Targeted dextran-b-poly(ε-caprolactone) micelles for cancer treatments

Abstract

In this study, targeted drug deliveries with excellent biocompatibility have been investigated to improve the efficacy and reduce the systemic toxicity of drugs. First, amphiphilic dextran-b-poly(ε-caprolactone) (Dex-PCL) copolymers were synthesized by a “click” reaction between α-alkyne terminated dextran and azido-terminated poly(ε-caprolactone). Then, the targeted molecules, such as folic acid and galactose, were conjugated to Dex-PCL. To verify their feasibility as drug delivery vehicles, DOX was loaded into Dex-PCL micelles with or without a targeted molecule. The in vitro release of DOX from DOX-loaded micelles demonstrated that there was a continuous release after burst release in the first 6 h. The cell viability assay of DOX-loaded micelles conjugated with targeting molecules against HeLa and HepG2 cells was investigated. Targeted DOX-loaded micelles showed significant bindings with tumor cells and efficient inhibition to corresponding targeted cells. Therefore, targeted DOX-loaded micelles provided an efficient drug delivery platform for the inhibition of cancer cells.

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1. Introduction

To improve the efficacy and reduce the systemic toxicity of drugs, nanoparticle-based drug delivery systems have been developed. Polymeric micelles based on amphiphilic copolymers, as one type of drug nanocarrier, have received considerable attention due to the greatly enhanced water solubility and the prolonged circulation in blood compartments.^1–4 The inner hydrophobic core of micelles provides a room for loading the hydrophobic drug, while the hydrophilic shell can enhance the stability of the micelle in an aqueous environment.^5–9 As drug carriers for cancer therapy, polymeric micelles show enhanced accumulation in tumor tissues by the enhanced permeability and retention (EPR) effect or active targeting.^10–13 Frequently, aliphatic polyesters have been chosen as hydrophobic blocks to build the amphiphilic copolymers because of their excellent biocompatibility and biodegradablility.^14,15 On the other hand, poly(ethylene glycol) (PEG) is often chosen as a hydrophilic block due to a number of advantages such as biocompatibility, non-toxicity, low immunogenicity and antigenicity.^16,17 More important, PEGylated nanoparticles can enhance blood circulation time and promote the tumor accumulation.^18,19 However, in some cases, the absence of reactive groups in the molecular chain for further modification or ligand coupling and nondegradability limited the applications of PEGylated nanoparticles.²⁰ Thus, the development of hydrophilic PEG analogs as nano-carriers of drugs with good biocompatibility, biodegradability, stealthy property and presence of abundant functional groups has received increasing attention.

Dextran, composed of anhydro-D-glucose, is a natural analog of PEG and has attracted great interest for use a as drug carriers due to its unique properties such as biodegradability and nonfouling property.^21–23 In addition, dextran can be easily chemically modified because of its 5% branching structure and plentiful hydroxyl groups on the chain.^24,25 Therefore, various targeting ligands, including RGD peptide²⁶ and galactose,²⁷ can be conjugated with dextran to build targeted drug delivery systems.

To date, only few researches about dextran-based block copolymers for cancer therapy have been reported. Sun et al. synthesized dextran-SS-poly(ε-caprolactone) diblock copolymer and used it for drug delivery. Evidently, a faster drug release was observed in the presence of reductive agents. These redox-sensitive micelles may be promising candidates as platforms for targeted intracellular drug delivery.²¹ Li et al. synthesized amphiphilic dextran-b-poly(ε-caprolactone) for an anticancer drug carrier by the end-to-end coupling of amino-terminated dextran and aldehyde-terminated poly(ε-caprolactone). The DOX release from the DOX-loaded Dex-PCL nanoparticles exhibited a sustained release manner with a certain amount of burst release in the first 9 h.²⁰

Though some anticancer drugs exhibit high efficacy with a variety of solid tumors, the delivery systems without a specific property will lead to substantial toxicity to normal tissue. It is still a challenge to deliver drugs on demand and to a specific location. In this study, targeted drug deliveries with excellent biocompatibility have been investigated. Dex-PCL diblock copolymers were first synthesized by a versatile “click” reaction between α-alkyne terminated dextran and azido-terminated poly(ε-caprolactone). Different bioactive molecules, folic acid and galactose were then conjugated to Dex-PCL copolymers to build a tumor-targeting nanocarrier. To verify the targeted release of an antitumor drug, DOX was chosen as a drug model to be loaded into Dex-PCL micelles with or without a targeted molecule. The in vitro DOX release behaviour and cell viability assays of targeted DOX-loaded Dex-PCL micelles against HeLa and HepG2 cells were investigated. These targeted DOX-loaded nano-carriers hold great potential for in vivo chemotherapeutics.

2. Materials and methods

2.1. Materials

Dextran (Dex, M_n = 6000, Sigma), propargylamine (98%, Sigma), sodium cyanoborohydride (95%, Sigma), sodium azide (Sigma), N,N,N′,N′,N′′-pentamethyldiethylenetriamine (PMDETA, 99%, Sigma) and polyethyleneimide (PEI, M_w = 25 kDa, Sigma) were used as received. 2-Bromoethanol was purified by vacuum distillation over CaH₂. ε-Caprolactone (Sigma) was purified by vacuum distillation. Folic acid (FA), galactosamine (Gal) and 3-(4,5-dimethyl-thiazol-2-yl)-2,5-diphenyl tetrazolium bromide (MTT) were purchased from Sigma-Aldrich. Doxorubicin hydrochloride (DOX·HCl) was purchased from Zhejieng Hisun Pharmaceutical Co., Ltd. Dimethyl sulfoxide (DMSO) was dried over calcium hydride (CaH₂) and purified by vacuum distillation with CaH₂. All the other reagents and solvents were purchased from Sinopharm Chemical Reagent Co., Ltd. and used as obtained.

2.2. Characterizations

¹H NMR spectra were recorded on a Bruker AV 400 NMR spectrometer in dimethyl sulfoxide-d₆ (DMSO-d₆) or deuterated chloroform (CDCl₃). FT-IR spectra were recorded on a Bio-Rad Win-IR instrument using the potassium bromide (KBr) method. Transmission electron microscopy (TEM) measurements were performed on a JEOL JEM-1011 transmission electron microscope with an accelerating voltage of 100 kV. A drop of the micelles solution (0.1 g L⁻¹) was deposited onto a 230 mesh copper grid coated with carbon and allowed to dry in air at 25 °C before measurements. Dynamic laser scattering (DLS) measurements were performed on a WyattQELS instrument with a vertically polarized He–Ne laser (DAWN EOS, Wyatt Technology). The scattering angle was fixed at 90°.

2.3. Synthesis of α-alkyne dextran (alkyne-Dex)

Dex (2.7 g, 0.454 mmol) was dissolved in 2% (w/v) acetate buffer (pH 5.0) in a flask at 50 °C. Propargylamine (2.5 g, 45.4 mmol) and sodium cyanoborohydride (2.85 g, 45.4 mmol) were added under stirring. The mixture was allowed to stir at 50 °C for 96 h. The solution was concentrated by rotavapor, and then dialyzed against deionized water for 4 days, and the product was collected by lyophilization (yield: 72%).

2.4. Synthesis of N₃-poly(ε-caprolactone) (N₃-PCL)

Typically, 2-bromoethanol (0.1 g, 0.8 mmol), ε-caprolactone (ε-CL, 5 g, 0.043 mol) and stannous octoate (0.017 g, 0.1 mol% with respect to ε-caprolactone) were added into a glass ampoule with a magnetic bar. The reaction was continued in an oil bath at 120 °C for 24 h. The product was collected by precipitation in 10-fold diethyl ether. Then, the product was re-dissolved into trichloromethane and precipitated thrice in diethyl ether. After vacuum drying for 24 h, PCL–Br was obtained with 70% yield.

Second, PCL–Br (1.2 g, 0.2 mmol) was dissolved in 10.0 mL of DMF. After heating to 80 °C, NaN₃ (0.126 g, 2 mmol) was added. After stirring the reaction mixture for 18 h at 80 °C, the reaction solution was cooled to room temperature and precipitated in 80.0 mL of diethyl ether. The crude product was re-dissolved in trichloromethane and washed at least twice with deionized water. Then, the solution was dried with MgSO₄ and precipitated in diethyl ether. After vacuum drying for 24 h, PCL-N₃ was obtained with 74% yield.

2.5. Synthesis of dextran-block-poly(ε-caprolactone) (Dex-PCL)

PCL-N₃ (0.6 g, 0.1 mmol), α-alkyne Dex (1.2 g, 0.2 mmol) and PMDETA (40 μL, 0.2 mmol) were dissolved in 30.0 mL of dried DMSO. The mixture was stirred for 10 min and degassed by three freeze–thaw cycles and transferred to another Schlenk flask, containing CuBr (30 mg, 0.2 mmol) via an N₂-purged syringe. The Schlenk flask was placed in an oil bath at 60 °C for 72 h. The reaction medium was dialyzed against deionized water (MWCO 10 kDa) for 4 days, and the product was obtained by lyophilization (yield: 78%).

2.6. Preparation of folic acid grafted Dex-PCL (FA–Dex-PCL)

Dex-PCL2 (1 g), folic acid (0.08 g, 0.17 mmol), EDC (0.07 g, 0.34 mmol) and DMAP (0.004 g, 0.034 mmol) were dissolved in 20 mL of DMSO under vigorous stirring. The reaction was performed for 72 h. After the reaction finished, the reaction medium was dialyzed against deionized water (MWCO 7 kDa) for 4 days, and the product was obtained by lyophilization (yield: 82%).

2.7. Preparation of galactose grafted Dex-PCL (Gal–Dex-PCL)

First, Dex-PCL2 (1 g) was dissolved in 20 mL of DMSO with vigorous stirring. After that, 0.05 mL of triethylamine and 0.04 g p-nitrophenyl chloroformate (NPC) was added sequentially. The reaction was performed at 20 °C for 24 h. After finishing the reaction, the reaction medium was dialyzed against deionized water (MWCO 7 kDa) for 4 days, and the product was obtained by lyophilization (yield: 79%). Second, 0.4 g NPC–Dex-PCL2 was dissolved in 20 mL of DMSO with vigorous stirring, and 0.24 g galactosamine was added. The reaction was performed at 40 °C for 48 h. After the reaction was completed, the reaction medium was dialyzed against deionized water (MWCO 7 kDa) for 4 days, and the product was obtained by lyophilization (yield: 81%).

2.8. In vitro drug loading and release

Doxorubicin (DOX) was used as a model drug for in vitro drug loading and release. DOX-loaded micelles were prepared by a simple dialysis technique. Typically, Dex-PCL2 (20.0 mg) and the drug (4.0 mg) were mixed in 2.0 mL of DMSO. The mixture was stirred at room temperature for 24 h and then added dropwise into 20.0 mL of PBS at pH 7.4. DMSO was removed by dialysis against water for 24 h. The dialysis medium was refreshed five times, and the entire procedure was performed in the dark. Then, the solution was filtered and lyophilized. Other DOX-loaded micelles were prepared similarly as DOX-loaded Dex-PCL2 micelle. To determine the drug loading content (DLC) and drug loading efficiency (DLE), the drug-loaded micelle was dissolved in DMSO and analyzed by fluorescence measurement (Perkin-Elmer LS50B luminescence spectrometer) using a standard curve method (λ_ex = 480 nm). The DLC and DLE of drug-loaded micelles were calculated according to eqn (1) and (2), respectively:

DLC (wt%) = amount of drug in micelle/amount of drug loaded micelle × 100 (1)

DLE (wt%) = amount of drug in micelle/total amount of feeding drug × 100 (2)

The in vitro drug release profiles of drug-loaded micelles were investigated in PBS at pH 7.4. The pre-weighed freeze-dried DOX-loaded micelles were suspended in 5.0 mL of release medium and transferred into a dialysis bag (MWCO 3500 Da). The release experiment was initiated by placing the end-sealed dialysis bag into 50.0 mL of release medium at 37 °C with continuous shaking at 70 rpm. At predetermined time intervals, 2.0 mL of external release medium was taken out and an equal volume of fresh release medium was replenished. The amount of released DOX was determined using fluorescence measurement (λ_ex = 480 nm). The release experiments were conducted in triplicate.

2.9. Intracellular drug release

The cellular uptake and intracellular release behaviors of DOX-loaded micelles were observed by confocal laser scanning microscopy (CLSM) and flow cytometric analyses toward HeLa and HepG2 cells.

2.9.1. CLSM. HeLa, HepG2 or L929 cells were placed into 6-well plates (2 × 10⁵ cells per well) and cultured in 2.0 mL of RPMI1640. After incubation for 24 h, the culture media were withdrawn and culture media containing DOX-loaded micelles were supplemented to confirm that the final DOX concentration is 10.0 mg L⁻¹. The cells were incubated for another 4 h and washed with PBS thrice. The cells were then fixed in 4% paraformaldehyde for 30 min and washed with PBS thrice. For staining the nuclei, the cells were incubated with 4′,6-diamidino-2-phenylindole (DAPI, blue) for 20 min. The images of cells were observed using a laser scanning confocal microscope.

2.9.2. Flow cytometric analyses. HeLa, HepG2 or L929 cells were seeded in 6-well plates at 2 × 10⁵ cells per well in 2.0 mL of RPMI1640 and cultured for 24 h. The cells were then washed by PBS and incubated at 37 °C for an additional 2 h with DOX-loaded micelles at a final DOX concentration of 5.0 mg L⁻¹ in complete DMEM. Thereafter, the culture medium was removed and the cells were washed with PBS thrice and treated with trypsin. Then, 1.0 mL of PBS was added to each culture well, and the solutions were centrifuged for 4 min at 3000 rpm. After the removal of supernatants, the cells were re-suspended in 0.3 mL of PBS. Data for 1 × 10⁴ gated events were collected, and analysis was performed by a flow cytometer (Beckman, California, USA).

2.10. Cell viability assays

The cytotoxicities of DOX-free micelles against HeLa and HepG2 cells were tested using a standard MTT assay. The cells were placed in 96-well plates (1 × 10⁴ cells per well) in 200 μL of RPMI1640, followed by incubation at 37 °C for 24 h. The culture medium was then withdrawn. Micelles solutions with different concentrations (0–10.0 g L⁻¹) in RPMI1640 were added. The MTT assay was performed after incubation for another 72 h. The absorbance at 490 nm was measured. Cell viability (%) was calculated by the following eqn (3):

Cell viability (%) = A_sample/A_control × 100 (3)

where A_sample and A_control are the absorbance of the sample and control wells, respectively.

The cytotoxicities of DOX-loaded micelles against HeLa, HepG2 or L929 cells were also measured by using a standard MTT assay. The cells were placed in 96-well plates (1 × 10⁴ cells per well) in 200 μL of RPMI1640, followed by incubation at 37 °C for 24 h. After washing with PBS, 180 μL of RPMI1640 and 20 μL of DOX-loaded micelle solutions in PBS were added to form culture media with different DOX concentrations (0–10.0 mg L⁻¹ DOX). The MTT assay was conducted after incubation for 24, 48 and 72 h. The absorbance was tested at 490 nm. Cell viability (%) was also calculated using eqn (3).

3. Results and discussion

3.1. Synthesis of Dex-PCL copolymer

As depicted in our previous reports, the chemical structure of alkyne-terminated dextran can be confirmed by ¹H NMR and FT-IR.^23,28 Azido-terminated PCL (N₃-PCL) was synthesized through the reaction between bromide-modified PCL (Br-PCL) and NaN₃. The chemical structure of azido-terminated PCL (N₃-PCL) was confirmed by ¹H NMR. As shown in Fig. S1,† the complete disappearance of peaks at 3.4 and 4.4 ppm, assigned to the methylene protons close to bromine (a, Br–CH₂CH₂– and b, Br–CH₂CH₂), strongly indicated the successful synthesis of N₃-PCL. Moreover, the molecular weights of PCL blocks were determined by ¹H NMR and GPC. According to the integration ratio between the proton of methylene of initiator at 3.4 ppm (b, Br–CH₂CH₂–) and the proton of ε-CL unit appearing at 1.3 ppm (f, –CH₂CH₂CH₂–), the molecular weight of PCL–Br could be calculated. As listed in Table S1,† the molecular weights of PCL–Br from ¹H NMR were coincident with that from the GPC measurement.

Dex-PCL copolymers were prepared by combining the clickable alkyne-Dex with N₃-PCL via a click reaction as depicted in Scheme 1. ¹HNMR and FT-IR confirmed the structure of Dex-PCL copolymers. The disappearance of the azide peak of N₃-PCL and α-alkyne peak of alkyne-Dex at about 2100 cm⁻¹ demonstrated that N₃-PCL had been completely consumed, suggesting the successful synthesis of Dex-PCL copolymers (Fig. S2†). Moreover, all the peaks assigned to Dex and PCL further confirmed the structure of Dex-PCL as shown in the ¹H NMR spectra (Fig. S3–S5†).

Scheme 1 Synthetic routes of Dex-PCL.

3.2. Synthesis of targeted Dex-PCL copolymer

The structure of FA–Dex-PCL2 was confirmed by ¹H NMR analysis. The peaks that appeared from 6.6 to 8.5 ppm were assigned to FA, indicating the successful synthesis of FA–Dex-PCL2 (Fig. S6†). Because the hydroxyl groups of Gal and dextran appear at the same peaks in the ¹H NMR spectrum, UV-vis assay was explored to confirm the structure of Gal–Dex-PCL. From the standard curve (y = 0.97x + 0.02, y represents absorbance values and x represents concentration of NPC–Dex-PCL2) absorbance at 280 nm, as shown in Fig. S7A,† we know that the modified rate of NPC in NPC–Dex-PCL2 was about 4% because the absorbance of NPC groups in NPC–Dex-PCL2 solution was 0.51, as shown in Fig. S7B.† Also, the raw material of NPC–Dex-PCL2 was 15.6 mg, and the disappearance of the characteristic peak of the benzene group at about 280 nm strongly demonstrated that Gal molecules substituted the NPC molecules, which means the successful synthesis of Gal–Dex-PCL2.

3.3. Micellization of block copolymers

It is well known that the amphiphilic block copolymers can self-assemble into micelles in selected aqueous media. Herein, a fluorescence method using pyrene as a probe has been explored to investigate the CMC value and further demonstrate the formation of micelles and the influence of the composition of block copolymers on the properties of micelles. The excitation spectra of pyrene with increased concentration of block copolymers were measured to confirm the self-assembly of the block copolymers. Typically, a red shift of the absorption band was observed with an increase in the concentration of the copolymer. The red shift indicated the formation of the micelles, which can be attributed to the transfer of the pyrene molecule from a water environment to a hydrophobic micellar core. Moreover, the CMC values were collected by the plot of fluorescence intensity ratio of I₃₄₂/I₃₃₅ versus log 10c of the copolymer. As listed in Table 1, the molecular weight of the PCL block had obvious influence on the CMC values due to the change of hydrophobicity. For an amphiphilic polymer, the ratio of hydrophobic parts decisively affect the CMC value. The higher ratio of the hydrophobic part results in a lower CMC value. The introduction of hydrophobic FA will improve the hydrophobic property of FA–Dex-PCL2, while for Gal, it is hydrophilic. Therefore, the CMC value of Gal–Dex-PCL2 is higher than that of Dex-PCL2, and the CMC value of FA–Dex-PCL2 is smaller. At the same time, with the increase in molecular weight of PCL, the CMC value becomes lower, indicating the increased hydrophobicity of the block copolymer.

Table 1 Characterizations of the micelles

Micelles	CMC^a (μg mL^−1)	Rh^a (nm)	DLC (wt%)	DLE (wt%)
a Determined at pH 7.4.
Dex-PCL1 (Dex-PCL2000)	60.32	70.9 ± 12.2	3.97	23.82
Dex-PCL2 (Dex-PCL4000)	2.81	112.3 ± 15.7	5.81	34.86
Dex-PCL3(Dex-PCL6000)	0.82	183.2 ± 21.8	7.98	47.88
FA–Dex-PCL2 (FA–Dex-PCL4000)	1.94	128.3 ± 18.8	6.44	38.64
Gal–Dex-PCL2 (Gal–Dex-PCL4000)	4.12	105.7 ± 14.9	4.97	29.82

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The hydrodynamic radii (R_h) of the amphiphilic aggregates were measured by DLS. With the increase of the PCL block length, the R_h values of Dex-PCL micelles increased gradually from 70.9 ± 12.2 to 183.2 ± 21.8 nm, as shown in Table 1. It should be attributed to the increase of hydrophobicity of the core, resulting from the increase of the molecular weight of the PCL block. The same results could also be certificated by transmission electron microscopy (TEM, Fig. S8†). The smaller values from TEM observations as compared with DLS should be due to the dehydration of the micelles in the process of TEM sample preparation.

3.4. In vitro DOX loading and release

Doxorubicin (DOX), a widely used antineoplastic drug, was chosen as the model drug to verify the feasibility of using the micelles for intracellular drug delivery in cancer chemotherapy. As shown in Table 1, the DLC of Dex-PCL micelles were in the range of 4.0–8.0% and the DLE were 23.8–47.9%. These data indicated that the drug loading capacity of Dex-PCL micelles would be improved by increasing the hydrophobicity of the core. The cumulative release percentages of DOX-loaded Dex-PCL micelles with different proportions versus time are plotted in Fig. 1. As depicted, a fast DOX release rate of DOX-loaded Dex-PCL micelles was observed in the first 6 h. After this stage, for all DOX-loaded Dex-PCL micelles the release of loaded-DOX from the micelles became slow and continuous in 72 h. In addition, it was found that the longer the PCL block existed in Dex-PCL, the slower the drug release rate was obtained. These results suggested that the strong hydrophobic interaction between a long PCL block and DOX could effectively hinder the release of the encapsulated drug in normal physiological conditions.

Fig. 1 In vitro DOX release from DOX-loaded Dex-PCL1, Dex-PCL2 and Dex-PCL3 in PBS at 37 °C at pH 7.4.

The influence of the targeting molecules on the DOX release rate was also investigated. The DLC and DLE of targeted Dex-PCL micelles are listed in Table 1. As shown in Fig. 2, the release rate of DOX-loaded micelles followed this order: Gal–Dex-PCL2 > Dex-PCL2 > FA–Dex-PCL2. The reason for this result is because the conjugation of the micelles with different targeting molecules can change the hydrophobicity of the micelles. In addition, the different hydrophilic materials hinged on the physical properties of micelles. As is well known, FA molecules are hydrophobic materials; the introduction of them will increase the hydrophobic parts of copolymers, which will bind more hydrophobic DOX. According to the similar principle, Gal molecules are hydrophilic materials, which will promote the copolymer to be more hydrophilic and thus bind less hydrophobic DOX.

Fig. 2 In vitro DOX release from DOX-loaded Dex-PCL2, FA–Dex-PCL2 and Gal–Dex-PCL2 in PBS at 37 °C at pH 7.4.

3.5. Cellular proliferation inhibition

To evaluate the cytotoxicity of DOX-loaded Dex-PCL micelles, in vitro cytotoxicity tests of DOX-loaded micelles and free DOX against HepG2 cells were conducted. The viability of HepG2 cells after incubation with free DOX, DOX-loaded Dex-PCL1, Dex-PCL2, and Dex-PCL3 for 24 h, 48 h, and 72 h at different concentrations are revealed in Fig. 3. As depicted, DOX-loaded Dex-PCL1 exhibited a higher cytotoxicity than DOX-loaded Dex-PCL2 and Dex-PCL3, which should be attributed to the slightly higher drug release speed of DOX-loaded Dex-PCL1.

Fig. 3 Cytotoxicities of DOX-loaded Dex-PCL3, Dex-PCL2 and Dex-PCL1 as well as free DOX towards HepG2 cells after incubation for 24, 48 and 72 h.

In addition, the potential toxicity of polymeric materials should be evaluated for drug delivery applications. The in vitro cytotoxicity of the Dex-PCL2 to HepG2 cells was evaluated using a MTT assay. PEI was selected as the positive control. As demonstrated in Fig. S9,† the viabilities of HepG2 cells treated with Dex-PCL2 for 72 h were over 90% at all test concentrations. These results suggested that Dex-PCL had low cytotoxicity and may be suitable candidates as carriers for drug delivery.

The targeted antitumor activity of DOX-loaded targeted Dex-PCL micelles was also investigated with HeLa and HepG2 cells using MTT assays. The cytotoxicities of DOX-loaded FA–Dex-PCL2, DOX-loaded Gal–Dex-PCL2, DOX-loaded Dex-PCL2, DOX-loaded FA–Dex-PCL2 with free FA and free DOX toward HeLa cells after incubation for 24 h, 48 h and 72 h were investigated. As depicted in Fig. 4, DOX-loaded FA–Dex-PCL2 micelles induced the higher cytotoxicity toward HeLa cells, a folate receptor positive tumor cell line, than that of DOX-loaded Dex-PCL2. In contrast, the cytotoxicity of DOX-loaded FA–Dex-PCL2 in the presence of 1 μg mL⁻¹ of free folic acid was obviously reduced because the folate receptors of HeLa cells were blocked by free FA. This phenomenon clearly indicated that the cytotoxicity of DOX-loaded FA–Dex-PCL2 against HeLa cells was caused mainly by the cellular internalization of the nanoparticles through folate-mediated targeting. Furthermore, DOX-loaded Gal–Dex-PCL2 and Dex-PCL2 micelles showed lower cytotoxicities than DOX-loaded FA–Dex-PCL2 micelles towards HeLa cells, which further confirmed the folate-mediated cellular uptake of the latter. Similarly, a receptor-enhanced cytotoxicity was also observed toward HepG2 cells, a galactose-receptor over-expressed cell line. As shown in Fig. S10,† DOX-loaded Gal–Dex-PCL2 had the highest cytotoxicity against HepG2 cells without free Gal. However, DOX-loaded Gal–Dex-PCL2 micelles showed a reduced cytotoxicity towards HepG2 cells in the presence of free galactose due to the suppression of the galactose receptor by the free galactose.

Fig. 4 Cytotoxicities of DOX-loaded Dex-PCL2, FA–Dex-PCL2, Gal–Dex-PCL2, FA–Dex-PCL2 with free FA and free DOX towards HeLa cells after incubation for 24, 48 and 72 h.

In addition, to further study DOX-loaded target micelles against cells effect, normal cells and L929 were used for comparison. Also, as is well known, either FA or Gal molecules have no specific interaction towards L929 cells. As shown in Fig. 5, DOX-loaded FA–Dex-PCL2, DOX-loaded Gal–Dex-PCL and DOX-loaded Dex-PCL showed nearly the same killing effects towards L929 cells. This result could be explained that due to lack of specific target molecules towards normal cells, all tested groups took the same status for treating L929 cells. Moreover, this result also confirmed that target molecules indeed could improve efficiency against tumor cells compared with Fig. 4.

Fig. 5 Cytotoxicities of DOX-loaded Dex-PCL2, FA–Dex-PCL2, Gal–Dex-PCL2, and free DOX towards L929 cells after incubation for 24, 48 and 72 h.

3.6. Cell uptake efficiency

The cell uptake efficiency of DOX-loaded targeted Dex-PCL2 micelles was investigated by confocal laser scanning microscopy (CLSM) assays toward HeLa and HepG2 cells. As shown in Fig. 6, a stronger intracellular DOX fluorescence was observed in HeLa cells after incubation with DOX-loaded FA–Dex-PCL2 micelles for 4 h. A weaker DOX fluorescence intensity in the nuclei was observed with the presence of free FA, which was a competing molecule and inhibited the receptor-mediated uptake of DOX-loaded FA–Dex-PCL2 micelles in HeLa cells. Similar results were also observed in HepG2 cells incubated with the DOX-loaded galactose-functionalized Dex-PCL2 micelles (Gal–Dex-PCL2). The fluorescence intensity of DOX in HepG2 cells incubated with DOX-loaded Gal–Dex-PCL2 micelles was higher than those incubated with DOX-loaded Dex-PCL2 micelles, but it was obviously reduced in the presence of free galactose (Fig. S11†).

Fig. 6 Representative CLSM images of HeLa cells incubated with DOX-loaded FA–Dex-PCL2 (A), DOX-loaded Dex-PCL2 (B) and DOX-loaded FA–Dex-PCL2 with free FA (C) for 4 h. For each panel, the images from left to right show differential interference contrast (DIC) image, DOX fluorescence in cells (red), cell nuclei stained by DAPI (blue), and overlays of the three images. The scale bar is 20 μm.

In addition, we carried out another CLSM assay to confirm the introduction of target molecules efficiency, as shown in Fig. 7; moreover, we could easily find that they displayed nearly the same fluorescent intensities. The result could be attributable to lack of specific interactions between DOX-loaded deliveries and L929 cells, which showed the same killing cells effect among all the tested groups.

Fig. 7 Representative CLSM images of L929 cells incubated with DOX-loaded FA–Dex-PCL2 (A), DOX-loaded Dex-PCL2 (B) and DOX-loaded Gal–Dex-PCL2 (C) for 4 h. For each panel, the images from left to right show differential interference contrast (DIC) image, DOX fluorescence in cells (red), cell nuclei stained by DAPI (blue), and overlays of the three images. The scale bar is 20 μm.

Flow cytometric analyses were also used to observe the cell uptake efficiency of DOX-loaded targeted Dex-PCL2 micelles. As shown in Fig. 8, the flow cytometric histogram for HeLa cells incubated with DOX-loaded FA–Dex-PCL2 obviously shifted to the highest fluorescence intensity region. The fluorescence intensity in HeLa cells decreased following this order: DOX-loaded FA–Dex-PCL2 > DOX-loaded FA–Dex-PCL2 with free FA > DOX-loaded Gal–Dex-PCL2 > DOX-loaded Dex-PCL2. A similar targeting molecule-enhanced fluorescence was observed in HepG2 cells. The fluorescence intensities of DOX in HepG2 cells were in the following order: DOX-loaded Gal–Dex-PCL2 > DOX-loaded Gal–Dex-PCL2 with free Gal > DOX-loaded FA–Dex-PCL2 > DOX-loaded Dex-PCL2 (Fig. S12†). In contrast, we observed that there was nearly no difference of cells uptake for L929 cells in terms of DOX-loaded Dex-PCL2 (b), DOX-loaded Gal–Dex-PCL2 (c), and DOX-loaded FA–Dex-PCL2 (d), as shown in Fig. S13.† Because of the lack of a special guide, regarded as target molecules, towards L929 cells, they displayed the same killing efficiency. These results demonstrated that the DOX-loaded micelles with targeting molecules showed significant bindings with tumor cells and efficient inhibition to targeted cells.

Fig. 8 Flow cytometric profiles of HeLa cells incubated with DOX-loaded micelles for 2 h. DOX-loaded Dex-PCL2 (a), DOX-loaded Gal–Dex-PCL2 (b), DOX-loaded FA–Dex-PCL2 (c), DOX-loaded FA–Dex-PCL2 with free FA (d) and free DOX (e).

4. Conclusions

In summary, the amphiphilic Dex-PCL copolymers were synthesized via a “click” reaction of α-alkyne terminated dextran and azido-terminated poly(ε-caprolactone). Then, the targeted molecules, folic acid and galactose, were conjugated to the Dex-PCL copolymers. The in vitro release of DOX from DOX-loaded Dex-PCL micelles showed a continuous release behavior after an initial burst release. Cellular proliferation inhibition and cell uptake efficiency of targeted DOX-loaded Dex-PCL micelles against HeLa and HepG2 cells depicted an enhanced and fast endocytosis due to the specific interaction between targeted DOX-loaded Dex-PCL micelles with corresponding tumor cells. Therefore, the targeted DOX-loaded nano-based carrier provides a platform to build an efficient drug delivery system for cancer therapy.

Acknowledgements

This research was financially supported by the National Natural Science Foundation of China (Projects 21474012, 51273037, 50903012, and 21174142), the Jilin Science and Technology Bureau (20130206074GX, International Cooperation Project 20120729), the New Century Excellent Talents in University of Jilin Province (2013-6, Jilin Provincial Education Department), the Jilin Human Resources and Social Security Bureau (201125020), and the Jilin Environmental Protection Bureau (201127).

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Footnote

† Electronic supplementary information (ESI) available. See DOI: 10.1039/c4ra15696j

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