values of n given in Table 1 suggest that FITC-dextran molecules have a similar diffusion behavior through the nanocapsule walls regardless of the molecular weight and the pH values of buffer considered. In every case, indeed the values of n are higher than 0.43, indicating that non-Fickian diffusion behavior rules the release of FITC-dextrans from the capsules. If we consider other release studies from LbL capsules, it emerges that every single system loaded with different molecules has its own peculiar release behavior, depending on the assembly conditions and on the multishell composition. The n values given by Yan et al.,31 in fact, for the release mechanism of 5-fluorouracil from poly(L-glutamic acid)/ chitosan microcapsules, indicated a Fickian type of transport which is in agreement with our consideration because 5fluorouracil is a very small molecule compared to dextran. On the other hand, Shu et al.43 investigated on the release of bovine serum albumin (BSA) from capsules of water-soluble chitosan and dextran sulfate encapsulating protein on the functionalized silica template, and despite the high molecular weight of the BSA, they observed a Fickian mechanism of transport. Diffusion coefficients of dextran through the polyelectrolyte multishell were calculated with the following equation, according to the model used for other similar systems45,46

Table 2. Best-Fit Values Calculated from Fitting of the Data to Equation 2 pH

(±0.0445) (±0.0289) (±0.0264) (±0.0524) (±0.0389) (±0.0416)

R2

⎛ Dappt ⎞0.5 C = 4⎜ 2 ⎟ C0 ⎝ πh ⎠

(2)

where Dapp is the apparent diffusion coefficient of the polymer through the multishell and h is the multishell thickness whose value of about 20 nm is obtained from the TEM images. Values of the diffusion coefficient calculated from eq 2 are reported in Table 2. As can be seen from the table, there is a difference of one order of magnitude between the apparent diffusion coefficients of 20 kDa FITC-dextran and 40 and 70 kDa FITC-dextran, which is in agreement with the difference inferred from the release profiles. Combining the information collected from the release profiles and from the calculated apparent diffusion coefficients, it seems that the multishell wall works as a filter for the passage of dextran with a molecular weight of 40 kDa or higher and that the polyelectrolyte network forms a membrane

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(9) Tan, Y.; Yildiz, U. H.; Wei, W.; Waite, J. H.; Miserez, A. Layer-byLayer Polyelectrolyte Deposition: A Mechanism for Forming Biocomposite Materials. Biomacromolecules 2013, 14, 1715−1726. (10) Ibarz, G.; Dahne, L.; Donath, E.; Mohwald, H. Smart micro- and nanocontainers for storage, transport, and release. Adv. Mater. 2001, 13, 1324−1327. (11) Rivera Gil, P.; del Mercato, L. L.; del-Pino, P.; Munoz-Javier, A.; Parak, W. J. Nanoparticle-modified polyelectrolyte capsules. Nano Today 2008, 3, 12−21. (12) Shutava, T. G.; Pattekari, P. P.; Arapov, K. A.; Torchilin, V. P.; Lvov, Y. M. Architectural layer-by-layer assembly of drug nanocapsules with PEGylated polyelectrolytes. Soft Matter 2012, 8, 9418−9427. (13) Tong, W. J.; Gao, C. Y. Layer-by-layer assembled microcapsules: Fabrication, stimuli-responsivity, loading and release. Chem. J. Chin. Univ. 2008, 29, 1285−1298. (14) Cuomo, F.; Lopez, F.; Ceglie, A.; Maiuro, L.; Miguel, M. G.; Lindman, B. pH-responsive liposome-templated polyelectrolyte nanocapsules. Soft Matter 2012, 8, 4415−4420. (15) Dejugnat, C.; Sukhorukov, G. B. PH-responsive properties of hollow polyelectrolyte microcapsules templated on various cores. Langmuir 2004, 20, 7265−7269. (16) Cuomo, F.; Ceglie, A.; Lopez, F. Temperature dependence of calcium and magnesium induced caseinate precipitation in H2O and D2O. Food Chem. 2011, 126, 8−14. (17) Cuomo, F.; Mosca, M.; Murgia, S.; Avino, P.; Ceglie, A.; Lopez, F. Evidence for the role of hydrophobic forces on the interactions of nucleotide-monophosphates with cationic liposomes. J. Colloid Interface Sci. 2013, 410, 146−151. (18) Lopez, F.; Cuomo, F.; Ceglie, A.; Ambrosone, L.; Palazzo, G. Quenching and dequenching of pyrene fluorescence by nucleotide monophosphates in cationic micelles. J. Phys. Chem. B 2008, 112, 7338−7344. (19) Medronho, B.; Romano, A.; Miguel, M. G.; Stigsson, L.; Lindman, B. Rationalizing cellulose (in)solubility: reviewing basic physicochemical aspects and role of hydrophobic interactions. Cellulose 2012, 19, 581−587. (20) Agarwal, A.; Lvov, Y.; Sawant, R.; Torchilin, V. Stable nanocolloids of poorly soluble drugs with high drug content prepared using the combination of sonication and layer-by-layer technology. J. Controlled Release 2008, 128, 255−260. (21) Strydom, S. J.; Otto, D. P.; Stieger, N.; Aucamp, M. E.; Liebenberg, W.; de Villiers, M. M. Self-assembled macromolecular nanocoatings to stabilize and control drug release from nanoparticles. Powder Technol. 2014, 256, 470−476. (22) Ai, H. Layer-by-layer capsules for magnetic resonance imaging and drug delivery. Adv. Drug Delivery Rev. 2011, 63, 772−788. (23) Ariga, K.; Lvov, Y. M.; Kawakami, K.; Ji, Q.; Hill, J. P. Layer-bylayer self-assembled shells for drug delivery. Adv. Drug Delivery Rev. 2011, 63, 762−771. (24) Becker, A. L.; Johnston, A. P. R.; Caruso, F. Layer-By-LayerAssembled Capsules and Films for Therapeutic Delivery. Small 2010, 6, 1836−1852. (25) Cuomo, F.; Lopez, F.; Ceglie, A. Templated globules  applications and perspectives. Adv. Colloid Interface Sci. 2014, 205, 124−133. (26) del Mercato, L. L.; Ferraro, M. M.; Baldassarre, F.; Mancarella, S.; Greco, V.; Rinaldi, R.; Leporatti, S. Biological applications of LbL multilayer capsules: From drug delivery to sensing. Adv. Colloid Interface Sci. 2014, http://dx.doi.org/10.1016/j.cis.2014.02.014. (27) Deshmukh, P. K.; Ramani, K. P.; Singh, S. S.; Tekade, A. R.; Chatap, V. K.; Patil, G. B.; Bari, S. B. Stimuli-sensitive layer-by-layer (LbL) self-assembly systems: Targeting and biosensory applications. J. Controlled Release 2013, 166, 294−306. (28) Acharya, S.; Sahoo, S. K. PLGA nanoparticles containing various anticancer agents and tumour delivery by EPR effect. Adv. Drug Delivery Rev. 2011, 63, 170−183. (29) Zauner, W.; Ogris, M.; Wagner, E. Polylysine-based transfection systems utilizing receptor-mediated delivery. Adv. Drug Delivery Rev. 1998, 30, 97−113.

with a cutoff appropriate for the release of molecules with a molecular weight smaller than 40 kDa.

4. CONCLUSIONS In this study, we demonstrated the formation of core−shell and hollow capsules, whose structures have been studied via DLS, the ζ potential, and transmission electron microscopy. Structures with submicrometer dimensions were produced, with a maximum size of 300 nm and a shell thickness of 20 nm. A fluorescently functionalized polymer with different molecular weights was entrapped in the liposomal template. After liposome rupture, the larger part or all polymer entrapped was still retained by the multishell carrier, thus demonstrating that the capsule wall was suitable to be a container for the loaded molecules. The polymer was successively slowly released because of the existence of a high gradient of concentration between the bulk and the capsules. The analysis of the data demonstrates that dextran was released from the capsule according to a non-Fickian mechanism. The advantage of this systems is its ability to retain molecules both at slightly acidic or neutral pH, eventually making the loaded molecule available to be internalized by cells through endocytosis. Moreover, the polyelectrolyte multishell seems to have a cutoff appropriate for a faster release of molecules with molecular weight smaller than 40 kDa since molecules of 40 kDa or higher are barely released.

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected]. *E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This work was financially supported by MIUR (PRIN 2010, 2010BJ23MN) and CSGI. REFERENCES

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