2.1 Introduction

The unique ability of carbon atoms to participate in robust covalent bonds with each other and also with other elements in diverse hybridization states (sp, sp² , sp³ ) allows them to form a wide range of structures, from small molecules to nanomaterials [1]. Carbon‐based nanomaterials have garnered a great deal of interest over the past three decades, starting with the debut of fullerene in 1985 [2], followed by carbon nanotubes [3] (CNTs) and graphene [4] in 1991 and 2004, respectively. Fullerene, CNTs, and graphene are primarily composed of sp² carbon atoms with a network of conjugated π electrons. In recent years, carbon dots (CDs) with mixed sp² and sp³ carbon atoms plus defects and heteroatoms, as well as nanodiamonds consisting of sp³ carbon atoms, have also received a great deal of attention [5, 6]. Due to the quantum confinement effect, these carbon nanomaterials (CNMs) possess many interesting physicochemical properties that are not attainable in bulk carbon materials like diamond and graphite [7, 8]. Over the past decade, there has been a flurry of development of biomedical research applications for graphene, CNTs, nanodiamonds, and CDs, owing to their small size, unique optical properties, and large surface area. In this chapter, we will focus on functionalization of CNMs (CNTs, graphene, nanodiamonds, and CDs) with polymers for biomedical applications. For these applications, a stable dispersion of CNMs in biological medium is a fundamental prerequisite that can be achieved by proper surface functionalization of the nanomaterial [9, 10]. Nonfunctionalized CNMs tend to form stable aggregates due to strong intermolecular interactions, such as van der Waals (vdW) forces and dipole–dipole interactions. Aggregation of nanomaterials results in unwanted changes to their physicochemical properties and makes homogeneous cell labeling very difficult. Polymer coating is an effective, convenient, and widely used approach for transferring CNMs into aqueous solution. Polymer coatings provide a number of distinct advantages over small‐molecule coatings, including ease of synthesis and processing, structural diversity, tunable surface functionality, and flexibility. Multiple contact points between the CNM surface and polymers provide a more robust surface coating than coating with small molecules, which typically have only one binding group. The strong binding also prevents desorption of ligands from the surface, which has been one of the sources of toxicity in practical applications. Most importantly, polymers can be designed to include reactive functional groups that can be further functionalized with biomolecules. The design and structure of polymers for functionalization of CNMs depend on the chemistry of the CNM and its potential applications in a biological environment. Polymer coating of CNMs not only increases the colloidal stability of the particles but also selectively changes or influences the CNM properties, such as reducing nonspecific adsorption of biomolecules or tuning the charge (for gene therapy applications). The polymer shell on the surface of CNMs, rather than the core, determines the major final properties of the coated CNMs. These properties are crucial for potential interactions of particles with a biological environment.

2.2 Colloidal Stability of CNMs

To successfully develop nanomaterials for biomedical applications, a detailed understanding of the factors driving the colloidal stability of nanoparticles (NPs) and their aggregation in biological medium is crucial. A classical theoretical approach based on Derjaguin–Landau–Verwey–Overbeek (DLVO) theory is commonly used to predict the stability of a colloidal system. This approach balances attractive vdW forces and repulsive forces caused by the electrostatic double layer (EDL) (Figure 2.1a,c) [12]. The attractive vdW forces are competed by the EDL which stabilizes the colloidal dispersion. In addition to electrostatic stabilization, other repulsive forces (e.g. hydration forces) and specific attractive forces (e.g. depletion and hydrophobic or magnetic forces) also contribute to the stabilization or destabilization of the colloidal system [13]. The DLVO theory has been extended (X‐DLVO) (Figure 2.1d) by adding these individual forces, assuming that each interaction is completely independent.

The colloidal stability of nanomaterials increases with increasing surface charge. The surface charge of NPs is commonly determined in terms of zeta potential. Zeta potential is the electrostatic potential of the particles measured at the so‐called shear plane, the distance from the surface at which ions are not bound to the particle. A zeta potential of ±30 mV is widely regarded as the borderline between stable and unstable colloids in aqueous solution. The net charge of a nanosystem is balanced by the counterions around the particle. Counterions adsorbed on the charged NP surface comprise the “Stern layer,” and a cloud of ions of opposite net charge surrounding the nanoparticles makes up the “diffuse layer” (Figure 2.1c). The ionic strength of the solution determines the radial size of the diffuse layer. Low ionic strength allows the diffuse layer to extend far from the particle surface, which repels particle–particle interactions. High ionic strength, as in electrolyte‐rich cell culture medium, induces the collapse of colloidal systems by compressing the EDL [14]. During endocytotic uptake by cells, NPs are exposed to varying conditions, such as pH changes from 7.4 (in the extracellular medium) to 5.5 (in late endosomes) to 4.5 (in lysosomes). Nanomaterials with amphoteric surface groups tend to have a neutral charge at pH values near the isoelectric point, which promotes aggregation due to decreased repulsion and enhanced attractive vdW forces [15]. In addition to their acidic pH, lysosomes also contain high levels of hydrolytic enzymes that can degrade the nanomaterial coating [16]. Thus, electrostatically stabilized NPs generally have poor stability in cellular medium and are not well‐suited for biomedical applications.

The influence of NPs on cellular interactions can be assessed by in vitro systems in which NPs interact with the components of the cell culture medium prior to any cellular contact. These culture media contain a complex mixture of proteins, such as serum albumin and globulins, and numerous biomolecules, such as amino acids and salts. The nonspecific adsorption of protein on the surface of NPs increases the hydrodynamic diameter and greatly reduces the mobility of the particles. Changes in the dispersion state of NPs have been shown to influence the cellular response and the fate of NPs in biological environments [17–19].

For successful design of the CNM biointerface, the following must be considered:

1. After functionalization, CNMs must be soluble in water and should not aggregate over a broad pH range, in the presence of high concentrations of electrolytes, nor in biological media.
2. The surface functionalization of the CNMs must promote biocompatibility and reduce nonspecific interactions.
3. Surface functionalization needs to be effective and controllable to permit adjustment of the number and nature of biomolecules attached to CNMs, thus facilitating their applications in targeting, biosensing, and bioimaging.

2.3 Functionalization of CNMs with Polymers

Surface polymers stabilize dispersion of CNMs by steric hindrance. Steric stabilization is a powerful tool to enhance the dispersion stability of CNMs under harsh conditions by preventing two particles from forming attractive vdW interactions due to osmotic pressure and elastic recoil effects (Figure 2.1b) [20]. Unlike pure electrostatic stabilization, steric stabilization also stabilizes NPs in the presence of high salt concentrations, which is important for biomedical applications. Using charged polymers, the stabilization of colloidal systems can be further improved due to combined electro‐steric repulsion [21, 22].

CNMs can be functionalized with polymers via two types of approaches: (i) noncovalent coating through electrostatic or hydrophobic interactions and (ii) covalent modification, which can be divided into “grafting to” and “grafting from” approaches.

2.4 Influence of Polymers on the Spectral Properties of CNMs

The physicochemical properties of CNMs are determined by their intrinsic nanostructure. The surface functionalization of CNMs with polymers can introduce defects into the nanostructures or alter the nano‐environment around the CNMs, which can lead to changes in their intrinsic properties. In this section, we will focus on the effect of polymer functionalization on the spectral properties of various CNMs.

The fluorescence of CNTs, which appears mainly in the second near‐infrared window (NIR‐II window, 1000–1700 nm), is distinct from that of other CNMs. CNTs such as single‐walled carbon nanotubes (SWCNTs) absorb a photon with the band gap of the E₂₂ transition (Figure 2.2a), which excites the electron from the second valence band to the second conduction band, leaving a hole behind. The excited electron and the hole form a bound pair called an exciton, which recombines to emit a photon in the NIR‐II region (Figure 2.2b). The excitonic nature of the photoluminescence properties of CNTs makes these nanomaterials highly sensitive to the environment, as well as to the length of the nanotube. Defects in the conjugated system of nanotubes act as discontinuities along their length. The defect sites in a SWCNTs can cause nonradiative recombination of diffusing excitons and lower the quantum yield. Noncovalent coating of SWCNTs with polymers avoids disruption of the π‐network of pristine CNTs and increases the luminescence quantum yield by one order of magnitude [29]. The photoluminescence peak of polymer‐coated nanotubes can be red‐shifted by creating a more polarizable environment around the CNTs due to π–π stacking [30]. Single‐stranded DNA can disperse CNTs in water by helically wrapping around the CNT surface through π‐stacking [31]. Kurnosov et al. found that cysteine doping (from 10⁻⁸ to 10⁻³ M) into an aqueous suspension of nanotubes wrapped with DNA leads to an increase in photoluminescence intensity [32]. The photoluminescence intensity was enhanced by 27% in the presence of 10⁻³ M cysteine, and the researchers attributed this increase to the passivation of p‐defects on the nanotube by the reactive thiol group of cysteine [32]. On the other hand, the fluorescence intensity of SWCNTs can be quenched upon noncovalent coating with a pyrene‐functionalized poly(3‐hexylthiophene) derivative, due to a strong photo‐induced electron transfer resulting from the noncovalent linkage of SWCNTs to the poly(3‐hexylthiophene) backbone via the pyrene bridge [33]. SWCNTs can be covalently functionalized with polymers via the carboxylic groups at the open ends of the SWCNTs. However, the damage to the conjugated π network during covalent functionalization of SWCNTs is a major concern. The defects created in the conjugated π network during covalent modification decrease the fluorescence quantum efficiency of NIR‐II emission, due to the increased probability of nonradiative recombination of excitons at the defect site [26].

A large, pristine graphene sheet is a zero band gap semiconductor with the conduction band and valence band touching at the K points. Graphene sheets typically are expected to be nonfluorescent. However, with decreasing size of the graphene sheet and introduction of functional groups at its edges, band gap opening is predicted due to quantum confinement and edge effects [34, 35]. The origin of fluorescence from graphene oxide (GO) remains poorly understood. Various mechanisms have been proposed, including nanometer‐sized aromatic domains in a GO sheet [36], strong electronic coupling between the carboxylic groups and the neighboring sp² graphitic carbon π‐electrons [37], and electronic transitions between the non‐oxidized pristine sp² graphitic carbon regions and oxidized carbon regions at the boundaries [38].

Exceptionally photostable, high‐quantum‐yield NIR fluorescence in nanodiamonds arises from color centers, such as nitrogen vacancy (NV) centers (Figure 2.2c) and silicon vacancy (SiV) centers, present in the diamond core. The emission state of NV centers in fluorescent nanodiamonds can be manipulated by controlling the nanodiamond's size or surface moieties or by changing the surface chemistry at the atomic level. The emission state is also sensitive to the action of the surrounding medium [39–41]. Bradac et al. demonstrated that it is possible to control the intermittency in the luminescence (blinking) of 5‐nm diamonds by embedding them in a poly(vinyl alcohol) film. The electronic discontinuity at the interface was altered, and the luminescence temporal behavior of individual NV centers was restored to a continuous luminescence signal [40]. Petrakova et al. reported modulating nanodiamond fluorescence through noncovalent interactions with charged polymers, such as poly(allylamine) and poly(diallyldimethylammonium chloride), by switching between NV¯ and NV⁰ energy states [42]. An analogous mechanism has been used to image DNA transfection in cells [43]. In this case, nanodiamonds were coated with cationic poly(ethyleneimine) and modified with DNA. After transfection into cells, the DNA detached from the particles, which resulted in switching between the NV¯ and NV⁰ energy states [43].

CDs have been widely studied as a green substitute for quantum dots due to their excellent photoluminescence and high biocompatibility. The mode of formation or the mechanism of photoluminescence origin (due to quantum effect, bandgap transition of conjugated π‐domains, different emissive traps on the CD surface, or molecular state) is a matter of debate. Polymers are increasingly becoming an important component for both synthesis and modification of CDs [44]. Surface passivation and functionalization with polymers lead to higher quantum confinement of emissive energy trapped on the CD surface, resulting in higher photoluminescence quantum yield [44]. Quantum‐sized CDs with a surface passivated with poly(propionylethylenimine‐co‐ethylenimine) (PPEI‐EI, with EI fraction ∼20%) showed strong luminescence with two‐photon excitation in the NIR spectral region [45]. Enhancement of emission by chemical crosslinking or physical immobilization of polymer chains on CDs is known as the crosslink‐enhanced emission (CEE) effect [46]. Tao et al. demonstrated the CEE effect by hydrothermally crosslinking polyacrylic acid (PAA) and ethylenediamine chains [47]. CDs developed from poly(thiophene phenylpropionic acid) demonstrated broad absorption (400–750 nm), red‐emission, a strong photoacoustic response, and high photothermal conversion efficiency (η ~ 38.5%) upon NIR laser irradiation, enabling them to act as multifunctional fluorescent, photoacoustic, and thermal theranostics for simultaneous diagnosis and therapy of cancer [48].

Raman spectroscopy provides useful information about the structure, electronic properties, composition, and homogeneity of CNMs, as well as information about the chemical environment around the nanomaterials. Changes in D‐band and G′‐band (D‐band overtone) Raman spectra can be used to probe and monitor structural modifications of the nanotube sidewalls that stem from the introduction of defects and the attachment of different chemical species [49]. Raman shift in the D* band (the second‐order overtone of the D band) is often reported for CNTs embedded in polymers, and can be used to measure the strain or stress applied to the CNTs from the surrounding media [50]. The D‐band in the Raman spectrum of GO arises from the stretching of sp³ carbons of graphene sheets, which occurs at 1366 cm⁻¹ . The G‐band arises from the stretching of sp² carbons and occurs at 1582 cm⁻¹ . The interaction between polymer and GO shifts the G‐band, due to the change in the electronic structure of graphene [51].

2.5 Functionalizing CNMs with Antifouling Polymers for Bioapplications

When nanomaterials are introduced into cell culture media, the constituent proteins coat the NP surface to form a new bio‐nano interface, the so‐called “protein corona” (Figure 2.3a) [52, 53]. The protein corona is composed of an inner layer of proteins with a lifetime of several hours in slow exchange with the environment (the hard corona) and an outer layer of weakly bound proteins characterized by a faster exchange rate with the free proteins (the soft corona) [53, 54]. It is the hard corona, rather than the pristine NP surface, that interacts with cellular receptors. To avoid nonspecific adsorption of proteins, dense polymeric shells with anti‐fouling properties are needed. Anti‐fouling polymers decrease the nonspecific interactions of nanoparticles with proteins and cells (Figure 2.3b), prolong the time the nanoparticle spends in the circulatory system, improve biodistribution, and can be further modified with a targeting moiety exposed to the surroundings. Coating polymers need to be hydrophilic and ideally neutral or negatively charged. They can be synthetic or natural (such as proteins and polysaccharides). For example, use of SWCNTs for biological research was pioneered by Hongjie Dai and colleagues, who introduced poly(ethyleneglycol) (PEG) on their surfaces [26]. Although PEG is most often used to coat the surfaces of various nanoparticles, other polymers also have excellent properties [55, 56].

Surface functionalization of nanoparticles can be covalent or noncovalent. As described in Section 2.4, covalent functionalization radically decreases the intrinsic optical properties of CNTs. Instead, amphiphilic polymers with a hydrophilic part exposed to solution and a hydrophobic part that interacts noncovalently with the nanotubes are typically used [7]. Phospholipids, the major components of cell membranes, are often used as an optimal hydrophobic component that strongly anchors onto CNTs and creates stable particles with low nonspecific adsorption to proteins [57–59]. Similar properties (i.e. reduction of nonspecific adsorption of proteins) have been observed for pluronic copolymers or PEG with the surfactant Tween, which shielded CNTs more efficiently than other coatings [60, 61]. Robinson et al. [62] introduced a large 90‐kDa amphiphilic poly(maleic anhydride‐alt‐1‐octadecene)‐methoxy‐block‐poly(ethylene glycol) coating to prepare CNTs with a longer blood circulation time than those with a phospholipid‐PEG (PL‐PEG) coating. Nevertheless, Liu et al. pointed out that a long circulation time leads to accumulation in the skin, underscoring the need to find an optimal balance between high uptake in the tumor and accumulation in skin [58].

The length of the PEG chain is an important parameter: longer PEG (5 kDa compared to 2 kDa) provides higher stability, lower nonspecific interaction with proteins, and lower reticuloendothelial system (RES) uptake, as well as longer circulation time and therefore higher tumor uptake upon modification with Arg‐Gly‐Asp (RGD)‐targeting peptide [59]. In addition to length, the branching structure of PEG chains can prolong circulation time. According to one study [57], increasing the PEG size to more than 7 kDa did not significantly improve the particle properties, but branching provided the particles with desirable characteristics for in vivo applications [57].

In addition, PEG‐coating methods need to be carefully optimized to ensure that CNTs possess the desired properties. Long periods of sonication create defects at the SWCNT sidewall, which leads to cutting at the defect site. The commonly used one‐hour sonication period shortens the CNTs [63] and degrades PEG into shorter fragments [64]. Welsher et al. proposed sonication of nanotubes in an excess of sodium cholate, followed by surfactant exchange for PL‐PEG, which protects the CNT fluorescence and targeting properties [63]. Zeineldin et al. showed that 2 kDa PL‐PEG is sufficient to decrease nonspecific binding and block the uptake into cells of nanotubes not fragmented by sonication for one hour. After subsequent modification with folic acid, nanotubes were specifically taken up only by cells expressing folic acid receptor [64].

GO needs to be modified with polymers primarily to reduce its cytotoxicity and increase biocompatibility and physiological stability. GO can be modified covalently by amidic coupling with six‐armed PEG, which leads to particles that are highly stable in medium or sera [65]. If GO is reduced to increase its fluorescence, it becomes water‐insoluble and has to be modified with an amphiphilic polymer. Similar to CNTs, reduced GO can be noncovalently modified with 5 kDa PEG with poly(maleic anhydride‐alt‐1‐octadecene), which ensures stability and prolongs circulation time [66]. Tan et al. focused on the behavior of nanoGO coated with 10 kDa amino‐terminated six‐arm‐branched PEG in serum and on reduction of protein adsorption onto the surface upon polymer modification. The polymer layer altered protein binding selectivity. Some of the proteins (immune‐related factors) adsorbed onto the surface at a higher rate, which could serve for eliminating immune factors evoked by other types of nanomaterials [67].

CDs are often covalently modified with PEG using amidic coupling [68]. Particles coated with PEG have lower in vitro and in vivo toxicity than uncoated particles [69]. In addition to increasing stability and biocompatibility, 5 kDa PEG prevents the creation of a protein corona, prevents interaction with opsonins, prolongs circulation time, and changes the biodistribution [70]. Such particles can be efficiently targeted by RGD peptide to tumor cells [71]. CDs can be created directly from the polymer by partial carbonization. CDs created from PEG do not need to be modified with a surface polymer layer, saving an extra functionalization step. These particles show good stability and prolonged circulation in blood, and after modification with a nuclear localization signal peptide, they can serve as nucleus‐imaging probes [72].

Nanodiamonds are considered the most biocompatible CNM. The purpose of polymer coating is therefore not to reduce their toxicity but to increase stability and reduce nonspecific interactions with proteins and cells to prolong circulation time and enhance the targeting effect. For example, detonation nanodiamonds (DNDs) can be covalently modified with Zonyl polymer (a combination of perfluoroalkyl chains and a PEG block). The Zonyl layer reduced adsorption of the protein bovine serum albumin (BSA) from 80% (of non‐modified particles) to 20% [73]. Covalent modification of high‐pressure high‐temperature (HPHT) nanodiamonds with 5 kDa PEG led to nanoparticles that were stable for at least two weeks, even in 1 M NaCl. These nanodiamonds were taken up by cells after a 24‐hour incubation. In contrast, nonmodified nanodiamonds aggregated during pre‐incubation in phosphate buffer saline (PBS) and stayed on the membrane. Nanodiamonds dispersed in serum with media before addition to cells were partially stabilized by proteins such as serum albumin and entered cells [74]. A PEG coating was used also as interface for nanodiamonds embedded in gold shells [75]. Wu et al. showed that coating HPHT nanodiamonds with PEG‐modified human serum albumin resulted in particles that were stable in 1 M NaCl with no change in size and that were internalized by cells [76].

However, many biomedical applications call for particles with proper anti‐fouling properties that do not enter cells at all. PEG‐coated nanoparticles bind various proteins, depending on the length and conformation of the PEG chain [67, 77, 78]. Moreover, PEG is an immunogenic polymer. PEG on GO stimulates potent cytokine responses [79] and induces deleterious effects and cell death [80]. Other possible surface modifications, such as proteins, polysaccharides, and various synthetic polymers, are becoming more popular. Proteins have turned out to be good stabilization agents, increasing the biocompatibility and stability of nanoparticles, but they usually do not reduce nonspecific interactions with other proteins nor internalization by cells. Engineered M13 bacteriophage modification of CNTs stabilized particles in both PBS and media with serum and enabled their further modification with an antibody against prostate‐specific membrane antigen [81]. Covering GO with complement factor H, an abundant protein in plasma, almost completely protects the particle from complement activation. Coating particles with serum albumin, which has a lesser shielding effect, is less protective, and coating with immunoglobulin G worsens complement activation [82]. Stabilization and enhanced intracellular uptake of serum albumin‐coated HPHT nanodiamonds has been demonstrated in vitro [83] and in vivo [84]. CNTs modified with polysaccharide dextran sulfate proved to be resistant to opsonins and bacterial cells. These nanotubes were modified by a site‐specific conjugation strategy (by oxime ligation) with antibodies that selectively target pathogenic bacteria [78]. GO with a cross‐linked chitosan/dextran coating proved to be stable in physiological conditions with low nonspecific interactions with human serum albumin [85]. Doxorubicin (DOX)‐loaded GO with noncovalently attached chitosan modified with cyclic RGD peptide specifically targets cancer cells [86]. Covalent amidic modification of GO with the polysaccharide hyaluronic acid leads to improved cytotoxicity and negligible hemolytic activity. A key advantage of hyaluronic acid as a polymer coating is its possibility to target cancer cells (with CD 44 determinants overexpressed on their surface). With this combined approach, there is no need to modify nanoparticles with other targeting moieties [87]. An even more straightforward approach is to use hyaluronic acid to prepare CDs, which are not fully carbonized and therefore are modified with hyaluronic acid as a shielding and targeting polymer on the surface [88]. Fructose‐based branched glycopolymer improved the stability of amino‐silica coated DNDs, targeted the particles to five glucose transporters [GLUT] that are expressed on breast cancer cells, and delivered DOX to selectively kill the cancer cells [89].

For exceptional results in targeted delivery and construction of selective biosensors, a perfect stealth coating is needed. All of the coatings described above were prepared by the “grafting to” method, in which polymers are prepared in advance in the solution and then noncovalently or covalently attached to the nanoparticle surface. Various synthetic polymers can polymerize in thicker and denser brushes directly on the surface of particles (“grafting from” approach) [90]. The “grafting from” approach cannot be recommended for CNTs because of their limitation in covalent modification. There are a few reports of using “grafting from” approaches to coat graphene and CDs, but the best‐studied material for preparation of dense brushes is nanodiamonds. GO modified with poly(acrylic acid) grown from graphene was found to have slightly better biocompatibility in vivo, less toxicity, reduced protein adsorption (especially the content of immunoglobulin G in the protein corona), and reduced membrane disruption, as well as a longer circulation time, than GO with 5 kDa PEG grafted to the surface [91]. CDs coated with polycation‐block‐polyzwitterion polymer by surface‐initiated ATRP have an excellent ability to suppress serum albumin adsorption. This type of polymer protects a DNA vector against nonspecific interactions with serum components and results in better transfection efficiency [92].

Hydrophilic, biocompatible hyperbranched polyglycerol grafted from the surface by anionic ROP is commonly used to prepare dense polymeric surfaces. CDs with such surfaces are water‐dispersible particles with low toxicity [93]. On nanodiamond surfaces, polyglycerol performs better than PEG, with lower nonspecific adsorption of lysozyme [94, 95]. These particles have excellent stability [96] and can be further modified with cyclic RGD peptide by azide‐alkyne cycloaddition, giving the particles reasonably high targeting efficiency to cancer cells [97, 98]. Exceptional properties, including no nonspecific interaction with proteins and very high particle stability (even in 1 M NaCl for at least two weeks), were achieved by preparing a dense poly[N‐(2‐hydroxypropyl)methacrylamide] layer on HPHT nanodiamonds. The shell was grown from silica with terminal methacrylate groups by radical polymerization (Figure 2.4a). Addition of azide‐ or alkyne‐ functionalized methacrylate monomer results in the possibility to further modify the nanodiamonds by azide‐alkyne cycloaddition catalyzed by Cu(I) ions [99, 100]. Such nanodiamonds with a dense polymer layer modified with cyclic RGD showed remarkably high specific interaction with integrin‐overexpressing cancer cells, thanks to the complete suppression of nonspecific interactions with the polymer shell (Figure 2.4b) [99]. A similar polymer shell was used to construct programmable quantum diamond nanosensors bearing densely loaded gadolinium(III) complexes [101].

2.6 Functionalization of CNMs with Stimuli‐Responsive Polymers

Stimuli‐responsive polymers respond to their environment by changing their physical and/or chemical properties. They have been designed to be responsive to a variety of stimuli, including pH, temperature, the wavelength or intensity of light, electric or magnetic fields, and the presence of various small molecules and biomolecules. These polymers can respond by various means, such as altering color, transparency, solubility, or shape. Typically, slight changes in the environment are sufficient to induce large changes in the properties of such polymers [102]. Here, we describe conjugates of CNMs with various stimuli‐responsive polymers.

2.6.2 pH‐Responsive Carbon Nanoparticles

Most pH‐responsive polymers are based on PAA or its analogues. Nanoparticles with a pH‐responsive graphene/PAA assembly formulated through noncovalent interactions exhibit reversible aggregation behavior, depending on the pH of the solution (Figure 2.5) [109].

Wang et al. prepared strongly fluorescent CDs with citric acid as a carbon source and hyperbranched polyethyleneimine (PEI) as a surface passivation agent. These particles showed good water solubility and low cell toxicity. PEI, a polycationic polymer, made the CDs exhibit pH‐responsive optical properties, and they were used as pH sensors due to their reversible pH‐responsive fluorescence [110].

Numerous pH‐responsive CNMs have been investigated as potential drug carriers capable of controlled active compound release, primarily for use in cancer therapy. Most of these drug‐delivery strategies take advantage of the difference between normal physiological pH and the much lower extra‐ and intracellular pH of solid tumors. Zhou et al. developed PEI‐treated GO particles coated with a pH‐responsive charge‐reversal polyelectrolyte and integrin monoclonal antibody for targeted delivery and controlled release of DOX into cancer cells [111]. Another strategy for DOX delivery made use of SWCNTs covered with different derivatives of the biopolymer chitosan and folic acid as a targeting moiety. While this system was stable at physiological pH 7.4 and 37 °C, most of the DOX was released at pH 5.5, which is close to the pH of the tumor microenvironment [112]. Similarly, highly fluorescent CDs rich in surface amino groups were coated with PEGylated oxidized alginate polymer and loaded with DOX via acid‐labile Schiff base linkage. The strong fluorescence of the nanoparticles could be utilized for cellular imaging [113]. Another drug delivery system based on CDs, delivering cisplatin(IV) prodrug to tumors, was developed by Feng et al. CDs were converted by dimethylmaleic acid (DMMA) and PEG functionalized poly(allylamine) into pH‐responsive charge‐convertible drug nanocarriers. The particles showed negative charge under normal physiological conditions and could be converted into positively charged species in the tumor extracellular microenvironment [114].

2.7 Functionalization of CNMs with Polymers for Delivery of Nucleic Acids

The development of nonviral, nonimmunogenic, and biocompatible vectors for efficient intracellular transfection of nucleic acids (NAs) is one of the challenges facing future gene therapy approaches [125]. Nonviral nanomaterial‐based gene delivery systems must fulfill the following requirements: (i) cell membrane penetration; (ii) endosomal escape; (iii) release of the NA cargo; (iv) successful gene expression [plasmid DNA (pDNA) delivery] or gene silencing [silencing structure delivery – short interfering RNA (siRNA), micro RNA (miRNA)] related to the effector molecule; and (v) high specificity and no off‐target effects associated with RNA interference (RNAi) [125, 126]. Polymer‐coated CNMs have proven to be one of the most promising nonviral platforms for delivery of genetic materials because of their ability to protect NAs against enzymatic cleavage, the possibility to attach accessory targeting molecules, their high binding capacity, and optical properties suitable for in vitro and in vivo tracking [127]. To construct delivery vectors, CNMs with positively charged surfaces are typically complexed with negatively charged NAs [128–131]. The surface of the CNMs can be functionalized by covalent or noncovalent attachment of macromolecular layer(s) to create reliable nonviral gene delivery vectors. These systems can subsequently be used for covalent or noncovalent bonding of NA molecules.

Liu et al. found that PEG‐coated CNTs are capable of siRNA delivery to afford efficient RNAi of CXCR4 and CD4 receptors on human T cells and peripheral blood mononuclear cells [131]. siRNA was incorporated onto the surface of CNTs via covalent disulfide bonds (Figure 2.6a), which are cleavable by thiol‐reducing enzymes aided by the acidic pH in lysosomes. A shorter PEG chain enhances the cellular uptake of SWCNTs. siRNA attached via a cleavable bond to the CNT exhibits superior transfection efficiency and silencing effects compared to conventional liposome‐based nonviral agents. The synergistic combination of RNAi with NIR photothermal therapy (using siRNA and a highly hydrophilic cationic PEI‐modified CNT complex; Figure 2.6b) greatly enhanced the therapeutic efficacy. siRNA internalization and transfection efficiency were significantly improved upon conjugating tumor‐targeting NGR (Cys‐Asn‐Gly‐Arg‐Cys‐) peptide with SWCNT/PEI/siRNA, which led to stronger suppression in proliferation of PC‐3 cells in vitro [132]. In tumor‐bearing mice, the SWCNT/PEI/siRNA/NGR delivery system exhibited higher antitumor activity due to greater accumulation in tumors without obvious toxicity in main organs. Moreover, there was a significant difference in transfection efficiencies in the order SWCNT/PEI/siRNA/NGR > SWCNT/PEI/siRNA > PEI/siRNA. More details on gene delivery using CNTs as a vehicle can be found in recent reviews [26, 136].

Nanodiamonds coated with polycationic polymers such as PEI and poly(allylamine) are a very promising platform as gene delivery vehicles. DNDs coated with low‐molecular‐weight PEI (0.8 kDa) (Figure 2.6c) exhibit higher transfection efficiency compared to amine‐terminated DNDs (DND‐NH₂) without PEI modification or PEI itself. The transfection efficiencies of pDNA varied as a function of the DND/pDNA weight ratio and decreased in the following order: DND/PEI (0.8 kDa)/pDNA > PEI/pDNA > DND‐NH₂/pDNA > DND/pDNA > pDNA. Even though both DND/PEI/pDNA and DND‐NH₂/pDNA complexes were successfully internalized, only DND/PEI/pDNA exhibited high transfection efficiency. This could be because DND‐NH₂/pDNA remained imprisoned in the endosomal space [133]. A similar approach was used for HPHT NDs, which were tested in a Ewing sarcoma cell model [137]. In contrast, a recently published study focusing on plasma hydrogenated cationic DNDs (DND‐H) with a primary size of 7 nm showed that cationic DNDs without a macromolecular coating conjugated with siRNA can potently inhibit EWS/FLI‐1 gene expression in a Ewing sarcoma cell line [138]. Zhang et al. demonstrated that a cationic nanodiamond‐polymer brush, synthesized by ATRP of 2‐(dimethylamino)ethyl methacrylate (DMAEMA) (Figure 2.6d), provides higher gene expression than PEI (25 kDa)/pDNA with lower cytotoxicity [134].

The abundance of carboxylic groups on GO enables simple functionalization with common polycationic polymers such as branched PEI (1.2 and 10 kDa) via electrostatic interactions [129, 139, 140]. Positively charged PEI‐coated GO/PEI complexes (Figure 2.6e) can bind and protect pDNA and also provide a higher gene expression level and lower cytotoxicity compared to PEI (10 kDa) [129]. Covalent conjugation of GO with branched PEG (10 kDa) and branched PEI (25 kDa) in a ~1 : 1 : 5 GO:PEG:PEI weight ratio helped avoid aggregation in the presence of sera (tested sera concentrations: 0%, 10%, 20%, 30%). This GO/poly(ethyleneglycol)‐polyethyleneimine (PEG‐PEI) complex was successfully used for pDNA and siRNA delivery. Compared with PEI and GO/PEI, GO/PEG‐PEI exhibits higher transfection efficiency and almost no interference from sera. Moreover, the strong absorption of GO in the NIR region allows synergistic increase in the delivery of siRNA under NIR laser irradiation at low power density, owing to photothermally induced local heating, which facilitates intracellular trafficking [139]. An extensive review of the application of graphene‐based nanosheets to gene delivery was recently published [141].

Fluorescent PEI‐coated CDs (carbon dot‐polyethyleneimine (CD‐PEI)) can be used as vectors for NA delivery. Positively charged CD‐PEI, synthesized via one‐step microwave pyrolysis of glycerol and branched PEI (25 kDa), allows pDNA binding via electrostatic interactions. The pDNA condensation capability and cytotoxicity of CD‐PEI/pDNA is a function of pyrolysis time, possibly due to the destruction of PEI during the formation of CDs. CD‐PEI/pDNA (Figure 2.6f) are able to mediate gene transfection with higher or comparable efficiency, and with lower cytotoxicity, than PEI alone (25 kDa). Furthermore, for cell transfection in serum‐containing media, some CD‐PEI/pDNA particles can achieve a gene expression level superior that obtained with PEI (25 kDa), although it is slightly lower than in the serum‐free case (tested sera concentrations: 0%, 10%). The proper pyrolysis time is the key to balance the transfection efficiency and cytotoxicity of CD‐PEI/pDNA [135]. The block polymer poly‐[DMAEMA]‐block‐poly[N‐(3‐(methacryloylamino)propyl)‐N,N‐dimethyl‐N‐(3‐sulfopropyl)ammonium hydroxide], covalently grafted from CDs by surface‐initiated ATRP, can serve as a condensing agent for pDNA (Figure 2.6g) [92]. Particles modified with this polycation/polyzwitterion copolymer showed less cytotoxicity and enhanced serum stability and cellular internalization compared with particles modified with non‐zwitterionic polymers.

In general, carbon nanomaterial‐polyethyleneimine (CNM‐PEI) complexes have better transfection efficiency than PEI alone (25 kDa) in serum‐containing media. The transfection efficiency of PEI (25 kDa) declines dramatically with increasing serum concentration (tested sera concentrations: 10%, 30%, 50%) [92]. Even though CNM‐based carriers do not offer as high transfection efficiency as viral‐based vectors, they are able to maintain high cellular uptake, reduce adverse effects, and specifically target cells. These results highlight the importance of polymer coating of CNMs for gene delivery applications.

2.8 Outlook

In this chapter, we have provided an overview of how to use polymer coating to control the interface between CNMs and biosystems. Different CNMs require different biointerfacing approaches to make them compatible with biological environments for use in biomedical applications. For example, noncovalent polymer coating is preferable for CNTs because it avoids the risk of losing the NIR‐II emission during covalent modification. A graphene surface can be functionalized with polymers by both covalent and noncovalent approaches; however, noncovalent functionalization preserves the extended π‐conjugation of the graphene sheet, while covalent functionalization creates sp³ defects in the graphene ring. On the other hand, nanodiamonds can be conveniently functionalized by both noncovalent and covalent surface treatments without affecting the intrinsic optical properties of the NV centers buried inside the diamond core. Although CD surfaces are rich in ─OH and ─COOH functional groups, their surface can be further functionalized with PEG or other polymers. The use of polymers as coatings for CNMs is now well‐established, and advances in synthetic and biointerfacing techniques have led to a wealth of literature describing in vitro and in vivo investigations of CNMs.

In addition to polymer coatings, another commonly used biointerfacing approach involves silica coating of CNMs [53, 90, 142]. Silica coating has been studied as a means to enhance colloidal stability of CNMs due to silica's high stability, low cost, chemical inertness, processability, and optical transparency. A silica layer can confer on CNM cores both steric and electrostatic protection. Another popular biointerfacing approach is micellar solubilization through noncovalent wrapping of CNMs with a surfactant [143–145]. Although micelle‐stabilized CNMs can be stable in aqueous solution for several months, the difficulty in purifying these particles and removing excess, usually toxic surfactant molecules raises biocompatibility concerns.

As illustrated in this chapter, intense efforts and interest have been invested by many laboratories to functionalize the CNMs and explore their potential application in medicine. However, it is important to recognize that translation of CNMs from an interesting nanomaterial to an effective biomedical product is still at the nascent stages. Most conclusions on biomedical applications of CNMs have been demonstrated in vitro and in vivo. It remains to be seen whether such strategies will be clinically realistic and efficacious by minimizing adverse reactions. The clinical translation can be impeded by fundamental limitations of human physiology (i.e. vessel pore size, renal and hepatic clearance, RES), potential toxicity, and/or interference with other medical tests [146]. To increase the likelihood of clinical translation of CNMs, at least the following three criteria [146] should be considered. (i) CNMs should composed of nontoxic materials and/or biodegradable to clearable (renal or hepatic) components. (ii) Surface coatings of the CNMs should minimize the nonspecific tissue/organ uptake, and binding to serum proteins. (iii) Hydrodynamic diameter of CNMs or degradation products should be ≤5.5 nm to facilitate complete renal elimination.

Polymer coatings can significantly increase the hydrodynamic diameter of CNMs. The large size of a polymer coating can make CNMs incompatible with biosensor applications, and targeting smaller organelles becomes difficult, as does renal excretion. The size of polymer‐coated CNMs (nanodiamonds, CNT, graphene) typically exceeds the renal filtration cut‐off size (≤5.5 nm) [147], which increases their chance of becoming entrapped inside the liver and spleen and raises long‐term toxicity concerns. In this regard, due to their compact size, single‐digit nanodiamonds, and CDs can afford favorable renal clearance. Therefore, for clinical applications, small‐sized CNMs (nanodiamonds, CNT, graphene) with a biodegradable, thin, compact polymer coating are desirable.

Another important issue is lack of commonly followed standard protocols for synthesis, surface characterization, and toxicity testing of CNMs, which lead to inconsistent results and hinder clinical translation. The broad distribution in diameter, impurities (mainly including amorphous carbon and catalyst particles) also affects the reproducibility. In addition, each new functionality elevates the complexity (e.g. multistep syntheses, purification, and characterization), cost, and regulatory barriers (e.g. owing to multicomponent, heterogeneous formulations) [148]. Therefore, it is recommended that the FDA's regulation be followed regarding nanotechnology products development and characterization and that the FDA be consulted early in the development process to facilitate a mutual understanding of the scientific and regulatory issues for the nanotechnology products (https://www.fda.gov/ScienceResearch/SpecialTopics/Nanotechnology/ucm301114.htm ). Hopefully, a coordinated effort involving all parties, including chemists, pharmacologists, toxicologists, clinicians, pharmaceutical companies, and regulatory authorities, will push the CNM research activity to achieve the ultimate aim of clinical translation to help patients.

Acknowledgments

The work of J.N. was supported by the Grant Agency of the Czech Republic (Project Number 18‐17071S). The work of P.C. was supported by European Regional Development Fund; OP RDE; Project: “Chemical biology for drugging undruggable targets (ChemBioDrug)” (Grant No. CZ.02.1.01/0.0/0.0/16_019/0000729).

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