Fe₃O₄@nanogel via UOx/HRP initiated surface polymerization for pH sensitive drug delivery

Abstract

This communication describes a new strategy to fabricate a nanogel layer around magnetic nanoparticles by surface free-radical polymerization triggered by the cascade reaction of urate oxidase and horseradish peroxidase, which showed high loading capacity, pH-responsive drug release and low cytotoxicity.

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For the biomedical application of inorganic nanoparticles, the creation of few atomic layers of organic molecules or polymer chains around the surface of nanoparticles has been well developed to endow them with biocompatibility and interactive functions at the new surface.^1–7 A newly emerging approach for the modification of inorganic nanoparticles is the creation of nanosized polymer gel layers around the inorganic cores.^8–13 The core–shell polymer nanogels can provide the additional performances, such as high loading capacity within the gel networks and stimuli-responsiveness via function polymer chains and cleavable crosslinkers.^14–18 In general, the formation of polymer gel layers around nanoparticles are realized via the bulk polymerization of vinyl- or initiator-modified nanoparticles with monomers and few crosslinkers or precipitation polymerization.^17,19–23 However, these methods have some limitations, such as harsh reaction conditions, use of catalysts and surfactants, residual toxicity and so on. Recently, enzyme-mediated radical polymerization as an excellent and mild approach has been successfully employed to construct bioactive hydrogels.^24–30 Furthermore, our group has utilized enzymes-modified nanoparticles as the “active” core, which can generate surface radical to initiate polymerization and form smart microgel.³¹ The surface enzyme-initiated polymerization can form core–shell hybrid gel with uniform size due to the short life time of surface generated radical, which can therefore inhibit the excessive polymerization and aggregation in bulk phase.^31,32

Herein, we develop a new strategy to fabricate multifunctional core–shell nanogels by the surface free-radical polymerization initiated by the cascade reaction of urate oxidase (UOx) and horseradish peroxidase (HRP). UOx is a peroxisomal enzyme which can catalyze the oxidation of uric acid to allantoin.³³ However, most microbial UOx present unsatisfactory enzyme activity and stability under physiological conditions.^33–36 Therefore, the catalytic activity of UOx would have a great degree reduction after the surface-initiated polymerization of nanogels, which will benefit the biomedical applications of the enzyme-initiated system because of no cytotoxicity from the further radical after polymerization. Magnetic nanoparticles (MNP_S) are selected as a very intriguing candidate of inorganic cores due to their easily separation characterization, while the intrinsic magnetic property can benefit the potential applications, such as magnetic-targeting, thermotherapy, and imaging.^37–41 The surface UOx around MNPs catalyzes the oxidative transformation of uric acid to allantoin with concomitant hydrogen peroxide (H₂O₂), which reacts with ACAC via HRP catalysis to generate ACAC radicals at the interface to initiate the polymerization and gelation. The functional monomer methacrylic acid is selected to provide pH-responsive property within gel network. The pH-responsive hybrid nanogels exhibit not only a high drug loading capacity but also a pH-controllable drug releasing.

The proposed preparation of the multifunctional hybrid nanogels, is shown in Scheme 1. The magnetic Fe₃O₄ nanoparticles (MNPs) were synthesized according to the literature.⁴² And the average diameter of the obtained MNPs was approximate 40 nm. The surface of the obtained MNPs was first amino-functionalized by treatment with 3-aminopropyltriethoxysilane (APTES) to yield MNPs–APTES. Subsequently, the MNPs–APTES was further functionalized by reacting with succinic anhydride in DMF to achieve carboxylic functionalized nanoparticles MNPs–COOH. The zeta potential of the Fe₃O₄ nanoparticles with different surface modification was displayed in Fig. S2 (in ESI†). The carboxyl modified MNPs was then activated by 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide/N-hydroxysuccinimide (EDC/NHS). After wash, the acquired magnetite nanoparticles were combined with excess enzyme UOx and HRP solution (molar ratio of UOx : HRP = 1 : 1) to covalently attach the UOx and HRP simultaneously to the carboxyl-modified MNPs. Afterwards, the as-prepared UOx/HRP-MNPs were redispersed in deionized water for further use after washed three times. The hydrogel was fabricated around the interface of UOx/HRP-MNPs with the addition of coating precursor solution, which was composed of PEGMA₃₆₀ (4%, v/v), methacrylic acid (MAA, 2%, v/v), PEGDA₂₅₀ (1%, v/v), ACAC (1%, v/v) and uric acid (0.02 M). The core–shell nanogels were acquired by magnetic separation from the solution after 12 h reaction and re-dispersed in water for further characterization.

Scheme 1 The scheme of the preparation of the core–shell nanogels by UOx/HRP mediated hydrogelation.

The surface gelation mechanism around MNPs is described as follows. The UOx surround the surface of the MNPs first catalyze oxidation of uric acid in the coating solution to generate of allantoic acid and H₂O₂, which interacts with the adjacent HRP on the MNPs and ACAC immediately in the bulk media to initiate the subsequent generation of ACAC radicals. Finally, the forming radicals around the UOx/HRP-MNPs can trigger the gelation of the hydrogel layer on the MNPs via the surface radical polymerization. The diffusion of ACAC radicals within the forming gel networks and short life time can prevent the likelihood of the polymerization in bulk phase. The easily deactivation of UOx during enzymatic reaction also reduce the bulk polymerization due to the low amount of radical. At the same time, the low residue activity of UOx after polymerization can reduce the cytotoxicity of the final nanogels in biological system that can interact with uric acid (Fig. S4 in ESI†). In consequence, the core–shell nanogels with high monodispersity were collected other than the aggregated bulk hydrogel.

As shown in Fig. 1, the morphology and size distribution of the collected core–shell nanogels were investigated using SEM, TEM and DLS measurements, respectively. The comparison of the SEM images of the air-dried Fe₃O₄ nanoparticles and the magnetic core–shell nanogels in Fig. 1a and b reveal that the nanogels are in a state of spherical shape and smooth surface with a diameter of 180 nm in dry state, while the magnetite Fe₃O₄ nanoparticles are coarse in surface. In addition, the nanogels exhibit more excellent monodispersity than the MNPs. The contrast of the surface appearance of MNPs and the core–shell structures indicates that the hydrogel layer has been successfully coated onto the surface of magnetite Fe₃O₄ nanoparticles. Furthermore, TEM images also demonstrate the differences of MNPs and nanogels in structure, as shown in Fig. 1d and S1 (in ESI†). The core–shell structure is quite explicit with a black core and a light contrast gray shell, while the initial MNPs have not this structure feature (Fig. S1 in ESI†), and hydrogel shell of the nanogels is about 10 nm in dry state according to the TEM image. The nanoparticles of both MNPs and nanogels were further characterized by size distribution (Fig. 1c). The average diameters of the MNPs and nanogels are approximate 80 nm and 200 nm, respectively, which is a little higher than that of SEM and TEM result in dried sample. The about 60 nm thickness in wet state also confirms the formation of apparent core–shell structure. Therefore, the pH responsive nanogels shells have been successfully coated on the MNPs through UOx/HRP initiated interfacial radical polymerization.

Fig. 1 (a and b) The SEM images of the single Fe₃O₄ nanoparticles and the core–shell nanogels. (c) The size distribution of Fe₃O₄ and the nanogels in buffer. (d) The TEM image of the core–shell nanogels.

These core–shell nanogels can be employed as carriers for intracellular drug delivery in future cancer chemotherapy. To verify the feasibility, positively charged doxorubicin hydrochloride (DOX·HCl) was chosen as a model drug to evaluate the effect of the drug carrier systems. The DOX was loaded by mixing aqueous solutions of the core–shell nanogels and DOX·HCl in Tris·HCl buffer at pH 8.0, which was kept in dark place for 24 h. Free DOX was removed by magnetic separation and the loading content of DOX was determined by UV/Vis spectroscopy by measuring the absorbance of DOX at 480 nm. The negatively charged nanogels showed extremely high drug loading content and drug-loading efficiency (DLC = 88.1%, DLE = 66.0%). This high drug loading may be attributed to the interaction of negatively charged nanogels with the positively charged DOX.

To evaluate the feasibility of the DOX-loading nanogels for cancer therapy, the cytotoxicity of DOX-loaded nanogels, nanogels and free DOX toward SH-SY5Y cells were tested by MTT assay in Fig. 2. The SH-SY5Y cells were treated with varying concentrations of free and nanogels-loading DOX (0–20 μg mL⁻¹) for 24 h and 48 h, respectively. At first, all the data showed that the blank core–shell nanogels without DOX exhibited >85% cell viability, suggesting unobvious cytotoxic effect of original nanogels on the cells. The cell viabilities with various form DOX and different concentration were investigated and shown in the Fig. 2. The cytotoxicity of DOX-loaded nanogels or pure DOX was observed by the significant growth inhibition of SH-SY5Y cells with the various drugs. The cell viability is about 28.4% when the 20 μg mL⁻¹ nanogels-loading DOX was treated with SH-SY5Y cells for 48 h. The low cell viability suggested that DOX-loaded nanogels were efficiently taken up by the SH-SY5Y cells and released DOX in acidic microenvironment inside the cancer cells, which inhibited the cellular proliferation. At the same time, the cell viability is about 14.5% when the 20 μg mL⁻¹ free DOX was treated with SH-SY5Y cells for 48 h. The lower cell viability of free DOX is reasonable because of the easily permeability of free DOX into cell. Our DOX-loaded nanogels are the better system than free one due to the selective uptake by tumour cells and pH responsive release to get better therapeutic efficiency.

Fig. 2 Cytotoxicities of nanogels, DOX-loaded nanogels, and free DOX toward SH-SY5Y cells after incubation for 24 h (a) and 48 h (b). Error bar represented means of three independent experiments.

The pH-triggered cumulative drug release of DOX from DOX-loaded nanogels in vitro was investigated at various pH values. Specifically, a certain amount of DOX-loaded nanogels including 600 μg DOX in dialysis bag was placed in PBS at different pH values of 4.5, 5.8, and 7.4. Fig. 3 shows the significant dependence of DOX release upon acidic pH values, and the release rate increased as the decreased pH value. As shown in Fig. 3, at endo-lysosomal acidity of pH 4.5, 55.5% DOX was released after 48 h incubation at 37 °C. In contrast, only 39.1% and 14.7% DOX were released at pH 5.8 and 7.4, respectively. The release rate of DOX enhanced with the decreased pH value, in the order of pH 4.5 > pH 5.8 > pH 7.4. The elevated released amount of DOX in low pH was attributed to the weakened electrostatic interaction between the positively charged DOX with negatively charged COO⁻ groups of MAA in nanogel networks, which amount was inversely proportional to proton concentration according to the ionization equilibrium of methacrylic acid. Therefore, the pH-sensitive nanogels can possibly reduce the leakage of their embedded drug during transportation to the tumour regions while accelerate the drug release toward cancer cell in tumour region with more acidic microenvironment. The high drug loading content could benefit cancer therapy by the long-time cumulative release of DOX.

Fig. 3 In vitro DOX releases from DOX-loaded nanogels at 37 °C in PBS at pH 4.5, 5.8, and 7.4.

The cellular uptake and intracellular drug release of the DOX-loaded nanogels was furtherly confirmed by flow cytometry analysis (FACS) and Laser Confocal Scanning Microscopy (LCSM). The samples of SH-SY5Y cells were treated by DOX-loaded nanogels with 0.5–24 h incubation time. The quantitative characterization of various form drugs uptaked into the cells over time was done by FACS analyses. The flow cytometric histogram of the SH-SY5Y cells with DOX-loaded nanogels confirmed the high cellular uptake of the DOX-loaded nanogels, whose fluorescence intensity gradually increased with the increasing incubation time (Fig. S5 in ESI†). The cells nuclei were stained with DAPI for elevated contrast and clarity. As shown in Fig. 4, after 0.5 h incubation, tiny bright DOX fluorescence was observed in the cells, indicating the successful endocytosis of the DOX-loaded nanogels. Significantly, a stronger red fluorescence signal was found in the cell nuclei after 2 h incubation, demonstrating the fast cellular internalization of nanogels and speedy release of DOX which migrated into cell nuclei within 2 h. Similar to FACS analysis, the intensity of the red fluorescence was enhancing with an increase in the incubation time, suggesting that more DOX reached to cell nuclei. Meanwhile, considerable red fluorescence was also observed in the cytoplasm of cells.

Fig. 4 LCSM images of the SH-SY5Y cells incubated with the DOX-loaded nanogels for respective 0.5 h, 2 h, 8 h and 24 h. In each panel, the images from left to right display cell nuclei stained using DAPI (blue), DOX fluorescence in cells (red) and their overlap. CLSM images of SH-SY5Y cells after 0.5–24 h incubation with DOX-loaded nanogels (DOX: 5 μg mL⁻¹).

Conclusions

In conclusion, we have demonstrated a mild and stable approach of UOx/HRP-mediated interfacial radical polymerization for fabricating the pH sensitive core–shell nanogels. The core–shell hybrid nanogels presented in this work incorporated the different functionalities of both the magnetic core and pH-responsive shell into one particle with fascinating properties, which was served as an intelligent drug delivery carrier. The DOX loaded nanogels exhibited extremely high drug loading content and achieved high DOX release rate under lower pH conditions. Additionally, the high cellular proliferation inhibition against SH-SY5Y cells was also displayed in cell experiment. Consequently, with the excellent biocompatibility and efficient intracellular drug delivery, primarily in the nucleus, the core–shell hybrid nanogels presented could act as a potential platform for drug delivery system and tumor therapy.

Acknowledgements

This research was supported by the National Natural Science Foundation of China (No. 21274111, 51473123 and 51402215), the Recruitment Program of Global Experts.

Notes and references

1. R. Ghosh Chaudhuri and S. Paria, Chem. Rev., 2011, 112, 2373–2433.
2. C. Tassa, S. Y. Shaw and R. Weissleder, Acc. Chem. Res., 2011, 44, 842–852.
3. R. Cheng, F. Meng, C. Deng, H.-A. Klok and Z. Zhong, Biomaterials, 2013, 34, 3647–3657.
4. P. D. Thornton and A. Heise, J. Am. Chem. Soc., 2010, 132, 2024–2028.
5. Y. Wang, S. Gao, W.-H. Ye, H. S. Yoon and Y.-Y. Yang, Nat. Mater., 2006, 5, 791–796.
6. X. Zhang, P. Yang, Y. Dai, P. a. Ma, X. Li, Z. Cheng, Z. Hou, X. Kang, C. Li and J. Lin, Adv. Funct. Mater., 2013, 23, 4067–4078.
7. X. Wang, Y. Yang, P. Gao, F. Yang, H. Shen, H. Guo and D. Wu, ACS Macro Lett., 2015, 4, 1321–1326.
8. M. Karg and T. Hellweg, J. Mater. Chem., 2009, 19, 8714–8727.
9. K. S. Soppimath, D. W. Tan and Y. Y. Yang, Adv. Mater., 2005, 17, 318–323.
10. W. Wu, J. Shen, P. Banerjee and S. Zhou, Biomaterials, 2010, 31, 7555–7566.
11. W. Wu, N. Mitra, E. C. Yan and S. Zhou, ACS Nano, 2010, 4, 4831–4839.
12. X. Wang, D. Niu, Q. Wu, S. Bao, T. Su, X. Liu, S. Zhang and Q. Wang, Biomaterials, 2015, 53, 349–357.
13. X. Wang, D. Niu, P. Li, Q. Wu, X. Bo, B. Liu, S. Bao, T. Su, H. Xu and Q. Wang, ACS Nano, 2015, 9, 5646–5656.
14. C.-Y. Quan, J.-X. Chen, H.-Y. Wang, C. Li, C. Chang, X.-Z. Zhang and R.-X. Zhuo, ACS Nano, 2010, 4, 4211–4219.
15. J. Zhang, F. Yang, H. Shen and D. Wu, ACS Macro Lett., 2012, 1, 1295–1299.
16. W. Wu, J. Shen, Z. Gai, K. Hong, P. Banerjee and S. Zhou, Biomaterials, 2011, 32, 9876–9887.
17. N. Singh, A. Karambelkar, L. Gu, K. Lin, J. S. Miller, C. S. Chen, M. J. Sailor and S. N. Bhatia, J. Am. Chem. Soc., 2011, 133, 19582–19585.
18. X.-q. Zhao, T.-x. Wang, W. Liu, C.-d. Wang, D. Wang, T. Shang, L.-h. Shen and L. Ren, J. Mater. Chem., 2011, 21, 7240–7247.
19. H. Jiang and F.-J. Xu, Chem. Soc. Rev., 2013, 42, 3394–3426.
20. P. Yan, N. Zhao, H. Hu, X. Lin, F. Liu and F.-J. Xu, Acta Biomater., 2014, 10, 3786–3794.
21. P. B. Zetterlund, S. C. Thickett, S. Perrier, E. Bourgeat-Lami and M. Lansalot, Chem. Rev., 2015, 115, 9745–9800.
22. T. M. Ruhland, P. M. Reichstein, A. P. Majewski, A. Walther and A. H. Müller, J. Colloid Interface Sci., 2012, 374, 45–53.
23. G. L. Li, H. Möhwald and D. G. Shchukin, Chem. Soc. Rev., 2013, 42, 3628–3646.
24. R. Shenoy and C. N. Bowman, Biomaterials, 2012, 33, 6909–6914.
25. T. Su, Z. Tang, H. He, W. Li, X. Wang, C. Liao, Y. Sun and Q. Wang, Chem. Sci., 2014, 5, 4204–4209.
26. T. Su, D. Zhang, Z. Tang, Q. Wu and Q. Wang, Chem. Commun., 2013, 49, 8033–8035.
27. W. Wang, J. Qian, A. Tang, L. An, K. Zhong and G. Liang, Anal. Chem., 2014, 86, 5955–5981.
28. J. Gao, W. Zheng, D. Kong and Z. Yang, Soft Matter, 2011, 7, 10443–10448.
29. L. Dong, Q. Miao, Z. Hai, Y. Yuan and G. Liang, Anal. Chem., 2015, 87, 6475–6478.
30. J. Li, Y. Kuang, J. Shi, J. Zhou, J. E. Medina, R. Zhou, D. Yuan, C. Yang, H. Wang, Z. Yang, J. Liu, D. M. Dinulescu, M. Daniela and B. Xu, Angew. Chem., Int. Ed., 2015, 54, 13307–13311.
31. Q. Wu, X. Wang, C. Liao, Q. Wei and Q. Wang, Nanoscale, 2015, 7, 16578–16582.
32. R. Shenoy, M. W. Tibbitt, K. S. Anseth and C. N. Bowman, Chem. Mater., 2013, 25, 761–767.
33. X. Wu, C. C. Lee, D. M. Muzny and C. T. Caskey, Proc. Natl. Acad. Sci. U. S. A., 1989, 86, 9412–9416.
34. X. Liu, M. Wen, J. Li, F. Zhai, J. Ruan, L. Zhang and S. Li, Appl. Microbiol. Biotechnol., 2011, 92, 529–537.
35. Y. Huang, Y. Chen, X. Yang, H. Zhao, X. Hu, J. Pu, J. Liao, G. Long and F. Liao, Biotechnol. Appl. Biochem., 2015, 62, 137–144.
36. J. Feng, L. Wang, H. Liu, X. Yang, L. Liu, Y. Xie, M. Liu, Y. Zhao, X. Li and D. Wang, Appl. Microbiol. Biotechnol., 2015, 99, 7973–7986.
37. S. Laurent, D. Forge, M. Port, A. Roch, C. Robic, L. Vander Elst and R. N. Muller, Chem. Rev., 2008, 108, 2064–2110.
38. M. Mahmoudi, S. Sant, B. Wang, S. Laurent and T. Sen, Adv. Drug Delivery Rev., 2011, 63, 24–46.
39. O. Veiseh, J. W. Gunn and M. Zhang, Adv. Drug Delivery Rev., 2010, 62, 284–304.
40. Y. Pan, X. Du, F. Zhao and B. Xu, Chem. Soc. Rev., 2012, 41, 2912–2942.
41. J. Ge, Y. Hu, M. Biasini, W. P. Beyermann and Y. Yin, Angew. Chem., Int. Ed., 2007, 46, 4342–4345.
42. S. Xuan, F. Wang, Y.-X. J. Wang, C. Y. Jimmy and K. C.-F. Leung, J. Mater. Chem., 2010, 20, 5086–5094.

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Footnote

† Electronic supplementary information (ESI) available: Experimental details and supplementary figures. See DOI: 10.1039/c6ra06331d

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