Phospholipid End-Capped Bioreducible Polyurea Micelles as a

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Phospholipid End-Capped Bioreducible Polyurea Micelles as a Potential Platform for Intracellular Drug Delivery of Doxorubicin in Tumor Cells Eun Jin Seo, Johnson V. John, Rimesh Augustine, Il Ho Jang, Dae Kyoung Kim, Yang Woo Kwon, Jae Ho Kim, and Il Kim ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.6b00256 • Publication Date (Web): 19 Sep 2016 Downloaded from http://pubs.acs.org on September 21, 2016

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ACS Biomaterials Science & Engineering

Phospholipid End-Capped Bioreducible Polyurea Micelles as a Potential Platform for Intracellular Drug Delivery of Doxorubicin in Tumor Cells Johnson V. John,† Eun Jin Seo,‡ Rimesh Augustine,† Il Ho Jang,‡ Dae Kyoung Kim, ‡ Yang Woo Kwon, ‡ Jae Ho Kim,‡* and Il Kim† * †

BK21 PLUS Center for Advanced Chemical Technology, Department Polymer Science and

Engineering, Pusan National University, Geumjeong-gu, Busan 609-735, Republic of Korea ‡

Department of Physiology, School of Medicine, Pusan National University, Yangsan 626-870,

Gyeongsangnam-do, Republic of Korea KEYWORDS. Degradation, Drug delivery, Nanocarrier, Phospholipid, Polyurea, Tumor cell

ACS Paragon Plus Environment

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ABSTRACT. Bioreducible polymeric nanocarriers bearing disulfide bonds have been widely used for intracellular therapeutic delivery, since they are quickly sliced or reduced in the reductive milieu of cytosol. Incorporation of hydrophobic phospholipid analogues to polymers improves the biocompatibility by reducing the protein adsorption and platelet adhesion on the cell membranes. In this study, we have developed a series of bioreducible polyurea (PUs) bearing disulfide linkages in their backbone and phospholipid moieties in their chain ends. The reducible PUs exhibit interesting self-assembly behavior and controlled release profiles at intracellular mimic conditions. The self-assembled hybrid nanocarriers with an average diameter of about 110 nm efficiently encapsulated the model anticancer drug doxorubicin (Dox). The in vitro Dox release profile demonstrated a good glutathione (GSH)-responsive release of Dox at 10 mM GSH. An in vitro cell viability assay was also performed with various cell lines. The antitumor activity tests using HCT15 and HCT116 cancer cells showed that Dox-loaded nanocarriers bearing disulfide linkages induced significantly higher cytotoxicity in cancer cells than those without disulfide linkages. Hence, the PU nanocarriers bearing disulfide linkers and α,ω-phospholipid moieties have a promising potential to trigger the drug into the intracellular compartment of cancer cells.


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INTRODUCTION Clinical results of various anticancer drugs were not satisfactory due to the poor solubility, high toxicity, low stability, permeability and non-targetability of the drugs. To address these issues, polymeric nanocarriers, composed of biocompatible polymers with variety of functionalities, have appeared as a promising choice for drug delivery. Unique architecture and functionalities of polymeric nanocarriers offer tailor-made properties like higher encapsulation of therapeutics, prolonged circulation in the blood stream, resistance to renal excretion, and preferential accumulation into tumor tissue via the enhanced permeation and retention (EPR) effect.1–3 Even though a high level of therapeutic uptake exhibited at the tumor tissue, the effective concentration of the anticancer drug is often inadequate within cancer cells due to slow drug release into the cytosol from the polymeric nanocarriers.2,4 Therefore, rapid and controlled intracellular drug release is necessary after the micelles reach at the cytoplasm to improve anticancer efficacy, as well as reducing drug resistance in cancer cells. To date, stimuli-responsive nanocarriers circumvent these challenges in anticancer therapy by triggering the therapeutic payload into specific diseased tissues based on the unique features of cancer tissues. Hence, the synthesis and fabrication of stimuli-responsive nanocarriers by intra or extracellular stimuli, such as temperature, pH, light, or glutathione (GSH) has been emerged as a topical area of research for the controlled delivery of therapeutics.5,6 Among these stimuli, GSH-responsive nanocarriers have been received much attention due to the significant variation in the redox potential between the extracellular environment (usually in the micromolar scale, ∼10 µM in the plasma) and the intracellular environment (approximately 1–10 mM in the cytoplasm). Several studies have mentioned that cancer cells usually show higher levels of cytosolic GSH than normal cells.7–9 This significant difference in GSH concentration facilitates


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the cleavage of disulfide bonds in the cytosol, suggesting that GSH-responsive nanocarriers are a good choice for intracellular therapeutic delivery against cancer. Over the last few years, redox-responsive or bioreducible polymer nanocarriers have been widely employed in anticancer therapy. Usually, the disulfide linker incorporated in backbone10– 12

and side chains13,14 of polymer architecture or using as reducible cross-linkers15–19 has been

widely used as a good candidate for the enhanced delivery of antitumor drugs into tumor cells. Considering the cleavage of disulfide bond results in the cleavage of polymer backbone or the break of micellar shell, the polymers bearing disulfide linkages in the backbone are smart enough to disintegrate the polymer nanocarriers for the enhanced therapeutic delivery. Most of the studies were held on the developments of di- or tri-block copolymers with disulfide bonds in between the hydrophilic and hydrophobic blocks 20–22, or using a disulfide bearing pendant group as the hydrophobic block.23–25 However, the multi-step synthesis and/or the difficult processing did not allow them to be used broadly, and more commercially tangible as well as effective approaches should be necessary to develop. Recently, a class of biodegradable multi-block polyurethanes (PUs) bearing different amounts of disulfide bond in their backbone have been studied.10–12 The redox-sensitive PU micelles enhance the endocytosis into tumor cells and efficiently trigger the encapsulated payloads into the cytosol. The good biocompatibility and the tailorability of PUs enable the synthesis of different functional groups and hydrophobic/hydrophilic segments into the polymer chains to fabricate versatile drug delivery systems.26–28 Block copolymers form nanostructures by self-assembly and their drug loading efficiency, cellular uptake and degradability can be tuned by the constitution of them.


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Natural phospholipids (PLs) play a significant role in the cell surface structural formation.1 Thus, the incorporation of PL analogues to polymers is one of the effective ways in improving biocompatibility of the polymer nanocarriers by suppressing protein adsorption and platelet adhesion on the surface of cell membranes.29–34 Recently Wang and coworkers reported that phosphorylcholine end-capped PUs could improve blood compatibility of the hybrid system by suppressing platelet adhesion and activation effectively.35 Herein, we report a class of bioreducible non-block PUs bearing disulfide linkage in the backbone and PL moieties at the chain ends for the redox-responsive delivery of doxorubicin (Dox). A simple condensation reaction of cystamine and isophorone diisocyanate (IPDI), followed by addition of PL (soybean-based phosphatidylethanolamine (SPE) and 1,2-bis(10,12tricosadiynoyl)-sn-glycero-3-phosphoethanolamine (DCPE)) gave target polymers. A PU bearing no disulfide linkage was also synthesized and compared. The physiochemical characteristics of the PL end-capped PU with disulfide linkage (PL–PU(SS)–PL) hybrid materials were determined using NMR, dynamic light scattering (DLS), and transmission electron microscopy (TEM). In vitro release behavior of Dox from self-assembled nanocarriers was measured in the absence/presence of GSH at pH 7.4. In addition, the cytotoxicity and the intracellular drug internalization behavior were evaluated using HEK293ft normal cells, and HCT15 and HCT116 cancer cells. EXPERIMENTAL SECTION Materials. Dibutyline tin diaulerate (DBTL), cystamine dichloride (Sigma–Aldrich), and IPDI (Acros Organics) were used without further purifications. DCPE was purchased from Avanti Polar Lipids and SPE was extracted from soybean lecithin by solvent extraction.36 N,N-


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Dimethylformamide (DMF) was distilled over sodium, and acetone and chloroform were distilled over calcium hydride. Instrumentation and measurements. 1H NMR (400 MHz), 13C NMR (100 MHz) and 31P NMR (161.9 MHz) spectra were recorded on a Varian INOVA 400 NMR spectrometer. Chemical shifts are presented in parts per million (ppm), relative to the residual solvent peaks as the internal standard. Peak multiplicities in 1H NMR spectra are abbreviated as s (singlet), d (doublet), t (triplet), m (multiplet), or br (broad). Fourier transform infrared (FT-IR) spectra were recorded on a Shimadzu IR prestige 21 spectrometer. The spectra were recorded using KBr discs at wavelengths ranging from 4000−600 cm−1. The molecular weight (MW) and polydispersity index (Đ) of the polymers were measured with a Waters GPC system (equipped with a Waters 1515 HPLC solvent pump), a Waters 2414 refractive index detector, and three Waters Styragel High Resolution columns (HR4, HR2, HR1, with effective MW ranges of 5000−500,000, 500−20,000, and 100−5000 g/mol, respectively) at 40 °C using high-performance liquid chromatography (HPLC)-grade DMF containing 0.1 N LiBr eluent at a flow rate of 1.0 mL/min. Monodisperse polystyrenes were used to generate the calibration curve. DLS measurements were carried out using a Nano ZS90 zeta potential analyzer (Malvern Instruments, Ltd., U.K.) with a He−Ne laser (633 nm), 90° collecting optics, and a thermoelectric Peltier temperature controller. Block copolymer solutions (2 mg/mL) were filtered through a 0.5µM filter prior to use. Particle morphology was analyzed by transmission electron microscopy (TEM). TEM was performed using a JEOL−1299EX electron microscope with an accelerating voltage of 80 keV. TEM samples were prepared in grids with formvar film and treated with oxygen plasma (from a Harrick plasma cleaner/sterilizer) for 15 s to make their surface hydrophilic. A TEM grid was then floated on top of the bead with the hydrophilic face in contact


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with the solution; the aqueous solution was blotted away with a strip of filter paper and the samples were dried overnight at room temperature. Synthesis of phospholipid end-capped PU. A series of PL−PU(SS)−PL polymers were synthesized according to the reported procedures15 with slight modifications. The brief reaction procedure is shown in Scheme 1 and the feed ratios of all original materials are listed in Table 1. For example, cystamine and a catalytic amount of DBTDL were completely dissolved in DMF at 60 °C under dry nitrogen atmosphere. Then, the prescribed amount of IPDI was added into the flask under stirring and the pre-polymerization reaction was allowed to proceed at 60 °C for 2−3 h. Phospholipid solution in CHCl3 was then added drop wise into the reaction mixture. The reaction mixture was further stirred for 2 h at 60 °C, and then the flask cooled to room temperature. The resulting polymer was isolated by precipitating in excess of diethyl ether, purified by repeated precipitation into diethyl ether from DMF solution, and eventually dried under vacuum at room temperature for 48 h. The yields of all resulting copolymers were over 80 %. The control sample bearing no disulfide linkage (PL-PU-PL) was prepared using hexane-1,6diamine (HDA) instead of cystamine in similar procedures. Preparation and characterization of PU nanocarriers. PL−PU(SS)−PL and PL−PU−PL nanocarriers were fabricated by simultaneous heating-cooling, followed by dialysis. Initially, 20 mg of PL−PU(SS)−PL or PL−PU−PL was dissolved in 7 mL DMSO, then 3 mL deionized water was added drop wise into the polymer solution; the resulting solution was heated to 50 °C for 1 h. The clear solution mixture was then allowed to cool slowly at a rate of 1 °C/min in an automated thermal cycler, to 25 °C. The resulting solution was transferred into a dialysis membrane (MW cut-off of 1000) at 25 °C and dialyzed against 10 mM phosphate-buffered saline (PBS) pH 7.4. The outer phase was replaced with fresh buffer solution at 1, 2, 4, 6, and 12 h.


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After 24 h, the solution containing PU nanocarriers inside the membrane were collected and subsequently lyophilized after filtering through an 800 nm syringe filter. The yield (62 wt%) of nanocarriers was calculated by weighing the lyophilized micelle powder. To

prepare

the

Dox-loaded

nanocarriers,

20

mg

of

DCPE−PU(SS)−DCPE,

SPE−PU(SS)−SPE or DCPE−PU−DCPE was dissolved in 4 mL of DMSO, then Dox (20 mg) was added to each solution and lipophilized by using triethylamine. 2 mL of deionized water was added drop wise into the polymer solution. The resulting solution was heated to 50 °C for 1 h and cooled slowly at a rate of 1 °C/min to 25 °C in an automated thermal-cycler, and then dialyzed against 10 mM PBS, pH 7.4, at room temperature for 24 h. The final mixed nanocarriers were used after filtration through an 800-nm syringe filter (61% yield). For quantitative analysis of drug encapsulation, aliquots of the drug-loaded micelle solution were lyophilized and dissolved in 2 mL of DMSO and analyzed using UV−Vis spectroscopy. The characteristic absorbance of Dox (485 nm) was recorded and compared to a five-point standard curve of Dox, from 0 to 50 mg/mL in DMSO. The percentages of drug loading content (DLC) and drug loading efficiency (DLE) were calculated using the following equations. DLC % =

DLE % =

weight of the drug in the micelle × 100 weight of drug loaded micelle

weight of the drug in the micelle × 100 weight of drug for drug loaded micelle preparation

The release profiles of Dox-loaded DCPE−PU(SS)−DCPE and DCPE−PU−DCPE nanocarriers were determined in PBS (100 mM, pH 7.4) at 37 °C at various concentrations of GSH form (0 to 20 mM GSH). The Dox-loaded hybrid micelle solutions were divided at various aliquots. Each aliquot was transferred to a dialysis tube (MWCO of 2000). The dialysis tube was immersed into GSH containing buffer 20 mL of buffer and was shaken at 37 °C. At desired time


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intervals, 1 mL of the release medium was taken out to measure the fluorescence and replaced with the fresh medium. The concentration of Dox was determined by UV measurements (excitation at 484 nm). To determine the loaded drug content, Dox-loaded micelles solutions were freeze-dried, then dissolved in DMSO, and analyzed with UV−visible spectroscopy. A calibration curve was obtained using Free Dox/DMSO solutions with different Dox concentrations. To determine the amount of Dox released, calibration curves were run with Dox/phosphate buffer solutions with different Dox concentrations at pH 7.4. The emission at 484 nm was recorded. Release experiments were conducted in triplicate. Reduction sensitivity of nanocarriers at various GSH concentrations. To investigate the reductive

degradation

of

Dox-loaded

DCPE−PU(SS)−DCPE

and

DCPE−PU−DCPE

nanocarriers, they were exposed at various GSH concentrations (0−20 mM) mimicking intracellular reduction environment of tumor cells. They were incubated at 37 ºC for 12 h and then were analyzed by using fluorescence spectrophotometer and fluorescence microscope for the confirmation of reduction sensitive character. Cell culture studies. Human dermal fibroblast (primary cell) and Human embryonic kidney (HEK293ft) cells were maintained in Dulbecco’s modified Eagle’s medium with 10% fetal bovine serum. Human HCT15 and HCT116 colorectal cancer cell lines were obtained from the American Type Culture Collection (ATTC) and were used between passages 20 and 40, while maintained

in

RPMI-1640

(Wellgene)

supplemented

with

10%

FBS

and

1%

penicillin/streptomycin. All cell lines were grown in a humidified incubator with 5 of carbon dioxide (CO2) and 95% air at 37 °C.


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In vitro cytotoxicity. The 3-[4,5-dimethylthiazol-2-yl]-2,5-diphenyl tetrazolium bromide (MTT) assay was utilized to assess the biocompatibility of PL end-capped PU nanocarriers. Viable cells were calculated by trypan blue staining (>90% cell viability for experiments) and seeded in 96-well plates at 10,000 cells/100 µL growth medium. Serial dilutions of DCPE−PU(SS)−DCPE, SPE−PU(SS)−SPE and DCPE−PU−DCPE were added to the plate (100 µL/well) at 37 °C in 5% CO2 for 48 h. The cells were then incubated with MTT stock solution (0.5 mg/mL in PBS; pH 7.4) at 37 °C for 4 h. Formazan granules generated cells were dissolved in 100 µL of DMSO. The absorbance of the solution at 562 nm was determined using a Power Wavex microplate spectrophotometer (Bio-Tek Instruments, Inc., Winooski, VT) after dilution to a linear range. Cellular uptake by confocal microscopy and Flow cytometry. Laser scanning confocal microscopy (Olympus FluoView FV1000) was used to observe both the intracellular uptake and distribution of Dox from Dox-loaded DCPE−PU(SS)−DCPE and DCPE−PU−DCPE nanocarriers. To determine the cellular uptake, Dox-loaded nanocarriers were incubated with HEK293FT human dermal fibroblast cells and HCT15 and HCT116 colorectal cancer cells overnight at a concentration of 25 µg/mL on a 12-mm circular glass coverslips with 10% DMEM in 24-well culture plates for 24 h. The cells in the cover slips were fixed with 4% paraformaldehyde for 10 min, then counterstained with DAPI, and fixed with DPX on clean glass slides. Slides were observed under a fluorescence confocal microscope (Nikon A1R, Nikon Instruments Inc.) and analyzed using NIS Elements software. The confocal microscopy settings were kept uniform between samples. Dox excitation and emission wavelengths were at 485 nm and 590 nm, respectively, whereas DAPI excitation and emission wavelengths were at 405 nm and 450 nm, respectively.


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Human dermal fibroblast, HEK293FT cells and HCT15 and HCT116 colorectal cancer cells (1×105 cells/well) were incubated in DMEM media (2 mL) with Dox-loaded nanocarriers for 1 h at 37 °C. The cells were washed with PBS (pH 7.4) solution, and incubated at 37 °C for another 12 h. After incubation, the cells were washed with cold PBS and harvested using trypsin-EDTA. The Dox fluorescence in the cell pellets was monitored by FACS Canto II (BD Bioscience, San Jose, CA) and FACSDiva software 6.1.3. In vitro cytotoxicity of Dox-loaded PL end-capped hybrid micelles. HEK293FT cells and HCT15 and HCT116 colorectal cancer cells (1×105 cells/well) were incubated with Dox-loaded nanocarriers in DMEM media (2 mL) for 60 minutes at 37 °C to evaluate anti-tumor activity of the micelles. Then the cells were washed with phosphate buffered saline (PBS) solution and incubated at 37 °C for another 12 h. After incubation, the cells were washed with cold PBS (pH 7.4) and harvested using trypsin-EDTA. After incubating the cells, the MTT solution was removed and the insoluble forzman crystals were dissolved in 100 µL DMSO. The absorbance was measured at the wavelength of nm. In vivo bio-distribution studies: Tumor-bearing nude mice were used for determining the dynamics and bio-distribution of NIR-dye-conjugated vesicles for tumor imaging. Initially, tumor-bearing nude mice were prepared by injecting a suspension of 1×108 HCT116 cells/mouse in PBS (pH 7.4, 200 µL) into the subcutaneous dorsa of athymic nude mice. After the tumor volume reached approximately 4–5 mm3, NIR-dye-loaded nanoparticles (DCPE-PU(SS)-DCPEIR820, 200 µL) were injected into the vein of the mouse and monitored for 72 h. The fluorescence image of the mouse before treatment was acquired and set as background. All NIR fluorescence images from the abdomen and the back were automatically acquired at 72 h after injections and displayed with the same fluorescence intensity scale after background subtraction


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at 782 nm using the Maestro 2 in vivo imaging system (Cambridge Research and Instruments, Inc., Woburn, MA, USA). Before starting experiments, mice were anesthetized by an intraperitoneal injection of 5 wt% chloral hydrate aqueous solution (7.5 mL/kg, anesthetic liquid/body weight). To confirm the bio-distribution of NIR-dye loaded nanoparticle in different organs, the mice were sacrificed 72 h injection. Various organs were harvested, rinsed in saline, and imaged for fluorescence. Additionally, to evaluate the specificity of micelles for normal or tumor tissues, NIR-dye-loaded nanoparticles were subcutaneously (s.c.) injected into a normal region and a tumor region on the backs of the mice. The fluorescence intensity of NIR-dyeconjugated vesicles within the normal and tumor regions was observed and compared at 782 nm using the Maestro 2 in vivo imaging system (Cambridge Research and Instruments, Inc., Woburn, MA, USA).

Results and Discussion Synthesis and characterization of phospholipid end-capped PU. In the past two decades, tremendous effort has been directed towards the preparation of various disulfide-bearing redoxresponsive nanocarriers end-capped with various polymers, such as polyethylene glycol, polycaprolactone, and other hydrophilic homopolymers, for the intracellular delivery of therapeutics in anticancer therapy.27,28 However, researchers continue to search for promising biodegradable and biocompatible carriers to improve cellular uptake and controlled delivery of therapeutics to pathological sites. In this study, phospholipid end-capped PU materials containing bioreducible disulfide units in the backbone were synthesized via a polycondensation of cystamine with a slight excess of IPDI, followed by end-capping with PL such as SPE or DCPE (Scheme 1). The control sample bearing no disulfide moieties were also synthesized by


12

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using HDA (instead of cystamine) and end-capped with DCPE. The structures of the resulting hybrid system were confirmed by NMR and FT-IR spectra.

Scheme 1. Synthesis route of phospholipid end-capped polyurethanes. Figure 1 shows the 1H NMR spectra of SPE and DCPE end-capped PL−PU(SS)−PL and DCPE−PU−DCPE (control). The resonance peaks at 7.93 ppm (a, a1, and a2) are ascribed to the –NH proton formed by the reaction of IPDI with SPE and DCPE. Multiple peaks found at 5.45 and 5.95 ppm (b, b1, and b2) were assigned to the -NH protons formed by the reaction of IPDI with cystamine and HDA. Usually, the −CH2SSCH2− resonance peak is found between 2 and 3 ppm. Unfortunately, it is not easy to differentiate PL−PU(SS)−PL from DCPE−PU−DCPE since methylene protons (−NH−CH2−CH2−) in both samples appear at the similar region. The chemicals shifts d, d1 and d2 represent the protons of −CH2−SS−CH2− in cystamine and –NH– CH2−C− in IPDI units and the chemical shifts e, e1 and e2 correspond to −NH−CH2− in cystamine linker and −NH−CH2− of HDA.


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Figure 1. 1H NMR spectra of (A) DCPE−PU(SS)−DCPE, (B) SPE−PU(SS)−SPE, and (C) DCPE−PU−DCPE (control) in DMSO as a solvent.

FT-IR spectra (Figure 2) confirm the incorporation of SPE and DCPE in both the NCOterminated PU and the DCPE end-capped control. A broad, stretched band was observed between 2200–2300 cm−1 in the FT-IR spectra of the NCO-terminated polyurethane urea; all the isocyanate groups were consumed during the reaction with the PL, including SPE and DCPE. Therefore, we couldn’t find any NCO stretching band was observed in the SPE or DCPE endcapped PL−PU−PL hybrid system after end-capping. P=O group peaks at 1180-1210 cm−1 were also clearly observed in the phospholipid incorporated PU series, but not in the NCO-terminated PU. The SPE−PU(SS)−SPE or DCPE−PU(SS)−DCPE sample can also be differentiated by comparing the peaks at 640-700 cm−1, representing –CS– stretching band. The disulfide bearing polymers clearly show this band at 661 cm−1, while DCPE−PU−DCPE show no peaks in this


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region. All the results collected by various analyses demonstrate the successful synthesis of the various hybrid polymeric materials. The Đ value of the hybrid materials were measured by GPC using polystyrene as the standard and DMF as the eluent. As summarized in Table 1, the number average MW (Mn) values range from 8000 to 10000 and the Đ values are around 1.4.

Figure

2.

FT-IR

spectra

of

(a)

OCN−PU(SS)−NCO,

(b)

SPE−PU(SS)−SPE,

(c)

DCPE−PU(SS)−DCPE, and (d) control (DCPE−PU−DCPE). Table 1. Compositions, size, and physical properties of PL−PU−PL-responsive nanocarriers. Feed ratio Mn Sample

a

PL/(SS or CC)/IPDI

a

b

Đ

Theor.

GPC

(mol/mol/mol)

(g/mol)

(g/mol)

DCPE−PU−DCPE

2/5/6

6800

9500

1.44

DCPE−PU(SS)−DCPE

2/5/6

6200

8900

1.41

SPE−PU(SS)−SPE

2/5/6

6300

9400

1.42

Average diameter determined by DLS. bPolydispersity index measured by GPC.


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Characterization of redox-responsive micelles: Recently, Ding et al. reviewed the potential of self-assembly of polyurethane/polyurea-based biodegradable nanocarriers as controlled drug delivery vehicles.37 Interestingly, PL or PL mimic moieties are influenced the self-assembly behavior of the polymer architecture.29−32,35 In our study, phospholipid end-capped control and disulfide-bearing hybrid micelles were prepared using a solvent exchange method. The size and shape of the micelles were studied by using DLS and TEM. The TEM images are shown in Figure 3. The control (DCPE−PU−DCPE) and disulfide-bearing SPE-PU(SS)−SPE and DCPE−PU(SS)−DCPE hybrid micelles were very similar in size (110±10 nm) and shape. The DLS results show the similar size distribution of micelles in the nanoscale range, as shown in the inset of each TEM images. Recently, Wang et al.35 reported the micellization of a lipid analogue phosphorylcholine end-capped PU, with polylactic acid as the hydrophobic core and a phosphoryl choline corona. They reported that the nanosize distribution of the micelles (≤100 nm) was influenced by the phosphorylcholine unit. Based on this report, we expected our novel hybrid micelles, bearing a hydrophobic cystamine and IPDI core and bound by PL corona, to form hybrid micelles with a diameter of ≤110 nm.

Figure 3. TEM images of hybrid micelles fabricated from (a) control, (b) SPE−PU(SS)−SPE, and (c) DCPE−PU(SS)−DCPE lipopolymers. Insets in (a) (b) and (c) are corresponding particle size distributions obtained by DLS measurements.


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Evaluation of reduction responsive behavior of the disulfide-bearing and control hybrid micelles was carried out by DLS measurements at different conditions (without GSH, 5 mM GSH, 10 mM GSH). As shown in Table 2, the aggregation behavior of the disulfide-bearing polyurethane micelles increased with increasing concentrations of GSH. Tunable cleavage and aggregation was observed in the disulfide-bearing micelles at 10 mM GSH. Similar degradation and aggregation scenarios have been reported by He et al.10 and Yu et al.12 in disulfide-bearing PEG/PCL end-caped micelles during intracellular delivery of hydrophobic drugs. However, the degradation and aggregation of PL−PU(SS)−PL micelles ensured due to the reducible nature of disulfide bearing back bone of the nanocarriers and the control (DCPE−PU−DCPE) micelles which doesn’t have disulfide linkage did not support any degradation at 10 mM GSH (Table 2).

Table 2. Size distribution of PL−PU−PL hybrid micelles fabricated by self-assembly before and after Dox encapsulation and the resulting drug loading content (DLC) and drug loading efficiency (DLE)

Sample

GSH (mM)

Blank Micelles

Dox-loaded micelles

Cumulant

Particle size

Cumulant

Particle size

diameter

distribution

diameter

distribution

DCPE−PU−

0

100±2

0.14±0.04

120±2

0.13±0.01

DCPE

5

108±4

0.21±0.03

128±6

0.25±0.02

10

115±5

0.25±0.05

135±5

0.31±0.05

DCPE−PU(SS)−

0

103±2

0.16±0.043

131±3

0.21±0.01

DCPE

5

124±5

0.25±0.02

140±5

0.32±0.07

10

153±9

0.35±0.09

175±6

0.43±0.09

SPE−PU(SS)−

0

112±5

0.15±0.04

135±3

0.19±0.03

SPE

5

130±5

0.28±0.09

147±5

0.34±0.03

10

152±6

0.39±0.03

186±6

0.49±0.06

DLC

DLE

(%)

(%)

8.2

37.0

9.3

39.3

8.9

38.5


17


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Page 18 of 47

Fabrication of Dox-loaded micelles and in vitro drug release. Significant research efforts have been put towards developing redox responsive nanocarriers for the controlled delivery of therapeutics, not only in vitro, but also in vivo.38 In the current study, we evaluated the drug loading behavior of PL end-capped hybrid micelles by encapsulating hydrophobic Dox, a potent drug for chemotherapy, into hybrid micelles through a solvent exchange dialysis method. In addition, we have encapsulated the same drug (Dox) into the control sample for reduction responsive comparative studies. Initially, we calculated the DLC and DLE of the DCPE−PU−DCPE micelles to be 8.2 % and 37 %, those of the DCPE−PU(SS)−DCPE to be 9.3 % and 39.3 %, and those of the SPE−PU(SS)−SPE micelles to be 8.9%, 38.5 %, respectively. To evaluate the drug release profile, Dox-loaded PL end-capped control and disulfidebearing hybrid micelles were incubated at 37 °C at various concentrations of GSH (from 0 to 20 mM GSH) in pH 7.4. The drug release profile was monitored using UV-visible spectroscopy, which covers the characteristic 484 nm absorption maximum of Dox in solution. The GSHresponsive cumulative release profile of Dox is plotted as a function of time in Figure 4. The significantly faster release of Dox was recorded from disulfide-bearing hybrid micelles at 10 mM GSH and above than from those with low GSH (Figure 4c). Additionally, we compared the significance of the disulfide linkage in the drug release behavior of DCPE−PU(SS)−DCPE with Dox-loaded control (DCPE−PU−DCPE) sample. For example, within 20 h, 42% of Dox was released at 10 mM GSH and 20 mM GSH and 85−90% Dox was released within 50 h from the Dox loaded DCPE−PU(SS)−DCPE. The Dox release was significantly low for the control nanocarriers with bearing less than 10 mM GSH. However, 20−30% of Dox was released from the Dox-loaded control (

Phospholipid End-Capped Bioreducible Polyurea Micelles as a

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