polyethylene glycol as bifunctional reinforcing

Cellulose (2017) 24:4461–4477 DOI 10.1007/s10570-017-1431-6

ORIGINAL PAPER

Cellulose nanocrystals/polyethylene glycol as bifunctional reinforcing/compatibilizing agents in poly(lactic acid) nanofibers for controlling long-term in vitro drug release Hou-Yong Yu

. Chuang Wang . Somia Yassin Hussain Abdalkarim

Received: 1 May 2017 / Accepted: 28 July 2017 / Published online: 3 August 2017 Ó Springer Science+Business Media B.V. 2017

Abstract Electrospun poly(lactic acid) (PLA) nanofibers would seem be poor carriers for drug delivery due totheir hydrophobicity, high crystallinity, weak mechanical strength and burst drug release due to poor compatibility with hydrophilic drugs. Thus, in this work, cellulose nanocrystals (CNC) as reinforcing agents and polyethylene glycol (PEG) as a compatibilizer to increase dispersion of both CNC and tetracycline hydrochloride (TH) in PLA matrix were successfully electrospun. The addition of CNC/PEG reduced fiber diameters, enhanced fiber uniformity, and the water contact angle was decreased from 117.3° for neat PLA to 98.0° for composite nanofibers with 10 wt% CNC/PEG. When drug-loading levels were further increased, the contact angle was decreased from 119.3° to 75.0°, and fiber diameter was decreased from 680 ± 34 nm for neat PLA with 3 wt% THs to 340 ± 17 nm for composite nanofibers with (TH 30%). More interestingly, more than 95.7% of drug contents were delivered within 1032 h, high drug loading efficiencies of composite nanofibers were more than 98%, and long-term sustained release behavior of composite nanofibers was obtained. The H.-Y. Yu (&) C. Wang S. Y. H. Abdalkarim The Key Laboratory of Advanced Textile Materials and Manufacturing Technology of Ministry of Education, College of Materials and Textiles, National Engineering Lab for Textile Fiber Materials and Processing Technology, Zhejiang Sci-Tech University, Hangzhou 310018, China e-mail: [email protected]

composite nanofibers showed good hydrophilicity and biocompatibility with MG-63 cells. Furthermore, compared to neat PLA, a 57.1% improvement in tensile strength and 240% increase in Young’s modulus were achieved for the composite nanofibers with (TH 15%). Additionally, the maximum decomposition temperature (Tmax) of the composite nanofibers with (TH 30%) was improved by 7.7 °C. The composite nanofibers with improved physical and hydrophilic properties, especially long-term drug release, showed good biocompatibility for use as a drug carrier for long-term sustained drug delivery systems and replacing the traditional medical dressings in biomedical applications. Keywords Polylactic acid Cellulose nanocrystals Polyethylene glycol Electrospinning Composite nanofibers Sustained release

Introduction The development of the electrospinning technique has proved to be a useful method for producing ultrafine fibers and nanofibers with high surface area, high porosity, and controllable degradation for drug release (cancer treatment), wound healing and tissue engineering (Cui et al. 2010; Hamman 2010; Li et al. 2007; Luo and Wang 2014; Perez et al. 2013). Of interest especially for controlled delivery systems, electrospinning synthetic polymers and biopolymers has the

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ability to vary fiber size, allowing the fibers to incorporate therapeutic drugs and reduce drug burst release for controlled drug delivery (Kenawy et al. 2002). However, electrospun synthetic polymers have shown poor biocompatibility during in vitro drug release (Billiet et al. 2012; Guo et al. 2013; Jiang et al. 2015; Rezwan et al. 2006; Rodrigues et al. 2015). Commercial polylactic acid (PLA), one of the ideal biopolymers, has been electrospun into PLA nanofibers with good biocompatibility and improved cell response in sustained drug release applications (Chou et al. 2015; Haddad et al. 2016; Xu et al. 2017; Yu et al. 2015). However, neat PLA nanofibers exhibit fewer polar groups (high hydrophobicity) and high crystallinity, which lead to low drug loading capacity of hydrophilic drugs, drug molecules located on the nanofiber surface and initial burst release of the drug. In addition, neat PLA nanofibers show poor mechanical strength and weak thermal stability, which result in structural instability during drug release (Li et al. 2012; Maretschek et al. 2008; Wei et al. 2015). In order to overcome these limits, the hydrophobic PLA nanofibers can be blended with hydrophilic nanofillers to control the drug release and improve some properties (Maretschek et al. 2008). Recently, cellulose nanocrystals (CNCs) with high mechanical strength (Young’s modulus of 150 GPa) and hydrophilic hydroxyl groups (Abdalkarim et al. 2017; French 2014), which are highly promising green nanofillers, have been used to tune the hydrophilicity, mechanical properties and drug release of electrospun PLA nanofibers (Jackson et al. 2011; Shi et al. 2012; Xiang et al. 2013). Shi et al. reported that by adding CNCs, the tensile stress and Young’s modulus of electrospun PLA were greatly increased, and the PLA degraded more rapidly in phosphate-buffered saline solution. However, high CNC content (10 wt%) would induce an obvious reduction in the tensile properties of electrospun PLA due to poor dispersion and compatibility of CNCs within the PLA matrix (Shi et al. 2012). Xiang et al. also found that CNCs could act as a nucleating agent to increase the PLA crystallization rate and the crystallinity of electrospun PLA nanofibers. Also, the addition of CNCs showed no significant effect on the fiber diameter, but improved degradation rate and drug release rate (Xiang et al. 2013). However, the increased CNC contents would induce a higher initial burst of the drug (Columbia Blue), because the strong interaction

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between CNCs and drugs resulted in a poor dispersion of drugs into the PLA matrix. Therefore, there is a big challenge in simultaneous compatibility among the CNCs, PLA, and drugs in the composite nanofibers. Polyethylene glycol (PEG) is a good compatibilizer in PLA, and has been used to develop functional composite nanofibers for effective drug delivery (Gui et al. 2012). Ten et al. (2010) used 30% PEG as a compatibilizer to efficiently improve dispersion of CNCs into a PHBV matrix; the resulting composite films showed an obvious improvement in mechanical properties. Moreover, the hydroxyl groups of PEG can simultaneously interact with CNCs and PHBV, and hydrophilic CNC has good affinity to TH drugs. Herein, we report our study on the incorporation of CNC/PEG as bifunctional reinforcing materials into PLA nanofibers to fabricate a serious of PLA/CNC/ PEG composite nanofibers by using electrospinning, in which the biodegradable CNCs as organic nanofillers and PEG as compatibilizer can increase the dispersion of CNCs in a polymeric matrix. We evaluated the morphology, microstructure, and hydrophilicity of the resulting PLA nanofibers. Furthermore, we investigated the effects of various drug loadings on morphology as well as hydrophilicity, mechanical and thermal properties, in vitro degradation, drug release behavior, and cytocompatibility of the composite nanofibers. As expected, the incorporation of CNC/PEG could improve the attachment of the cells on PLA/CNC/PEG, and open the way for a new strategy for drug carriers.

Experimental section Materials All reagents were analytical grade and used as received without further purification. Microcrystalline cellulose (MCC, about 10 lm of size) was purchased from Sinopharm Chemical Reagent Co, Ltd. Poly (lactic acid) (PLA, Mn = 1.0 9 105) was supplied by Bright China Industrial Co. Ltd (Shenzhen, China). Polyethylene glycol-400 (PEG) was purchased from Guoyao Group Chemical Reagent Co, Ltd. Sulfuric acid was obtained from Guoyao Group Chemical Reagent Co, Ltd. Chloroform, methanol, acetone, N,N-dimethylformamide (DMF), disodium hydrogen (Na2HPO412H2O) and potassium dihydrogen

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phosphate (KH2PO4) were purchased from Guoyao Group Chemical Reagent Co, Ltd. KH2PO4 and Na2HPO412H2O were used without further purification to prepare 0.1 M phosphate buffer solutions (PBS) of pH 7.4. Human MG-63 cells (an osteosarcoma cell line) were purchased from the Shanghai Institute of Biochemistry and Cell Biology. Fetal bovine serum (FBS) and Dulbecco’s modified Eagle’s medium (DMEM) were obtained from Gibco (USA). The Cell Counting Kit-8 was purchased from Beyotime Institute of Biotechnology (China). The ActinTracker Green was provided from Biyuntian Biology Technology Co, Ltd. All other chemicals were of analytical grade and purchased from Hangzhou Mike Chemical Agents Company (China). Tetracycline hydrochloride (TH) was purchased from Sigma and used as received.

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on the PLA. For the preparation of composite nanofibers, the specific amount of CNC was 5 wt% based on the PLA. The final solution was stirred for at least 24 h to form a homogeneous solution. The mass ratio of tetracycline hydrochloride (TH) to PLA/CNC/ PEG composites was 3, 10, 15, 25 and 30 wt%. For comparison, 3 wt% TH was added into the PLA solution. The electrospinning process for the preparation of (drug-loaded) neat PLA and composite nanofibers with different drug loadings was carried out according to the following conditions: applied voltage was 18 kV, the temperature was kept at room temperature, the feeding rate was 1 mL/h, the tip-tocollector distance was 15 cm, a rotating mode was utilized and the rotating speed of the collector was 220 rpm, and the collecting time was 8 h. Characterization

Extraction of cellulose nanocrystals Preparation of CNCs was carried out according to the previous work (Yu et al. 2012). Briefly, the MCC was first hydrolyzed with 64 wt% sulfuric acid solution at 50 °C for 1 h under strong mechanical stirring. After cooling to room temperature, the resultant suspension was repeatedly washed by successive centrifugations (12,000 rpm at 10 °C for 20 min) with deionized water until approximate neutrality was attained. Finally, the sample was freeze-dried for 48 h to obtain dry CNCs. The obtained CNCs with rod-shaped morphology were about 20 nm in diameter and 230 nm in length, and the aqueous suspension before freeze-drying was about 0.1 wt% with the amount of sulfate half-ester groups (around 50 mmol/kg) determined by using oxime reaction and conductivity titration (Yu et al. 2016). Preparation of electrospinning solution Composite electrospinning solution 10 wt% PLA solution in 9:1 (w/w) chloroform/DMF co-solvent was stirred at 60 °C for 30 min to make certain that the molecular chains of PLA were well entangled. The asprepared solutions were continuously stirred for 24 h until clear solutions were obtained. The weight ratio of PEG to CNCs powder was 1:1. The PLA solutions were mixed with a specific amount of mixed CNC/ PEG (1–10 wt% based on PLA). For the control samples the specific amount of PEG was 5 wt% based

The morphology of neat PLA and composite nanofibers with different drug loadings was observed on a field emission scanning electron microscope (FESEM, HITACHIS-4800) at 2.0 kV. The elemental analysis of the prepared composite nanofiber was conducted by an energy-dispersive X-ray spectrometer (EDX). The surface roughness of PLA and composite nanofibers were measured by a Park System XE-100E atomic force microscope (AFM). FT-IR spectra were recorded on a Nicolet IS50 FTIR spectrophotometer, operating at 64 scans and 4 cm-1 resolution in the region between 4000 and 650 cm-1 at room temperature. Pellets of dried PLA and composite nanofibers were made with KBr. Non-isothermal crystallization and melting curves were determined using a DSC Q20 differential scanning calorimeter from TA instruments. Around 5–10 mg of the sample was placed in a DSC cell in a glove box. Each sample was first heated at room temperature to 200 °C at a rate of 20 °C min-1 and maintained at 200 °C for 3 min to remove previous thermal history. The sample was then cooled to 0 °C at a rate of 10 °C min-1 and heated again to 200 °C at a rate of 10 min-1. The thermal stability of the samples (5–10 mg) was tested in the range of 30–600 °C (nitrogen atmosphere) at a rate of 20 °C min-1 by using a Netzsch TG209 F1 thermo-gravimetric analyzer (TGA). The tensile properties of the samples with rectangles 80 9 20 mm2 and thickness of 50–60 lm were

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measured on mechanical testing equipment (Instron 2712-018) with a crosshead speed of 1 mm min-1. The tests were conducted at a temperature of 20 °C and relative humidity of 65%, after equilibration under a controlled environment for at least 24 h. Three replicates were tested for each sample. The contact angles were measured on the air surface of PLA and composite nanofibers using a pendant drop method on a Data physics OCA40 contact angle analyzer at room temperature. About 2 lL of deionized water was dropped onto the surface at a contact time of 5 s. Twenty independent determinations at different sites of the sample were averaged. The in vitro degradation was tested according to the previous method. The porous drug-loaded nanofibers were cut into rectangles 2 9 5 cm2 and 50–60 lm in thickness. Each specimen was placed in a test tube containing 20 mL of phosphate-buffered saline (PBS, pH = 7.4) and incubated at 37 °C for a period of 35 days. The samples were removed every 5 days, washed with deionized water and then dried. The weight of the dried samples was recorded, and the percentage weight loss was calculated using the following formula. Wight loss ð%Þ ¼

m0 mt 100 m0

ð1Þ

where m0 is the initial mass, mt is the mass after a given time of hydrolysis. Drug loading and drug loading efficiency were determined by calculating the drug weight from the absorbance data. In brief, various amounts of THs (0.0625, 0.125, 0.25, 0.5, and 1 mg mL-1) were dissolved in a PBS buffer solution, and the absorbance was recorded by a TU-1901 spectrophotometer (Beijing Purkinje General Instrument). A standard curve with a linear correlation (R2 = 0.99988) was obtained. Neat PLA (100 mg) and composite nanofibers with various drug loadings were dissolving in chloroform (15 mL). To extract the TH in the samples, 32 mL 0.1 M PBS was added into the solution, and then solvents were removed by an air pump. The drug was completely dissolved by sonication and centrifugation at a speed of 10,000 r min-1 for 20 min. The resultant TH solutions in PBS were analyzed spectrophotometrically (at 360 nm) and the drug weight was calculated from the absorbance data according to the standard curve. Accordingly, the drug loading and drug loading efficiency can be calculated as: drug loading

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(%) = (weight of drug measured)/(weight of drug unloaded nanofibers) 9 100; drug loading efficiency (%) = (weight of drug measured)/(weight of drug added) 9 100. In vitro drug release profiles of the composite nanofibers were evaluated in PBS (pH 7.4). Briefly, 100 mg of composite nanofibers were sealed in a dialysis bag. The dialysis bag was immersed in 32 mL of PBS and maintained at 37 °C. At designated time intervals, 4 mL of the buffer solution was withdrawn from the samples, and an equal volume of fresh buffer was added to the flasks. The amount of TH drug in the buffer solution was determined by measuring the UV absorbance at 360 nm using a TU-1901 spectrophotometer. The release of TH was determined by monitoring the absorbance at 360 nm as a function of time. The results of triplicate tests were used to calculate the cumulative release percentage of drugloaded composite nanofibers according to the drug release standard curve equation. The human MG-63 cells, an osteosarcoma cell line, were maintained in Dulbecco’s modified Eagle’s medium (DMEM) with 10% fetal calf serum (FCS) and antibiotics (100 U mL-1 penicillin, 100 U mL-1 streptomycin) at 37 °C and in a 5% CO2 atmosphere. MG-63 cells were digested using 0.25% trypsin for about 1 min and resuspended in the medium. Cell numbers were determined by counting via a hemocytometer and then diluted to the concentration of 1 9 105 cells mL-1. All the composite nanofibers were disinfected with 75% ethanol and washed with PBS solution, and then placed into 24-well cell culture plates, followed by adding 1 mL cell suspension to each well. After culturing MG-63 cells for 2 h, the culture medium with unattached cells was carefully removed from the wells by a pipette. The samples were then washed with PBS buffer. Subsequently, the MG-63 cells were identified by Actin-Tracker Green, which allowed the measurement of the cytoplasmic area and consequent estimation of cellular spreading. Fluorescence images were photographed by a fluorescence microscope (Olympus IX71-22FL/PH). The cell initial attachment was evaluated after cultivation for 2 h. The cell culture medium was discarded and the samples were washed with PBS buffer three times. The attached cells were stained with Actin-Tracker Green and observed qualitatively using fluorescence microscopy. A mean value for three places of each sample was determined. This test was carried out as

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described elsewhere (Yu et al. 2014; Zhang et al. 2017). The porosity (e) of electrospun composite nanofiber was calculated according to our previous work (Abdalkarim et al. 2017). The true volume was measured by the pycnometric method, and the bulk density (qbulk) was calculated by a gravimetric method. The porosity was calculated according to the following equation: qbulk eð%Þ ¼ 1 100% ð2Þ qtrue where qtrue is the true density of the nanofibers. Results and discussion Morphology and hydrophilicity The morphology of nanofibers was investigated by using a field emission scanning electron microscope. Typical morphologies of neat PLA and the composites nanofibers reinforced with various CNC/PEG contents produced under optimum conditions are shown in Fig. 1a, d–f. The nanofiber surfaces appeared similarly smooth compared with those for PLA nanofibers showing no evidence of CNC/PEG agglomerates on the fiber surfaces. The diameter histograms of the neat PLA and composite nanofibers and the average diameter of each fiber are shown in Fig. 1g. The diameter of PLA nanofiber was about 2.5 ± 0.12 lm. With the addition of 10 wt% CNC/PEG, the average diameters of composite nanofibers decreased to 1.2 ± 0.1 lm. Nevertheless, in the case of the composite nanofiber of PLA/CNC (5 wt%) and PLA/PEG (5 wt%), the average diameters of composite nanofibers decreased to 0.7 ± 0.04 and 0.9 ± 0.8 lm, respectively (Fig. 2b, c). The reduction in diameter could be because of the weakening effects of CNC and PEG on the PLA chain entanglement and inter- or intramolecular interactions (Huan et al. 2016). In addition, the increase in charge density was due to changes in conductivity of the electrospinning solution, and the electric field increased during the electrospinning process. The enhanced conductivity of the electrospinning solution could lead to a reduction in fiber diameter (Lu and Hsieh 2010). Moreover, the composite nanofibers, including neat PLA nanofibers, demonstrated porosities in the range

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of 60.7–74.8%. The porosity increased with the addition of CNC/PEG contents (Fig. 1h), indicating that the change in the fiber diameter could cause a great difference in the microstructure and hydrophilicity properties. The hydrophilicity behavior of neat PLA and composite nanofibers with various CNC/PEG concentrations was investigated by using contact angle measurements. Incorporation of CNC/PEG into PLA nanofibers, the water contact angle was significantly decreased from 117.3° for neat PLA to 98.0° with 10 wt% CNC/PEG content as shown in Table 1. However, the water contact angle was decreased from 117.3° for neat PLA to 94.7° and 97.3° for PLA/CNC (5 wt%) and PLA/PEG (5 wt%), respectively. An improved hydrophilicity was possibly due to the mainly hydrophilic nature of the PEG structure molecule chain by hydrogen bonding interactions with the presence of water, and thus improved water wettability. In addition, the composite nanofibers were composed of CNCs with more hydrophilic groups, which were beneficial to absorb water molecules to obtain enhanced hydrophilicity of composite nanofibers (Narain et al. 2002; Zhang et al. 2015). Also, the reduced surface roughness and increased porosity values (Fig. 1h) of composite nanofibers caused lower contact angles. Morphology and hydrophilicity Tetracycline hydrochloride (TH) was selected as a model drug due to interest in the treatment of inflammation diseases. TH is insoluble in chloroform, and thus it was solubilized in a small amount of methanol (0.05 g/mL) and then added to the composite solution (CNC/PEG, 10 wt%). The resulting solutions were yellow but clear, as shown in Fig. 2, indicating good dispersion of TH in the PLA/CNC/ PEG composite solution. Figure 3 shows the SEM morphologies of neat PLA and composite nanofibers with various drug loading. It was found when both tetracycline hydrochloride (TH) and CNC/PEG were successfully incorporated in PLA nanofibers, that the composite nanofibers were becoming homogeneous in the nanoscale range. The corresponding diameter frequency distribution is illustrated in Fig. 3g. With the increase of the drug-loaded contents, the fiber diameter was decreased from 680 ± 34 nm for neat PLA with (TH 3%) to

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Fig. 1 FE-SEM images and size distributions for neat PLA (a) and PLA/CNC (5 wt%) (b), PLA/PEG (5 wt%) (c), and composite nanofibers with the CNC/PEG contents of 1 (d), 5 (e), and 10 wt% (f), average fiber diameter (g), contact angles values and porosity (h)

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spinning matrix could enhance the spinnability of the drug and the quality of the resultant nanofiber. Nevertheless, the use of the cationic drug (TH) further increased the solution conductivity and the electric field during the electrospinning process, and then reduced the fiber diameter (Haroosh and Dong 2013). The EDX analysis confirmed the deposit of tetracycline hydrochloride (TH) as a hydrochloride salt on the surface and the inside of the PLA and PLA/CNC/PEG nanofibers, which could be supported by appearances of the signals of C, O, and Cl in Fig. 3a, b, d, e and the insert EDX images of the fractured surface. Generally, during the process of nanofiber formation, the drugs were easily excluded from the inner crystalline region of PLA and composite nanofibers, leading to the deposit of TH on the nanofiber surface. Polylactic acid (PLA) has been used for tissue engineering, because of its excellent biocompatibility, biodegradability, and non-toxicity. By modulating surface properties of PLA, the hydrophilicity for PLA composite nanofibers might directly improve in vitro/vivo biological performance such as tissue response, cell attachment, and protein adsorption (Armentano et al. 2013; Guo et al. 2013). The hydrophilicity of neat PLA nanofiber and composite nanofibers with various drug loadings was revealed by water contact angle measurements. As shown in Table 1, the contact angle of neat PLA nanofibers was hydrophobic, with a contact angle of 119.3°. This hydrophobicity was due to surface roughness. Indeed,

340 ± 17 nm with composite nanofibers with (TH 30%). The average fiber diameter of PLA/CNC/PEG nanofibers with various TH contents was found to decrease and the distribution of diameters became broadened after addition of the drug into the PLA/ CNC/PEG nanofiber. It is obvious that all nanofibers, including neat PLA with the addition of 3–25% TH, showed a bimodal size distribution due to the uneven distribution of the charges with a high viscosity of electrospinning solution, which resulted in uneven of the splitting of the jet. However, for PLA/CNC/PEG with 30% TH, the fibers exhibited a unimodal distribution with a mean diameter of 340 ± 17 nm, and this phenomenon indicated that the PLA as a co-

Fig. 2 Drug-loaded electrospinning solution of (a) PLA (TH 3%), PLA/CNC/PEG (TH 3%) (b), PLA/CNC/PEG (TH 10%) (c), PLA/CNC/PEG (TH 15%) (d), PLA/CNC/PEG (TH 25%) (e), and PLA/CNC/PEG (TH 30%) (f)

Table 1 Drug loading, drug loading efficiency, contact angle, porosity and surface roughness of drug-loaded PLA and composite nanofibers with various drug loadings Sample

FaH–CO

PLA (TH 3%)

–

PLA/CNC/PEG (TH 3%)

Drug loading efficiency (%)

Contact angle (°)

Porosity (%)

Surface roughness (nm)

2.91

97.0

119.3

62.5

16.7

0.39 ± 0.05

2.95

98.3

117.5

65.8

14.8

PLA/CNC/PEG (TH 10%) PLA/CNC/PEG (TH 15%)

0.40 ± 0.06

9.90

99.0

115.2

69.3

9.5

0.42 ± 0.03

14.82

98.8

109.7

74.8

6.3

PLA/CNC/PEG (TH 25%)

0.42 ± 0.07

24.96

99.8

96.3

77.2

6.1


0.45 ± 0.08

29.85

99.5

75.0

83.5

6.0

a

Experimental drug loading (%)

FH–CO was obtained from de-convoluted FT-IR spectra (Yu et al. 2016)

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Fig. 3 FE-SEM images for drug-loaded PLA (a, 3% TH) and PLA/CNC/PEG composite nanofiber with various drug loadings: 3% TH (b), 10% TH (c), 15% TH (d), 25% TH (e), and 30% TH (f), and size distributions (g), Insert is EDX spectrum

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surface roughness played a very important role in hydrophilicity (Yoon et al. 2008). As already mentioned above, by addition of hydrophilic groups on the CNCs and PEG, the hydrophilicity was improved due to the hydrophilic nature of the PEG structure and also more hydrophilic groups of CNCs. However, when drug-loadings furthered increased, the contact angle was decreased from 119.3° to 75°, illustrating an enhanced hydrophilicity of the composite nanofibers. This indicates that the drugs played the main role in hydrophilicity of composite nanofibers, because TH as a hydrochloride salt with high charge density could increase the conductivity of the electrospinning solution to reduce the diameter, and then enhance the porosity of the composite nanofibers. Also, the water contact angle of neat PLA without the drug was 117.3°, but it was slightly increased to 119.3° for PLA (TH 3%). This behavior might be attributed to higher surface roughness (16.7 nm), small diameter and weak interaction between the drug and the polymer matrix during the electrospinning process (Haroosh and Dong 2013). In general, high surface roughness can induce an increase of the water contact angle (Yoon et al. 2008). Compared with neat PLA, the composite nanofibers with various drugs loaded showed a decrease in surface roughness (Table 1), which also contributed to smaller contact angle value. Chemical structure The FT-IR spectra of the drug tetracycline hydrochloride (TH), CNCs, neat PLA and composite nanofibers with various drug loadings are shown in Fig. 4. The CNC showed characteristic bands in the region of 3700–3000 cm-1 corresponding to O–H stretching vibrations. In the range of 2800 and 3000 cm-1 to stretching vibrations of C–H of methylene groups, the peak at 1061 cm-1 we assign to asymmetric C–O–C bond stretching from the pyranose ring, and 1145 cm-1 to C–O–C glycosidic linkage (Cui et al. 2006; Xiang et al. 2013). The characteristic bands of THs in the region between 1200–1800 cm-1, the peaks at 1682 and 1516 cm-1 were attributed to the carbonyl and amino groups of the amide in ring A, respectively, while the bands at 1617 and 1591 cm-1 were assigned to the carbonyl groups in the A and C rings (see insert of Fig. 4a, He et al. 2009). The FT-IR spectra for all composite nanofibers were very similar to each other, even with various drug loadings. The peak at

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3000–2980 cm-1 corresponded to C-H stretching vibration, and the peaks in the range of 1100–1370, such as 1088, 1130, and 1184, 1270 and 1383 cm-1 , corresponded to C–O–C stretching vibrations. Furthermore, the peaks at 1455 and 868 cm-1 were related to – CH3 bending vibration and the C–C bending vibration of PLA, respectively. More importantly, the peak at 1617 cm-1 assigning to the carbonyl groups in C rings from TH (insert of Fig. 4a) was found, indicating that drugs were successfully encapsulated in the PLA composite nanofiber. In Fig. 4b, strong peaks at 1754 and 1775 cm-1 were ascribed to the ordered and amorphous regions (free C=O groups); the former peak was easily affected by CNC loading levels and the interactions between the CNCs and PLA (Liao et al. 2011; Yu et al. 2012), and thus was denoted as hydrogen bonded C=O groups. With the incorporation of CNC-PEG, the band position for hydrogen-bonded components shifted slowly from 1754 cm-1 for neat PLA to 1762 cm-1 for the composite nanofibers with 10 wt% CNC/PEG, whereas that for the free components remained almost the same. The combined Gaussian/Lorentzian spectral function was investigated to curve-fit the FT-IR spectra of the carbonyl band in the range from 1850 and 1650 cm-1 to calculate the hydrogen bond fraction (FH–CO) according to our previously reported equation (Yu et al. 2012; Yu and Yao. 2016). The FH–CO of the composite nanofibers are listed in Table 1. The FH–CO value for all composite nanofibers was gradually increased from 0.39 to 0.45. It shows that the number of hydrogen bonds in the composite nanofibers was increased by the addition of more TH drugs. The composite nanofibers with 30 wt% THs had the highest amounts of intermolecular hydrogen bonds due to efficient dispersion CNC/PEG and TH drugs within the polymer matrix. Drug loading and drug loading efficiency Previous studies have shown that surface area, pore size, and morphology play important roles to improve the drug loading efficiency of hydrophobic materials (Armentano et al. 2013). In this study, neat PLA and composite nanofibers with 10 wt% CNC/PEG content (with small fiber diameter and improved hydrophilicity, as a model sample) were selected to evaluate the viability as long-term sustained release drug delivery systems. Drug loading efficiency and drug loading of

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Fig. 4 FT-IR spectra (a, b) carbonyl stretching region (vC=O) in the infrared spectra for the composite nanofibers with various drug loadings and (c) peak de-convolution for the composite nanofibers with (TH 15%)

neat PLA and the composite nanofibers are summarized in Table 1. It was found that compared with neat PLA, the drug loading efficiencies of composite nanofibers were more than 98%. This can be due to the interaction between polar groups of TH drug and polar hydroxyl groups of CNC/PEG, and thus efficiencies of composite nanofibers were higher than neat PLA. Crystallization behavior and thermal stability Differential scanning calorimetry (DSC) was used to study the crystallization behavior of neat PLA and composite nanofibers, as illustrated in Fig. 5a, b. Neat PLA and all composite nanofibers showed no obvious

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melt-crystallization peak during the first cooling scan. In contrast, in the second heating scan, the cold crystallization peaks appeared for neat PLA and composite nanofibers with various drug loadings (Fig. 5b). Compared to neat PLA without the drug, cold crystallization temperature (Tcc) of the PLA nanofibers with (TH 3%) was decreased from 108.6 to 100.5 °C. Meanwhile, the glass transition temperature (Tg) was reduced from 66.5 to 64.6 °C (Table 2). It indicates that TH can act as a plasticizer to prompt mobility of crystalline PLA chains, leading to lower Tcc and Tg. A similar phenomenon was observed in drug-loaded composite nanofibers by other researchers (Cui et al. 2006; Zong et al. 2002). The incorporation of 5 wt% CNC or 5 wt% PEG can also induce

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Fig. 5 Non-isothermal DSC first cooling (a), and second heating curves (b), TGA (c), and the thermal parameter (d) of neat PLA and composite nanofibers with various drug loadings

obvious reductions in Tcc and Tg values, compared with neat PLA nanofibers. It showed that both CNC and PEG can prompt the chain mobility for PLA. Moreover, the Tg values for composite nanofibers, except for the composite nanofiber with high TH contents of 25 and 30%, were lower than that of PLA/ CNC/PEG (10 wt%). This indicates that the drug TH showed a plasticizing effect on PLA crystallization (Table 2). The low molecular weight of TH would decrease packing density and crystallization ability of crystalline PLA chains, leading to lower Tg values of the drug-loaded composite nanofibers (Haroosh and Dong 2013). However, the addition of too much TH might hinder the mobility of crystalline PLA chains, leading to increasing Tg and Tcc values, like PLA/ CNC/PEG (TH 25%). With the incorporation of

10 wt% CNC/PEG, the Tcc of the PLA nanofibers was reduced greatly from 108.6 to 80.3 °C, indicating that PLA crystallization became easier due to the heterogeneous nucleation effect of CNCs (Cacciotti et al. 2014; He et al. 2009). At the same loading of CNC/ PEG, with the increase of drug-loaded contents from (TH 3–30%), the cold crystallization temperature was slightly increased to 86.2–91.3 °C, but still lower than that of PLA nanofibers. This suggests that the PLA crystallization was becoming easy because of the effect of both TH and CNC/PEG, thus the composite nanofibers could undergo enough crystallization in the first cooling scan, leading to an obvious reduction in cold crystallization peak for the PLA composite nanofibers. Compared with drug-loaded PLA nanofibers, the melting peak (Tm) value for the

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Table 2 Thermal pentameters for PLA nanofibers, drug-loaded PLA and composite nanofibers with various drug loadings Sample

Tg (°C)

Tcc (°C)

DHcc (J/g)

Tm (°C)

DHm (J/g)

Xc (%)a

PLA

66.5

108.6

17.4

149.6

29.1

31.3

PLA/CNC (5 wt%)

64.2

96.3

18.1

149.6

29.6

33.5

PLA/PEG (5 wt%) PLA/CNC/PEG (10 wt%)

60.1 58.4

87.4 80.3

18.7 19.0

149.1 150.4

30.1 30.5

34.0 36.4

PLA (TH 3%)

64.6

100.5

3.94

152.4

28.9

32.0

PLA/CNC/PEG (TH 3%)

54.4

86.2

3.91

151.2

30.9

38.1


54.1

87.6

3.68

151.4

30.2

40.1


57.2

86.2

3.33

151.8

26.1

36.7


61.5

151.2

25.6

40.8


–

–

–

151.6

24.5

41.8

CNC

–

–

–

–

–

84.3

a

Xc = DHm/[DHm,100 (Cacciotti et al. 2014)

%

91.3

PLA 9 (1 - (wCNC ? wPEG ? wTH)], where DHm is melting enthalpies, DHm,100

composite nanofibers was slightly decreased. Table 2 illustrates that the degree of crystallinity (Xc) was increased gradually from 31.3% for neat PLA to 33.5% for PLA/CNC (5 wt%), 34.0% of PLA/PEG (5 wt%) and 36.4% for PLA/CNC/PEG (10 wt%). Meanwhile, compared with neat PLA (TH 3%) the degree of crystallinity (Xc) of composite nanofiber with various drug loading shifted gradually from 32.0 to 41.8% for composite nanofibers with (TH 30%). This demonstrates that the introduction of more TH drugs was helpful in forming imperfect PLA crystals and then improving crystallinity in the composite nanofibers. The thermal stability of the drug-loaded PLA and composite nanofibers was investigated by using TGA (Fig. 5c). The main thermal parameters, including thermal degradation onset temperature (T0), maximum decomposition temperature (Tmax) and complete decomposition temperature (Tf), are illustrated in Fig. 5d. With the increase of TH drug contents, Tmax increased slowly from 355.4 °C for neat PLA with (TH 3%) to the maximum value of 368.9 °C for composite nanofibers with (TH 30%). A similar trend can also be seen for T0 and Tf. Compared to neat PLA nanofiber, the T0, Tmax and Tf values of composite nanofibers loaded with 30 wt% THs were improved by 7.3, 13.5 and 7.6 °C, respectively. This may be due to the formation of more hydrogen bonds between the nanofillers-matrix and increased crystallinity in composite nanofibers (Yu et al. 2012; Yu and Yao. 2016; Zhang et al. 2017).

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3.21

%

PLA = 93 J g-1

Mechanical properties It is important to study the mechanical properties of PLA composite nanofibers for uses in tissue engineering applications, because these materials are required to have a certain strength, flexibility, and toughness. The mean mechanical parameters are summarized in Fig. 6a, b. It was found that with the increase of CNC/ PEG and drug-loading contents, both the tensile strength and Young’s modulus gradually increased. Compared with neat PLA nanofibers, 57.1% improvement in tensile strength, 240% increase in Young’s modulus, and 24.4% reduction in elongation at break were obtained for the composite nanofiber with (TH 15%). This result can be explained by the smaller fiber diameter of composite nanofibers due to the strong interaction between CNCs and the PLA matrix. Indeed, hydrogen bonding interactions led to a reorientation of the molecules to the intermolecular chain, resulting in increased rigidity of composite nanofibers and thus enhancement in the mechanical strength of the composite nanofibers. In addition, TH drugs with low molecular weight acting as a compatibilizer may improve chain mobility and thus the orientation of crystalline PLA chains (Haroosh and Dong 2013). It is well known that the mechanical properties of the composite nanofibers were dependent on many factors, including nanofiber structure, the interaction between nanofillers and PLA matrix, and the density (porosity) of the electrospun mats (Xu et al. 2014). Hence, the porosities of composite nanofibers have

Cellulose (2017) 24:4461–4477

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Fig. 6 Tensile strength, Young’s modulus (a), and elongation at break (b) as a function of drug loadings for neat PLA and composite nanofibers

been found in the range of 62.5–83.5% as listed in Table 1, based on the densities of the nanofibers for the fiber diameter of (680–340 nm) for neat PLA with (TH 3%) and composites nanofibers with various drug-loadings, respectively. The average fiber diameter decreased with increasing TH content and their porosity increased, indicating that the change in the fiber diameter could cause large variation in the mechanical properties. Upon further increasing of the drug loading contents (TH 25–30 wt%), decreases in tensile strength and Young’s modulus of the PLA/ CNC/PEG composite nanofiber with (TH 25–30%) were observed. This phenomenon may be due to the reduced frictional forces between nanofibers caused by loading more drugs. Also, the imperfect PLA crystals could weaken the mechanical strength. In vitro degradation behavior, drug release behavior and cytocompatibility Biodegradation behavior of neat PLA (TH 3 wt%) and the composite nanofibers with various drug loadings are shown in Fig. 7a. All samples were immersed in PBS for up to 35 days. The weight loss of all samples was increased with the immersion period of 35 days. The contact angle results showed that the incorporation of hydrophilic CNC/PEG into PLA nanofibers improved the surface hydrophilicity of PLA composite nanofibers, and thus could easily speed up their degradation rate. A degradation rate of about 52% weight loss could be obtained for the composite nanofibers with 30 wt% TH after 35 days.

Additionally, as already mentioned above, the drugs played the main role in hydrophilicity of the composite nanofibers, because TH as a hydrochloride salt with high charge density can increase the conductivity of electrospinning solution to reduce the diameter, and then enhance the porosity. Therefore, more TH drugs can increase the amounts of reactive groups for composite nanofibers to improve degradation ability, because the composite nanofibers can directly contact the degradation medium. By contrast, the degradation rate of neat PLA nanofibers was very slow, only 18% weight loss of neat PLA nanofibers occurred after 35 days, which was due to the higher crystallinity of PLA and the hydrophobic nature of PLA chains. The water contact angle of neat PLA nanofiber (TH 3%) was slightly larger than those of composite nanofibers with various drug loadings; as a result, a slow degradation rate of neat PLA was observed (Kouhi et al. 2015; Sebe et al. 2015). Figure 7b illustrates drug release behavior of neat PLA nanofiber and composite nanofibers with different drug loadings. The release profiles of all nanofibers with various drug loadings were categorized by two typical phases. An initial burst release or steady release phase was observed. During the first 120 h, the PLA nanofibers gave about 12% drug release, but only 13% drug release during the subsequent 1032 h. This result indicated that minimum stained release and a larger initial burst of PLA (TH 3 wt%) were observed, comparing to composite nanofibers with (TH 3–30%). Furthermore, during the later 1032 h, the composite nanofibers with 15, 25, and 30 wt% TH delivered

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Cellulose (2017) 24:4461–4477

Fig. 7 In vitro degradation behavior (a), accumulative drug release of PLA nanofiber and composite nanofibers with (TH 15%) drug loading (b), and fluorescence microscopy images of neat PLA (TH 3%) and composite nanofiber with 15 wt% THs (c)

more than 80% of their drug contents, but the composite nanofibers with the drug at low loading content (TH 3 and 10%) delivered only 16 and 27.5%, respectively. In the same time frame, neat PLA nanofiber delivered only about 13% of its drug content. This result indicates that the limited drug release of neat PLA and composite nanofibers with low drug loadings during the later period was due to fiber size and crystallinity. It is easy to understand that the crystalline region can hinder the diffusion of the aqueous environment into the polymer inner layers, and thus hinder the diffusion of the TH drugs from the nanofibers (Kenawy et al. 2002; Sebe et al. 2015). With the increase of TH drug contents from (15–30%), the accumulated drug release was boosted from 77.8 to 95.7%. From the observations above, this hints that

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drug release behavior of typical composite nanofibers can be controlled by drug loading amounts, fiber size, porosity and hydrophilicity of nanofiber or added drugs. The smaller fiber size (diameter) was beneficial for increasing drug release amounts of the composite nanofibers. As already mentioned above, when drugloading further increased, the contact angle was decreased from 119.3° to 75°; as a result, an enhanced hydrophilicity of the composite nanofibers with smaller fiber size was found. In comparison with low drug content, a significant increase in the accumulated drug release of the composite nanofibers with high drug content was ascribed to the unique structure and properties of such composite nanofibers systems with higher porosity, and the possibility of controlling their hydrophilicity by introducing CNC/PEG and TH, to

Cellulose (2017) 24:4461–4477

make liquid enter into the side of PLA nanofibers for sufficient contact with drugs, then resulting in more accumulated drug release. We can conclude that such composite nanofibers show great potential uses as a pharmaceutical application (Wang et al. 2012). Furthermore, the adhesion and proliferation of human MG-63 cells were cultured on neat PLA nanofiber and composite nanofibers with (TH 15%) to investigate the cytocompatibility of the biomaterials for potential uses in biomedical applications. Fluorescent images were used for the qualitative attachment analyses of MG-63 on neat PLA and the composite nanofibers, respectively. In Fig. 7c, the amounts of cells cultured on the composite nanofibrous were more numerous than those for neat PLA nanofiber. This result was ascribed to the increased hydrophilicity of composite nanofibers and good cytocompatibility of CNC, which directly improved in vitro/vivo biological performance, such as tissue response and cell attachment. It has been reported that incorporation of CNC could improve efficiently cell–matrix interactions and thus cellular uptake of the biopolymer composites (Yu and Yao 2016). Moreover, it has been reported previously that the random-oriented scaffolds were better for cell growth than other scaffolds (Lu et al. 2012). Such composite nanofibers with excellent biocompatibility exhibit great potential for biomedical applications.

Conclusion In summary, electrospun PLA/CNC/PEG nanofibers were successfully prepared via an electrospinning technique. The increasing contents of CNC/PEG reduced fiber diameters, and enhanced fiber uniformity. The water contact angle of composite nanofibers was decreased from 117.3° for neat PLA to 98.0° for composite nanofibers with 10 wt% CNC/PEG. However, when drug-loadings furthered increased to (TH 30%), the contact angle was decreased from 119.3° for neat PLA (TH 3%) to 75.0° for composite nanofibers with (TH 30%), and fiber diameter was decreased from 680 ± 34 to 340 ± 17 nm. As a result, enhanced hydrophilicity of the composite nanofibers was achieved. More interestingly, high drug loading efficiencies of composite nanofibers were more than 98%, and long-term sustained release behavior of composite nanofibers was obtained. Meanwhile, the composite

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nanofibers with (TH 15–30%) drug loadings delivered more than 95.7% of their drug content within 1032 h. In contrast, neat PLA nanofiber gave about 13% drug release during the same period. Furthermore, the composite nanofibers showed good biocompatibility with MG-63 cells by incorporating CNC/PEG and drugs. With increasing drug contents, both Young’s modulus and tensile strength of the composites nanofibers with (TH 15%) were improved by 240 and 57.1%, respectively. Additionally, the T0, Tmax, and Tf values were improved by 7.3, 13.5 and 7.6 °C respectively. The composite nanofibers with improved physical and hydrophilic properties, especially longterm drug release, displayed good biocompatibility and exhibited great potential for application in fields such as long-term sustained drug delivery systems and other biomedical materials. Acknowledgments The financial supports from the National Natural Science Foundation of China (51403187), and ‘‘521’’ Talent Project of Zhejiang Sci-Tech University and Open fund in Top Priority Discipline of Zhejiang Province in Zhejiang SciTech University (2016YXQN07) are greatly acknowledged.

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