Graphene-based nanosheets for delivery of chemotherapeutics and biological drugs

    Graphene-based nanosheets for delivery of chemotherapeutics and biological drugs Gayong Shim, Mi-Gyeong Kim, Joo Yeon Park, Yu-Kyoung Oh PII: DOI: Reference:

S0169-409X(16)30103-X doi: 10.1016/j.addr.2016.04.004 ADR 12961

To appear in:

Advanced Drug Delivery Reviews

Received date: Revised date: Accepted date:

15 January 2016 17 March 2016 7 April 2016

Please cite this article as: Gayong Shim, Mi-Gyeong Kim, Joo Yeon Park, Yu-Kyoung Oh, Graphene-based nanosheets for delivery of chemotherapeutics and biological drugs, Advanced Drug Delivery Reviews (2016), doi: 10.1016/j.addr.2016.04.004

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Graphene-based nanosheets

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for delivery of chemotherapeutics and biological drugs

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Gayong Shim†, Mi-Gyeong Kim†, Joo Yeon Park, and Yu-Kyoung Oh*

College of Pharmacy and Research Institute of Pharmaceutical Sciences,

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Seoul National University, 1 Gwanak-ro, Gwanak-gu, Seoul 151-742, Korea

* Corresponding author (Tel: 82-2-880-2493; Fax: 82-2-882-2493)

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E-mail addresses: [email protected] (Y.-K. Oh)

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† These authors equally contributed to this work.

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ACCEPTED MANUSCRIPT Abstract

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Graphene-based nanosheets (GNS), including graphenes, graphene oxides and reduced graphene oxides, have properties suitable for delivery of various molecules. With their twodimensional structures, GNS provide relatively high surface areas and capacity for noncovalent - stacking and hydrophobic interactions with various drug molecules. Currently, GNS-based delivery applications extend to chemotherapeutics as well as biological drugs, including nucleic acid drugs, proteins, and peptides. Surfaces of GNS have been modified with various polymers, such as polyethylene glycol and biopolymers, which enhance biocompatibility and increase drug loading. Anticancer drugs are prominent among chemotherapeutic agents tested, and have been loaded onto GNS with relatively high loading capacities compared with other nanocarriers. For enhanced distribution to specific tissues, GNS have been covalently or non-covalently modified with targeting ligands, including folic acid, transferrins, and others. In this review, we cover the current status of GNS for delivery of anticancer chemotherapeutics and biological drugs, with a focus on nucleic acid drugs. Remaining challenges for the application of GNS for drug-delivery systems and future perspectives are also addressed.

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Keywords: graphene-based nanosheets, drug delivery, chemotherapeutics, biological drugs, surface modification, targeting ligands.

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ACCEPTED MANUSCRIPT Contents Introduction

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Physicochemical properties of GNS for drug delivery

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Quenching of fluorescent drugs

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2.2.

Surface modification to enhance biocompatibility Modification with synthetic polymers

3.2.

Modification with natural polysaccharides

3.3.

Modification with proteins

3.4.

Modification with phospholipids

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3.1.

GNS for delivery of chemotherapeutics

Athracycline antitumor antibiotics

4.2.

Anticancer quinolone alkaloids

4.3.

Taxanes

4.4.

Platinum complexes

4.5.

Nitrosourea compounds

4.6.

Pyrimidine analogues

4.7.

Polyphenolic compounds

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Interaction with drug molecules

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4.8.

Quinone compounds

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Other chemotherapeutics

4.10.

Co-delivery of anticancer drugs with photosensitizers

GNS for delivery of biological drugs 5.1.

Delivery of plasmid DNA

5.2.

Delivery of siRNA

5.3.

Delivery of protein and peptide drugs

Targeting moieties for GNS 6.1.

Small chemicals 3

ACCEPTED MANUSCRIPT Monoclonal antibodies and aptamers

6.3.

Proteins and peptides

6.4.

Polysaccharides

Pharmacokinetics and biodistribution

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Current challenges and future perspectives

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6.2.

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ACCEPTED MANUSCRIPT 1. Introduction

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Graphene is a kind of carbon allotrope, exemplified by fullerenes and carbon nanotubes, with a flat single layer of graphite containing stacked layers of carbon atoms with a honeycomb lattice [1]. Graphene is reported to have various extraordinary properties, including a unique structure and mechanical, thermal, and optical characteristics [2,3]. Graphene-based nanosheets (GNS) with nanoscale lateral sizes include graphene nanosheets [4] (Fig. 1A), graphene oxide (GO) nanosheets [5] (Fig. 1B), and reduced graphene oxide (rGO) nanosheets [6] (Fig. 1C). GNS have been intensively exploited as nanocarriers for a variety of drugs, including chemicals [7] and biological drugs [8]. GNS in particular have been proposed as a potential nanocarrier of water-insoluble anticancer chemotherapeutics [9,10] .

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Although the initial physical exfoliation method for preparing graphene single layers—a novel approach using adhesive tape—was very simple, its applicability to large-scale, industrial production of graphene layers is limited [11]. Therefore, chemical exfoliation methods, including oxidation of graphite resulting in GO and rGO [12,13], have become more commonly used for bulk scale production. GO can be dispersed as a single layer in water owing to the hydroxyl (-OH), epoxide (-O-), and carboxylic acid (COOH) functional groups on their surface and edges, but this impairs the planar hydrophobic structure of the graphene surface, resulting in the loss of its intrinsic properties. Reduction of GO to produce rGO restores a surface that mimics the original planar graphene layer with carboxyl groups on their edges [6].

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In this review, we provide an overview of significant advances in graphene-based nanomaterials for delivery of chemical and biological drugs, address current challenges, and highlight future perspectives.

Fig. 1. Structures of GNS. Structures of graphene (A)[4], GO (B)[5], and rGO (C)[6] are presented.

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ACCEPTED MANUSCRIPT 2. Physicochemical properties of GNS for drug delivery 2.1. Interaction with drug molecules

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Among the various types of delivery systems, GNS features a unique two-dimensional planar surface. This two-dimensional feature of GNS confers a higher capacity for drug loading than other types of drug-delivery systems [14]. GO and rGO have been more extensively studied for delivery of drugs than pure graphene. The relative paucity of studies on graphene nanosheets for drug delivery likely reflects the poor water dispersity of these nanosheets. Graphene nanosheets are composed of sp2-hybridized carbon in a honeycomb lattice with nanoscale lateral sizes. Aksay and colleagues reported that the mean thickness of a stack of 140 graphene sheets was ~1.75 nm [14]. The surface area of graphene as a dry powder, measured by the Brunauer-Emmett-Teller method, has been found to be 600–900 m2/g [14]. Although the high surface area of graphene nanosheets makes it possible to load drugs with high capacity, the hydrophobicity of these nanosheets and difficulties associated with handling them present limitations to their application in drug delivery. Because they lack functional groups on their surfaces [4], pure graphene nanosheets mainly interact with drug molecules through  stacking and hydrophobic interactions.

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Although GO and rGO have been actively studied for anticancer drug delivery, the differences in physicochemical properties between them may affect drug interaction patterns and loading capacities (Table 1). GO, an oxidized form of graphene, has sp2-hybridized carbons in the aromatic network, sp3-hybridized carbons with hydroxyl (-OH) and epoxide (O-) groups in the basal plane, and carboxylic acid (COOH) groups at the edges [15]. These functional groups of GO can interact with electronegative atoms of drug molecules through hydrogen bonding. The ionizable carboxylic acid group at the edge of GO allows electrostatic interactions with drug molecules. Moreover, the ionizable carboxylic acid groups can affect pH-dependent release of drug molecules from GO. Despite the presence of these functional groups, the basal plane of GO is mainly composed of polyaromatic networks, which allow drug molecules to bind through  stacking and hydrophobic interactions.

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rGO, a reduced form of GO, has fewer oxygen-containing functional groups than GO [16] and is more hydrophobic. The removal of a significant number of oxygen functional groups in rGO reduces the capacity of drug binding through hydrogen bonding or electrostatic interactions compared with GO. Instead,  stacking and hydrophobic interactions with drug molecules are more dominant in rGO. Overall, the hydrophobicity of rGO is higher than that of GO, but less than that of graphene layers [17]. One caveat in the study of rGO is the extent of GO reduction in a given preparation, which can vary from one laboratory to another. The extent of reduction can affect the density of oxygen functional groups and interaction with drug molecules. A recent study reported that the extent of reduction of rGO could affect the adsorption of serum proteins [18]. GO nanosheets have been used for loading various hydrophobic chemical molecules, including anticancer agents [9,19,20] and photosensitizers. As noted above, these hydrophobic molecules can be loaded onto GNS simply and efficiently by virtue of noncovalent  stacking and hydrophobic interactions. Dai and colleagues [9] first demonstrated that water-insoluble and aromatic drugs such as camptothecin analogues attach to the graphene surface through non-covalent van der Waals interactions. After loading onto 6

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polyethylene glycol (PEG)-modified GO (PEGylated GO), the camptothecin analogue SN38 exhibited enhanced solubility and 2–3 orders of magnitude greater anticancer efficacy than the free SN38 prodrug, irinotecan. Doxorubicin is a potent anticancer agent used clinically as part of various chemotherapy regimens, but it is associated with several adverse effects. To minimize the side effects of doxorubicin, researchers have created various nano-formulations, including liposomes [21,22]. Owing to its aromatic structure, doxorubicin has also been demonstrated to interact with GO mainly via  stacking [19]. In addition to  stacking, hydrogen bonds can also be formed between each -OH group of GO and doxorubicin or the -OH group of GO and the -NH2 group of doxorubicin at neutral pH [20]. The  stacking-based loading mechanism results in GO having a much higher loading efficiency for doxorubicin than other nanocarriers [23]. It is also known that the photosensitizer, chlorin e6, is loaded onto GO via  stacking as well as hydrophobicinteractions [24].

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In addition to small molecules, biomolecules such as proteins [25], peptides [26], and even nucleic acids [27-29] can be delivered by GO [25, 27, 28], rGO [26], and graphene nanosheets [29]. Although nucleic acids are highly hydrophilic macromolecules, the ring structure of nucleobases is capable of  stacking interactions with GO [27,28]. Among nucleic acids, single-stranded DNA (ssDNA) and ssRNA have much higher binding affinity for GO than double-stranded DNA (dsDNA) or ssDNA with secondary and tertiary structure, whose nucleobases are surrounded by the phosphate backbone [28]. The interaction energies of nucleobases with graphene nanosheets exhibit the rank order of G>A≈T>C, similar to the interaction energy order of nucleobases with single-walled carbon nanotubes [29]. Notably, cationic polyethylenimine-functionalized GO has been used for loading small interfering RNA (siRNA) via electrostatic interactions [30].

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Although GO has been more extensively studied for delivery of nucleic acids, a recent study compared the attractive van der Waals interaction energy and loading of single stranded RNA between GO and rGO [31]. This study showed that PEGylated rGO provided higher interaction energy and loading of single stranded RNA than PEGylated GO. Moreover, using computational simulation, this study also provided evidence that the interaction of single stranded RNA with rGO is stronger than that with GO. The physicochemical diversity and ease of uniform chemical synthesis of graphene also create a favorable platform for loading peptides [32]. It has been shown that immobilization of peptides on GNS is attributable to aromatic or cationic amino acid residues [33-36]. Kawazoe and coworkers [34] reported preferential binding of aromatic amino acids (tryptophan, tyrosine, phenylalanine, and histidine) to graphene and carbon nanotubes via  stacking interactions. Other researchers have reported that GO adsorption strength is higher for arginine, histidine, and lysine relative to tryptophan, tyrosine, and phenylalanine because of their electrostatic interactions with negatively charged GO [35]. Protein adsorption onto GO is associated with intrinsic surface functional groups [33]. Zhang and colleagues demonstrated that GO-based protein-delivery systems exhibit an extremely high capacity to load protein through physisorption compared with other nanocarriers [25]. Efforts to achieve effective protein loading generally benefit from a detailed analysis of the target protein and modification of the protein’s surface to increase its ability to bind GNS. 7

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2.2. Quenching of fluorescent drugs

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In addition to its drug-loading capacity, GO is able to serve as a fluorescence resonance energy transfer-sensing tool for biomedical applications by virtue of its ability to interact noncovalently [37]. Because of the powerful quenching capacity of fluorescent molecules after adsorption, GO has been widely utilized to develop sensors for biomolecules [27]. In the drug-delivery field, the quenching effects of GO and rGO have been effectively used to evaluate the loading capacity and cellular liberation of cargo drugs that possess their own fluorescence.

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Numerous studies have demonstrated doxorubicin loading onto GO by simple mixing, and shown that loading can be evaluated by monitoring quenching of the intrinsic fluorescence of doxorubicin (excitation wavelength, 480 nm; emission wavelength, 580 nm). Doxorubicinloaded GO exhibits a remarkable extinction of the fluorescence emission peak, demonstrating strong adsorption of doxorubicin onto GO [20]. Because chlorin e6 exhibits fluorescence with an excitation wavelength at 400 nm and an emission wavelength at 670 nm [38],  interactions between chlorin e6 and GO can also be confirmed by fluorescence quenching [24]. Drug release from GO can be similarly evaluated in vitro [24] and in vivo [39] by detecting the recovered fluorescence of drugs that have separated from GO.

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Even if the drug loaded onto GO does not have inherent fluorescence, labeling the drug with a fluorescent tag enables evaluation of drug loading, release, and delivery rate using the fluorescence-quenching effect. In this context, the ability of PEGylated GO to act as a nanocarrier for the intracellular delivery of proteins was previously investigated by labeling proteins with fluorescein isothiocyanate, and then assessing adsorption, release profile, and cellular trafficking [25]. Notably, the aromatic structures of fluorescence markers conjugated with proteins potentiate the strong  interactions between proteins and PEGylated GO.

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Quenching of fluorescent compounds has been observed after loading of a cationic dye to gelatin-modified graphene nanosheets [40]. In this study, the model cationic fluorescence probe, R6G, was loaded onto negatively charged gelatin-modified graphene nanosheets. The complex of R6G and gelatin-modified graphene nanosheets showed complete quenching, indicating fluorescence energy transfer from the dye to gelatin-modified graphene nanosheets. Using fluorescence recovery of R6G to evaluate the release from gelatin-modified graphene nanosheets, the authors of this study showed that this process was pH-dependent, exhibiting 28% and 48% release at pH 4.6 and 2.0, respectively. Given the acidic conditions of lysosomes, the enhanced release of drug from graphene nanosheets in acidic conditions would be a useful feature that would allow the liberation and diffusion of drug within cells.

3. Surface modification to enhance biocompatibility Although one of the main applications of GNS is as a drug-delivery vehicle [41], biocompatibility concerns can still be an issue. Several studies have explored the cytotoxicity of GNS. One such study reported that graphene obtained by the chemical vapor-deposition 8

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technique induces caspase-3 (apoptosis-related cysteine peptidase) activation, reactive oxygen species generation, and lactate dehydrogenase release in PC12 pheochromocytomaderived cells [42]. It has also been shown that the cytotoxicity of GO towards human fibroblasts is significantly increased at concentrations above 50 mg/l [43]. In addition, GO has been shown to accumulate in the lungs for long periods of time, inducing dose-dependent pulmonary toxicity after intravenous injection (10 mg/kg) in rats or mice [44]. To improve biocompatibility, researchers often functionalize graphene and its derivatives with stabilizers that prevent graphene aggregation under physiological conditions. Among such stabilizers are synthetic polymers, natural polysaccharides, and proteins (Fig. 2). Examples of materials used for surface modification of GNS are summarized in Table 1. Previous studies have demonstrated that stabilized graphene can be a promising drug-delivery nanocarrier owing to its relatively low toxicity, stability in the circulation, and ability to load anticancer drugs.

Fig. 2. Schematic illustration of surface modifications to enhance biocompatibility. Examples of surface modification with six-arm PEG (A) [9], poloxamer 407 (B) [51], chitosan (C) [55], low-molecular-weight heparin derivative (D) [58], gelatin (E) [40], or bovine serum albumin (F) [61].

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3.1. Modification with synthetic polymers

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3.1.1. Modification with PEG

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Several studies have reported that PEG grafting can reduce the cytotoxicity of GO, resulting in increased biocompatibility and physiological stability. For example, it has been shown that PEG-functionalized GO is not toxic against HCT-116 cells, having no effect on cell viability even at concentrations up to 100 mg/l [9]. Conjugation of six-arm, branched PEG to GO nanosheets via amide bond formation (Fig. 2A) was shown to result in improved biocompatibility. Plain GO was found to precipitate in phosphate buffered saline (PBS), cell culture medium and serum, whereas PEGylated GO showed stable dispersion in all those conditions. Amine-terminated, six-arm PEG has been conjugated to GO and loaded with either bovine serum albumin, ribonuclease A, or cyclic-AMP dependent kinase A. Following treatment with PEGylated GO, more than 80% and 90% of HeLa and MCF-7 cells, respectively, remained viable [25].

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In another study, GO was modified with methoxy PEG by conjugating amino-terminated PEG containing a disulfide bond (methoxyPEG-SS-NH2) to the carboxyl group of GO. The methoxy PEG-tethered GO was found to be stable in 0.1 mM PBS and cell culture medium for at least 24 hours, whereas the size of plain GO began to increase by 2 hours after dispersion [45].

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Moreover, blood biochemistry and hematological analyses and histological examinations have revealed that PEG-functionalized GO (20 mg/kg) intravenously administered into mice does not cause any considerable organ damage or body weight changes within 40 days after dosing. In BALB/C mice, the blood circulation half-life of PEG-functionalized GO was found to be 1.5 hours [46].

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Linear PEG-grafted GO has been used to deliver both anticancer drugs and photosensitizers [47]. In this application, PEG with a molecular weight (MW) of 2,000 Da was used for modification of GO. Compared with plain GO nanosheets, linear PEGconjugated GO was found to be 10-fold less toxic in acute toxicity tests, with 100% of mice surviving a dose of 80 mg/kg PEG-conjugated GO and only 10% of mice surviving following treatment with the same dose of plain GO. Another study assessed the fate of PEGylated GO after intravenous, intraperitoneal, or oral administration in mice [48]. These researchers found that PEGylated GO was weakly adsorbed in the gastrointestinal tract, resulting in rapid excretion through urine and feces and a lack of overt toxicity. Although grafting PEG onto graphene can decrease toxicity, it prevents the enzyme-triggered degradation of graphene [49]. Therefore, favorable in vivo behavior and toxicity of functionalized graphenes are associated with controlled ultra-small sizes and well-designed surface coating materials that promote rapid excretion and/or degradation. Although functionalization of graphene with PEG has been shown to improve biocompatibility, this modification can inhibit the release of loaded anticancer drugs. To solve this problem, Shi and coworkers grafted redox-responsive PEG onto the surface of GO 10

ACCEPTED MANUSCRIPT nanosheets [45] and demonstrated that the resulting redox-responsive PEGylated GO exhibit increased physiological solubility and stability in the circulation.

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3.1.2. Modification with other polymers

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In another study, GO [50] and rGO [51] nanosheets were non-covalently functionalized by grafting the block copolymer, poloxamer 407, onto nanosheets (poloxamer 407/graphene) (Fig. 2B) [50]. Poloxamer 407/rGO exhibited higher stability in cell culture medium. At a concentration of 0.2 g/l, rGO formed agglomerates in cell culture medium, whereas poloxamer 407/rGO yielded a stable suspension in medium [50]. Another study reported the stability of poloxamer 407-coated GO in PBS and cell culture media containing 10% serum. Whereas plain GO was found to form agglomerates in buffer and serum-fortified media, no sign of agglomeration was observed for poloxamer 407-coated GO [51].

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Stabilization of GO nanosheets by covalent functionalization with poloxamer 108, Tween-80, or maltodextrin has also been used for anticancer drug-delivery applications [52]. This latter study found little cytotoxicity of the three polymers against HT29 cells, reporting more than 80% cell viability at a concentration of 400 mg/l.

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Poly(lactide)-PEG copolymers have been used to modify the surfaces of GO [53]. In this study, poly(lactide)-PEG copolymers were adsorbed onto GO through hydrophobic interactions. Poly(lactide)-PEG-coated GO showed colloidal stability in saline solutions at concentrations up to 0.1 M NaCl, whereas GO exhibited an increase in size at concentrations of 0.05 M NaCl and above. Moreover, poly(lactide)-PEG-coated GO was found to be stable upon storage for at least 2 months at 8°C.

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Li and colleagues demonstrated the systematic functionalization of GO with the welldefined poly(N-isopropylacrylamide) using “click” chemistry [54]. Unlike GO, poly(Nisopropylacrylamide)-grafted GO yielded stable dispersions in PBS at a concentration of 200 mg/l. The size of poly(N-isopropylacrylamide)-grafted GO dispersed in water did not significantly change over a temperature range of 25°C to 40°C.

3.2. Modification with natural polysaccharides Among various polysaccharides, chitosan and heparin have been most extensively used to modify the surfaces of GO or rGO. Although hyaluronic acid belongs to the family of natural polysaccharides, it has been used to modify GO to enhance targeting capability, and will be addressed in the targeting ligand section of this review. Chitosan has been covalently grafted to GO via an amide linkage (Fig. 2C) for delivery of insoluble anticancer drugs [55]. Unlike GO, chitosan-grafted GO was found to be stable in PBS and culture media at a concentration of 300 mg/l. Chitosan-grafted GO showed a favorable cytotoxicity profile toward human hepatic (HepG2) cells, which retained more than 80% of their viability at a chitosan-grafted GO concentration of 100 mg/l. Chitosan-coated magnetic rGO has been shown to exhibit decreased cytotoxicity compared with plain rGO 11

ACCEPTED MANUSCRIPT [56]. PC3 cell viability following treatment with rGO and chitosan-coated magnetic rGO at the same concentration (160 g/l) was 40% and almost 100%, respectively.

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Chitosan functionalization has been used to enhance the biocompatibility of GO [57]. Chitosan-functionalization of GO was shown to increase the viability of exposed cells compared with those treated with plain GO. However, the degree of cytotoxicity of chitosanfunctionalized GO was found to depend on cell type. Chitosan-functionalized GO did not reduce the viability of CEM cancer cells at concentrations up to 400 mg/l, whereas plain GO reduced the viability of CEM cells to less than 30% at the same concentration—a more than 3-fold improvement in viability with chitosan functionalization. In MCF-7 cancer cells, the viability-enhancing effect of chitosan-functionalized GO was less pronounced, with chitosanfunctionalized GO decreasing cell viability to 50% compared with 27% for plain GO.

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Low-molecular-weight heparin derivatives have been exploited as a coating material for rGO (Fig. 2D) to provide dispersion stability and consequent tumor accumulation of cargo drug [58]. LHT7, a taurocholate derivative of low-molecular-weight heparin and a potent angiogenesis inhibitor, interacts with rGO and has been used to non-covalently coat it. LHT7coated rGO was found to retain its stable dispersion in PBS for up to 3 days at a concentration of 50 mg/l, whereas plain rGO formed aggregates under the same conditions. In vivo molecular imaging revealed much higher tumor accumulation of LHT7-coated rGO, showing that the surface coating of LHT7 conferred stability in the circulation. In addition, delivering doxorubicin using LHT7-coated rGO provided a synergistic anticancer effect in KB tumor-bearing mice.

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3.3. Modification with proteins

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Zhu and colleagues reported an environmentally friendly and facile method for preparing gelatin-functionalized graphene nanosheets from GO (Fig. 2E). They showed that gelatin acted not only as a reducing agent, but also as a functionalizing reagent [40]. Gelatinfunctionalized graphene nanosheets were stable in PBS (0.01 M), cell culture media, and serum at a concentration of 200 mg/l for 24 hours, whereas GO showed agglomeration under the same conditions. MCF-7 cells treated with gelatin-functionalized graphene nanosheets at a concentration of 200 mg/l retained more than 90% of their viability. Wang and coworkers reported gelatin-functionalized graphene nanosheets for methotrexate delivery [59]. Consistent with the previous study of Zhu and colleagues [40], they observed that gelatin-functionalized graphene nanosheets were stable in PBS (0.01 M), cell culture media, and serum at a concentration of 200 mg/l for 24 hours, whereas GO agglomerated under the same conditions. Gelatin-functionalized graphene nanosheets exerted almost no cytotoxicity against A549 cells at a concentration of 300 mg/l. Mahanta and Paul synthesized GO nanosheets using the controlled pyrolysis of citric acid [60]. The resulting GO nanosheets were further functionalized by cross-linking with the small protein, bovine α-lactalbumin, to increase biocompatibility. Bovine α-lactalbumin-conjugated GO nanosheets showed negligible hemolysis of red blood cells at a concentration of 333 mg/l, whereas GO nanosheets caused more than 20% hemolysis at a more than 10-fold lower 12

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concentration (30 mg/l). Moreover, bovine α-lactalbumin-functionalized GO nanosheets exhibited 20-fold lower cytotoxicity than GO toward MCF-7 and MDAMB231 cells. In these cells, the concentration of GO required to cause more than 80% cell death was 12 mg/l, but 240 mg/l of bovine α-lactalbumin-functionalized GO was required to cause the same decrease in viability.

3.4. Modification with phospholipids

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Another study reported that bovine serum albumin facilitated the reduction of GO, resulting in a non-aggregating, bovine serum albumin-functionalized rGO (Fig. 2F) [61]. This study revealed that a toxic reductant (e.g., hydrazine) is not needed to produce rGO nanosheets. In MCF-7 cells, 80% viability was retained after treatment with bovine serum albumin-functionalized rGO at a concentration of 40 mg/l. In mice, no histological changes in organs were observed 30 days after injection of a dose of 20 mg/kg.

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Only a few studies have reported modification of GO with phospholipids. Wang and colleagues developed GO functionalized with a phospholipid monolayer [62], and used neutral, anionic, and cationic liposomes to investigate the interactions between GO nanosheets and liposomes. They found that functionalization with negatively charged liposomes increased stability in aqueous solution compared with functionalization with neutral or cationic liposomes. The diminished stability of cationic liposome-coated GO was attributed to neutralization of the cationic charge by the negative charge on GO nanosheets. GO modified with anionic liposomes composed of 1,2-dimyristoyl-sn-glycero-3-phospho-(1rac-glycerol) was found to be stable at 4°C for 2 months, owing to electrostatic repulsion of the negatively charged liposome-GO complex. Moreover, anionic liposome-coated GO exhibited an increased doxorubicin loading ratio owing to synergistic loading to both GO and the lipid membrane.

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As mentioned above, several studies have reported that graphene and its derivatives functionalized with different biocompatible materials have enhanced biocompatibility and physiological stability, with no severe toxicity in vitro or in vivo. However, a more detailed characterization, including mechanistic toxicity studies, will be important for evaluating their biocompatibility and toxicology profile. Although a major thrust of the research on surface modification of GO has been on enhancing biocompatibility, another goal of surface modification is to enable the loading of active substances onto GO. This type of surface modification can be applied to nucleic acid drugs, which cannot be loaded onto GO through hydrophobic or  interactions. For example, plasmid DNA is unable to bind GNS and requires surface modifications of GO with cationic polymers, such as cationic polyethyleneimine.

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ACCEPTED MANUSCRIPT Table 1. Materials used for surface modification of GNS to enhance biocompatibility.

Poloxamer 108

SN38

10

[9]

Bovine serum albumin

350

[25]

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GO, 5–50

Stable in serum, 100% cell viability (100 mg/l)

Conjugation

Sonic/ Cent

GO, <200

>80% cell viability

Conjugation

Cent

GO, <200

Doxorubicin

N/A

[45]

Conjugation

Cent

Graphene, 10–50

No precipitation in 0.1 mM PBS (24 h) No organ damage in mice (20 mg/kg)

N/A

N/A

[46]

Conjugation

Cent/ Filt

GO, <200

10-fold higher survival than GO (80 mg/kg)

Doxorubicin chlorin e6

12 16

[47]

Hydrophobi c interactions Hydrophobi c interactions Conjugation

Cent

Graphene, 80

Stable in media (0.2 g/l)

Doxorubicin

289

[50]

GO, 41

Stable in media with 10% serum

Methylene blue

22

[51]

Filt

GO, 20–120

>80% cell viability (400 mg/l)

Ellagic acid

100

[52]

Conjugation

Sonic/ Filt

GO, 100

75% cell viability (0.75 mg/l)

Doxorubicin

143

[117]

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Cent

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Poly(amido amine)

Ref.

Drug

Conjugation

Sonic/ Filt

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Poloxamer 407

Drug loading capacity (wt%)

Effect

D

PEG

GNS type, lateral size (nm)

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Materials

Size control method

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Modification

Poly(lactide) -PEG

Hydrophobic interactions

N/A

GO, 455– 534

Stable in 0.1 M NaCl

Paclitaxel

11

[53]

Poly(Nisopropyl acrylamide) Chitosan

Conjugation

Filt

GO, 190

Stable in PBS (200 mg/l; 25°C, 37°C)

Camptotheci n

16

[54]

Conjugation

Filt

GO, 170

Stable in media (300 mg/l)

Camptotheci n

20

[55]

Conjugation

N/A

rGO, 94

2.5-fold higher cell viability than rGO

Doxorubicin

12

[56]

Conjugation

Cent

>3-fold higher cell viability than GO (400 mg/l)

Fluorouracil

5

[57]

GO, 350

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Sonic/ Filt

rGO, 100–200

Stable in PBS (50 mg/l, 3 days)

Doxorubicin

33

[58]

Gelatin

Reduction

Cent

rGO, 80–3000

Stable in PBS, media, serum (200 mg/l)

Doxorubicin

13

[40]

Reduction

Sonic / Cent

rGO, N/A

Stable in PBS, media, serum (200 mg/l)

Methotrexate

28

[59]

Conjugation

Pyrolysi s of citric acid

GO, 300

Negligible hemolysis (333.3 mg/l) 20-fold lower cytotoxicity than GO

N/A

N/A

[60]

No histologic changes in organs of mice (20 mg/kg)

N/A

N/A

[61]

Stable in water (4°C, 2 months)

Doxorubicin

70

[62]

Hydrophobic interactions

Cent

Graphene, N/A

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1,2dimyristoylsn-glycero3-phospho(1-racglycerol) sodium salt

rGO, 70

NU

Hydrophobic Sonic/ interactions Cent

Bovine serum albumin

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Lowmolecularweight heparin

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Filt, filtration; Sonic, sonication; Cent, centrifugation; N/A, information not available. 4. GNS for delivery of chemotherapeutics

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GNS have been studied for delivery of various chemotherapeutics, including anticancer drugs and photosensitizers. Among the tested anticancer drugs are doxorubicin, camptothecin, paclitaxel, cisplatin analogues, 1,3-bis(2-chloroethyl)-1-nitrosourea, fluorouracil, methotrexate, lucanthone, -lapachone, and ellagic acid (Fig. 3; Table 2). Photosensitizers codelivered with anticancer chemotherapeutics include hypocrellin A and chlorin e6. These chemotherapeutics are loaded onto GNS via physical adsorption or chemical conjugation.

4.1. Anthracycline antitumor antibiotics The anthracycline antitumor antibiotic, doxorubicin (Fig. 3), has been the most extensively studied for loading onto GO. Because of the anthracycline structure, which contains aromatic rings and amino groups, doxorubicin can be loaded onto GO via dual driving forces— stacking interactions and electrostatic interactions—providing for high loading capacity [4, 15, 19, 20, 63]. Doxorubicin-loaded GO has been studied for doxorubicin delivery to various human tumor cells, including breast cancer cells [3], nasopharyngeal carcinoma cells [64], and cervical cancer cells [45]. 15

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Fig. 3. Examples of anticancer chemotherapeutics delivered using GNS.

It has been shown that doxorubicin delivered on GO is taken up to a greater extent and shows higher efficacy against doxorubicin-resistant breast cancer MCF-7 cells than free doxorubicin [3]. Both endocytosis-dependent and -independent pathways have been proposed to contribute to cellular entry of doxorubicin on GO. Physically adsorbed doxorubicin on GO surfaces exhibits pH-sensitive release under acidic conditions. GO, acting as a physical cross-linker, has been used to form an injectable doxorubicin gel [64]. The addition of a small amount of GO to a doxorubicin solution was found to form a gel matrix that possessed considerable viscosity and mechanical strength suitable for injection. It was found that doxorubicin was released from the gel matrix in a sustained manner, and exerted anticancer effects against CNE1 human nasopharyngeal carcinoma cells. Various strategies have been reported to facilitate intracellular doxorubicin release from GO, including glutathione treatment, and redox- or pH-responsive systems. Simple treatment with glutathione has been shown to facilitate intracellular doxorubicin release from GO [39,45], promoting more effective release of doxorubicin compared with doxorubicin-loaded 16

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PEGylated GO alone [39]. Although endogenous levels of glutathione can induce the release of doxorubicin from PEGylated GO, exogenous addition of glutathione has been shown to enhance the release rates of doxorubicin in vitro and in vivo. In these studies, the release of doxorubicin from GO was detected by monitoring recovery of its fluorescence. For redoxresponsive detachment of PEG, Li and colleagues [45] conjugated amino-terminated PEG bearing a disulfide bond to carboxylated GO. Treatment with external glutathione accelerated the liberation of PEG from GO, resulting in rapid release and enhanced anticancer effects of doxorubicin against HeLa human cervical cancer cells.

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To exploit the ability of lower pH in endosomes or lysosomes to enhance the intracellular release of doxorubicin, researchers have loaded doxorubicin onto GO through electrostatic interactions. This was accomplished by conjugating doxorubicin to the anionic carboxylatefunctional polyelectrolyte, citraconic anhydride-functionalized poly(allylamine), and loading it onto positively charged polyethyleneimine-modified GO [65]. Under the acidic conditions of lysosomes or endosomes, it has been found that doxorubicin is released from polyethyleneimine-modified GO owing to repulsion between cationic polyethyleneimine and the charge-reversed cationic polyelectrolyte at the reduced pH.

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Gold nanoclusters and doxorubicin have been loaded onto rGO for multimodal cancer cell imaging and treatment [66]. Gold nanoclusters exhibit near-infrared photoluminescence, which allows cell imaging, and can be loaded onto rGO through non-covalent interactions. Doxorubicin delivered using gold nanocluster-decorated rGO showed a diffuse cytoplasmic distribution pattern in HepG2 cells. In contrast, free doxorubicin was localized to the cell membrane. Doxorubicin delivered using gold nanocluster-modified rGO showed a higher delivery rate compared to free doxorubicin.

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Although most studies have reported the use of GO for delivery of doxorubicin, a few studies have investigated graphene nanosheets [62], including phospholipid-functionalized graphene nanosheets, or rGO for delivery of doxorubicin. Formation of an rGO and gold nanocomposite hybrid has been reported for delivery of doxorubicin [66]. rGO is known to interact with metal nanoparticles through - stacking, cation-, or van der Waals interactions. The resulting rGO/gold nanocomposite hybrid was further loaded with doxorubicin and shown to exert an anticancer effect in HepG2 hepatoma cells [66].

4.2. Anticancer quinoline alkaloids Among anticancer alkaloids, the plant-derived quinoline alkaloid camptothecin, which has an aromatic ring structure (Fig. 3), has been loaded onto GO decorated with a water-soluble polymer [54, 67, 68] or graphene nanosheets [69] through hydrophobic interactions and  stacking interactions. Using poly(N-isopropylacrylamide)-conjugated GO, Li and coworkers achieved a camptothecin loading capacity of 15.6% [54]. Since camptothecin attaches strongly to graphene surfaces, the rate of camptothecin release from poly(Nisopropylacrylamide)-modified GO was found to be lower than that from other nanoparticles, making this an effective system for the stable and sustained release of camptothecin. Camptothecin-loaded, poly(N-isopropylacrylamide)-conjugated GO was shown to exert increased growth-inhibitory effects against A-5RT3 metastatic skin tumor cells compared 17

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with free camptothecin [54]. Park and colleagues [67] further showed that poly(N-vinyl caprolactam)-grafted GO loaded with camptothecin was internalized into KB cells by endocytosis and exhibited robust cancer cell-killing effects through the topoisomeraseinhibiting mechanism of camptothecin.

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Poly(vinyl alcohol)-functionalized GO has been compared with multiwalled carbon nanotubes as a nanocarrier for camptothecin [68]. This study showed that the amount of camptothecin loaded onto PVA-functionalized GO and poly(vinyl alcohol)-functionalized multiwalled carbon nanotubes was similar: 0.12 and 0.1 g per gram of nanocarrier, respectively. Cytotoxicity tests performed using MDA-MB-231 cells showed that 80% of cells treated with poly(vinyl alcohol)-functionalized GO (1 g/l) remained viable compared with controls.

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To achieve synergistic anticancer effects, researchers have co-loaded camptothecin and doxorubicin onto folic acid-modified GO [41]. The adsorption of doxorubicin onto GO was found to be higher relative to that of camptothecin. The release rate of doxorubicin after 48 hours in acidic conditions (pH 5.0) was 2-fold higher than that of camptothecin. The higher release rate of doxorubicin from GO was interpreted as reflecting the greater hydrophilicity of doxorubicin conferred by its amine group at the lower pH. Folic acid-modified GO coloaded with camptothecin and doxorubicin exhibited synergistic anticancer activity in MCF-7 cells compared with GO loaded with either drug alone.

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Starch-functionalized graphene nanosheets have been used as delivery vehicles for the camptothecin analogue hydroxycamptothecin [69]. In this application, starch-functionalized graphene nanosheets were synthesized by reduction of GO using soluble starch as a reductant. The large surface area of starch-functionalized graphene nanosheets allows hydroxycamptothecin to be physisorbed through – interactions, achieving a loading capacity of 5.7 g per 100 g of starch-functionalized graphene nanosheets. Cytotoxicity tests showed that treatment of SW-620 cells with up to 200 mg/l starch-functionalized graphene nanosheets did not significantly reduce cell viability.

4.3. Taxanes

Taxane family drugs, with a diterpene structure, are usually difficult to formulate owing to their poor water solubility. The taxane family includes paclitaxel (Fig. 3) and docetaxel, which exert anticancer effects by inhibiting microtubules. GO modified with the waterdispersible poly(lactide)-PEG copolymer, a US Food and Drug Administration-approved copolymer, was investigated as a means for enhancing the stability of GO in aqueous dispersions and improving paclitaxel delivery [53]. This poly(lactide)-PEG copolymerfunctionalized GO exhibited a paclitaxel loading capacity of 9–11%, and showed greater colloidal stability on storage at 8°C over 2 months in the presence of NaCl than did the corresponding plain GO. Paclitaxel delivered using poly(lactide)-PEG copolymerfunctionalized GO exhibited anticancer activity against A549 human lung alveolar basal carcinoma epithelial cells.

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Platinum complexes include cisplatin, carboplatin, oxaliplatin, and their analogues. A PEGylated GO functionalized with an immobilized apoptosis indicator has been designed for concurrent delivery of the cisplatin analogue Cis-Pt(NH3)2Cl2 and noninvasive, real-time monitoring of cell apoptosis [70]. In this application, Cis-Pt(NH3)2Cl2 was covalently linked to PEGylated GO for glutathione-responsive reduction and release within the intracellular environment. Immobilization of a fluorescent probe cleavable by caspase-3 was introduced for apoptosis monitoring. Cellular damage caused by the cisplatin analogue induced apoptosis and caspase-3 up-regulation, resulting in liberation of the GO-bound, quenched fluorescent probe and generation of a fluorescence signal that could be imaged.

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GO functionalized with polyacrylic acid has been investigated for delivery of carmustine (1,3-bis(2-chloroethyl)-1-nitrosourea), an anticancer alkylating agent [71]. The drug was conjugated to the carboxyl groups of polyacrylic acid through amide bonding, and this covalent tethering of carmustine to polyacrylic acid-functionalized GO was found to enhance the thermal stability of the drug. Moreover, it also increased the half-life of the drug from ~19 hours (free drug) to 43 hours. The drug delivered using polyacrylic acid-functionalized GO exhibited an IC50 that was 4.3-fold lower than that of free drug in GL261 cells. The enhanced anticancer effect of 1,3-bis (2-chloroethyl)-1-nitrosourea following delivery using polyacrylic acid-functionalized GO was attributed to enhanced stability and more efficient endocytosis.

4.6. Pyrimidine analogues

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GO decorated with magnetite (Fe3O4) nanoparticles has been studied as a delivery system for fluorouracil [72] (Fig. 3), a pyrimidine analogue that exerts an anticancer effect by inhibiting thymidylate kinase. The pyrimidine structure of fluorouracil allows loading onto Fe3O4/GO through stacking interactions and hydrogen bond formation between the strong electronegative atoms of fluorouracil and H-atoms of carboxyl groups in GO. Acidic pH was shown to promote greater release of fluorouracil from Fe3O4/GO than neutral pH. This pH-responsive release is thought to reflect pH-dependent changes in hydrogen bonding between fluorouracil and Fe3O4/ GO. GO functionalized with chitosan has also been used for delivery of fluorouracil [57], which was loaded through  stacking. The release of fluorouracil from chitosanfunctionalized GO was found to be sustained over 3 days and was promoted at lower pH. Fluorouracil delivered using chitosan-functionalized GO was effective, causing greater than 80% toxicity toward CEM and MCF-7 cancer cells at a concentration of 400 mg nanosheets/l.

4.7. Polyphenolic compounds 19

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Ellagic acid (Fig. 3) is a natural polyphenolic compound that acts as an anticancer agent by inducing reactive oxygen species-mediated apoptosis. Ellagic acid has been loaded onto GO modified with various biocompatible moieties, such as poloxamer 108, Tween-80, and maltodextrin [52]. Ellagic acid, whose polyphenolic aromatic structure allows non-covalent interactions with GO, showed a loading capacity of 1, 1.22, and 1.14 g per gram of GO functionalized with poloxamer 108, Tween-80, and maltodextrin, respectively. All three forms of ellagic acid-loaded, modified GO showed higher cancer cell-killing effects in MCF7 breast cancer and HT-29 colon carcinoma cells than did free ellagic acid.

4.8. Quinone compounds

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-Lapachone (Fig. 3), a poorly water-soluble o-naphthoquinone that acts as an antitumor agent by inhibiting DNA topoisomerase I, has been loaded onto GO or rGO decorated with Fe3O4 nanoparticles [73]. Both Fe3O4-decorated GO and rGO were loaded with -lapachone through non-covalent  interactions. Increasing the time for hydrazine reduction of rGO increased the -lapachone loading capacity. The loading capacity of -lapachone on Fe3O4/rGO was 46%, much higher than that on Fe3O4/GO, which showed little ability to load -lapachone. This dramatic difference in loading capacity between Fe3O4/rGO and Fe3O4/GO was attributed to the planar structure of rGO. Since most oxygen-containing groups of GO are eliminated with an increase in sp2 clusters, rGO has a much more planar structure than GO, allowing more favorable ππ stacking and hydrophobic interactions with the neutral, aromatic -lapachone. -Lapachone delivered to MCF-7 breast cancer cells using Fe3O4/rGO caused cell shrinkage and exerted greater anticancer activity than the free drug.

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Li and colleagues [74] described the subcellular-targeted delivery of -lapachone following functionalization of rGO with an optical fiber nanoprobe. The surfaces of the optical fiber nanoprobe were modified to contain positively charged aminopropyltriethoxysilane, and the nanoprobe assembled with the negatively charged rGO through electrostatic interactions. Because the optical nanoprobes have no capacity to load lapachone, rGO was assembled onto the nanoprobes to support -lapachone loading and provide access to the subcellular compartments that support the biochemical responses of βlapachone.

4.9. Other chemotherapeutics 4.9.1. Methotrexate Gelatin-functionalized graphene nanosheets have been used for the controlled release of methotrexate [59]. This study reported that methotrexate (Fig. 3) was adsorbed onto gelatinfunctionalized graphene nanosheets through non-covalent interactions and showed a saturated loading capacity of 0.28 g methotrexate per gram of gelatin-functionalized graphene nanosheets. Acidic conditions accelerated the release of methotrexate, reflecting the higher solubility of methotrexate under acidic conditions, dissociation of hydrogen bonds, and decomposition of gelatin under acidic conditions. Gelatin-functionalized graphene nanosheets 20

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4.9.2. Lucanthone

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Lucanthone (Fig. 3), a prodrug of the apurinic endonuclease-1 inhibitor, hycanthone, has been loaded onto PEGylated GO nanoribbons and delivered to glioblastoma multiforme cell lines [75]. Cellular uptake of lucanthone loaded onto PEGylated GO nanoribbons was greater than 60% in apurnic endonuclease-1-overexpressing U251 cells, but less than 40% in MCF-1 cancer cells and rat glial progenitor cells. In line with these uptake patterns, lucanthone delivered using PEGylated GO nanoribbons exhibited higher anticancer efficacy than free lucanthone in APEX1-overexpressing U251 cells, but lower efficacy than the free form in MCF-7 cells.

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4.10. Co-delivery of anticancer drugs with photosensitizers

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Photodynamic therapy uses a photoresponsive agent—termed a photosensitizer—and light. Exposure to a specific wavelength of light generates reactive oxygen species, which kill nearby cells. However, most photosensitizers tend to readily form aggregates in aqueous solutions owing to their hydrophobic properties [76]. Thus, various types of drug-delivery carriers, including liposomes [77], solid lipid nanoparticles [78], polymeric nanoparticles [79], silica nanoparticles [80] and gold nanoparticles [81], have been used to deliver photosensitizers. Recently, GNS have been studied for the delivery of photosensitizers alone [24,82-84] or co-delivered with anticancer drugs [47,85].

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GO has been used as a carrier for photosensitizers such as hypocrellin A [82] and hypocrellin B [83], alone or combination with chemotherapeutics. In this application, hydrophobic hypocrellin A was loaded onto GO through - stacking interactions with the aromatic ring moiety and hydrogen bonding with hydroxyl and carbonyl groups of hypocrellin A [82] . Loading of hypocrellin A onto GO was found to enhance the stability of hypocrellin A in aqueous solution. Free hypocrellin A lost much of its singlet oxygengenerating ability, which was decreased to one third that of fresh hypocrellin A after 3 days in aqueous solution. However, hypocrellin A on GO retained its stability for 3 days. Hypocrellin A-loaded GO exerted photodynamic effects in HeLa cells, reducing cell viability to less than 50% at a hypocrellin A concentration of 2.2 mg/l after 2.5 minutes of light irradiation. The cell-killing effect of hypocrellin A-loaded GO following light irradiation was attributed to intracellular reactive oxygen species generated by hypocrellin A.

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Chemotherapeutics Doxorubicin

Doxorubicin-resistant MCF-7, CNE1 cells

Ref. [3] [64]

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Efficacy test

Doxorubicin-resistant MCF-7cells

[39]

HeLa cells U87MG, MCF-7 cells

[45] [65]

HepG2 cells

[66]

A-5RT3 cells

[54]

Camptothecin

KB cells

[67]

Camptothecin

MDA-MB-231 cells

[68]

Hydroxycamptothecin Doxorubicin, Camptothecin Paclitaxel Cisplatin analogue Fluorouracil Fluorouracil Methotrexate

SW-620 cells MCF-7 cells

[69] [41]

A549 cells 4T1-bearing mice HepG2 cells MCF-7 cells A549 cells

[53] [70] [72] [57] [59]

Polyacrylic acid-functionalized GO PEGylated graphene Fe3O4/rGO, Fe3O4/GO

1, 3-bis (2-chloroethyl)1-nitrosourea Lucanthone

GL261 cells

[71]

U251 cells MCF-7 cells

[75] [72]

rGO (modified nanoprobe)

-Lapachone Ellagic acid

MCF-7 cells

[74]

MCF-7, HT-29 cells

[52]

Camptothecin

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Doxorubicin

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Poly(vinyl alcohol)functionalized GO

Doxorubicin Doxorubicin

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Citraconic anhydridefunctionalized poly(allylamine)/ polyethyleneimine-GO Gold nanocluster-decorated rGO Poly (N-isopropylacrylamide)GO Poly (N-vinyl caprolactam)grafted GO

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Poly(lactide)–PEG PEGylated GO Fe3O4/graphene nanosheets Chitosan-functionalized GO Gelatin-functionalized graphene nanosheets

Poloxamer 108-GO, Tween80GO, Maltodextrin-GO

-Lapachone

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Notably, the same group studied the major drug loading mechanisms by testing hypocrellin B, which has a subtly different structure than hypocrellin A [83]. Compared with hypocrellin A, hypocrellin B has one fewer hydroxyl groups and one more double bond. Hypocrellin B, for which - interactions are more dominant than is the case for hypocrellin A, showed approximately a 2-fold higher drug-loading capacity (2 mg/mg GO) than hypocrellin A (1 mg/mg GO) [83]. Hypocrellin B delivered on GO was found to exert photodynamic anticancer effects, reducing cell viability to less than 20% in several tumor cell lines, including HeLa, SMMC-7721, SGC-7901 and A549 cells, upon laser irradiation at 470 nm for 45 seconds [83].

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GO has also been used to deliver both photosensitizers and anticancer agents with the goal of achieving a synergistic effect through combined photodynamic therapy and chemotherapy [85]. In this application, hypocrellin A, used as a photosensitizer, and 7-ethyl10-hydroxycamptothecin (SN-38), used as an anticancer drug, were both loaded onto GO. Hypocrellin A and SN-38 co-delivered using GO exerted synergistic anticancer effects compared with single treatment with hypocrellin A/GO and SN-38/GO. Synergy between codelivered hypocrellin A and SN-38 was determined based on the synergistic index (the expected growth-inhibition rate divided by the observed growth-inhibition rate), where an index value higher than 1.0 is considered to indicate synergistic effects. The combination of hypocrellin A and SN-38 at the same concentration resulted in synergistic indices greater than 1.0 over the concentration range 0.5–6.0 M in A549 cells after irradiation with 470 nm light-emitting diodes.

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Chlorin e6 alone has been loaded onto folic acid-conjugated GO to increase the effects of anticancer photodynamic therapy [24]. In MGC 803 human gastric cancer cells, treatment with chlorin e6 on folic acid-conjugated GO decreased the viability of cells to ~10% after laser irradiation, whereas cell viability was maintained at ~80% in the absence of laser irradiation.

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Chlorin e6 has been co-delivered with doxorubicin using PEGylated GO nanosheets [47]. Chlorin e6 and doxorubicin co-delivered by PEGylated GO improved in vitro anticancer activity compared with chlorin e6/PEGylated GO or doxorubicin/PEGylated GO alone following illumination at 660 nm. In this study, synergistic effects of co-delivery were evaluated by calculating a combination index [86], where a combination index value less than 1.0 is considered to indicate synergy. The synergistic effects of co-delivered chlorin e6 and doxorubicin were found to depend on the specific combination used. Whereas doxorubicin and chlorin e6 combined at a molar ratio of 1:8 showed no synergy, molar ratios of 1:2 and 1:4 exhibited combination index values less than 0.4, indicating strong synergy. Treatment of SCC7 cells with doxorubicin (2.5 M) and chlorin e6 (5.0 M) reduced cell viability to less than 20% after irradiation. Moreover, following intravenous administration, chlorin e6-loaded PEGylated GO showed greater accumulation in tumor tissues of SCC7 tumor-bearing mice than did free chlorin e6, and suppressed tumor growth more effectively than chlorin e6 on GO or doxorubicin on GO. Although non-covalent loading of drugs onto graphene is common in the drug-delivery field, GO chemically conjugated with chlorin e6 for photodynamic therapy has also been reported [84]. Here, chlorin e6 was covalently conjugated to GO via a disulfide (SS) bond, which is 23

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cleaved by redox agents, such as glutathione or dithiothreitol. Incubation with dithiothreitol released chlorin e6, detected as the recovery of chlorin e6 fluorescence. Moreover, chlorin e6 cleaved from the GO conjugates showed strong ability to generate singlet oxygen, whereas GO-bound chlorin e6 did not. Thus, the expectation is that Ce6 would selectively exert a photodynamic effect in cancer cells that had taken up chlorin e6-SS-GO, since the intracellular concentration of glutathione is known to be much higher than that of the extracellular environment.

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Co-delivery of anticancer drugs and photosensitizers using GO has been suggested for synergistic enhancement of anticancer effects. To date, however, most such studies have been done in cultured cells. Thus, strategies for applying graphene for co-delivery of photosensitizers and anticancer drugs in an in vivo model will require further investigation. One limitation of co-delivery in in vivo applications is the relatively short penetration of light used for photodynamic therapy. With progress in optical device technology, the co-delivery approach is expected to see application in various in vivo setting in the near future.

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Most drugs have been loaded onto GO or rGO by physisorption, rather than through chemical conjugation. Although this commonly applied approach for loading drug on GNS has several advantages, such as the simplicity of the process and the potential for high loading capacity, there still issues that must be addressed for effective drug loading. As shown in Table.1, the loading capacity of a given drug onto GNS differs according to the type of GNS, charge of surface modifiers, and structure of drug molecules. First, regarding GNS type, rGO may facilitate higher loading capacity of hydrophobic drugs than GO because of its larger uncharged planar area for hydrophobic and π-π interactions [26]. Second, the functionalization of GO with negatively charged polysaccharide [58, 87] and negative phospholipids [62] could allow electrostatic interactions with positively charged drugs. It is also possible that cargo drug and surface modifier might compete for the same surface area if both molecules adsorb onto GNS via π-π stacking interactions. Third, the structure of drug molecules is crucial. If the compound has electronegative atoms that can interact with GO through hydrogen bonding, drug loading onto GO may be increased. The ionizable carboxyl groups at the edge region of GO can provide pH-dependent drug loading capacity. Therefore, for optimal drug loading, the selection of GNS type and coating material should be considered based on the physicochemical properties of cargo drugs.

5. GNS for delivery of biological drugs 5.1. Delivery of plasmid DNA Recently, GO has been studied for delivery of large nucleic acid-based biomolecules, including plasmid DNA and siRNA (Fig. 4). To enable plasmid DNA delivery, Liu and colleagues complexed negatively charged GO with branched cationic polyethylenimine of different molecular weights (1.2 and 10 kDa) (Fig. 4A) through electrostatic interactions [88]. Branched polyethylenimine-modified GO was found to be stable in physiological solutions, such as saline and cell culture medium, without aggregating. In addition, surface coating of GO with polyethylenimine increased the zeta potential of GO from -50 mV to approximately +30 to +50 mV. The positive charge of polyethylenimine-modified GO enables complexation 24

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with negatively charged plasmid DNA through electrostatic interactions. Plasmid DNA encoding enhanced green fluorescent protein- was complexed with free polyethyleneimine or polyethyleneimine-modified GO. In the case of 1.2-kDa polyethylenimine, plasmid DNA delivered using polyethylenimine-modified GO, but not free 1.2-kDa polyethyleneimine alone, resulted in expression of the encoded enhanced green fluorescent protein. However, in the case of 10-kDa polyethylenimine, both naked polyethylenimine and polyethyleneiminemodified GO effectively transfected plasmid DNA. However, unlike polyethylenimine on GO, naked 10-kDa polyethylenimine exhibited significant cytotoxicity.

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Branched polyethyleneimine has been chemically conjugated to GO for delivery of plasmid DNA [89]. In this application, the amino group of branched, 25-kDa polyethyleneimine was conjugated to the carboxyl groups of GO using N-ethyl-N’-(3dimethylaminopropyl)carbodiimide hydrochloride. Branched polyethylenimine-conjugated GO conjugates showed a zeta potential of +49 mV and exhibited gel retardation owing to the associated plasmid DNA encoding green fluorescent protein. Although the branched polyethyleneimine conjugated to GO transfected plasmid DNA to an extent similar to that of free polyethyleneimine, chemical conjugation of polyethylenimine to GO reduced the cytotoxicity of free polyethylenimine. Upon treatment of HeLa cells with polyethyleniminemodified GO at a concentration of 10 mg/l for 1 hour, cell viability remained at ~90%. However, the viability of cells exposed to branched, 25-kDa polyethylenimine under the same conditions was less than 50%. One day post transfection, the majority of Cy3-labeled plasmid DNA was located in the nucleus. Other researchers have modified GO with branched, 25-kDa polyethyleneimine for delivery of plasmid DNA encoding green fluorescent protein [90]. They observed that more than 45% of cells were transfected with the plasmid DNA delivered using the branched polyethylenimine-conjugated GO at a molar ratio of amine groups in polyethylenimine-conjugated GO to plasmid DNA phosphate groups (N/P ratio) of 7.6:1. Moreover, this study observed the nuclear entry of plasmid DNA in complex with the nanocarrier in H293T cells.

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Plasmid DNA encoding luciferase has also been delivered using polyethylenimineconjugated GO [91]. In this study, gel retardation decreased as the N/P ratio increased, suggesting that the amount of polyethylenimine on GO is a crucial factor in determining the capacity for plasmid DNA complexation. Plasmid DNA-complexed polyethylenimineconjugated GO showed efficient transfection of HeLa and PC-3 cells. Similar to the findings of Zhang and colleagues [89], this study observed that the transfection efficiency of plasmid DNA with polyethylenimine-conjugated GO was comparable to that of the free form of polyethylenimine, but cytotoxicity was reduced. Branched polyethylenimine has been linked to rGO using hydrophilic PEG as a spacer [8]. Low-molecular-weight (1.8 kDa), branched polyethylenimine and PEG-modified rGO were shown to form complexes with plasmid DNA and support expression of encoded genes in PC-3 and NIH3T3 cells. Notably, this study combined photothermal effects to facilitate transfection, showing that irradiation of cells with near-infrared irradiation (808 nm) enhanced transfection efficiency. This enhanced efficiency was attributed to heat-accelerated cytosolic escape of plasmid DNA from endosomes.

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Fig. 4. Schematic illustration of biological drug-loaded GO. Plasmid DNA was loaded onto GO modified with cationic polyethylenimine (A) [88] or chitosan (B) [96]. siRNA was loaded onto GO modified with cationic polyethylenimine (C) [30] or 1-pyrenemethylamine hydrochloride (D) [105].

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Linear polyethylenimine-modified GO has also been used for plasmid DNA delivery [92]. In this application, amine groups of linear, 25-kDa polyethylenimine were chemically conjugated to reactive epoxides of GO. Linear polyethylenimine-conjugated GO protected plasmid DNA from nucleases. Specifically, plasmid DNA complexed to linear polyethylenimine-modified GO remained intact after a 2-hour incubation with DNase I, whereas naked plasmid DNA was degraded within 15 minutes.

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Linear, 60-kDa polyethyleneimine has been used for transfection of plasmid DNA encoding green fluorescent protein in cells and a zebrafish model [93]. At an N/P ratio of 9.2:1, linear polyethylenimine-conjugated GO showed transfection efficiencies greater than 90% in H293T cells and human osteosarcoma U2OS cells and was minimally toxic, as evidenced by cell viabilities higher than 80% at 3 days post treatment. This study also reported that the transfection efficiency of linear polyethylenimine alone was lower and cytotoxicity was higher compared with linear polyethylenimine-modified GO at similar N/P ratios. Moreover, microinjection of zebrafish embryos with plasmid DNA complexed to linear polyethylenimine-conjugated GO was found to result in the expression green fluorescent protein with low toxicity.

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Several studies have investigated the cytotoxicity of branched or linear polyethylenimine in free form and conjugated to GO [30,88,89,91]. The cytotoxicity of polyethylenimine and polyethylenimine-modified GO was found to be dependent on molecular weight, structure, and concentration of polyethylenimine [88,92]. The viability of HeLa cells remained high following treatment with free, branched, 1.2-kDa polyethylenimine or the GO-conjugated form at a concentration of 100 mg/l [88]. In the case of 10-kDa polyethylenimine, the viability of HeLa cells was more than 4-fold higher following delivery using the GOconjugated form than was the case using the free form at the same polyethylenimine concentration (100 mg/l) [88]. The viability of HeLa cells treated with branched, 25-kDa polyethylenimine-conjugated GO was about 2-fold higher than that of cells treated with the same concentration (8 mg/l) of branched, 25-kDa polyethylenimine [30,89]. Another study reported that the viability of PC-3 cells was about 1.5-fold higher after treatment with branched, 25-kDa polyethylenimine in GO-conjugated form compared to treatment with the free form [91]. Linear, 25-kDa polyethylenimine-conjugated GO was found to be nontoxic to HeLa and HEK 293 cells [92]. However, in this study, only a single concentration was tested, and cell viability after treatment with free linear polyethyleneimine was also greater than 90% [92]. A similar improvement in toxicity profile was observed for linear, 65-kDa polyethylenimine-grafted GO, with H293T and U2OS cells treated with the GO-grafted form exhibiting more than 3-fold higher viability than cells treated with the free form [93]. To improve gene delivery efficiency, Li and colleagues [94] incorporated a nuclear localization signal (NLS) peptide (PKKKRKV) onto 10-kDa polyethyleneimine-modified GO through electrostatic and hydrogen bond interactions. Tethering of the NLS peptide to polyethylenimine-modified GO was found to improve transfection efficiency and enhance the nuclear delivery of plasmid DNA. Polyamidoamine dendrimers and oleic acid-functionalized graphene have been studied for delivery of plasmid DNA [95]. In this study, oleic acid was first adsorbed onto graphene nanosheets to provide a carboxyl group for chemical modification with polyamidoamine dendrimers. Plasmid DNA encoding green fluorescent protein was complexed to cationic 27

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polyamidoamine dendrimers and oleic acid-functionalized graphene through electrostatic interactions and transfected into HeLa human cervical cancer and MG-63 human osteosarcoma cells. In HeLa cells, polyamidoamine dendrimers and oleic acid-functionalized graphene were found to exert cytotoxicity similar to that of polyamidoamine dendrimerconjugated GO, with more than 70% of HeLa cell remaining viable after treatment with 100 mg nanocarriers/l. Polyamidoamine dendrimers and oleic acid-functionalized graphene showed higher transfection efficiencies than polyamidoamine dendrimer-conjugated GO.

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Another platform that has been tested for the delivery of plasmid DNA is GO conjugated with chitosan, a cationic natural polysaccharide. In this latter study [55], chitosan was conjugated to GO using carbodiimide as a coupling agent. The resulting chitosan-conjugated GO formed complexes with plasmid DNA that showed complete gel retardation at N/P ratios higher than 4. At an N/P ratio of 8:1, chitosan-conjugated GO and plasmid DNA complexes were found to be 160 nm in size. Although chitosan-grafted GO showed a lower transfection efficiency of plasmid DNA than free branched 25-kDa polyethylenimine at an N/P ratio of 10, it was less toxic in HepG2 cells, even at a concentration of 100 mg/l.

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Chitosan derivatives have been introduced onto GO to allow electrostatic loading of plasmid DNA [96] (Fig. 4B). In this study, which focused on characterization of modified GO after complexation with plasmid DNA, GO surfaces were coated with folate-conjugated trimethyl chitosan and loaded with plasmid DNA. Folate- and trimethyl chitosan-modified GO exhibited a diameter of 112 nm, a thickness of 3.0 nm, and a zeta-potential of 30.9 mV. It was found that plasmid DNA was gradually released from folic acid/trimethyl chitosanconjugated GO up to 30% over the course of 80 hours in vitro in phosphate buffer. In vitro toxicity tests showed that folate/trimethyl chitosan-modified GO at concentrations up to 80 mg/l did not reduce the viability of HeLa cells.

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Chitosan has also been used as an anchor for surface modification of GO with other targeting ligands. Taking advantage of non-covalent functionalization of negatively charged GO with positively charged chitosan, chitosan was conjugated to folic acid [97] and cyclic RGD (Arg-Gly-Asp) [98]. The attachment of chitosan derivatives of folic acid and cyclic RGD provided enhanced drug delivery to target receptor-overexpressing cells. Graphene and carbon nanotube composites have been designed for transfection of plasmid DNA encoding green fluorescent protein [99]. In this application multiwalled carbon nanotubes were prepared, and graphene and carbon nanotube nanocomposites were synthesized by oxygen plasma etching technology.

5.2. Delivery of siRNA Polyethyleneimine-conjugated GO has been used for co-delivery of siRNA and doxorubicin [30]. In this study, branched, 25-kDa polyethylenimine was covalently conjugated to GO (Fig. 4C) using carbodiimide as a coupling agent. Polyethylenimineconjugated GO showed a zeta potential of +55.5 mV and electrostatically bound siRNA specific for the anti-apoptotic protein, Bcl-2. Treatment of HeLa cells with polyethylenimineconjugated GO/siRNA complexes significantly decreased Bcl-2 expression levels. This study 28

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also employed sequential co-delivery of siRNA and doxorubicin. For this, HeLa cells were first treated with polyethyleneimine-conjugated GO/siRNA complexes for 48 hours, then treated with polyethylenimine-conjugated GO loaded with doxorubicin for an additional 24 hours. Co-delivery of doxorubicin with polyethyleneimine-conjugated GO complexed with Bcl-2–specific siRNA sequences increased the potency of doxorubicin compared with doxorubicin-loaded polyethylenimine-conjugated GO complexed with control (scrambled) siRNA, decreasing the IC50 of doxorubicin from 1.3 mg/l to 0.52 mg/l.

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Co-delivery of siRNA and doxorubicin using polyethylenimine-modified GO has been shown to overcome multidrug resistance of breast cancer cells [100]. Branched, 25-kDa polyethylenimine and poly(sodium 4-styrenesulfonates) were assembled onto GO using a layer-by-layer assembly method. siRNA targeting the multidrug resistance-associated micro RNA, miR-21, was complexed through electrostatic interactions, and doxorubicin was loaded by physical adsorption. In doxorubicin-resistant MCF-7 cells, co-delivery of anti-miR-21 siRNA and doxorubicin using the functionalized GO resulted in higher cellular accumulation compared with free doxorubicin. The co-delivery system was shown to enter cells via caveolae- and clathrin-mediated endocytosis.

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Plasmid DNA encoding short hairpin RNA (shRNA) has been delivered using polyethylenimine and PEG-functionalized GO [101]. In mouse malignant melanoma B16 cells, treatment with plasmid DNA encoding a Stat3 oncogene-specific shRNA induced activation of the pro-apoptotic caspase 3 and down-regulated the oncoproteins Bcl-2 and cmyc. After intratumoral injection (20 g/mouse) into B16 tumor-bearing C57BL/6 mice, Stat3 shRNA-encoding plasmid DNA delivered using polyethylenimine and PEGfunctionalized GO was shown to more effectively inhibit tumor growth than scrambled shRNA-encoding plasmid DNA delivered by the same nanocarrier.

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Polyethyleneimine and PEG-functionalized GO have been used for photothermal enhancement of plasmid DNA and siRNA delivery [102]. Near-infrared laser irradiation at 808 nm (power, 0.5 W/cm-2) was found to enhance the transfection efficiency of plasmid DNA encoding green fluorescent protein in HeLa cells and enhance siRNA-mediated knockdown of mRNA for the target, polo-like kinase 1, in MDA-MB-435s cells. This enhanced nucleic acid delivery was attributed to an increase in cell membrane permeability caused by a photothermal effect under mild irradiation conditions. PEGylated GO and poly(2-dimethyl aminoethyl methacrylate) nanohybrids have been used for siRNA delivery [103]. In this study, PEG was chemically conjugated to GO via amide bonds, and poly(2-dimethyl aminoethyl methacrylate) was introduced onto PEGylated GO via hydrophobic interactions. Luciferase-specific siRNA delivered in complexes with PEGylated GO/poly(2-dimethyl aminoethyl methacrylate) showed a silencing efficiency in luciferase-expressing HeLa cells comparable to that observed following delivery of the siRNA using the commercial transfection reagent, Lipofectamine 2000. PEGylated GO and gold composite have been studied for siRNA delivery [104]. In this study, GO was loaded with branched, 25-kDa polyethyleneimine via charge-charge interactions, coated with AuCl4-, and chemically conjugated with PEG. Bcl-2–specific siRNA was complexed to PEGylated GO and gold composite or to Lipofectamine 2000, as a control. siRNA delivered on PEGylated GO and gold composites was shown to reduce the expression 29

ACCEPTED MANUSCRIPT of Bcl-2 to a similar degree as Lipofectamine 2000-mediated siRNA transfection in HL-60 human promyelocytic leukemia cells.

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Folic acid-conjugated GO has also been used for siRNA delivery [105]. To load siRNA, these researchers introduced 1-pyrenemethylamine hydrochloride onto folic acid-conjugated GO (Fig. 4D). 1-Pyrenemethylamine hydrochloride has aromatic ring structures, which are suitable for - stacking with GO, and a primary amine group, which is positively charged to allow electrostatic siRNA loading. Transfection of HeLa cells with siRNA targeting telomerase reverse transcriptase using 1-pyrenemethylamine hydrochloride/folic acidconjugated GO complexed with siRNA reduced mRNA and protein expression levels of human telomerase reverse transcriptase.

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Co-delivery of siRNA and doxorubicin has been studied using GO chemically modified with folic acid-conjugated oligochitosan [106]. In this application, GO was grafted with folic acid-conjugated oligochitosan and further modified with a quaternary ammonium group. The folic acid- and quaternary ammonium-modified cationic GO was complexed with siRNA or adsorbed with doxorubicin. Sequential delivery of MDR1-specific siRNA-loaded GO followed by doxorubicin-loaded GO reduced the IC50 value of doxorubicin from 15.6 mg/l to 10.3 mg/l in doxorubicin-resistant MCF-7 cells. This improvement in the anticancer effects of doxorubicin was attributed to the downregulation of P-glycoprotein by MDR1-specific siRNA.

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Plain forms of GO have not been used for nucleic acid delivery. Using GO as a nucleic acid delivery system requires modification with cationic polymers, such as polyethylenimine or chitosan, to allow electrostatic complexation with plasmid DNA or siRNA. Examples of cargo and modifying materials loaded onto GNS are summarized in Table 3. Delivering nucleic acid drugs using unmodified GO with possibly toxic cationic materials may remain as a future challenge.

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Moreover, although progress has been made in the delivery of siRNA using GO, most such studies have investigated these issues at the in vitro level (Table 3). Because siRNA is unstable in the blood and faces penetration barriers under in vivo conditions, many siRNA delivery systems that have proved satisfactory in vitro have not been suitable for in vivo use. For in vivo applications, siRNA delivery systems must satisfy biocompatibility and nonimmunogenicity criteria. Moreover, siRNAs must be protected from serum nucleases and effectively delivered to target tissues [107]. Future in vivo evidence of effective siRNA delivery using GO may support the feasibility of further applications of GO for siRNA delivery systems.

5.3. Delivery of protein and peptide drugs To date, there has been little progress in applying GNS for delivery of protein or peptide drugs. However, a basic study on the interaction of GO with amino acids, peptides, and proteins has been done [108]. In this study, the adsorption of amino acids, peptides, and proteins onto GO was measured using the fluorescence-quenching effect. The efficiency of tryptophan quenching caused by binding to GO was found to be pH-dependent, exhibiting 30

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diminished quenching at pH 9.0 compared with that at pH 5.6. These results suggest the involvement of electrostatic interactions between GO and tryptophan. Moreover, surface coating of GO with poloxamer 407 was shown to reduce the quenching efficiency of tryptophan owing to competitive effects. The fact that poloxamer 407 coats the GO surface via hydrophobic interactions suggests that such interactions also play a role in tryptophan binding to GO. These data provide evidence that peptides and proteins containing tryptophan residues can bind to GO through both electrostatic and hydrophobic interactions. To allow non-covalent loading of bioactive peptides on rGO, Oh and colleagues designed a chimeric peptide [26]. In this study, the cell-penetrating peptide, buforin IIb, was tethered to rGO through  interactions using a seven-phenylalanine spacer as an anchoring moiety and tetra-aspartate as a repulsion spacer.

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PEG-modified GO [25] and graphene multilayers [109-111] have been used for protein delivery. Zhang and colleagues [25] conjugated amine-terminal, six-arm PEG to GO using carbodiimide-mediated coupling and non-covalently loaded proteins onto the PEGylated GO via - stacking, hydrophobic interactions, and electrostatic interactions. Protein loading onto GO was confirmed by monitoring fluorescence quenching of fluorescein isothiocyanatelabeled bovine serum albumin, which is highly quenched when adsorbed onto GO. Protein adsorbed onto GO was found to be stable for 6 hours after exposure to trypsin; by comparison, naked protein was degraded after a 1-hour incubation with trypsin. Functional proteins, such as ribonuclease A and protein kinase A, have been delivered using PEGylated GO and shown to retain their activity in cells.

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GO multilayer systems have been reported as a sustained release system for protein delivery [109]. Using ovalbumin as a model protein, the authors of this report prepared a layer-by-layer assembly using GO and cationic poly(β-amino ester). As the number of GO layer increased, the time required to release ovalbumin from the GO multilayer assembly also increased. For example, a 20-layer GO and poly(-amino ester) assembly required 62 days to release 75% of bound ovalbumin. Multilayer systems composed of poly(β-amino ester), GO, and rGO have also been used for controlled protein release via electrical stimulation [110]. The release of ovalbumin from poly(-amino ester) and multilayer assembly was 50 times higher in the electrical-stimulation group compared with the unstimulated control group. Poly(allylamine hydrochloride)-modified GO multilayer capsules have been studied for protein delivery [111]. Here, poly(allylamine hydrochloride)-modified GO multilayer capsules were assembled by layer-by-layer adsorption of cationic poly(allylamine hydrochloride), and fluorescein isothiocyanate-conjugated bovine serum albumin and fluorescein isothiocyanate-conjugated insulin were encapsulated in the shell of multilayer nanocapsules.

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ACCEPTED MANUSCRIPT Table 3. Examples of nucleic acid-based therapeutics and proteins loaded onto GNS.

Luciferase Survivin shRNA

GO GO

GFP

GO

siRNA

Bcl-2 miR-21

GO GO

Stat3

GO GO

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Polo-like kinase 1

[88], [89], [90]

Branched polyethyleneimine

PC-3, NIH/3T3, HeLa cells

[8], [91]

Linear polyethyleneimine

HEK293, H293T, U2Os, HeLa cells

[92], [93]

Nuclear localized signal peptide (PKKKRKV), branched polyethyleneimine Polyamidoamine dendrimer, oleic acid

HeLa, 293T cells

[94]

HeLa, MG-63 cells HeLa cells HeLa cells

[95]

NIH-3T3, NG97 cells HeLa cells Doxorubicinresistant MCF-7 cells B16 cells

[99]

Chitosan Folate conjugated trimethyl chitosan Carbon nanotube

Branched polyethyleneimine Branched polyethylenimine, poly(sodium 4styrenesulfonates) Branched polyethylenimine, PEG Branched polyethylenimine, PEG PEG, poly(2-dimethyl aminoethyl methacrylate) Branched polyethylenimine, gold, PEG 1-pyrenemethylamine hydrochloride

Luciferase

GO

Bcl-2

GO

Human telomerase reverse transcriptase

GO

MDR1

GO

Folic acid conjugated chitosan oligosaccharide

Ribonuclease A, Protein kinase A Ovalbumin Ovalbumin Bovine serum albumin TRAIL

GO

Six-arm PEG

GO GO, rGO GO

Poly(β-amino ester) Poly(β-amino ester) Poly(allylamine hydrochloride) PEG

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HeLa, H293T cells

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GFP

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Plasmid DNA

Modification materials

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GFP

Type of GNS GO GO GO rGO GO GO GO GO

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Target gene

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Cargo

GO

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[55] [96]

[30] [100]

[101]

HeLa cells

[102]

HeLa cells

[103]

HL-60 cells

[104]

HeLa cells

[105]

Doxorubicinresistant MCF-7 cells HeLa cells

[106]

In vitro release In vitro release Encapsulation of protein A549-xenografted mice

[109] [110]

[25]

[111] [112]

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Gu and colleagues [112] first reported sequential co-delivery of protein and anticancer agent using GO in an animal model. In this application, PEGylated GO was conjugated with the anticancer protein, TRAIL (tumor necrosis factor-related apoptosis-inducing ligand), and noncovalently loaded with doxorubicin. Notably, taking advantage of furin on the cell membrane, these researchers inserted a furin-cleavable peptide between PEG and TRAIL to trigger the release of TRAIL in the tumor environment after passive delivery to tumor tissues via the enhanced permeability and retention effect. In A549-xenografted mice, intravenous co-delivery of furin-cleavable TRAIL and doxorubicin using GO was shown to inhibit tumor growth more effectively than co-delivery of furin-noncleavable TRAIL and doxorubicin on GO, and delivery of TRAIL and doxorubicin alone using GO.

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Delivery of proteins or peptides using GNS is still in its infancy. As shown in Table 3, most studies to date have used GO for protein or peptide delivery, and only a few studies have used rGO. Moreover, few studies have attempted to load therapeutic peptides and proteins on GO and demonstrate their functional activity following delivery. However, previous studies showing adsorption of proteins and peptides onto GO and enhanced stability against proteolytic enzymes indicate the potential of GO as a new carrier for therapeutic protein and peptide cargoes in cases where effective intracellular delivery systems are required.

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GNS have been modified with various targeting ligands in order to achieve target tissue or cell-directed delivery of therapeutics. Most targeting ligand approaches employed to date have focused on cancer. Although nanomaterials tend to selectively accumulate in tumor tissue owing to the enhanced permeability and retention effect, retention does not guarantee specific uptake by tumor cells. To improve cancer cell-specific uptake, researchers have modified GNS with various targeting ligands.

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Cancer cell-targeting strategies are based on the use of ligands for receptors or surface molecules that are overexpressed on cancer cells compared with normal cells [113]. Targeting ligands include folic acid, monoclonal antibodies, DNA aptamers, transferrin, peptides, and hyaluronic acid (Fig. 5). Ligands used for targeted delivery of GNS and their modification methods are described in Table 4. Recently, tumor microenvironment conditions have been used for tumor tissue-specific activation of targeting ligands on GNS [114]. Surface modification of GO with chitosan has been used to anchor other targeting ligands. Taking advantage of non-covalent functionalization of negatively charged GO with positively charged chitosan, researchers conjugated chitosan to folic acid [97] and cyclic RGD peptide [98], showing that attachment of chitosan derivatives of folic acid and cyclic RGD peptide provided enhanced drug delivery to cells overexpressing target receptors.

6.1. Small chemicals It has been reported that folate receptors are overexpressed on the surfaces of numerous 33

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cancer cells, including breast, ovary, lung, kidney, head and neck, brain, and myeloid cancers [115]. Given this expression pattern, it is not surprising that folic acid has long been used as a targeting ligand in various nanocarriers. This concept of folic acid-mediated tumor delivery has been exploited in a host of drug-delivery applications using GO or rGO.

Fig. 5. Schematic illustration of GO or rGO modified with targeting moieties. GO was modified with a cholesterol derivative of folic acid (A) [41], a monoclonal antibody (B) [19], NGR peptide (C) [140], or transferrin (D) [135]. rGO was modified with a DNA polyaptamer (E) [132], or a cholesterol derivative of hyaluronic acid (F) [87]. 34

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In one such application, rGO [116] or GO [117] was covalently or non-covalently modified with folic acid for enhanced drug delivery to tumors. Chemical reduction of a mixture of GO and folic acid using hydrazine monohydrate was found to allow non-covalent functionalization of rGO with folic acid through van der Waals interactions [116]. Noncovalent modification of rGO with folic acid enhanced the delivery of doxorubicin to MDAMB-231 cells. Folic acid has also been covalently conjugated to poly(aminoamine) dendrimer-modified GO for enhanced doxorubicin delivery to HeLa cells [117].

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In another study, folic acid was first conjugated to chitosan, and then the folic acidchitosan conjugates were used to modify the surfaces of GO [106]. Folic acid has been attached to GO functionalized with sulfonic acid groups for dual drug delivery [41] (Fig. 5A). In this application, Zhang and colleagues functionalized the surfaces of GO with sulfonic acids, and conjugated folic acid for co-delivery of doxorubicin and camptothecin.

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Several groups have reported enhanced cellular uptake of GO [24, 117-119] or rGO [120] modified with folic acid. Cellular uptake efficiency of GO covalently conjugated with folic acid via amide bonds was tested by confocal Raman scattering micro-spectroscopy in folate receptor-positive HeLa human cervical cancer cells and folate receptor-negative A549 cells. These studies revealed that folic acid-modified GO was internalized in folate receptorpositive HeLa cells by receptor-mediated endocytosis [118]. Others have also reported enhanced drug delivery using folic acid-conjugated hybrid GO in HeLa cells [117,119]. Similarly, uptake of chlorin e6 photosensitizers in MGC803 human stomach cancer cells was reported to be higher after delivery using folic acid-conjugated GO [24]. Using a fluorescent dye, Wang and coworkers showed that delivery of folic acid-conjugated rGO was enhanced in CBRH7919 mouse liver cancer cells [120].

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Folic acid-modified PEGylated GO has also been investigated for the diagnosis and treatment of hepatocarcinoma [23]. In this study, gadolinium-diethylenetriamine-pentaacetic acid-poly(diallyl dimethylammonium) chloride (MagnevistTM) was adsorbed onto GO for magnetic resonance imaging. Following covalent conjugation of folic acid and PEG to GO, doxorubicin was loaded onto the folic acid-modified PEGylated GO, achieving a loading capacity of 1.4 mg drug per carrier. Magnetic resonance imaging showed that the signal intensity of folic acid-modified PEGylated GO in HepG2 human hepatocellular carcinoma cells was higher than that of unmodified PEGylated GO. In keeping with this, doxorubicin delivered using folic acid-modified PEGylated GO was more effective in killing HepG2 cells than was observed upon co-treatment with free folic acid in competition assays. In an application of folic acid-conjugated GO for delivery of photosensitizers [24], folic acid-conjugated GO was found to load chlorin e6 with a loading efficiency of 80%, an efficiency higher than that achieved by other nanocarriers, such as polymeric micelles and silica nanoparticles. This higher loading efficiency is thought to be attributable to the large surface area of GO, which provides two accessible sides for drug loading. Uptake of folic acid-conjugated GO loaded with chlorin e6 by MGC803 human gastric cancer cells was high in the absence of folic acid in the culture media, but little uptake was observed in the presence of folic acid, suggesting that delivery into the cell is folic acid receptor-mediated. 35

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Folic acid has been conjugated to a superparamagnetic GO-Fe3O4 nanohybrid for dualtargeted drug delivery through magnetic stimuli and a chemical ligand [121,122]. In this application, a superparamagnetic GO-Fe3O4 nanohybrid was first prepared using a chemical precipitation protocol. As a targeting ligand for breast cancer cells, folic acid was conjugated onto the GO-Fe3O4 nanohybrid. Doxorubicin delivered using folic acid-modified GO-Fe3O4 nanohybrid was shown to exert a greater anticancer effect in the SK3 human breast cancer cell line than the GO-Fe3O4 nanohybrid alone [121]. Doxorubicin-loaded, acid-modified GOFe3O4 nanohybrid was internalized into cancer cells through two mechanisms: an external magnetic stimulus and receptor-mediated endocytosis.

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Although a number of studies have modified GO [23,24,117-119] or rGO [116, 120] with folic acid through covalent conjugation, non-covalent functionalization methods have several advantages over covalent conjugation from the standpoint of minimizing chemical reaction and purification steps [123]. “Host-guest” interaction was used for non-covalent functionalization of folic acid to the surfaces of GO [124]. Specifically, adamantine-grafted porphyrin was first loaded onto GO with the adamantine moiety exposed to the surface. The physical mixing of folic acid-modified -cyclodextrin with adamantine-grafted porphyrin allowed the ring structure of -cyclodextrin (host) to interact with adamantine (guest), resulting in non-covalent modification of the surfaces of GO with folic acid. Host-guest interaction-based modification of folic acid was shown to enhance the delivery and anticancer effect of doxorubicin compared to doxorubicin-loaded non-modified GO in folate receptorpositive HeLa cervical carcinoma cells, but not in folate receptor-negative OCT-1 mouse osteoblast cells.

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Non-covalent attachment of folic acid on the surfaces of rGO has also been reported using simple chemical reduction of a mixture of GO and folic acid [116]. In this study, incubation of folic acid and GO in the presence hydrazine monohydrate at 80°C resulted in the formation of rGO non-covalently modified with folic acid through van der Waals interaction. Folic acid-modified rGO was found to be stable in aqueous media under various pH conditions. Folic acid non-covalently attached on rGO was effectively delivered into folate receptor-expressing MBA-MB 231 human breast cancer cells. Another group reported folic acid-modified GO using chitosan as a mediator [122]. Although this study did not test the cellular uptake of folic acid and chitosan-modified Fe3O4/GO in folate receptor-overexpressing cell lines, it did report the differential loading capacity of doxorubicin on folic acid and chitosan-modified Fe3O4/GO as a function of pH. These researchers observed that folic acid and chitosan-modified Fe3O4/GO loaded 0.98 mg/mg of doxorubicin at pH 7.4, but only 0.74 mg/mg at pH 5.3. The authors attributed this pH-dependent loading of doxorubicin onto the carrier to pH-responsive hydrogen bond interactions between the drug and carrier, suggesting that protonation of amine groups in the drug reduced hydrogen bonding with GO carriers and thereby reduced loading capacity. pHresponsive interactions between amine group-containing drugs and graphenes could be invoked to explain the intracellular release of drug from graphenes in endosomal or lysosomal compartments. A recent study described the in vivo biodistribution of folic acid-modified GO [125]. In this study, GO was functionalized with folic acid-conjugated poloxamer 407 (MW, 12.6 kDa). 36

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Since receptor-mediated endocytosis is reported to depend on the ligand surface density [126], different ratios of folic acid-conjugated poloxamer 407 were used to decorate the surfaces of GO. In KB human epithelial mouth carcinoma cells, GO internalization increased steadily with an increase in the density of folic acid-conjugated poloxamer 407. In addition, by controlling the coating ligand density of folic acid-conjugated poloxamer 407, the distribution to the tumor could be increased by more than 2-fold compared with that to the liver.

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Numerous studies have reported enhanced delivery of folic acid-modified GNS in various cell lines, including HeLa human cervical cancer cells [117-119], HepG2 human hepatocellular carcinoma cells [23], MGC803 human stomach cancer cells [24], MBA-MB 231 human breast cancer cells [116], CBRH7919 mouse liver cancer cells [120], and SK3 human breast cancer cells [121]. However, few studies have clearly demonstrated the biodistribution and anticancer effects of these folic acid-modified graphenes in tumor xenograft animal models.

6.2.1. Integrin v3 antibody

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An integrin v3 antibody has been used for integrin receptor-mediated delivery of doxorubicin on GO [65]. Among integrins, integrin v3 plays the most important role during angiogenesis. The suppression of integrin v3 with targeting antibodies, antagonists, or peptides has been shown to produce clear anticancer effects in various cancers [127]. For surface modification of GO, amino groups of an integrin v5 monoclonal antibody were covalently linked to carboxylated GO [65]. Integrinv5 antibody-modified GO was found to increase the uptake of doxorubicin in human glioblastoma U87MG cells overexpressing integrin v3 compared with plain GO.

6.2.2. CD20 antibody An ultra-small GO has been covalently conjugated with rituximab, a chimeric monoclonal antibody against the CD20 molecule found on the plasma membrane of B cells [128]. In this application, thiolated Rituxan was conjugated to the amine groups on PEGylated GO through a sulfosuccinimidyl 4-(N-maleimidomethyl)cyclohexane-1-carboxylate linker [19] (Fig. 5B). A photoluminescence study revealed that rituximab-modified PEGylated GO recognized CD20-positive B cells, but not CD20-negative T cells.

6.2.3. Protein tyrosine kinase 7-specific aptamer DNA aptamers, which specifically bind to target molecules with high affinity, have been utilized as targeting ligands to enhance drug delivery [129]. A number of aptamers capable of recognizing target receptors have been reported [130]. One such DNA aptamer suitable for targeted anticancer drug delivery is protein tyrosine kinase 7, a biomarker for leukemia cells 37

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[131]. Oh and colleagues [132] synthesized a protein tyrosine kinase 7-specific DNA polyaptamer using the rolling-circle amplification method. Given the single-stranded oligonucleotide-binding capability of rGO, polyaptamer sequences of protein tyrosine kinase 7 were designed to include an oligoT bridge that acts as an rGO nanosheet anchoring moiety (Fig. 5E). Surface coating of rGO nanosheets with protein tyrosine kinase 7-specific DNA polyaptamers increased the efficiency of cellular uptake in CCRF-CEM leukemia cells overexpressing protein tyrosine kinase 7. In vivo molecular imaging demonstrated enhanced distribution of DNA polyaptamer-modified rGO to CCRF-CEM tumor tissues compared with plain rGO. In keeping with this enhanced tumor distribution, doxorubicin delivered using DNA polyaptamer-modified rGO exhibited a greater antitumor effect against CCRF-CEM tumor growth in nude mice compared with doxorubicin delivered using plain rGO or rGO modified with scrambled-sequence polyaptamers.

6.3.1. Transferrin

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DNA aptamers have also been studied for GO-mediated delivery of specific moleculeresponsive drugs [133]. In this application, an adenosine-5-triphosphate-binding DNA aptamer and two ssDNA linker sequences were mixed with GO to form a layered structure of DNA-GO nanoaggregates. Upon recognition of cytoplasmic adenosine-5-triphosphate by the aptamers, the nanoaggregate structures dissociate through detachment of the DNA aptamer from its target, releasing doxorubicin from GO at an intracellular site where adenosine-5triphosphate levels are high.

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Transferrin is the natural ligand for the transferrin receptor. Because expression of the transferrin receptor is highly upregulated on the surface of several types of tumor cells, the transferrin-transferrin receptor complex has been considered to provide a valuable mechanism for enhancing anticancer drug delivery and improving cancer diagnosis [134]. Xing and colleagues [135] reported dual modification of GO with transferrin and dihydroartemisinin, an anti-malarial drug, in which dihydroartemisinin was conjugated onto carboxylated GO (Fig. 5D). For transferrin modification, a PEGylated phospholipid derivative of transferrin was used to provide hydrophobic interactions with the phospholipid moiety and GO. Dihydroartemisinin/transferrin-modified GO was shown to be taken up by EMT6 murine mammary tumor cells. The reduction of dihydroartemisinin/transferrinmodified GO in the presence of transferrin was offered as support of transferrin receptormediated endocytosis. Molecular imaging of EMT6 tumor-bearing mice revealed the accumulation of dihydroartemisinin/transferrin-modified GO in tumor tissues. Dihydroartemisinin/transferrin-modified GO exhibited dramatic in vivo antitumor activity, resulting in 100% survival of treated mice. In contrast, mice treated with only dihydroartemisinin-modified GO showed only 50% survival. This enhanced antitumor efficacy of dihydroartemisinin/transferrin-modified GO was attributed not only to targeted delivery, but also to the cascade activity of ferric ion and dihydroartemisinin. After cellular uptake, ferric ion (Fe III) can be released from dihydroartemisinin/transferrin-modified GO because of the low pH of lysosomes and converted to ferric ion (Fe II) by intracellular ferric reductase. The reaction between ferric ion and dihydroartemisinin has been shown to produce 38

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6.3.2. Vascular endothelial growth factor-121

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Inhibition of tumor angiogenesis using targeted nanoparticles has been considered to be an excellent strategy for treating cancer. Vascular endothelial growth factor receptor is remarkably overexpressed in the vasculature of diverse solid tumors, where it regulates angiogenesis. Among the various molecules used to target vascular endothelial growth factor receptor are synthetic small ligands, anti-vascular endothelial growth factor receptor antibodies, and peptides [136].

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A recent report described the conjugation of vascular endothelial growth factor 121, a natural vascular endothelial growth factor receptor ligand, to GO [137]. In this application, PEGylated GO was functionalized with 1,4,7-triazacyclononane-1,4,7-triacetic acid and reacted with vascular endothelial growth factor 121-SH, resulting in vascular endothelial growth factor 121-conjugated GO.

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6.3.3. RGD peptide

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The RGD peptide, composed of the three amino acids Arg-Gly-Asp, is known to be a ligand for integrin receptors that binds preferentially to v3 and v5 integrin receptors with high affinity and selectivity. Targeting tumor cells or tumor vasculature using RGD peptidebased strategies has been considered a promising approach for delivering drugs or imaging agents for cancer therapy and diagnosis [138]. Cyclic RGD peptide-modified chitosan-GO has been tested for integrin receptor-directed delivery [98]. In this study, GO was noncovalently modified with cyclic RGD peptide-conjugated chitosan, and labeled with fluorescein isothiocyanate to visualize cellular uptake. In cancer cells expressing low levels of v3 integrin, the fluorescence signal was weak. By contrast, strong fluorescence was observed in cancer cells expressing high levels of v3 integrin.

6.3.4. NGR peptide

The NGR (Asn-Gly-Arg) peptide motif specifically recognizes the CD13 molecule. CD13, also known as aminopeptidase N, is a transmembrane ectopeptidase that cleaves proteins and peptides containing an N-terminal neutral amino acid. This enzyme is associated with the growth of diverse human cancers and has been suggested as a potential target for cancer therapy. NGR peptide-based drug-delivery systems and imaging-related applications have drawn increasing recent attention [139]. NGR peptide-conjugated GO has been reported as a tumor-targeting drug carrier [140]. In this report, researchers synthesized a GO-Ag nanocomposite by chemically depositing Ag nanoparticles onto GO using a hydrothermal reaction, and linked it to drug through ester bonds. Drug-loaded GO-Ag nanocomposites were further functionalized with a PEGylated 39

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phospholipid derivative of the NGR peptide (Fig. 5C). Consistent with the effective targeting ability of NGR, NGR peptide-modified GO/Ag nanocomposites exhibited specific delivery of drug to tumor cells.

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6.4. Polysaccharides

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Hyaluronic acid has been widely utilized as a nanomaterial targeting moiety for cancer diagnosis and treatment [2,141]. Hyaluronic acid is present in abundance in extracellular matrices, where it is known to contribute to cell motility and proliferation through interactions with a number of cell surface receptors, including CD44. Since CD44 receptors are overexpressed in various types of tumor cells, targeting CD44 receptors using hyaluronic acid has been widely studied for the selective delivery of drugs and diagnostics to tumor tissues [142]. This CD44 receptor-binding ability of hyaluronic acid has been exploited to modify the surfaces of GNS.

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Hyaluronic acid has been covalently conjugated to GO through formation of amide bonds [143]. Hyaluronic acid-grafted GO was found to exhibit enhanced cellular uptake in CD44-overexpressing HeLa cells. This compares with the negligible uptake of hyaluronic acid-grafted GO observed in L929 cells, in which expression of CD44 receptors is low. Treatment of HeLa cells in vitro with up to 200 mg/l hyaluronic acid-grafted GO did not significantly affect cellular viability. The study also tested the in vivo toxicity of hyaluronic acid-grafted GO following intravenous injection, reporting no abnormal hematological or biochemical effects 10 days after administration at a single dose of 10 mg/kg. The loading capacity of doxorubicin was found to be 0.82 g per gram of hyaluronic acid-grafted GO. Doxorubicin-loaded hyaluronic acid-grafted GO, administered intravenously at a dose of 6 mg/kg every 3 days, exerted higher antitumor effects against HeLa cell-bearing nude mice than mice treated with hyaluronic acid-grafted GO alone.

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Hyaluronic acid has been conjugated to GO together with a fluorescent dye using carboxymethyl chitosan as an intermediate linker [144]. In this application, GO was first conjugated with carboxymethyl chitosan using carbodiimide as a coupling agent, after which the carboxymethyl chitosan on GO was chemically conjugated with hyaluronic acid and fluorescein isothiocyanate. The resulting GO was loaded with doxorubicin, achieving a loading efficiency of 95%. Consistent with other studies, hyaluronic acid/fluorescein isothiocyanate/carboxymethyl chitosan-modified GO exhibited pH-dependent release of doxorubicin, showing higher drug release under acidic conditions (pH 5.8) than under neutral conditions (pH 7.4). An in vitro toxicity study revealed that the viability of HeLa cells treated with up to 6 mg/l of hyaluronic acid/fluorescein isothiocyanate/carboxymethyl chitosanmodified GO remained greater than 80%. Moreover, cellular uptake of this triply modified GO was higher in CD44 receptor-positive HeLa cells than in CD44 receptor-negative L929 cells.

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ACCEPTED MANUSCRIPT Table 4. Ligands used for targeted delivery of GNS.

Electrostatic interactions

GO GO GO

HepG2 cells MGC803 cells MCF-7 cells HeLa cells HeLa cells HeLa cells HeLa cells CBRH7919 cells SK3 cells

[23] [24] [41] [96] [117] [118] [119] [120] [121]

Doxorubicin release

[97] [122] [124]

Van der Waals interactions

rGO

HeLa-xenografted mice MDA-MB 231 cells

Hydrophobic interactions

GO

KB-xenografted mice

[125]

Integrin av3 antibody

Conjugation

GO

U87MG-xenografted mice

[65]

CD20 antibody

Conjugation

GO

Raji B cells

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DNA polyaptamer specific for protein tyrosine kinase 7

Hydrophobic interactions

rGO

CCRF-CEMxenografted mice

[132]

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Host-guest interactions

GO GO GO GO GO GO GO rGO GO

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Chemical conjugation

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Folic acid and its derivatives

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Modification method

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[116]

Transferrin

Hydrophobic interactions

GO

EMT6-bearing mice

[135]

VEGF121

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GO

[137]

Cyclic RGD

Electrostatic interactions

GO

U87MG-xenografted mice Bel-7402, SMMC7721, HepG2 cells

NGR peptide

Hydrophobic interactions

GO

S180-bearing mice

[140]

Hyaluronic acid and its derivatives

Conjugation

GO

[143]

Conjugation

GO

HeLa-xenografted mice HeLa cells

Hydrophobic interactions

rGO

KB-xenografted mice

[87]

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[98]

[144]

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A derivative of hyaluronic acid containing cholesterol, as a moiety for anchoring to rGO through non-covalent interactions, has been used to target CD44-overexpressing cancer cells [87] (Fig. 5F). Doxorubicin delivered using rGO modified with this cholesterol derivative of hyaluronic acid was shown to exhibit greater uptake in CD44 receptor-overexpressing KB cells. Modification of the surface of rGO with this derivative reduced the toxicity of rGO. Following intravenous administration of plain rGO at a dose of 40 mg/kg, only 40% of mice survived, whereas 100% of mice survived following administration of the same dose of hyaluronic acid cholesterol derivative-modified rGO. This study suggests that the acute in vivo toxicity of rGO can be controlled by modulating its surface properties.

7. Pharmacokinetics and biodistribution

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The in vivo fates of GNS are important issues for further development of GNS for delivery of various drugs. Pharmacokinetic profiles define the blood circulation and residence times of GNS in the body. Biodistribution provides information essential for predicting the safety of GNS for drug delivery. Pharmacokinetic studies have been done for PEGylated GO [145], 1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid-functionalized GO [146], poloxamer 407-functionalized graphene [147], and doxorubicin loaded onto polyethyleneimine-modified GO [65]. Radiolabeling and molecular imaging techniques have been used for detection of GO in the blood and various organs.

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Biodistribution studies of GO have been performed separately or in parallel with pharmacokinetic studies. Gamma counting [145], fluorescence imaging [47], single-photon emission computed tomography imaging [148], and positron emission tomography [149] have been used to study biodistribution. It has been shown that biodistribution patterns of GO are influenced by several factors, including geometric properties, lateral size, surfacemodifying polymers, and targeting ligands.

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I-radiolabeled, PEGylated GO has been used to evaluate the pharmacokinetics of sixarm-PEG–modified GO following intravenous administration at a dose of 4 mg/kg with 20 Ci 125I in BALB/c mice [145]. 125I-radiolabeled PEGylated GO showed a distribution halflife of 0.39 hours and an elimination half-life of 6.97 hours in a two-compartment model. Values for the pharmacokinetic parameters, area under the curve and volume of distribution, were 4.64 mg∙min/ml and 3.76 l, respectively. 125I-radiolabeled PEGylated GO, delivered at an intravenous dose of 4 mg/kg, was found to distribute mainly to the liver and spleen 1 hour post administration [145]. The accumulation of 125I-radiolabeled PEGylated GO in the liver and spleen gradually decreased over 60 days. In most other organs, the levels of 125Iradiolabeled PEGylated GO were less than 5% of the injected dose/g at 3 days post dose. Hematoxylin and eosin staining of major organs, including liver, spleen and kidney, showed gradual clearance of PEGylated GO, possibly via both renal and fecal elimination. Excretion in urine was higher than that in feces. The first day after dosing, excretion in urine was about 1.5-fold higher than that by feces, but increased to more than 7.2-fold higher beginning on the second day. Blood biochemistry and histology studies performed over 90 days post dose revealed no significant disorders, and histological analyses after a single intravenous injection of 20 mg/kg showed no obvious organ damage. 42

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Bianco and colleagues [146] used 111In-radiolabeled GO for a pharmacokinetic and biodistribution study in mice. These researchers functionalized GO with the radiometal chelating agent, 1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid, and radiolabeled 111 In. After intravenous administration at a dose of 50 g 111In-radiolabeled GO (~5–6 MBq) per C57BL mouse, the blood concentrations of GO showed an initial rapid decrease (within 1 hour) followed by much slower secondary elimination phase. At 1 hour post dose, the organ distribution of 111In-radiolabeled GO was highest in the liver, followed by the spleen and the lung. The amounts of 111In-radiolabeled GO in the liver tended to decrease over 24 hours. No specific organ accumulation was observed. Although the distribution to the kidney was not notable compared with that in the liver or spleen, the excreted amounts of 111In-radiolabeled GO was more than one order of magnitude higher in urine than in feces, indicating that the main route of excretion is renal rather than via the hepatobiliary pathway. The major role of the renal route for excretion of GO has been consistently reported in this [146] and a previous study [145].

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Oh and colleagues [147] reported structure-dependent differences in pharmacokinetic profiles of poloxamer 407-modified graphene nanosheets and single-walled carbon nanotubes. Graphene nanosheets or single-walled carbon nanotubes were modified with poloxamer 407 by non-covalent adsorption, and intravenously injected at a carbon dose of 5 mg/kg. In this study, a non-compartmental analysis was used to generate pharmacokinetic parameters. The mean residence time of poloxamer 407-modified graphene nanosheets was 21.0 hours—a duration 2.2-fold higher than that of single-walled carbon nanotubes.

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Although most studies have focused on the pharmacokinetics of GO or graphene nanosheets, a recent study reported the pharmacokinetics of doxorubicin after intravenous delivery using polyethyleneimine-modified GO [65]. Doxorubicin, covalently conjugated with citraconic anhydride-functionalized poly(allylamine) and electrostatically loaded onto polyethyleneimine-modified GO, was intravenously injected into BALB/c mice at a dose of 5 mg/kg. Doxorubicin delivered on GO exhibited a half-life of 2.1 hours, whereas that of free doxorubicin was 0.52 hours. Moreover, a histological analysis revealed no obvious toxicity in major organs, including liver, kidney, spleen and lung, after intravenous administration of doxorubicin (5 mg/kg)-loaded, polyethyleneimine-modified GO. Fluorescence imaging has been used to study the biodistribution of doxorubicin delivered by NGR peptide-modified GO/Ag nanocomposites [140]. In this study, the fluorescence of doxorubicin was used to trace the in vivo fate of the drug. At 3 hours after intravenous injection in mice, the distribution of the free drug to the heart and kidney was higher (>2-fold) than that of the drug given in plain GO/Ag nanocomposites or NGR peptidemodified GO/Ag nanocomposites. By contrast, doxorubicin delivered on plain GO/Ag nanocomposites or NGR peptide-modified GO/Ag nanocomposites preferentially accumulated in the liver, spleen, lung and tumor, where its distribution was higher than that of the free drug. Moreover, whereas doxorubicin delivered on GO/Ag nanocomposites showed a 2-fold higher distribution to the liver than to tumor tissues, the fluorescence intensity of doxorubicin delivered on NGR peptide-modified GO/Ag nanocomposites was more than 1.2-fold greater in the tumor than in the liver. The administration of doxorubicin in GO/Ag nanocomposites also increased the tolerated dose. Following intravenous administration of doxorubicin at a dose of 5 mg/kg as plain or NGR peptide-modified GO/Ag 43

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nanocomposites, 100% of mice survived, whereas only 50% of mice survived following administration of the same dose of free doxorubicin. Moreover, mice treated with doxorubicin as plain or NGR peptide-modified GO/Ag nanocomposites showed no reduction in body weight over 10 days post dose.

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The biodistribution of chlorin e6 delivered on PEGylated GO after intravenous injection in mice has been reported [47]. In SCC7 tumor-bearing mice, chlorin e6-loaded PEGylated GO exhibited a greater distribution to the tumor at 1 and 48 hours post injection than the same dose (10 mg/kg) of the free form of chlorin e6. This greater tumor distribution of chlorin e6 was attributed to the enhanced retention and permeability effect.

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Several biodistribution studies have been done to evaluate the performance of targeting ligands incorporated on the surfaces of GNS. Using single-photon emission computed tomography imaging, Vallis and colleagues [148] investigated the biodistribution of anti111 HER2 antibody (trastuzumab)-conjugated GO labeled with In-benzyldiethylenetriaminepentaacetic acid. Following intravenous injection into 231/H2N tumorxenografted mice, trastuzumab-conjugated GO accumulated in tumor tissues, reaching a concentration of 15% of the injected dose/g on day 3 post dose. The biodistribution of anti-CD105 antibody-conjugated and PEGylated GO labeled with Ga has been investigated in tumor-bearing mice [149]. Positron emission tomography imaging showed that 66Ga-labeled anti-CD105 antibody-conjugated/PEGylated GO intravenously injected into 4T1 tumor-bearing BALB/c mice accumulated in tumor tissues to a greater extent than in all other major organs at day 1 post administration. Notably, the distribution of 66Ga-labeled anti-CD105 antibody-conjugated/PEGylated GO to tumor tissues was 1.4-fold higher than that of 66Ga-labeled/PEGylated GO at this same time point. At 24 hours post dose, the distribution of 66Ga-labeled anti-CD105 antibody-conjugated/PEGylated GO was highest in liver followed by the spleen and the kidney for both 66Ga-labeled antiCD105 antibody-conjugated/PEGylated GO and 66Ga-labeled/PEGylated GO. Notably, this study tested the distribution to the brain, and found a negligible background level in the brain regardless of GO modification and time point studied. 64

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Cu-radiolabeling has been used to study the in vivo fate of PEGylated and modified GO [137]. Positron emission tomography imaging has also been used to track the fate of vascular endothelial growth factor121-modified and 64Cu-labeled six-arm PEGylated GO. Following intravenous injection into U87MG tumor-bearing BALB/c mice, radiolabeled and modified GO exhibited a 1.9-fold enhancement in tumor distribution (8% of injection dose/g) at 3 hours post injection compared with 64Cu-labeled PEGylated GO. However, the distribution of 64Cu-labeled vascular endothelial growth factor-modified/PEGylated GO to the liver, intestine, and lung was higher than that to the tumor. In particular, the distribution to the liver was still notable for both 64Cu-labeled vascular endothelial growth factormodified/PEGylated GO (20% of injection dose/g) and 64Cu-labeled PEGylated GO (17%of injection dose/g) at 3 hours post dose. The overall tissue levels decreased over time. At 2 days post dose, the organ with the highest distribution remained the liver, but unlike 3 hours post dose, the organ with the second-highest distribution of both 64Cu-labeled vascular endothelial growth factor121-modified/PEGylated GO and PEGylated GO was the kidney. The biodistribution of folate-modified GO has been investigated in a tumor-bearing mouse 44

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model [125]. For molecular imaging, these researchers intravenously injected folate-modified GO (10 mg/kg) labeled with the fluorescent dye Cy5.5 into KB tumor-bearing C3H/HeN mice. Compared with unmodified GO, folate-modified GO showed an increasing distribution to the tumor that positively correlated with the folate coating density.

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Hyaluronic acid derivative-modified rGO has also shown greater distribution to tumor tissues [87]. In this study, rGO modified with a Cy5.5-labeled cholesterol derivative of hyaluronic acid was intravenously injected (15 mg/kg) into a KB tumor xenograft nude mouse model. The distribution to KB tumor tissues was higher at day 1 and 2 post dose in mice treated with the hyaluronic acid derivative-modified rGO compared with those injected with unmodified rGO.

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The biodistribution of integrin 3 antibody-modified and anionic doxorubicin conjugate-loaded polyethyleneimine-modified GO has also been studied following intravenous injection into U87MG tumor-bearing nude mice [65]. This study, which measured doxorubicin in homogenates of organs and tumor tissues by fluorospectroscopy, showed that the distribution of doxorubicin to tumor tissues was higher after delivery with the antibody-modified GO than with the free form.

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To date, most biodistribution studies have been performed using normal or tumor-bearing mice, mainly addressing distribution patterns to tumor tissues in the case of target ligandmodified GNS. However, several studies have reported that even GNS modified with target ligands is also distributed to other organs, including the liver and spleen. Possible directions of future studies include biodistribution studies in non-rodent animal models and simultaneous investigation of biodistribution patterns of both drug and GNS.

8. Current challenges and future perspectives

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Although significant progress has been made in the field of drug delivery using graphenebased nanomaterials, several issues remain to be resolved and additional improvements must be made for effective clinical translation. One reason GNS have been the focus of such keen interest is the dual side-available planar form of the graphene surface, which is unlike that of rectangular or spherical nano-delivery systems. Because their ultrahigh surface area is composed purely of carbon, GNS offer the excellent advantage of enabling drug loading with high efficiency compared with other nanocarriers. However, this large surface area can also be a disadvantage, fostering interactions with blood components and formation of precipitates with red blood cells or serum proteins after in vivo administration. A key design feature essential for minimizing nonspecific interactions with blood components in the circulation and improving hemocompatibility is surface functionalization of GO with hydrophilic materials [47, 150]. PEGylation, which has been used to increase in vivo stability and circulation time of conventional nanoparticles, has been one approach for minimizing the interaction of GO with blood components [47]. Although such PEGylated GO exhibits in vitro cytotoxicity similar to that of non-modified GO, it exhibits dramatically reduced toxicity in vivo. For example, a single acute toxicity study demonstrated that all mice 45

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intravenously injected with 80 mg/kg PEGylated GO survived, whereas all mice injected with the same dose of plain GO died [47]. Moreover, PEGylation has been shown to enhance the in vivo stability of GO and thereby increase the delivery of cargo drug to tumor tissues [47]. In addition to PEG, biocompatible materials such as hyaluronic acid [87], chitosan [63], heparin [58], and serum albumin [49] have been shown to provide enhanced in vivo stability of GO. Judging from these studies, additional investigations of surface modifications of GO will provide further improvement in blood circulation and biocompatibility for systemic injection.

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Although several reports have demonstrated successful in vivo therapeutic efficacy of GO or rGO-based drug-delivery systems, the in vivo safety will need to be assured for further clinical applications. A previous study reported that GO and aggregated graphene induces severe lung injury and apoptosis, whereas well-dispersed graphene with Pluronic exhibits remarkably reduced toxicity [151]. Chu and colleagues [152] demonstrated that unmodified rGO can cause reproductive dysfunction and interfere with the development of the offspring of pregnant mice. Specifically, these researchers showed that mice injected with a high dose of rGO (25 mg/kg) at an early stage of pregnancy bore mostly healthy offspring, but some fetuses were born with a deformity. However, when injected late in pregnancy, low and medium doses of rGO (6.25 or 12.5 mg/kg) caused all pregnant mice to abort fetuses, and a high dose of rGO caused the death of all pregnant mice. Xu and colleagues [153] investigated the short- and long-term effects of orally administered rGO. With short-term treatment (5 days), they found that rGO-treated mice became slightly lethargic and showed blunted neuromuscular coordination. Liu and colleagues [145] used 125I-radiolabeled PEGylated GO to assess long-term toxicity in vivo. They monitored animals for 3 months following a single bolus injection of a high dose (20 mg/kg) of PEGylated GO and found no evidence of significant in vivo toxicity. The results of safety evaluation studies to date suggest that surface-modified GO have a favorable toxicity profile.

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The biodegradation and metabolism of GO used for drug delivery are crucial issues. Surface modification of GO has been reported to affect horseradish peroxidase-induced oxidization and degradation. Specifically, GO modified with PEG via a disulfide bond was found to be biodegradable by horseradish peroxidase-induced oxidization [49]. Thus, the design of biodegradable GO through suitable surface modifications is an important step towards clinical applications. The nanoscale lateral sizes of GNS, which have been reported to range from 5 to 3,000 nm (Table 1), are crucial for physiochemical stability and biological responses. Compared with micron size GO, nanoscale GO has been known to form stable dispersion due to the higher charge density derived from the higher edge-to-area ratios [154]. Nanoscale has been reported to have direct effects on protein binding affinities [155], cellular internalization mechanisms, metabolism, and toxicity [156-158]. Su and colleagues reported that micron-scale (2 μm) and nanoscale (350 nm) GO were taken up equally by macrophages through active Fc receptormediated phagocytosis, but only micron-size GO caused severe inflammatory responses after cellular uptake [156]. Another study demonstrated that the adsorption of micron-size GO (750–1300 nm) to the cell membrane induced M1 macrophage polarization [157]. GO with an average lateral size of 780 nm was shown to promote the generation of reactive oxygen species in A549 lung carcinoma cells [158]. Moreover, for the delivery of anticancer drugs to 46

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tumor tissues, passive targeting through the enhanced permeability and retention effect has been reported to be effective for nanocarriers less than 400 nm in size [159]. Thus, the lateral size of GNS is important for minimizing unwanted biological responses and enhancing the desired effects.

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In addition to the importance of size per se, methods for controlling the sizes of nanosheets should be studied further and standardized. Currently used methods for controlling size vary among laboratories. Although centrifugation has been the most commonly used method for obtaining GO, sonication and filtration have also been used (Table 1). In addition to average lateral size, the acceptable range of polydispersity should be considered. Although lateral size is a more important factor for the biological responses to GO, thickness needs to be evaluated as a quality control parameter for industrial production. For charged drug molecules that are bound to oppositely functionalized GO, zeta potentials can serve as another quality control parameter for confirming the loading of drugs.

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Despite the development of GO or rGO for biomedical applications, methods for manufacturing GNS differ from laboratory to laboratory [10,160]. The main methods for preparing GO are Brodie, Staudenmaier and Hummers’ methods, as well as variations on these themes [161,162]. rGO has also been prepared using a variety of reducing agents, including hydrazine hydrate, sodium borohydrate, and sodium hydrosulfite [163,164]. One recently reported method for preparing GO is through the pyrolysis of citric acid [165]. In this study, GO or graphene quantum dots were produced by controlling the degree of carbonization of citric acid. Moreover, methods for convenient and environmentally friendly preparation of various GNS need to be developed. Facile methods have been reported for preparing functionalized graphene nanosheets using bovine serum albumin or gelatin as reducing and stabilizing agents. These methods can avoid the need for toxic chemicals in the manufacturing process, and are thus more environmentally friendly.

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All of these methods can cause various levels of oxidation of graphene and differences in lateral sizes (Table 1), with the result being that GO and rGO reported in different studies are not the same. It is likely that differences in surface properties, including the density of functional groups and lateral sizes of GO or rGO prepared by different methods can affect drug-loading capacity and biological behavior. This is especially true in the case of GO, where the extent of reduction can affect hydrophobicity and interactions with hydrophobic drugs. A recent study reported that particle size, particulate state, and oxygen content/surface charge of GO have strong impacts on biological and toxicological responses in red blood cells and fibroblasts [166]. The type of biocompatible reducing agent, the reduction time, and standardized protocols used to quantify the extent of reduction need further study. The development of manufacturing and quality control standards for GNS is thus urgently needed to ensure the production of uniform GNS, which could enhance the reproducibility of laboratory studies and consequently support translation of GNS-based drug-delivery systems to the clinic. Numerous studies of GNS as drug-delivery systems are currently in progress, and the various merits of GNS-based nanomaterials are continuously being demonstrated by diverse research groups. However, some problems, such as biocompatibility and developmental feasibility as drug carriers, still await solutions. Although safety studies of graphene-based nanomaterials have consistently shown favorable results, there is still considerable room for 47

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improvement owing to the promising and unique characteristics of various types of GNS. Progress in the GNS field is being watched with keen interest to gauge whether the potential of GNS can be realized and ultimately lead to a new era of nanomedicine.

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Acknowledgment

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This work was supported by a research grant from the Ministry of Science, ICT and Future Planning (NRF-2015R1A2A1A01005674), Republic of Korea.

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