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Polypeptide-Grafted Nanodiamonds for Controlled Release of Melittin to Treat Breast Cancer Haiwang Lai, Fan Chen, Mingxia Lu, Martina H. Stenzel,* and Pu Xiao* Centre for Advanced Macromolecular Design, School of Chemistry, The University of New South Wales, Sydney, Australia S Supporting Information *

ABSTRACT: A peptide vector consisting of nanodiamonds (NDs) and PEGylated polyglutamic acid (ND@PLGPEG-co-PLGA) has been designed and developed. The negative charges at the surface of the vector were exploited to bind a positively charged peptide drug melittin via electrostatic interaction. The surface was saturated when the weight ratio of ND@ PLGPEG-co-PLGA to melittin (MEL) was 5 to 1. The desorption of melittin from the surface was controlled by pH, with almost no melittin released from the nanoparticles under physiological pH conditions in 2 days. However, steady release was detected in an acidic environment. The preserved structure and activity of bound melittin were demonstrated by the HPLC and 2D MCF-7 cell culture models, respectively. The bound melittin exhibited improved toxicity toward MCF-7 cells dependent on the concentration of MEL in NDs. Our results suggested that the negatively charged polymercoated NDs were able to release the cargo upon exposure to breast cancer cells.

T

coupled to allow the nanoparticles to be applied in diagnostic imaging application.14 Incorporation of peptide drug into nanoparticles via electrostatic attraction has rarely been reported despite the fact that formation of a polyelectrolyte complex between polymer and protein is widely applied in the literature.15 This is due to their lower surface charge that is essential to retain the peptide drug in the polymer matrix by electrostatic interaction.7 MEL in contrast has a relatively high density of positive charges, which are mostly concentrated toward one end of the peptide (Scheme S1).1 The positive charges in MEL can be utilized to incorporate MEL into negatively charged nanoparticles, and the charge effects can be reversed to trigger the release of MEL from the carrier. Nanomaterials with an intrinsic ability to be used for imaging purposes are increasingly being loaded with drugs and implemented for therapy and real-time readout on the treatment efficacy. Among these nanomaterials, nanodiamonds (NDs) are emerging as a promising platform for nanomedicine because of their biocompatibility and stable fluorescence. NDs have been extensively used to deliver peptides and proteins,16−19 which are predominantly absorbed via electrostatic interaction. However, the employment of unmodified NDs may encounter difficulties due to colloidal instability. Furthermore, proteins or peptides deposited on naked NDs may form multilayers at the surface which may lead to a loosely bound outer layer, conformational changes, and dehydration of

he principal toxin component of bee venom (BV), melittin (MEL), is an α-helical hemolytic toxin with MW of 2846 Da and contains 26 amino acid residues, including 50% hydrophobic residues and 6 positive charges (Scheme S1).1 Each MEL chain contains two α-helical segments and ultimately takes the form of a bent rod.1 MEL has demonstrated anticancer activity against several cancer cells, including breast cancer, prostate cancer, and gastrointestinal cancer, due to its potent membrane-disrupting activity.2−5 Despite the cytotoxicity to a variety of tumor cells, MEL is also toxic to normal cells due to its hemolytic property and vulnerable to enzymatic degradation.5,6 Therefore, a suitable delivery vehicle is required that can deliver MEL safely to tumor sites.5,7 Peptide or protein drugs can either be covalently bound or physically encapsulated in a drug carrier. However, the former method requires the presence of reactive functional groups. Moreover, chemical modification of the drug can cause deactivation.7 Physical encapsulation of peptide drugs can be achieved by traditional means by using a nanoparticle matrix that is compatible with the drug or by taking advantage of the surface charges. The potent hemolytic ability of MEL prevents the employment of liposomes as a delivery vehicle.8 Previous attempts to package MEL into nanoparticles include lipidic assembly,9,10 perfluorocarbon-core surfactant-coated nanoparticles,11,12 and rigid-cored polymeric nanoarchitecture.13 However, none of these methods were developed upon the controlled dissociation and triggered drug release mechanism, which is of great importance in minimizing side effects during circulation and efficient drug release in tumor cells. Moreover, imaging agents including florescent tags are necessary to be © XXXX American Chemical Society

Received: May 28, 2017 Accepted: July 10, 2017

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DOI: 10.1021/acsmacrolett.7b00389 ACS Macro Lett. 2017, 6, 796−801

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ACS Macro Letters Scheme 1. Reaction Route to ND@PLGPEG-co-PLGA

the proteins.20 Polymer coating of NDs integrates the merits of polymer and NDs into a single entity, not only improving the colloidal stability of the core material but also enabling the placement of therapeutic drugs or targeting ligands at the shell polymer.21−27 For instance, the application of cationic polymercoated NDs to deliver plasmid DNA has been demonstrated to enhance the transfection efficiency.21 Positively charged MEL requires the use of negatively charged polymers to create strong electrostatic charges. This can be achieved by coating the NDs with polypeptides based on glutamic acid that have received attention due to their biocompatibility, biodegradability, and intriguing secondary structure.28,29 Herein, we report the polypeptide-functionalized NDs as a promising carrier to deliver MEL. The surface of NDs was designed to be coated with a negatively charged PEGylated polyglutamic acid. The electrostatic interaction between charged synthetic polymer and charged peptide or protein drug has been demonstrated to be an efficient way to condense the drug into nanoparticles.30,31 Upon pH conversion from basic or neutral to acidic, the negative charges in the polymer shell reduced, leading to an attenuation of electrostatic interaction with the peptide drug which was eventually released from the nanoparticles. The conformation change induced by pH also plays an important role in the binding and release of MEL from the polypeptide-coated NDs. Surface grafting of synthetic polypeptide to the surface of nanoparticles has garnered scientific interest in the design of functional biocompatible interfaces due to the biocompatibility, biodegradability, and structural similarity of synthetic polypeptides to natural peptides.28 Generally, surface grafting of polypeptide can be achieved by grafting from and grafting onto approaches. The former method relies on the attachment of primary amine onto the surface and subsequent ring-opening polymerization of N-carboxyanhydride (NCA). However, this technique is limited by the poor control over the molecular weight and molecular weight distribution resulting from difficulties in controlling the initiation and propagation rates.28 Moreover, the comprehensive characterization of the grafted polypeptide remains a challenge. Nonetheless, this is not a problem in the grafting onto approach as polymer is

synthesized and characterized prior to surface conjugation. As the structure and composition of the grafted polymer are essential to understand the interaction between the surface and therapeutics to be delivered, we adopted the grafting onto method to synthesize the polypeptide-coated NDs in the current work. The EDC/NHS coupling was employed here instead of the single end functional group coupling technique such as click chemistry32 due to the facile reaction and purification for EDC/NHS coupling (Scheme 1). The synthesis of the polymer started with the synthesis of benzyl glutamate NCA (BLG-NCA) and subsequent ring-opening polymerization initiated by tert-butyl amine (Scheme 1). The synthesis and polymerization of BLG-NCA were widely reported in the literature to prepare polyglutamic acid based material.33,34 The NMR spectrum of the BLG-NCA (Figure S1) showed that the monomer was successfully synthesized with a high purity. After the polymerization, the chain length of the resulting polymer PBLG was calculated to be 140 by comparing the peak areas for terminal methyl groups at 1.3 ppm and the methine proton at 4.6 ppm (Figure S2). However, it did not correspond to the feed ratio and conversion, which can be ascribed to the coexisting normal amine mechanism and activated monomer mechanism.35,36 The polymerization proceeded via a normal amine mechanism, while the activated monomer mechanism acted as side polymerization which influenced the kinetics and molecular weight distribution.36 The SEC of the PBLG140 revealed a relatively broad distribution (Đ = 1.59, Figure S3). Deprotection of the benzyl groups with HBr in trifluoroacetic acid solution afforded PLGA140. The successful removal of benzyl groups was indicated by the disappearance of peaks at 7.3 ppm in the NMR spectrum (Figure S2). The carboxyl groups were then partially PEGylated by DCC coupling. Based on NMR integral analysis, it was calculated that 110 out of the 140 repeating units in the PLGA precursor were derivatized with PEG by comparing the peak area for methyl groups in PEG (3.2 ppm) and all the glutamic acid side methylene peaks (from 1.5 to 2.6 ppm). The obtained polymer was denoted as PLGPEG110-coPLGA30. 797

DOI: 10.1021/acsmacrolett.7b00389 ACS Macro Lett. 2017, 6, 796−801

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in cargo loading and structure preservation. The adsorption of ND@P with MEL was first attempted in pH 7.4 buffer solution with a weight ratio of 10:1 (ND@P:MEL). HPLC was used to quantify the binding efficiency after mixing. It was found that at this ratio NDs were in excess; the binding of MEL with NDs was so tight that the surface charges were slightly neutralized, and no free MEL was detected in the supernatant (Figure 1). Based on the grafting density of polymer chains on the ND surface, we calculated that on each ND@P cluster the number of MEL molecules loaded was 966 (SI, grafting density calculation), which corresponded to 17 MEL molecules per polypeptide chain. When the ratio of MEL was increased, more MEL was adsorbed at the surface. The anionic carboxyl groups were gradually neutralized as illustrated in Table 1 and Figure 1A, and the surface was eventually saturated. As illustrated in Figure 1B, when the ratio of ND@P to MEL was 10−1 and 6− 1, no free MEL was found in the supernatant after mixing. However, a characteristic MEL peak in the HPLC graph (Figure 1B,d) appeared after adsorption onto NDs when the ratio increased to 5−1, indicating that the surface of NDs was saturated with MEL. As the repulsion between ND particles were attenuated by the decrease of charges, an increase in size was found as the ratio of MEL increased (Table 1). The release of MEL from the NDs was investigated in a similar manner as adsorption assays. As shown in Figures S5− S8, the bound MEL was eluted out with the same retention time as unbound MEL, which suggested the binding with NDs did not alter the structure of MEL. There is a significant difference in desorption in different pH conditions (Figure 2). Specifically, more than 80% of the adsorbed MEL was released from the surface in 10 h under acidic environment, and about 95% was released in 2 days. The half-life for adsorbed MEL is about 3 h. However, less than 5% of MEL (10−1 ratio) was desorbed in physiological pH of 7.4 in 2 days. The pKa of the carboxyl groups of PLGPEG110-co-PLGA30 was determined to be 5.4 (Figure S9), which can explain the different affinities of ND@P to MEL in different pH conditions. The carboxyl groups were calculated to be 37% dissociated when pH was 5.0 by the Henderson−Hasselbalch equation, and they were almost all ionized when pH was 7.4.37 The pI value of MEL is 12.06, and MEL is always positively charged when pH is below 12.06.38 Therefore, a decrease in pH value resulted in the neutralization of the carboxyl groups and gave rise to less

The functionalization of NDs with polypeptide was accomplished by a grafting onto strategy via the EDC/NHS coupling reaction between the amino groups on NDs and carboxyl groups in PLGPEG-co-PLGA (Scheme 1). The aminofunctionalized NDs were modified from carboxylated NDs with ethylenediamine by EDC/NHS coupling in water. The polymer-coated NDs displayed a negative charge of −36.6 mV, opposite to the +31.2 mV of the amino-functionalized NDs (Table 1), suggesting the successful conjugation of Table 1. DLS and Cytotoxicity Characterization of NDs and MEL Sample ND-NH2 ND@PLGPEG110-co-PLGA30 (ND@P) ND@MEL:10−1 ND@MEL:6−1 ND@MEL:5−1 MEL a

Hydrodynamic diameter (nm)a

Zeta potential (mv)

IC50 of MEL (μg mL−1)

155 220

+31.2 −36.6

NA NA

263 291 305 NA

−27.3 −13.1 −3.56 NA

2.0 2.1 1.3 3.2

z: average mean.

carboxyl-containing polymer onto the surface. The quantification by TGA revealed that 11% of polymer content was grafted (Figure S4). However, the control experiment without adding EDC and NHS gave 4% of polymer adsorption, which verified the covalent conjugation of the polymer. The slight increase in size (from 155 to 220 nm) was probably caused by a small extent of cross-linking during the conjugation procedure (Table 1). TEM images (Figure S12) showed that the particles were aggregates of small primary NDs. Based on the DLS and TEM sizes, the number of individual ND particles in each ND@P cluster was estimated to be around 22. It was calculated that about 55 polymer chains were grafted onto each cluster (SI, grafting density calculation). For drug loading and release study, a variety of nanoscale vehicles have been reported based on the electrostatic interaction of charged synthetic polymers with peptide or protein drugs.15 We assume that the combination of NDs with synthetic polymers as peptide vectors may bring improvement

Figure 1. (A) Zeta potential of NDs obtained after mixing with different ratio of MEL; (B) HPLC chromatograms of (a) MEL solution before adsorption (10−1); (b) supernatant after adsorption (10−1); (c) supernatant after adsorption (6−1); and (d) supernatant after adsorption (5−1). 798

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biocompatibility as a drug carrier. MEL is a very promising candidate for cancer chemotherapy due to its nonspecific cytolytic attack of all lipid membranes and lower possibility for cancer cells to develop resistance.5 The IC50 of MEL was found to be 3.2 μg mL−1. Moreover, the IC50 for BV against MCF-7 cells was reported to be about 6 μg mL−1.2 MEL is the principal toxin component of BV; therefore, it is reasonable that the IC50 of MEL is lower than BV raw extract. Our number is also consistent with the results reported previously (3.64 μg mL−1).38 In all cases of ND@MEL, lower IC50 values were found than the unbound MEL. The reason may be that free MEL exerts anticancer effect by a cell membrane disrupting mechanism. However, the bound MEL is internalized into cancer cells via endocytosis and affects various intracellular targets including the membranes of internal organelles.41,42 There is also a trend that the higher ratio of MEL to NDs leads to higher toxicity (Figure 3B). Flow cytometry was used to compare the cell uptake of NDs with different MEL content. The polymer shell was labeled with Cy5. Figure 4 shows the fluorescence profile and relative uptake

Figure 2. Cumulative release of MEL versus time under pH 5.0 and 7.4 conditions.

negative zeta potential (Figure S10), which weakened the binding with MEL. This can also account for the faster release from higher MEL content particles (Figure 2). With more MEL bound, the NDs were less negatively charged (Figure 2) which led to weaker binding and faster desorption. Moreover, a decrease in α-helix content from 87% to 18% was observed when the pH of PLGPEG110-co-PLGA30 solution (Figure S11) increased from 5.0 to 7.4. The transition from α-helix to random coil breaks the intramolecular hydrogen bonds to fold the α-helix and allows the amide bonds to form intermolecular hydrogen bonds with surrounding molecules including MEL, which also partially accounts for the stronger binding to MEL in higher pH condition.39 This conformation change is also induced by the increased charge repulsion when the pH increased. This pH-controlled release was very desirable for minimizing the side effect and protecting the peptide drug from metabolism during circulation in blood.15 The change of other parameters (e.g., the concentration of salts or other cationic compounds such as amino acids) can also lead to the release of the adsorbed MEL and will be investigated in the future research. NDs and polypeptides are attracting increasing attention as a drug delivery platform due to their excellent biocompatibility.28,40 However, the cytotoxicity of the association of NDs and polypeptide still needs to be evaluated. The cell viability study revealed nontoxicity to MFC-7 cells for the ND@P up to a concentration of 62.5 μg mL−1 for 48 h (Figure 3A). These results suggested that ND@P particles display desirable

Figure 4. Flow cytometry analysis of Cy5-labeled nanoparticle uptake into MCF-7 cells after 3 h incubation.

rate after 3 h incubation. It is obvious that the almost saturated ND vectors were internalized significantly more than the slightly loaded vector. It has been reported that surface charge plays a pivotal role in the cellular uptake process of nanoparticles.43,44 The cell surface is dominated by proteoglycans which consist of a core protein anchored to the membrane and linked to highly anionic glycosaminoglycan side chains.44 The negatively charged cell surface tends to facilitate the attachment of cationic or less negatively charged particles. Hence, the more neutralized NDs which contained more MEL show higher cell uptake than those more negative NDs.

Figure 3. Cytotoxicity study of (A) ND@P and (B) MEL and complexes of NDs and MEL with different ratios. 799

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(6) Torchilin, V. P.; Lukyanov, A. N. Peptide and protein drug delivery to and into tumors: challenges and solutions. Drug Discovery Today 2003, 8, 259−266. (7) Du, A. W.; Stenzel, M. H. Drug Carriers for the Delivery of Therapeutic Peptides. Biomacromolecules 2014, 15, 1097−1114. (8) Huang, C.; Jin, H.; Qian, Y.; Qi, S.; Luo, H.; Luo, Q.; Zhang, Z. Hybrid Melittin Cytolytic Peptide-Driven Ultrasmall Lipid Nanoparticles Block Melanoma Growth in Vivo. ACS Nano 2013, 7, 5791− 5800. (9) Lundquist, A.; Wessman, P.; Rennie, A. R.; Edwards, K. Melittin− Lipid interaction: A comparative study using liposomes, micelles and bilayerdisks. Biochim. Biophys. Acta, Biomembr. 2008, 1778, 2210− 2216. (10) Zetterberg, M. M.; Reijmar, K.; Pränting, M.; Engström, Å.; Andersson, D. I.; Edwards, K. PEG-stabilized lipid disks as carriers for amphiphilic antimicrobial peptides. J. Controlled Release 2011, 156, 323−328. (11) Pan, H.; Soman, N. R.; Schlesinger, P. H.; Lanza, G. M.; Wickline, S. A. Cytolytic peptide nanoparticles (‘NanoBees’) for cancer therapy. Wiley Interdiscip. Rev. Nanomed. Nanobiotechnol. 2011, 3, 318−327. (12) Jallouk, A. P.; Moley, K. H.; Omurtag, K.; Hu, G.; Lanza, G. M.; Wickline, S. A.; Hood, J. L. Nanoparticle Incorporation of Melittin Reduces Sperm and Vaginal Epithelium Cytotoxicity. PLoS One 2014, 9, e95411. (13) Misra, S. K.; Ye, M.; Kim, S.; Pan, D. Defined Nanoscale Chemistry Influences Delivery of Peptido-Toxins for Cancer Therapy. PLoS One 2015, 10, e0125908. (14) Winter, P. M. Perfluorocarbon Nanoparticles: Evolution of a Multimodality and Multifunctional Imaging Agent. Scientifica 2014, 2014, 10. (15) Gu, Z.; Biswas, A.; Zhao, M.; Tang, Y. Tailoring nanocarriers for intracellular protein delivery. Chem. Soc. Rev. 2011, 40, 3638−3655. (16) Shimkunas, R. A.; Robinson, E.; Lam, R.; Lu, S.; Xu, X.; Zhang, X.-Q.; Huang, H.; Osawa, E.; Ho, D. Nanodiamond−insulin complexes as pH-dependent protein delivery vehicles. Biomaterials 2009, 30, 5720−5728. (17) Swiecicki, J.-M.; Tailhades, J.; Lepeltier, E.; Chassaing, G.; Lavielle, S.; Mansuy, C. Peptide-coated nanoparticles: Adsorption and desorption studies of cationic peptides on nanodiamonds. Colloids Surf., A 2013, 431, 73−79. (18) Kuo, Y.; Hsu, T.-Y.; Wu, Y.-C.; Chang, H.-C. Fluorescent nanodiamond as a probe for the intercellular transport of proteins in vivo. Biomaterials 2013, 34, 8352−8360. (19) Moore, L.; Gatica, M.; Kim, H.; Osawa, E.; Ho, D. Multi-protein Delivery by Nanodiamonds Promotes Bone Formation. J. Dent. Res. 2013, 92, 976−981. (20) Aramesh, M.; Shimoni, O.; Ostrikov, K.; Prawer, S.; Cervenka, J. Surface charge effects in protein adsorption on nanodiamonds. Nanoscale 2015, 7, 5726−5736. (21) Zhang, P.; Yang, J.; Li, W.; Wang, W.; Liu, C.; Griffith, M.; Liu, W. Cationic polymer brush grafted-nanodiamond via atom transfer radical polymerization for enhanced gene delivery and bioimaging. J. Mater. Chem. 2011, 21, 7755−7764. (22) Huynh, V. T.; Pearson, S.; Noy, J. M.; Abboud, A.; Utama, R. H.; Lu, H.; Stenzel, M. H. Nanodiamonds with Surface Grafted Polymer Chains as Vehicles for Cell Imaging and Cisplatin Delivery: Enhancement of Cell Toxicity by POEGMEMA Coating. ACS Macro Lett. 2013, 2 (3), 246−250. (23) Zhao, L.; Nakae, Y.; Qin, H.; Ito, T.; Kimura, T.; Kojima, H.; Chan, L.; Komatsu, N. Polyglycerol-functionalized nanodiamond as a platform for gene delivery: Derivatization, characterization, and hybridization with DNA. Beilstein J. Org. Chem. 2014, 10, 707−713. (24) Lai, H.; Lu, M.; Lu, H.; Stenzel, M. H.; Xiao, P. pH-Triggered release of gemcitabine from polymer coated nanodiamonds fabricated by RAFT polymerization and copper free click chemistry. Polym. Chem. 2016, 7, 6220−6230. (25) Zhao, J.; Lai, H.; Lu, H.; Barner-Kowollik, C.; Stenzel, M. H.; Xiao, P. Fructose-Coated Nanodiamonds: Promising Platforms for

Eventually, the higher cellular uptake gave lower IC50 of encapsulated MEL. In summary, we have prepared ND@PLGPEG110-co-PLGA30 through multiple but still facile reactions as a vector for melittin (MEL). The ND@PLGPEG110-co-PLGA30 particles possess negative zeta potential, enabling the immobilization of positively charged MEL. The affinity can be tuned by changing the pH value of the dispersion media, resulting in a pHtriggered release. The bound MEL displayed enhanced toxicity to MFC-7 cells than the free MEL in a concentrationdependent manner. More MEL encapsulated resulted in lower IC50, which is due to the higher cell uptake resulting from the more neutralized surface. Our results demonstrated the potential application of ND@PLGPEG110-co-PLGA30 as protein and peptide vectors for the treatment of breast cancers.

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsmacrolett.7b00389. Additional information regarding the synthesis and characterization of grafted polymer; HPLC investigation of adsorption and release of surface-bound MEL; pKa measurement of carboxylic acids and surface zeta potential measurement versus pH; circular dichroism of the synthesized polypeptide and TEM images of ND@P (PDF)

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected]. *E-mail: [email protected]. ORCID

Martina H. Stenzel: 0000-0002-6433-4419 Pu Xiao: 0000-0003-3077-451X Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS P.X. acknowledges funding from the Australian Research Council’s Discovery Early Career Researcher Award (DE140100318). M.H.S. acknowledges funding from the Australian Research Council’s Discovery Project (DP140100240).

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REFERENCES

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