Dual-Sensitive Hydrogel Nanoparticles Based on Conjugated

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Cite This: ACS Appl. Mater. Interfaces 2018, 10, 19336−19346

Dual-Sensitive Hydrogel Nanoparticles Based on Conjugated Thermoresponsive Copolymers and Protein Filaments for Triggerable Drug Delivery Roujin Ghaffari,† Niloofar Eslahi,*,§ Elnaz Tamjid,∥ and Abdolreza Simchi*,†,‡

ACS Appl. Mater. Interfaces 2018.10:19336-19346. Downloaded from pubs.acs.org by OPEN UNIV OF HONG KONG on 01/25/19. For personal use only.

†

Department of Materials Science and Engineering and ‡Institute for Nanoscience and Nanotechnology, Sharif University of Technology, Azadi Avenue, P.O. Box 11365/8639, Tehran 14588-89694, Iran § Department of Textile Engineering, Science and Research Branch, Islamic Azad University, P.O. Box 14515/775, Tehran 1477893855, Iran ∥ Department of Nanobiotechnology, Faculty of Biological Sciences, Tarbiat Modares University, Tehran 1411713116, Iran ABSTRACT: In this study, novel hydrogel nanoparticles with dual triggerable release properties based on fibrous structural proteins (keratin) and thermoresponsive copolymers (Pluronic) are introduced. Nanoparticles were used for curcumin delivery as effective and safe anticancer agents, the hydrophobicity of which has limited their clinical applications. A drug was loaded into hydrogel nanoparticles by a single-step nanoprecipitation method. The drug-loaded nanoparticles had an average diameter of 165 and 66 nm at 25 and 37 °C, respectively. It was shown that the drug loading efficiency could be enhanced through crosslinking of the disulfide bonds in keratin. Crosslinking provided a targeted release profile under reductive conditions using an in vivo agent, glutathione (GSH), or in the presence of trypsin. Cytocompatibility assay using HeLa and L929 fibroblast cells exhibited no adverse effect of nanoparticles on cell viability up to 1 mg/mL. Besides, the green fluorescence of curcumin confirmed the uptake of drug-loaded nanoparticles by cancer cells. The redox and temperature-sensitive nanoparticles are potentially useable for the efficient delivery of hydrophobic drugs to targeted regions having a triggerable release profile. KEYWORDS: drug delivery, hydrogel nanoparticles, stimuli-responsive copolymer, keratin biopolymer, Pluronic conjugates

1. INTRODUCTION In the past decade, advancement in smart multitargeting drug delivery systems using nanotechnology has gained considerable attraction due to their more efficient drug release.1 Originally, polymers were used to stabilize or solubilize drugs for controlled release, however, progress in new polymer designs with the development of synthetic strategies has made it possible to formulate various polymeric drug delivery systems.2 Controlled drug delivery systems are designed to deliver drugs at predetermined rates for desirable times or specific sites to overcome the shortcomings of conventional drug formulations.3,4 Nanogels or hydrogel nanoparticles are three-dimensional networks of crosslinked polymer chains. These nanocarriers are very promising systems for target-specific delivery of drugs, because they have high drug loading (DL) capacity and enhanced cellular uptake efficiency5−7 with the ability to protect the drug and postpone its degradation.8,9 Recent advances deal with the development of stimuli-responsive polymeric hydrogels.10,11 For this aim, thermoresponsivetriblock copolymers such as Pluronic (PEO−PPO−PEO) has extensively been studied, especially in drug delivery due to their © 2018 American Chemical Society

approved use in medical devices by the US Food and Drug Administration.12 The self-assembly of amphiphilic copolymers through hydrophobic interactions of poly(propylene oxide) (PPO) blocks in the inner core and hydrophilic interactions of poly(ethylene oxide) (PEO) blocks in the outer shell forms water soluble micelles above the lower critical solution temperature.13 Hydrophobic drugs can easily be loaded inside the micelles to improve their therapeutic activity, metabolic stability, and circulation time.14 A good example is loading curcumin, as a hydrophobic drug, in the polymeric micelle to be transferred into the tumor cells.15 Although the drug has a wide range of pharmacological effects such as anti-inflammatory, antioxidant, antimutagenic, and antitumor properties,16 its application in treatments of many diseases is confined due to low solubility in aqueous solutions and rapid degradation at physiological pH.15 Therefore, encapsulation of the drug in the nanocarriers is an effective method for efficient anticancer treatment, owning to the presence of large hydrophobic blocks Received: January 24, 2018 Accepted: May 17, 2018 Published: May 17, 2018 19336

DOI: 10.1021/acsami.8b01154 ACS Appl. Mater. Interfaces 2018, 10, 19336−19346

Research Article

ACS Applied Materials & Interfaces in Pluronic, as a cargo space.13,17 For instance, Sahu et al.16 showed that the encapsulation yield of curcumin-loaded Pluronic micelles can be tuned with the drug to polymer ratio. Nevertheless, the copolymers suffer from dissociation upon dilution, owing to the relatively high critical micelle concentration (CMC).18 Therefore, to stabilize drugs within the nanocarrier and obtain more controllable release, chemical crosslinking or conjugation with other polymers has been examined.19 For instance, a redox-responsive drug delivery system using heparin−Pluronic conjugates showed a more controllable release profile with higher efficiency.20 On the basis of this concept, Nahain et el.21 have recently prepared a dualsensitive chemically crosslinked polymer having a pHresponsive covalent benzoic−imine bond and redox-sensitive disulfide to deliver the anticancer drug Taxol. In this study, we introduce a new type of dual-sensitive hydrogel nanoparticles for sustained and triggered drug delivery. The nanoparticles are temperature and redox sensitive and are prepared by chemical grafting of triblock copolymers (Pluronic) with protein filaments (keratin). Compared with synthetic polymers, natural polymers with thiol groups, such as keratin, are more suitable for biomedical applications. Keratin is a natural protein which can be extracted from wool, hair, nail, and feathers.22 Its ability to self-assemble into various physical shapes, excellent biological compatibility, biodegradability, and low toxicity toward cells has made keratin a potential candidate for biomedical applications, particularly as carriers for targeted drug delivery.23 Moreover, keratin is rich in lysine and arginine which are cleavable in vivo by trypsin, an essential protease found in the body that is generally overexpressed in inflamed and tumorous tissues.19 Thiol groups can be used as disulfide linkages to facilitate the selective release of drugs in reductive environments. For instance, a redox-sensitive drug delivery system based on keratin-g-poly(ethylene glycol) loaded with doxorubicin exhibited a controlled release profile with better efficiency.19 Therefore, we utilized keratin to prepare hydrogel nanoparticles aiming to attain a more controllable drug release. For this purpose, keratin as a redox-sensitive polymer was conjugated with Pluronic as a temperature-sensitive copolymer to produce novel dual-sensitive nanoparticles. The conjugated nanoparticles were used for the encapsulation of an antiinflammatory and antitumor drug (curcumin). It is shown that a high drug loading efficiency with a controllable release profile under reductive conditions and in the presence of trypsin is obtained. The physicochemical properties and cytotoxicity of the nanohydrogel are also demonstrated.

2.2. Preparation of Polymeric Nanoparticles. To prepare the nanoparticles, carboxylation of the triblock copolymer (3 g) was accomplished in dioxane (15 mL) in the presence of succinic anhydride (62.5 mg), DMAP (65 mg), and TEA (75 μg) at room temperature for 24 h. The solvent was then removed in a rotary evaporator (IKA, Germany) and the product was filtered in cold diethyl ether and dried in a vacuum (200 torr) for 12 h. Afterwards, the carboxylated copolymer (3 g) was activated by EDC (0.22 g) and NHS (0.14 g) in 18 mL phosphate buffer solution (PBS) for 1 h. Separately, 0.3 g keratin was dissolved in 12 mL PBS at pH = 5.8. This solution was then added to the carboxylated copolymer drop-by-drop for 24 h on a magnetic stirrer (Heidolph, Germany). The product was dialyzed against distilled water using a dialysis tubing cellulose membrane (MW cut-off 12 kDa) and finally lyophilized to obtain Pluronic−keratin conjugates (grafting ratio: 72%). 2.3. Drug Loading. Curcumin was encapsulated into the hydrogel nanoparticles at different ratios through a single-step nanoprecipitation method.25 For a sample run, 100 mg of the conjugated copolymer and 10 mg of curcumin were dissolved in 5 mL acetone. The solution was added to 5 mL deionized (DI) water (Millipore, 18 MΩ) drop-bydrop while stirring. During this process, the copolymer was selfassembled into nanoparticles with core-encapsulated curcumin. For crosslinking by disulfide bonds through the oxidation of the thiol groups of keratin, 1 wt % H2O2 was added to the final solution and stirred for 3 h. The solution was then dialyzed in DI water for 24 h to separate the unencapsulated drug. The resulting solution was freezedried and kept in a dry place for further characterization. To determine the efficiency of encapsulation (EE) and drug loading (DL), samples with various drug to polymer ratios were analyzed by UV−vis spectrometry (LAMBADA 35 Spectrophotometer, PerkinElmer), and the efficiency and loading percentages were calculated by25

EE =

weight of the drug in micelles × 100 weight of the feeding drug

(1)

DL =

weight of drug in micelles × 100 weight of the feeding polymer and drug

(2)

2.4. Physicochemical Characterization of Nanoparticles. Fourier transform infrared (FTIR) spectroscopy (ABB Bomem MB100) was employed in transmission mode using KBr pellets in the wavenumber range of 4000−400 cm−1 at a resolution of 4 cm−1. 1 H NMR spectra were obtained on an NMR 500 MHz spectrometer (Bruker, Germany) at room temperature. Deuterated water (D2O) was used as the solvent and the chemical shifts of the copolymers were measured in parts per million (ppm) using D2O as the internal reference. The critical micelle concentration (CMC) of the conjugated copolymer was analyzed by fluorescence spectroscopy (Varian Cary Eclipse, Agilent) using pyrene as a probe.15 Aliquots of a pyrene solution in acetone (5 × 10−7 M) were added into different copolymer concentrations (0.01−2.5 g/L). All samples were excited at 335 nm and the fluorescence spectra were recorded between 350 and 450 nm. The critical concentration was determined through the intersection of the tangent at the inflection with the horizontal tangent through the points at low concentrations.26 The size distribution of the nanoparticles was examined by dynamic light scattering (DLS, ZEN3600, Malvern, U.K.) at two temperatures of 25 and 37 °C. The ζ-potential of the drug-loaded nanoparticles was also examined by the same instrument. Transmission electron microscopy (TEM) was carried out using a CM120, Philips TEM (Netherlands), at an acceleration voltage of 100 kV. The specimens were prepared by dropping a small drop of the copolymer aqueous solution on a copper grid and air-dried prior to visualization. The stability of the drug-loaded nanoparticles was examined in PBS at 4 and 37 °C for 30 days. After the respective storage periods, the samples were centrifuged for 10 min (at 6000 rpm) and examined not only qualitatively with the naked eye, but also quantitatively by UV− vis spectroscopy. The presence of precipitations in the solutions

2. MATERIALS AND METHODS 2.1. Materials. Pluronic 127 (PEO99−PPO65−PEO99), with a molecular weight of 12.5 kDa, was purchased from Sigma-Aldrich. Succinic anhydride, 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC), N-hydroxysuccinimide (NHS), 1,4-dioxane, triethylamine (TEA), 4-dimethylaminopyridine (DMAP), hydrogen peroxide, glutathione (GSH), urea, sodium dodecyl sulfate (SDS), and curcumin were obtained from Merck (Germany). Tween-80 was purchased from Kimiya-Pakhsh (Iran). Trypsin solution was obtained from Pasteur Institute (Iran). Keratin was extracted from wool fibers on the basis of our previous study.24 In brief, cleaned, defatted wool fibers were mixed with aqueous solution containing 8 M urea, 0.5 M sodium pyrosulfite, and 0.05 M SDS at 65 °C and stirred for 24 h. After filtration and dialysis, the extracted solution was lyophilized for 48 h by a freezedryer (Lyotrap/Plus, U.K.) at −40 °C to obtain keratin powder (molecular weight (MW): 45−65 kDa). 19337


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ACS Applied Materials & Interfaces

Figure 1. (a) FTIR study shows carboxylation of the block copolymer by succinic anhydride and its interactions with the peptide bonds of protein filaments through EDC/NHS chemistry. Drug loading does not change the characteristic bands of the conjugated copolymer. (b) 1H NMR spectra of the carboxylated and conjugated copolymer. 2.7. Statistical Analysis. All data were reported as mean ± standard deviation, and each experiment was performed in triplicate. Multifactorial one-way analysis of variance was performed for the comparison of groups, and p value ≤0.05 was considered to be statistically significant.

indicates instability of the nanoparticles, while a uniformly transparent solution indicates their stability. 2.5. In Vitro Release Study. To determine the concentrationdependent kinetics of drug release in vitro, the drug-encapsulated nanoparticles (2 mg/mL) were loaded into a dialysis bag and immersed in 200 mL phosphate buffer solution (PBS, pH 7.4) containing Tween-80 (0.5% w/w) at 37 °C with constant stirring. At certain time intervals, the incubation medium (3 mL) was withdrawn and replaced with fresh PBS. The amount of released drug in the incubation medium was then quantified by UV−vis spectroscopy. Since keratin is rich in lysine and arginine, it is cleaved by trypsin protease in the body. Therefore, drug release in the presence of trypsin (0.04 M) was also studied using a similar procedure. The redox sensitivity of the synthesized nanoparticles was evaluated by employing GSH (10 mM) as a reducing agent, corresponding to GSH concentration in the cytoplasm.19 The mechanism of drug release was analyzed by curve fitting of the experimental results using kinetic models.20,27 2.6. In Vitro Cell Viability and Cellular Uptake Assay. The possible cytotoxicity of free curcumin and drug-loaded nanoparticles was evaluated using the 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazoliumbromide (MTT) method. Human epithelial carcinoma cells (HeLa cells) and L929 fibroblast cells were seeded in 96-well plates (1000 cells each). The cells were cultured in Roswell Park Memorial Institute-1640 medium containing 10% (v/v) fetal bovine serum, which was supplemented with 40−50 U/mL penicillin and 50 U/mL streptomycin. The cells were kept for 24 h at 37 °C in a 5% CO2 humidified sterile incubator. Afterwards, the medium was changed and a fresh cell culture medium containing the materials of study (blank cells, drug-loaded nanoparticles, pristine drug, and the conjugated copolymer with different concentrations (0.01, 0.1, 1 mg/mL)) was replaced. After incubation for 1 or 3 days, 20 μL of 5 mg/mL MTT solution was added to each well. After 4 h, the MTT solution was replaced with 150 μL dimethyl sulfoxide and kept under shaking for 10 min. A microplate reader was used to measure the optical density of each well at 490 nm. To visualize the cellular uptake of drug-encapsulated nanoparticles, HeLa cell line was seeded into 12-well plates and incubated overnight. Then, the growth medium was replaced with curcumin-loaded nanoparticles (0.1 mg/mL) and incubated for 24 and 48 h. The delivery of curcumin into HeLa cells was visualized using a fluorescence microscope (Eclipse TE2000-S, Nikon Instruments Inc.). In addition, the nuclei were stained with 4′,6-diamidino-2phenylindole (DAPI) and the cells were imaged after 48 h of incubation using a Leica TCS SPE confocal laser scanning microscope (Germany) with an excitation at 488 nm for curcumin.

3. RESULTS AND DISCUSSION 3.1. Conjugation of the Block Copolymers with Protein Filaments. Figure 1a shows FTIR spectrum of the conjugated copolymer in comparison with the pristine block copolymer (Pluronic) and extracted keratin. The absorbing bands at 1108 and 1243 cm−1 are ascribed to the C−O−C stretching of an aliphatic ether group and twisting vibration of −CH2 in Pluronic, respectively (Figure 1a,D). A new peak appeared at 1730 cm−1, after carboxylation of the copolymer by means of succinic anhydride (Figure 1a,E), which is assigned to the carboxyl (COOH) stretching vibration.24 The spectrum of keratin shows characteristic absorption bands assigned mainly to the peptide bonds, known as amide bonds (Figure 1a,A). The amide I band is connected mainly with the CO stretching vibration and it occurs in the range of 1700−1600 cm−1, while the amide II band falls at 1534 cm−1 which is related to N−H bending and C−H stretching vibration. The amide III band at 1237 cm−1 results from the in-phase combination of C−N stretching and N−H in-plane bending.28 The broad absorption band region from 3600 to 3200 cm−1 is attributed to the stretching vibration of N−H and O−H bonds. Peaks that fall in the 3000−2800 cm−1 range are related to C− H stretching modes. After conjugation, a peak at around 1650 cm−1 appeared (Figure 1a,C), which corresponds to the CO stretching vibration of amide formed between the carboxylated block copolymer and primary amine groups of the protein by EDC/NHS chemistry.20 The disappearance of the peak at 1730 cm−1, which is associated with the carboxylic groups, supports this interaction. After drug loading, no major changes in the IR spectra occur (Figure 1a,B); only a small peak near 1725 cm−1 is visible, which is assigned to CO stretching vibrations present in the carbonyl (ketone) groups of curcumin.17 To support the conjugation of the block copolymer with the protein filament, 1H NMR spectroscopy was employed. The results are shown in Figure 1b. For the carboxylated block copolymer, the peaks of ethylene protons (CH2−CH2) of PEO 19338


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ACS Applied Materials & Interfaces Scheme 1. Proposed Mechanism of Copolymer Conjugation

Figure 2. (a) Fluorescence spectroscopy shows the effect of copolymer concentration on the fluorescence intensity of pyrene in the wavelength range of 350−450 nm. (b) Fluorescence intensity as a function of the copolymer concentration at λex = 335 nm.

particles, the fluorescence intensity of pyrene was plotted at a wavelength of 390 nm as a function of copolymer concentration in a semi-log scale (Figure 2b). The intersection point of the straight lines yields a value of 0.38 mg/mL for CMC. Notably, this concentration is lower than that of CMC values reported for Pluronic micelles (CMC = 1−25 g/L).14,31 Therefore, the conjugated copolymer is easily self-assembled into nanoparticles and the associated micelle is more stable than Pluronic in aqueous solution due to the chemical interactions with keratin. 3.3. Effect of Crosslinking on Drug Loading. The drug was loaded into the conjugated copolymer micelles through the nanoprecipitation method. As shown in Figure 1a, curcumin loading did not change the major IR characteristic peaks of the copolymer, indicating encapsulation of the drug in the nanoparticles without the formation of strong chemical bonds. Figure 3a shows a representative UV−vis spectrum of the drug-loaded copolymer nanoparticles before and after crosslinking. The spectrum of the pristine copolymer is shown for comparison. The peak around 270 nm is assigned to keratin relating to the presence of aromatic amino acids including tryptophan, tyrosine, and phenylalanine in the protein chain.32 The distinct peak at 420 nm is for curcumin caused by electron transition from the π bonding orbital to the antibonding orbital, due to the existence of a benzene ring and carbonyl groups.16 Interestingly, the characteristic peak of the drug is intensified after crosslinking. In agreement with the previous study on curcumin-loaded F127 micelles,16 it appears that the oxidation of keratin’s thiol groups and crosslinking with disulfide bonds enhance the efficiency of encapsulation and drug loading.

and methyl protons (OCH2CH(CH3)O) of PPO are seen at δ ∼ 3.5−3.8 and 1.10 ppm, respectively.24 The weak resonance peak at δ ∼ 2.6 ppm corresponds to the methylene protons (CH2CH2COOH) of succinic groups, revealing carboxylation of terminal hydroxyl groups of the block copolymer. According to the peak area of the methyl protons (δ ∼ 1.1 ppm) and the methylene protons (δ ∼ 2.6 ppm) in the activated Pluronic, about 60% of the hydroxyl groups in Pl were converted to carboxyl groups.29 The spectrum of the conjugated copolymer contains the major peaks of the block copolymer with an additional resonance peak at δ ∼ 2.8 ppm which determines successful conjugation of keratin to the activated Pluronic through an amide bond formed by EDC/NHS chemistry, as shown in Scheme 1, and reported by others.14,29 3.2. Micellization Behavior of Conjugated Nanoparticles in Aqueous Solution. Amphiphilic copolymers containing hydrophobic and hydrophilic segments can selfassemble into nanoparticles in aqueous solution.30 The micellization behavior of the nanoparticles was studied by CMC determination using fluorescence spectroscopy. Entrapment of pyrene molecules into the assembled hydrophobic regions of the micelle changes the polarity in the surrounding environment enhancing the fluorescence intensity of pyrene.26 Figure 2a shows the fluorescence spectra of the conjugated copolymers at different concentrations. No major change in the fluorescence intensity was observed in the concentration range of 0.01−0.1 mg/mL. Nevertheless, at higher concentrations (from 0.5 up to 2.5 mg/mL), the intensity was significantly enhanced, indicating the micellization of the conjugated copolymers. To determine the CMC value for the nano19339


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Figure 3. (a) UV−vis spectra of nanoparticles show the effect of crosslinking on drug-loading efficiency. (b) DLS graphs illustrate the size distribution of nanoparticles at 25 and 37 °C. TEM images show the prepared nanoparticles (c) before and (d) after drug loading and crosslinking.

Potential for high entrapment efficiency is due to the large hydrophobic blocks in Pluronic.16 Curcumin is incorporated into the hydrophobic inner core of copolymer micelles via noncovalent interactions. Similar curcumin loading efficiencies were also observed for other Pluronic-based nanocarriers.13−15 To determine the hydrodynamic size and size distribution of the nanoparticles, DLS was employed. Figure 3b shows representative DLS graphs of the nanoparticles in aqueous solution measured at two different temperatures of 25 and 37 °C. DLS graphs show the distribution of nanoparticles at different sizes (r). The effect of drug loading and crosslinking on the average size (Z-average) of the nanoparticles is reported in Table 2. The results indicate that the average hydrodynamic size of the nanoparticles increased from 152 to 165 nm after

Table 1 shows EE and DL values for different drug to copolymer ratios. It can be concluded that the 1:10 drug to Table 1. Efficiency of Drug Loading in the Conjugated Copolymer Nanoparticles Depending on the Copolymer to the Drug Ratio drug to copolymer ratio drug loading (%) efficiency of encapsulation (%)

1:10 7.4 82

1:20 3.6 76

1:40 0.83 34

polymer ratio has the highest encapsulation efficiency and drug loading. Meanwhile, the drug content of the curcumin-loaded nanoparticles decreases significantly at higher copolymer ratios.

Table 2. Effect of Temperature and Crosslinking on the Average Hydrodynamic Size of Nanoparticles 25 °C

polydispersity index Z-average (nm)

37 °C

conjugated nanoparticle

drug-loaded nanoparticle after crosslinking

drug-loaded nanoparticle after reduction

conjugated nanoparticle

drug-loaded nanoparticle after crosslinking

drug-loaded nanoparticle after reduction

0.327

0.299

0.307

0.473

0.262

0.240

152

165

194

62

66

75

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ACS Applied Materials & Interfaces Scheme 2. Proposed Mechanism of Disulfide Bond Formation

drug loading at 25 °C. It is supposed that the hydrophobic drug is encapsulated into the polymeric nanoparticles by physical entrapment which mainly results from hydrophobic interactions, hydrogen bonds, and van der Waals forces.15 Representative TEM images of the nanoparticles are shown in Figure 3c,d. The morphology of the self-assembled nanoparticles is roughly spherical in shape with clear boundaries. There is an increase in the particle size after drug loading. However, the size of the nanoparticles is smaller than those estimated by DLS, because the TEM images were taken at dry state, whereas the nanoparticles are in a hydrated state during DLS measurements.19,20 The copolymer conjugates can form nanoparticles with a keratin rich core stabilized by Pluronic chains on the shell surface in aqueous media. The keratin rich core of the nanoparticles can be crosslinked via disulfide bonds upon oxidation of the remaining thiol groups (using H2O2), by which the nanoparticles can be stabilized, as shown in Scheme 2. To show the redox sensitivity of the drug-loaded nanoparticles, the disulfide bonds were cleaved by a reducing agent (GSH) and the hydrodynamic size was measured. As shown in Table 2, the average hydrodynamic size is increased from 165 to 194 nm indicating the redox sensitivity of the conjugated copolymer. It is also interesting to note the effect of temperature on the nanoparticle size. The results reveal that at higher temperature, smaller nanoparticles are attained. This observation could be attributed to the dehydration process and strong hydrophobic interactions among the PPO domains at 37 °C resulting in shrinkage in size.33,34 This behavior, which is linked to the thermosensitivity of Pluronic, has been observed in other studies as well.15,16,20 3.4. Stability Study. An ideal nanocarrier is desirable to be stable in aqueous media and capable of releasing the cargo in response to the changes in the environment under physiological

conditions.19 Stability of the nanoparticles was evaluated in PBS at 4 and 37 °C for 30 days. It was found that the solutions remained uniformly transparent in both temperatures during the storage time without any aggregate formation. As shown in Figure 4, the drug-loaded nanoparticles were completely stable

Figure 4. Stability of drug-loaded nanoparticles over different storage times and temperatures.

at 4 °C, as no significant changes were observed in the UV absorbance of the formulations over the time-period. On the other hand, curcumin retention decreased at 37 °C after 5 days, and remained constant afterwards. By the end of 30 days, curcumin retention was higher in the sample stored at 4 °C (about 90%), whereas the sample stored at 37 °C had lower curcumin retention (about 65%). Evidently, the ζ-potential has an important effect on the storage stability of the colloid dispersion system. Particle aggregation is likely to occur if the ζpotential of particles is too low to provide sufficient electric repulsion or steric barriers between each other. The measured ζ-potential of the drug-loaded nanoparticles was −23.6 mV, providing good physical stability to the formulation. The 19341


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Figure 5. Release profile of curcumin from the conjugated nanoparticles in PBS at 37 °C. * indicates significant difference after 48 h (p < 0.05).

Table 3. Model Analysis of the Release Mechanisma kinetic models equation (Mt/M∞)

a

zero order K0t

first order 1 − exp(−K1t)

Higuchi KHt1/2

drug-loaded nanoparticles

R = 0.7534 K0 = 0.06132

R = −1.071 K1 = 1.196

R = 0.8368 KH = 0.7716

drug-loaded nanoparticles + reducing agent

R2 = 0.9895 K0 = 0.3551

R2 = −0.7745 K1 = 0.4562

R2 = 0.8619 KH = 4.325

drug-loaded nanoparticles + trypsin

R2 = 0.8116 K0 = 0.1092

R2 = −1.61 K1 = 5.334

R2 = 0.991 KH = 1.393

drug-loaded nanoparticles + reducing agent + trypsin

R2 = 0.9778 K0 = 0.4256

R2 = −0.9696 K1 = 5.429

R2 = 0.8981 KH = 5.158

2

2

2

power law Ktn R2 = 0.8413 K = 0.5628 n = 0.5647 R2 = 0.9895 K = 0.3375 n = 0.9987 R2 = 0.9914 K = 1.298 n = 0.5144 R2 = 0.9846 K = 0.8762 n = 0.8591

M, K, t, and n are concentration, kinetics constant, time, and kinetics exponent, respectively.

concentration should be preserved between the minimum effective therapeutic level and the maximum tolerable period in circulation.16 Figure 5 also shows that trypsin accelerates the release kinetics. As can be seen, the release is initiated by trypsin. Trypsin is an essential protease found in the body that is generally overexpressed in inflamed and tumorous tissues.19 Under in vivo conditions, the existence of lysine and arginine amino acids in keratin leads to the possibility of cleavage by trypsin. Meanwhile, in the presence of both reducing agent and trypsin, the release kinetics is accelerated even more, revealing the synergistic effect of the reducing agent and the proteolytic action of trypsin. It can be seen that the drug release in the presence of both GSH and trypsin after 192 h is 9 times more than the control sample. Besides, the samples containing GSH or trypsin individually have released their drug 8 times and 2 times more than the control sample, respectively. Statistical analysis also demonstrated a significant difference between the samples after 48 h. To investigate the mechanism of drug release from the nanoparticles, the experimental data were fitted to different kinetic models. Table 3 shows the calculated values of release constants together with the regression coefficients (R2) on the basis of regression analysis. In these models, Mt/M∞ indicates the fraction of the released drug and t is the time. K0, K1, KH, and K are the kinetic constants and are determined by the

negative charge of nanoparticles could be attributed to the carboxylic groups of keratin as well as to the phenolic groups of curcumin. 3.5. Drug Release Studies. Figure 5 shows the release profile of curcumin from the crosslinked nanoparticles in the absence (control) and presence of the reducing agent (GSH) and trypsin. The release kinetics follows a typical two-phase profile. The initial burst release may be ascribed to the fast release of surface-located drugs, whereas the second phase is mainly associated with drug diffusion.13 As it can be seen, 15% of curcumin was released from the nanoparticles containing GSH within the first 48 h, while a lower amount of drug (5%) was released from the control sample. By the end of incubation time, the control sample released about 10% of the total drug, while the sample containing the reducing agent (GSH) released 70% of its total drug. This observation indicates the effect of reducing agent on accelerating the release kinetics through cleaving of the disulfide bonds. It is worth noting that the release of the hydrophobic drug from the nanoparticles depends on the hydrophobic interactions between the drug and the inner core. Stronger interactions result in a slower release of drug from the nanocarriers. Therefore, in the presence of the reducing agent, weaker interactions between the encapsulated drug and nanoparticles result in a higher amount of drug release, displaying redox-dependent kinetics.19 For effective targeting of cancerous tissues, the anticancer drug 19342


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Figure 6. Cell viability of (a, b) HeLa cells and (c, d) L929 fibroblast cells exposed to different concentrations of free and encapsulated curcumin after 24 h incubation, and different formulations after 24 and 72 h of incubation, respectively. * indicates significant difference between samples, ** significant decrease in cell viability with prolonging incubation time (p < 0.05).

experimental data fitting. The exponent of the power equation, n, determines the release mechanism. It can be concluded that the power law equation was best fitted with the experimental release data for the drug-loaded conjugated copolymer (R2 = 0.84−0.99). The results also determine that the addition of trypsin and reducing agent does not affect the release mechanism. For spherical swellable polymeric systems, release exponents between 0.43 and 0.85 are assigned to anomalous transport.35 The calculated n values suggest that the dominant release mechanism behind curcumin release from the nanoparticles is probably controlled by diffusion, in agreement with other studies on block copolymer micelles.16 3.6. Cellular Assay. The biocompatibility of the polymeric carriers is of significant importance for drug delivery application. We employed MTT assay to determine potential cytotoxicity of the drug-loaded nanoparticles after incubation with HeLa and L929 fibroblast cells. The cell viability at different concentrations and times are shown in Figure 6. Pristine curcumin was examined as a control. As seen in Figure

6a,c, the cell viability after 24 h of incubation is >80% for the drug-loaded nanoparticles even at relatively high concentrations up to 1 mg/mL. This finding indicates the good biocompatibility of the nanoparticles as a safe carrier for cancer therapy. Notably, there is no significant difference in cytocompatibility between the drug-encapsulated nanoparticles and pristine curcumin except at 1 mg/mL concentration for HeLa cells (Figure 6a). At longer incubation time, however, cytocompatibility decreased significantly for HeLa cells. There is also significant difference between free drug and conjugated as well as drug-loaded nanoparticles at 72 h (Figure 6b). However, no significant differences can be seen in samples incubated with L929 fibroblast cells (Figure 6d). The enhanced cytotoxicity of the drug-loaded nanoparticles is attributed to the higher uptake of curcumin, which effectively inhibits HeLa growth.13 The anticancer activity of curcumin, through directly killing tumor cells as well as inhibiting angiogenesis, has been reported by others as well.25,36 It should be noted that free curcumin as a hydrophobic drug is insoluble and incapable of cellular uptake. 19343


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Figure 7. Fluorescence microscope images of HeLa cells incubated with curcumin-loaded nanoparticles at 37 °C after an incubation time of (a) 24 h and (b) 48 h. Intracellular delivery of curcumin-loaded nanoparticles into HeLa cells by confocal laser scanning microscopy: (c) DAPI, (d) curcumin, (e) merged curcumin with DAPI, (f) bright field, (g) merged curcumin with bright field, and (h) merged curcumin with bright field and DAPI.

florescence with the corresponding bright field or DAPI image shows that the nanoparticles mainly reside inside the cells. Therefore, the designed nanoparticles are efficient vehicles to deliver the drug molecules to the cytoplasm of the cancer cells. Our results are consistent with previous studies.13,16

Therefore, the drug need to be incorporated into a stable nanocarrier. The obtained results reveal that the designed polymeric nanoparticles might serve as a potential carrier to improve in vitro cytotoxicity of the drug. The intracellular uptake of nanoparticles by HeLa cells was analyzed by fluorescence microscopy after 24 and 48 h of incubation. Curcumin is naturally fluorescent, which makes direct visualization of drug uptake by cells possible. Figure 7a,b show intrinsic green fluorescence of the drug, validating the cellular uptake of curcumin-loaded nanoparticles into the cells. Greater fluorescence intensity with increasing time corresponds to more efficient curcumin delivery to cells. To better show the intracellular delivery of the curcumin-loaded nanoparticles into HeLa cells, confocal laser scanning microscopy was also employed (Figure 7c−h). DAPI (blue fluorescence) was used to stain the nuclei of the cells. The overlay of the curcumin

4. CONCLUSIONS Novel nanoparticles based on thermosensitive copolymers conjugated with protein filaments were prepared for triggerable drug delivery. The nanoparticles were prepared by conjugation of block copolymers (Pluronic) and wool-extracted protein filaments (keratin) by EDH/NHS chemistry. Curcumin was encapsulated inside the copolymer micelles through the nanoprecipitation method. The nanoparticles had an average diameter of 152 nm at 25 °C but they shrunk to 62 nm at 37 19344


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ACS Applied Materials & Interfaces °C. To control the release profile, disulfide crosslinking of keratin was performed. It was shown that the crosslinking enhanced the stability of the nanoparticles and controlled the drug delivery in the redox environment. The addition of a reducing agent (GSH) could promote the release kinetics through breaking of the disulfide bonds. It was also shown that trypsin improved the release kinetics. Dose-dependent cytocompatibility assay determined good biocompatibility of the nanoparticles (80%) even at high concentrations (≤1 mg/ mL). Fluorescence microscope images confirmed the cellular internalization of drug-loaded nanoparticles. The prepared nanoparticles with dual redox and temperature sensitivity have the potential to meet some of the challenges of hydrophobic drug formulation for delivery to cancer cells.

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected] (N.E.). *E-mail: [email protected]. Tel: +98 (21) 6616 5226. Fax: +98 (21) 6600 5717 (A.S.). ORCID

Abdolreza Simchi: 0000-0002-9111-2977 Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS Dr. Eslahi wishes to acknowledge the National Elites Foundation of Iran for Allameh post-doctoral fellowship. Dr. Simchi thanks funding support of Sharif University of Technology (Grant Program No. G930305) and Iran National Science Foundation (INSF, Grant No. 95-S-48740). They also acknowledge Arash Ramedani (Sharif University of Technology), Ali Dinari, and Behnam Hajipour (Tarbiat Modares University), and the personnel of the Cell Bank of Iran Pasteur Institute for their cooperation and useful consultation on biological assay.

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