Dual drug release from core-shell nanoparticles with distinct release profiles.

RESEARCH ARTICLE – Pharmaceutics, Drug Delivery and Pharmaceutical Technology

Dual Drug Release from Core–Shell Nanoparticles with Distinct Release Profiles YANG CAO,1,2 BOCHU WANG,1 YAZHOU WANG,1 DESHUAI LOU1 1

Key Laboratory of Biorheological Science and Technology, College of Bioengineering, Ministry of Education, Chongqing University, Chongqing 400030, China 2 Chongqing Key Laboratory of Ultrasound Molecular Imaging, Department of Ultrasound, Institute of Ultrasound Imaging, The Second Affiliated Hospital of Chongqing Medical University, Chongqing 400010, China Received 4 April 2014; revised 6 July 2014; accepted 15 July 2014 Published online 12 August 2014 in Wiley Online Library (wileyonlinelibrary.com). DOI 10.1002/jps.24116

ABSTRACT: Multiple drug combination is a promising strategy in biomedical fields, such as cancer chemotherapy and tissue engineering. With the aim of codelivering multiple drugs with different characteristics, immiscible and miscible liquids were utilized to fabricate nanoparticles of polyvinylpyrrolidone/poly(lactic-co-glycolic acid) (PLGA) and poly(ε-caprolactone)/PLGA with distinct core–shell structure by coaxial electrospray. Each kind of nanoparticles can encapsulate the hydrophilic rhodamine B and hydrophobic naproxen in one single step efficiently. Encapsulation efficiency was over 85%. The different release patterns of dual-drug encapsulated were demonstrated when the drug location swapped, attributing to the distinct core–shell structures of nanoparticles and the interaction between drug molecules and carrier polymers. Meanwhile, the release profiles of encapsulated drugs with different loading amount were investigated as well. Dual drug release profiles from nanoparticles were affected by the unique architecture of nanocarriers (porous and core–shell structure), physical properties of polymers, and drugs. In addition, polymer–drug and drug–drug molecular interaction may take an important role in drug release behaviors. The results suggested that the distinct release kinetics of multiple drugs fabricated by coaxial electrospray can be C 2014 Wiley Periodicals, Inc. and the American Pharmacists obtained and tuned to fulfill the clinical requirement in combination therapy. Association J Pharm Sci 103:3205–3216, 2014 Keywords: distinct release profiles; core–shell nanoparticles; dual-drug encapsulation; Biomaterials; Controlled release/delivery; Nanoparticles; Nanotechnology; Polymeric drug carrier

INTRODUCTION Antitumor treatment is a multistage process that is regulated by a serial of inhibition factors. It is known that cancer cells are able to acquire defense mechanisms by over expressing drug efflux pumps, enhancing self-repairing ability, increasing drug metabolism, or expressing altered drug targets.1,2 Therefore, the efficacies of cancer therapy with one single antitumor drug delivery are diminished. One of promising approaches is the codelivery of various therapeutic agents in the same delivery vehicles for the cancer treatment.3–5 It has been developed as an effective, well-known regimen used in the daily treatment of tumors clinically, with the advantages of a synergistic effect, suppressed drug resistance and the ability to tune the drug dosage at the target sites.6 By delivering multiple therapeutic agents to the target sites simultaneously, the multidrug delivery systems can promote drug synergism and pave the way to precision design and tailoring in cancer chemotherapeutics.1,2 Since the inhibition factors exhibit their specific time-dependent concentration profiles in the antitumor treatment, therapeutic drugs or factors should be used at optimal dosages for different periods to optimize their effects.7,8 In order to tackle these challenges, drug combinations releasing multiple drugs in a controlled manner are highly demanded. Unlike controlled release Correspondence to: Bochu Wang (Telephone: +86-023-65112840; Fax: +86023-65112877; E-mail: [email protected]) Bochu Wang’s present address is Shazheng Street No. 174, Shapingba, Chongqing, China. Journal of Pharmaceutical Sciences, Vol. 103, 3205–3216 (2014)

C 2014 Wiley Periodicals, Inc. and the American Pharmacists Association

systems that deliver a single drug, multidrug delivery systems with distinct release profiles are more difficult to fabricate. Several strategies have been utilized to codeliver multiple drugs into a single carrier, such as physical loading into the particle core,9,10 chemical conjugation to the particle surface,3–5 and covalent linkage between polymer and therapeutic agents.11,12 However, it is a major challenge for chemical conjugation to control the ratios of different drugs in the same systems, resulting from steric hindrance between drug and polymers, and batch-to-batch heterogeneity.1,2 Physical encapsulation is a promising drug loading strategy that has been widely used in multiple growth factors delivery in tissue engineering. The polymeric scaffold developed by Richardson could provide multiple growth factor delivery with a distinct release rate for each factor.13 The release rate was adjusted in this system by simply altering the amount of factor incorporated into scaffolds as well as polymer degradation time. In addition, the polymers’ characteristics such as cross-linking and polyionic complexation have been reported to affect the product’s loading efficacy and release kinetics.14–16 Multilayered or multicompartmentalized microcapsules also allows for simultaneous multiple drug delivery in biomedicine.17,18 The microcapsules shell provided additional barriers to the diffusion of loaded drug in the core. Despite the advancement in nanoparticle drug delivery, most research efforts focus on microscale delivery systems in tissue engineering, whereas delivering multiple drugs with a single nanovesicle remains largely unexplored. Meanwhile, the large size of the mircoparticles is a hindrance for intravenous injection, and surgical operations become a necessity for implantation, which is inconvenient for patients.

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Compared with conventional drug formulations, polymeric micro/nanoparticles have the potential to better control of delivery with large molecules such as proteins and nucleic acids as well as small molecule entrapment.19 Moreover, biocompatible and biodegradable polymeric micro/nanoparticles have attracted great attention because they have the potential to prolong drug circulation half-life, reduce nonspecific uptake, and better accumulate at the tumors through enhanced permeation and retention effect or active targeting.20 Over the past few decades, several technologies have been developed to deliver multiple drugs or therapeutic agents by incorporating them into matrices, liposomes, polymeric microparticles, nanoparticles, and so on.21 The most popular technologies for multiple drug delivery are emulsion-based techniques and electrospun/spraying.22 The application of emulsion-based methods is limited since the long time exposure of bioactive agents to harsh organic solvent would make the loaded drug denatured to lose biological activities. Moreover, there are some other disadvantages such as numerous steps involved in the processing, additives or surfactants resulted in a broad particle size distribution and poor drug loading capacity.23,24 Some attempts have been made to produce polymeric delivery systems using electrospun/spraying for pharmaceutical purposes.25–27 The great advantage of the electrospun/spraying method is that it is simple and inexpensive. This method can produce monodisperse particles or fibers without the need of surfactants or elevated temperatures.28 In addition, it can be operated at ambient temperature and pressure conditions.29 The small diameter of the electrospray particles and electrospun fibers with a high surface area is helpful for mass transfer and efficient drug release.30 The drug release profiles of electrospray particles and electrospun fibers can finely tailored by morphology, composition, and porosity of the systems.31–34 However, it is difficult for the aforementioned technologies to load different drugs with dissimilar physical properties into the same carrier.35 Thus, coaxial electrospinning with two separate feeding capillary channels is introduced. This technology would provide an alternative and simple means to produce core–shell systems, which can be utilized to deliver multiple drugs simultaneously.36 It has the potential to encapsulate therapeutic agents with different hydrophilic properties inside a core–shell polymeric particle, which supersedes other methods requiring two or more steps to achieve the encapsulated product.37 Understanding the release mechanisms, as well as how to alter drug release profiles, is important for the successful tumor inhibition. Meanwhile, in nanoscale drug systems, drug release behaviors are mainly affected by the interactions among solvent molecules, polymers, and drug molecules. One of the main challenges of multiple drug combination is to choose the corresponding type of nanoparticles to get the required distinct drug release behaviors. With the aim of better know the mechanisms of multiple drug release properties from different nanocarriers, two kinds of core–shell nanoparticles were fabricated by coaxial electrospray in this paper. According to the affinity toward water, hydrophilic and hydrophobic polymers were chosen as carrier materials. Poly(lactic-co-glycolic acid) (PLGA), approved by United States Food and Drug Administration (US FDA) for commercial application, is the most widely studied because of its sustained release characteristics, biocompatibility, and biodegradability.38–40 It was chosen as the shell polymer. The hydrophobic poly(g-caprolactone) (PCL) and Cao et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:3205–3216, 2014

hydrophilic polyvinylpyrrolidone (PVP) were chosen as the core carriers. PCL also has been approved by US FDA to be used in vivo. It has been reported to have lots of advantageous properties, such as good biodegradability and biocompatibility, good drug permeability and antifatigue capability, as well as low cost relative to other biodegradable polyesters. PVP is widely used as an additive to tune the drug release behaviors. It has been reported that the drug release behaviors could be adjusted by changing the ratio of the PVP in drug carriers.41 Hydrophilic rhodamine B and hydrophobic naproxen were chosen as the model drugs encapsulated in nanoparticles (PVP/PLGA and PCL/PLGA). They were encapsulated in individual regions of nanoparticles respectively, and the distribution of model drugs is adjustable to alter the different release behaviors. Moreover, distinct release profiles of dual-drug with different drug loading contents are investigated and the releasing mechanisms of the core–shell nanoparticles are discussed as well.

EXPERIMENTAL Materials Poly (D,L-lactide-co-glycolide) copolymer (PLGA 50:50, Mw: 50,000 Da) and PCL (Mw: 80,000 Da) were purchased from Jinan Daigang Bio-Tech. Inc., Jinan, China. PVP (K30, Mw: 50,000 Da) was obtained from Sinopharm Group Company Ltd., Tianjin China. Rhodamine B and naproxen were purchased from Sangon Biotech Company, Ltd., Tianjin China and Galaxy Chemical Company, Ltd. (Wuhan, China), respectively. Nanoparticle Fabrication by Coaxial Electrospray The coaxial setup was used to produce nanoparticles (Fig. 1). The core capillary had an inner diameter of 0.3 mm and outer diameter of 0.5 mm. The inner diameter of outer capillary was 1.0 mm. The inner and outer fluid liquids were injected into two flow channels by separate syringe pumps (TJ-3A; Longer, China). PLGA was used as outer material. PCL or PVP was chosen as inner polymer. A combination of tetrahydrofuran and acetonitrile (2:8, v/v) was used as the organic solvent to dissolve PLGA. PCL was dissolved in acetonitrile. N,N-dimethyl formamide (DMF) was added into the PVP ethanol solution

Figure 1. Schematic diagram of the coaxial electrospray system for core–shell nanoparticle. DOI 10.1002/jps.24116


(DMF–ethanol = 1:2) to increase the electrical conductivity of the solvent. A voltage generator supplied a high voltage to the nozzle by means of a crocodile clip. The potential was 16 kV and the distance between the needle tip and the collector was 9 cm. The outer PLGA copolymer solution was 6% (w/v). The inner copolymer solution was 6% (w/v) PCL or 4% (w/v) PVP, respectively. These two systems were named as PCL/PLGA and PVP/PLGA. The flow rates of inner and outer fluid liquids were 0.2 and 0.8 mL/h.

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Statistical Analysis All the experiments were measured in triplicates and the results were presented as mean ± SD. Statistical analyses of the experimental data from different groups were performed by applying one-way ANOVA (p < 0.05).

RESULTS Characterization of Electrospraying Particles

Characterization of Nanoparticles The core–shell nanoparticles were collected and suspended in anhydrous ethanol. Dynamic light scattering (DLS) methods were used to measure particle diameters and size distribution at 20◦ C by Zetasizer Nano ZS 90 instrument (Malvern Instruments Ltd., Malvern, UK). The morphologies of nanoparticle surface were observed by field emission scanning electron microscopy (FE-SEM, Nova 400; FEI). For transmission electron microscope (TEM; FEI Tecnai 10; Philips Electron Optics, Holland), nanoparticle dispersions were deposited and dried on formvar film-coated copper grids after stained by phosphomolybdic acid hydrate. In order to examine the core–shell structures of the nanoparticles, FITC (fluorescein isothiocyanate) was mixed with core polymer solutions. The FITC-loaded particles were produced as mentioned above. A fluorescence optical microscopy was used to observe the structure of the electrospraying nanoparticles.

In order to obtain smooth and uniform particles, and to operate the system in the stable cone-jet mode, the solution and processing parameters of electrospraying were optimized. The surface morphologies and core–shell structure of particles were illustrated by SEM and TEM images (Fig. 2). SEM images revealed that both PVP/PLGA and PCL/PLGA were generally spherical and nearly monodisperse in size. According to the results of DLS measurements, the average diameter of PVP/PLGA and PCL/PLGA nanoparticles was 394.7 ± 35.2 nm (Pdi = 0.093) and 606.5 ± 39.9 nm (Pdi = 0.289), respectively. Because RhB and Nap changed the properties of the electrospray liquid solutions, size diameters of nanoparticles were increased after mixing dual drugs with the polymers (Fig. 2). The average diameter of PVP-R/PLGA-N, PVP-N/PLGA-R, PCL-R/PLGA-N, and PCL-N/PLGA-R was 404.2 ± 32.7, 419.2 ± 23.5, 665.2 ± 45.9, and 623.3 ± 53.2 nm, respectively. As shown in TEM and LSCM (Fig. 2), these two systems had the desired core–shell structures.

In Vitro Drug Release Measurements

In Vitro Drug Release Profile Studies

The drug-loaded particles were collected and incubated in a conical flask with 20 mL of prewarmed phosphate-buffered saline (PBS) (60–120 rpm/min) at 37◦ C. At designated time intervals, a 2-mL release medium was taken. The same amount of fresh buffer was then added in order to maintain the sink condition. The concentrations of the rhodamine B and naproxen were determined using the dual-wavelength spectrophotometer by UV spectrometry (Lambda 900; PerkinElmer). Accumulated release percentage of drug was determined as

With modern nanotechnology developed, some of methods have been used to load two or more therapeutic agents successfully. Owing to the advantages of coaxial-capillary electrospraying, it allows the encapsulation of multiple drugs in one single step, including drugs with different characteristics in hydrophilic properties. By choosing various polymer carriers, distinct drug release profiles can be achieved to fit the requirement in clinical. In this study, hydrophilic RhB and hydrophobic Nap were chosen as model drugs. Two core–shell particles with different core materials (PCL/PLGA, PVP/PLGA) were fabricated by coaxial-capillary electrospray method. In group A, 1% (w/w) RhB and Nap were mixed with carrier polymers as outer liquid or inner liquid (Table 1). With different characteristics of core polymers in hydrophilic properties in nanoparticles, PVP/PLGA and PCL/PLGA demonstrated variable release profiles. In the mono-drug systems, both RhB and Nap were affected by the drug distribution significantly (Fig. 3). Nap had a different degree of burst release of these two systems. It showed faster and higher cumulative release content for Nap than RhB in PCL/PLGA systems, despite of no matter where the drug located. In PVP/PLGA nanoparticles, it showed that RhB had faster release rate and higher cumulative release when located in PVP core, whereas Nap showed the opposite trends. It had higher release rate and cumulative release in PLGA shell. It displayed an interesting phenomenon that the release sequences of these two drugs were swapped when located in different regions in mono-drug systems. The initial burst release of RhB was inhibited when entrapped in PCL/PLGA. It suggested that different drugs exhibited different release behaviors in different systems. Distinct release kinetics of RhB was observed in these four samples. It had the

Cn · V + Vi Q(%) =

n−i i=0

mdrug

Ci × 100%

(1)

Here, Q (%) was the amount of accumulated release drugs. V (mL) was the total volume of samples. Cn (mg/mL) and Vi (mL) were the concentration and volume of samples taken at n and i time point. mdrug (mg) was the mass of drug in particles. The mass of drug was calculated as the total amount of drug dissolved in the fluid liquids in electrospray process. The number of times of drug release media replacements was numbered as n. The amount of drugs encapsulated in the composite particles was measured by dissolved the drug-loaded particles in 0.5 mL DMF. Add 2.5 mL PBS (pH 7.4) to the mixtures. Rhodamine B was determined at 554 nm and naproxen was measured using 330 nm as a characteristic wavelength and 369 nm as a reference one, respectively. The entrapment efficiency tests were performed according to Peltonen et al.42 DOI 10.1002/jps.24116


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Figure 2. (a and b) SEM images of nanoparticles. (a1) PCL/PLGA, (a2) PCL-R/PLGA-N, (a3) PCL-N/PLGA-R, (b1) PVP/PLGA, (b2) PVPR/PLGA-N, (b3) PVP-N/PLGA-R; (c and d) TEM images of nanoparticles. (c1 and c2) PCL/PLGA, (d1 and d2) PVP/PLGA; (e and f) Light microscope and fluorescence optical microscope images of core–shell nanoparticles. (e1) Light microscope image of PCL–PLGA nanoparticles. (e2) Fluorescence optical microscope image of PCL–PLGA nanoparticles. (e3) The combination of images from (e1) and (e2). (f1) Light microscope image of PCL–PLGA nanoparticles. (f2) Fluorescence optical microscope image of PVP–PLGA nanoparticles. (f3) The combination of images from (f1) and (f2). Scar bar = 5 :m.

highest cumulative release in PVP-R/PLGA. However, RhB release behaviors were delayed in PCL/PLGA nanoparticles obviously. When two drugs were both entrapped in the nanoparticles, drug release profiles were quite different. Figure 4 showed the experimental drug release profiles from PVP/PLGA and PCL/PLGA systems. The sequences of two agents did not change when the drug distribution was swapped for PCL/PLGA nanoparticles. PCL-R/PLGA-N presented a significant initial burst effect, in which more than 50% of Nap was released out within the first 12 h. The release in the following time was linear until 72 h. Similarly, Nap was initial burst released from PCL-N/PLGA-R with 45% in first 12 h and nearly 60% after 72 h. In contrast, the initial burst of RhB was minimal in these two systems, only 30% was released in the beginning 12 h for PCL-R/PLGA-N and nearly 2% for PCL-N/PLGA-R. When RhB was located in outer shell (PCL-N/PLGA-R), it had a slower release rate than in inner parts of nanoparticles (PCLR/PLGA-N), indicating that the release profiles of RhB were influenced by the drug distribution as well as the characteristics of the carrier polymer. This phenomenon, however, was not founded when Nap was encapsulated. In PVP/PLGA systems, Nap showed the highest cumulative release when entrapped in the core. However, RhB had the similar release profiles in the core and shell parts of the nanoparticles.


Compared with drug release behaviors from mono-drug systems (PVP-R/PLGA), the cumulative release of RhB was decreased in dual-drug systems (PVP-R/PLGA-N) with nearly 50% drug release in the first 12 h, whereas 80% for the former sample (Fig. 5). The accumulation of Nap release was much more when encapsulated in core (PVP-N/PLGA-R) than shell (PVP-R/PLGA-N) after 72-h incubation. When located in PVP core, the accumulation of Nap release from mono-drug systems was much higher than that of dual-drug nanoparticles. As for RhB, it did not show much difference between these two systems. The results indicated that Nap presented different release kinetics when located in different regions, and RhB showed a similar release behavior no matter which part it located in. In PCL/PLGA systems, the initial release of Rhb was hindered (Fig. 5). RhB of mono-drug and dual-drug systems had the similar release profiles when entrapped in PLGA shell. The accumulative of RhB in PCL-R/PLGA-N was higher than that of PCL/PLGA-R by 20%. Nap showed different release kinetics from RhB in PCL/PLGA nanoparticles. These all four samples exhibited the similar releasing curves, indicating release behaviors of Nap could not be alternated by changing its distribution. The distinct release profiles were observed from these two systems (Fig. 6 and Table 2). The effective concentrations of the model RhB and Nap were considered at the level when the

DOI 10.1002/jps.24116


Table 1.

Definition of Samples in Coaxial Electrospray Experiments Inner Fluid

Samples

Outer Fluid

Polymers

Drug

Polymers

Drug

PCL

Rhodamine B Naproxen / / Rhodamine B Naproxen Rhodamine B Naproxen / / Rhodamine B Naproxen

PLGA

/ / Rhodamine B Naproxen Naproxen Rhodamine B / / Rhodamine B Naproxen Naproxen Rhodamine B

Group A PCL-R/PLGA PCL-N/PLGA PCL/PLGA-R PCL/PLGA-N PCL-R/PLGA-N PCL-N/PLGA-R PVP-R/PLGA PVP-N/PLGA PVP/PLGA-R PVP/PLGA-N PVP-R/PLGA-N PVP-N/PLGA-R Group B PCL-1/PLGA-10 PCL-5/PLGA-10 PCL-5/PLGA-5 PCL-5/PLGA-1 PCL-10/PLGA-10 PVP-1/PLGA-10 PVP-5/PLGA-10 PVP-5/PLGA-5 PVP-5/PLGA-1 PVP-10/PLGA-10

PVP

PCL

PVP

PLGA

Rhodamine B (w/v) Naproxen (w/v) 1% PLGA 10% 5% 10% 5% 5% 5% 1% 10% 10% 1% PLGA 10% 5% 10% 5% 5% 5% 1% 10% 10%

concentration kept constant in this study, according to ideal drug dynamics in vivo. It presented the same trends that Nap was released out of the nanoparticles in the first 0.5 h. RhB showed a different release profile. In PCL/PLGA, an initial effect was observed when RhB loaded in the shell. When encapsulated in the core, a sustained release was obtained for PCL-N/PLGA-R. As for PVP/PLGA nanoparticles, RhB both had a peak concentration level for core and shell located. It kept steady afterwards. Compared with PCL-R/PLGA-N, RhB release was greatly inhibited when located in outer shell (PCLN/PLGA-R). However, it exhibited a slightly higher release rate of Nap when encapsulated in outer shell (PCL-R/PLGA-N). The observed curves of drug release suggested that the release behaviors of RhB were more sensitivity to drug distribution than that of Nap in PCL/PLGA nanoparticles. In PVP/PLGA systems, the release of RhB kept the same trends when loaded in core or shell. Nap, however, showed different release rate when the distribution was alternated. The differences between RhB and Nap release kinetics were various for different samples. The dosage forms can be adjusted to fit the clinical demand. It has been reported the drug release profiles may be facilitated by the amount of drugs encapsulated in nanoparticles.22,43,44 The higher drug loading means that nanoparticles can carry more chemotherapy drugs. However, an increase in drug loading may result in rapid drug release. To better control the drug release rate of polymeric nanoparticles, samples with different drug loading were prepared and the drug release profiles were investigated. RhB and Nap were encapsulated in core and shell, respectively. Figure 7 plotted the release profiles of RhB and Nap from the various nanoparticles. The trends of RhB and Nap release profiles did not change DOI 10.1002/jps.24116

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obviously in different samples of PCL/PLGA systems. The release rate of Nap was faster than that of RhB. On the contrary, PVP/PLGA nanoparticles yielded a different release pattern. When different amount of drug encapsulated, the release rates of drugs were alternated. Nap was released from the polymer shell firstly, and then RhB diffused out of the nanoparticles in PVP-5/PLGA-10 and PVP-1/PLGA-10 systems. However, in PVP-10/PLGA-10 nanoparticles, the drug release sequences were alternated. RhB was released at the beginning. Nap was released afterward. It exhibited the same drug release trends for the samples when the 5% RhB loaded in the core. Generally, the results suggested that Nap in the shell was released firstly and RhB in core released afterwards for all PCL/PLGA nanoparticles. As for PVP/PLGA, an interesting phenomenon was observed. The drug with higher encapsulation amount was released faster than the other one, despite of their location. Comparing with emulsion single capillary electrospraying, high encapsulation efficiency of encapsulated drugs was observed in coaxial-capillary electrospraying (Table 3). In this study, both RhB and Nap had high encapsulation efficiencies over 85%. It seems that the hydrophobic or hydrophilic properties of the drugs did not affect the amount of encapsulation with the coaxial-capillary electrospray method, as reported previously.31–33 In another words, both hydrophobic and hydrophilic drugs can be encapsulated into the coaxial electrospray particles effectively. The efficient encapsulation of drugs, particularly the hydrophilic ones, into the hydrophobic polymer is very promising.

DISCUSSION Over the past decades, many pharmaceutical researches have witnessed the boost of multiple drug delivery systems. Multiple drug combination is one of promising strategies in several diseases, such as HIV/AIDS, malaria, and cancer chemotherapy. To optimize their effects, therapeutic drugs or factors should be used at optimal dosages for different periods, resulting in an additive or synergistic effect.3–5,7,8 The specific distinct drug release of multiple drugs or factors with temporal profiles is required in a successful tissue regeneration or tumor inhibition treatment.45 In order to deliver the individual drug at optimal dosages, the release behaviors of each drug in multidrug systems should be controlled independently. With the aim of delivering and releasing multiple drugs in various release stages, two kinds of nanoparticles with core–shell structure (PVP/PLGA and PCL/PLGA) were fabricated by coaxial electrospraying, which is a one-step method and has the potential to generate narrow size distributions of nanoparticles, with limited agglomeration and high yields.46 As shown in the Figure 1, electrospray nanoparticles looked uniform and presented porous structure on the surface. It is the typical morphology of electrospraying and electrospinning products, attributing to the rapid evaporation of organic solvent during electrospraying or electrospinning process.41 It is believed that solvent evaporation and polymer diffusion are the two main mechanisms determining particle formation and their resulting properties.47 Because of the different evaporation rate and diffusion rate between PCL solution and PVP solution, the different size distributions of these two systems were presented. Because of the particle aggregation behaviors, Cao et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:3205–3216, 2014

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Figure 3. The release profiles of RhB and Nap from mono-drug-loaded nanoparticles.

Figure 4. The release profiles of RhB and Nap from dual drug-loaded nanoparticles.

particle size distribution obtained from DLS measurement was higher than that of SEM and TEM measurements.29,35 The difference in the diameter and surface of nanoparticles produced by coaxial electrospray may be due to the removal rate of solvent from the PVP ethanol solution and PCL acetonitrile solution.48 Since the dissolution of PLGA in ethanol was negligible, the polymer–solvent configuration was able to strongly Cao et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:3205–3216, 2014

reduce the diffusion between outer solution and inner solution at the tip of the coaxial needle.45 Meanwhile, the breakup process is greatly influenced by fluid properties such as electrical conductivity, viscosity, and interfacial tension.49 The average diameter of PVP/PLGA nanoparticles was smaller than that of PCL/PLGA because the addition of hydrophilic PVP caused higher conductance of electrospraying solution.41 In addition, DOI 10.1002/jps.24116


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Figure 5. The release profiles of RhB and Nap from mono-drug-loaded and dual drug-loaded nanoparticles.

the miscibility of the core and shell liquid is important for coaxial-capillary electrospraying. Generally, smaller interfacial surface tension of outer liquid than inner liquid was required, since it had good controllability and clear boundary. Because of the different surface tension of immiscible solutions, phase separation would occur during electrospray process. In another case, it has been reported that a stable compound cone-jet mode was easily established for miscible and partially miscible liquid pairs since the interfacial surface tension between the inner and outer liquids was minimized.45 In order to check the performances of these systems, both immiscible and miscible solutions were fabricated in this study. As shown in TEM images, these two systems had the desired core–shell structures. Several factors were reported to affect drug release rate from core–shell structure nanoparticles, such as shell thickness, drug physical properties, drug–polymer interactions, and drug–drug interactions. Initially, the water molecules diffused into the nanopores in the shell layer, and then the surfaceloaded drug and encapsulated drug in the shell diffused out through the polymer shell and/or nanopores into the buffer solution, resulting in burst release in the first 12 h. There was an obvious burst release in the beginning of the release profiles of mono-drug- and dual-drug-loaded systems. These initial burst releases were probably resulted from the surface or near surface loading of the drug, judging from the quick re-

DOI 10.1002/jps.24116

lease observed in the beginning.47 Entrapped drugs tended to move to the surface of the nanocarriers in the process of electrospraying along with solvent evaporation.41 Meanwhile, since the drug and carrier polymer were well-mixed before electrospray, the initial burst release occurred because of the large surface to volume ratio of the nanoparticles fabricated in this study, resulting from the nano-scale size diameters and high porosity of the matrix. The porous structures and pores after solvent evaporation in the nanoparticle surface made it possible for encapsulated drugs to diffuse in the initial incubation time.31–34 It was shown that Nap exhibited the similar release behaviors both in core or shell (Fig. 3). Although Nap was encapsulated in core region, the drug with small molecules weight could diffuse through pores of the matrix easily and quickly, which were formed due to the fast evaporation of solvent during electrospraying. In addition, it has been reported that the layer or the shell can provide additional barriers to the diffusion of the entrapped drugs in the core.50 When the drug was encapsulated in the core, there were two more barriers, which might affect its release profiles. One was thick PLGA shells, leading to long diffusion lengths. When the drug encapsulated in the outer shell was released, a porous shell was formed and the internal porosity would act as channels to allow more rapid penetration of water into the core and thus faster dissolution


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Figure 6. Drug concentration of dual-drug-loaded particles.

and release of encapsulated drugs.22 The other was the interactions between the core drug and the shell drug during release, such as formation of physical or covalent aggregate.43,44 As for core–shell structure systems, the diffusion rate was affected by the ratio of the thickness of the shell to that of the core.30 The thickness of shell layer can be adjusted by the flow ratio of inner liquid and outer liquid in electrospray process. Because the electrospray parameters were kept constant in this study, the shell thickness had no obvious difference for these drug-loaded systems. Meanwhile, the thickness of the nanoscale particles was too small as shown in Figure 2 (TEM). The effects of shell thicknesses on drugs release from the electrospray particles with the porous structure were diminished. The drug release behaviors were depended on particle porosity and hydrophobicity, as well as molecular interaction between polymer and drugs after this initial burst.51 Because of Table 2. Differences Between the Time when Reaching Effective Concentration of RhB and Nap Samples PCL-R/PLGA-N PCL-N/PLGA-R PVP-R/PLGA-N PVP-N/PLGA-R

RhB

Nap

t

2h Over 12 h 2h 4h

0.5 h 0.5 h 0.5 h 0.5 h

1.5 h Over 11.5 h 1.5 h 3.5 h


the large surface to volume ratio of the nanoparticles, high degree of particle swelling may occur, increasing the interface of polymer–water and water penetration rate.52 Since PCL is a hydrophobic polymer with good penetrability and poor water swelling capacity,53 water penetrated in PCL may take an important role in drug release. In PVP/PLGA, the hydrophilicity enhancer PVP can sharply increase the uptake of water molecules into nanoparticles.54 Thus, drug molecules can escape from the systems more easily. In addition, pin nanopores were presented on the surface of polymer shell, resulting from polymer swelling and solvent evaporation in electrospray process. It has been reported that changes in the polymer concentration also influenced drug release rates.55 The swelling and bulk erosion time of the nanoparticles produced with high polymer concentrations can be delayed by dense chain entanglement, resulting in prolonged diffusion times. In this step, the release rate was believed to depend on the rate of water penetration and drug diffusion. A diffusion-driven release might dominate the drug release profiles. The different drug release rates and release patterns from the core–shell nanoparticles developed were mainly attributed to the distinct core–shell structures of nanoparticles and the difference of two drugs in hydrophilic properties.37 It is well known that the drug feature is one of the significant effects that can influence drug release rate, including drug dissolution, polymer–drug interactions, and drug–drug DOI 10.1002/jps.24116


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Figure 7. Release profiles of dual-drug-loaded nanoparticles with different drug loading amount.

interactions. The drug must be dissolved in water before being released, and this process could decrease the overall release rate.43,44 Drug dissolution and diffusion out of the matrix could be promoted by the water penetration, attributing to the little affinity between encapsulated drugs and carrier polymers.51 In this study, hydrophilic RhB and hydrophobic Nap had the different solubility in PBS release medium. It showed various release rates in core–shell nanoparticles. Since the drug was diffused out of the particles through the hydrophilic porous DOI 10.1002/jps.24116

passages inside the polymer matrix, the different behaviors between the release of RhB and Nap were attributed to the different partition coefficients of the two substances within the polymer matrix.56 As suggested in mono-drug-loaded systems, Nap with better partition coefficient within PLGA can be entrapped with the chains of the polymers in nanoparticles. On the contrary, RhB entrapping in hydrophilic porous regions inside the PLGA matrix usually exhibited burst release effect in PVP-R/PLGA nanoparticles. In another words, the strong Cao et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:3205–3216, 2014

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Table 3.

Encapsulation Efficiencies and Drug Loading Rates of RhB and Nap in Nanoparticles

Samples

Encapsulate Efficiency (%)

Samples

Encapsulate Efficiency (%)

Group A PCL-R/PLGA-N PCL-N/PLGA-R PCL-R/PLGA PCL /PLGA-R PCL-N/PLGA PCL /PLGA-N

RhB 92.5 ± 0.9 82.8 ± 1.4 95.3 ± 2.1 88.3 ± 1.2 – –

Nap 78.4 ± 3.1 90.3 ± 2.1 – – 92.2 ± 0.8 85.3 ± 1.4

PVP-R/PLGA-N PVP-N/PLGA-R PVP -R/PLGA PVP /PLGA-R PVP -N/PLGA PVP /PLGA-N

RhB 99.3 ± 0.7 90.1 ± 1.4 99.4 ± 0.7 93.1 ± 2.4 – –

Nap 90.2 ± 3.2 94.3 ± 1.3 – – 89.4 ± 3.2 79.2 ± 2.4

Group B PCL-1/PLGA-10 PCL-5/PLGA-10 PCL-5/PLGA-5 PCL-5/PLGA-1 PCL-10/PLGA-10

94.3 ± 2.4 92.1 ± 1.1 95.8 ± 0.3 94.3 ± 0.4 96.3 ± 0.6

90.3 ± 3.2 89.2 ± 0.8 87.3 ± 3.2 85.3 ± 4.3 90.3 ± 0.4

PVP-1/PLGA-10 PVP-5/PLGA-10 PVP-5/PLGA-5 PVP-5/PLGA-1 PVP-10/PLGA-10

98.1 ± 1.5 99.1 ± 0.8 99.2 ± 1.2 98.9 ± 0.5 99.8 ± 0.9

93.1 ± 2.1 91.4 ± 1.3 90.1 ± 1.6 88.3 ± 1.7 93.1 ± 0.8

affinity with carrier polymers can prevent the diffusion of the encapsulated drug out of the polymer matrix. Besides, different from PCL/PLGA core–shell nanoparticles, PVP/PLGA had a hydrophilic core, which was benefited to load the hydrophilic agents. Moreover, the hydrophilic of the systems was changed, increasing the uptake of water.54 The loaded drugs can diffuse from the systems easily. In addition, because of the hygroscopic and hydrophilic properties, PVP was reported that it can accelerate the drug release rate, resulting from the high polymer– solvent interactions.57 Therefore, the polymer chain can absorb more water molecules and encapsulated drugs can diffuse from the matrix of greater porosity. Moreover, the sustained release behaviors of entrapped drug in dual-drug loading were possibly related with the presence the other drug on the shell of the carriers.7,8 When distinct drug loading content alternated, the release rates of encapsulated drugs changed (Fig. 7). Because of a high amount of drug loading, the thickness of polymeric matrix was reduced and the porosity was increased, resulting from the space left vacant after drug release. Therefore, a faster drug release rate occurred. Because of the interactions between RhB and PCL, which had the more hydrophobicity than PLGA, the release of RhB was hindered by both PCL core and PLGA shell. Thus, it presented sustained release profiles for RhB in PCL/PLGA nanoparticles. When it came to PVP/PLGA dual-drug-loaded systems, it showed different release kinetics. The faster release of core RhB was occurred when the amount of RhB was larger than that of Nap. It attributed to the affinity between drug and carrier polymers. Because of the improved affinity, hydrophilic RhB would tend to disperse better in hydrophilic PVP polymers. The mount of drug encapsulated had a great effect on drug release. The higher showed a quicker release finally. Meanwhile, it has reported that the effect of drug loading amount on the in vitro release profile could be attributed to different drug diffusion rate caused by the different drug loading levels.58 In general, the drug release rate of dual-drug-loaded systems depends on: (1) nanoparticles: size diameters, architecture, porosity; (2) carrier polymers: hydrophilic/hydrophobic properties, drug–polymer interaction; (3) drug: dissolubility, hydrophilic/hydrophobic properties, molecular weight, drug– drug interaction.


CONCLUSIONS Two typical nanoparticles of PVP/PLGA and PCL/PLGA with distinct core–shell structure were fabricated by coaxial electrospraying in this study. It had a narrow size distribution and smooth surface morphology, confirmed by DLS and SEM measurements. It enabled successful encapsulation of two model drugs (rhodamine B and naproxen) with different characteristics in hydrophilic properties in one single step. Encapsulation efficiency of both the hydrophobic and hydrophilic drug into the particles was more than 85%. Different temporal release behaviors of each drug were obtained, attributing to the distinct core– shell structures and the difference of two drugs and polymers in hydrophilic properties. The drug release profiles can be explained as following steps: water penetration, surface drug release, drug dissolution, outer drug diffusion with shell polymer swell, inner drug diffusion, and nanoparticle collapse. Meanwhile, the release profiles of encapsulated drugs with different loading amount were investigated as well. In combination therapy, multiple therapeutic agents with difference in hydrophilic properties can be codelivered by core–shell nanoparticles fabricated by coaxial electrospraying. Specific requirement of individual drug release patterns can be fulfilled and the release rates and characteristic sequences of dual-drugs can be tailored and tuned. Therefore, the synergistic therapeutic effects of the encapsulated drugs are expected in treating diseases in vivo.

ACKNOWLEDGMENT This project was financially supported by the National Natural Science Foundation of China (Project No.: 31200713).

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