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Injectable Thermally Responsive Mucoadhesive Gel for Sustained Protein Delivery Laura Mayol,*,† Marco Biondi,† Fabiana Quaglia,† Sabato Fusco,‡ Assunta Borzacchiello,§ Luigi Ambrosio,§ and Maria I. La Rotonda*,† Department of Pharmaceutical and Toxicological Chemistry, University of Naples Federico II, Via D. Montesano, 49, 800131 Naples, Italy, IIT@CRIB, Italian Institute of Technology, via Morego 30, 16163 Genoa, Italy, and Institute of Composite and Biomedical Materials, National Research Council (IMCB-CNR), P.le Tecchio, 80, 80125 Naples, Italy Received August 3, 2010; Revised Manuscript Received November 4, 2010

Poloxamer thermoresponsive gels are widely explored in controlled drug delivery. Nevertheless, these gels possess inadequate mechanical properties, poor bioadhesiveness, and high permeability to water. To overcome these issues, we blended mucoadhesive hyaluronic acid (HA) with poloxamer analogs. This study aimed to investigate the features affecting the microscopic properties of the gels, which determine their macroscopic properties and capability to control/sustain protein release. Results showed that HA hampers water-poloxamer interactions, thus, strongly influencing physicochemical properties of poloxamer gels. This leads to gels with improved mechanical properties in which the diffusion kinetics of macromolecular active molecules are drastically slowed down. Poloxamer-HA gels can sustain the delivery of proteins, such as insulin, and may allow the modulation of its release kinetics by modifying HA content within the gels in the administration sites in which the active molecule release mechanism is mainly governed by its diffusion.

Introduction Poloxamers or pluronics are triblock copolymers made of poly(ethylene oxide)-b-poly(propylene oxide)-b-poly(ethylene oxide) (PEO-PPO-PEO), widely used in many biomedical applications, such as tissue engineering1,2 and drug delivery systems,3,4 thanks to their good tolerability, low irritancy/ toxicity, and positive therapeutic effect.5-8 The amphiphilic properties of poloxamers are due to the hydrophilic nature of PEO block between 0 and 100 °C, while PPO solubility in water strongly decreases when the temperature exceeds 15 °C.9 Indeed, concentrated aqueous solutions of poloxamers form thermoresponsive hydrogels that undergo a sol-to-gel phase transition at a critical temperature, namely, lower critical gelation temperature (LCGT). Below LCGT and in a suitable concentration range, aqueous solutions of poloxamers exist as low-viscosity liquids, while, above LCGT, their viscosity sharply increases for a slight increase in temperature. A possible mechanism of poloxamer gelation is driven by the self-assembly of polymer chains in solution due to polymer-polymer interactions, which trigger the formation of a semisolid phase. In particular, with increasing temperature, PPO blocks undergo dehydration, and this in turn promotes unimer-to-micelle aggregation events if the micellar volume fraction exceeds a critical value around 0.539.10-15 When the temperature exceeds LCGT, the dehydration of the PPO block is more significant,10,16 thus, enhancing the entanglement and packing of micelles,17 which are the building blocks of a semisolid gelled phase. In particular, PEO/PPO molar ratio and poloxamer molecular weight directly affect the gelling properties of poloxamers and, * To whom correspondence should be addressed. Tel.: +39(0)81678667. E-mail: [email protected] and [email protected]. † University of Naples Federico II. ‡ IIT@CRIB. § IMCB-CNR.

hence, the LCGT. Thus, the latter can be tailored by properly selecting the suitable poloxamer(s) to trigger the sol-to-gel transition around the physiologic temperature.18 This ability, together with their amphiphilic nature, which allows them to be loaded with both hydrophilic and hydrophobic drugs, prompted their use as injectable in situ forming drug delivery systems. However, poloxamer hydrogels show many drawbacks, such as inadequate mechanical properties, poor bioadhesiveness, and high permeability to water.18,19 To overcome these issues, many attempts were done to modify poloxamer gels both by chemical and physical entanglements.4,19-22 In this context, we have previously designed thermosensitive and mucoadhesive polymeric platforms for sustained drug delivery by blending poloxamer with hyaluronic acid (HA),23 which is a biodegradable and highly biocompatible natural mucoadhesive polysaccharide, being a main constituent of the extracellular matrix of connective tissues.24 This study demonstrated that the blending of low molecular weight HA into poloxamer gels allows an improvement of viscoelastic properties and mucoadhesive force of the gels without influencing their gelation behavior. Here we aim to get insight into the features affecting microscopic properties of poloxamer/HA platforms which, consequently, determine their ability to control and sustain the release of macromolecular active molecules. Indeed, in situ forming drug delivery platforms are of special interest for the controlled release of biotechnological drugs, such as protein and peptides, which can be administered basically by parenteral routes due their poor pharmacokinetic profiles. To this aim, thermosensitive gels made of poloxamers and hyaluronic acid (HA) were prepared and, as a model protein, insulin was loaded into the obtained platforms. Gel physicochemical features were evaluated by means of differential scanning calorimetry (DSC) and thermogravimetric analysis (TGA). Moreover, transport properties of fluorescein isothiocyanate-labeled insulin (Ins-

10.1021/bm1008958  2011 American Chemical Society Published on Web 12/13/2010

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Table 1. Formulations Prepared: Bound/Free Water Percentages and Diffusion Coefficients of Ins-FITC in Poloxamer-Based Gels formulation

F68 (% w/w)

F127 (% w/w)

HA (% w/w)

bound water (% ( SD)a

free water (% ( SD)a

DIns × 108 (cm2/s ( SD)b,c

P3 P3H1 P3H2

15 15 15

15 15 15

0 1 2

39.7 ( 0.8 26.7 ( 1.6 15.9 ( 0.5

60.3 ( 0.8 73.3 ( 1.6 84.1 ( 0.5

19.95 ( 5.02 5.65 ( 4.92 2.12 ( 0.29

a SD calculated on at least three experiments. b SD calculated on at least 10 experiments. c Diffusion coefficient of fluorescein isothiocyanate-labeled insulin (Ins-FITC).

FITC) within the gels were studied by the fluorescence recovery after photobleaching (FRAP) technique. Diffusion studies of insulin within the gels, together with gel dissolution kinetics and in vitro insulin release from the gels, were performed to get insight on the release mechanism of the protein from the gels.

Materials and Methods Materials. Poloxamers (PEOa-PPOb-PEOa) are a group of amphiphilic triblock polymers, designed with variable numbers of oxyethylene (a) and oxypropylene (b) units. In this study, poloxamer F127 (a ) 100 and b ) 65) and F68 (a ) 76 and b ) 29) were obtained from Lutrol (Basf, Germany). Low molecular weight (150 kDa) HA was supplied by Fab (Abano Terme, Italy). Insulin and fluorescein isothiocyanate-labeled insulin (Ins-FITC) from bovine pancreas, trifluoroacetic acid (TFA), phosphate buffer salts, and sodium and potassium chloride were purchased from Sigma (Milano, Italy), while HPLC grade acetonitrile was supplied by Carlo Erba (Milano, Italy). Gel Preparation. Poloxamer/HA-based formulations were prepared by mixing poloxamers F127 and F68 in distilled water under continuous stirring at 4 °C until a clear solution was obtained. Afterward, HA was added at room temperature and, for rheological, FRAP, and in vitro release tests, insulin was simply dispersed into the gel. The formulations studied are reported in Table 1. Rheological Experiments. The viscoelastic properties of the insulinloaded gels (at 5 mg/mL) were investigated to assess whether the presence of the protein could influence the viscoelastic properties of the gels and their ability to undergo gelation around LCGT. Rheological tests were performed by small-amplitude oscillatory shear experiments using a rotational rheometer (Bohlin GEMINI) as previously reported.23 Briefly, experiments were performed at 37 °C in the 0.1-10 Hz oscillation frequency range, and a strain amplitude at which linear viscoelasticity is attained. The shear storage or elastic modulus (G′) as well as the shear loss or viscous modulus (G′′) were measured as a function of frequency. Thermal Analyses. Thermoanalytical tests were performed to study the interactions between water and poloxamer gels and, more specifically, to assess the influence of HA on the physicochemical features of gels. In particular, the heat involved in the solid-to-liquid phase transition of the water within poloxamer-based gels was determined by a differential scanning calorimeter (DSC; DSC Q1000, TA Instruments, U.S.A.), calibrated with a pure indium standard. The samples were placed in hermetically sealed aluminum pans, equilibrated at -40 °C, and heated to 20 °C at 2.5 °C/min under an inert atmosphere, maintained by dry nitrogen purging at a flow rate of 50.0 mL/min. The heat evolved by the fusion of water within the gels was calculated from the recorded DSC thermograms by integrating the endothermic melting peaks. To evidence the differences among the interplay between the gels and water, thermogravimetric analyses (TGA; TGA 2950 apparatus, TA Instruments, U.S.A.) were also performed. The gels were placed in the TGA instrument pan under an inert nitrogen atmosphere. Samples were equilibrated at 60 °C and kept in isothermal conditions until a constant weight was reached. To eliminate the dependence of TGA results on the initial sample mass m0, which affects experiment duration, results were normalized with respect to m0 and expressed in term of a normalized evaporation rate, defined as the ratio between water

evaporation rate (V) at time t and that at initial time point (V0). Moreover, the obtained curves were depicted as a function of a normalized time defined as the ratio t/tf, where tf is the time point corresponding to the plateau of water evaporation rate. All the experiments were run in triplicate. Diffusion Studies. The diffusion coefficients of fluorescein isothiocyanate-labeled insulin (Ins-FITC) in water and within the gels were studied using the fluorescence recovery after photobleaching (FRAP) technique.25-27 FRAP is based on the study of the kinetics of fluorescence recovery in a small region of the analyzed sample, in which fluorescence extinction (photobleaching) is caused. The homemade FRAP apparatus is composed of (i) a direct microscope (AX60 Olympus) to observe the sample; (ii) a mercury lamp (100W; USH02D Ushio), supplied with a shutter, to excite the fluorescent molecule during image acquisition; (iii) a monochromatic argon laser (488 nm; Innova 90-2) supplied with shutters and a spatial filter (100 µm; M900 Newport) to bleach the fluorescence; (iv) a charge coupled device (CCD) camera (Pentamax, Princeton Instruments) to acquire images; and (v) a PC to store and analyze images. FRAP experiments were performed as described in the following: before photobleaching, the samples are uniformly fluorescent. Then, a portion of the sample, small compared to the overall sample dimensions, is briefly exposed to a high-intensity monochromatic laser beam, which causes local photobleaching in the beam region. In this bleached area, fluorescence is gradually recovered due to the spontaneous diffusion of the molecules that did not undergo loss of fluorescence from the surrounding, nonbleached regions. The process of fluorescence redistribution is recorded as a series of digital images and diffusion coefficients are evaluated by spatial frequency analysis (SFA) of images as described elsewhere.28 Experiments are automated and controlled by procedures properly written in Metamorph software (Universal Imaging Corp.), and the images are analyzed by a specifically developed Matlab program (MathWorks, Inc.). For FRAP experiments, Ins-FITC loaded gels were prepared. Protein concentration was optimized at 0.2 mg/mL, which could allow highcontrast images, thus enabling an accurate determination of the diffusion coefficients of Ins-FITC. The nongelled poloxamers were loaded in microslides (VitroCom Inc., U.S.A.) and the temperature was set at 37 °C to induce platform gelation. Results were averaged on at least 10 experiments. Release and Dissolution Kinetics. In vitro release rate of insulin and dissolution kinetics of the obtained gels were evaluated by immersing dialysis membranes (Spectra/Por Biotech Cellular ester, molecular cut off 12 kDa) loaded with 0.5 mL of gels in 30 mL of phosphate buffer saline (PBS; 120 mM NaCl, 2.7 mM KCl, 10 mM phosphate salts; pH ) 7.4) at 37 °C. To find out whether the release mechanism was controlled by protein diffusion into the gels or by platform dissolution, gel dissolution kinetics were monitored by recovering and weighing, at predetermined time intervals, the gel-loaded membranes. As for release studies, insulin was loaded into the gels at a therapeutic dose of 5 mg/mL and, at scheduled time intervals, the release medium was withdrawn, replaced with the same volume of fresh medium, and analyzed by reversed-phase high-performance liquid chromatography (RP-HPLC). The chromatograph was equipped with a HPLC LC-10AD pump (Shimadzu, Milano, Italy), a 7725i injection valve (Rheodyne), a SPV-10A UV-vis detector (Shimadzu) set at the wavelength of 220 nm and a C-R6A integrator (Shimadzu). RP-HPLC experiments were

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Figure 1. Mechanical spectra of P3 and Insulin-loaded P3 (5 mg/ mL) at 37 °C.

carried out using a Jupiter 5 µm C18 (250 × 4.6 mm, 300 Å porosity) (Phenomenex, Klwid, USA) column. The mobile phase was a mixture of water and acetonitrile (7:3 v/v) containing 0.1% (v/v) TFA. The flow rate was set at 1 mL/min, and the run time was 10 min. Insulinloaded gel/PBS volume ratio was chosen to guarantee sink conditions. Both dissolution and release kinetics results were averaged on three independent batches. One-way analysis of variance (ANOVA) was performed on gel dissolution and protein release data; a p < 0.05 was considered significant.

Results Preliminary rheological analyses were performed on insulinloaded gels to verify that the protein did not influence gel rheological behavior. Figure 1 shows the mechanical spectra of P3 and insulin-loaded P3, obtained at 37 °C, the reported rheograms being representative of the behavior of all gel formulations. Actually, results demonstrated that the presence of the protein did not modify the capability of the platforms to undergo reverse phase gelification. Indeed, both the platforms displayed a gel-like behavior, with both moduli being roughly constant, and G′ higher than G′′, over the entire frequency range. In Figure 2, DSC thermograms of all the formulations are reported. In all cases, two distinct endothermic peaks, relative to the fusion of water, were detected around -10 and 0 °C. In particular, the first peak (around -10 °C) is associated with the fraction of water interacting with the polymer network (i.e., bound water), while the second one refers to the free water fraction, not interacting with the polymer. DSC results showed that the area associated to the first peak decreases with increasing HA content in poloxamer gel. The total heat evolved during the fusion of water (∆Htot) was calculated by integrating the two peaks, while the percentage of bound and free water was calculated as the ratio between the enthalpies associated to the first (∆H1) and the second peak (∆H2), and ∆Htot. In Table 1 the fractions of bound and free water of the formulations P3, P3H1, and P3H2 are reported. The percentage of bound (free) water was found to be decreasing (increasing) with increasing HA content. In particular, free water percentage was found to be 60.3 ( 0.8, 73.3 ( 1.6, and 84.1 ( 0.5% for P3, P3H1, and P3H2 platforms, respectively. Figure 3 reports TGA results, expressed as normalized evaporation rate (V/V0) as a function of normalized time (t/tf), as explained in Materials and Methods. TGA findings indicate that the presence of HA into gels did influence water evaporation

Figure 2. DSC thermograms of poloxamer-based platforms: (A) P3; (B) P3H1; (C) P3H2. Heating rate: 2.5 °C/min.

capability of the gels, and in particular, V/V0 was found to be decreasing with increasing HA percentage in the platform. More specifically, at time 0.5, V/V0 was found to be around 0.2 for P3H1 and P3H2 platforms, while in the case of the P3 platform, it was approximately 0.5. FRAP results indicated that the addition of HA to poloxamerbased platforms causes a significant hindrance to Ins-FITC transport within the gels. In particular, increasing HA percentage within the platforms is associated to decreasing values of InsFITC diffusion coefficients. As reported in Table 1, the diffusion coefficient of Ins-FITC in P3 platform is 1.99 ( 0.52 × 10-7 cm2/s, while in the case of P3H1 and P3H2 systems it was (5.65 ( 4.92) × 10-8 cm2/s and (2.12 ( 0.29) × 10-8 cm2/s, respectively. The dissolution kinetics of the platforms are expressed in terms of the residual weight percentage of the gels as a function of the time of immersion into PBS at 37 °C, as reported in Figure 4. After 3-4 days, the dissolution of all the platforms was substantially complete and in particular, after 24 h, all the platforms showed a mass loss higher than 50%. ANOVA tests

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Figure 3. Normalized water evaporation rate as a function of time for poloxamer-based platforms during an isothermal TGA test at 60 °C.

Figure 6. Proposed mechanism of poloxamer gelation in the presence of HA or not. HA presence into poloxamer gels hampers the interaction between water and poloxamer macromolecules, thus, favoring the interactions among macromolecular species.

Figure 4. Dissolution kinetics of the platforms P3, P3H1, and P3H2.

Figure 5. Release kinetics of insulin loaded at 5 mg/mL into platforms P3, P3H1, and P3H2.

revealed that there was not a statistically significant difference between the dissolution data of the different platforms. The release of insulin obeyed zero order release kinetics from P3, P3H1, and P3H2 platforms (r2 > 0.98 for all formulations), as shown in Figure 5. In particular, the release of insulin from P3H2 gels was slightly slower with respect to the delivery from P3 and P3H1 platforms. However, it must be pointed out that release profiles of insulin from the different platforms do not show statistically significant differences from each other.

Discussion The influence of the addition of low molecular weight HA into poloxamer gels was investigated to get an understanding

on the features affecting the physicochemical properties of the platforms which, in turn, can determine the transport properties of macromolecular active molecules into the gels in view of their possible application as sustained protein delivery systems. DSC thermograms (Figure 2) showed that the fraction of water interacting with poloxamer macromolecules is decreasing with increasing HA content in poloxamer gel. This result strongly suggests that the addition of HA into poloxamer gels hinders the interactions between water and poloxamer molecules. Moreover, results of TGA tests exhibited a decreasing evaporation rate of water from the gels with an increasing percentage of HA in the platform, thus indicating that in poloxamer-HA platforms water is strongly bound to the gel network. Indeed, in physical self-assembling gels these microscopic changes can lead to a gel with different macroscopic properties. Actually, when a polymer is dissolved in water, polymer-polymer, polymer-water, and water-water interactions are possible. For polymers exhibiting a sol-gel transition at a defined threshold temperature, namely, lower critical gelation temperature (LCGT), a temperature increase results in a negative free energy (∆G) of the system. This makes polymer-water association unfavorable, while facilitating the polymer-polymer and water-water interactions. The negative free energy is attributed to the higher entropy term (∆S) compared to the increase in the enthalpy term (∆H) in the thermodynamic relation ∆G ) ∆H - T∆S. In particular, the entropy increase is mainly due to water-water association, also known as hydrophobic effect,29 which allows the packing of polymeric micelles at LCGT. Therefore, HA presence into poloxamer gels hampers the interaction between water and poloxamer macromolecules, thus favoring poloxamer-poloxamer interactions, as schematically presented in Figure 6. More specifically, the gelation process driven by poloxamer micelle assembly and packing is facilitated, as the temperature increases. Actually, we have previously demonstrated that the addition of HA leads to a lower gelation

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temperature of the platforms and could improve the poloxamer rheological properties.23 The increase of both viscoelastic moduli values of HA-containing gels suggests possible HA interactions with poloxamer micelles through secondary bonds (such as hydrogen bonds) and the formation of larger micelle aggregates embedded into HA random coils. This hypothesis is also supported by photon correlation spectroscopy (PCS) analyses, which evidenced, in HA containing gels, the presence of aggregates having diameters higher compared to micelle dimension.23 FRAP analyses showed that the presence of HA into poloxamer gels leads to a strong decrease (more than 1 order of magnitude) of the diffusion coefficient of fluorescein isothiocyanate-labeled insulin (Ins-FITC) within the gels. Actually, protein diffusion in a gel is best described by Fick’s laws or by Stefan-Maxwell equations, which correlate the solute flux with the gradient of its chemical potential in the system. At a fixed temperature, the general expression of a solute diffusion coefficient through hydrogels is:

D ) f(r, φ, ξ) D0 Here, D0 is the diffusion coefficient of solute in the pure solvent phase, r is the size of solute molecules, φ is the polymer volume fraction within the gel, and ξ is the mesh size of the gel.30 Thus, the decrease of Ins-FITC diffusion coefficient in HA-containing gels can be ascribed to a tighter polymeric network with improved mechanical properties. Moreover, the interaction among solutes, gel polymers and solvents has an important effect on the diffusion process.30 In particular, the addition of a highly hydrophilic anionic mucopolysaccharide, such as HA, into the poloxamer gel is expected to cause secondary interactions with insulin, thus, contrasting solute transport. These results are in line with previous findings which demonstrated that HA addition into hydrogel scaffolds leads to a slower protein diffusion within gels.31,32 Insulin in vitro release kinetics from the gels evidenced the ability of the platforms to control and sustain insulin delivery for 3-4 days, which could be useful for therapies in which insulin basal level should be attained.33 Nevertheless, no statistically significant differences among the release profiles of insulin from the different gels were found, even if FRAP experiments evidenced an important effect of HA presence on protein transport. This discrepancy can be explained by taking into account that diffusion studies and in vitro release tests were performed in different experimental conditions. Actually, FRAP tests were carried out on nondiluted samples loaded with FITClabeled insulin. Contrariwise, release tests were performed in 30 mL of release medium to ensure protein sink conditions and with gels in which nonlabeled insulin was dispersed. The seemingly different findings obtained in FRAP and in vitro release results could not be ascribed to the to possible FITCinsulin interactions because in vitro release profiles of both labeled and nonlabeled insulin from all the gels were found to be basically the same (data not shown). Thus, the differences between FRAP and release tests can be reasonably ascribed to the strong dilution to which the gels undergo during in vitro release tests. This hypothesis is also supported by platform dissolution tests which evidenced that gel dissolution kinetics were very rapid and did not evidence statistically significant differences between the different platforms. This can be explained by considering that the phenomenon of poloxamer gelation is basically concentration-dependent. Therefore, when

Mayol et al.

the volume ratio between release medium and gel is high, the gel undergoes a strong dilution that causes the rapid formation and release of a sol phase. Gel dissolution results allow to deduce that, in the case of in vitro experiments, the dissolution is prevailing on the diffusive mechanism. Nevertheless, it must be pointed out that, in subcutaneous tissues, the volume ratio between biological fluids and gel is considerably lower compared to the ones used for in vitro dissolution/release experiments. Therefore, in the case of in vivo applications, the diffusive contribution is expected to prevail on the gel dissolution, and this should allow a more important sustaining of protein release. This hypothesis is also supported by some studies in which it was demonstrated that the residence time of poloxamer analogs-based gels in vivo is significantly longer compared to the one observed during in vitro membrane model experiments, after both subcutaneous and intramuscular administration.34-36

Conclusions We demonstrated that polymeric platforms made up of poloxamers and HA are potentially useful as vehicles for sustained protein delivery. Thermal analyses revealed that HA strongly affects the physicochemical properties of poloxamer gels by hampering the interactions between poloxamer molecules and water, while promoting the interactions among macromolecular species. Moreover, transport experiments showed a strong effect of HA on protein diffusion due to a substantial reinforcement of the gel in terms of improved mechanical/ mucoadhesive properties and possible secondary interactions between HA and insulin. The results of dissolution and release tests, on the contrary, suggest a weak effect of HA addition within poloxamer-based gels. This can probably be ascribed to the strong dilution imposed when setting up in vitro release experiments so that a more important contribution of HA on gel dissolution and protein release kinetics could be envisaged in the case of in vivo applications where the active molecule release mechanism is mainly governed by diffusion.

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