European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx

Contents lists available at ScienceDirect

European Journal of Pharmaceutics and Biopharmaceutics

journal homepage: www.elsevier .com/locate /e jpb

Research paper

Chitosan/glucose 1-phosphate as new stable in situ forming depotsystem for controlled drug delivery

http://dx.doi.org/10.1016/j.ejpb.2014.05.0150939-6411/� 2014 Elsevier B.V. All rights reserved.

Abbreviations: CS, chitosan; DCF, diclofenac potassium; DD, deacetylationdegree; DSC, differential scanning calorimetry; EDTA, ethylene diamine tetraaceticacid; EY, eosin Y; G0 , storage modulus; G00 , loss modulus; G1-P, Glucose-1-Phosphate; GPC, gel permeation chromatography; H&E, hematoxylin and eosin;HCl, hydrochloric acid; I.D., inner diameter; ISFD, in situ forming depot; MB,methylene blue; Mw, molecular weight; O.D., outer diameter; PBS, phosphatebuffered saline; PLA, poly lactic acid; PLGA, poly lactic-co-glycolic acid; RBC, redblood cells; s.c., subcutaneous; Tgel, gelation temperature; tgel, gelation time; WBC,white blood cells; b-GP, b-glycerophosphate.⇑ Corresponding author. University of Strasbourg, CNRS 7199, Laboratoire de

Conception et Application de Molécules Bioactives, équipe de Pharmacie Biogalé-nique, route du Rhin No. 74, F-67401 Illkirch Cedex, France. Tel.: +33 3 68 85 42 51;fax: +33 3 68 85 43 06.

E-mail address: nanton@unistra.fr (N. Anton).

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphate as new stable in situ forming depot system for controlled drug dEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

Stephanie Supper a,b,c, Nicolas Anton b,c,⇑, Julie Boisclair a, Nina Seidel a, Marc Riemenschnitter a,Catherine Curdy a, Thierry Vandamme b,c

a Novartis Pharma AG, Basel, Switzerlandb University of Strasbourg, Faculty of Pharmacy, Illkirch Cedex, Francec CNRS UMR 7199, Laboratoire de Conception et Application de Molécules Bioactives, équipe de Pharmacie Biogalénique, Illkirch Cedex, France

a r t i c l e i n f o

Article history:Received 30 January 2014Accepted in revised form 14 May 2014Available online xxxx

Keywords:ChitosanIn situ forming depotHydrogelGlucose-1-PhosphateSol/gel transitionSustained release

a b s t r a c t

Chitosan (CS)-based thermosensitive solutions that turn into semi-solid hydrogels upon injection at bodytemperature have increasingly drawn attention over the last decades as an attractive new type of in situforming depot (ISFD) drug delivery system. Despite the great potential of the standard CS/b-glycerophos-phate (b-GP) thermogelling solutions, their lack of stability over time at room temperature as well as atrefrigerated conditions renders them unsuitable as ready-to-use drug product. In the present study, weinvestigated Glucose-1-Phosphate (G1-P) as an alternative gelling agent for improving the stability ofCS-based ISFD solutions. The in vitro release performance of CS/G1-P formulations was assessed usingseveral model compounds. Furthermore, the local tolerance of subcutaneously implanted CS/G1-Phydrogels was investigated by histological examination over three weeks. The thermogelling potentialof CS/G1-P solutions, determined by rheology, is dependent on the polymer molecular weight (Mw)and concentration as well as on the G1-P concentration. Differential scanning calorimetry (DSC)measurements confirmed that sol/gel transition takes place at around body temperature and is not fullythermo-reversible. The long term storage stability was evaluated through the appearance, pH, viscosityand gelation time at 37 �C of the solution. The results emphasized an enhanced stability of the CS/G1-Psystem compared to the standard CS/b-GP. CS solution with 0.40 mmol/g G1-P is stable for at least9 months at 2–8 �C, versus less than 1 month when using b-GP as gelling agent. Furthermore, the solutionis easy to inject, as evidenced from injectability evaluation using 23–30 G needles. In vitro release exper-iments showed a sustained release over days to weeks for hydrophilic model compounds, demonstratingthereby that CS/G1-P may be suitable for the prolonged delivery of drugs. The inflammatory reactionobserved in the tissue surrounding the hydrogel in rats was a typical foreign body reaction, similar tothe one observed for CS/b-GP hydrogels. These features confirm the potential of CS/G1-P solutions asan injectable ready-to-use in situ forming hydrogel.

� 2014 Elsevier B.V. All rights reserved.

1. Introduction

Among ISFD systems, thermogelling solutions which formsemi-solid hydrogels after administration in response to tempera-ture increase from ambient to physiological temperature are ofparticular interest. Such injectable in situ forming hydrogel, basedon CS as biocompatible and biodegradable polymer, was firstdescribed by Chenite et al. in 2000 [1], who reported an aqueoussolution of CS and b-GP at physiological pH which had the propertyof turning into an semisolid gel at temperature around 37 �C andconcluded that this polyol-phosphate salt is able to turn thepolymeric solution into a temperature-controlled pH-dependent

elivery,

2 S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx

gelling system. The sol/gel transition of the CS/b-GP solution, thegelation mechanisms, and their control for potential biomedicalapplications were extensively investigated in the last decade [2–7]. These studies revealed that gelation of the CS/b-GP solution isa consequence of several intermolecular interactions. In brief, theeffect of b-GP on CS solutions (solubilized in acidic aqueousmedium) is to partially neutralize the polymer chains along withpreventing precipitation by the formation of a protective hydrationlayer around the CS molecules. Then, increasing the temperaturehas two effects: (i) it induces a proton transfer from CS to b-GP,thereby promoting CS neutralization, and (ii) it results in breakinghydrogen bonds at the origin of the protective shell cohesionaround the CS molecules. As a result, polymer chains becomehydrophobic and create physical crosslinking between each other,forming a hydrogel.

CS/glycerophosphate thermogels have been studied as drugdelivery systems for the controlled release of various types ofdrugs (e.g. low Mw hydrophilic and hydrophobic compounds,proteins and peptides) in vitro as well as in vivo and showedpotential to achieve prolonged release over days to weeks [7–11].Furthermore, the hydrogels were also reported as promising inject-able scaffolds for cell encapsulation and tissue engineering, espe-cially for bone and cartilage repair, due to the biocompatibility,biodegradability, high water content and porosity of the threedimensional matrix [12–16].

The feasibility of developing analogous in situ thermogellingsystems using other gelling agent than b-GP, like inorganicphosphate salts, polyols or polyoses, was evaluated in severalstudies, with more or less success [17–20]. However, storage sta-bility of CS-based thermogelling solutions has received only littleattention and the few results reported, based on viscosity changesover the storage period, revealed a lack of long term stability atroom temperature as well as at refrigerated conditions. The factthat gelation occurs with time even at refrigerated conditions isa concern with CS/b-GP systems since other storage or formula-tion approaches, like freezing or freeze-drying respectively, areless suitable and would not allow the thermogelling system tobe presented as a convenient ready-to-use product in pre-filledsyringes [21]. Storage stability is an essential condition in viewof a pharmaceutical application as in situ forming hydrogels,and this point still remains very challenging for such pharmaceu-tical purpose [19,22].

In a previous study, we investigated the impact of the chemicalnature of the gelling agent on the rheological properties andgelation mechanisms. G1-P salt showed very similar ability asthe commonly used b-GP, to turn CS solutions into thermosensitivesystems [23]. In addition, we showed that G1-P allowed increasingthe ordering of the protective water layer around CS at low tem-peratures, which could be likely related to the stability of the ther-mogelling solution. It followed that G1-P could be actually analternative gelling agent to b-GP in order to prepare in situ forminghydrogels with optimized physicochemical properties with realpotential for drug delivery applications. It is to be noted that oneearly previous report [23,24], presents the gelling properties ofCS/G1-P. In [24], the authors propose this system for ocular appli-cation with an approach basically different compared to the onewe adopted here, i.e. applied to in situ forming depot system fors.c. administration. Therefore, the objectives of the present studyare to evaluate (i) the feasibility of using G1-P as alternative gellingagent for CS-based ISFD system for subcutaneous (s.c.) applica-tions, through the determination of the sol/gel transition tempera-ture, gelation time at physiological temperature, gel strength andinjectability of CS/G1-P solutions; (ii) the storage stability of CS/G1-P solutions at room temperature and refrigerated conditions,in comparison with the standard CS/b-GP, (iii) the in situ formedCS/G1-P hydrogel’s ability to prolong the release of several model

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

compounds and (iv) the histological biocompatibility of the CS/G1-P hydrogel.

2. Materials and methods

2.1. Materials

Two different CS of technical grade: Chitoscience� 90/50(deacetylation degree (DD) 89.7%, Mw: 122 kDa) and Chitoscience�

90/200 (DD 89.7%, Mw: 267 kDa), and two CS of pharmaceuticalgrade: Chitoceutical� 90/50(1) (DD 90.7%, Mw: 135 kDa) andChitoceutical� 90/50(2) (DD 88.7%, Mw 149 kDa) were used in thisstudy. All CS are derived from shrimp shell and were purchasedfrom Heppe Medical Chitosan (Halle, Germany). b-Glycerophos-phate disodium salt hydrate (b-GP, Mw: 288.10 g/mol),a-D-glucose 1-phosphate disodium salt hydrate (G1-P, Mw:376.16 g/mol), methylene blue (MB), eosin Y disodium salt (EY)and phosphate buffered saline (PBS) were purchased from Sigma.Hydrochloric acid was purchased from Merck. Diclofenacpotassium (DCF) was provided from Novartis Pharma AG (Basel,Switzerland). All other reagents were of analytical grade and wereused without further purification. Ultrapure� water was obtainedusing a MilliQ� Millipore filtration system (Millipore, Molsheim,France).

2.2. Physical characterization of chitosans using triple detection GPC

All CS were characterized by triple detection GPC on a ViscotekTriple Detector Array max system (Viscotek, USA) using two Visco-Gel A6000M (mixed Bed) columns (Malvern Instruments GmbH,Germany). The setup consisted of a size exclusion chromatographconnected to a light scattering cell; a refractive index detectorand a viscometer, allowing for simultaneous determination of theabsolute polymer Mw, hydrodynamic radius and intrinsic viscosity(g). The analysis were performed at 30 �C with a 0.3 M acetic acid,0.3 M sodium acetate, 1% ethylene glycol mobile phase and using aflow rate of 0.7 mL/min and a dn/dc of 0.163 mL/g. CS sampleswere prepared at concentrations of 1 mg/mL, dissolved 24 h underagitation in the same buffer and then filtered using a 0.2 lm filterprior to analysis. The injection volume was 100 lL. Calibration ofthe system was made using a Pullulan standard (Mw = 133 kDa,[g] = 0.515 dL/g, polydispersity index Mw/Mn = 1.08) obtainedfrom Viscotek. For the evaluation of the physical data OmniSECsoftware from Viscotek was used. Each characterization of CS sam-ples was performed in duplicate.

2.3. Preparation of CS/G1-P solutions

CS solutions were obtained by dissolving 3.75 wt.% CS in hydro-chloric acid (HCl) according to a molar ratio for chitosan aminegroups/HCl equal to 0.9/1. G1-P solutions were prepared at variousconcentrations in MilliQ water. CS and G1-P solutions wereseparately cooled to 4 �C for 15 min. Then, G1-P solution wasadded drop-by-drop into the CS solution placed in an ice-bathunder magnetic stirring. The obtained CS/G1-P solution was fur-ther stirred for 15 min. The resulting formulations were stored at2–8 �C. CS/b-GP solution was prepared according to the sameprocedure.

2.4. Determination of gelation time

Gelation times of CS/G1-P thermogelling solutions weredetermined in PBS at 37 �C. After 7 min equilibration at roomtemperature, 1 mL CS/G1-P solution was injected in a 10 mL vial(inner diameter: 22 mm) containing 5 mL PBS pH 7.4 at 37 �C (in

te as new stable in situ forming depot system for controlled drug delivery,

S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx 3

a temperature-controlled bath). In order to control whether thesolution underwent gelation, the vial was inverted horizontallyevery 30 s (without removing it from the bath) over 20 min, andthe gelation time was defined as the time at which the solutionno longer flowed.

2.5. Rheological measurements

Rheological measurements were performed using a HaakeRheostress1 rheometer (Thermo Fisher Scientific, Karlsruhe) witha circulating environmental system for temperature control. Thesample volume was about 0.5 mL, introduced between a cone-plate geometry (2� cone angle, 35 mm diameter). A solvent trapwas used to minimize water evaporation during all tests. Theviscoelastic properties of the thermogelling solutions wereassessed by measuring the storage modulus (G0) and loss modulus(G00), representing the elastic behavior and the viscous behavior ofthe material, respectively. The value of the strain amplitude wasset at 1% to ensure that all measurements were carried out withinthe linear viscoelastic range. Three kinds of tests were performed:(i) Temperature sweep tests were carried out at a constantfrequency of 1 Hz, with a temperature increase from 7 to 50 �Cand a heating rate of 1 �C/min to determine gelation temperatures.(ii) Time sweep tests, to determine gelation times, were carried outat constant angular frequency of 1 Hz and constant temperature of37.00 ± 0.20 �C. The sol/gel transition temperature (Tgel) or time(tgel) corresponds to the intersection of the curves of G0 and G00.(iii) The strength of the hydrogels was evaluated by performing fre-quency sweep tests on hydrogels formed either in contact with airor in PBS. Gels were made as follows: 1 mL CS/G1-P solution waspoured in round shaped molds and was incubated 6 h, 24 h and48 h at 37 �C to allow formation of hydrogels, either in an atmo-sphere saturated in moisture to avoid water evaporation or in1 mL PBS pre-heated at 37 �C. Then, dynamic frequency sweepstests were performed on the hydrogels at 37.00 ± 0.20 �C and G0

and G00 were measured over frequencies decreasing through loga-rithmic steps from 30 to 0.1 Hz. The gel strength was evaluatedin terms of the value of G0 at low frequencies. Measurements weredone in triplicate.

2.6. DSC analysis

The sol/gel transition and thermoreversibility of a CS/G1-Psolution containing 1.5 wt.% Chitoceutical� 90/50(2) and0.40 mmol/g G1-P was analyzed using a differential scanning calo-rimeter Q2000 from TA Instruments (Alzenau, Germany). The sam-ple was hermetically sealed into a DSC sample pan and an emptypan was used as reference. The scanning range of temperaturewas 0–50 �C at a heating rate of 1.0 �C/min, followed by a coolingcycle from 50 to 0 �C at a rate of 10.0 �C/min and a second heatingcycle from 0 to 50 �C at a rate of 1.0 �C/min. All measurementswere performed in duplicate.

2.7. Stability of chitosan/gelling agent solutions

Stability tests were carried out with different CS solutions: (i) apure CS solution, 1.5 wt.% Chitoceutical� 90/50(2); two CS/G1-Psolutions (ii) 1.5 wt.% Chitoceutical� 90/50(1)/0.27 mmol/g G1-Pand (iii) 1.5 wt.% Chitoceutical� 90/50(1)/0.40 mmol/g G1-P, and(iv) a CS/b-GP solution, 1.5 wt.% Chitoceutical� 90/50(1)/0.40mmol/g b-GP. These solutions were filtered through 0.22 lm filters,filled in 2R glass vials and purged with nitrogen prior to hermeticsealing. Mean headspace oxygen content determined by frequencymodulated spectroscopy using the Lighthouse FMS-760 Instru-ment (Lighthouse Instruments, LLC) was 0.70 ± 0.07%. Sampleswere stored at refrigerated conditions (i.e. 2–8 �C) and at room

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

temperature (i.e. 20–25 �C), and analyzed after t = 1, 7, 14, 30, 60,90, 135, 180 and 270 days for the samples stored at 2–8 �C and att = 1, 7, 14, 30, 60 and 180 days for the samples stored at 20–25 �C respectively. At each time point, a panel of different charac-teristics were followed for each samples: (i) the formulation wasvisually inspected for any change in the physical appearance, i.e.,color, turbidity, consistency (liquid or gel state); (ii) the pH valuewas determined using a 691 pH meter (Metrohm, Herisau, Switzer-land); (iii) the viscosity at 20.00 ± 0.20 �C was measured byperforming rotational tests at controlled shear rates, during whichthe shear rate was kept constant for 60 s and then increased step-wise from 5 to 103 s�1. The apparent shear viscosity values werecalculated as the mean of the apparent shear viscosities determinedafter 30 s equilibration at each shear rate; (iv) the gelation time wasdetermined by a time sweep test as defined in ‘‘Rheologicalmeasurements’’. All measurements were performed in triplicate.

2.8. Evaluation of injectability

The injection force for a CS/G1-P solution containing 1.5 wt.%Chitoceutical� 90/50(1) and 0.40 mmol/g G1-P was measured usingan electronic tensile tester (Zwick Z 2.5, Ulm, Germany) and ana-lyzed with the testXpert II software (version 3.0, Zwick, Ulm, Ger-many). 1 mL thermogelling solution was injected into an emptyglass vessel, using a 1 mL syringe (1 mL Luer-Lok Tip syringe, BD,Franklin Lakes, NJ, USA). Several needles were tested: (i) 23 G(HSW FINE-JECT� 23 G � 100, Henke-Sass Wolf GmbH, Tuttlingen,Germany), (ii) 25 G (Terumo� Hypodermic Needles, 25 G � 100 ThinWall Needle, Terumo Medical Corporation, Somerset, NJ, USA) and(iii) 30 G (BD Microlance™ 3, 30 G �½00, BD, Franklin Lakes, NJ,USA). Before measurement, all syringes were equilibrated 10 minat room temperature and primed to expel the air bubbles until asmall amount of the solution came out. Injection force was mea-sured at two different injection speeds: 80 and 100 mm/min, foreach needle size. All measurements were replicated 10 times.

2.9. In vitro release studies

In vitro release tests were performed in triplicate. CS/G1-Psolutions loaded with three different model compounds: methyleneblue (MB), diclofenac potassium (DCF) and eosin Y (EY) wereprepared. The model compounds were first solubilized in the G1-Psolution prior to mixing it with the CS solution (Chitoceutical 90/50(1)). The final formulations contained 1.5 wt.% CS, 0.40 mmol/gG1-P and either 0.06 wt.% MB or 0.06 wt.% DCF or 0.06, 0.3 or0.6 wt.% EY. As positive controls, CS/MB, CS/DCF and CS/EY solutionswere prepared with the same protocol as described above but with-out including the gelling agent (i.e. G1-P), i.e. MB, DCF or EY were firstsolubilized in water and subsequently added to the CS solution.Then, 0.5 g thermogelling solution was injected into a 50 mL tubecontaining 25 mL PBS pH 7.4, in a reciprocating water bath main-tained at 37 �C. After 1 h at rest, to allow the gel formation, mildshaking (10%) was started. At defined time points, 10 mL sampleswere removed from the release medium and replaced with freshPBS to maintain the volume constant. The amount of compoundreleased was monitored by spectrophotometry using an Agilent8453 UV–visible spectrophotometer, at 665 nm for MB, and276 nm for DCF and 517 nm for EY. All measurements were per-formed in duplicate and a CS/G1-P hydrogel (without MB, DCF orEY) was used as a blank control for absorbance correction.

2.10. Local tolerability of CS/G1-P hydrogel

2.10.1. Animal studyAnimal studies have been carried out in accordance with the EC

Directive 86/609/EEC. The local tolerability of CS/G1-P hydrogels

te as new stable in situ forming depot system for controlled drug delivery,

4 S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx

over three weeks was examined in healthy 10 week old HansWistar male rats. 0.5 mL sterile CS/G1-P thermogelling solution(1.5 wt.% Chitoceutical� 90/50(2), 0.40 mmol/g G1-P), equilibratedat room temperature, was subcutaneously injected using a 23 Gneedle in the interscapular region of nine rats under generalanesthesia. The same volume of sterile saline solution was subcu-taneously injected in their caudal dorsal area as a negative control.The health status and body weight of the animals were monitored.Blood samples were collected under isoflurane by sublingualpuncture at day �3 (predose) for hematology analysis and by venacava puncture on the day of necropsy for hematology andbiochemistry analysis. After 1, 7 and 21 days, three rats were sac-rificed via exsanguination from vena cava puncture after exposureto isoflurane and the tissues at the injection sites were excised forhistological examination.

2.10.2. Hematology and clinical chemistryBlood samples for hematology were collected into EDTA antico-

agulant (BD Microtainer� tube, K2EDTA, reference 365975), andstandard hematological parameters were examined using anADVIA 120� analyzer. Blood samples for clinical chemistry werecollected into plain tubes (SARSTEDT PC Micro Tube, reference41.1500.005), and the varying parameters disclosed were only tri-glycerides and glucose, and were analyzed using a Synchron CX5�

analyzer.

2.10.3. Histological examinationTissue samples used for histological examination were fixed in

10% v/v neutral buffered formaldehyde, embedded in paraffin andsectioned at 3 lm thickness. The sections were then stained withhematoxylin and eosin (H&E) or with 0.01% Fast Green in conjunc-tion with Safranin-O [25]. For histopathological assessment, thestained sections were examined using light microscopy. Eachanimal served as their own control. Microscopic findings wereentered in Pristima software and graded using a semiquantitativescale (minimal to severe); minimal (+) change contained one or afew small foci/cells; mild (++) change was composed of small tomedium size foci/cells; moderate (+++) change contained frequentand moderately sized foci/cells, marked (++++) change had exten-sive, confluent foci affecting most of the tissue and severe(+++++) change was effacing completely the normal structure ofthe tissue. Histological examinations were performed with 3 ani-mals by time point (i.e. in total 9 animals for the experiment),and thus since some responses were different between theanimals, the representation, for instance, ‘‘+++/++++’’ shows therange of the severity among animals from the same group, that isto say inflammation was moderate to marked.

2.11. Statistical analysis

Data were recorded as mean ± standard deviation. The signifi-cance of differences between groups was assessed using a t testwhen two groups were compared. ANOVA followed by Tukey’stests were applied for comparisons involving 3 or more groups.All significance tests were performed in Minitab 16 (Minitab, StateCollege, PA), and a p value less than 0.05 indicated a statisticallysignificant difference.

3. Results and discussion

3.1. Physical characterization of CS by triple detection GPC

A preliminary physical characterization of the chitosan batcheswas performed and reported in Supplementary information.Chitoscience� 90/50 and the two batches of Chitoceutical� 90/50

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

were supposed to have approximately the same Mw. Indeed, theviscosity of a polymeric solution depend on the Mw of the dis-solved polymer [26] and, according to the supplier, the viscosityof all CS 90/50 batches (1% solution in 1% acetic acid at 20 �C) isabout 50 mPa s. As expected, they showed Mw in the samerange of 122, 135 and 149 kDa respectively, whereas Chitoscience�

90/200, for which the viscosity is stated about 200 mPa s, revealeda higher Mw of 267 kDa. The Mw distribution, characterized by thepolydispersity index Mw/Mn, is rather broad, ranging from 1.57 to1.90 for all samples. Furthermore, the polydispersity index as wellas the intrinsic viscosity and the hydrodynamic radius increasedslightly with increasing polymer size.

3.2. Sol/gel transition behavior

CS/G1-P thermogelling solutions should undergo sol/gel transi-tion in situ due to body temperature; therefore their gelation timewas determined in PBS at 37 �C. The sol/gel transition should notbe too fast to avoid needle clogging during injection, but shouldbe rapid enough to avoid drug burst from the liquid formulation[27]. The influence of the CS Mw and of the CS and G1-P concentra-tions on the gelation time (Table 1) was studied. First, consistentlywith literature [28], decreasing the polymer Mw, from 267 to122 kDa, resulted in a significant decrease of the gelation timefrom 15 to 3.5 min. The same effect was observed with the CS con-centration, i.e. concentration variation from 2 to 1.5 wt.%, gave riseto a strong drop of the gelation time from 20 to 2.5 min. However,it is to be noted that a further decrease of the CS content to1.25 wt.% did not further accelerate the gelation. Finally, increasingthe G1-P concentration from 0.20 to 0.40 mmol/g resulted inreducing the gelation time from 7 to 2.5 min, in agreement withprevious reports on CS/b-GP systems [6,29]. Actually, the effect ofthe CS concentration and Mw on sol/gel transition time was likelylinked to viscosity effects: the higher the concentration or Mw, thehigher the solution viscosity, and the lower the mobility of the gel-ling agent and polymer chains, slowing down the gelation kinetics[30]. It is to be noted that in this reference, authors do not specif-ically study the effect on the molecular weight on the gelationkinetics, but as the viscosity increases with the CS molecularweight, it is evident that the same result is expected. Whenincreasing the G1-P concentration, consequently higher pHinduced a reduction of the degree of ionization of CS. Electrostaticinteraction opportunities between the polymer and the gellingagent were lowered, thereby reducing the protective hydrationlayer around CS. Furthermore, this leads also to a reduction ofthe electrostatic repulsion between CS chains, therefore promotinggelation at 37 �C.

3.3. Viscoelastic properties of CS/G1-P solutions

The viscoelastic behavior of CS solutions with different G1-Pconcentration is summarized in Fig. 1. The sol/gel transition tem-perature (Tgel) is defined as the temperature at which the storagemodulus (G0) is equal to the loss modulus (G00). As shown inFig. 1(a), at temperatures below Tgel, CS/G1-P systems exhibitedviscoelastic liquid behavior, with the curves showing the domi-nance of the loss modulus (G00 > G0). Upon temperature increase,at T > Tgel, the elastic portion of the system became dominant(G0 > G00), clearly indicating that the solutions turned into gels. Fol-lowing a similar trend, and corroborating the results of Table 1, thetime sweep tests at 37 �C (Fig. 1(b)) show a decrease of the gelationtime along with the increase of G1-P concentration. Furthermore,once the gelation process was initiated, the two moduli increasegradually up to the end of the test, indicating that the gel networkformation is progressive and was not completed during the timeof the measurements. The decrease of the gelation temperature

te as new stable in situ forming depot system for controlled drug delivery,

Table 1Gelation time at 37 �C and pH value of CS/G1-P solution as a function of CS Mw, CSconcentration and G1-P concentration.

Chitosan G1-P CS/G1-P

Mw (kDa) Concentration(wt.%)

Concentration(mmol/g)

tgel (min) pH

267 1.5 0.33 15 7.1122: 267 (1:1) 1.5 0.33 7 7.2122 1.5 0.33 3.5 7.4122 2 0.40 >20 7.3122 1.75 0.40 11 7.2122 1.5 0.40 2.5 7.4122 1.25 0.40 3 7.4122 1.5 0.27 4.5 7.3122 1.5 0.20 7 7.2

Fig. 1. (a) Temperature and (b) time dependence at 37 �C of storage modulus (G0 , d)and loss modulus (G00 , s) of CS/G1-P solutions (1.5 wt.% CS) for G1-P concentrationsfrom 0.27 to 0.53 mmol/g. The data are shifted along the vertical axis by a factor of10a to avoid overlapping (n = 3).

Fig. 2. (a) Frequency dependence of storage modulus (G0 , d) and loss modulus (G00 ,s) of CS/G1-P hydrogel (1.5 wt.% CS and 0.40 mmol/g G1-P) after 6 h at 37 �C. (b)Strength of CS/G1-P hydrogels (1.5 wt.% CS and 0.27 mmol/g G1-P) after 6 h, 24 hand 48 h incubation either in a saturated water vapor atmosphere or in PBS at 37 �C(n = 3).

S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx 5

of CS/G1-P systems goes along with a reduction in gelation time atphysiological temperature and therefore can be linked to theincrease in the pH value of the thermogelling solution when addingmore gelling agent. These results confirm the trend observed abovein vials. The differences in the gelation times observed by rheologyand those determined in vials can be attributed to the differentexperimental approaches and methodologies. Additional measure-ments on the sample viscosity are reported in the Supplementaryinformation section Fig. S1 and revealed that increasing the gelling

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

agent results in increasing the viscosity of the solutions, howeverremaining in values compatible with the aimed application.

Frequency sweep experiments provided the value of the hydro-gels strength (after gel formation, under saturated water vaporatmosphere at 37 �C). As shown in Fig. 2(a), the system displayG0 > G00 over the entire frequency range, with a value of G0 indepen-dent of the frequency. Furthermore, G0 and G00 curves exhibit almostparallel lines in the low frequency range, and the G0:G00 ratio isherein higher than 20:1, which typically indicates a gel-like char-acter [26]. In addition, the strengths of the gel formed after 6 h,24 h and 48 h in a saturated water vapor atmosphere at 37 �C,appeared to significantly increase over time (Fig. 2(b)), in align-ment with the observations made from the temperature and timesweep tests. This phenomenon is consistent with the moleculargelation mechanisms, induced by the temperature increase, andattributed to the gradual removing of the protective hydrationlayer formed by the gelling agent around the CS chains. Therefore,even if the first hydrophobic attractive interactions occur betweenthe polymer chains, inducing the formation of a network, the num-ber of CS-gelling agent interactions is still further reduced withtime, leading to a gradual increase of the gel strength. Furthermore,once the gelation process was initiated, the increased viscosity ofthe system slows down the diffusion of G1-P away from the poly-mer and the self-association of the CS chains. Therefore, the num-ber of junction points between CS molecules rises with time untilcomplete gelation is achieved.

Additionally, the strengths of CS/G1-P hydrogels formed eitherin a saturated water vapor atmosphere or in PBS at 37 �C werecompared (Fig. 2(b)). Gelation in PBS led to significantly strongergels, with G0 values in the range of 1000 Pa versus 200–400 Pa ina saturated water vapor atmosphere. Indeed, in the aqueous mediaat pH 7.4, G1-P can diffuse freely out of the gel during the gelationprocess while the pH is maintained higher than the pKa of CS

te as new stable in situ forming depot system for controlled drug delivery,

6 S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx

[2,31]. Thus, CS–CS attractive interactions are no longer preventedby the protective hydration layer and the steric hindrance causedby G1-P, thereby creating more physical junctions of the polymernetwork. CS/G1-P thermogelling solutions are intended to beadministered subcutaneously, where the gelling agent is likely todiffuse out of the depot and into the surrounding tissue, since fluidflow in s.c. tissue usually ranges from 7 to 53 mL/100 g per min[32,33]. Thus, strong hydrogels are expected to form progressivelyat the site of injection.

3.4. Differential scanning calorimetry

DSC has been used to characterize the thermoreversibility of CS/G1-P solutions. The first heating scan (Fig. 3, thermogram a) showsa broad endothermic peak (DH = 0.30 J g�1) between 33 and 48 �C,with a maximum temperature of 37.1 �C. This peak can be relatedto sol/gel transition of the CS/G1-P solution. This endothermic pro-cess could correspond to both the disruption of the hydrogenbonds forming the hydration shell surrounding the CS chains,and to the formation of intra- and inter-polymeric hydrophobicinteractions [34,35]. This is confirmed by the fact that the sol/geltransition temperature disclosed here is consistent with the gela-tion temperature given by rheological experiments at the sameheating rate (see Fig. 1(a)).

Likewise, a second DSC scan (i.e. this time, after gel formation)was performed on the CS/G1-P sample under identical conditions(Fig. 3, thermogram b). The transition appears thermo-irreversiblesince no endothermic peak appears for the second measurement.Complementary rheological experiments, presented in the Supple-mentary information section Fig. S2, corroborate these DSC resultsand disclose that once gel formed, it does not turns back to solutionstate even at low temperatures.

3.5. Storage stability of CS/gelling agent solutions

Storage stability is a crucial parameter that needs to be assessedduring pharmaceutical development of an in situ forming hydrogelsystem. This type of formulations should have an acceptable shelflife, i.e. their physical thermogelling properties should be main-tained throughout defined storage conditions up to administration.This section presents the stability study of two representative CS/G1-P thermogelling solutions (with 0.27 and 0.40 mmol/g G1-P)along with the comparison with the classical CS/b-GP thermogel-ling solution (with 0.40 mmol/g b-GP) and pure CS solution (atthe same polymer concentration: 1.5 wt.%). The study included

Fig. 3. DSC thermograms of a CS/G1-P solution (1.5 wt.% CS and 0.40 mmol/g G1-P)recorded at a heating rate of 1 �C/min. Thermogram a and b correspond to the firstand second heating scan respectively (n = 2).

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

following-up over 6–9 months, for two different storage tempera-tures (2–8 �C and 20–25 �C), the appearance, pH, viscosity andgelation time of the solution as well as CS Mw.

3.5.1. Appearance and pHOverall, no change in the appearance (liquid state) nor in the pH

values of the CS and CS/G1-P solutions was observed over the 6–9 months storage (for both temperatures). The pH of the pure CSsolution remained 5.2 ± 0.1 while those of CS/G1-P solutions were7.1 ± 0.1 and 7.2 ± 0.1, for 0.27 and 0.40 mmol/g G1-P, respectively.On the contrary, CS/b-GP solution turned into a turbid gel in lessthan 30 days at 2–8 �C and after only 1 day at 20–25 �C, with a dropof the pH value from 7.4 to 7.0.

3.5.2. ViscosityApparent shear viscosities of the CS solution and CS/gelling

agent systems are reported in Fig. 4. Results show that for storageperiods as long as 6–9 months and at refrigerated conditions, theapparent shear viscosity of the pure CS and CS/G1-P solutionsremained unchanged. On the contrary, the viscosity of CS/b-GPincreased significantly within 30 days, from about 180 mPa s upto 7500 mPa s at a shear rate of 5 s�1. Likewise, the viscosity ofCS/G1-P solutions was stable at room temperature over 2 monthsand only slightly increases at low shear rates after 6 months,whereas the viscosity of the pure CS solution very slightly reducedover time and the viscosity of the CS/b-GP solution raised drasti-cally after only 7 days. These results are fully aligned with formerobservations, i.e. the pure CS solution as well as the CS/G1-P solu-tions remained in the liquid state over time for both storage condi-tions while the CS/b-GP solution turned into gel with time. Storageunder refrigerated conditions allowed maintaining the CS/b-GPformulation in the liquid state for just over 14 days versus less than7 days at room temperature. Even, as previously observed by mon-itoring the gel strengths (Fig. 2(b)), once the sol/gel transitionoccurred, the viscosity of the CS/b-GP gels continue to increaseup to stabilization (more appreciable at 2–8 �C in Fig. 4(g)).

3.5.3. Gelation timeThe sol/gel transition times of the different CS/gelling agent

formulations setup at 37 �C, were followed over the same storageperiod, and reported in Fig. 5. It appears that the gelation time ofthe solution composed of 0.27 mmol/g G1-P is not stable over time,and decreases from about 40 min initially, to 35 min and 27 minafter 6 months storage at room temperature and 2–8 �C respec-tively, with significant variations observed over time under 20–25 �C conditions (e.g. at 60 days gelation times appearssignificantly different). These variations vanish when the G1-Pconcentration is increased to 0.40 mmol/g. Indeed, the gelationtime observed for the CS/G1-P solution containing 0.40 mmol/gG1-P stored at 20–25 �C was stable over 2 months, but slightlydecreased afterward. However, when stored under refrigeratedconditions, this solution displayed a constant gelation time of12.2 ± 1.0 min over 9 months. With regard to the CS/b-GP solu-tions, the samples underwent a spontaneous gelation in less than30 days at 2–8 �C and 1 day at 20–25 �C. These observations wereconfirmed by the rheological measurements of the gelation time,since the CS/b-GP system exhibited a gel-like behavior (G0 > G00)from the beginning of the time sweep test. However already beforethese time points, the gelation times were nevertheless very fast, inthe range of 1–2 min. So the CS/b-GP solution was not stable whenstored either at room temperature or under refrigerated condi-tions. Conversely, the addition of 0.40 mmol/g G1-P as alternativegelling agent to the CS solution significantly improved the stabilityof the system, in particular when stored at 2–8 �C.

The lack of stability of CS/b-GP and other CS/polyol thermogel-ling solutions, even stored at refrigerated conditions, was already

te as new stable in situ forming depot system for controlled drug delivery,

Fig. 4. Apparent shear viscosity at 20.0 ± 0.2 �C for CS solution and CS/gelling agent systems stored at 2–8 �C or at 20–25 �C (n = 3). CS concentration is 1.5 wt.% and the typeand concentration of gelling agent is indicated on the figure.

S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx 7

reported by Ruel-Gariepy et al. [22] and Schuetz et al. [19]. Theyconcluded that the most appropriate storage method consisted inlyophilization of the formulations, to prevent any evolution tothe gel state. However, this approach is expensive and wouldrequire a reconstitution step before injection. In addition, it alsoinduced substantial changes in the gelation times and viscositycompared to freshly prepared solution. In the present study,increasing the molecular size of the gelling agent polyol part, thatis to say using G1-P instead of the commonly used b-GP (the glu-cose-moiety is substituted for the glycerol-moiety), was shownto improve the stability of the CS-based thermosensitive solutions.

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

In view of our previous studies on the gelation mechanism [23],increasing the size of the gelling agent polyol part likely enhancesthe stability of the protective hydration layer surrounding the CSchains, thus allowing extending the shelf life of the CS/gellingagent solution. This hypothesis is confirmed with the follow-upof the chitosan degradation (through the molecular weight) overtime, presented in the Supplementary information section Fig. S3,and showing that pure CS shows a decrease of its Mw whereasCS forming hydrogels remains unaffected.

To sum-up, investigating the stability of CS/gelling agents at dif-ferent storage temperatures showed the superior protective effect

te as new stable in situ forming depot system for controlled drug delivery,

Fig. 5. Evolution of the gelation time, determined by rheology at 37 �C, of CS/gellingagent solutions stored at 2–8 �C and at 20–25 �C (n = 3).

8 S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx

of G1-P compared to the classical b-GP. It allowed identifying theoptimized formulation and storage conditions to provide anacceptable shelf life as well. Actually, CS/G1-P solution with agelling agent concentration of 0.40 mmol/g was stable for at least9 months when stored at refrigerated conditions. Thus, this ther-mogelling solution has the potential to be presented in the solutionstate, without requiring any reconstitution step before administra-tion, and ready-to-use pre-filled syringes can therefore beconsidered.

3.6. Determination of injectability

In the development of ISFD systems, an important issue also liesin the ease of injection through the thin needles commonly usedfor s.c. administration [36,37]. Therefore, the FDA Guidance forIndustry on container closure systems for packaging human drugsand biologics requires the evaluation of the performance of asyringe by establishing the force necessary to inject the dosageform. Needle size recommended for s.c. injection range between25 G (outer diameter (O.D.) = 0.51 mm, inner diameter(I.D.) = 0.26 mm) and 30 G (O.D. = 0.31 mm, I.D. = 0.16 mm), andthe volume of dosage form injected should be between 0.5 and2.0 mL [38–40]. In this section, we measured the force requiredto inject 1 mL thermogelling solution containing 1.5 wt.% CS and0.40 mmol/g G1-P (chosen as the optimized formulation) through23, 25 and 30 G needles. The pressure resistance of s.c. tissue atthe injection site was not taken into consideration. Two injectionspeeds were studied, 80 and 100 mm/min, which corresponds to40 and 33 s for completing the injection, respectively. It is also tonotice that the upper limit of 20 N was set as acceptance criteriafor s.c. injections, according to Schoenhammer et al. [33]. Theresults presented in Supplementary information show that, in theconditions of our study, only the needle size significantly impactedthe injection force. The force necessary to inject the CS/G1-P solu-tion was below 5 N when using 23 G and 25 G needles, which maybe related to very smooth injections. Using a thinner needle of 30 Grequired an increased injection force of about 17.5 N, which stilldid not exceed the upper limit of 20 N. In contrary to other ISFDsolutions which were generally administered through large nee-dles of 18–23 G due to their high viscosity [41–43], the CS-basedformulation can be considered as easy to inject through needlesas thin as 30 G. Administration via small needle size is known tocause less pain at the injection site, thereby improving patientcomfort and compliance [44].

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

3.7. In vitro release of model compounds

Due to the polycationic nature of CS, electrostatic interactions(attractive or repulsive) may occur between –NH3+ groups of CSand negatively or positively charged substances. MB (pKa of 3.8)and DCF (pKa of 4.15), respectively positively and negativelycharged at pH 7.4 were selected as model compounds to assessthe impact of the charge on the compound’s release profile.Fig. 6(a) compares the release profiles of these two compoundsin similar experimental conditions from formulations preparedwith and without gelling agent. In presence of gelling agent, therelease profile shows a high burst (>50%) in the first 6 h followedby a sustained release of the compound over about two days. Thehigh initial burst can be explained by the lag time between theinjection in the release medium and the complete gelation, whilethe following sustained release can be attributed to the diffusionof the compounds through the polymeric matrix. In contrast, with-out gelling agent, the complete release of the compound isachieved in less than 4 h. Despite the charge difference betweenMB and DCF, their release kinetics are rather similar, which doesnot allow at first sight to conclude regarding the influence of theelectrostatic interactions on the release profiles.

It is noteworthy that hydrogels like CS/G1-P systems areconsidered as homogeneous matrix systems in which the drug issolubilized in a continuous phase. As a consequence, the distanceof diffusion of the solute is not constant but increases with time,following thereby a nonsteady state diffusion regime generallydescribed by the second Fick’ law displayed in Eq. (1).

dCdt¼ D

d2C

dx2 ð1Þ

where C is the solute concentration, D the diffusion coefficient and xthe diffusion distance [45,46]. Eq. (2), which is derived from Eq. (1)[47], describes the case of homogeneous matrix systems like CS/G1-P systems and well fits the first part of the profile (0 6Mt/M1 6 0.6).

Mt

M1¼ 2S

VDtp

� �1=2

ð2Þ

where Mt is the mass of solute released at time t, M1 the total massof solute, S the surface of the gel and finally V the volume of thesample (including release medium). Accordingly, this model wasused to fit MB and DCF in vitro release data (see Fig. 6 inset). Corre-lation coefficients of R = 0.993 and R = 0.998 were determinedrespectively. It follows from the fits the relative comparison of thevalues of D, assuming that the geometrical conditions arestrictly the same for all the samples, and finally giving hereDMB = 1.44 � DDCF. The release mechanism of these two model com-pounds from the hydrogel matrix is effectively the diffusion (goodcorrelation with the model), but the diffusion of MB appears slightlyfaster than DCF, which might be due their different charges in thepresent experimental conditions (at pH 7.4). However, releasebehaviors of the monovalent anion DCF and the monovalent cat-ionic MB overall remain comparable. Thus, single ionic interactionsbetween DCF and CS did not seem to affect substantially the releasemechanism, in agreement with previous reports on CS-based ther-mogelling systems [22,48].

On the other hand, release profiles of divalent anionic EYreported in Fig. 6(b) show a more extended kinetics dependingon the concentration. Here also without gelling agent the soluteshows a very quick release (completed within 4–6 h) comparedto sustained release over about 12 days when the CS/G1-P hydro-gel is formed in presence of the gelling agent. Three different EYconcentrations were encapsulated, 0.06, 0.3, 0.6 wt.%, and weobserved a decrease of the kinetics as the concentration isincreased. Likewise, the release profiles were fitted with the

te as new stable in situ forming depot system for controlled drug delivery,

Fig. 6. (a) Release profiles of MB and DCF at 0.06 wt.%, from pure CS solutions andCS/G1-P hydrogels (mean ± SD, n = 3). (b) Release profiles of EY at 0.06 wt.% frompure CS solutions, and at 0.06, 0.3 and 0.6 wt.% from CS/G1-P hydrogels (mean ± SD,n = 3). Solid lines represent the curve fits from Eq. (2).

S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx 9

diffusion model adapted to homogeneous matrix systems as doneabove, giving correlation coefficients of R = 0.996, R = 0.987 andR = 0.973, for EY concentrations of 0.06, 0.3, 0.6 wt.%, respectively.Here also the correlation is relatively good but decreases slightlywith increasing the concentration, maybe due to higher EY/hydro-gels interactions impacting on the pure diffusion. The curve fitsgive, between the different concentrations of EY, the following rel-ative variations of diffusion coefficients: D0.06 = 1.66 � D0.3 andD0.06 = 5.61 � D0.6. These results show clear influence of theconcentration on the diffusion mechanism, meaning there areattractive intermolecular interactions slowing down the diffusionof EY. This idea is even corroborated when we compare the releaseof monovalent cationic MB or monovalent anionic DCF with thedivalent anionic EY at the same concentration (0.06 wt.%), givingthe diffusion of MB and DCF respectively around 89 times and 62times the one of EY (DMB = 89.28 � DEY and DDCF = 62.09 � DEY).

Fig. 7. (a) Temperature and (b) time dependence at 37 �C of storage modulus (G0 ,closed symbols) and loss modulus (G00 , opened symbols) of CS/G1-P/EY solutions forEY concentrations from 0 to 1.2 wt.%. The data are shifted along the vertical axis bya factor of 10a to avoid overlapping (n = 1).

3.8. Rheological behavior of CS/G1-P/EY

To examine the influence of EY loading on the thermogellingbehavior of the CS/G1-P solution, the changes in gelation

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

temperature and gelation time with the EY content was investi-gated. In Fig. 7(a), a clear sol/gel transition can be observed forthe reference CS/G1-P solution without EY, which exhibits a typicalsolution-like behavior with G0 < G00 below 39 �C and a gel behaviorwith G0 > G00 above this temperature. Addition of 0.03 wt.% EY led toa sol–gel behavior from 5 to 28.7 �C, with G0 � G00, and to the forma-tion of a weak gel upon further temperature increase. For higher EYloadings (0.3 and 1.2 wt.%), the formulations exhibited gel behavior(G0 > G00) over the whole temperature range. It is to be noted, even ifthe system has crossed the sol/gel transition when loaded with EYabove concentrations of 0.03% EY, the viscosity of the gels at roomtemperature still remains low enough to allow its easy injectionthrough 23 G needles comparable to unloaded systems. Actually,the strength of the hydrogel (evaluated in terms of the value ofG0) increases with the temperature, from 11.9 Pa at 20 �C to70.1 Pa at 42 �C for the system loaded with 0.3% EY, and from22.3 Pa at 20 �C to 92.3 Pa at 42 �C for the system loaded with1.2% EY. This can be explained by the fact that the CS/EY electro-static attraction compete with the CS/G1-P one, giving rise in fastersol/gel transition. Corroborating these results, the time sweep testsat 37 �C (Fig. 7(b)) show a decrease of the gelation time from 8.8 to0.5 min along with an increase of EY concentration from 0 to0.03 wt.%. A further increase of the EY loading to 0.3 and 1.2 wt.%led to weak gel behavior from the beginning of the test. Formationof ionic bridges between CS molecules via the use of negatively

te as new stable in situ forming depot system for controlled drug delivery,

(c) 1 day, x 1

(a) (b)

Fig. 8. (a and b) Local tolerance study in the subcutaneous tissue of rats at 1 day. (c) Histology of the CS/G1-P hydrogel and surrounding tissues at 1 day. The arrows indicatethe location of the CS/G1-P hydrogel. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

Fig. 9. Histopathological examination of inflammatory response (H&E staining) at the interface between the CS/G1-P hydrogel and the subcutaneous tissue at different timepoints, 1 day, 7 days and 21 days. Legend: V: vessels; E: eosinophil; M: macrophage; N: neutrophil; F: fibroblast. Dashed line on 21 days 20� image indicates the delimitationbetween the loose connective tissue and the fibrous capsule surrounding the hydrogel, composed of mature fibroblasts and dense collagen fibers. (For interpretation of thereferences to color in this figure legend, the reader is referred to the web version of this article.)

10 S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx

charged compounds is a well-known process for the preparation ofCS hydrogels for pharmaceutical and biomedical applications,especially by using tripolyphosphate or citrate as cross linker[49–51]. Due to the multivalent anionic nature of EY at the studied

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

pH, this compound can also behave as a cross linker, bindingreversibly the polycationic CS chains through ionic interactions.This ionic crosslinking leads to the formation of a weak network,as demonstrated by the viscoelastic behavior of the EY-loaded

te as new stable in situ forming depot system for controlled drug delivery,

Fig. 10. Histopathological examination of inflammatory response (0.01% Fast Greenin conjunction with Safranin-O staining) at the interface between the CS/G1-Phydrogel and the subcutaneous tissue after 7 days, and 21 days. Arrows showsphagocytosed hydrogel. Dashed line indicates the delimitation between the looseconnective tissue and the fibrous capsule surrounding the hydrogel, composed ofmature fibroblasts and dense collagen fibers. (For interpretation of the references tocolor in this figure legend, the reader is referred to the web version of this article.)

Table 2Histological examination of s.c. tissue of rats implanted with CS/G1-P hydrogels. Forinflammation, ‘‘++’’, ‘‘+++’’ and ‘‘++++’’ indicates mild, moderate and marked inflam-matory response respectively. For fibrosis, ‘‘�’’, ‘‘+’’ and ‘‘++’’ represents no, minimaland slight fibrosis respectively.

Time point post-injection Inflammation Fibrosis

1 day +++/++++ acute �7 days +++/++++ subacute +/++ immature21 days ++ chronic +/++ mature

S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx 11

hydrogels versus blank CS/G1-P solution. However, the CS/G1-P/EYsystems could still be easily injected through 23 G needles.

3.9. Histological examination

In order to investigate the local tolerance of implanted CS/G1-Phydrogels, the solution was injected into the s.c. tissue of nine rats.The injected solution turned into a gel in situ in few minutes underphysiological temperature (Fig. 8). The histopathological effectsresulting from the contact of the s.c. tissue with the hydrogel weremonitored at 1, 7 and 21 days post-injection. As expected, theresults show a typical foreign body reaction due to the implanta-tion of the hydrogel. On the first day post-injection, moderate tomarked acute inflammatory response could be observed (seeFig. 9(a and b)). Inflammatory cells, i.e. neutrophils, eosinophilsand fewer macrophages, were present in the tissue surroundingthe hydrogel. Furthermore, minimal to slight focal to multifocalhemorrhage accompanied the inflammation. Neutrophils and

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

macrophages containing intracytoplasmic granular material werefound at the interface with the hydrogel, showing that these cellsphagocytized the CS-based depot.

After one week post-injection, the depot was surrounded bynecrotic neutrophils and cellular debris admixed with fibrin.Moderate to marked subacute inflammation was observed, charac-terized by the presence of macrophages and fewer neutrophils,eosinophils and lymphocytes in the s.c. tissues, as shown in(Fig. 9(c and d)). Minimal to slight fibrosis consisting in animmature granulation tissue, composed of sprouting capillariesand activated fibroblasts, was located in the surrounding area ofthe hydrogel. Minimal focal multifocal hemorrhage accompaniedthe inflammation. Phagocytosis of the depot was indicated by thepresence of granular material in the cytoplasm of a few neutrophilsand macrophages, as displayed in Fig. 10(a).

Finally, after 21 days post-implantation, the degree of inflam-mation of the s.c. tissue declined significantly compared to theearlier time points (Fig. 9(e and f)). A slight chronic inflammationcomposed of lymphocytes and smaller amount macrophages waspresent around the hydrogel. Minimal to slight fibrosis, consistingin a mature less vascularized granulation tissue, encapsulated thedepot (Fig. 10(b)). As previously, minimal focal to multifocalhemorrhage accompanied the inflammation and intracytoplasmicphagocytized material was present in few neutrophils andmacrophages.

The histopathology results, summarized in Table 2, showed thatthe inflammatory reaction reduced with time, while slight fibrosisprogressively surrounded the implant site. These reactions arecomparable to the tissue responses typically observed for otherimplantable drug delivery systems, like biodegradable PLA andPLGA microspheres, or standard CS/b-GP hydrogels, which alsoshow acute inflammation followed by chronic inflammation, gran-ulation tissue formation and fibrosis [52,53]. This suggests therebyreasonably good compatibility of the hydrogel in the s.c. tissues.For all animals, the hydrogel persisted over the 3 weeks of study.Furthermore, no detrimental effect of the hydrogel on the bodyweight or animal behavior was observed.

4. Conclusion

To conclude, through a panel of different physico-chemicalcharacterization methods, including injectability evaluation,thermo-gelation and stability studies, the present work demon-strates the feasibility of CS/G1-P solution as interesting alternativeto the standard CS/glycerophosphate thermosensitive systemforming a depot at physiological conditions. Rheological thermo-analysis and DSC scan confirmed the sol/gel transition temperatureof the CS solution containing 0.40 mmol/g G1-P at around bodytemperature and demonstrated that the strong gel formed wasnot thermoreversible. Moreover, the storage stability problem ofthe formulation, actually one of the most important issue of stan-dard CS-based thermogelling systems, could be solved by usingG1-P as alternative gelling agent. The CS/G1-P in situ depot formingsolution kept its thermogelling properties for minimum 2 months

te as new stable in situ forming depot system for controlled drug delivery,

12 S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx

at room temperature and over 9 months at refrigerated conditions.Therefore, and in view of the easy injectability of the formulationsthough thin needles, a convenient presentation in ready-to-useprefilled syringes could be considered.

Evaluation of the release profiles obtained with different typesof low Mw hydrophilic compounds indicates that CS/G1-P thermo-sensitive hydrogels were able to retard their release for one toseveral days by physical entrapment. No decrease in release wasobserved by loading the hydrogel with the monoanionic DCFcompared to the monocationic MB. Thus, reversible ionic bindingsbetween MB and CS did not seem to influence the diffusioncontrolled release mechanism. However, as confirmed by the vis-coelastic behavior of CS/G1-P/EY systems, EY behaves as a crosslinker between CS molecules due to its dianionic nature. Conse-quently, the in vitro release of this low Mw compound displayedonly a mild initial burst and was extended over more than oneweek.

The inflammatory reaction observed around the CS/G1Pformulation injected subcutaneously in rats was a typical foreignbody reaction, similar to the tissue response reported for other par-enteral depot systems, like standard CS/glycerophosphate hydro-gels. Following the injection of the formulation, acute to chronicinflammation occurred in a sequential order and inflammationtended to fade with time. Fibrosis accompanied the inflammatoryresponse and surrounded the implant site; it started 7 dayspost-injection with a well vascularized immature granulationtissue and toward 21 days, it formed a slightly vascularized thickcapsule composed of fibroblasts and collagen fibers. Thus, thehydrogel exhibited acceptable tissue biocompatibility.

Acknowledgments

The authors would like to thank Baskim Ajdini for the technicalsupport on GPC and DSC measurement, and Karine Bigot, YvonneHager (Novartis Pharma, Basel, Switzerland), Thanh Truongand Yoav Timsit (Novartis Institutes for BioMedical Research,Cambridge, USA) for the in vivo tolerability study.

Appendix A. Supplementary material

Supplementary data associated with this article can be found, inthe online version, at http://dx.doi.org/10.1016/j.ejpb.2014.05.015.

References

[1] A. Chenite, C. Chaput, D. Wang, C. Combes, M.D. Buschmann, C.D. Hoemann, J.C.Leroux, B.L. Atkinson, F. Binette, A. Selmani, Novel injectable neutral solutionsof chitosan form biodegradable gels in situ, Biomaterials 21 (2000) 2155–2161.

[2] M. Iliescu, C.D. Hoemann, M.S. Shive, A. Chenite, M.D. Buschmann,Ultrastructure of hybrid chitosan–glycerol phosphate blood clots byenvironmental scanning electron microscopy, Microsc. Res. Tech. 71 (2008)236–247.

[3] J. Cho, M.C. Heuzey, A. Begin, P.J. Carreau, Chitosan and glycerophosphateconcentration dependence of solution behaviour and gel point using smallamplitude oscillatory rheometry, Food Hydrocolloids 20 (2006) 936–945.

[4] D. Filion, M. Lavertu, M.D. Buschmann, Ionization and solubility of chitosansolutions related to thermosensitive chitosan/glycerol-phosphate systems,Biomacromolecules 8 (2007) 3224–3234.

[5] M. Lavertu, D. Filion, M.D. Buschmann, Heat-induced transfer of protons fromchitosan to glycerol phosphate produces chitosan precipitation and gelation,Biomacromolecules 9 (2008) 640–650.

[6] F. Ganji, M. Abdekhodaie, S.A. Ramazani, Gelation time and degradation rate ofchitosan-based injectable hydrogel, J. Sol-Gel Sci. Technol. 42 (2007) 47–53.

[7] H.Y. Zhou, X.G. Chen, M. Kong, C.S. Liu, D.S. Cha, J.F. Kennedy, Effect ofmolecular weight and degree of chitosan deacetylation on the preparation andcharacteristics of chitosan thermosensitive hydrogel as a delivery system,Carbohydr. Polym. 73 (2008) 265–273.

[8] E. Ruel-Gariepy, M. Shive, A. Bichara, M. Berrada, D. Le Garrec, A. Chenite, J.C.Leroux, A thermosensitive chitosan-based hydrogel for the local delivery ofpaclitaxel, Eur. J. Pharm. Biopharm. 57 (2004) 53–63.

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

[9] M. Berrada, A. Serreqi, F. Dabbarh, A. Owusu, A. Gupta, S. Lehnert, A novel non-toxic camptothecin formulation for cancer chemotherapy, Biomaterials 26(2005) 2115–2120.

[10] E. Khodaverdi, M. Tafaghodi, F. Ganji, K. Abnoos, H. Naghizadeh, In vitro insulinrelease from thermosensitive chitosan hydrogel, AAPS PharmSciTech 13(2012) 460–466.

[11] Y. Peng, J. Li, J. Li, Y. Fei, J. Dong, W. Pan, Optimization of thermosensitivechitosan hydrogels for the sustained delivery of venlafaxine hydrochloride, Int.J. Pharm. 441 (2013) 482–490.

[12] M.S. Shive, C.D. Hoemann, A. Restrepo, M.B. Hurtig, N. Duval, P. Ranger, W.Stanish, M.D. Buschmann, BST-CarGel: in situ chondroinduction for cartilagerepair, Oper. Tech. Orthop. 16 (2006) 271–278.

[13] C.D. Hoemann, J. Sun, A. Legare, M.D. McKee, M.D. Buschmann, Tissueengineering of cartilage using an injectable and adhesive chitosan-basedcell-delivery vehicle, Osteoarthr. Cartilage 13 (2005) 318–329.

[14] S.M. Richardson, N. Hughes, J.A. Hunt, A.J. Freemont, J.A. Hoyland, Humanmesenchymal stem cell differentiation to NP-like cells in chitosan–glycerophosphate hydrogels, Biomaterials 29 (2008) 85–93.

[15] C. Marchand, G.E. Rivard, J. Sun, C.D. Hoemann, Solidification mechanisms ofchitosan–glycerol phosphate/blood implant for articular cartilage repair,Osteoarthr. Cartilage 17 (2009) 953–960.

[16] Q.F. Dang, J.Q. Yan, H. Lin, X.G. Chen, C.S. Liu, Q.X. Ji, J.J. Li, Design andevaluation of a highly porous thermosensitive hydrogel with low gelationtemperature as a 3D culture system for Penaeus chinensis lymphoid cells,Carbohydr. Polym. 88 (2012) 361–368.

[17] L.S. Nair, T. Starnes, J.W.K. Ko, C.T. Laurencin, Development of injectablethermogelling chitosan-inorganic phosphate solutions for biomedicalapplications, Biomacromolecules 8 (2007) 3779–3785.

[18] H.T. Ta, H. Han, I. Larson, C.R. Dass, D.E. Dunstan, Chitosan-dibasicorthophosphate hydrogel: a potential drug delivery system, Int. J. Pharm.371 (2009) 134–141.

[19] Y.B. Schuetz, R. Gurny, O. Jordan, A novel thermoresponsive hydrogel based onchitosan, Eur. J. Pharm. Biopharm. 68 (2008) 19–25.

[20] L. Casettari, M. Cespi, G.F. Palmieri, G. Bonacucina, Characterization of theinteraction between chitosan and inorganic sodium phosphates by means ofrheological and optical microscopy studies, Carbohydr. Polym. 91 (2013) 597–602.

[21] S. Makwana, B. Basu, Y. Makasana, A. Dharamsi, Prefilled syringes: aninnovation in parenteral packaging, Int. J. Pharm. Invest. 1 (2011) 200–206.

[22] E. Ruel-Gariepy, A. Chenite, C. Chaput, S. Guirguis, J.-C. Leroux,Characterization of thermosensitive chitosan gels for the sustained deliveryof drugs, Int. J. Pharm. 203 (2000) 89–98.

[23] S. Supper, N. Anton, N. Seidel, M. Riemenschnitter, C. Schoch, T.F. Vandamme,Rheological study of chitosan/polyol-phosphate systems: influence of thepolyol part on the thermo-induced gelation mechanism, Langmuir 29 (2013)10229–10237.

[24] X. Chen, X. Li, Y. Zhou, X. Wang, Y. Zhang, Y. Fan, Y. Huang, Y. Liu, Chitosan-based thermosensitive hydrogel as a promising ocular drug delivery system:preparation, characterization, and in vivo evaluation, J. Biomater. Appl. 27(2012) 391–402.

[25] E. Rossomacha, C.D. Hoemanni, M.S. Shive, Simple methods for stainingchitosan in biotechnological applications, J. Histotechnol. 27 (2004) 31–36.

[26] T.G. Mezger, The Rheology Handbook, 3rd revised ed., Vincentz, Hannover,Germany, 2006. pp. 171–174.

[27] S. Kempe, K. Mäder, In situ forming implants – an attractive formulationprinciple for parenteral depot formulations, J. Controlled Release 161 (2012)668–679.

[28] M.L. Tsai, H.W. Chang, H.C. Yu, Y.S. Lin, Y.D. Tsai, Effect of chitosancharacteristics and solution conditions on gelation temperatures of chitosan/2-glycerophosphate/nanosilver hydrogels, Carbohydr. Polym. 84 (2011) 1337–1343.

[29] S. Kempe, H. Metz, M. Bastrop, A. Hvilsom, R.V. Contri, K. Mader,Characterization of thermosensitive chitosan-based hydrogels by rheologyand electron paramagnetic resonance spectroscopy, Eur. J. Pharm. Biopharm.68 (2008) 26–33.

[30] J.Y. Cho, M.C. Heuzey, A. Begin, P.J. Carreau, Physical gelation of chitosan in thepresence of beta-glycerophosphate: the effect of temperature,Biomacromolecules 6 (2005) 3267–3275.

[31] D. Filion, M.D. Buschmann, Chitosan – glycerol-phosphate (GP) gels releasefreely diffusible GP and possess titratable fixed charge, Carbohydr. Polym. 98(2013) 813–819.

[32] M.L. Shively, B.A. Coonts, W.D. Renner, J.L. Southard, A.T. Bennett, Physico-chemical characterization of a polymeric injectable implant delivery system, J.Controlled Release 33 (1995) 237–243.

[33] K. Schoenhammer, H. Petersen, F. Guethlein, A. Goepferich,Poly(ethyleneglycol) 500 dimethylether as novel solvent for injectablein situ forming depots, Pharm. Res. 26 (2009) 2568–2577.

[34] P.L. Privalov, N.N. Khechinashvili, A thermodynamic approach to the problemof stabilization of globular protein structure: a calorimetric study, J. Mol. Biol.86 (1974) 665–684.

[35] G. Bruylants, J. Wouters, C. Michaux, Differential scanning calorimetry in lifescience: thermodynamics, stability, molecular recognition and application indrug design, Curr. Med. Chem. 12 (2005) 2011–2020.

[36] F. Cilurzo, F. Selmin, P. Minghetti, M. Adami, E. Bertoni, S. Lauria, L. Montanari,Injectability evaluation: an open issue, AAPS PharmSciTech 12 (2011) 604–609.

te as new stable in situ forming depot system for controlled drug delivery,

S. Supper et al. / European Journal of Pharmaceutics and Biopharmaceutics xxx (2014) xxx–xxx 13

[37] A. Jaber, G. Bozzato, L. Vedrine, W. Prais, J. Berube, P. Laurent, A novel needlefor subcutaneous injection of interferon beta-1a: effect on pain in volunteersand satisfaction in patients with multiple sclerosis, BMC Neurol. 8 (2008) 38.

[38] L.A Gatlin, C.A. Brister Gatlin, Formulation and administration techniques tominimize injection pain and tissue damage associated with parenteralproducts, in: Injectable Drug Development, CRC Press, 1999, pp. 401–421.

[39] K. Strauss, I. Hannet, J. McGonigle, J.L. Parkes, B. Ginsberg, R. Jamal, A. Frid,Ultra-short (5 mm) insulin needles: trial results and clinicalrecommendations, Pract. Diab. Int. 16 (1999) 218–222.

[40] D.M. Robb, Z. Kanji, Comparison of two needle sizes for subcutaneousadministration of enoxaparin: effects on size of hematomas and pain oninjection, Pharmacother.: J. Hum. Pharmacol. Drug Ther. 22 (2002) 1105–1109.

[41] L.R. Asmus, B. Kaufmann, L. Melander, T. Weiss, G. Schwach, R. Gurny, M.Moller, Single processing step toward injectable sustained-releaseformulations of Triptorelin based on a novel degradable semi-solid polymer,Eur. J. Pharm. Biopharm. 81 (2012) 591–599.

[42] F. Kang, J. Singh, In vitro release of insulin and biocompatibility of in situforming gel systems, Int. J. Pharm. 304 (2005) 83–90.

[43] S. Kempe, H. Metz, K. Mader, Do in situ forming PLG/NMP implants behavesimilar in vitro and in vivo? A non-invasive and quantitative EPR investigationon the mechanisms of the implant formation process, J. Controlled Release 130(2008) 220–225.

[44] L. Arendt-Nielsen, H. Egekvist, P. Bjerring, Pain following controlled cutaneousinsertion of needles with different diameters, Somatosens. Mot. Res. 23 (2006)37–43.

Please cite this article in press as: S. Supper et al., Chitosan/glucose 1-phosphaEur. J. Pharm. Biopharm. (2014), http://dx.doi.org/10.1016/j.ejpb.2014.05.015

[45] W.I. Higuchi, Diffusional models useful in biopharmaceutics. Drug release rateprocesses, J. Pharm. Sci. 56 (1967) 315–324.

[46] N.A. Peppas, Mathematical modeling of diffusion processes in drug deliverypolymeric systems, in: Controlled Drug Bioavailability, vol. 1, New York, 1983.

[47] P. Buri, F. Puisieux, E. Doelker, J.P. Benoit, Formes pharmaceutiques nouvelles:aspects technologique, biopharmaceutique et médical, Eds. Tec et Doc,Lavoisier, 1985, pp. 105–133.

[48] H.Y. Zhou, X.G. Chen, M. Kong, C.S. Liu, Preparation of chitosan-basedthermosensitive hydrogels for drug delivery, J. Appl. Polym. Sci. 112 (2009)1509–1515.

[49] J. Berger, M. Reist, J.M. Mayer, O. Felt, N.A. Peppas, R. Gurny, Structure andinteractions in covalently and ionically crosslinked chitosan hydrogels forbiomedical applications, Eur. J. Pharm. Biopharm. 57 (2004) 19–34.

[50] T.T. Ching, H.C. Chih, D.L. Yu, F.W. Ming, P.L. Chun, L.H. Jin, H.C. Szu, H.H. Kuo,Development of chitosan/dicarboxylic acid hydrogels as wound dressingmaterials, J. Bioactive Compat. Polym. 26 (2011) 519–536.

[51] Z. Lu, J. Steenekamp, J. Hamman, Cross-linked cationic polymer microparticles:effect of N-trimethyl chitosan chloride on the release and permeation ofibuprofen, Drug Dev. Ind. Pharm. 31 (2005) 311–317.

[52] J.M. Anderson, M.S. Shive, Biodegradation and biocompatibility of PLA andPLGA microspheres, Adv. Drug Deliv. Rev. 28 (1997) 5–24.

[53] J. Sun, G. Jiang, T. Qiu, Y. Wang, K. Zhang, F. Ding, Injectable chitosan-basedhydrogel for implantable drug delivery: body response and induced variationsof structure and composition, J. Biomed. Mater. Res. 95A (2010) 1019–1027.

te as new stable in situ forming depot system for controlled drug delivery,