Biomedical applications of magnetite nanoparticles

CHAPTER

Biomedical applications of magnetite nanoparticles

13

Greter Ortega1,2 and Edilso Reguera1 1

Center for Applied Science and Advanced Technology, Legaria Unit, of the National Polytechnic Institute, Del. Miguel Hidalgo, Mexico 2Faculty of Chemistry, Department of General and Inorganic Chemistry, University of Havana, Havana, Cuba

13.1 INTRODUCTION Recently, great attention has been devoted to the synthesis of various magnetic nanoparticles (MNPs) due to their extensive applications in biomedicine area. The main representatives of the MNPs are iron-oxide nanoparticles (IONPs), mainly magnetite (Fe3O4) and maghemite (γ-Fe2O3). In this chapter, magnetite nanoparticles will be referred to as IONPs. The synthesis conditions and surface functionalization of IONPs are critical to leading their physicochemical properties, as well as their colloidal stability, and biological behavior/fate. Since the biomedical applications of nanoparticles are mainly dependent on their size, IONPs have been classified as follows: (1) micrometer-sized paramagnetic iron oxide (MPIO; several micrometers), (2) superparamagnetic iron oxide (SPIO; hundreds of nanometers), and (3) ultrasmall superparamagnetic iron oxide (USPIO; less than 50 nm) (Bulte and Kraitchman, 2004). In addition, IONPs must combine very small size, narrow size distribution, and optimal surface coating with optimal magnetic susceptibility and negligible remanent magnetization. IONPs have received significant attention recently in the medical and pharmaceutical fields such as for drug delivery, hyperthermia, magnetic resonance imaging (MRI), tissue engineering, biosensors and bioanalysis, and biomolecule separations due to their biocompatibility and biodegradability. Also, the introduction of “theranostics,” allowing simultaneous drug delivery and imaging, represents an important nanotechnology contribution.

13.2 SYNTHESIS AND STABILIZATION OF MAGNETITE NANOPARTICLES FOR BIOMEDICAL APPLICATIONS The synthesis procedures for the obtention of magnetite nanoparticles are classified as follows: (1) physical methods, (e.g., gas-phase deposition and electron Materials for Biomedical Engineering: Nanomaterials-based Drug Delivery. DOI: https://doi.org/10.1016/B978-0-12-816913-1.00013-1 © 2019 Elsevier Inc. All rights reserved.

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beam lithography); (2) wet chemical methods (e.g., chemical coprecipitation, hydrothermal, thermal decomposition and sol 2 gel reactions, synthesis inside nanoreactors, electrochemical pathways, aerosol/vapor-phase method, sonochemical decomposition reactions, supercritical fluid method, and so on); and (3) microbial methods. Some technical and experimental key aspects of the most widely used procedures to synthesize Fe3O4 for medical applications are briefly described in Fig. 13.1A (Laurent et al., 2008). The coprecipitation technique introduced by Massart (1981) is probably the simplest and most efficient chemical procedure to obtain a large amount of USPIO nanoparticles by using a stoichiometric ratio of 2:1 (Fe31/Fe21) in a nonoxidizing (free of oxygen) environment and basic media. However, the traditional coprecipitation reaction results in the formation of polydisperse and poorly crystalline nanoparticles with moderate magnetic properties declining the biomedical application. On the other hand, hydrothermal synthesis of magnetite nanoparticles is another very common method reported in the literature (Wan et al., 2005). These reactions are carried out in reactors or autoclaves in aqueous media and pressures and temperatures around 2000 psi and 200
C, respectively. Despite the fact that this method is widely used to obtain USPIOs with monodisperse and highly crystalline nanoparticles for biomedical application, this process must be amended for suitable industrial performance, especially in the safety of the reactant and the necessary high pressures. USPIO nanoparticles can also be produced via decomposition of metal precursors such as organometallic complexes under controlled heat (Zhao et al., 2012). This is one of the most effective strategies to obtain magnetite nanoparticles for further biomedical applications, mainly as an RMI contrast agent. On the other hand, the sol 2 gel method is based on the hydrolysis of precursors such as iron alkoxide solutions to yield a sol of nanoparticles. After that, the condensation and polymerization processes induce the formation of a gel by networking a three-dimensional iron-oxide structure. Because these reactions are carried out at room temperature, further heat treatments are needed to achieve the final crystalline state of SPIO and USPIO nanoparticles (Lemine et al., 2012). The achievements of this method rely on a good control of the predetermined structure of the particle, size monodispersity, as well as the opportunity to insert molecules, that endure their chemical-physical properties within the solgel matrix. Also, the polyol process, which can also be understood as a solgel method, offers an alternative instead of traditional surfactants (Fig. 13.1B). The use of the polyols [e.g., ethylene glycols such as mono-, di-, tri-, and poly(ethylene glycol), propylene glycol, etc.] draw upon their usefulness as a solvent, reducing agent, and capping agent, allowing well-defined shapes and sizes by controlling the particle growth, ensuring an appropriate degree of crystallinity of the NPs, and preventing interparticle interaction with minimal aggregation (Cai and Wan, 2007).

FIGURE 13.1 (A) Schematic representation of the most widely used procedures to synthesize Fe3O4 nanoparticles for medical applications. (B) Example of transmission electron micrographof magnetite nanoparticles synthesized through the polyol method. (C) Schematic representation of magnetite surface chemistry and the different capping stabilizers.

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Finally, the synthesis of USPIO using the confined space conditions resulting in restricted “nano-reactors” in which the particle growth can be perfectly controlled have been developed (Laurent et al., 2008). Another important aspect is the stabilization of IONPs for biomedical applications to avoid aggregation in a biological medium and under a magnetic field. The stability of such a colloidal suspension results from the competition between attractive and repulsive forces which rely on the nanoparticle surface features, as well as the magnetic performances. Four classes of forces mainly contribute to the interparticle interactions: van der Waals (attractive), electrostatic interactions (attractive and repulsive), steric repulsion forces, and magnetic interacting forces: ferromagnetic, dipolar, and RKKY (Ruderman, Kittel, Kasuya, and Yosida) interactions. Controlling the strengths of these forces is the key to obtaining particles with suitable colloidal stabilities. In aqueous systems, IONPs perform as Lewis acids and adsorb and/or coordinate water molecules or hydroxyl groups (Begin-Colin and Felder-Flesch, 2012). The hydroxyl groups at the surface are amphoteric and they may be singly, doubly, and triply coordinated to Fe atoms. Around the IONP isoelectric point (pH 5 6.8), the surface charge is not sufficient and the particles are not stable enough in water and flocculate. For this reason, the surface covering improves the colloidal and physical stability of the particles, increasing their water-dispersibility and providing suitable functionalization for further bioconjugation with targeting ligands. The capping agents that provide the colloidal stabilization can be classified as monomeric, polymeric, inorganic, and liposomal stabilizers (Fig. 13.1C). The monomeric stabilizers include functional groups such as carboxylates (Tomba´cza et al., 2013), phosphates (Daou et al., 2007), etc. (Cornell and Schwertmann, 2003). The nanoparticles can be functionalized during or after (through ligand exchange) the synthesis procedures and the ionic group that remains exposed to the solvent is responsible for providing charge and hydrophilic surface. However, for many applications in medicine, polymer-coated IONPs are preferred over functionalization with small organic compounds. The adsorption of polymers onto nanoparticles via two alternative approaches: grafting “onto” and “from” (Basuki et al., 2013) confers protective steric repulsion (Khoee and Hemati, 2013). In addition, stabilizations using amphiphilic copolymers or charged polymer chains are enhanced. In the last case, extra electrostatic repulsion is provided, combining an ionic and steric (electrosteric) stabilization effect. For biomedical applications, some properties of the polymer must be taken into consideration. In this sense, the length and molecular weight, chemical structure, hydrophilic feature, steric conformation, surface coverage packing, and binding strength to the particle surface lead to the pharmacokinetics, biodistribution, and biological fate of the nanoparticles. In the literature, the most common capping agents are dextran and its derivatives, polyethylene glycol (PEG), starch, arabinogalactan, glycosaminoglycan, sulfonated styrene-divinylbenzene, polyvinyl alcohol (PVA), polyvinylpyrrolidone

13.3 Bioconjugation of Magnetite Nanoparticles

(PVP), poloxamers, polyoxamines, polydopamine, chitosan, polyethylenimine (PEI), and many others (Boyer et al., 2010). Among the above-mentioned polymers, PEG and dextran are undoubtedly the most widely used polymers for IONP coating for nanomedicine applications. PEG and its carboxylic derivates present excellent solubility and stability in aqueous solutions, as well as in physiological media. The PEG-coated IONPs are barely recognized by the macrophage system, increasing their plasma half-life due to their notable steric repulsive and stealth properties (Ruiz et al., 2013). Also, a fashionable tendency is the molecular imprinting technology that relies on the formation of specific sites in a polymer matrix starting from the memory of a template. Therefore, the combining of IONPs with molecular imprinting technology results in a powerful analytical tool toward selective separation of a large variety of target molecules assisted by magnetic handling (Sierra-Martin and Fernandez-Barbero, 2016). On the other hand, magnetite nanoparticles can be coated with inorganic shells, mainly silica (Ciriminna et al., 2013), gold (Silva et al., 2016), or gadolinium (III) complex (Bae et al., 2010) as stabilizers and binding ligands. Finally, liposomes (phospholipid bilayer vesicles) have been used to encapsulate a large amount of IONPs achieving liposomal magnetic systems known as magnetoliposomes. The magnetoliposomes are considered as encouraging candidates for various biomedical applications, such as MRI, targeted drug and gene delivery, tumor hyperthermia, and magnetic cell separation (Han and Zhou, 2016). These systems have also served as mimetics of virus-like particles with high magnetic susceptibility by using them as a scaffold for the self-assembly of a viral protein cage.

13.3 BIOCONJUGATION OF MAGNETITE NANOPARTICLES One of the major concerns regarding the use of magnetite nanoparticles for bioapplications is the conjugation of biomolecules such as enzymes, antibodies (Abs), or their fragments, aptamers, oligosaccharides, proteins, peptides, peptidomimetics, and small targeting ligands without compromising their functionality once attached. In this sense, a considerable number of chemical approaches have been used (Sapsford et al., 2013). The success of a conjugation strategy is determined for several parameters that the conjugated nanoparticles must guarantee: (1) colloidal stability in physiological media, (2) suitable magnetic features, (3) biocompatible and nontoxic fate, and (4) specific recognition of drug and enzymatic activity according to the case. In the literature, there are many reviews and papers summarizing and classifying the different conjugation strategies. In this section, two major groups are distinct: one in which biomolecules are linked to nanoparticles without their previous derivatization and the other when earlier treatment and functionalization of biomolecules are required. Also, for both cases, there are

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three main strategies: physical association, covalent binding, and using adapter biomolecules. Tables 13.113.3 summarizes the principle examples for these. Physical adsorption (See Table 13.1) is generally based on hydrophobic, electrostatic, hydrogen bonding, and van der Waals attractive forces between the Table 13.1 Different Strategies for the Conjugation of Biomolecules on Magnetite Nanoparticles by using Physical Adsorption

13.3 Bioconjugation of Magnetite Nanoparticles

Table 13.2 Different Strategies for the Conjugation of Biomolecules on Magnetite Nanoparticles by Covalent Binding

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Table 13.3 Different Strategies for the Conjugation of Biomolecules on Magnetite Nanoparticles by using Adapter Biomolecules

pristine biomolecules and the surface of the NPs (Puertas et al., 2011). However, despite the fact that physical adsorption is the simplest and most straightforward immobilization process, it often suffers from denaturation of biomolecules, yielding poor reproducibility, scale-up of the process, and stability in biological media. On the other hand, covalent binding provides a number of distinct advantages, such as higher stability of the bioconjugate and better reproducibility than physical adsorption. For covalent linking, depending on the coupling reaction selected, the binding can occur with or without chemical modification of the biomolecules (Montenegro et al., 2013), by almost chemistry a` la carte (see Table 13.2). Generally speaking, the main disadvantage of some of these procedures is the usually poor control of oriented attachment of biomolecules. For example, when the antigen-binding sites of antibodies are sterically blocked upon conjugation, the bioactivity, avidity, and targeting efficiency of the functionalized NPs are diminished. Oriented immobilization can also be accomplished using adapter biomolecules previously linked to the IONP surface in which several options are described in the literature as follows (See Table 13.3): (1) biotin-binding proteins (avidin, streptavidin), (2) Abs or proteins that specifically recognize the Fc portion of a

13.4 Physical-Chemical-Biological Features of Magnetite Nanoparticles

secondary Abs, (3) nucleic acid-mediated hybridization, and (4) engineered Abs with terminal linkerpeptide residues. Each of these alternatives has its own advantages and disadvantages and, in general, the binding is much stronger than that achieved by physical adsorption.

13.4 PHYSICAL-CHEMICAL-BIOLOGICAL FEATURES OF MAGNETITE NANOPARTICLES 13.4.1 PHYSICAL FEATURES OF MAGNETITE NANOPARTICLES LEADING TO BIOMEDICAL APPLICATIONS Once IONPs are synthesized and functionalized the physical features such as crystalline phase, size, polydispersity, shape, capping efficiency, shell thickness, colloidal stability surface area, and charge define their chemical and physical properties as well as their biological fate (Ne´methova´ et al., 2017) and, therefore, the possible application. By far, magnetite is the most used IONP for biomedical applications in which magnetism is the lead property. In iron oxides, unpaired electrons of, for example, orbitals of Fe31, couple magnetically with the electrons of the p orbitals of O22 through the super-exchange process. For the particular case of bulk magnetite, the structure of which is type inverse spinels, below Tc (Curie temperature) the Fe31 ions are divided equally between tetrahedrally coordinated (A) sites, and octahedrally coordinated (B) sites, whereas Fe21 occupy B sites. As a result, the spin moments of all the Fe31 ions on B sites are aligned parallel to one another but directed opposite to the spin moments of the Fe31 ions occupying the A positions. Therefore, the magnetic moments of all Fe31 ions cancel, making no net contribution to the magnetization of the solid. However, all the Fe21 ions have their moments aligned parallel to one another, which is responsible for the net ferrimagnetic magnetization. However, because the particle sizes of suitable magnetite colloids for most biomedical application (SPIO and USPIO) are much smaller than the size of one magnetic domain, they behave as nanomagnets made of single domains fully magnetized and characterized by the anisotropic energy. The magnetic energy of a nanomagnet hinges on the direction of the magnetization vector regarding the crystallographic directions, and it is minimal for anisotropy directions or easy axes. The anisotropic energy is directly proportional to the crystal volume and enriched with the crystal radius increases. There are four contributions to the anisotropy field: (1) the bulk magnetocrystalline anisotropy field, which depends upon the chemical composition and the crystallographic structure of the material; (2) the demagnetizing field, which is determined by the shape of the crystal; (3) the anisotropy constant, which also depends upon the

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surface structure of the crystal; and (4) the dipolar interaction between neighboring crystals of agglomerated structures. Another important parameter of the magnetic behavior of a single nanodomain particle is the Ne´el relaxation time (τN). The Ne´el relaxation time depends on the anisotropy energy and is described as a function characterized by the time constant of the return to equilibrium of the magnetization after a perturbation. Said differently, τN defines the fluctuations that arise from the jumps of the magnetic moment between different easy directions. In addition, the return of the magnetization to equilibrium for colloidal systems is lead by the contributions of Ne´el and the Brownian relaxation, which, in the last one, the viscous rotation of the entire particle is taken into account. For SPIO and USPIO particles, τB is shorter than τN, therefore the net relaxation is dominated by viscous rotation of the particle. In these conditions, the magnetization curve is reversible due to the rapid magnetic relaxation allowing the system to be at thermodynamic equilibrium. This behavior has been named “superparamagnetism.”

13.4.1.1 Basic principles of magnetite nanoparticles as MRI contrast agents and hyperthermia Magnetic resonance imaging (MRI) is one of the most powerful diagnostic tools in clinical practice. "Contrast" refers to the signal differences between adjacent ill-health regions as a result of the interaction between the contrast agents and neighboring water protons. This technique is based on an external magnetic field applied over a biological system with a radiofrequency (RF) pulse at an exactly Larmor frequency, in which the hydrogen nuclei of water are excited and align antiparallel to the external magnetic field. As they relax to the aligned state, the nuclei emit the energy they absorbed. There are two different relaxation pathways. The first, called longitudinal or T1 relaxation, involves the decreased net magnetization (Mz) due to energy transfer, whereas the second, called transverse or T2 relaxation, involves induced magnetization on the perpendicular plane (Mxy) vanishing by the dephasing of the spins. The local environment around the proton nuclei will alter their relaxation time. Therefore, transverse relaxation is affected by inhomogeneous magnetic fields produced by tissue-inherent factors or external sources. As was explained above, SPIOs and USPIOs exhibit strong magnetization under the external magnetic field, which leads to microscopic field inhomogeneity, and the increase in the relaxation rate of the dephasing of protons due to the dipolar interaction between proton spins and the nanoparticle magnetic moment. In other words, the high magnetic susceptibility of magnetic cores shortens the transverse proton relaxation time at the tissue regions where the nanoparticles are localized and generate a darker MR image of the targeted area in contrast to the biological background (Na et al., 2009). The efficiency of the contrast agents is evaluated in terms of the relaxivity (r1 or r2) according to the following equation: RiðobsÞ 5 T1o 1 ri C; where i 5 1; 2 i (for T1 and T2 contrast agents, respectively); RiðobsÞ is the observed relaxion rate (s21); Tio is the relaxation time before adding contrast agents (s); ri is the

13.4 Physical-Chemical-Biological Features of Magnetite Nanoparticles

relaxivity coefficient (mM21 s21) usually determined by calculating the slope of a plot of 1/T1 and 1/T2 versus the concentration of the contrast agents; and C (mM) is the concentration of contrast agents. Because the T2 contrast effect is highly dependent on the magnetization of a particle, the r2 values can be increased by enhancing the magnetic moment of a nanoparticle. For these reasons, the control of magnetic properties of the nanoparticles by controlling (1) the intrinsic material properties, such as material composition and crystal structure, and (2) extrinsic factors, such as size, shape, and organic capping, are extremely necessary to obtain efficient T2 contrast agents. Monodisperse and highly crystalline magnetite nanoparticles, synthesized by thermal decomposition, are usually used. Also, as was explained in Section 13.4.1, the magnetic spins of only Fe21 ions in the octahedral sites of magnetite contribute to the net magnetic moment. Replacing Fe21 ions with other divalent transition metal ions (e.g., Mn21, Co21, Ni21, and Zn21 ions) is a widely used option to control the magnetic properties of such nanoparticles. For example, the r2 value of Fe3O4 nanoparticles is 218 mM21 s21, whereas for analog nanoparticles of MnFe2O4 it is 358 mM21 s21 (Shokrollahi, 2013). Because the magnetization of nanoparticles is strengthened as the size increases, larger nanoparticles exhibit higher r2 relaxivity until a plateau of the maximum r2 for larger nanoparticles, which exhibit ferromagnetic behavior. The aggregation by the remanent magnetization of ferrimagnetic nanoparticles hinders their biomedical applications, but some studies have tried to control such an aggregation process (Lee et al., 2012). For a clustering system, each one behaves as a large magnetized sphere, and its overall magnetic moment is proportional to the cluster size. In this sense, the synthesis of water-dispersible clusters of iron-oxide nanoparticles has been actively tracked by using crosslinked iron-oxide (CLIO) nanoparticles, dextran coatings, amphiphilic block copolymers, silica-coated clusters, etc. (Laurent et al., 2008). On the other hand, the surroundings of superparamagnetic nanoparticles exposed to an RF alternating current (AC) magnetic field are heated due to the Ne´el and Brownian relaxation processes. The Ne´el mechanism addresses internal heating generated by internal friction between the crystal lattice and the rotating magnetic spins, in which the dipole overcomes an energy barrier altering the direction. The Brownian mechanism generates and releases mechanical heat from friction between the rotating nanoparticles and solution. Ne´el heating is dominant at smaller nanoparticle sizes (below 1516 nm), whereas at larger diameters, Brownian relaxation, also characterized by the medium viscosity, prevails. The heating performances are modulated by nanoparticle size, crystal structure, polydispersity, shape, and magnetocrystalline anisotropy. For a large size distribution in polydispersive samples, the heating process is unfavorably due to combined heating mechanisms.

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13.4.2 PHARMACOKINETICS, BIODISTRIBUTION, AND BIOLOGICAL FATE OF IONPS USED IN BIOMEDICAL APPLICATIONS For many applications, the IONPs are intravenously injected. In this sense, nanoparticles stabilized with such polymers as PVP, chitosan, dextran, or PEG, as well as vectorized nanoparticles (that mean conjugated with biomolecules able to recognize a biological target) may be used. After the nanoparticles are injected, they can be transported through the blood vessels depending on the hydrodynamic forces. At the same time, the IONPs may be or not be magnetically guided to the targeted tissue, aimed for the targeting, imaging, or therapeutic bonded moieties. The Food and Drug Administration (FDA) established that adult patients exposed under 8 T of induced fields lack significant physiological risk. Immediately after injection, a process known as opsonization in which the plasma proteins are adsorbed onto IONPs occurs, depending on the attractive and repulsive forces discussed above. The opsonized nanoparticles are rapidly captured by the macrophages of the liver (i.e., Ku¨pffer cells), of the spleen, and of the bone marrow, which are tissues belonging to the reticuloendothelial system (RES). The adsorption of plasma proteins onto the IONP surface depends on the size, hydrophilicity, shape, and surface charge of the particles (Reddy et al., 2012). Generally, smaller nanoparticles adsorb a lower amount of proteins and enable an increase in the plasma half-life of the IONPs. In addition, IONPs with size smaller than 10 nm are rapidly removed by the renal clearance process, while IONPs with diameters around 50 2 100 nm accumulate in the hepatic parenchyma, whereas particles with sizes higher than 200 nm result in spleen uptake, and the bigger particles (more than 4 μm diameter) are mainly captured and retained in the lungs, bringing about the risk of embolism. At the cellular level, the mechanism of penetration of the opsonized NPs is dramatically influenced by their size. Hence, larger IONPs are only captured by the cells capable of phagocytosis, whereas smaller nanoparticles than 150 nm could have cellular access through pinocytosis, but the diversity of entry mechanisms is quite variable. Following cellular uptake, the IONPs are clustered within the lysosomes with posterior degradation via a variety of hydrolytic enzymes under low pH. On the other hand, the opsonization occurs more rapidly with the hydrophobic particles than with the hydrophilic ones. There is some controversy regarding the optimum geometry for the oncoming of the nanoparticles to the cell membrane and the entire internalization (Delgado et al., 2014). Some authors report that spherical shapes are rapidly captured because of their symmetry and others report the best results for other geometries. Also, the surface charge strongly influences the opsonization process. Neutral charge interacts minimally with the plasma proteins, extending the circulation time of the IONPs, whereas a high surface charge (positive or negative) enhances the phagocytosis process. Also, contrary to the expected behavior of repulsion among anionic IONPs and negative cell membranes, the nanoparticles are able to interact with the cells by adsorptive

13.4 Physical-Chemical-Biological Features of Magnetite Nanoparticles

endocytosis with unfavorable implications (Wilhelm et al., 2002). Strongly negative nanoparticles surface increased their hepatic uptake, whereas positively charged particles tend to nonspecifically stick to the cells. In order to distribute the IONPs on the tissues other than the RES organs, the nanoparticle surfaces have been provided with shielding capping (also known as a “stealth” property). At present, PEG is the most widely used hydrophilic polymer with a stealth property (Jenkins et al., 2016) retarding IONP opsonization by the steric hindrance effect, which depends on the chain length and the packing density of PEG on the surface. Finally, the administration route may also influence IONP biodistribution. For example, when injected locally (e.g., subcutaneously or intratumorally) at the diseased site, small magnetic particles undergo passive infiltration and are gradually absorbed by the lymphatic capillary system. In addition, the nanoparticles administrared by the inhalation route cross the pulmonary epithelium or the blood 2 brain and blood 2 testis endothelium, whereas the intraperitoneally injected ones are accumulated in liver and spleen, followed by kidneys, heart, testes, and uterus, while very low accumulation was observed in lungs (Kwon et al., 2008).

13.4.2.1 Toxicity and Biocompatibility The in vitro (in various cell lines) and in vivo (in different animal species) toxicity measurements of IONPs may reveal preliminary and essential information concerning the safety of these nanoparticles (Markides et al., 2012; Szalay et al., 2011). However, the toxicity of IONPs may depend on numerous factors including: (1) nanoparticle physical-chemical features such as chemical composition, size, structure, solubility, surface chemistry; (2) nanoparticle biological behavior like biodegradability, pharmacokinetics, and biodistribution; and (3) administration protocol concerning the dose and route. In addition, when IONPs are used as nanocarriers, the toxicity profile of the drug itself may be changed as a consequence of a modification of its cell/tissue biodistribution and clearance/metabolization. For these reasons, an in-depth evaluation of the cytotoxicity is therefore required for each case regarding the proposed application. Furthermore, comparisons of toxicities between IONPs with different physical and chemical features can be nonrigorous. Incubation of the IONPs with cells may alter the nature of cell adhesion, which may further affect their morphology, cytoskeleton, proliferation, differentiation, migration, and survival. Generalizing carefully, in in vitro experiments, IONPs with a hydrodynamic diameter smaller than 30 nm and relatively high doses, show higher toxic effects than larger particles (up to around 500 nm) (Karlsson et al., 2009). Despite the fact that the surface-coated IONPs perform, in general, a slight interaction with the components of the cell culture media, the NPcell interaction changes depending on the capping stabilizers. In these senses, for neutral ligands, PEG is the polymer with an improved stealth property and dextran-coated IONPs exhibit a variety of toxicities and cell interactions depending mainly on the dose and size (Ding et al., 2010). In general, positively

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charged IONPs present more interactions with plasma proteins through nonspecific binding (causing cell lysis) than the neutral and negatively charged IONPs. Thus, the surface coatings, even if providing IONPs with a similar surface charge, may importantly alter their cytotoxicity, depending on the nature of the surface coating in a dose-dependent manner (de Freitas et al., 2008). Notable is the fact that, in in vitro assays, the cytotoxicities are increased because the IONPs and their degradation products remain in continuing interaction with the cells in the culture. In contrast, in in vivo experiments, NPs are continuously eliminated from the body in case they are biodegradable. Moreover, once intravenously injected IONPs are taken up by the cells, they are metabolized in the lysosomal sacs at low pH by hydrolytic enzymes and proteins. The metabolized free iron is then incorporated into the body (B4000 mg of iron in a normal male adult) as hemoglobin, linked to transferrin, and then mostly eliminated via the fecal route. Depending on the dose administered, the IONPs show a satisfactory safety profile in different animal species (Bourrinet et al., 2006). Thus, higher doses result in an increased plasmatic iron concentration, which may lead to oxidative stress and various toxicities including hematological parameters, cardiovascular, renal, and hepatic toxicities. Also, the toxicity of IONPs is strongly dependent on the different routes of administration such as intravenous, intraperitoneal, subcutaneous, and oral feeding. However, the divergences on the sizes, capping agents, and doses of the administrated IONPs do not allow a regular behavior to be established. Generally, the IONPs could be considered as biocompatible. However, the safety of IONPs varies with their physicochemical properties and administration parameters, through the determination of the dosing range and schedule to avoid side effects and toxicity.

13.5 MAGNETITE NANOPARTICLES IN DISEASE THERAPY 13.5.1 DRUG DELIVERY The limitations of conventional drug-delivery systems, such as poor specificity to the site of action and reduced drug diffusion through biological barriers, could be potentially overcome by using IONPs with polymeric, lipidic, or inorganic coverings as drug-delivery carriers for systemic administration. In this sense, a rational design of the shell of the IONPs acquiescent to the physicochemical properties of the drug could retain the drug molecules in the carrier system until it reaches the target site to which it is released. Generally, depending on the therapeutic doses, affinities, release triggers, and so on, drugs can be loaded just by adsorption onto the functionalized NPs or incorporated into an organic/inorganic matrix built over the magnetic core, resulting in a pharmacokinetic/pharmacodynamic profile (Liu et al., 2017). For example, different mesoporous silica nanoparticles (MSNPs) are combined with IONPs in different ways, as follows (Colilla et al., 2013):

13.5 Magnetite Nanoparticles in Disease Therapy

IONP@mesoporous silica as coreshell structures, rattle-type IONP@hollow mesoporous silica spheres, IONPs encapsulated within mesoporous silica nanospheres, and mesoporous silica nanoparticles decorated with IONPs. On the other hand, engineered stabilized-vectorized IONPs with suitable physical-chemical properties and pharmacokinetics, biodistribution, and biological fate allow: (1) enhanced half-life and exposure of the NP-encapsulated drug; (2) magnetic guidance using an applied extracorporeal magnetic field or implanted magnet (Hournkumnuard and Natenapit, 2013); and (3) stimulisensitive and selective carriers for drug delivery to cells, organs, or tissues, under endogenous or exogenous stimulus (i.e., temperature, pH, ionic strength, redox potential or a magnetic gradient in the specific case of magnetic targeting) (Karimi et al., 2016). The magnetic manipulation of nanoparticles as translational vectors is useful for delivery purposes (Pankhurst et al., 2003). A distinctive advantage of the magnetomechanical property of magnetic nanoparticles is their ability to be externally guided to the desired cells or tissues by means of a remote magnetic influence. In this sense, magnetic nanoparticles can transform an external magnetic field into a mechanical force. The direction and amount of mechanical forces from magnetic nanoparticles can be modulated by saturation magnetization (Ms) of magnetic nanoparticles, the distance of the magnetic particle from the magnet, and the desired magnetic field parameters (e.g., field strength, gradient, and direction). Because the magnetic force of IONPs is proportional to its volume, nanoparticles with a larger core are desirable to obtain higher-saturation magnetization. However, the nanoparticles should be small enough to retain superparamagnetic properties (SPIO, USPIO), preventing nanoparticle aggregation in the absence of an external magnetic field, and to guarantee suitable blood circulation time of nanoparticles, as well as internalization processes. Once the nanoparticles reach the target site, they should release the therapeutic molecules. Recently, stimuli-responsive polymer (SRP)-coated SPIOs become very popular in the field of targeted delivery and cancer therapy (Wang et al., 2017). Among the reported SRPs, temperature and pH-responsive polymers are the most popular, where the drug release depends upon the temperature change and pH values, which vary in different tissues and cellular compartments. The pioneering family of the polymer poly(N-isopropylacrylamide) (PNIPAm) is the widely used temperature-responsive polymer which at below a critical solution temperature (CST) is soluble in water and after heating becomes insoluble. For example, a novel device based on MSNPs with IONPs encapsulated inside the silica matrix and covered with a thermo-responsive copolymer of poly(ethylenimine)-β-poly(N-isopropylacrylamide) (PEI/NIPAM) which also retain proteins in the polymer shell by physical interactions has been reported (Baeza et al., 2012). The nanodevice avoids the early release of cargo at low temperature (20
C), whereas the trapped molecules depart when the temperature exceeds 35
C40
C.

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Magnetic fields can also serve as an external stimulus to trigger the release of molecules from magnetic nanoparticles. A very illustrative example is the reported reversible stimuli-responsive drug-delivery systems under an alternating magnetic field (Ruiz-Herna´ndez et al., 2011). In such a system, SPIOs encapsulated in oligonucleotide-modified MSNPs, and loaded with fluorescein (as a model drug), are subsequently self-assembled with IONPs functionalized with the complementary strands, which perform as gatekeepers. A DNA double-stranded structure was designed to be cleaved at a temperature of 47
C, which corresponds to the upper limit of therapeutic magnetic hyperthermia. The magnetic-responsive release of this system was tested by the exposition of fluorescein-loaded MSNPs to an alternating magnetic field. Once the temperature reached 47
C the pore channels were uncapped and the cargo molecule was released.

13.5.1.1 Drug delivery for cancer treatment For drug delivery in cancer treatments, the integration into target-specific IONP formulations can limit unwanted side effects, while increasing the dosage at the diseased tissue (Tietze et al., 2015). The gathering of NPs at tumor sites is typically achived in two ways: passive and active targeting using or not using magnetic guidance (Fig. 13.2). Passive targeting is characteristic for NPs with sizes less than approximately 100 nm which can pass through leaky vessels into the tumor vasculature microenvironment due to the enhanced permeability and retention (EPR) effect. However,

FIGURE 13.2 Schematic representation of drug delivery by using therapeutic IONPs for cancer treatment.

13.5 Magnetite Nanoparticles in Disease Therapy

passive targeting is limited because not all tumors exhibit a suitable and homogeneous EPR effect across the whole site. Attempts at overcoming these limitations are made with active targeting by modifying NPs through the conjugation of targeting ligands possessing high affinity toward the unique molecular signatures found on malignant cells. In general, various categories of targeting ligands have been employed for tissue-specific targeting of drug nanocarriers (Ulbrich et al., 2016): proteins such as transferrin, aptamers that target specific antigen types expressed/overexpressed on the cancer cells (e.g., prostate-specific membrane antigen), cell-penetrating agents such as the transactivator of transcription (TAT)peptide, sugar moieties such as galactose, small molecule-targeting ligands such as anisamide and folic acid (FA), peptides such as chlorotoxin for brain tumor, and antibodies like Trastuzumab (breast cancer) and Rituximab (lymphoma). In addition, there are peptides and antibodies also present, eliciting therapeutic effects in a cell-specific manner, by inhibiting or stimulating various cellular pathways (e.g., activation of apoptotic/necrotic pathways, function blocking, and immune response stimulation). Once the vectorized IONPs are internalized into the tumor, the release of the drugs is allowed by the stimuli-responsive gatekeepers based mainly on their different biochemical characteristics compared to normal tissues (Liu et al., 2017). For example, tumors show a reductive environment due to the hypoxia and the overproduction of reductive biomolecules (e.g., reductase and glutathione). The pH of the tumor environment (6.2 2 7.2) is slightly lower than that of the normal tissue environment (7.4) due to the presence of acidic metabolites. The altered expression of specific enzymes (e.g., matrix metalloproteinases or cathepsins) is another characteristic that is found in pathological conditions. Moreover, externally controlled physical stimuli, such as light and magnetic field, can be promising options due to the spatiotemporal controllability of target-specific gene and drug release on the tumor. Cancer treatment is based mainly on the use of chemotherapeutics and therapeutic genes or nucleic acids. Chemotherapeutics encompass a broad category of small-molecule drug formulations, which have been developed to initiate a therapeutic response via cytotoxic, cytostatic, or antineoplastic effects. Currently, several stimuli-sensitive and selective IONP carriers for drug delivery have been combined with chemotherapeutics (Revia and Zhang, 2016), including paclitaxel (Sattarahmady et al., 2016), cisplatin (Voulgari et al., 2016), doxorubicin (Gautier et al., 2013), methotrexate (Farjadian et al., 2016), gemcitabine (Parsıan et al., 2016), docetaxel (Oshima et al., 2010), idarubicin (Gunduz et al., 2014), epirubicin (Voicu et al., 2014), 5-fluorouracil (Sagira et al., 2016), etc. For example, a magnetic cisplatin delivery nanosystem has been reported (Voulgari et al., 2016). In this sense, cisplatin was physically loaded on poly (methacrylic acid)-g-poly(ethyleneglycol methacrylate) polymers coating on magnetite nanoparticles synthesized by a coprecipitation method. Their antitumoral effect was tested in vitro by using a cisplatin-resistant HT-29 human colon adenocarcinoma model and in vivo in mice, and by applying magnetic field gradients.

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For the cisplatin-loaded nanocarriers, a major anticancer activity and lower toxicity than the free cisplatin were recorded, and also enhanced by applying a magnetic field gradient at the tumor site. Another illustrative example is based on the inhibition of MGMT (O6-methylguanine-DNA methyltransferase) overexpressed in glioblastoma multiforme (GBM) tumors by using the inhibitor O6-benzylguanine (BG) (Stephen et al., 2014). They reported superparamagnetic IONP capping with a biocompatible and redox-responsive chitosan-PEG copolymer (NPCP). Also, the NPCP was modified through covalent attachment of BG and tumor-targeting peptide chlorotoxin (NPCP-BG-CTX). Once the nanoparticles reach the tumor, the cargo (BG) is released under reductive intracellular conditions. The in vitro and in vivo assays demonstrated the improved biodistribution of BG and a threefold increase in survival in comparison with untreated animals.

13.5.2 HYPERTHERMIA The hyperthermia treatment of cancer with magnetic particles was introduced in 1957 (Gilchrist, 1957). Hyperthermia can be classified according to temperature ranges: mild hyperthermia (39
C42
C) to increase drug perfusion and oxygenation which can sensitize cells to radiotherapy or chemotherapy; moderate hyperthermia (41
C46
C) to stimulate degradation of proteins, interfering with essential cell processes and causing programmed cell death or apoptosis; and thermoablation (above 45
C) to kill the cells causing carbonization, coagulation, and necrosis (Cherukuri et al., 2010). The cells of tumors are irrevocably damaged under temperatures in the range of 42
C245
C. Also, heating the tumor increases the blood flow as well as its microvascular permeability and oxygenation. Such an altered tumor microenvironment becomes more permeable and accessible to drugs. Moreover, there is no evidence of a difference in heat sensitivity between normal and cancerous tissue. However, the vasculature of cancerous tumors sets up regions of lower pH and deprived oxygen concentration (hypoxia) that sensitize tumors to hyperthermia at temperatures between 40
C and 44
C. This range of temperature can be achieved utilizing magnetic particles in an external radiofrequency (RF) alternating current (AC) magnetic field (AMF), as was explained in Section 13.4.2.1. For a certain superparamagnetic material, the heating generated is spatially selective to the area in which these crystals are localized. The rate at which electromagnetic energy is absorbed by a unit mass of a biological material is defined as the specific absorption rate (SAR). For small anisotropy and crystal size nanoparticles, the SAR is proportional to the relaxation time (τ). Considering the progression of τ with the crystal volume, an ideal SAR for magnetite relates to a diameter range from 7 to 20 nm, which is also rapidly decreased by increasing the size polydispersity (Deatsch and Evans, 2014). Hyperthermia has been tested in in vitro and in vivo conditions by using magnetic fluid alone (MFH) (Guibert et al., 2017), MNPs coated with stabilizers, or

13.5 Magnetite Nanoparticles in Disease Therapy

by using encapsulated MNPs in delivery nanocarriers, such as liposomes (Gogoi et al., 2016). In addition, magnetic hyperthermia has also been combined with chemotherapy for gaining a more efficient antitumor response. For example, magnetic nanoparticle polymeric micelles (MNP-PMs) synthesized via the conjugation of magnetite nanoparticles with poly(ethylene glycol)-poly(lactide) (PEG-PLA) were evaluated as nanocarriers of doxorubicin (DOX), allowing hyperthermia and chemotherapy to be combined (Kim et al., 2015). When AMF was applied to heat the DOX-loaded MNP-PMs, the temperature was increased to the conditions of hyperthermia. Afterward, a magnetic stimulus resulted in a change in the motion of PLA chains, producing a rapid discharge of the drug molecules. When A549 cells were treated with MNP-PMs encapsulated with DOX and combined with hyperthermia under AMF, they discovered that 78% of the cells were destroyed by both the effects of the heat and the drug. Also, and more importantly, the synergic effect was more effective than separate chemotherapy or hyperthermia treatment. Also, guided magnetic fluid hyperthermia has been developed to allow homogeneous heat induction in the specific area where the tissue is compromised for tumor cells, in a less invasive way. In this case, NPs are functionalized with specific tumor cell-targeting ligands, as was explained in Section 13.5.1.1. For instance, the antitumor drug tamoxifen (TMX) was loaded as an inclusion complex inside surface-coated β-cyclodextrin (βCD) polymer on superparamagnetic magnetite nanoparticles (size of 12 nm diameter) and using folic acid (FA) as a targeting ligand for breast cancer tumor (Hayashi et al., 2010). When an alternating current (AC) magnetic field is applied to them (SAR of 132 W g1 at 230 kHz and 100 Oe), the induced heating generated the drug delivery from the βCD cavity in a controllable on/off process by switching the frequency of the magnetic field. In addition, improving the tumor cell specificity allows the enhancement of the efficacy and safety of the magnetic thermal ablation technique (e.g., up to B70
C) with necrosis of the tumors and intensive cell destruction (Hilger et al., 2002). A few studies report a clinical evaluation of magnetic fluid hyperthermia (MFH). In prostate cancer, MFH treatment was carried out by transperineal injection of six cycles of aminosilane-coated IONPs at weekly intervals, followed by heating the tumor tissue using AMF. The treatment was found to be welltolerated, but with temporary symptoms, such as urinary problems and tiredness (Johannsen et al., 2010). On the other hand, patients with recurrent glioblastoma multiforme were treated with aminosilane-coated iron-oxide nanoparticles using a 3D image-guided intratumoral injection and exposed to an alternating magnetic field, inducing a median maximum intratumoral temperature of 44.6
C (42.4
C49.5
C). The treatment was well countenanced for all patients with slight effects and there was evidence of local tumor reduction (Maier-Hauff et al., 2007).

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On the other hand, photothermia has been receiving more attention in recent years due to it requiring a lower concentration of nanoparticles than does magnetic hyperthermia, and it could be effective even after intravenous injection. The most-used phototermal agents are nanoparticles (e.g., plasmonic gold-based nanorods, nanostars, or nanoshells; graphene nanocomposites, or semiconductor nanomaterials) with strong optical absorptions of the near-infrared radiation (NIR) under low powers of laser and in-depth penetration of biological tissues. In recent years, magnetite was reported as a potential photothermal agent due to considerable absorbance in the NIR spectrum and their efficient biodegradation. The key issue to solve is the increase in the magnetite concentration in the tumor area. In this sense, targeting agents [e.g., arginine-glycine-aspartic acid (RGD) peptide] and magnetosomes are commonly used. One of the most efficient agents for magnetic hyperthermia and photothermia is magnetosomes synthesized by magnetotactic bacteria. In this case, a pure magnetite core is enveloped by a biological lipid bilayer membrane providing excellent biodistribution and biocompatibility, with less cytotoxic effects than chemically synthesized IONPs at equivalent concentrations. Also, magnetosomes conjugated with RGD peptide (magnetosomes@RGD) are synthesized by translational coupling of the RGD peptide with the magnetosomal protein MamC. A recent report demonstrated that magnetosomes@RGD exhibited much more efficient photothermia than magnetic hyperthermia, by inhibiting tumor progression in vivo in mice bearing PC3 tumors after systemic administration (Sangnier et al., 2018).

13.5.3 MAGNETOFECTION Gene and nucleic acid therapies rely on the introduction of nucleic acids into cells to modulate the overexpression of the desired protein, for the downregulation of endogenous gene expression, for bypassing or repairing endogenous genetic defects, or for activating the innate immune system. In this sense, magnetofection is defined as nucleic acid [e.g., DNA, single-stranded deoxyribonucleotide sequences (AODN), and small interfering double-stranded ribonucleotides (siRNA)] delivery under the influence of a magnetic field acting on nucleic acid vectors that are associated with magnetic nanoparticles (mainly, magnetite). Nucleic acid delivery can be carried out using viral vectors (transduction), or nonviral vectors (transfection). Magnetofection encompasses the following steps: (1) the synthesis of the DNA-loaded MNPs; (2) in vitro incubation with cells or in vivo systemic administration; and (3) guiding the vector near the target through the application of an external magnetic allowing the transfection. Many works have reported the significantly better-quality and confined transduction of assembling virus with magnetic particles, both in vitro and in vivo, when administered either intramuscularly or intravenously with magnetic guidance (Tresilwised et al., 2010). In this sense, biotin can be chemically, metabolically, or genetically coupled to the virus and (strept)avidin to the surface of

13.5 Magnetite Nanoparticles in Disease Therapy

magnetic particles or streptavidin-modified viral particles are captured by biotinylated magnetic particles (Lesch et al., 2010). For transfection, conjugates have been achieved by the association of nucleic acids on a magnetic core coated with a cationic polymer (e.g., PEI, PEI-oleate, and chitosan-coated SPION) (Huang et al., 2015) or by means of antigen 2 antibody or avidin 2 biotin interactions (Amjad et al., 2017). However, simple physical adsorption of nucleic acid molecules onto the particle surface may lead to their partial degradation and removal under a biological environment. Thus, a load of nucleic acids in magnetic cationic liposomes (MCLs) is generally chosen to safeguard them until intracellular delivery (Govindarajan et al., 2013). The benefits of magnetofection are summarized in several reviews and method papers (Plank et al., 2011). Some of these are: (1) for viral applications (e.g., viral titers; rapid characterization of the resistance of mutant viruses against antiviral agents; to synchronize viral infection; in transducing airway epithelial cells with lentiviral vectors; for establishing models and identifying mechanisms of infection with viral pathogens); (2) for neurosciences applications; (3) to define biochemical and signaling pathways in a broad variety of research field; (4) to identify mechanisms of cancer development; and (5) in gene silencing upon transfection of exogenous siRNA. Reports on nucleic acid delivery with magnetic nanoparticles in vivo are less abundant than the numerous publications on cell culture applications. One of the most outstanding works is the synthesis of a new formulation known as LipoMag combined with transplantable magnets, based on oleic acid-coated magnetite nanoparticles covered with cationic lipid suspended in a carrier liquid without agglomeration (Namiki et al., 2009). LipoMag can be used as an efficient magnetic vector for delivering siRNA into cells providing antitumor effects when systemically intravenously injected into mice bearing gastric tumors, with no evident associated effects. Lastly, Smolders et al. (2018) recently compared various transfection methods resulting in superior transfection efficiency for Glial-Mag magnetofection of BV2 cells.

13.5.4 SCAFFOLD-BASED TISSUE ENGINEERING Tissue engineering deals with the construction of biocompatible scaffolds based on two- or three-dimensional microstructures that provide the cells with attachment, growth, differentiation, and proliferation properties. Cell colonization of these scaffolds to regenerate a damaged organ/tissue can take place either in in vitro or in vivo conditions. For this purpose, a technique known as magnetic force-based tissue engineering (Mag-TE) has been reported based on nontoxic and biocompatible magnetic nanodevices [e.g., magnetic gelatin NPs, MNPloaded hydroxyapatite, and collagen (Bock et al., 2010), silk-fibroin electrospun composite fibers with different amounts of magnetite and magnetic cationic

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liposomes (MCLs) (Ishii et al., 2011)] to provide magneto-responsive features to the cells. Generally, this technology has been addressed by (1) loading the cells with, e.g., MCLs or chitosan-coated MNPs; (2) hydrating the scaffold with the suspension of the cells; and (3) applying a magnetic field, allowing highly packed cells into the scaffold. In addition, magnetic nanoparticles functionalized with cell adhesion peptide sequence RGD have been used for enabling efficient cell patterning (Ito et al., 2007). Moreover, employing magnetic force-based tissue engineering, different patterns of cell layers, three-dimensional configurations with complex shapes such as tubular structures for vascular tissue construction (Shimizu et al., 2007a), threedimensional tissues for bone and tissue reconstruction (Shimizu et al., 2007b), etc., are allowed. For instance, bone marrow-derived mesenchymal stromal cells (MSCs) labeled with ferumoxide (dextran-coated superparamagnetic iron-oxide nanoparticles approved by FDA for RMI application) were percutaneously injected into a rabbit ulnar defect (Oshima et al., 2010). This technique significantly facilitated the infiltration of ferumoxide-labeled cells into ceramic and significantly contributed to the enhancement of bone formation, even in the chronic phase by means of an external magnetic targeting system in a simulated clinical environment.

13.6 MAGNETITE NANOPARTICLES FOR MRI CONTRAST AGENTS: ORGAN/TISSUE/CELLULAR IMAGING Recently, various medical imaging modalities, including MRI, optical fluorescence, computed tomography (CT), and positron emission tomography (PET) have shown enormous potential for sensitive diagnosis. In particular, MRI has several advantages, including excellent anatomic detail, enhanced soft tissue contrast, and high spatial resolution, allowing the differentiation between malignant and healthy tissues. As was explained in Section 13.4.1.1 SPIO and USPIO mainly magnetite nanoparticles are extensively used as T2 contrast agents, resulting in dark signals. As was explained in Section 13.4.2, the biodistribution and internalization processes of magnetite nanoparticles into cells are strongly dependent on size. After administration, SPIOs smaller than 50 nm in diameter are rapidly taken up by hepatic and splenic macrophages/ having clinically used T2 contrast agent for imaging of liver tumors and metastases. Some commercially available SPIOs are: for liver tumor imaging Endorem or Feridex (Fe3O4 NPs of 150 nm in diameter coated with dextran) and Resovist (Fe3O4 NPs with a carboxydextran coating); for gastrointestinal tract imaging Lumirem (dextran-coated IONPs with a diameter of 300 nm). Therefore, USPIOs have been applied in MRI studies of several pathologies and have been tried in both preclinical animal models and clinical

13.6 Magnetite Nanoparticles for MRI Contrast Agents

assays (Liu et al., 2016). Ferumoxtran-10 (30 nm diameter, exhibiting human blood half-life of 24 2 36 hours) is allied to various diseased tissues. For instance, USPIO coated with dextran offers higher diagnostic precision for the imaging of lymph node metastases from various primary tumors; atherosclerotic plaques on the arterial wall in human patients suffering from carotid stenosis; the inflammatory response in central nervous system disorders; brain tumors and multiple sclerotic lesions; macrophage infiltration in organ-transplant situations such as kidney (for humans), lung (rat model), and knee joints (rabbit model) (Reddy et al., 2012). Similarly to therapeutic applications of IONPs, the specificity toward the target tissues may be improved by NP functionalization with targeting ligands for imaging purposes such as antibodies, peptides, and small ligands. An illustrative example is the conjugation of magnetite nanoparticles with the commercial antiHER2 monoclonal antibody trastuzumab (TMNC) by using protein A for in vivo application (Corsi et al., 2011). The TMNC conjugate allows (1) to determine the targeting efficiency for selectively recognizing the human epidermal growth factor receptor 2 (HER2) expressed in MCF-7 breast cancer cells in cultures and (2) biodistribution by combining MRI monitoring, highly sensitive epifluorescence tracking of TMNC in Balb/c nude mice bearing MCF-7 cells, and accurate tissue analyses. Another successful application of IONPs in MRI is specific cell tracking. The possibility of magnetic particles being efficiently cell-internalized via vectorized nanoparticles or transfection agents, as wells as the electrostatic interaction among the charged surfaces of magnetic particles and cell membranes, has provided a useful alternative for tracking cells after transplantation or transfusion in vivo by MRI. For instance, the most outstanding results are obtained for (Reddy et al., 2012): (1) imaging of transplanted ferumoxide-labeled human hematopoietic progenitor and embryonic stem cells; (2) imaging of transplanted feridex-labeled pancreatic islet cells for in vivo monitoring in diabetic rats; (3) imaging of SPIO-labeled dendritic cells to follow the migration of these cells to the adjacent lymph nodes; and (4) imaging and detection of apoptotic cells achieved with annexin V-conjugated IONPs. On the other hand, because different information about the same region of interest is obtained by each imaging modality, combining various imaging techniques provided with a single injection of contrast agent has been an attractive goal (Lee et al., 2015). Therefore, diverse multimodal imaging probes are based on multifunctional nanoparticles with SPIO. Some examples are: (1) dual T1-T2 MR imaging by introducing T1 contrast materials (e.g., conjugating Gd31 complexes), forming inorganic shells such as Gd2O(CO3)2, Mn-containing metal 2 organic framework (MOF) and tunneling magnetite nanoparticle morphology; (2) radionuclide-MR imaging including single-photon emission computed tomography (SPECT) and PET by combining radioactive metal ions (99mTc, 111In, 131I for SPECT and 18F, 64Cu, 124I for PET) with magnetic nanoparticles capped with chelating ligands; (3) CT-MR imaging by using heterostructured nanocrystals

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composed of radiopaque elements and magnetic nanoparticles (e.g., iron oxide 2 tantalum oxide core 2 shell nanoparticles); (4) optical-MR imaging with IONPs conjugated with fluorescent dyes [e.g., fluorescein isothiocyanate (FITC), rhodamine B, Cy5.5], gold, Eu-complex, quantum dots, and upconversion nanoparticles (UCNPs); and (5) ultrasound-MR imaging via the incorporation of IONPs within US contrast agent such as stabilized gas microbubbles.

13.7 MAGNETITE NANOPARTICLES AS THERANOSTIC PLATFORMS The term “theranostic” was defined as “a material that combines the modalities of therapy and diagnostic imaging” (Funkhouser, 2002). Nowadays, the main goal of the theranostic field is to optimize the ability to monitor diseased lesions, delivery kinetics, and drug efficacy, that allow rigorous therapy control to generate very specific and individual diagnoses of conditions, providing individualized and personalized medicine (Crawley and Thompson, 2014). Magnetite nanoparticles are perhaps the most-used platform for theranostic systems. The major benefits of biocompatible stabilized-vectorized IONP platforms are: (1) very high drug loading according to their physicochemical properties; (2) suitable and tunable pharmacokinetics, biodistribution, and biological fate with enhanced half-life and exposure of the NP-encapsulated drug; (3) ability for magnetic guidance with combined stimuli-sensitive and selective triggers; and (4) the possibility to use the particle itself for therapeutic function through hyperthermia treatment while producing MRI contrast for monitoring drug delivery and the therapeutic response (Han et al., 2016). It is noteworthy that most of the reports on IONPs for theranostic purposes have only been verified as proof-of-concept in in vitro conditions rather than in in vivo ones. To illustrate recently reported theranostic systems some examples will be described. For instance, Lee et al. (2013) reported IONPs conjugated with gemcitabine (G) allowing MRI contrast enhanced with the controlled intratumoral delivery of the mentioned chemotherapeutic agent. For this purpose, ATF peptides were used for specific recognition of uPAR-overexpressing pancreatic cancer target and tumor stromal cells. In addition, an enzymatic release trigger was used by linking the drug with the IONPs with a lysosomal enzyme-sensitive peptide. Furthermore, the systemic release of ATF-IONP-G significantly inhibited tumor growth in an orthotopic human pancreatic cancer xenograft model tracked by MRI signals. On the other hand, encapsulated gemcitabine (G) and fluorescent iron oxide (FIO) in poly(lactide-co-glycolide) (PLGA) nanospheres, functionalized with HER-2 antibody (Trastuzumab, Herceptin) yielding multifunctional nanospheres (PGFIO), have been reported. In this case dual therapeutic [chemotherapy,

13.7 Magnetite Nanoparticles as Theranostic Platforms

magnetic hyperthermia (MHT)] and imaging (MRI and fluorescence) modalities have been combined (Jaidev et al., 2017). The therapeutic efficacies of HERPGFIO and HER-PGFIO 1 MHT were studied in 2D, 3D in vitro MIAPaCa-2 cell culture models, and in vivo using a subcutaneous human pancreatic cancer xenograft model in SCID mice. The tumor regression in animals treated with HER- PGFIO 1 MHT was about 86% monitored by MRI due to the synergistic effect of the chemotherapy and magnetic hyperthermia against pancreatic cancer (Fig. 13.3).

FIGURE 13.3 (A) T2 MRI images showing contrast in tumors in different groups. [Group I (controluntreated), Group II (MHT treated), Group III (HER-PGFIO treated), and Group IV (HERPGFIO 1 MHT).] (B) Excised tumor from the respective animal at the end of the study. (C) Tumor volume regression plot across different groups. Permission from Jaidev, L.R., Raj Chellappan, D., Vasanth Bhavsar, D., Ranganathan, R., Sivanantham, B., Subramanian, A., et al., 2017. Multi-functional nanoparticles as theranostic agents for treatment & imaging in pancreatic cancer. Acta Biomater. 49, 422433 with license number: 4044870948072.

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13.8 MAGNETITE NANOPARTICLES FOR BIOANALYSIS The unique properties of magnetite nanomaterials have been utilized in the bioanalysis field, in magnetic separation techniques allowing sample preparation prior to biological analysis, as well as for the fabrication of nanosensors and nanobased immunoassays with enhanced specificity and sensitivity.

13.8.1 IN VITRO BIOSEPARATION Magnetic separation represents an attractive alternative method compared to the traditional methods (e.g., electrophoresis, ultrafiltration, precipitation, and affinity chromatography) for the selective and reliable capture of specific proteins, peptides, DNA, cells, and viruses in complex biological samples (He et al., 2014). Also, magnetic separation could be coupled with subsequent analytical techniques for detection and quantification, such as liquid chromatography (LC), mass spectrometry like matrix-assisted laser desorption ionization mass spectrometry (MALDI-MS), Fourier transform infrared spectroscopy (FTIR), surface-enhanced Raman spectroscopy (SERS), polymerase chain reaction (PCR), flow cytometry (FCM), and electron and fluorescence microscopy. The basic principle of magnetic separation is very simple. Magnetic nanoparticles (mainly magnetite) previously functionalized with an immobilized affinity tag, or ion-exchange groups, or hydrophobic/hydrophilic ligands, are mixed with the mixture sample (e.g., crude cell lysates, whole blood, plasma, urine, biological fluid, or fermentation broth) containing the desired molecules. After a suitable incubation time and experimental conditions, the affinity complexes are isolated by magnetic decantation and the contaminants washed out. Finally, after proper elution procedures, the purified target molecules are recovered. For proteins (antibodies and enzymes) and peptides the most-used affinity tags are: (1) MNPs functionalized with Ni21-chelating species, which allow for the selective binding of His-tagged proteins preventing nonspecific adsorption of undesired entities (Shao et al., 2012); (2) MNPs functionalized with protein A or G, which are anti-Fc for IgA and IgG antibodies, allowing selective affinities and reversible capture (Ma et al., 2012); (3) alkyl-functionalized MNPs, (meso)porous silicon and carbon-based magnetic materials enriched by n-alkyl (C8- or C18-) chains to capture peptides or proteins through hydrophobic interactions (Chen et al., 2009); (4) magnetic molecularly imprinted polymers (MMIPs) in which surface imprinting generated on magnetic clustering provides recognition cavities (in shape, size, and chemical functionalities) complementary to template molecules, which also correspond to the target molecules (Wang et al., 2011; Turan and Sahin, ¸ 2016); (5) grafted magnetic nanoparticles with highly positively charged metal ions (Fe31, Ga31, Zr41, Ti41, and Ce41) for rapid enrichment of phosphorylated peptides (Xu et al., 2006) and metal oxide (TiO2, Al2O3, ZrO2, Nb2O5, SnO2, CeO2, ZnO, Ga2O3, and Ta2O3) for enrichment of phosphopeptides

13.8 Magnetite Nanoparticles for Bioanalysis

(Nath et al., 2015 ); and (6) lectin-coupled magnetic beads, boronic acid-modified MNPs, hydrazide-functionalized magnetic beads, and hydrophilic magnetite surface (amine, zwitterionic) for capturing glycopeptides/glycoproteins (Hage et al., 2012). For separation of nucleic acids (including DNA and RNA) the principle is based on specific oligonucleotides immobilized on MNPs binding to specific sites of target nucleic acids in crude cell lysates and then being magnetically isolated (Zhang et al., 2012). The separation of cells (mainly bacterial and mammalian) from complex media or from mixtures of various cell types is accomplished by MNPs labeled with cell-specific aptamers, oligosaccharides, or antibodies able to recognize specific antigens expressed/overexpressed at the cell surface (Karumanchi et al., 2002). Separation of viral particles from virus contained in biological media has been achieved using virus-specific globulin proteins coupled with the MNPs and then magnetically separated (Borlido et al., 2013). Magnetic sorting of MNPlabeled cells and viruses has emerged as a promising technique. In addition, isolating rare cells, such as circulating tumor cells and endothelial cells from vast backgrounds of cells has been done by integrating magnetic actuation onto microfluidic chips for dip-stick methods and point-of-care diagnostics (Ozkumur et al., 2013).

13.8.2 BIOSENSORS AND IMMUNOASSAYS In biosensor and immunoassay designs, magnetite nanoparticles are widely used as transducer elements and/or magnetic platforms for immobilization, separation, and concentration of the analyte target prior to detection (Fig. 13.4). The magnetic transducer principles are classified into two main groups: magnetic resonance-based detection and magnetoresistive sensors (Sobczak-Kupiec et al., 2016).

FIGURE 13.4 Scheme of the application of magnetite nanoparticles in different bioanalysis techniques and by combining with other nanoparticles. Inside: Principle of magnetic resonance-based detection.

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Magnetic resonance detection is based on the inhomogeneous magnetic fields of magnetic nanoparticles that accelerate dephasing of surrounded water protons, thus leading to a change in proton relaxation times (T2). For instance, there are two main magnetic resonance assays: (1) cellular labeling, in which MNPs are targeted to surface biomarkers on cells and unbound MNPs washed away with further MRI detection and (2) magnetic relaxation reversible switches by using magnetic resonance imaging (MRI) or relaxometry techniques. In the last case, reversible aggregation and dissociation of magnetic nanoparticle clusters are induced by target antigens, leading to a change in proton relaxation times (ΔT2) (Colombo et al., 2012). On the other hand, magnetoresistive sensors are based on magnetoresistance phenomena in which the changeable electronic resistance of the material is obtained by exposing it to a magnetic field. For instance, a giant magnetoresistive (GMR) sensor array for sensitive and specific multiplexed food allergen detection in pg\mL ranges has been reported (Ng et al., 2016). In this work, binding of a magnetically labeled biomolecule on an immunoassay structure generates a localized magnetic field that changes the resistance of the GMR sensor chips. The most widespread use of magnetite nanoparticles and beads is as immunomagnetic platforms. Magnetite nanoparticles provide a solid support with a remarkable surface area, increasing the conjugated biomolecule amount and avidity, allowing for improvment of analytical sensor performance. In addition, washing protocols can be automated with magnets. The most common studies are based on a sandwich assay supported by magnetite microbeads and superparamagnetic nanoparticles in: (1) colloidal suspension (Kourilov and Steinitz, 2002); (2) cellulose paper (Ortega et al., 2017); (3) conventional and printed electrodes (Stefano et al., 2014); (4) thin-film transistor (Hu et al., 2016); (5) microfluidic system (Tsaia et al., 2010); and (6) automatic portable and screening devices (Jeong et al., 2016) for the separation and detection of proteins (Kriz et al., 2005), antibodies (Ortega et al., 2016), nucleic acids (Bruno et al., 2014), viruses (Hung et al., 2014), bacteria (Martı´n et al., 2015), toxins (Lin et al., 2011), and cancer biomarkers (Ho et al., 2016). In addition, immuno-magnetic platforms are combined with metal (e.g., gold and silver) (Liang et al., 2015), semiconductors (quantum dots) (Suqin and Yunsheng, 2016), and carbon nanoparticles (carbon nanotubes, graphenes) (Demeritte et al., 2015) as transducers and signal amplification in new detection methods like optical [e.g. (1) colorimetric (Fu et al., 2016), (2) surface plasmon resonant (SPR) (Singh et al., 2016), (3) local (LSPR) (Tang et al., 2013), (4) surface-enhanced Raman scattering (SERS) (Wang et al., 2015), and (5) luminescent sensors (Liu et al., 2014)], and electrochemical nano-based immunoassays (Cortina et al., 2016). Generally, in the last case, magnetic nanoparticles aid the localization of the immunocomplexes on the electrode surface (Martin et al., 2014) as well as allowing large electrochemical response-amplified signals by acting as an electrocatalytic (Krishnan and Walgama, 2013) and nanocarriers of electroactive compounds (Otieno et al., 2014).

13.9 Final Remarks

Among all procedures that combine magnetic and metal nanoparticles, the barcode strategy first reported by Mirkin et al. exhibits powerful amplification ability and has been widely used in sensitive multiplexing detection of biomarkers (Nam et al., 2003). The technique relied on magnetic particles functionalized with specific antibodies (Abs) for antigen capture and gold NPs functionalized with secondary Abs (to recognize the antigen for a different region) and barcode oligonucleotides that can sandwich the target captured by magnetic nanoparticles. The resulting sandwich layout is magnetically removed from the medium and then, the barcode DNA is isolated to be quantified by using real-time polymerase chain reaction (PCR), chip-based scanometric methods, and fluorescence (Wei et al., 2016).

13.9 FINAL REMARKS Magnetite nanoparticles have been investigated for a wide range of biomedical and healthcare applications based on their surface chemistry, high biocompatibility, and, most outstandingly, the high magnetic susceptibility and the loss of magnetization at zero fields preventing their agglomeration in biological systems. In this sense, the magnetic properties of magnetite nanoparticles allow their use in several biomedical applications, such as: drug delivery with magnetic guidance and external stimulus, chemotherapy-based hyperthermia, magnetofection, magnetic force-based tissue engineering, MRI contrast agents, in vitro magnetic bioseparation, bioanalysis like immunomagnetic platforms and magnetic transducers, and theranostic platforms. Furthermore, nowadays the main challenge in magnetite biomedical applications is to achieve a real advancement in temporal and spatial site-specific drug delivery, local hyperthermia, and imaging in human patients with no secondary effects. For those purposes, some key aspects must be improved: (1) a controllable and scaling-up synthesis with tunable and narrow size distribution of magnetite nanoparticles with whole shell protection allowing optimal colloidal stability in physiological media, but minimally affecting their magnetic properties; (2) an effective and oriented bioconjugation of a large amount of targeting ligands on the nanoparticle surface without affecting their biological activity to enhance their magnetic guidance into the body; (3) an improvement in the efficiency of nanoparticles as contrast agents by the correct design of the structure of the magnetic core; (4) an enhancment of the magnetic moment of nanoparticles by controlling the intrinsic and extrinsic material properties, such as material composition and crystal structure, size, shape, and organic capping for hyperthermia applications; and (5) well-established toxicologic studies in the human body of such nanoparticles, as well as the drugs loaded in the carrier systems. Despite the fact that the real consolidation of the use of IONPs in the biomedical field is still in its first stages, with just a few examples of those nanoparticle

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systems in clinical trials, the promissory results in recent research predict that these nanomaterials will be the most outstanding in the near future in medicine and theranostics.

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