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hydrogel composites with pH and redox sensitivity for combined release of anticancer drugs
Chemical Engineering Journal 287 (2016) 20–29
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Injectable shell-crosslinked F127 micelle/hydrogel composites with pH and redox sensitivity for combined release of anticancer drugs Nannan Gao a, Shaoyu Lü a,⇑, Chunmei Gao a, Xinggang Wang a,b, Xiubin Xu a, Xiao Bai a, Chen Feng a, Mingzhu Liu a,⇑ a State Key Laboratory of Applied Organic Chemistry, Key Laboratory of Nonferrous Metal Chemistry and Resources Utilization of Gansu Province, Department of Chemistry, Lanzhou University, Lanzhou 730000, People’s Republic of China b Research Institute of Lanzhou Petrochemical Corporation, Petrochina Lanzhou Petrochemical Company, Lanzhou 730060, People’s Republic of China
h i g h l i g h t s A novel dual drug carrier is constructed by injectable micelle/hydrogel composite. This drug carrier has redox-sensitivity and pH-sensitivity. The preparation is inexpensive, facile and fast. The system can load and deliver hydrophobic and hydrophilic drugs simultaneously. The system can be used in combination chemotherapy for cancer treatment.
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Article history: Received 28 July 2015 Received in revised form 15 October 2015 Accepted 6 November 2015 Available online 14 November 2015 Keywords: Micelle/hydrogel composite Shell-crosslinked micelle Injectable Stimuli-response Dual drug carrier
a b s t r a c t Combination chemotherapy as a preferred strategy for cancer treatment has been widely developed recently. However, single drug carriers of the combination chemotherapy systems are difficult to address the performance requirement that delivers hydrophilic and hydrophobic drugs simultaneously. In this work, a smart dual drug delivery system of injectable micelle/hydrogel composite was developed. This smart drug delivery system consists of curcumin (Cur) loaded shell-crosslinked F127 micelle and 5fluorouracil (Fu) dispersed chitosan/oxidized dextran (CS/ODex) hydrogel. Both of the shell-crosslinked F127 micelle and CS/ODex hydrogel were prepared based on Schiff base bonds, which provides pHsensitivity and biodegradability for the system. In addition, disulfide bonds were introduced in the shell of F127 micelle, allowing for redox-sensitivity of the system. The system was detailed characterized by 1H NMR, FTIR, TEM, SEM, DLS and fluorescence spectroscopy. Independent release behaviors of the two drugs are observed under conditions of different precursor concentration, different CS/ODex proportion, different pH and different concentration of DTT. The results show that the drug carrier has pH and redox sensitivity and the release rate can be adjusted by the concentration of the precursor solution. In vitro degradation experiment and cytotoxicity assay determine that the system has good biodegradability and biocompatibility. These results indicated that the injectable shell-crosslinked F127 micelle/hydrogel composite may load and deliver hydrophobic and hydrophilic drugs simultaneously and find application in combination chemotherapy for cancer treatment. Ó 2015 Elsevier B.V. All rights reserved.
1. Introduction Chemotherapy has been widely used to treat with established cancer in the clinic along with surgical excision and irradiation [1]. However, traditional chemotherapy is always accompanied by many side effects: most drugs are of high toxicity but lack of ⇑ Corresponding authors. Tel.: +86 931 8912387; fax: +86 931 8912582. E-mail addresses: [email protected] (S. Lü), [email protected] (M. Liu). http://dx.doi.org/10.1016/j.cej.2015.11.015 1385-8947/Ó 2015 Elsevier B.V. All rights reserved.
site-specific, drug level is uncontrolled and therefore need repeated administration and so on. The facts above bring tremendous physical and mental harm to patients [2]. Due to these obstacles, combination chemotherapy as a preferred strategy for cancer treatment has replaced traditional chemotherapy, which is characterized by the synergistic effects using multiple drugs of different therapeutic effects, and thus improves the curative effect [3]. Drug delivery system (DDS), as an useful adjunct to combination chemotherapy, has been widely researched over the last few
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decades [4–8], as DDS could overcome the problems of conventional administration by enhancing drug solubility, prolonging duration time, minimizing side effects, protecting drug bioactivity and so on [9–11]. Accordingly, a variety of drug delivery systems (DDSs), such as lipids [12], microspheres [13], nanoparticles [14,15], micelles [16] and gels [17] have been developed. However, single drug carriers of the combination chemotherapy systems are difficult to address the performance requirement that delivers different drugs simultaneously. Therefore, developing a dual drug delivery system (DDDS) which can simultaneously load multiple drugs and controlled release of each drug is desired. Zhu et al. designed composite hydrogels based on chitosan and mesoporous silica which could simultaneously release gentamicin and bovine serum albumin in a sustained manner [18]. Gong et al. developed poly-(3-caprolactone)–poly(ethylene glycol)–poly(3-ca prolactone)-based micelle/hydrogel carrier to deliver paclitaxel and fluorouracil simultaneously for the treatment of colorectal peritoneal carcinomatosis [4]. Despite a few examples of DDDS for hydrophobic and hydrophilic drug delivery application, the possible toxicity of the agents these systems used, undesirable side products from gel degradation, and the complex synthesis process compromise their real use. Here, we developed an injectable micelle/hydrogel dual drug delivery system based on natural polysaccharide through an inexpensive, facile and fast method. Injectable hydrogels are appealing for entrapping hydrophilic drugs with the web-like structure and release drugs directly to the target site (local drug delivery) with longer drug retention and larger drug accumulation in target cells than intravenous administration [19–21]. Micelles are one of the ideal vehicles for poorly water-soluble anticancer drugs with their intrinsic core–shell structure and could be passively targeted to the tumor site due to enhanced permeability and retention (EPR) effect [22–24]. 5-Fluorouracil (Fu), as a hydrophilic anticancer drug, was loaded in hydrogel. Curcumin (Cur), as a hydrophobic anticancer drug, was loaded in micelles. Multiple drugs of different therapeutic effects were used to improve the curative effect. The injectable micelle/hydrogel delivery system consists of chitosan (CS), dextran (Dex), and Pluronic F127 (PEO–PPO–PEO) micelles. Chitosan, a positively charged natural polysaccharide, composed of b-(1,4)-linked D-glucosamine
and N-acetyl-D-glucosamine units [25], has been widely used as hydrogel material for a long time due to excellent biocompatibility, biodegradability and antimicrobial activity [26,27]. Dextran (Dex) is another natural biocompatible material widely used in biomedical applications. The micelle is prepared using F127, which is commercially available and contains hydrophilic block poly(ethylene oxide) (PEO) and hydrophobic block poly(propylene oxide) (PPO) [28]. The gelation of the composite system is attributed to the Schiff base reaction between aldehydes or amine groups on the three components. The Schiff base reactions have drawn great attention in preparation of DDSs with their attractive properties: aldehydes and amino groups on polysaccharides can be easily acquired through facile methods [29,30]; the reactions take place at physiological conditions without toxic chemical cross-linking agents and water is the only byproduct [31]; hydrogels crosslinked by Schiff base bonds could gradually degrade because of the hydrolysis of Schiff base bonds in vitro, and acid environment could accelerate the degradation. Therefore, the Schiff base based micelle/injectable hydrogel would be a biodegradable and pH-sensitive delivery system, which is preferable for anticancer drugs carriers because of most cancer tissues hold a much lower pH compared to the normal tissues [32]. As we know, above the critical micelle concentration (CMC), F127 can self-assemble into micelle. PPO serves as a hydrophobic core providing a microenvironment for loading of hydrophobic
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drugs, PEO serves as a hydrophilic outer shell making micelle stable and unrecognizable by the reticuloendothelial system [33]. Researchers have demonstrated that Pluronic F127 is a promising multidrug resistance (MDR) reversal material because it can inhibit P-glycoprotein (P-gp) efflux, inhibit the glutathione (GSH)/glutathione (GST) detoxification system, enhance proapoptotic signaling, decrease anti-apoptotic defense in MDR cells and so on [34]. However, there are some defects in the use of Pluronic F127 micelle in the blood circulation. F127 is not able to keep micelle structure at low concentration, then drugs entrapped in the core will leak out. In addition, F127 tends to be prematurely excreted by the kidneys or penetrate into healthy endothelia, due to its relatively small particle size. Moreover, F127 does not have environmental sensitivity to smart release of drug [35]. Based on the above considerations, we prepared shellcrosslinked F127, using cystamine dihydrochloride as a crosslinking agent. After shell crosslinking, the structure of F127 micelle would be fixed and consequently more stable than before. In addition, the disulfide bonds on the shell of F127 micelle allow for redox-sensitivity of micelle, which could be cleaved by GSH in vivo. As known, GSH is abundant (mM) in the cytosol and subcellular compartments and much more in cancer cells, but very low (lM) in plasma [36,37]. Such micro-environmental differences allow that F127 micelles keep stable under physiological conditions and trigger sudden burst of drugs through the shell breaking in tumor cells. Therefore, the smart injectable micelle/hydrogel dual drug delivery system could simultaneously load hydrophobic and hydrophilic drugs, and local deliver drugs through injectable hydrogel and intracellular release of drugs through shellcrosslinked F127 micelle. The pH and redox potential stimuli responsive drug delivery behaviors relying on the significant pH and redox potential differences between the tumor and normal tissues were determined. 2. Materials and methods 2.1. Materials Chitosan (CS, degree of deacetylation was 88.0%) was purchased from Sinopharm Chemical Reagent Co., Ltd. and used without further purification. Dextran (Dex, Mw 70 k) and Dess–Martin periodinane was provided from Aladdin biological technology Co., Ltd. (Shanghai, China). Pluronic F127 was obtained from Shanghai Sangon Biological Engineering Technology & Services Co., Ltd. (Shanghai, China). Dithiothreitol (DTT), cystamine dihydrochloride, sodium periodate and other chemicals were used as received. 2.2. Preparation of shell-crosslinked F127 micelle The shell-crosslinked F127 micelle was carried out according to a previously reported procedure in the literature [38]. The Dess– Martin periodinane was employed in the process of chemical conversion of the terminal alcohols on F127 into aldehydes (F121CHO). The Dess–Martin periodinane (0.5 g) was added to a solution of F127 (5.16 g) in methylene chloride (410 mL) and the resulting solution was stirred for 24 h at 40 °C. After removal of most of the solvent, the product was isolated by precipitation into hexane. The extent of oxidation, which defined as the number of oxidized F127-CHO residues per 100 original F127 feeds, was determined by the proton nuclear magnetic resonance (1H-NMR) analysis. F127-CHO (1.8 g) was dissolved in tetrahydrofuran (THF, 100 mL). This solution was added dropwise to deionized water (350 mL) and stirred for 30 min. THF was then removed using a rotary evaporator at 40 °C. Subsequently, cystamine dihydrochloride (0.21 mmol) was added dropwise to the above micelle
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solution. The mixtures were stirred for 1 h at 30 °C to crosslink the aldehydes termini on F127-CHO. 10 mL of the resulting reaction mixture was used to determine the degree of crosslinking according to the previous literature [31]. The unreacted cystamine dihydrochloride was determined, and the crosslinking degree of micelle was calculated. The products were dialyzed for 3 days and lyophilized. 2.3. Preparation of CS solution First, CS (3 g) was dispersed in deionized water (80 mL) while stirring. Then, acetic acid was added dropwise to the solution until a clear solution was obtained. Afterwards, the mixture was cooled to 4 °C and the pH was adjusted to 6.2 using ice aqueous sodium bicarbonate. Finally, a corresponding amount of deionized water was added to the reaction system and the final concentration of CS was 2 wt%. The CS solution was stored at 4 °C before use. 2.4. Preparation of oxidized dextran (ODex) Sodium periodate (13.2 g in aqueous solution) was added to a Dex solution (1% (w/v), 200 mL) and stirred at room temperature for 2 h in the dark. Diethylene glycol (4 mL) was added to quench any unreacted sodium periodate. After that the solution was stirred for 1 h and dialyzed for 3 days and lyophilized. The oxidation degree of ODex, which defined as the percentage of Dex structure units that have been oxidized, was determined by hydroxylamine hydrochloride titration. First, hydroxylamine hydrochloride (3 g), ethanol (8 mL), deionized water (5 mL), sodium hydroxide (0.84 g) in ethanol (20 mL) and indicator bromophenol blue solution (0.45% (w/v), 5 mL) were mixed to obtain hydroxylamine hydrochloride solution. Then, ODex (0.100 g), hydroxylamine hydrochloride solution (5 mL), ethanol (5 mL) and deionized water (5 mL) were mixed in conical flask and stirred for 15 min. Finally, the solution was titrated by standard hydrochloric acid solution until a yellow color was obtained. Dex was employed as a control group and the values were average of three experiments. The formula of oxidation degree of ODex are as follows:
Degree of oxidation ð%Þ ¼
162 ðV 2 V 1 Þ C 103 100% 2 MW
where the C (mol L1) is the concentration of the standard hydrochloric acid solution, MW is average molecular weight of Dex, the V1 and V2 (mL) are the volume of the standard hydrochloric acid solution used by the test group and the control group, respectively.
2.6. Characterization The morphological of the micelle was examined by transmission electron microscope (TEM, JEM-1200EX/S, Hitachi, Japan) operating at an accelerating voltage of 200 kV. The samples were prepared by the micelle solution on a formvar coated copper grid and then incubated in oven for 15 min at 37 °C. The size of the micelle was analyzed via dynamic light scattering (DLS) using a laser light scattering spectrometer (ALV/SP-125). Scattered light was collected at a fixed angle of 90° for duration of 10 min. All data were averaged over three measurements. The zeta-potential (n) was measured by Malvern ZEN3600, England. The critical micelle concentration (CMC) of the shell-crosslinked micelle was evaluated by fluorescence spectroscopy (LS55, Perkin– Elmer, America), using pyrene as a probe. The excitation spectra of micelle solutions were scanned from 360 to 400 nm with an emission wavelength of 390 nm. The CMC was estimated as the crosspoint when extrapolating the intensity ratio of I384/I373. FTIR absorption spectra of CS, Dex, ODex and CS/ODex (Gel 3) were taken on a Fourier transform infrared (FTIR) spectrometer (Nicolet 670 FTIR, USA) over the region from 4000 to 400 cm1. The gelation time of the hydrogel was visually measured by the vial inversion method when the sample solutions were unable to be pipetted up and down. The morphology of the hydrogel after lyophilization was observed by a JSM-5600LV scanning electron microscopy (SEM), Japan. Before observation, all samples were fixed on aluminum stubs and coated with gold.
2.7. Cytotoxicity test The cytotoxicity of the injectable micelle/hydrogel dual drug delivery system was assessed by 3-(4,5-dimethylthiazol-2-yl)-2,5 -dipheny-ltetrazolium bromide (MTT) assay with HeLa cells. The hydrogel precursor solutions of Gel 1, Gel 2 and Gel 3 (CS solution, ODex solution and micelle solution) which had been sterilized just before use were mixed and injected into a 96-well plate to form hydrogels in the bottom of the plate, respectively. The total volume of the precursor solutions was 20 lL. Then 160 lL of culture medium containing HeLa cells was added into each well at a density of 10,000 cells/well. After culturing for 48 h in a humidified atmosphere with 5% CO2 at 37 °C, the medium was removed carefully and replaced by MTT reagent for another 4 h. Then the cells were dissolved by 200 lL of DMSO. Cell viability was determined by the absorbance values at 490 nm which was measured with a microplate reader (Tecan, Mannedorf, Switzerland).
2.5. Preparation of injectable micelle(Cur)/hydrogel(Fu) dual drug delivery system To prepare Cur-loaded micelle, Cur and shell-crosslinked F127 micelle (0.1 g) were dissolved in 5 mL of distilled water with sonication for 30 min and subsequently stirred for 48 h. The free Cur was then removed by centrifugation. The loading amount of Cur in micelle was determined by testing the weight and the resulting drug-loaded micelle solution was stewing before use. The micelle(Cur)/hydrogel(Fu) delivery system was prepared as follows. First, Fu was dissolved in ODex solution. Then, Fu-loaded ODex solution, CS solution and Cur-loaded micelle solution were mixed at 4 °C (the concentration ratio of CS and ODex was 1/1). Finally, the mixture was heated to 37 °C for gelation. The final concentration of Fu and micelle was 2.5 mg mL1 and 3.3 mg mL1, respectively. The concentration of the hydrogel was 1.0 wt% (Gel 1), 1.5 wt% (Gel 2) and 2.0 wt% (Gel 3), respectively.
2.8. Drug release study of the injectable micelle/hydrogel dual drug delivery system in vitro Gel 1, Gel 2 and Gel 3 were directly immersed in 8 mL buffer (ethanol/phosphate buffer solution (80:20, v/v)), and then incubated at 37 °C under oscillation at 80 rpm. At a specific time, the release medium (3 mL) was withdrawn and replaced with fresh buffer. UV/vis absorbance (Lambda 35, Perkin–Elmer, America) was recorded at 268 nm for Fu and 430 nm for Cur. The release amount of drugs was determined via the standard curves of each drug in the same buffer. Gel 3 with various CS/ODex mass ratio (1/1, 1/2 and 2/1), various pH values (pH 7.4, pH 5.0 and pH 3.0) and various concentration of DTT (0 M, 5 mM, 10 mM) were also tested.
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2.9. Degradation behavior of the injectable micelle/hydrogel system in vitro The in vitro degradation of the micelle/hydrogel dual drug delivery system was examined as follows. 3 mL of the different hydrogels were immersed in tubes with 10 mL PBS (pH 7.4) at 37 °C. At predetermined time intervals, the PBS was taken out and excess surface water was gently wiped with filter paper. After that, the hydrogel was weighed, and fresh buffer was replenished. The weight loss ratio of the system was calculated.
2.10. Statistical analysis Statistical analysis of data was performed by one-way analysis of variance (ANOVA), assuming confidence level of 95% (p < 0.05) for statistical significance.
3. Results and discussion 3.1. Preparation of injectable micelle/hydrogel composite In order to improve the therapeutic outcome of combined therapy in treatment of cancer, we developed a smart dual drug delivery system of injectable micelle/hydrogel composite. As shown in Scheme 1, this smart drug delivery system consists of shellcrosslinked F127 micelle and CS/ODex hydrogel. To obtain shellcrosslinked F127 micelle, parts of terminal alcohols on F127 were oxidized to aldehydes using the Dess–Martin periodinane. Then the F127-CHO micelle was shell-crosslinked by linking up the aldehydes with cystamine dihydrochloride. The final micelle was redox-responsive with the disulfide bonds. ODex was prepared by oxidizing Dex with sodium periodate, which could cleave the vicinal glycols and convert them to aldehydes. There were many amino groups on CS and some free amino groups on the shellcrosslinked F127 micelle that could crosslink with the aldehydes
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on ODex and form hydrogel. The micelle/hydrogel composite was generated by mixing micelle solution, CS solution and ODex solution via Schiff connection. The system would be pH-responsive due to the carbon–nitrogen double bonds.
3.2. Characterization of the shell-crosslinked F127 micelle Fig. 1 shows the 1H NMR spectra of F127-CHO and shellcrosslinked F127. The peak at 9.7 ppm in the F127-CHO spectrum was attributed to the hydrogen in aldehydes (ACHO), further demonstrating the primary alcohols of F127 were converted to aldehydes. From the peak areas of 9.7 ppm and 1.18 ppm (OCH2OH(CH3)O of F127-CHO) in the F127-CHO spectrum, we can estimate that about 85% of the proportion of alcohols in F127 was converted to aldehydes. The peak at 9.7 ppm in the shellcrosslinked F127 spectrum was weaker than that in the F127CHO spectrum, confirming that the F127-CHO was successfully crosslinked by cystamine dihydrochloride. The crosslinking degree of micelle was determined by the ninhydrin assay and about 60% of the aldehydes were crosslinked. The quantity of the crosslinker we used is 1.5 times the molar of F127-CHO. We tried some experiments, the maximum degree of crosslinking of micelles was obtained at this molar ratio. F127 self-aggregates to micelle in aqueous solution because of the central hydrophobic PPO block and hydrophilic PEO blocks. The TEM micrograph (Fig. 2a) shows F127 forms micelle at 37 °C with a particle size of 26 nm. The shell-crosslinked F127 was prepared by solvent exchange. The TEM image (Fig. 2b) suggests that the shell-crosslinked micelle has a spherical morphology and well dispersed. DLS result (Fig. 2c) conforms that the micelle has an average diameter of 94 nm with narrow distribution (PDI = 0.07). F127-CHO micelles were shell-crosslinked by linking up the terminal aldehydes (on the PEO) with cystamine dihydrochloride. Therefore, crosslinking may occurred both in inter-micelles and intramicelles. In addition, cystamine dihydrochloride was not only the shell-crossliker but also chain extender. The shell-crosslinked
Scheme 1. Synthetic route of F127-CHO, shell-crosslinked F127, ODex and micelle/hydrogel composite and the schematic representation of drug loaded shell-crosslinked F127 micelle/hydrogel composites.
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Fig. 1. The 1H NMR spectra of F127-CHO and shell-crosslinked F127.
F127 micelles with more molecular chain formed the close packing of micelles, therefore crosslinked micelles have larger volume than the uncrosslinked one. The size obtained from TEM is a little smaller than that from DLS because of the shrinkage of micelle. It was reported that the nanocarriers with a diameter smaller than 200 nm could passively accumulated in the tumor sites via the EPR effect [39,40], which provides the possibility for the targeted delivery of drugs to specific tumor sites. Zeta potential is defined as the potential difference between solution and the outer layer of micelle, a significant measurement of the stabilization provided by electrostatic stabilization. The zeta potential of the shell-crosslinked micelle and Cur-loaded micelle in distilled water are 2.4 and 18.8, respectively. The relative higher zeta potential provides better suspension stability and the negative potential provides micelle with long term stability. A fluorescence technique using pyrene was monitored to measure the critical micelle concentration (CMC) of shell-crosslinked micelle and demonstrates the ability of loading hydrophobic drugs after crosslinking. As we know, when micelles form, pyrene gives priority to be distributed in the hydrophobic core instead of outer shell, and therefore the environment of pyrene turns from polar to nonpolar. Fig. 3 shows the change in the intensity ratio of I384/I373 in pyrene excitation spectra which are highly dependent upon the hydrophobicity of the environment. There is no obvious change in the intensity ratio at lower concentration, but emerges an increases above 0.0015 mg mL1, which indicating that the incorporation of pyrene in the hydrophobic core of the shellcrosslinked micelle. The CMC (about 0.0015 mg mL1) of the shell-crosslinked micelle is smaller than that of the uncrosslinked one (about 0.035 mg mL1) [35]. Shell-crosslinked F127 micelle shows better thermodynamic stabilities and would be still intact undergoing dilution.
3.3. Characterization of CS/ODex hydrogel Sodium periodate was used to prepare ODex by cleaving the vicinal glycols and converting them to dialdehydes. Then ODex could be a macromolecular cross-linker for those polymers bearing free amino groups to form hydrogels. The oxidation degree of ODex was about 47% according to the result of hydroxylamine hydrochloride titration, that is, there were 47 per 100 sugar units had been oxidized into aldehydes. The FTIR spectra of CS, Dex, ODex and CS/ODex are presented in Fig. 4. In the spectrum of CS, the characteristic peak at 1599 cm1 is attributed to ANH2 bending vibration. Compared with the spectrum of Dex, the spectrum of ODex shows a new peak at 1735 cm1 (C@O stretching of the aldehydes), which indicates that some of the primary alcohols have been successfully oxidized to aldehydes. As the aldehydes groups formed hemiacetal, the peak at 1735 cm1 is inconspicuous. In the spectrum of CS/ODex, the peaks at 1735 and 1599 cm1 almost disappear while the characteristic peak at 1576 cm1 (ACNA) appears, suggesting the aldehydes groups in the ODex have reacted with the amino groups in the CS. The scanning electron morphologies (SEM) of the hydrogel (Gel 3) by virtue of the freeze-drying step are presented in Fig. 5. Fig. 5a shows the surface image and the interior morphology of the hydrogel, respectively. The surface of the hydrogel is not very smooth with many pores, and the internal of the hydrogel shows a continuous and porous structure with an average pore size of 50–80 nm, resembling other natural macromolecular hydrogel system structure [41,42]. The porous structure offers hydrogel enough sufficient internal space for the encapsulation of drugs and the pores will act as channels for drugs release. The other freeze-dried hydrogels with a relatively same crosslink density display the same morphology (data not shown).
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Fig. 2. TEM images of F127 micelles (a) and shell-crosslinked F127 micelles (b); size distributions of shell-crosslinked F127 (c); digital photographs (d) of water (1), shellcrosslinked F127 aqueous solution (2), Cur aqueous solution (3) and Cur-loaded shell-crosslinked F127 aqueous solution (4).
Fig. 3. Intensity ratios (I384/I373) from pyrene excitation spectra as a function of shell-crosslinked F127 concentration in aqueous.
Fig. 4. FTIR spectra of CS (a), Dex (b), ODex (c) and CS/ODex (Gel 3) (d).
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Fig. 5. SEM images of the surface (a) and the interior (b) of the Gel 3 (2 wt%).
3.4. Gelation time of the injectable micelle/hydrogel dual drug delivery system An ideal injectable hydrogel can be injected through a syringe and then gel rapidly at a specific location. Therefore, the gelation time is an important index for injectable hydrogel. Fast gelation would make the injection difficult by clogging the needle, while slow gelation would result in the gel precursor be diluted with body fluids and form delocalized gel [43,44]. The gelation time of the micelle/hydrogel system was determined at 37 °C and the results are displayed in Fig. 6. The gelation of the micelle/hydrogel system occurred in a short time within 2 min. The higher the concentration of CS or gel precursor are, the shorter the gelation time took. With increasing of the gel precursor concentration, the reactive groups (aldehydes and amino groups) increased, promoting a fast reaction. For the same gel precursor concentration (such as 2 wt%) used, changing the volume ratio of CS and ODex has little effect on the concentration of reactive groups as there are sufficient reactive groups (aldehydes and amino) on the macromolecular chains. However, the viscosity of gel precursor solution was greatly influenced. The higher the proportion of CS is, the higher the viscosity of the precursor solution, which results in slow flow of the hydrogel and an apparent rapid gelation. Considering the potential application of the micelle/hydrogel system in clinic, hydrogels with concentration of 1.0 wt%, 1.5 wt% and 2.0 wt% are prepared in this study, as an ideal injectable hydrogel should be injected through a syringe and then gel rapidly at a specific location.
Fig. 6. Gelation times of Gel 1 (1.0 wt%), Gel 2 (1.5 wt%) and Gel 3 (2.0 wt%) with different ratio of CS/ODex (1/1, 1/2 and 2/1) at 37 °C.
3.5. Cytotoxicity As the biocompatibility of the biomaterials is extremely important for their future biomedical applications, cell viability assay against HeLa cells was used to preliminarily evaluate the cytotoxicity of the hydrogel. The cell viability was 87.4%, 91.3% and 89.9% for Gel 1, Gel 2 and Gel 3, respectively. These results indicate that the cells with different hydrogel concentration are more than 85% viable after 48 h incubation, demonstrating the hydrogel is low toxic for living cells. This result is reasonable because CS and Dex have been widely used as biomedical materials due to their excellent biocompatibility, and F127 is also a low-toxic synthetic material which has been approved by FDA. 3.6. Drug delivery behavior of the injectable micelle/hydrogel dual drug delivery system The entrapment efficiency (EE) of Cur was about 39.46% and the drug-loading (DL) was as much as 7.59%. The final concentration of Fu was 2.5 mg mL1 in the hydrogel. We first examined the release behaviors of Cur and Fu from the hydrogels with different concentration (Gel 1, Gel 2 and Gel 3) in PBS at 37 °C. From Fig. 7a, Cur shows a sustained and relatively slow release behavior. In 9 d, about 11.87%, 16.84% and 24.86%% of Cur were released from Gel 1, Gel 2 and Gel 3, respectively. From Fig. 7b, Fu shows a sustained but relatively fast release behavior. In 9 d, about 51.87%, 63.48% and 75.80% of Fu were released from Gel 1, Gel 2 and Gel 3, respectively. The release rate of the two drugs decreases with the increase of the hydrogel concentration which can be explained by the increase of crosslinking degree. A higher crosslinking density makes a decrease in gel pore size of hydrogel, resulting in slow diffusion of drugs in the hydrogel. Fu was encapsulated directly in the pores of the hydrogel, it can be easily released from the hydrogel; while Cur was loaded in the micelle, the release of Cur needs overcome two barriers (micelle and hydrogel). Therefore the release rate of Fu is faster than that of Cur. For the same concentration of the hydrogel (2 wt%), we also examined the release behaviors of the two drugs as a function of the ratio of CS/ODex and the release profiles are shown in the Fig. 8. As there are sufficient reactive groups (aldehydes and amino) to form the hydrogel network with high crosslinking density, changing the ratio of CS/ODex has little effect on the crosslinking density, the release behaviors of the two drugs are similar at different ratios of CS/ODex. Therefore, the release rate of the drugs can be adjusted by the concentration of the hydrogel. Both the shell of the micelle and the hydrogel was crosslinked by Schiff base. It is known that the Schiff base bond is pHsensitive with the hydrolysis in an acidic environment, therefore, the drugs release experiments of the hydrogels were tested at pH 7.4, 5.0 and 3.0, respectively. Fig. 9 reveals that the release rate of the drugs from the hydrogels is influenced by the pH value.
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Fig. 7. The release behaviors of Cur (a) and Fu (b) from Gel 1 (1.0 wt%), Gel 2 (1.5%) and Gel 3 (2.0 wt%) at 37 °C.
Fig. 8. The release behaviors of Cur (a) and Fu (b) from Gel 3 at 37 °C as a function of the ratio of CS/ODex (1/1, 1/2 and 2/1).
Fig. 9. The release behaviors of Cur (a) and Fu (b) from Gel 3 at 37 °C as a function of pH (pH 7.4, pH 5.0 and pH 3.0).
The lower the pH value is, the faster the drugs were released. For Cur, about 17.66% (pH 5.0) and 18.80% (pH 3.0) have been released in 9 d, more than the amount released at pH 7.4 (about 11.87%). The similar situation occurred for Fu. Therefore, the hydrogel prefers to release drugs in an acid environment, which is highly desirable for treatment of cancer, since tumor tissues are known to be acidic [45].
Finally, we tested the release behaviors of the two drugs exposed to different concentration of DTT at 0 M, 5 mM and 10 mM, as the presence of redox-sensitive disulfide bonds on the shell of the micelle. From Fig. 10a, it is apparent that the release rate of Cur increased as the amount of DTT in the media increased. With higher DTT concentration of 5 mM and 10 mM, about 19.52% and 27.83% were released in 9 d, compared to 11.87% of the group
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Fig. 10. The release behaviors of Cur (a) and Fu (b) from Gel 3 at 37 °C as a function of the concentration of DTT (0 M, 5 mM, 10 mM).
with 0 mM DTT. DTT as a reducing agent, could damage the structure of micelle by cleaving the disulfide bonds, therefore reductive stimulus could accelerate the release of Cur. In order to study the influence of DTT to micelles, Cur release experiments from micelles with or without DTT were explored. The results (Fig. S1) showed that Cur release was faster in solution with 10 mM DTT than that without DTT, indicating that DTT influenced the drug release behavior. The release behavior of Fu (Fig. 10b) also shows redoxdependence. The fastest release rate of Fu from the hydrogel appears in the higher DTT concentration. It is reasonable because media would be acidic in the presence of DTT, which promotes the degradation of the gel. The micelle/hydrogel dual drug delivery system is highly desirable for treatment of cancer, since the concentration of glutathione in tumor cells is higher than the level in the bloodstream and healthy cells, which can cleave disulfide bonds like DTT [46]. In vitro drug delivery results reveal that this smart drug carrier has a long-term and sustained drug release behavior and has pH and redox sensitivity. This carrier is highly desirable for treatment of cancer, since tumor tissues are known to be acidic, and the concentration of glutathione in tumor cells is higher than the level in the bloodstream and healthy cells, which can cleave disulfide bonds like DTT. Moreover, the release rate of the drugs could be adjusted by changing the hydrogel concentration. Therefore, it has potential applications for clinical treatment.
3.7. Degradation behavior of the injectable micelle/hydrogel system To test whether the micelle/hydrogel system is biodegradable, the weight loss of the system was examined with incubation time at 37 °C. As shown in Fig. 11, all tested hydrogels showed biodegradability. After 10 d of incubation, the weight loss of Gel 1, Gel 2 and Gel 3 was 45.8%, 54.87% and 64.93%, respectively. The degradation of the hydrogels is attributed to the hydrolysis of Schiff base bonds. By our observations, the degradation process of the gels was from outside to inside. Firstly, the surface of the gels became rough, then the volume of the gels decreased gradually, nearly maintaining the original shape. As acidic environment is helpful for the hydrolysis of Schiff base bonds and medium containing DTT is acidic, other two hydrogels cultured at pH 3.0 or with 20 mM DTT have rapid degradation behaviors in contrast to the group which cultured at pH 7.4 without DTT. Therefore, the micelle/hydrogel will be subject degradation in vivo as modeled by the in vitro degradation.
Fig. 11. The degradation behaviors of the shell-crosslinked F127 micelle/hydrogel composites under different conditions. All the tests were carried out at 37 °C.
4. Conclusions In attempts to develop a smart dual drug delivery system to improve therapeutic outcome of combined therapy in treatment of cancer, we have prepared an injectable micelle/hydrogel composite, which consists of curcumin (Cur) loaded shell-crosslinked F127 micelle and 5-fluorouracil (Fu) dispersed CS/ODex hydrogel. Both of the shell-crosslinked F127 micelle and CS/ODex hydrogel were prepared based on Schiff base bonds, which provides pHsensitivity and biodegradability for the system. In addition, disulfide bonds were introduced in the shell of F127 micelle, allowing for redox-sensitivity of the system. Drug delivery results reveal that the drug carrier has pH and redox sensitivity. In vitro degradation experiments show that the hydrogel is able to degrade by the hydrolysis of Schiff base bonds. MTT assay shows that the hydrogels display excellent compatibility with cells. These studies indicate that the micelle/hydrogel dual drug delivery system could load and deliver hydrophobic and hydrophilic drugs simultaneously and have potential application in cancer treatment. Acknowledgments The authors gratefully acknowledge the financial support of the National Natural Science Foundation of China (Grant Nos.
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51273086, 51473072, 51503091) and Special Doctorial Program Fund from the Ministry of Education of China (Grant No. 20130211110017), and the Fundamental Research Funds for the Central Universities (Grant No. lzujbky-2015-26). Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.cej.2015.11.015. References [1] D.J. Beebe, J.S. Moore, J.M. Bauer, Q. Yu, R.H. Liu, C. Devadoss, B.-H. Jo, Functional hydrogel structures for autonomous flow control inside microfluidic channels, Nature 404 (2000) 588–590. [2] M.N.R. Kumar, A review of chitin and chitosan applications, React. Funct. Polym. 46 (2000) 1–27. [3] M. Fassnacht, M. Terzolo, B. Allolio, E. Baudin, H. Haak, A. Berruti, S. Welin, C. Schade-Brittinger, A. Lacroix, B. Jarzab, Combination chemotherapy in advanced adrenocortical carcinoma, N. Engl. J. Med. 366 (2012) 2189–2197. [4] C. Gong, C. Wang, Y. Wang, Q. Wu, D. Zhang, F. Luo, Z. Qian, Efficient inhibition of colorectal peritoneal carcinomatosis by drug loaded micelles in thermosensitive hydrogel composites, Nanoscale 4 (2012) 3095–3104. [5] Z. Liu, P. Yao, Versatile injectable supramolecular hydrogels containing drug loaded micelles for delivery of various drugs, Polym. Chem. 5 (2014) 1072–1081. [6] H. Zhang, G. Wang, H. Yang, Drug delivery systems for differential release in combination therapy, Expert Opin. Drug Deliv. 8 (2011) 171–190. [7] H. Zhang, C. Zhao, H. Cao, G. Wang, L. Song, G. Niu, H. Yang, J. Ma, S. Zhu, Hyperbranched poly(amine-ester) based hydrogels for controlled multi-drug release in combination chemotherapy, Biomaterials 31 (2010) 5445–5454. [8] Z. Cao, H. Wu, J. Dong, G. Wang, Quadruple-stimuli-sensitive polymeric nanocarriers for controlled release under combined stimulation, Macromolecules 47 (2014) 8777–8783. [9] J. Liu, Y. Huang, A. Kumar, A. Tan, S. Jin, A. Mozhi, X.-J. Liang, pH-sensitive nanosystems for drug delivery in cancer therapy, Biotechnol. Adv. 32 (2014) 693– 710. [10] F. Kratz, A. Warnecke, Finding the optimal balance: challenges of improving conventional cancer chemotherapy using suitable combinations with nanosized drug delivery systems, J. Control. Release 164 (2012) 221–235. [11] S. Chen, F. Jiang, Z. Cao, G. Wang, Z.-M. Dang, Photo, pH, and thermo tripleresponsive spiropyran-based copolymer nanoparticles for controlled release, Chem. Commun. 51 (2015) 12633–12636. [12] A.A. Khan, J. Mudassir, N. Mohtar, Y. Darwis, Advanced drug delivery to the lymphatic system: lipid-based nanoformulations, Int. J. Nanomed. 8 (2013) 2733. [13] J.A. Floyd, A. Galperin, B.D. Ratner, Drug encapsulated polymeric microspheres for intracranial tumor therapy: a review of the literature, Adv. Drug Deliv. Rev. (2015). [14] C.E. Probst, P. Zrazhevskiy, V. Bagalkot, X. Gao, Quantum dots as a platform for nanoparticle drug delivery vehicle design, Adv. Drug Deliv. Rev. 65 (2013) 703–718. [15] P. Galvin, D. Thompson, K.B. Ryan, A. McCarthy, A.C. Moore, C.S. Burke, M. Dyson, B.D. MacCraith, Y.K. Gun’ko, M.T. Byrne, Nanoparticle-based drug delivery: case studies for cancer and cardiovascular applications, Cell. Mol. Life Sci. 69 (2012) 389–404. [16] L. Chen, T. Jiang, C. Cai, L. Wang, J. Lin, X. Cao, Polypeptide-based ‘‘smart” micelles for dual-drug delivery: a combination study of experiments and simulations, Adv. Healthcare Mater. 3 (2014) 1508–1517. [17] D. Ma, J. Lin, Y. Chen, W. Xue, L.-M. Zhang, In situ gelation and sustained release of an antitumor drug by graphene oxide nanosheets, Carbon 50 (2012) 3001–3007. [18] M. Zhu, Y. Zhu, L. Zhang, J. Shi, Preparation of chitosan/mesoporous silica nanoparticle composite hydrogels for sustained co-delivery of biomacromolecules and small chemical drugs, Sci. Technol. Adv. Mater. 14 (2013) 045005. [19] J.-K. Cho, K.-Y. Hong, J.W. Park, H.-K. Yang, S.-C. Song, Injectable delivery system of 2-methoxyestradiol for breast cancer therapy using biodegradable thermosensitive poly(organophosphazene) hydrogel, J. Drug Target. 19 (2011) 270–280. [20] G. Chang, T. Ci, L. Yu, J. Ding, Enhancement of the fraction of the active form of an antitumor drug topotecan via an injectable hydrogel, J. Control. Release 156 (2011) 21–27. [21] L. Li, J. Gu, J. Zhang, Z. Xie, Y. Lu, L. Shen, Q. Dong, Y. Wang, Injectable and biodegradable pH-responsive hydrogels for localized and sustained treatment of human fibrosarcoma, ACS Appl. Mater. Interfaces 7 (2015) 8033–8040.
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Injectable shell-crosslinked F127 micelle/hydrogel composites with pH and redox sensitivity for combined release of anticancer drugs Nannan Gao a, Shaoyu Lü a,⇑, Chunmei Gao a, Xinggang Wang a,b, Xiubin Xu a, Xiao Bai a, Chen Feng a, Mingzhu Liu a,⇑ a State Key Laboratory of Applied Organic Chemistry, Key Laboratory of Nonferrous Metal Chemistry and Resources Utilization of Gansu Province, Department of Chemistry, Lanzhou University, Lanzhou 730000, People’s Republic of China b Research Institute of Lanzhou Petrochemical Corporation, Petrochina Lanzhou Petrochemical Company, Lanzhou 730060, People’s Republic of China
h i g h l i g h t s A novel dual drug carrier is constructed by injectable micelle/hydrogel composite. This drug carrier has redox-sensitivity and pH-sensitivity. The preparation is inexpensive, facile and fast. The system can load and deliver hydrophobic and hydrophilic drugs simultaneously. The system can be used in combination chemotherapy for cancer treatment.
a r t i c l e
i n f o
Article history: Received 28 July 2015 Received in revised form 15 October 2015 Accepted 6 November 2015 Available online 14 November 2015 Keywords: Micelle/hydrogel composite Shell-crosslinked micelle Injectable Stimuli-response Dual drug carrier
a b s t r a c t Combination chemotherapy as a preferred strategy for cancer treatment has been widely developed recently. However, single drug carriers of the combination chemotherapy systems are difficult to address the performance requirement that delivers hydrophilic and hydrophobic drugs simultaneously. In this work, a smart dual drug delivery system of injectable micelle/hydrogel composite was developed. This smart drug delivery system consists of curcumin (Cur) loaded shell-crosslinked F127 micelle and 5fluorouracil (Fu) dispersed chitosan/oxidized dextran (CS/ODex) hydrogel. Both of the shell-crosslinked F127 micelle and CS/ODex hydrogel were prepared based on Schiff base bonds, which provides pHsensitivity and biodegradability for the system. In addition, disulfide bonds were introduced in the shell of F127 micelle, allowing for redox-sensitivity of the system. The system was detailed characterized by 1H NMR, FTIR, TEM, SEM, DLS and fluorescence spectroscopy. Independent release behaviors of the two drugs are observed under conditions of different precursor concentration, different CS/ODex proportion, different pH and different concentration of DTT. The results show that the drug carrier has pH and redox sensitivity and the release rate can be adjusted by the concentration of the precursor solution. In vitro degradation experiment and cytotoxicity assay determine that the system has good biodegradability and biocompatibility. These results indicated that the injectable shell-crosslinked F127 micelle/hydrogel composite may load and deliver hydrophobic and hydrophilic drugs simultaneously and find application in combination chemotherapy for cancer treatment. Ó 2015 Elsevier B.V. All rights reserved.
1. Introduction Chemotherapy has been widely used to treat with established cancer in the clinic along with surgical excision and irradiation [1]. However, traditional chemotherapy is always accompanied by many side effects: most drugs are of high toxicity but lack of ⇑ Corresponding authors. Tel.: +86 931 8912387; fax: +86 931 8912582. E-mail addresses: [email protected] (S. Lü), [email protected] (M. Liu). http://dx.doi.org/10.1016/j.cej.2015.11.015 1385-8947/Ó 2015 Elsevier B.V. All rights reserved.
site-specific, drug level is uncontrolled and therefore need repeated administration and so on. The facts above bring tremendous physical and mental harm to patients [2]. Due to these obstacles, combination chemotherapy as a preferred strategy for cancer treatment has replaced traditional chemotherapy, which is characterized by the synergistic effects using multiple drugs of different therapeutic effects, and thus improves the curative effect [3]. Drug delivery system (DDS), as an useful adjunct to combination chemotherapy, has been widely researched over the last few
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decades [4–8], as DDS could overcome the problems of conventional administration by enhancing drug solubility, prolonging duration time, minimizing side effects, protecting drug bioactivity and so on [9–11]. Accordingly, a variety of drug delivery systems (DDSs), such as lipids [12], microspheres [13], nanoparticles [14,15], micelles [16] and gels [17] have been developed. However, single drug carriers of the combination chemotherapy systems are difficult to address the performance requirement that delivers different drugs simultaneously. Therefore, developing a dual drug delivery system (DDDS) which can simultaneously load multiple drugs and controlled release of each drug is desired. Zhu et al. designed composite hydrogels based on chitosan and mesoporous silica which could simultaneously release gentamicin and bovine serum albumin in a sustained manner [18]. Gong et al. developed poly-(3-caprolactone)–poly(ethylene glycol)–poly(3-ca prolactone)-based micelle/hydrogel carrier to deliver paclitaxel and fluorouracil simultaneously for the treatment of colorectal peritoneal carcinomatosis [4]. Despite a few examples of DDDS for hydrophobic and hydrophilic drug delivery application, the possible toxicity of the agents these systems used, undesirable side products from gel degradation, and the complex synthesis process compromise their real use. Here, we developed an injectable micelle/hydrogel dual drug delivery system based on natural polysaccharide through an inexpensive, facile and fast method. Injectable hydrogels are appealing for entrapping hydrophilic drugs with the web-like structure and release drugs directly to the target site (local drug delivery) with longer drug retention and larger drug accumulation in target cells than intravenous administration [19–21]. Micelles are one of the ideal vehicles for poorly water-soluble anticancer drugs with their intrinsic core–shell structure and could be passively targeted to the tumor site due to enhanced permeability and retention (EPR) effect [22–24]. 5-Fluorouracil (Fu), as a hydrophilic anticancer drug, was loaded in hydrogel. Curcumin (Cur), as a hydrophobic anticancer drug, was loaded in micelles. Multiple drugs of different therapeutic effects were used to improve the curative effect. The injectable micelle/hydrogel delivery system consists of chitosan (CS), dextran (Dex), and Pluronic F127 (PEO–PPO–PEO) micelles. Chitosan, a positively charged natural polysaccharide, composed of b-(1,4)-linked D-glucosamine
and N-acetyl-D-glucosamine units [25], has been widely used as hydrogel material for a long time due to excellent biocompatibility, biodegradability and antimicrobial activity [26,27]. Dextran (Dex) is another natural biocompatible material widely used in biomedical applications. The micelle is prepared using F127, which is commercially available and contains hydrophilic block poly(ethylene oxide) (PEO) and hydrophobic block poly(propylene oxide) (PPO) [28]. The gelation of the composite system is attributed to the Schiff base reaction between aldehydes or amine groups on the three components. The Schiff base reactions have drawn great attention in preparation of DDSs with their attractive properties: aldehydes and amino groups on polysaccharides can be easily acquired through facile methods [29,30]; the reactions take place at physiological conditions without toxic chemical cross-linking agents and water is the only byproduct [31]; hydrogels crosslinked by Schiff base bonds could gradually degrade because of the hydrolysis of Schiff base bonds in vitro, and acid environment could accelerate the degradation. Therefore, the Schiff base based micelle/injectable hydrogel would be a biodegradable and pH-sensitive delivery system, which is preferable for anticancer drugs carriers because of most cancer tissues hold a much lower pH compared to the normal tissues [32]. As we know, above the critical micelle concentration (CMC), F127 can self-assemble into micelle. PPO serves as a hydrophobic core providing a microenvironment for loading of hydrophobic
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drugs, PEO serves as a hydrophilic outer shell making micelle stable and unrecognizable by the reticuloendothelial system [33]. Researchers have demonstrated that Pluronic F127 is a promising multidrug resistance (MDR) reversal material because it can inhibit P-glycoprotein (P-gp) efflux, inhibit the glutathione (GSH)/glutathione (GST) detoxification system, enhance proapoptotic signaling, decrease anti-apoptotic defense in MDR cells and so on [34]. However, there are some defects in the use of Pluronic F127 micelle in the blood circulation. F127 is not able to keep micelle structure at low concentration, then drugs entrapped in the core will leak out. In addition, F127 tends to be prematurely excreted by the kidneys or penetrate into healthy endothelia, due to its relatively small particle size. Moreover, F127 does not have environmental sensitivity to smart release of drug [35]. Based on the above considerations, we prepared shellcrosslinked F127, using cystamine dihydrochloride as a crosslinking agent. After shell crosslinking, the structure of F127 micelle would be fixed and consequently more stable than before. In addition, the disulfide bonds on the shell of F127 micelle allow for redox-sensitivity of micelle, which could be cleaved by GSH in vivo. As known, GSH is abundant (mM) in the cytosol and subcellular compartments and much more in cancer cells, but very low (lM) in plasma [36,37]. Such micro-environmental differences allow that F127 micelles keep stable under physiological conditions and trigger sudden burst of drugs through the shell breaking in tumor cells. Therefore, the smart injectable micelle/hydrogel dual drug delivery system could simultaneously load hydrophobic and hydrophilic drugs, and local deliver drugs through injectable hydrogel and intracellular release of drugs through shellcrosslinked F127 micelle. The pH and redox potential stimuli responsive drug delivery behaviors relying on the significant pH and redox potential differences between the tumor and normal tissues were determined. 2. Materials and methods 2.1. Materials Chitosan (CS, degree of deacetylation was 88.0%) was purchased from Sinopharm Chemical Reagent Co., Ltd. and used without further purification. Dextran (Dex, Mw 70 k) and Dess–Martin periodinane was provided from Aladdin biological technology Co., Ltd. (Shanghai, China). Pluronic F127 was obtained from Shanghai Sangon Biological Engineering Technology & Services Co., Ltd. (Shanghai, China). Dithiothreitol (DTT), cystamine dihydrochloride, sodium periodate and other chemicals were used as received. 2.2. Preparation of shell-crosslinked F127 micelle The shell-crosslinked F127 micelle was carried out according to a previously reported procedure in the literature [38]. The Dess– Martin periodinane was employed in the process of chemical conversion of the terminal alcohols on F127 into aldehydes (F121CHO). The Dess–Martin periodinane (0.5 g) was added to a solution of F127 (5.16 g) in methylene chloride (410 mL) and the resulting solution was stirred for 24 h at 40 °C. After removal of most of the solvent, the product was isolated by precipitation into hexane. The extent of oxidation, which defined as the number of oxidized F127-CHO residues per 100 original F127 feeds, was determined by the proton nuclear magnetic resonance (1H-NMR) analysis. F127-CHO (1.8 g) was dissolved in tetrahydrofuran (THF, 100 mL). This solution was added dropwise to deionized water (350 mL) and stirred for 30 min. THF was then removed using a rotary evaporator at 40 °C. Subsequently, cystamine dihydrochloride (0.21 mmol) was added dropwise to the above micelle
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solution. The mixtures were stirred for 1 h at 30 °C to crosslink the aldehydes termini on F127-CHO. 10 mL of the resulting reaction mixture was used to determine the degree of crosslinking according to the previous literature [31]. The unreacted cystamine dihydrochloride was determined, and the crosslinking degree of micelle was calculated. The products were dialyzed for 3 days and lyophilized. 2.3. Preparation of CS solution First, CS (3 g) was dispersed in deionized water (80 mL) while stirring. Then, acetic acid was added dropwise to the solution until a clear solution was obtained. Afterwards, the mixture was cooled to 4 °C and the pH was adjusted to 6.2 using ice aqueous sodium bicarbonate. Finally, a corresponding amount of deionized water was added to the reaction system and the final concentration of CS was 2 wt%. The CS solution was stored at 4 °C before use. 2.4. Preparation of oxidized dextran (ODex) Sodium periodate (13.2 g in aqueous solution) was added to a Dex solution (1% (w/v), 200 mL) and stirred at room temperature for 2 h in the dark. Diethylene glycol (4 mL) was added to quench any unreacted sodium periodate. After that the solution was stirred for 1 h and dialyzed for 3 days and lyophilized. The oxidation degree of ODex, which defined as the percentage of Dex structure units that have been oxidized, was determined by hydroxylamine hydrochloride titration. First, hydroxylamine hydrochloride (3 g), ethanol (8 mL), deionized water (5 mL), sodium hydroxide (0.84 g) in ethanol (20 mL) and indicator bromophenol blue solution (0.45% (w/v), 5 mL) were mixed to obtain hydroxylamine hydrochloride solution. Then, ODex (0.100 g), hydroxylamine hydrochloride solution (5 mL), ethanol (5 mL) and deionized water (5 mL) were mixed in conical flask and stirred for 15 min. Finally, the solution was titrated by standard hydrochloric acid solution until a yellow color was obtained. Dex was employed as a control group and the values were average of three experiments. The formula of oxidation degree of ODex are as follows:
Degree of oxidation ð%Þ ¼
162 ðV 2 V 1 Þ C 103 100% 2 MW
where the C (mol L1) is the concentration of the standard hydrochloric acid solution, MW is average molecular weight of Dex, the V1 and V2 (mL) are the volume of the standard hydrochloric acid solution used by the test group and the control group, respectively.
2.6. Characterization The morphological of the micelle was examined by transmission electron microscope (TEM, JEM-1200EX/S, Hitachi, Japan) operating at an accelerating voltage of 200 kV. The samples were prepared by the micelle solution on a formvar coated copper grid and then incubated in oven for 15 min at 37 °C. The size of the micelle was analyzed via dynamic light scattering (DLS) using a laser light scattering spectrometer (ALV/SP-125). Scattered light was collected at a fixed angle of 90° for duration of 10 min. All data were averaged over three measurements. The zeta-potential (n) was measured by Malvern ZEN3600, England. The critical micelle concentration (CMC) of the shell-crosslinked micelle was evaluated by fluorescence spectroscopy (LS55, Perkin– Elmer, America), using pyrene as a probe. The excitation spectra of micelle solutions were scanned from 360 to 400 nm with an emission wavelength of 390 nm. The CMC was estimated as the crosspoint when extrapolating the intensity ratio of I384/I373. FTIR absorption spectra of CS, Dex, ODex and CS/ODex (Gel 3) were taken on a Fourier transform infrared (FTIR) spectrometer (Nicolet 670 FTIR, USA) over the region from 4000 to 400 cm1. The gelation time of the hydrogel was visually measured by the vial inversion method when the sample solutions were unable to be pipetted up and down. The morphology of the hydrogel after lyophilization was observed by a JSM-5600LV scanning electron microscopy (SEM), Japan. Before observation, all samples were fixed on aluminum stubs and coated with gold.
2.7. Cytotoxicity test The cytotoxicity of the injectable micelle/hydrogel dual drug delivery system was assessed by 3-(4,5-dimethylthiazol-2-yl)-2,5 -dipheny-ltetrazolium bromide (MTT) assay with HeLa cells. The hydrogel precursor solutions of Gel 1, Gel 2 and Gel 3 (CS solution, ODex solution and micelle solution) which had been sterilized just before use were mixed and injected into a 96-well plate to form hydrogels in the bottom of the plate, respectively. The total volume of the precursor solutions was 20 lL. Then 160 lL of culture medium containing HeLa cells was added into each well at a density of 10,000 cells/well. After culturing for 48 h in a humidified atmosphere with 5% CO2 at 37 °C, the medium was removed carefully and replaced by MTT reagent for another 4 h. Then the cells were dissolved by 200 lL of DMSO. Cell viability was determined by the absorbance values at 490 nm which was measured with a microplate reader (Tecan, Mannedorf, Switzerland).
2.5. Preparation of injectable micelle(Cur)/hydrogel(Fu) dual drug delivery system To prepare Cur-loaded micelle, Cur and shell-crosslinked F127 micelle (0.1 g) were dissolved in 5 mL of distilled water with sonication for 30 min and subsequently stirred for 48 h. The free Cur was then removed by centrifugation. The loading amount of Cur in micelle was determined by testing the weight and the resulting drug-loaded micelle solution was stewing before use. The micelle(Cur)/hydrogel(Fu) delivery system was prepared as follows. First, Fu was dissolved in ODex solution. Then, Fu-loaded ODex solution, CS solution and Cur-loaded micelle solution were mixed at 4 °C (the concentration ratio of CS and ODex was 1/1). Finally, the mixture was heated to 37 °C for gelation. The final concentration of Fu and micelle was 2.5 mg mL1 and 3.3 mg mL1, respectively. The concentration of the hydrogel was 1.0 wt% (Gel 1), 1.5 wt% (Gel 2) and 2.0 wt% (Gel 3), respectively.
2.8. Drug release study of the injectable micelle/hydrogel dual drug delivery system in vitro Gel 1, Gel 2 and Gel 3 were directly immersed in 8 mL buffer (ethanol/phosphate buffer solution (80:20, v/v)), and then incubated at 37 °C under oscillation at 80 rpm. At a specific time, the release medium (3 mL) was withdrawn and replaced with fresh buffer. UV/vis absorbance (Lambda 35, Perkin–Elmer, America) was recorded at 268 nm for Fu and 430 nm for Cur. The release amount of drugs was determined via the standard curves of each drug in the same buffer. Gel 3 with various CS/ODex mass ratio (1/1, 1/2 and 2/1), various pH values (pH 7.4, pH 5.0 and pH 3.0) and various concentration of DTT (0 M, 5 mM, 10 mM) were also tested.
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2.9. Degradation behavior of the injectable micelle/hydrogel system in vitro The in vitro degradation of the micelle/hydrogel dual drug delivery system was examined as follows. 3 mL of the different hydrogels were immersed in tubes with 10 mL PBS (pH 7.4) at 37 °C. At predetermined time intervals, the PBS was taken out and excess surface water was gently wiped with filter paper. After that, the hydrogel was weighed, and fresh buffer was replenished. The weight loss ratio of the system was calculated.
2.10. Statistical analysis Statistical analysis of data was performed by one-way analysis of variance (ANOVA), assuming confidence level of 95% (p < 0.05) for statistical significance.
3. Results and discussion 3.1. Preparation of injectable micelle/hydrogel composite In order to improve the therapeutic outcome of combined therapy in treatment of cancer, we developed a smart dual drug delivery system of injectable micelle/hydrogel composite. As shown in Scheme 1, this smart drug delivery system consists of shellcrosslinked F127 micelle and CS/ODex hydrogel. To obtain shellcrosslinked F127 micelle, parts of terminal alcohols on F127 were oxidized to aldehydes using the Dess–Martin periodinane. Then the F127-CHO micelle was shell-crosslinked by linking up the aldehydes with cystamine dihydrochloride. The final micelle was redox-responsive with the disulfide bonds. ODex was prepared by oxidizing Dex with sodium periodate, which could cleave the vicinal glycols and convert them to aldehydes. There were many amino groups on CS and some free amino groups on the shellcrosslinked F127 micelle that could crosslink with the aldehydes
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on ODex and form hydrogel. The micelle/hydrogel composite was generated by mixing micelle solution, CS solution and ODex solution via Schiff connection. The system would be pH-responsive due to the carbon–nitrogen double bonds.
3.2. Characterization of the shell-crosslinked F127 micelle Fig. 1 shows the 1H NMR spectra of F127-CHO and shellcrosslinked F127. The peak at 9.7 ppm in the F127-CHO spectrum was attributed to the hydrogen in aldehydes (ACHO), further demonstrating the primary alcohols of F127 were converted to aldehydes. From the peak areas of 9.7 ppm and 1.18 ppm (OCH2OH(CH3)O of F127-CHO) in the F127-CHO spectrum, we can estimate that about 85% of the proportion of alcohols in F127 was converted to aldehydes. The peak at 9.7 ppm in the shellcrosslinked F127 spectrum was weaker than that in the F127CHO spectrum, confirming that the F127-CHO was successfully crosslinked by cystamine dihydrochloride. The crosslinking degree of micelle was determined by the ninhydrin assay and about 60% of the aldehydes were crosslinked. The quantity of the crosslinker we used is 1.5 times the molar of F127-CHO. We tried some experiments, the maximum degree of crosslinking of micelles was obtained at this molar ratio. F127 self-aggregates to micelle in aqueous solution because of the central hydrophobic PPO block and hydrophilic PEO blocks. The TEM micrograph (Fig. 2a) shows F127 forms micelle at 37 °C with a particle size of 26 nm. The shell-crosslinked F127 was prepared by solvent exchange. The TEM image (Fig. 2b) suggests that the shell-crosslinked micelle has a spherical morphology and well dispersed. DLS result (Fig. 2c) conforms that the micelle has an average diameter of 94 nm with narrow distribution (PDI = 0.07). F127-CHO micelles were shell-crosslinked by linking up the terminal aldehydes (on the PEO) with cystamine dihydrochloride. Therefore, crosslinking may occurred both in inter-micelles and intramicelles. In addition, cystamine dihydrochloride was not only the shell-crossliker but also chain extender. The shell-crosslinked
Scheme 1. Synthetic route of F127-CHO, shell-crosslinked F127, ODex and micelle/hydrogel composite and the schematic representation of drug loaded shell-crosslinked F127 micelle/hydrogel composites.
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Fig. 1. The 1H NMR spectra of F127-CHO and shell-crosslinked F127.
F127 micelles with more molecular chain formed the close packing of micelles, therefore crosslinked micelles have larger volume than the uncrosslinked one. The size obtained from TEM is a little smaller than that from DLS because of the shrinkage of micelle. It was reported that the nanocarriers with a diameter smaller than 200 nm could passively accumulated in the tumor sites via the EPR effect [39,40], which provides the possibility for the targeted delivery of drugs to specific tumor sites. Zeta potential is defined as the potential difference between solution and the outer layer of micelle, a significant measurement of the stabilization provided by electrostatic stabilization. The zeta potential of the shell-crosslinked micelle and Cur-loaded micelle in distilled water are 2.4 and 18.8, respectively. The relative higher zeta potential provides better suspension stability and the negative potential provides micelle with long term stability. A fluorescence technique using pyrene was monitored to measure the critical micelle concentration (CMC) of shell-crosslinked micelle and demonstrates the ability of loading hydrophobic drugs after crosslinking. As we know, when micelles form, pyrene gives priority to be distributed in the hydrophobic core instead of outer shell, and therefore the environment of pyrene turns from polar to nonpolar. Fig. 3 shows the change in the intensity ratio of I384/I373 in pyrene excitation spectra which are highly dependent upon the hydrophobicity of the environment. There is no obvious change in the intensity ratio at lower concentration, but emerges an increases above 0.0015 mg mL1, which indicating that the incorporation of pyrene in the hydrophobic core of the shellcrosslinked micelle. The CMC (about 0.0015 mg mL1) of the shell-crosslinked micelle is smaller than that of the uncrosslinked one (about 0.035 mg mL1) [35]. Shell-crosslinked F127 micelle shows better thermodynamic stabilities and would be still intact undergoing dilution.
3.3. Characterization of CS/ODex hydrogel Sodium periodate was used to prepare ODex by cleaving the vicinal glycols and converting them to dialdehydes. Then ODex could be a macromolecular cross-linker for those polymers bearing free amino groups to form hydrogels. The oxidation degree of ODex was about 47% according to the result of hydroxylamine hydrochloride titration, that is, there were 47 per 100 sugar units had been oxidized into aldehydes. The FTIR spectra of CS, Dex, ODex and CS/ODex are presented in Fig. 4. In the spectrum of CS, the characteristic peak at 1599 cm1 is attributed to ANH2 bending vibration. Compared with the spectrum of Dex, the spectrum of ODex shows a new peak at 1735 cm1 (C@O stretching of the aldehydes), which indicates that some of the primary alcohols have been successfully oxidized to aldehydes. As the aldehydes groups formed hemiacetal, the peak at 1735 cm1 is inconspicuous. In the spectrum of CS/ODex, the peaks at 1735 and 1599 cm1 almost disappear while the characteristic peak at 1576 cm1 (ACNA) appears, suggesting the aldehydes groups in the ODex have reacted with the amino groups in the CS. The scanning electron morphologies (SEM) of the hydrogel (Gel 3) by virtue of the freeze-drying step are presented in Fig. 5. Fig. 5a shows the surface image and the interior morphology of the hydrogel, respectively. The surface of the hydrogel is not very smooth with many pores, and the internal of the hydrogel shows a continuous and porous structure with an average pore size of 50–80 nm, resembling other natural macromolecular hydrogel system structure [41,42]. The porous structure offers hydrogel enough sufficient internal space for the encapsulation of drugs and the pores will act as channels for drugs release. The other freeze-dried hydrogels with a relatively same crosslink density display the same morphology (data not shown).
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Fig. 2. TEM images of F127 micelles (a) and shell-crosslinked F127 micelles (b); size distributions of shell-crosslinked F127 (c); digital photographs (d) of water (1), shellcrosslinked F127 aqueous solution (2), Cur aqueous solution (3) and Cur-loaded shell-crosslinked F127 aqueous solution (4).
Fig. 3. Intensity ratios (I384/I373) from pyrene excitation spectra as a function of shell-crosslinked F127 concentration in aqueous.
Fig. 4. FTIR spectra of CS (a), Dex (b), ODex (c) and CS/ODex (Gel 3) (d).
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Fig. 5. SEM images of the surface (a) and the interior (b) of the Gel 3 (2 wt%).
3.4. Gelation time of the injectable micelle/hydrogel dual drug delivery system An ideal injectable hydrogel can be injected through a syringe and then gel rapidly at a specific location. Therefore, the gelation time is an important index for injectable hydrogel. Fast gelation would make the injection difficult by clogging the needle, while slow gelation would result in the gel precursor be diluted with body fluids and form delocalized gel [43,44]. The gelation time of the micelle/hydrogel system was determined at 37 °C and the results are displayed in Fig. 6. The gelation of the micelle/hydrogel system occurred in a short time within 2 min. The higher the concentration of CS or gel precursor are, the shorter the gelation time took. With increasing of the gel precursor concentration, the reactive groups (aldehydes and amino groups) increased, promoting a fast reaction. For the same gel precursor concentration (such as 2 wt%) used, changing the volume ratio of CS and ODex has little effect on the concentration of reactive groups as there are sufficient reactive groups (aldehydes and amino) on the macromolecular chains. However, the viscosity of gel precursor solution was greatly influenced. The higher the proportion of CS is, the higher the viscosity of the precursor solution, which results in slow flow of the hydrogel and an apparent rapid gelation. Considering the potential application of the micelle/hydrogel system in clinic, hydrogels with concentration of 1.0 wt%, 1.5 wt% and 2.0 wt% are prepared in this study, as an ideal injectable hydrogel should be injected through a syringe and then gel rapidly at a specific location.
Fig. 6. Gelation times of Gel 1 (1.0 wt%), Gel 2 (1.5 wt%) and Gel 3 (2.0 wt%) with different ratio of CS/ODex (1/1, 1/2 and 2/1) at 37 °C.
3.5. Cytotoxicity As the biocompatibility of the biomaterials is extremely important for their future biomedical applications, cell viability assay against HeLa cells was used to preliminarily evaluate the cytotoxicity of the hydrogel. The cell viability was 87.4%, 91.3% and 89.9% for Gel 1, Gel 2 and Gel 3, respectively. These results indicate that the cells with different hydrogel concentration are more than 85% viable after 48 h incubation, demonstrating the hydrogel is low toxic for living cells. This result is reasonable because CS and Dex have been widely used as biomedical materials due to their excellent biocompatibility, and F127 is also a low-toxic synthetic material which has been approved by FDA. 3.6. Drug delivery behavior of the injectable micelle/hydrogel dual drug delivery system The entrapment efficiency (EE) of Cur was about 39.46% and the drug-loading (DL) was as much as 7.59%. The final concentration of Fu was 2.5 mg mL1 in the hydrogel. We first examined the release behaviors of Cur and Fu from the hydrogels with different concentration (Gel 1, Gel 2 and Gel 3) in PBS at 37 °C. From Fig. 7a, Cur shows a sustained and relatively slow release behavior. In 9 d, about 11.87%, 16.84% and 24.86%% of Cur were released from Gel 1, Gel 2 and Gel 3, respectively. From Fig. 7b, Fu shows a sustained but relatively fast release behavior. In 9 d, about 51.87%, 63.48% and 75.80% of Fu were released from Gel 1, Gel 2 and Gel 3, respectively. The release rate of the two drugs decreases with the increase of the hydrogel concentration which can be explained by the increase of crosslinking degree. A higher crosslinking density makes a decrease in gel pore size of hydrogel, resulting in slow diffusion of drugs in the hydrogel. Fu was encapsulated directly in the pores of the hydrogel, it can be easily released from the hydrogel; while Cur was loaded in the micelle, the release of Cur needs overcome two barriers (micelle and hydrogel). Therefore the release rate of Fu is faster than that of Cur. For the same concentration of the hydrogel (2 wt%), we also examined the release behaviors of the two drugs as a function of the ratio of CS/ODex and the release profiles are shown in the Fig. 8. As there are sufficient reactive groups (aldehydes and amino) to form the hydrogel network with high crosslinking density, changing the ratio of CS/ODex has little effect on the crosslinking density, the release behaviors of the two drugs are similar at different ratios of CS/ODex. Therefore, the release rate of the drugs can be adjusted by the concentration of the hydrogel. Both the shell of the micelle and the hydrogel was crosslinked by Schiff base. It is known that the Schiff base bond is pHsensitive with the hydrolysis in an acidic environment, therefore, the drugs release experiments of the hydrogels were tested at pH 7.4, 5.0 and 3.0, respectively. Fig. 9 reveals that the release rate of the drugs from the hydrogels is influenced by the pH value.
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Fig. 7. The release behaviors of Cur (a) and Fu (b) from Gel 1 (1.0 wt%), Gel 2 (1.5%) and Gel 3 (2.0 wt%) at 37 °C.
Fig. 8. The release behaviors of Cur (a) and Fu (b) from Gel 3 at 37 °C as a function of the ratio of CS/ODex (1/1, 1/2 and 2/1).
Fig. 9. The release behaviors of Cur (a) and Fu (b) from Gel 3 at 37 °C as a function of pH (pH 7.4, pH 5.0 and pH 3.0).
The lower the pH value is, the faster the drugs were released. For Cur, about 17.66% (pH 5.0) and 18.80% (pH 3.0) have been released in 9 d, more than the amount released at pH 7.4 (about 11.87%). The similar situation occurred for Fu. Therefore, the hydrogel prefers to release drugs in an acid environment, which is highly desirable for treatment of cancer, since tumor tissues are known to be acidic [45].
Finally, we tested the release behaviors of the two drugs exposed to different concentration of DTT at 0 M, 5 mM and 10 mM, as the presence of redox-sensitive disulfide bonds on the shell of the micelle. From Fig. 10a, it is apparent that the release rate of Cur increased as the amount of DTT in the media increased. With higher DTT concentration of 5 mM and 10 mM, about 19.52% and 27.83% were released in 9 d, compared to 11.87% of the group
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Fig. 10. The release behaviors of Cur (a) and Fu (b) from Gel 3 at 37 °C as a function of the concentration of DTT (0 M, 5 mM, 10 mM).
with 0 mM DTT. DTT as a reducing agent, could damage the structure of micelle by cleaving the disulfide bonds, therefore reductive stimulus could accelerate the release of Cur. In order to study the influence of DTT to micelles, Cur release experiments from micelles with or without DTT were explored. The results (Fig. S1) showed that Cur release was faster in solution with 10 mM DTT than that without DTT, indicating that DTT influenced the drug release behavior. The release behavior of Fu (Fig. 10b) also shows redoxdependence. The fastest release rate of Fu from the hydrogel appears in the higher DTT concentration. It is reasonable because media would be acidic in the presence of DTT, which promotes the degradation of the gel. The micelle/hydrogel dual drug delivery system is highly desirable for treatment of cancer, since the concentration of glutathione in tumor cells is higher than the level in the bloodstream and healthy cells, which can cleave disulfide bonds like DTT [46]. In vitro drug delivery results reveal that this smart drug carrier has a long-term and sustained drug release behavior and has pH and redox sensitivity. This carrier is highly desirable for treatment of cancer, since tumor tissues are known to be acidic, and the concentration of glutathione in tumor cells is higher than the level in the bloodstream and healthy cells, which can cleave disulfide bonds like DTT. Moreover, the release rate of the drugs could be adjusted by changing the hydrogel concentration. Therefore, it has potential applications for clinical treatment.
3.7. Degradation behavior of the injectable micelle/hydrogel system To test whether the micelle/hydrogel system is biodegradable, the weight loss of the system was examined with incubation time at 37 °C. As shown in Fig. 11, all tested hydrogels showed biodegradability. After 10 d of incubation, the weight loss of Gel 1, Gel 2 and Gel 3 was 45.8%, 54.87% and 64.93%, respectively. The degradation of the hydrogels is attributed to the hydrolysis of Schiff base bonds. By our observations, the degradation process of the gels was from outside to inside. Firstly, the surface of the gels became rough, then the volume of the gels decreased gradually, nearly maintaining the original shape. As acidic environment is helpful for the hydrolysis of Schiff base bonds and medium containing DTT is acidic, other two hydrogels cultured at pH 3.0 or with 20 mM DTT have rapid degradation behaviors in contrast to the group which cultured at pH 7.4 without DTT. Therefore, the micelle/hydrogel will be subject degradation in vivo as modeled by the in vitro degradation.
Fig. 11. The degradation behaviors of the shell-crosslinked F127 micelle/hydrogel composites under different conditions. All the tests were carried out at 37 °C.
4. Conclusions In attempts to develop a smart dual drug delivery system to improve therapeutic outcome of combined therapy in treatment of cancer, we have prepared an injectable micelle/hydrogel composite, which consists of curcumin (Cur) loaded shell-crosslinked F127 micelle and 5-fluorouracil (Fu) dispersed CS/ODex hydrogel. Both of the shell-crosslinked F127 micelle and CS/ODex hydrogel were prepared based on Schiff base bonds, which provides pHsensitivity and biodegradability for the system. In addition, disulfide bonds were introduced in the shell of F127 micelle, allowing for redox-sensitivity of the system. Drug delivery results reveal that the drug carrier has pH and redox sensitivity. In vitro degradation experiments show that the hydrogel is able to degrade by the hydrolysis of Schiff base bonds. MTT assay shows that the hydrogels display excellent compatibility with cells. These studies indicate that the micelle/hydrogel dual drug delivery system could load and deliver hydrophobic and hydrophilic drugs simultaneously and have potential application in cancer treatment. Acknowledgments The authors gratefully acknowledge the financial support of the National Natural Science Foundation of China (Grant Nos.
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