Xiaorui Li,
Pengcheng Du and
Peng Liu*
State Key Laboratory of Applied Organic Chemistry and Institute of Polymer Science and Engineering, College of Chemistry and Chemical Engineering, Lanzhou University, Lanzhou 730000, China. E-mail: pliu@lzu.edu.cn; Fax: +86-931-8912582; Tel: +86-931-8912582
First published on 17th October 2014
Core–shell microspheres have attracted intense interest as drug delivery system (DDS) because of their integrated advantages of the core and shell materials. Here the structural effects, such as the crosslinking degree of the cores and the thickness of the polyelectrolyte complex shells, on the dipyridamole (DIP) loading and release performance were investigated in detail for the first time with the core–shell poly(methacrylic acid)@(chitosan/alginate)n microsphere as a drug carrier model, fabricated by encapsulating poly(methacrylic acid) (PMAA) nanogels in the layer-by-layer (LbL) engineered chitosan/alginate (CS/AL) multilayer shells. The core–shell microspheres with two-bilayer chitosan/alginate shells (PMAA@(CS/AL)2) were selected for the in vitro controlled release of the water-insoluble anticancer drug (doxorubicin (DOX)) in simulated body fluids (SBF). After the encapsulation, the DOX-loading capacity increased from 32.18% to 40.12% and the effect of the media pH values on the cumulative release from the PMAA@(CS/AL)2 was more remarkable than the core material, indicating the encapsulation with the polyelectrolyte multilayer shells was favorable for the drug loading and pH-responsive controlled release. Furthermore, the polyelectrolyte multilayer shells also improved the cytocompatibility of the drug carriers. This understanding will lead to better design of smart core–shell DDS for controlled release.
The stimuli-responsive polymeric shells could be coated onto various cores via the in situ polymerization,3,6 microfluidic,9 or layer-by-layer (LbL) assembly technique.10,11 The LbL assembly technique, via stepwise deposition of opposite charged polyelectrolytes onto a core, has been acknowledged as a convenient and versatile method for the core–shell microspheres.12 Compared with other relevant techniques, the drug-loading capacity and the controlled release performance of the LbL engineered core–shell drug-carriers could be conveniently and precisely controlled by tailoring the assembly parameters such as layers of adsorption, ionic strength, and pH of dipping solution.13
As for the core materials, the polymeric cores might expand upon the external stimuli or drug-loading,14,15 whereas the volumes of the mesoporous inorganic cores remain constant under the same conditions. The volume expansion of the cores might be restricted by the polymeric shells in the core–shell drug-carriers. An efficient approach to overcome the disadvantage is to design the yolk–shell microspheres.16,17 Unfortunately, the synthetic procedure is complicated and only the crosslinked synthetic polymers could be used.
Up to now, many LbL engineered core–shell drug-carriers with responsive polymer nanogels as cores have been reported,18,19 but there is no reference on their structural effect. In the present work, the poly(methacrylic acid) (PMAA) nanogels with different crosslinking degrees were synthesized via the facile distillation–precipitation polymerization in acetonitrile with 2,2′-azodiisobutyronitrile (AIBN) as initiator and divinylbenzene (DVB45) as crosslinking reagent. In order to realize the better controlled releasing performance, the chitosan/alginate multilayer shells ((CS/AL)n) with different thicknesses were coated onto the PMAA nanogels by the LbL assembly technique via electrostatic interaction between the amino groups of CS and the carboxyl groups of AL (Scheme 1). The effect of the cross-linking degrees of the PMAA nanogels and the thicknesses of the polyelectrolyte multilayer shells of the core–shell PMAA@(CS/AL)n microspheres on their drug-loading and pH-responsive controlled release performance were optimized with dipyridamole (DIP) as a model drug. Finally, the optimized core–shell microspheres were used as DDS for the in vitro controlled release of the hydrophobic anticancer drug (doxorubicin (DOX)) in simulated body fluids (SBF).
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Scheme 1 Schematic illustration of the drug loading and controlled release of the PMAA nanogels and the core–shell PMAA@(CS/AL)4 microspheres. |
DVB-45 were purchased from Tianjin Guangfu Chemical Co. Ltd., Tianjin, China and used without further purification. Methacrylic acid (MAA) (Tianjin Chemical Co. Ltd., Tianjin, China) was distilled under vacuum before use. Acetonitrile (analytically pure) was provided by Tianjin no. 3 Chemical Plant, Tianjin, China. AIBN (Tianjin Guangfu Chemical Co. Ltd., Tianjin, China) was recrystallized from ethanol before use. Deionized water was used throughout.
The PMAA nanogels were separated by centrifugation and washing thoroughly with acetonitrile and water and dried at 50 °C in a vacuum oven. The other PMAA nanogels with different crosslinking degrees were synthesized by the same procedure except for altering the DVB45 feeding ratios (10%, 15%, 20%, and 25%) for comparison.
A 10 mL buffer solution containing 10 mg DIP-loaded PMAA nanogels or DIP-loaded PMAA@(CS/AL)n microspheres was transferred into dialysis tube with a weight cutoff 14000 and immersed into 150 mL buffer solution at pH 1.8 or 7.4 at 37 °C, respectively. 5.0 mL dialysates were taken out at certain time intervals and analyzed with UV-vis spectrometry to detect the release rate. Furthermore, 5.0 mL fresh buffer solution was added after each sampling to keep the total solution volume constant. The cumulative release is expressed as the percentage of the released drug over time.
Higuchi model: Mt = kt1/2 |
Korsmeyer–Peppas model: Mt/M∞ = ktn |
The zeta potentials of the PMAA nanogels and the core–shell PMAA@(CS/AL)4 microspheres were determined with Zetasizer Nano ZS (Malven Instruments Ltd., UK), by adjusting the pH of sample solutions with NaOH or HCl solution.
The mean hydrodynamic diameters (Dh) of the PMAA nanogels and the core–shell PMAA@(CS/AL)n microspheres were determined by dynamical mode (dynamic light scattering (DLS)) on the “Light Scattering System BI-200SM, Brookhaven Instruments” device equipped with the BI-200SM Goniometer, the BI-9000AT Correlator, Temperature Controller and the Coherent INOVA 70C argon ion laser at 20 °C. DLS measurements are performed using 135 mW intense laser excitation at 514.5 nm and at a detection angle of 90° at 25 °C. The pH of sample solutions was adjusted with NaOH or HCl solution.
The drug loading and controlled release performance of the PMAA nanogels and the core–shell PMAA@(CS/AL)n microspheres were tracked by a Lambda 35 UV/vis spectrometer (Perkin-Elmer Instruments, USA) at room temperature.
The mean hydrodynamic diameters (Dh) of the swollen PMAA nanogels (PMAA-5, PMAA-10, PMAA-15, PMAA-20, and PMAA-25) were 364.9 ± 4.9, 375.3 ± 5.8, 394.4 ± 8.7, 505.7 ± 11.7, and 522.5 ± 4.2 nm from DLS, respectively. Comparing the diameters determined by TEM and DLS, which represented the particle sizes in their dried and swollen states respectively, the volume swelling ratios (volume ratio of the swollen state and dried state) of the PMAA-5, PMAA-10, PMAA-15, PMAA-20, and PMAA-25 nanogels could be calculated to be 227.4%, 193.8%, 133.7%, 130.0%, and 119.2%, respectively. The PMAA-5 nanogels exhibited the most remarkable swelling behavior. It indicated that the swelling properties of the PMAA nanogels prepared with lower feeding ratios of crosslinker were better than those with higher ones, due to that the polymer chains are more rigid and the crosslinked network is more compact when higher feeding ratios of crosslinker are used. The higher crosslinking degree might result in the lower drug-loading capacity, due to the two main factors: the drug-loading space and the diffusion of the drug molecules into the hydrogels.23 Thus the PMAA-10 nanogels with a medium crosslinking degree were chosen for the further experiments.
The Dh of the PMAA-5, PMAA-10 and PMAA-15 nanogels gradually decreased from 395, 443 and 467 nm to 323, 336 and 349 nm with decreasing the pH value from 10.0 to 3.0, respectively (Fig. 2(A)). Evidently, the increased swelling at higher pH values was induced by ionization of the carboxylic groups in the nanogels.24 The protonation/deprotonation of the pendent carboxylic acid groups also influences the swelling properties of the weak acidic PMAA chains, which are hydrophobic at lower pH media and hydrophilic at higher pH media.25 At higher pH values, the PMAA chains are deprotonated, resulting in the swelling of the nanogels; at lower pH values, the PMAA chains are protonated and the nanogels are in their collapsed state. So the PMAA nanogels are rigid microspheres, showing the smaller diameter (Fig. 2(A)).
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Fig. 2 pH dependence of the Dh of the PMAA nanogels with different crosslinking degrees (A) and the core–shell PMAA@(CS/AL)4 microspheres. |
The loaded drugs could be discharged by the shrinking of the nanogels when they are exposed to the acidic media, such as accumulation in the tumor sites from the normal tissues. So the phase transition induced by the protonation/deprotonation could be used as a pH-controlled switching for the pH-stimuli responsive controlled drug release.26
The zeta potentials of the core–shell PMAA@(CS/AL)n microspheres were measured to track the LbL assembly (Fig. 3(A)). The half bilayer numbers correspond to the CS deposition and the whole bilayer numbers represent the AL adsorption, the zero layer number corresponds to the PMAA-10 nanogel cores with the zeta potential of −37.6 mV. After the first CS layer was deposited, the zeta potential was +28.0 mV, indicating that the deposition of CS could completely cover the surface of the PMAA-10 cores. Then the zeta potentials changed to −21.9 mV after the adsorption of AL over the PMAA@CS1. With increasing the adsorbed layers, the symmetrically alternating zeta potentials indicated that the polyelectrolyte multilayer shells had been successfully deposited on the PMAA-10 cores by the LbL assembly.27
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Fig. 3 Zeta potentials (A) and the mean hydrodynamic diameters (B) of the PMAA-10 nanogels encapsulated by CS/AL multilayer shells with different thicknesses (PMAA@(CS/AL)n). |
The successful assembly could also be revealed by the TEM analysis (Fig. 1). Obviously, as the bilayer number of the polyelectrolyte multilayer shell increased, the average diameter of the (PMAA@(CS/AL)n (n = 1, 2, 3, 4)) microspheres increased gradually from 302.0 ± 4.4 nm of the PMAA-10 cores to 311.7 ± 4.2, 315.6 ± 3.4, 322.7 ± 3.6 and 334.9 ± 3.3 nm of the core–shell (PMAA@(CS/AL)1, PMAA@(CS/AL)2, PMAA@(CS/AL)3, and PMAA@(CS/AL)4) microspheres, respectively. It indicated that the thickness of the polyelectrolyte shells could be efficiently tuned by controlling the numbers of the CS/AL bilayer coated.
The mean hydrodynamic diameters (Dh) of the PMAA-10 cores and the core–shell PMAA@(CS/AL)1, PMAA@(CS/AL)2, PMAA@(CS/AL)3 and PMAA@(CS/AL)4 microspheres were 375.3, 415.7, 456.7, 479.8, and 498.2 nm, respectively (Fig. 3(B)). The Dh increased 40.4, 41.0, 23.1, and 18.4 nm while coating one, two, three, and four polyelectrolyte bilayers respectively, also indicating that the increased shell thickness with increasing the numbers of the polyelectrolyte bilayers coated. Their volume swelling ratios increased from 191.9% of the PMAA-10 cores to 237.2%, 303.0%, 328.7%, and 329.2% of the core–shell (PMAA@(CS/AL)n (n = 1, 2, 3, 4)) microspheres, respectively. It meant that the polyelectrolyte multilayer shells exhibited the higher swelling ability due to their non-covalent interlayer interaction. And the swelling ability tended to be steady with the CS/AL bilayer number more than 2. It demonstrated that the less additional polyelectrolyte could be adsorbed with more polyelectrolyte complex layers coated, as shown as the decreased amplitude in the zeta potential results (Fig. 3(A)).
The zeta potentials of the PMAA-10 nanogels and the core–shell PMAA@(CS/AL)4 microspheres at media with different pH values are compared in Fig. 4. In the whole range, the zeta potentials of the PMAA-10 nanogels were negative and their absolute values increased with increasing the pH values, due to that more neutral carboxylic acid groups of the PMAA nanogels deprotonated into the negative carboxylate ions in the higher pH media. As for the core–shell PMAA@(CS/AL)4 microspheres, their zeta potentials changed from positive to negative near pH 5, indicating that the core–shell PMAA@(CS/AL)4 microspheres have pH-dependent charge-conversional feature, which might be favor to the cellular internalization in the acidic tumor extracellular environments (as acidic as pH 5.7).28 Thus, their negatively charged surface in the neutral or basic media can maintain their stealth character during blood circulation, spontaneously and precisely trigger responsive drug release or enhance interaction between nanocarriers and targeting cells at the targeting tumor sites.
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Fig. 4 Zeta potentials of the PMAA-10 nanogels and the core–shell PMAA@(CS/AL)4 microspheres at different pH values. |
The zeta potential gap between the PMAA-10 nanogels and the core–shell PMAA@(CS/AL)4 microspheres was 53.8 mV at pH 3.0. As the media pH values increased, the gaps decreased rapidly to only 0.3 mV at pH 8.0 due to the polyelectrolyte multilayer shells coated. It could be ascribed that the CS/AL could form a dense shell to hide the influence of the charges from the surface beneath.29 It showed that the polyelectrolyte multilayer shells might act as a switch to tune the drug release from the PMAA cores.
As for the core–shell PMAA@(CS/AL)n microspheres, the drug-loading capacities at pH 4.0 at 25 °C were 58.76, 55.26, 51.28, and 49.78 mg g−1 for the PMAA@(CS/AL)1, PMAA@(CS/AL)2, PMAA@(CS/AL)3 and PMAA@(CS/AL)4, respectively. All the data were higher than that of the PMAA-10 nanogels (45.30 mg g−1) under the same condition, indicating that the encapsulation with the polyelectrolyte complex shells could enhance the drug-loading capacity due to the electrostatic interaction between the amino group-containing drugs and the carboxylic acid groups of AL,33 as well as the hydrogen bonds between the drug and the polyelectrolytes. However, the drug-loading capacities decreased with increasing the thickness of the polyelectrolyte complex shells. The coated polyelectrolyte shells might have two opposite effects on the drug-loading. The drug-loading capacity could be enhanced via the interaction between drug and the polyelectrolytes via the electrostatic interaction or hydrogen bonds. But also, contrastingly, the volume expansion of the PMAA nanogels during the drug-loading is suppressed by the shells coated, thus the drug-loading might be hindered. The combination of these two effects results in the decline in drug-loading capacity with increasing the thickness of the polyelectrolyte shells.
The cumulative DIP release from the DIP-loaded PMAA nanogels (PMAA-5, PMAA-10, and PMAA-15) were 61.97%, 42.97% and 30.50% at pH 1.8 and 55.11%, 25.83% and 17.23% at pH 7.4 at 37 °C, respectively (Fig. 5(A)). The faster releasing of DIP from all the three DIP-loaded PMAA nanogels were achieved at pH 1.8 due to the shrinkage of the PMAA nanogels and the dissolution of DIP in acidic media. It was also found that the cumulative DIP release decreased with increasing their crosslinking degree in the same releasing condition.
More than 80% of the total cumulative DIP release was finished within 750 min for all the DDS regardless of the crosslinking degrees of the nanogels and the pH values of the releasing media. However, the long-term releasing rates were obviously faster in the acidic media than in the basic media. The drug releasing rates and the cumulative release from the DIP-loaded PMAA-5 nanogels within 750 min were similar in both pH 1.8 and pH 7.4 media, indicating that the most drug release (about 80% of the total cumulative DIP release) depended on the shrinkage of the PMAA nanogels with low crosslinking degree, while the long-term releasing was resulted from the dissolution of the drug loaded. As the crosslinking degrees of the PMAA nanogels increased to 10% and 15% of the PMAA-10 and PMAA-15 nanogels, the cumulative release within 750 min in basic media was only about 70% of those in the acidic media. It indicated that the cumulative release had been mainly affected by the shrinkage of the PMAA nanogels and the diffusion of the drug molecules through the nanogels.34
The time dependence of the cumulative DIP release from the drug-loaded PMAA@(CS/AL)n (n = 0, 1, 2, 3, 4) microspheres at pH 1.8 and 7.4 at 37 °C are compared in Fig. 5(B). Different from the releasing performance of the PMAA nanogels, only about 70% of the total cumulative DIP release was finished within 750 min for all the PMAA@(CS/AL)n microspheres regardless of the thicknesses of the polyelectrolyte multilayer shells (CS/AL)n (n = 1, 2, 3, 4) and the pH values of the releasing media. It indicated that the polyelectrolyte (CS/AL)n multilayer shells could slow down the releasing rate of the PMAA-10 nanogels. It was also found that the release rate and the cumulative DIP release from the drug carriers at pH 1.8 were faster and higher than those at pH 7.4 at the same temperature. It could be speculated that the lower release rate might be ascribed to the DIP solubility in different pH media. DIP became insoluble at high pH values because it is an alkaline molecule which could only be dissolved in acidic media. That is to say, the difference in release rates also should be partly attributed to the solubility of the drugs in different pH media.
The cumulative DIP release increased with the increasing of the thicknesses of the polyelectrolyte multilayer shells (CS/AL)n at pH 1.8 due to the fast release of the drug loaded onto the polyelectrolyte multilayer shells, while the cumulative DIP release decreased with increasing the thicknesses of the polyelectrolyte multilayer shells (CS/AL)n in pH 7.4 media. In other words, the polyelectrolyte multilayer (CS/AL)n (n = 1, 2, 3, 4) shells could accelerate the drug release in the acidic media but retard it in the basic media. It indicated that the polyelectrolyte multilayer (CS/AL)n shells were favorite to the pH-responsive controlled release of drugs from the core–shell drug carriers. As shown in Fig. 2(B), the size of the core–shell PMAA@(CS/AL)4 microspheres shrank significantly with increasing the media pH values. The drug molecules were difficult to diffuse throughout the polyelectrolyte (CS/AL)4 shells. Furthermore, the difference between the drug release from the drug carriers at pH 1.8 and 7.4 decreased with increasing the thicknesses of the polyelectrolyte (CS/AL)n shells. Compromising on the drug-loading capacity and the controlled release performance, the two polyelectrolyte bilayer (CS/AL)2 was selected for the core–shell drug carriers.
Due to the acidic intracellular microenvironments such as inside endosomes and lysosomes (pH 4.5–6.5) after nonspecific adsorptive endocytosis,35 the DOX release was compared in the simulated normal tissue (pH 7.4) and the simulated tumor intracellular microenvironments (pH 5.0). The pH-dependent DOX release was observed that the cumulative DOX release from DOX-loaded PMAA@(CS/AL)2 core–shell microspheres at pH 5.0 (52.44% and 72.14%) were higher than those at pH 7.4 (24.91% and 21.31%) within 50 h (Fig. 6). The cumulative DOX release from the PMAA@(CS/AL)2 microspheres at pH 5.0 was faster than at pH 7.4, partly due to that the permeability of the polyelectrolytes complex shell is better at lower pH values than at higher pH values, because that the decreasing pH values weaken the interaction between CS and AL,36 so that the Dh of the core–shell PMAA@(CS/AL)4 microspheres increased from 376 nm to 592 nm with decreasing pH values from 11 to 3, as shown in Fig. 2(B). Additionally, the acid-soluble nature of DOX might be another reason. The pH-dependent DOX release has a great advantage in curing the cancer cells with acidic microenvironments, especially with the polyelectrolyte multilayer shells.
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Fig. 6 Cumulative DOX release from the DOX-loaded PMAA-10 and the DOX-loaded PMAA@(CS/AL)2 core–shell microspheres at pH 5.0 and pH 7.4 at 37 °C. |
The release data were analyzed on the basis of the Korsmeyer–Peppas equation and Higuchi kinetics. The coefficients of correlation (R2) of the Korsmeyer–Peppas equations were better than that of the Higuchi equation, and the n values of the Korsmeyer–Peppas equations were less than 0.5 (Fig. 1S†), so it could be concluded that the DOX releasing mechanism from the DOX-loaded PMAA-10 nanogels or the DOX-loaded PMAA@(CS/AL)2 microspheres at pH 5.0 and 7.4 should follow the non-Fickian model.37
Footnote |
† Electronic supplementary information (ESI) available. See DOI: 10.1039/c4ra05066e |
This journal is © The Royal Society of Chemistry 2014 |