Dual-Sensitive Hydrogel Nanoparticles Based on Conjugated

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Biological and Medical Applications of Materials and Interfaces

Dual-sensitive hydrogel nanoparticles based on conjugated thermoresponsive copolymers and protein filaments for triggerable drug delivery Roujin Ghaffari, Niloofar Eslahi, Elnaz Tamjid, and Abdolreza Arash Simchi ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.8b01154 • Publication Date (Web): 17 May 2018 Downloaded from http://pubs.acs.org on May 17, 2018

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ACS Applied Materials & Interfaces

Dual-sensitive hydrogel nanoparticles based on conjugated thermoresponsive copolymers and protein filaments for triggerable drug delivery Roujin Ghaffari1, Niloofar Eslahi2*, Elnaz Tamjid3, Abdolreza Simchi1,4* 1

Department of Materials Science and Engineering, Sharif University of Technology, Azadi

Avenue, P.O. Box 11365/8639, Tehran, Iran 2

Department of Textile Engineering, Science and Research Branch, Islamic Azad University,

P.O. Box 14515/775, Tehran, Iran 3

Department of Nanobiotechnology, Faculty of Biological Sciences, Tarbiat Modares University,

Tehran, Iran 4

Institute for Nanoscience and Nanotechnology, Sharif University of Technology, Azadi Avenue,

P.O. Box 11365/8639, Tehran, Iran *Corresponding authors: [email protected]; [email protected] Tel: +98 (21) 6616 5226; Fax: +98 (21) 6600 5717

ACS Paragon Plus Environment

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Abstract

In this study, a novel hydrogel nanoparticles with dual triggerable release properties based on fibrous structural proteins (Keratin) and thermoresponsive copolymers (Pluronic) is introduced. The nanoparticle was used for Curcumin delivery, as an effective and safe anticancer agent, hydrophobicity of which has limited its clinical application. The drug was loaded into hydrogel nanoparticles by a single-step nanoprecipitation method. The drug-loaded nanoparticles had an average dimeter of 165 nm and 66 nm at 25ºC and 37ºC, respectively. It was shown that the drug loading efficiency could be enhanced through crosslinking of the disulfide bonds in Kertain. Crosslinking provided a targeted release profile under reductive condition using an in vivo agent, glutathione (GSH), or in presence of trypsin. Cytocompatibility assay using HeLa and L929 Fibroblast cells exhibited no adverse effect of the nanoparticles on the 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 efficient delivery of hydrophobic drugs to targeted regions having triggerable release profile.

Keywords: Drug delivery; Hydrogel nanoparticles; Stimuli-responsive copolymer; Keratin biopolymer, Pluronic conjugates

1. Introduction In the past decade, advancement in smart multi-targeting drug delivery systems using nanotechnology have gained considerable attraction due to their more efficient drug release.1


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Originally, polymers were used to stabilize or solubilize drugs for a controlled release, however, progress in new polymer designs with the development of synthetic strategies have 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 in order to overcome the shortcomings of conventional drug formulations.3, 4 Nanogels or hydrogel nanoparticles are three-dimensional (3D) networks of cross-linked polymer chains. These nanocarriers are very promising systems for target-specific delivery of drugs, because they have high drug loading capacity and enhanced cellular uptake efficiency 5-7 with an ability to protect the drug and postpone its degradation.8, 9 Recent advances deal with development of stimuli responsive polymeric hydrogels.10,11 For this aim, thermoresponsive triblock copolymers such as Pluronic (PEO–PPO–PEO) has extensively been studied, especially in drug delivery due to their approved use in medical devices by the US Food and Drug Administration (FDA).12 Self-assembly of the amphiphilic copolymer through hydrophobic interactions of PPO block in the inner core and hydrophilic interactions of PEO blocks in the outer shell forms water soluble micelles above the lower critical solution temperature (LCST).13 Hydrophobic drugs can easily be loaded inside the micelles in order 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, anti-oxidant, antimutagenic, and anti-tumor 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 in Pluronic as cargo


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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 more controllable release profile with higher efficiency.20 Based on this concept, Nahain et el.21 have recently prepared a dual-sensitive chemically crosslinked polymer having a pH responsive covalent benzoic-imine bond and redox sensitive disulfide in order 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 is temperature and redox sensitive and is prepared by chemical grafting of triblock copolymers (Pluronic) with protein filaments (Keratin). As 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 to cells 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 The presence of thiol groups can be used as disulfide linkages to facilitate selective release of drugs in reductive environments. For instance, a redox-sensitive drug delivery system based on Keratin-g- polyethylene glycol loaded with doxorubicin exhibited controlled release


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profile with better efficiency.19 Therefore, we utilized Keratin to prepare hydrogel nanoparticles aiming to attain more controllable drug released. For this purpose, Keratin as a redox-sensitive polymer was conjugated with Pluronic as a temperature-sensitive copolymer to produce a novel dual sensitive nanoparticles. The conjugated nanoparticles were used for encapsulation of an anti-inflammatory and anti-tumor drug (Curcumin). It is shown that a high drug loading efficiency with controllable release profile under reductive condition and in the presence of trypsin is obtained. The physicochemical properties and cytotoxicity of the nanohydrogel are also demonstrated. 2. Materials and Methods 2.1 Materials Pluronic 127 (PEO99–PPO65–PEO99), with molecular weight of 12.5 kDa, was purchased from

Sigma-Aldrich

(USA).

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 based on our previous study.24 In brief, cleaned defatted wool fibres 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 freeze-dryer (Lyotrap/Plus, UK) at −40 °C to obtain Keratin powder (MW: 45-65 kDa). 2.2 Preparation of polymeric nanoparticles


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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 vacuum (200 torr) for 12 h. Afterwards, the carboxylated copolymer (3 g) was activated by EDC (0.22g) and NHS (0.14g) 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 singlestep nano-precipitation 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 Mῼ) drop by drop while stirring. During this process, the copolymer was self-assembled into nanoparticles with core-encapsulated curcumin. For crosslinking by disulfide bonds through 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 unencapsulated drug. The resulting solution was freeze-dried 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, USA), and the efficiency and loading percentages were calculated by 25:


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= =

(1)

∗ 100

∗ 100

(2)

2.4 Physicochemical characterizations of nanoparticles Fourier transform infrared spectroscopy (FTIR, ABB Bomem MB100, USA) was employed in transmission mode using KBr pellets in the wavenumber range of 4000–400 cm-1 at a resolution of 4 cm-1. 1H NMR spectra was obtained on a NMR 500 MHz (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, US) 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 fluorescence spectra were recorded between 350 nm 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 (ZEN3600, Malvern, UK) at two temperatures of 25ºC and 37ºC. Zeta 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 drug-loaded nanoparticles was examined in PBS at 4°C and 37°C for 30 days. After the respective storage periods, the samples were centrifuged for 10 minutes (at 6000 rpm) and examined not only qualitatively by naked eye, but also quantitatively by UV/Vis


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spectroscopy. Presence of precipitations in the solutions indicates instability of the nanoparticles, while a uniformly transparent solution indicates their stability. 2.5 In vitro release study In order to determine the concentration-dependent 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 ºC with a 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 cleaves by trypsin protease in 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 in cytoplasm.19 The mechanism of drug release was analyzed by curve fitting of the experimental results using kinetics 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-diphenylte-trazoliumbromide (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 RPMI-1640 medium containing 10% (v/v) FBS, which was supplemented with 40-50U/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


8

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(0.01, 0.1, 1 mg/ml)) was replaced. After incubation for one or three 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 (DMSO) and shacked 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 h and 48 h. The delivery of curcumin into HeLa cells was visualized using a fluorescent microscope (Eclipse TE2000-S, Nikon Instruments Inc.). In addition, the nuclei were stained by 4′,6-diamidino-2-phenylindole (DAPI) and the cells were imaged after 48 incubation using Leica TCS SPE confocal laser scanning microscope (CLSM, Germany) with excitation at 488 nm for curcumin. 2.7 Statistical analysis All data were reported as mean ± SD, and each experiment was done in triplicate. Multifactorial one-way analysis of variance (ANOVA) was performed for the comparison of groups and p value ≤ 0.05 was considered to be statistically significant. 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 Ketrain. The absorbing bands at 1108 and 1243 cm-1 are ascribed to the C–O–C stretching of aliphatic ether group and twisting vibration of –CH2 in Pluronic, respectively (Fig. 1a-D). A new peak is appeared at 1730 cm-1 after carboxylation of the copolymer by means of succinic anhydride (Fig. 1a-E), which is assigned to the carboxyl (COOH) stretching vibration.24 The spectrum of Keratin shows characteristic absorption bands


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assigned mainly to the peptide bonds, known as Amide bonds (Fig. 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 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 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 is appeared (Fig. 1a-C) that 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 occurrs (Fig. 1a-B); only a small peak near 1725 cm-1 is visible, which is assigned to C=O stretching vibrations present in carbonyl (ketone) groups of curcumin.17 To support conjugation of the block copolymer with the protein filament,

1

H 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 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


10

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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

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


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loading does not change the characteristic bands of the conjugated copolymer. (b) 1H NMR spectra of carboxylated and conjugated copolymer.

Scheme 1 Proposed mechanism of copolymer conjugation.

3.2 Micellization behavior of conjugated nanoparticles in aqueous solution Amphiphilic copolymers containing hydrophobic and hydrophilic segments can self-assemble 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 copolymer. In order to determine the CMC value for the nanoparticles, the fluorescence intensity of pyrene was plotted at a wavelength of 390 nm as a


12

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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) Florescence intensity as a function of the copolymer concentration at λex=335 nm.


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function of copolymer concentration in 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 the 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 nano-precipitation method. As shown in Figure 1a, Curcumin loading did not change the major IR characteristics peaks of the copolymer, indicating encapsulation of the drug in the nanoparticles without formation of strong chemical bonds. Figure 3a shows 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 π bonding orbital to antibonding orbital due to existence of benzene ring and carbonyl groups.16 Interestingly, the characteristic peak of the drug is intensified after crosslinking. In agreement with previous study on Curcumin-loaded F127 micelles,16 it appears that oxidation of Keratin’s thiol groups and crosslinking with disulfide bonds enhances the efficiency of encapsulation and drug loading.


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(a)

(b)

(c)

(d)

Figure 3 (a) UV-Vis spectra of nanoparticles show the effect of crosslinking on the drug loading efficiency. (b) DLS graphs illustrate the size distribution of nanoparticles at 25ºC and 37°C. TEM images show the prepared nanoparticles (c) before and (d) after drug loading and crosslinking. Table 1 shows EE and DL values for different drug to copolymer ratios. It can be concluded that 1:10 drug to polymer ratio has the highest encapsulation efficiency and drug loading.


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Meanwhile, the drug content of the Curcumin loaded nanoparticles decrease significantly at higher copolymer ratio. 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 non-covalent interactions. The similar Curcumin loading efficiencies were also observed for other Pluronic-based nanocarriers.13-15 Table 1 Efficiency of drug loading in the conjugated copolymer nanoparticles depending on the copolymer to the drug ratio Drug to copolymer ratio

1:10

1:20

1:40

Drug loading (%)

7.4

3.6

0.83

Efficiency of encapsulation (%)

82

76

34

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 25ºC and 37ºC. DLS graphs shows 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 indicates that the average hydrodynamic size of the nanoparticles is increased from 152 nm to 165 nm after drug loading at 25ºC. 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 Figures 3c and 3d. The morphology of the self-assembled nanoparticles are roughly spherical in shape with clear boundaries. There is an increase in the particles size after drug loading. However, the size


16

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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 Table 2 Effect of temperature and crosslinking on the average hydrodynamic size of nanoparticles. Temperature

25 °C

37 °C

Nanoparticle

Conjugated nanoparticle

Drug loaded nanoparticle after crosslinking

Drug loaded nanoparticle after reduction

Conjugated nanoparticle

Drug loaded nanoparticle after crosslinking

Drug loaded nanoparticle after reduction

PDI

0.327

0.299

0.307

0.473

0.262

0.240

Z-average (nm)

152

165

194

62

66

75

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 under 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 Table 2 shows, the average hydrodynamic size is increased from 165 nm to 194 nm indicating the redox sensitivity of the conjugated copolymer. It is also interesting to note the effect of temperature on the nanoparticles size. The results reveal that at the 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 ºC resulting in shrinkage in size.33 This behavior, which is linked to the thermosensitivity of Pluronic, has been observed in other studies as well.15, 16, 20


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Scheme 2 Proposed mechanism of disulfide bonds formation.

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°C and 37°C for 30 days. It was found that the solutions remained uniformly transparent in both temperature during the storage time without any aggregate formation. As shown in Figure 4, the drug loaded nanoparticles were completely stable at 4°C, as no significant changes was observed in 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 is higher in sample stored at 4°C (about 90%), whereas the sample stored at 37°C has lower curcumin retention


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(about 65%). Evidently, zeta potential has an important effect on the storage stability of colloid dispersion system. Particle aggregation is likely to occur if the zeta-potential of particles is too low to provide sufficient electric repulsion or steric barriers between each other. The measured zeta potential of the drug loaded nanoparticles was -23.6 mV, providing good physical stability to the formulation. The negative charge of nanoparticles could be attributed to the carboxylic groups of keratin as well as phenolic groups of Curcumin.

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

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 lower amount of drug (5%) was released from the control sample. By the end of


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incubation time, control sample released about 10% of the total drug, while the sample containing reducing agent (GSH) has 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 noteworthy 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 anti-cancer drug 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 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 condition, the existence of lysine and arginine amino acids in Keratin provides 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 drug release in 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 samples after 48 h.


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Figure 5 Release profile of Curcumin from the conjugated nanoparticles in PBS at 37 ºC. * indicates significant difference after 48 h (p < 0.05). To investigate the mechanism of drug release from the nanoparticles, the experimental data was 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 released drug and t is time. K0, K1, KH and K are the kinetics constant and determined by the experimental data fitting. The exponent of the power equation, n, determines the release mechanism. It can be concluded that 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


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dominant release mechanism behind Curcumin release from the nanoparticles is probably controlled by diffusion, in agreement with other studies on block copolymer micelles.16 Table 3 Model analysis of the release mechanism. M, K, t, and n are concentration, kinetics constant, time, and kinetics exponent, respectively. Kinetics model

Zero order

First order

Higuchi

Power law

K 0t

1 – exp (–K1t)

KHt1/2

Ktn

Drug loaded

R2= 0.7534

R2= -1.071

R2= 0.8368

R2= 0.8413

nanoparticles

K0= 0.06132

K1= 1.196

KH = 0.7716

K = 0.5628

Equation (Mt/ M∞)

n= 0.5647 Drug loaded

R2= 0.9895

R2= -0.7745

R2= 0.8619

R2= 0.9895

nanoparticles +

K0 = 0.3551

K1= 0.4562

KH = 4.325

K = 0.3375 n= 0.9987

Reducing Agent Drug loaded

R2= 0.8116

R2= -1.61

R2= 0.991

R2= 0.9914

nanoparticles +

K0 = 0.1092

K1= 5.334

KH = 1.393

K = 1.298 n= 0.5144

Trypsin Drug loaded

R2= 0.9778

R2= - 0.9696

R2= 0.8981

R2= 0.9846

nanoparticles +

K0= 0.4256

K1=5.429

KH = 5.158

K = 0.8762 n= 0.8591

Reducing Agent + Trypsin

3.6 Cellular assay The biocompatibility of the polymeric carriers is of significant importance for the drug delivery application. We employed MTT assay to determine potential cytotoxicity of the drugloaded 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


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control. As seen in Figures 6a and 6c, the cell viability after 24 h 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 the cytocompatibility between the drugencapsulated nanoparticles and pristine Curcumin except at 1 mg/ml concentration for HeLa cells (Figure 6a). At longer incubation time, however, cytocompatibility is 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 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. 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 the nanoparticles by HeLa cells was analyzed by fluorescence microscopy after 24 h and 48 incubation. Curcumin is naturally fluorescent, which makes direct visualization of drug uptake by cells possible. Figure 7a and 7b show intrinsic green fluorescence of the drug, validating the cellular uptake of Curcumin-loaded nanoparticles into the cells. Greater fluorescent intensity with increasing time corresponds to more efficient curcumin delivery

to

cells.

In order to better show the intracellular delivery of the curcumin-loaded nanoparticles into HeLa


23


Page 24 of 41

cells, confocal laser scanning microscopy was also employed (Figures 7c-7h). DAPI (blue fluorescence) was used to stain the nuclei of the cells. The overlay of the curcumin 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 in the cytoplasm of the cancer cells. Our results are consistent with previous studies.13,16

(a)

(b)

(c)

(d)

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


24

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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).

(a)

(b)

(c)

(d)

(e)

(f)

(g)

(h)


25


Page 26 of 41

Figure 7 Fluorescence microscope images of HeLa cells incubated with curcumin-loaded nanoparticles at 37ºC after 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. 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/NS chemistry. Curcumin was encapsulated inside the copolymer micelles through the nanoprecipitation method. The nanoparticles had an average diameter of 152 nm at 25 ºC but they shrunk to 62 nm at 37 ºC. In order to control the release profile, disulfide-crosslinking of Keratin was performed. It was shown that the crosslinking enhanced the nanoparticles stability and controlled drug delivery under 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 cellular internalization of the 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. Acknowledgments


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Dr. Eslahi wishes to acknowledge the National Elites Foundation of Iran for Allameh postdoctoral fellowship. Dr. Simchi thanks funding support of Sharif University of Technology (Grant Program No. G930305) and Iran National Science Foundation (INSF, Grant No. 95-S48740). They also acknowledge Mr. Arash Ramedani (Sharif university of Technology), Ali Dinari and Behnam Hajipour (Tarbiat Modares University), and the personnel of Cell Bank of Iran Pasteur Institute for their cooperation and useful consultation on biological assay. References 1.

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Table of Contents (TOC)


31

ACS Applied Materials & Interfaces 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54


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Page 33 ACS of 41 Applied Materials & Interfaces 1 2 3 4




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ACS Applied Materials & Interfaces Page 36 of 41 1 2 3 4 5 6 7 8


Page 37 ACS of 41 Applied Materials & Interfaces 1 2 3 4 5 6


ACS Applied Materials & Interfaces Page 38 of 41 1 2 3 4 5 6 7


Page 39 of 41 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60



ACS Applied Materials & Interfaces 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35


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Page ACS 41 ofApplied 41 Materials & Interfaces 1 2 3 4


Dual-Sensitive Hydrogel Nanoparticles Based on Conjugated

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