Multicompartment Microgel Beads for Co-Delivery of Multiple Drugs at

Dec 31, 2015 - Multidrug therapy may yield higher therapeutic effects as compared to monotherapy, yet its wide application has been hampered by the li...
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Multicompartment Microgel Beads for Co-Delivery of Multiple Drugs at Individual Release Rates Wing-Fu Lai,*,†,‡ Andrei S. Susha,† and Andrey L. Rogach† †

Department of Physics and Materials Science and Centre for Functional Photonics (CFP), City University of Hong Kong, Kowloon, Hong Kong ‡ Department of Mechanical Engineering, The University of Hong Kong, Pokfulam, Hong Kong S Supporting Information *

ABSTRACT: Multidrug therapy may yield higher therapeutic effects as compared to monotherapy, yet its wide application has been hampered by the limitations of conventional drug delivery systems, in which not only incompatible drugs cannot be co-delivered but also the release rates of individual codelivered drugs cannot be tuned separately. Regarding these limitations, we adopt the microfluidic electrospray technology to fabricate alginate-based multicompartment microgel beads. By using cadmium−telluride (CdTe) quantum dots (QDs) and a quenching agent as a model pair, the beads are shown to effectively separate incompatible drugs during co-delivery, and significantly prolong the time of observable fluorescence emission from QDs co-delivered with a quenching agent. Moreover, the drug release rates from different compartments can be tuned using the polymer blending technique to achieve a variety of drug release patterns. This study is one of the first to adopt the microfluidic electrospray technology to generate microgel beads with such versatility for co-delivery of multiple drugs. Our results provide evidence for the promising potential of our beads to be further developed as a carrier for multidrug therapy and other applications that require co-administration of multiple bioactive agents. KEYWORDS: alginate, co-delivery, electrospray, microgels, multidrug therapy

1. INTRODUCTION Co-administration of multiple drugs may offer higher therapeutic efficacy than monotherapy.1−6 This is supported by an earlier study,6 which used double-walled polymeric microspheres to co-deliver doxorubicin with a therapeutic transgene (encoding the p53 tumor suppressor protein, and being complexed with chitosan nanoparticles) to human hepatocellular carcinoma HepG2 cells. Not only was the combined treatment found to enhance cytotoxicity more effectively than either of the two therapies alone,6 but also the overexpression of p53 could promote caspase-3 activation, further enhancing the antiproliferative effect of doxorubicin.6 More recently, the promising potential of multidrug therapy is shown in vivo by the observation that administration of a therapeutic transgene, using adenoviraltype 5(dE1/E3) (Cytomegalovirus promoter), before administration of aerosol cisplatin can increase the expression of ATPbinding cassette (ABC) proteins in the respiratory system to enhance the therapeutic effect of the drug.5 These studies have demonstrated the potential of the multidrug treatment regimen in enhancing the therapeutic efficacy. Over the years, different drug delivery vehicles have been developed for co-delivery of multiple drugs;7−17 however, practical applications of many of these vehicles have been restricted by a number of limitations. For instance, the release rate of each of the co-delivered drugs can hardly be tuned individually, limiting the possibility of delivering multiple drugs © 2015 American Chemical Society

with different desired release kinetics at the same time. In addition, due to the possible interactions among co-delivered drugs, incompatible drugs usually are not able to be delivered concomitantly. This is shown by the observation that when a plasmid is co-delivered with a chemotherapeutic drug using a polymeric vector, the action of the drug may substantially suppress the expression of the transgene.18 This suggests that the incompatibility of chemotherapeutic drugs with gene drugs may reduce the efficiency of gene therapy in practice. Regarding the limitations of multidrug therapy, in this study we adopt the microfluidic electrospray technology and the polymer blending technique to fabricate alginate (Alg)-based multicompartment (MC) microgel beads, which can co-deliver multiple incompatible drugs; in addition, the release rate of each of the co-delivered drugs can be tuned separately to meet practical needs. Together with their size tunability and negligible cytotoxicity, our beads have created new opportunities for future development of multidrug therapy.

2. EXPERIMENTAL SECTION 2.1. Materials. Polyethylenimine (PEI) (Mw = 10000 Da) and carmellose (sodium salt, Mw = 250 kDa, 1500−3100 cP, degree of Received: October 27, 2015 Accepted: December 17, 2015 Published: December 31, 2015 871

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ACS Applied Materials & Interfaces substitution = 1.2) were obtained from Aladdin (Shanghai, China). Rhodamine 6G was purchased from Lambda Physik (Göttingen, Germany). Methylene blue (MB), calcium chloride (CaCl2 ), 3-mercaptopropionic acid (MPA), tetracycline hydrochloride (TH), alginic acid (sodium salt from brown algae), and various other chemicals were purchased from Sigma-Aldrich (St. Louis, MO). Dulbecco’s Modified Eagle’s Medium (DMEM; Gibco, Grand Island), penicillin G-streptomycin sulfate (Life Technologies Corporation, USA), and fetal bovine serum (FBS; Hangzhou Sijiqing Biological Engineering Materials Co., Ltd., China) were used as the cell culture medium. Trypsin-EDTA (0.25% trypsin-EDTA) was obtained from Invitrogen. 2.2. Synthesis of Cadmium−Telluride (CdTe) Quantum Dots (QDs). CdTe QDs were synthesized on the basis of a previously described method with slight modifications.19 Briefly, 1 mmol of cadmium acetate dehydrate was dissolved in 500 mL of deionized water, followed by the addition of 1.12 mmol of MPA. The pH of the reaction mixture was adjusted to 10.5 using a 1 M sodium hydroxide solution. 50 mL of a sodium tellurite solution (0.44 mg/mL), which was used as a tellurium source (Cd:Te = 5:1), was added into the solution mixture. Finally, 10 mmol of sodium borohydride was added to reduce the tellurium salt to Te2−. The reaction mixture was refluxed at 100 °C under open-air conditions. After the reaction, the purification of CdTe QDs was achieved by two precipitation−washing−centrifugation cycles with isopropyl alcohol. The purified CdTe QDs were redispersed in deionized water, and were stored at 4 °C for further experiments. 2.3. Spectroscopic Characterization. The solution of QDs was diluted in PBS (pH = 7.4). The UV−vis absorption spectrum was recorded using a Cary 50 UV−vis spectrophotometer (Varian, Inc., USA). Photoluminescence (PL) measurement was performed at room temperature using a FluoroMax-2 spectrofluorimeter (Instruments SA, Inc., USA). The PL spectrum was collected between 420 and 780 nm using an excitation wavelength of 400 nm. The PL quantum efficiency of the QDs was estimated by comparison to a reference (Rhodamine 6G in ethanol, PL quantum efficiency = 95%) as previously described.20,21 2.4. Treatment of QDs with Different Reagents. To determine the effect of an Alg solution on fluorescence emission from CdTe QDs, an aqueous solution of purified QDs was mixed with an equal volume of a 2% (w/v) Alg solution. The mixture was incubated for a preset time interval, followed by estimation of the PL quantum efficiency. The same approach was adopted to determine the effect of a 2% (w/v) CaCl2 solution and a 2% (w/v) PEI solution on fluorescence emission from QDs. 2.5. Fabrication of MC Microgel Beads. 1 g of alginic acid was solubilized in 100 mL of PBS to form the precursor solution. The solution was subsequently mixed with Direct Red 80 (DR80), CdTe QDs, PEI, MB, or TH to a final concentration of 1% (w/v) to form different solution mixtures. The solution mixtures were driven by syringe pumps (Model Lsp01-2A, Baoding Longer Precision Pump Co., Ltd., China), and were pumped through different metal needles before the mixtures merged into one single stream in the microfluidic nozzle. The nozzle was fabricated by tapering a capillary using a micropipette puller (P-97, Sutter Instrument, Inc., USA). The tip of the tapered capillary was polished to a desired diameter using sand paper, and was hydrophobized using n-octadecyltrimethoxysilane before use. The solution mixtures flowing toward the microfluidic nozzle were ionized using a high-strength electric field, which was formed between the nozzle and a ground circular electrode connected to a high-voltage power supply. A tapered tip driven by the electrostatic force was formed before microdroplets were generated from jet breakup. The microdroplets with multiple compartments were dropped into a collection bath containing a 1% (w/v) CaCl2 solution to form MC microgel beads. The beads were retrieved by centrifugation for 10 min at a relative centrifugal force of 1000 × g, where g is the gravitational constant. In addition to microgel beads generated using Alg alone, beads were generated using 1% (w/v) Alg/carmellose blends with different Alg weight percentages. The beads with 25, 50, and 75 wt % of Alg were designated as AC25, AC50, and AC75, respectively. On the other hand, the beads generated using Alg alone were designated as AC100. 2.6. Characterization of Microgel Beads. The morphology of the microgel beads was observed with an optical microscope (Nikon Eclipse

80i, Nikon, Japan) and a scanning electron microscope (SEM, Leo 1530, Germany). For the latter, the microgel beads were sputter-coated with gold before SEM analysis. After encapsulation of the QDs in the microgel beads, changes in the PL quantum efficiency of the QDs were determined by analyzing the intensities of the fluorescence signals from the beads in optical microscopic images using the open-source image analysis software, ImageJ. 2.7. Fourier Transform Infrared (FT-IR) Spectroscopy. FT-IR was performed using an FT-IR spectrometer (Spectrum 2000, Perkin Elmer, USA) at ambient conditions. The potassium bromide (KBr) disk technique was used for analysis. Spectra were obtained at a resolution of 2 cm−1, and were reported as an average of 16 scans. 2.8. Cytotoxicity Assay. HEK293 cells and 3T3 mouse fibroblasts were cultured in DMEM supplemented with 10% FBS, 100 UI/mL penicillin, 100 μg/mL streptomycin, and 2 mM L-glutamine. 24 h before the assay, cells were seeded separately in a 96-well plate at an initial density of 5000 cells per well, and were incubated under a humidified atmosphere of 5% CO2 at 37 °C. During the experiment, the growth medium in each well was replaced with 100 μL of the fresh cell culture medium containing Alg, carmellose, or microgel beads, at a desired concentration. After 5-h incubation at 37 °C, the medium was replaced with the fresh growth medium. The CellTiter 96 AQueous nonradioactive cell proliferation assay (MTS assay; Promega Corp., USA) was performed, according to the manufacturer’s instructions, either immediately or after 24 h of post-treatment incubation. Cell viability (%) in each well was determined as previously reported.22 2.9. Determination of the Drug Encapsulation Efficiency. MB and TH were used as model drugs. Encapsulation of the model drugs in the microgel beads was performed as described in section 2.5. The CaCl2 solution was collected from the collection bath. The concentrations of unencapsulated MB and TH were determined at 665 and 360 nm, respectively, using a UV−vis spectrophotometer (Varian, Inc., USA) as previously described to determine the drug encapsulation efficiency of the microgel beads.23 2.10. Drug Release Evaluation. After fabrication of drugencapsulated microgel beads, 5 mL of PBS (pH = 7.4) was added to the beads. At a preset time interval, 1 mL of the buffer solution was removed for testing, and was replaced with 1 mL of PBS. The amount of the drug released from the beads was determined using a UV−vis spectrophotometer (Varian, Inc., USA). The cumulative drug release was calculated using the following formula t

cumulative drug release (%) =

∑t = 0 mt m∞

× 100%

(1)

where mt is the mass of the drug released from the beads at time t, and m∞ is the mass of the drug loaded into the beads. 2.11. Determination of the Water Content. The dried and preweighed microgel beads (0.05 g) were immersed in 100 mL of PBS. At a preset time interval, the beads were retrieved by centrifugation for 1 min at a relative centrifugal force of 10000 × g, followed by the removal of the supernatant. The water content of the beads was calculated using the following formula m − md water content (%) = s × 100% ms (2) where ms and md represent the masses of the swollen and dried beads, respectively. 2.12. Statistical Analysis. All data were presented as the means ± standard deviations of triplicate experiments. Student’s t test was performed to assess the statistical significance. Differences with a p-value < 0.05 were considered to be statistically significant.

3. RESULTS AND DISCUSSION 3.1. Generation of MC Microgel Beads. Alg is an anionic polysaccharide consisting of two monomeric units: α-L-guluronic acid (G) and β-D-mannuronic acid (M) (Figure 1A). In an earlier study, Alg has been used as a vaccine delivery system for diphtheria toxoid.24 Compared to diphtheria toxoid alone, Alg 872

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Figure 1. Fabrication of the MC microgel beads. (A) Chemical structures of the two monomeric units of Alg, and the intermolecular network of Alg formed in the presence of Ca2+. (B) Schematic diagram showing fabrication of MC microgel beads using the microfluidic electrospray technology. (C) Photos showing (i) the outlets for the parallel streams, and (ii) the MC beads generated. The scale bars in i and ii represent 1 cm and 500 μm, respectively. (D) Size distribution of the MC beads. The flow rate of one stream is 250 μL/h, and that of the other is 1000 μL/h. The electric field strength is 7 kV/cm.

Figure 2. Size tunability of the MC microgel beads. (A) Optical images of the MC beads generated at different flow rates of the combined stream: (i) 3000, (ii) 2000, (iii) 1000, and (iv) 500 μL/h. The electric field strength is kept constant at 7 kV/cm. The scale bar is 500 μm. (B) A plot of the size of the beads as a function of the flow rate of the combined stream (*p < 0.05). The electric field strength is 7 kV/cm. (C) Optical images of the MC beads generated at different electric field strengths: (i) 3, (ii) 5, (iii) 7, and (iv) 9 kV/cm. The flow rate of the combined stream is 2000 μL/h. The scale bar is 500 μm. (D) A plot of the size of the beads as a function of the electric field strength. The flow rate of the combined stream is 2000 μL/h. 873

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Apart from the size of the beads, the size of the compartments of the beads can be tuned. To distinguish different compartments of the microgel beads, DR80 is added to one of the two parallel streams during microgel bead fabrication. By keeping the flow rate of the unloaded stream (Qu) constant at 1000 μL/h, the volume percentage of the compartment encapsulating DR80 increases from 30% to 90% when the flow rate of the dye-loaded stream (QL) is changed from 50 to 500 μL/h (Figure 3A,B).

nanoparticles loaded with the toxoid have led to a higher humoral immune response in guinea pigs.24 More recently, inorganic/ organic hybrid Alg/CaCO3 nanoparticles have been examined for co-delivery of doxorubicin hydrochloride (DOX) and paclitaxel (PTX).25 In HeLa cells and MCF-7/ADR cells, the dual drug-loaded nanoparticles have exhibited higher cell inhibitory effects,25 and have shown the potential to be used in combination chemotherapy to overcome multidrug resistance. All these findings provide evidence for the promising potential of Alg in drug delivery. In fact, Alg has an established track record of applications in delivery of cells and drugs,26−32 partially due to its low toxicity, biocompatibility, biodegradability, and ease of hydrogel formation with divalent ions (e.g., Ca2+) which form cross-links in Alg by binding to the guluronic residues (Figure 1A).33 Regarding its favorable properties in drug delivery, Alg is used in this study for fabrication of MC microgel beads. The beads are generated using the microfluidic electrospray technology because this technology allows for production of monodisperse microgel beads in a totally aqueous environment, thereby enhancing the clinical potential of the beads formed. During the fabrication process, a jet formed by the combination of parallel streams of different aqueous dispersed phases is broken into droplets (Figure 1B,C). In the absence of an electric field, formation of droplets is driven mainly by the interplay between the surface tension and gravity. Upon the application of an electric field, the increasing electrostatic force stretches the fluid dispensed through the nozzle, forming a tapered jet. In general, when the magnitude of the electrostatic force is comparable to that of the gravitational force, an unstable fluctuating jet will be formed, resulting in the formation of satellite droplets and hence a broad size distribution of the resultant microgel beads. However, as the electric field strength increases, the electrostatic force, rather than the gravitational force, dominates the pulling force against the surface tension. This forms a stable tapered jet, leading to the formation of more monodisperse droplets. Microscopic evaluation reveals that the diameter of the beads generated (flow rate of the first stream = 250 μL/h, flow rate of the second stream = 1000 μL/h, electric field strength = 7 kV/cm) is in a range 300−400 μm. The average diameter is 358.6 μm, and the polydispersity is approximately 4.77% (Figure 1D). Because of the laminar flow behavior, parallel streams of liquids can first flow toward the microfluidic nozzle without significant mixing, and subsequently merge into a single stream which is sprayed into air under an electric field. Upon ionic gelation with calcium ions, MC microgel beads can be formed which can encapsulate different chemical entities concomitantly in different compartments. 3.2. Tunability of MC Microgel Beads. The size tunability of the beads is a desirable factor in drug delivery because it provides versatility to practical applications. The size of the beads formed can be adjusted by changing the flow rate of the combined stream, and increases from around 290 to 440 μm when the flow rate of the combined stream increases from 500 to 3000 μL/h (Figure 2A,B). Apart from changing the flow rate, the size of the beads can be tuned by changing the electric field strength. In general, the size of the beads decreases as the electric field strength increases. This is evidenced by the observation that, at a constant flow rate of the combined stream, the size of the beads formed at the electric field strength of 7 kV/cm is around 330 μm, which is roughly 4 times as small as the size of those generated at the electric field strength of 3 kV/cm (Figure 2C,D).

Figure 3. Tunability of the size of the compartments of the MC microgel beads. (A) Photos of the MC beads encapsulating DR80. Different flow rates are used for the stream containing DR80: (i) 50, (ii) 100, (iii) 250, and (iv) 500 μL/h. The flow rate of the other stream is 1000 μL/h. The electric field strength is 7 kV/cm. The scale bar is 500 μm. (B) The volume percentages of the DR80-loaded compartment, as controlled by changing the ratio of the flow rate of the dye-loaded stream (QL) to the flow rate of the other stream (QU), which is kept constant at 1000 μL/h. The electric field strength is 7 kV/cm. (C) The average diameters of MC beads, as controlled by changing the QL/QU ratio, in which QU is kept constant at 1000 μL/h throughout the experiment. The electric field strength is 7 kV/cm.

Changing the flow rate of the dye-loaded stream as mentioned above has no apparent effect on the size of the beads formed (Figure 3C). This is due to the fact that the relatively high Qu dominates over QL in determining the flow rate of the resultant combined stream, which influences the size of the beads generated. 3.3. Co-Delivery of Incompatible Drugs. The capacity of the beads to co-deliver multiple incompatible drugs is evaluated using QDs as a model. QDs are tiny nanocrystals of a semiconducting material.34−38 Their fluorescence emission is easily quenched by different chemical agents including metal ions,39−41 874

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ACS Applied Materials & Interfaces carbon nanotubes,42 and organic dyes.43 This property may not be favorable for imaging applications, but has made QDs an excellent model to demonstrate the capacity of our MC beads to co-deliver incompatible drugs. In this study, CdTe QDs are synthesized in an aqueous environment using MPA as a stabilizer. The PL quantum efficiency of the QDs is estimated to be around 23%. The structural features of the QDs as well as the microgel beads are investigated by FT-IR (Figure 4).

MPA is chemisorbed as a carboxylate onto the CdTe surface. These signals are also found in the spectrum of the beads alone at 1640 and 1465 cm−1, due to the COO− asymmetric and symmetric stretching vibrations of Alg, respectively. As our QDs are stabilized with MPA, the S−H stretching vibration of MPA is supposed to be detected at around 2500 cm−1.45,46 However, this signal is absent in the spectra of the QDs and QD-loaded microgel beads. This suggests that MPA is bound to the surface of the QDs through Cd−S bonding. The quenching agent selected to pair with QDs is PEI, which is one of the extensively studied nonviral vectors for gene delivery.18,47−49 As shown in Figure 5A, the PL quantum efficiency of the QDs immediately drops from 23% to less than 1% when the QDs are mixed with a PEI solution. PEI and CdTe QDs can, therefore, be used as a model pair of incompatible drugs. To confirm that any subsequent quenching of QDs loaded into the beads is solely due to PEI, the quenching effects of all constituents of the microgel system on QD fluorescence are evaluated. No significant loss of the PL quantum efficiency of the QDs is observed when the QDs are mixed with the Alg or CaCl2 solutions (Figure 5A). This shows that constituents of the microgel beads have no immediate quenching effects on QD fluorescence. After 3 h of treatment with Alg, reduction in the PL quantum efficiency of the QDs is negligible. CaCl2 is found to cause aggregation of the QDs, leading to a drop in the PL quantum efficiency from around 20% to 2%. However, when microgel beads are fabricated in real practice, they are retrieved by centrifugation immediately after gelation. Even when QDs are encapsulated in the beads, the time of their exposure to CaCl2 should be short, and the quenching effect of CaCl2 on their PL quantum efficiency should be limited. This is confirmed by the observation that, 3 h after fabrication of microgel beads encapsulating QDs, changes in the intensity of fluorescence emission from the QDs are negligible (Figure 5B).

Figure 4. FT-IR spectra of (A) CdTe QDs, (B) plain microgel beads, and (C) QD-loaded microgel beads.

The spectrum of the MPA-capped CdTe QDs shows signals at 1569 and 1395 cm−1. These signals are attributed to the asymmetric and symmetric stretching vibrations of the COO− group, respectively.44 The presence of these signals suggests that

Figure 5. Treatment of QDs with different reagents. (A) Changes in the PL quantum efficiency of the QDs treated for different time periods with solutions of different reagents. The untreated QDs are used as the control. The UV−vis absorption and PL spectra of the untreated QDs are shown in the top right-hand corner. (B) Representative phase-contrast and fluorescence images of the QD-loaded microgel beads captured at different time points. The scale bar is 500 μm. (C) (i, ii) Optical and (iii, iv) SEM images of the beads half-loaded with the QDs. The flow rate of the stream containing QDs is 250 μL/h. The flow rate of the other stream is 1000 μL/h. The electric field strength is 7 kV/cm. The scale bar is 200 μm. 875

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Figure 6. Encapsulation of QDs in microgel beads. (A) Representative (i, iii, and v) phase-contrast and (ii, iv and (vi) fluorescence images of (i and ii) SC beads encapsulating QDs alone, (iii and iv) SC beads encapsulating both PEI and QDs, and (v and vi) MC beads encapsulating PEI and QDs in separate compartments. The images are taken immediately after QD encapsulation. The scale bar is 500 μm. (B) The PL quantum efficiency of the QDs immediately after co-encapsulation with PEI in SC and MC microgel beads. The QDs encapsulated in the SC beads in the absence of PEI are used as the control. (C) Representative (i, iii, and v) phase-contrast and (ii, iv, and vi) fluorescence images of the MC beads encapsulating QDs and PEI in separate compartments. The images are taken (i, ii) 0 h, (iii, iv) 3 h, and (v, vi) 6 h after QD encapsulation. The scale bar is 500 μm. (D) Time-dependent changes in the PL quantum efficiency of the QDs co-encapsulated with PEI in the MC and SC beads. The QDs encapsulated in the SC beads in the absence of PEI are used as the control.

Figure 7. Drug encapsulation and release from the microgel beads. (A) The drug encapsulation efficiency of microgel beads with different Alg weight percentages (*p < 0.05). (B) Changes in the water content of the beads as a function of time at pH 7.4. (C) Drug release profiles of (i) MB and (ii) TH from microgel beads with different Alg weight percentages.

MC beads half-loaded with QDs are generated using the microfluidic electrospray technology (Figure 5C). After being loaded into the MC beads in which QDs and PEI are encapsulated in separate compartments, the PL quantum efficiency of the QDs remains around 20%, which is comparable to that of the QDs encapsulated in microgel beads in the absence of PEI (Figure 6A,B). On the contrary, QDs encapsulated in PEI-loaded single-compartment

(SC) beads, in which PEI and QDs are in direct contact, exhibit a remarkable reduction in their PL quantum efficiency. This demonstrates the success of the MC beads in co-delivering potentially incompatible agents. Fluorescence emission from the QDs coencapsulated with PEI in MC beads is detectable even after 6−7 h (Figure 6C,D). Although a time-dependent reduction in fluorescence emission is observed, partially due to the diffusion of PEI 876

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ACS Applied Materials & Interfaces molecules from one compartment to another to cause the gradual quenching of QD fluorescence, the MC beads significantly prolong the time of observable fluorescence emission from the QDs, as compared to the SC beads in which the PL quantum efficiency of the QDs drops to less than 1% right after QD/PEI co-encapsulation. 3.4. Co-Delivery of Multiple Drugs at Tunable Release Rates. While the capacity to co-deliver incompatible drugs is

important for further development of multidrug therapy, the tunability of the release rates of individual co-delivered drugs is also desired as this allows the therapy to be tailored to fit practical needs. MB and TH are used in this section as model drugs. These two model drugs have been extensively used in the literature for drug release experiments.50,51 Furthermore, the maximum UV− vis absorption peak of MB is at a wavelength where absorption by TH is minimal (Figure S1); this is also true for the other way around. The absorption interference between the two model drugs is therefore negligible when the release profiles of the two model drugs co-delivered by the microgel beads are examined. To tune the release rates of individual drugs co-delivered by the MC beads, the polymer blending technique, in which carmellose is blended with Alg to modulate the drug release capacity of individual compartments of the MC beads, is employed. Carmellose is a derivative of cellulose.52 Due to its chemical stability,53,54 water solubility,55 biodegradability,56 and biocompatibility,53,54 carmellose has been used in various applications, ranging from food production57 to pharmaceutical applications.58 Owing to the long history of human use of carmellose, the safety of the beads generated from the Alg/carmellose blend can be secured. Above 80% of the model drugs can be encapsulated in the microgel beads (Figure 7A). No significant difference in the drug encapsulation efficiency is found among microgel beads with different Alg weight percentages. Despite this, due to the hydrophilic nature of carmellose, the microgel swelling capacity, as indicated by the water content of the beads,59 is enhanced when the Alg weight percentage in the beads decreases (Figure 7B). Changes in the microgel swelling capacity may affect the release rate of the encapsulated drug because water in the hydrogel matrix is the medium through which the drug diffuses.59 The release profiles of MB and TH from microgel beads with different Alg weight percentages are shown in Figure 7C. An increase in the weight percentage of Alg in the microgel beads causes a decrease in the drug release rate. Compared to TH, the release rate of MB is much faster. This is attributed to the smaller molecular weight of MB (319.85 Da) as compared to that of TH (480.90 Da).

Figure 8. Photos of the MC microgel beads generated using Alg/carmellose blends, and the profiles of drug release from those beads: (A) AC100 for both the MB-loaded and TH-loaded compartments; (B) AC25 for both the MB-loaded and TH-loaded compartments; (C) AC100 for the MB-loaded compartment, and AC25 for the TH-loaded compartment. The scale bar is 500 μm.

Figure 9. Viability of (A) HEK293 cells and (B) 3T3 mouse fibroblasts after 5-h treatment with the microgel beads and their constituents, (i) with or (ii) without a 24-h post-treatment incubation prior to the MTS assay. 877

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The smaller molecular weight of the drug may lead to a higher rate of drug diffusion.60 The ability to fine-tune the release profiles of individual drugs co-encapsulated in MC beads is demonstrated in Figure 8. By changing the Alg weight percentage in different compartments of the MC beads, the release patterns of the co-delivered drugs can be modulated. For instance, by decreasing the Alg weight percentage in the beads as a whole from 100% to 25%, the overall drug release pattern is maintained, but the release rate of each of the two co-delivered drugs increases. On the other hand, by generating MC beads using AC100 for one compartment and AC25 for the other, MB is tuned to be released in phase with TH. The capacity to tune the release profiles of individual codelivered drugs allows the multidrug therapy to be tailored to the needs of individual patients. Here it is worth mentioning that when the drug release profiles are tuned, properties of the drug (including surface charge, molecular weight, aqueous solubility, and affinity to the hydrogel matrix) may play a role; however, the effect of the drug properties on drug release profiles can be accommodated by changing the Alg weight percentage in different compartments. As the ability of the MC beads to separately control the drug release rates in individual compartments results merely from modulation of the microgel composition, this approach to co-deliver multiple drugs at individual release rates is basically applicable to microgel beads at all sizes. Finally, the toxicity of a drug carrier is an important factor to be considered for practical applications. The toxicity of the microgel beads is evaluated in HEK293 cells and 3T3 mouse fibroblasts. No apparent loss of cell viability is observed after treatment with different concentrations of Alg, carmellose, and microgel beads (Figure 9). This finding is consistent with the low cytotoxicity of Alg and carmellose as previously reported.23,61 The high safety profiles of the microgel beads indicate that the beads have potential to be developed as safe and well-tolerated carriers for drug delivery.

AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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REFERENCES

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4. CONCLUSIONS This study reports the fabrication of MC microgel beads, using the microfluidic electrospray technology and the polymer blending technique, for co-administration of multiple drugs. The beads cannot only co-deliver incompatible drugs, but also allow the release rates of individual co-delivered drugs to be tuned separately. The versatility provided by the MC beads is expected to bring new opportunities for multidrug therapy. In this study, only Alg and its polymer blend with carmellose are adopted to generate MC beads; however, the method of microgel fabrication is expected to be applicable to other materials (such as chitosan,62 κ-carrageenan,63 and gellan gum64) that may form hydrogels through ionic gelation. We, therefore, hope that this study not only provides a tunable microgel system for tailoring multidrug regimens to meet the needs of individual patients, but also sheds light on methods for development of carriers for multidrug co-delivery in the future.

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S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsami.5b10274. UV−vis absorbance spectra of MB and TH in deionized water (PDF) 878

DOI: 10.1021/acsami.5b10274 ACS Appl. Mater. Interfaces 2016, 8, 871−880

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ACS Applied Materials & Interfaces

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DOI: 10.1021/acsami.5b10274 ACS Appl. Mater. Interfaces 2016, 8, 871−880