Introduction: Smart Materials in Biomedicine

1 Introduction: Smart Materials in Biomedicine Elisa Mele LOUGHBOROUGH UNIVERSITY, MATERI A L S D E P A R T M E NT , L OU G H B OR OU G H , UNITED K INGDOM

1.1. Historical Evolution of Smart Nanomaterials Over the last decade, advances in material science and nanotechnology have led to the development of highly engineered materials that are able to cleverly modify and adapt their physicochemical properties in response to external stimuli. They are often referred to as “smart materials” [1]. For biomedical applications, smart materials with bespoke functionalities and responsiveness have been manufactured into a multitude of nanostructures, including nanoparticles (NPs), nanorods, nanogels, and micelles (Fig. 1.1A) [2–4]. They have been designed to respond to a wide variety of physical and chemical stimuli such as temperature, light, electromagnetic fields, mechanical stresses, pH, enzymatic activity, sugar concentration, and oxidative reactions (Fig. 1.1B) [2]. Those diverse classes of smart nanomaterials are the result of an evolutionary process, which has been initiated and strongly driven by studies on how biological systems interact with nanostructures [5]. Structures with sub-100 nm size are indeed prone to penetrate across in vivo barriers and to be internalized by cells. Initially, biocompatibility and cell uptake tests were conducted on water-stable quantum dots, gold, and iron oxide NPs, in order to verify their potential for biomedicine [5]. However, in vivo tests revealed a rapid renal clearance of this first generation of NPs due to the lack of suitable surface functionalization. Surface chemical treatments were then implemented in order to prolong blood circulation half-life and to achieve targeted delivery. The second generation of NPs relied mostly on poly(ethylene glycol) (PEG) and other ligands for anchoring themselves to cellular receptors. They passively accumulated inside target organs and tissues due to the enhanced permeation and retention effect [1]. Limitations associated with this passive and transient retention, and with the inability of the functionalized NPs to progress beyond the first few layers of cells in a tissue, moved the research toward stimuli-responsive nanomaterials. Smart nanomaterials have received great attention in recent years because they are able to overcome passive retention mechanisms and nonspecific cellular uptake. In fact, they exploit the physiological conditions of the target site or artificial environmental cues to

Smart Nanoparticles for Biomedicine. https://doi.org/10.1016/B978-0-12-814156-4.00001-X © 2018 Elsevier Inc. All rights reserved.

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FIG. 1.1 Schematic of (A) the diverse classes of stimuli-responsive nanomaterials and of (B) the variety of triggering mechanisms. Reproduced with permission from Ref. [2]; copyright (2017) Elsevier.

trigger therapeutic actions. They have been widely used in diverse biomedical fields, including drug, gene, and protein delivery, tissue engineering, biological imaging, and sensing. As case of study, this introductive chapter discusses one of the most promising biomedical applications of stimuli-responsive nanomaterials, which is the triggered release of drugs, proteins, and genes [6,7]. Smart nanocarriers offer the advantage of sitespecific release of therapeutic agents, and they also provide temporal and dosage control. Chemical composition, size, and shape of the nanocarriers, as well as preparation procedures and interactions with bioactive compounds, can be precisely modified in order to tailor cargo capacity, entrapment efficiency, and release profile [8]. Thanks to those properties, stimuli-responsive nanocarriers are nowadays regarded as potential candidates for the fabrication of high-efficient drug delivery systems (DDSs) that can respond to endogenous or exogenous stimuli or to multiple combinations of those. Here, an overview of the recent progress in the design and synthesis of nanocarriers of therapeutic molecules that respond to endogenous and exogenous stimuli will be provided.

1.2. Endogenous Stimuli Differently from healthy tissues, diseased tissues (tumor, inflamed and infected tissues, etc.) present changes in angiogenesis and in the structure of proteins and enzymes, abnormalities in pH conditions (acidic environment) and metabolic states, increased temperature, enhanced production of reactive oxygen species (ROS), and hypoxic response (inadequate oxygen supply) [1,6]. Those naturally occurring conditions, in

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particular gradients of pH and temperature and redox processes, have been exploited as endogenous stimuli for triggering the release of drugs from nanocarriers.

1.2.1. pH Gradients pH-sensitive nanocarriers, which are stable at physiological pH (7.4) but undergo modifications when exposed to pH imbalances, utilize acidic microenvironments within the human body to enable drug release and to enhance the bioavailability of therapeutic payloads, as shown in Fig. 1.2 [7,9]. The triggering mechanism can be based on protonation/ deprotonation of functional groups or on the cleavage of acid bonds [10]. Nanocarriers that respond to pH gradients along the gastrointestinal (GI) tract, from pH 1–3 in the stomach to pH 6–8 in the intestine [11], have been reported for oral drug delivery applications [12]. For example, pH-sensitive and biodegradable NPs of methoxy PEG-block-(poly(caprolactone)-graft-poly(methacrylic acid)) (mPEG-b-(PCL-g-PMAA)) have been synthesized for the release of hydrophobic drugs such as ibuprofen (IBU, an antiinflammatory drug) [13]. The block copolymer was designed to contain PMAA blocks, whose carboxylic groups accept protons at low pH and loose protons at neutral and high pH. In the stomach (low pH), PMAA chains form a protective layer around the core of NPs (containing the drug), limiting the release of ibuprofen (Fig. 1.3A). Once the NPs reach the intestine (neutral pH), the PMAA chains stretch and the drug is released. The release profile of IBU from mPEG-b-(PCL-g-PMAA) NPs was studied in vitro at 37°C for pH 3.0 and pH 7.4 (Fig. 1.3B). It was observed that about 55% of IBU was released at pH 3.0 and about 85% at pH 7.4, within 12 h. Importantly, the pH sensitivity of the NPs was tailored by modifying the MAA ratio in the block copolymer. When the number of MAA monomers increased, IBU release of 40% and 95% were achieved at pH 3.0 and 7.4, respectively, within 12 h. The ability of PMAA to respond to pH changes has also been exploited to control the oral release of other drugs [15,16], such as insulin for the treatment of diabetes [17], doxorubicin (DOX) for chemotherapy [18,19], and antibiotics such as metronidazole [20] and amoxicillin [21]. Together with PMAA, pH-responsive materials for oral drug delivery include Eudragit® (poly(methacrylic acid-co-methyl acrylate)), modified chitosan, and porous silica [12]. This class of smart NPs helps to protect the active therapeutic compounds from lowpH environments in the GI tract and to achieve organ-specific release. Nanocarrier systems with acid labile chemical bonds are widely utilized to deliver anticancer drugs with temporal and spatial specificity by targeting the acidic microenvironment of tumor tissues (pH values from 5.7 to 7.8 [22]) [10,23]. This strategy is based on the conjugation of drugs with both organic and inorganic nanomaterials. For instance, pH-sensitive hydrazone bonds have been exploited to link DOX to hydrophobic poly(L-aspartate) segments in gold NPs [14]. The NPs were stabilized with a monolayer of folate-conjugated poly(L-aspartate-doxorubicin)-b-poly(ethylene glycol) copolymer (Au-P(LA-DOX)-b-PEG-OH/FA) (Fig. 1.3C). Because the hydrazone linkages experience hydrolysis under acidic conditions [24], once the Au-P(LA-DOX)-b-PEG-OH/FA) NPs are

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FIG. 1.2 Schematic representation of pH variation in the body at organ, tissue, and cellular level and representative classes of pH-responsive DDSs. Reproduced with permission from Ref. [9]; copyright (2017) Elsevier.

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FIG. 1.3 (A) Scheme of the formation of mPEG-b-(PCL-g-PMAA) NPs in water and of the release of ibuprofen at pH 3.0 and 7.4; (B) in vitro release profile of ibuprofen from mPEG113-b-(PCL92.5-g-PMAA81) (black symbols) and mPEG113-b-(PCL91-g-PMAA155) (red symbols) NPs at 37°C, at pH 3.0 (square symbols), and 7.4 (triangular symbols). (C) Schematic representation of the pH-triggered release of DOX from Au-P(LA-DOX)-b-PEG-OH/FA NPs, and (D) release profile of DOX at 37°C, at pH 5.3 (circles), 6.6 (triangles), and 7.4 (squares). (A, B) Reproduced with permission from Ref. [13]; copyright (2013) The Royal Society of Chemistry. (C, D) Reproduced with permission from Ref. [14]; copyright (2009) Elsevier.

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internalized by tumor cells, the low pH of the endosomal intracellular compartments prompts the release of DOX. As shown in Fig. 1.3D, in conditions simulating the physiological bloodstream (37°C and pH 7.4), a slow release of DOX was measured (15% within 45 h). On the contrary, when the NPs were exposed to the simulated environment of the tumor tissue (pH 6.6) and of the endocytic compartment of the tumor cells (pH 5.3), the rate of DOX release drastically increased (83% at pH 6.6 and 94% at pH 5.3, within 45 h). Therefore, the use of pH-responsive linkages enhanced the efficacy of cancer therapy by ensuring an appropriate level of DOX release at the tumor site and a minimum one during blood circulation. The efficacy of nanostructures containing pH-sensitive hydrazone linkages has been demonstrated by in vivo studies on hepatocarcinoma [25] and melanoma [26]. Micelles of N-(2-hydroxypropyl) methacrylamide (HPMA) copolymer conjugated with DOX via hydrazone linkages have been synthesized and tested for an H22 mouse xenograft model of hepatocarcinoma [25]. Micelles with and without glutaraldehyde crosslinking (diameter of 10–20 nm) were radiolabeled and intravenously injected in mice. Their persistence in blood and accumulation in tumor tissue and major organs (brain, liver, kidney, and heart) were measured, showing that the cross-linked micelles had a greater accumulation in tumor tissue (2.34-fold higher), longer biological half-life (1.89-fold higher), and slower clearance rate (1.45-fold lower) compared to the non-cross-linked ones. The cross-linked micelles exhibited also a high antitumor activity with a significant inhibition of the tumor growth (71.8%), thanks to hydrolysis of the hydrozone linkages that rapidly and efficiently release the drug at the tumor site. An emerging application of pH-responsive materials is the treatment of inflammations and infections in chronic wounds, where pH is between 5.4 and 7.4 [27]. The current research in this area aims at producing dressings that are sensitive to changes of pH depending on the type and on the healing stage of the wound and on the presence of bacteria. So far, pH-sensitive hydrogels and fibrous scaffolds have been proposed for the controlled release of antibacterial and antiinflammatory compounds to the wound bed [28–31].

1.2.2. Temperature Changes DDSs based on temperature-sensitive nanomaterials exploit the hyperthermia of inflamed tissues and tumors (temperature in the range of 40–42°C) to achieve spatiotemporal-controlled release [7,32]. Thermoresponsive nanocarriers are able to reversibly alter their phase or volume in response to temperature changes. Consequently, the payloads are retained inside the nanocarriers at body temperature (35–37°C) during blood circulation and released to the target site where temperature is higher. Poly(N-isopropylacrylamide) (PNIPAAm) and its derivatives are widely studied as thermoresponsive polymers because they undergo phase transition at temperatures very close to physiological human temperature [7,32]. They are in a hydrate state (polymer chains in a hydrophilic coil state) below 32°C, which corresponds to their lower critical solution temperature (LCST), whereas they become less water soluble above LCST (polymer chains in a hydrophobic globule state) [33]. This behavior is due to the presence of both

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hydrophilic and hydrophobic moieties in the repeating unit of PNIPAAm, and it can be tuned by the addition of specific functional molecules. Biodegradable micelles of PNIPAAm and poly(lactide) with a LCST varying from 32.3°C to 39.1°C have been synthesized by introducing hydrophilic dimethyl acrylamide (DMAAm) in the polymer chain [34]. They have been used for targeting the delivery of amphotericin B (AmpB), a poorly water-soluble antimicrobial agent, and of curcumin, an antitumor natural compound [34,35]. The LCST of P(NIPAAm-co-DMAAm)-b-PLLA-b-P(NIPAAm-co-DMAAm) triblock copolymer increased linearly by increasing the content of DMAAm, reaching a value of 39.1°C when 24% of DMAAm was used. Importantly, the micelles were sensitive to small temperature changes. For micelles with a LCST of 37.8°C, from 32% to 50% of AmpB was released at 37°C after 3 days (depending on the drug content), whereas a much faster release (75%–90%) was measured at the same time point at 38°C, which is above the LCST. The ability of synthesizing PNIPAAm-based materials with the desired LCST has allowed the development of DDSs that selectively target the tissue/organ of interest. Examples are amphiphilic micelles of P-(N,N-isopropylacrylamide-co-N-hydroxymethylacrylamide)b-caprolactone (P-(NIPAAm-co-NHMAAm)-b-PCL), used for the treatment of cholangiocarcinoma, which is an epithelial cancer of the bile ducts [36]. P-(NIPAAm-coNHMAAm)-b-PCL micelles with a LCST of 38°C (temperature of the bile duct) and loaded with DOX showed a significant release of the chemotherapeutic agent only at temperature higher than LCST. Moreover, when they were injected into the tumor site of animals, they induced a 21.5% reduction in the growth of cholangiocarcinoma. Smart materials with high sensitivity to temperature have been applied in diverse biomedical areas, including therapies for vascular, cutaneous, and uterine diseases, corneal treatments, and cancer chemotherapy [7]. They have also been designed for the delivery of genes, nucleic acids, and proteins [2].

1.2.3. Redox Processes Differences in concentration of glutathione (GSH), which is an antioxidant tripeptide of glutamate, cysteine, and glycine, exist between healthy and diseased tissues (GSH levels are at least fourfold higher in tumor than in normal tissue) and between intracellular (2–10 mM) and extracellular (2–20 μM) microenvironments [37]. The resulting redox potential is used as stimulus for the design of DDSs that are stable in extracellular spaces and release their payloads in reducing conditions (within cytosol, nucleus, and mitochondria). Recently, silica NPs containing redox-triggered disulfide (DS) bonds have been developed for the delivery of genes and drugs [38–41]. DS linkers have been used to functionalize the surface of mesoporous SiO2 nanoparticles (MSNs) with a dendronized chitosan derivative (CP), in order to achieve codelivery of DOX and p53 gene [38]. When the MSN-DS-CP NPs were exposed to a high concentration of GSH (10 mM), a release of DOX of 80% was observed after 24 h, with a fast release already after 6 h. On the contrary, for a GSH concentration of 10 μM, 35% of DOX release was measured after 24 h. This was due to the reductive cleavage of DS bonds by GSH, leading to detachment of end-capping

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CP molecules from the surface of NPs and to free DOX diffusion in-and-out of the pores. In vitro studies on human cervical carcinoma cells (HeLa) showed that MSN-DS-CP/DOX and MSN-DS-CP/p53 induced a cell apoptosis of 22.6% and 15.5%, respectively, whereas cytotoxicity values reached 42.1% when the cells were treated with MSN-DS-CP/DOX/p53 complexes, thanks to the synergic effect of DOX and p53. A plethora of studies have demonstrated the role played by redox-sensitive DS linkages in the controlled release of DOX and of other antitumor agents such as paclitaxel (PTX), cyclopamine, camptothecin [42–47]. Together with SdS bonds, diselenide and ditellurium groups have also been analyzed [48], highlighting the significant clinical potential of this class of smart nanomaterials for cancer therapy.

1.3. Exogenous Stimuli Nanomaterials that are activated by stimuli external to the human body, such as electromagnetic fields, light, mechanical forces, and temperature, offer the advantages of high spatiotemporal selectivity in cargo delivery, invariance from patient physiological conditions, and minimal collateral damages to surrounding healthy tissues [1,32]. This section will particularly focus on magnetic nanomaterials because they are nowadays regarded as promising theranostic tools and they have been approved for clinical trials [32]. Magnetic nanoparticles (MNPs) for biomedicine are mainly based on ferrite colloids, magnetite (Fe3O4), and maghemite (γ-Fe2O3), and are currently used for drug release but also for magnetic resonance imaging (MRI), hyperthermia therapy, biosensing, and tissue engineering [49–51]. Local external magnetic fields can either guide MNPs toward the diseased site or generate thermal energy through magnetic hyperthermia. In a recent work, superparamagnetic iron oxide nanoparticles (SPIONs) have been conjugated with β-cyclodextrin (β-CD) and polymerized PTX (pTX), in order to achieve high stability of the drug during blood circulation [52]. In vitro studies on HeLa, human breast carcinoma (MCF-7) cells, and mouse invasive colon cancer (CT26) cells showed that the cytotoxicity of pPTX/CD-SPIONs increased with the application of a magnetic field, due to the accumulation of the NPs at the targeted cancer cells. The anticancer efficacy of pPTX/CD-SPIONs was tested on CT26-bearing mice by locally exposing the tumor region to an external magnetic field. A significant reduction of the tumor volume was observed 16 days after the intravenous injection of the MNPs. SPIONs have also been modified with PEG, poly(ethylene imine) (PEI), and polysorbate 80 (Ps 80) coatings, and loaded with DOX for targeting gliobastoma multiforme (a malignant brain cancer) [53]. The coating provided sites for drug loading via DOX-polymer hydrogen bonding (Fig. 1.4A). When DOX@Ps 80-SPIONs were exposed to the magnetic field, the bioavailability of DOX was enhanced due to the direct contact of the MNPs with the C6 glioma cells. As a consequence, DOX@Ps 80-SPIONs showed higher cytotoxicity than just DOX after 48 h of incubation with cells, and the highest apoptosis level (42.2%) in the presence of magnetic field. Real-time MRI of gliomas in rats revealed the antitumor efficacy of DOX@Ps 80-SPIONs,

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FIG. 1.4 (A) Schematic representation of the working principle of SPIONs loaded with DOX and coated with a polymer layer. The drug is released at the tumor site in the presence of an external magnetic field. (B) Temporal evolution of tumor volume monitored by MRI in animals treated with a saline solution, DOX, DOX@Ps 80-SPIONs, with and without magnetic field (MF). (C) MRI monitoring of tumor in mice treated with SPIONs and exposed to MHT cycles. The area highlighted in yellow is the tumor tissue; SPIONs are located in the black area within the tumor. (A, B) Reproduced with permission from Ref. [53]; copyright (2016) The Royal Society of Chemistry. (C) Reproduced with permission from Ref. [54]; copyright (2016) Elsevier.

that were able to remarkably slow down the tumor growth, particularly in the presence of the applied magnetic field (Fig. 1.4B). SPIONs have been explored for magnetic hyperthermia therapy (MHT) of human prostate cancer and glioblastoma multiforme in clinical trials [55,56]. MHT is a minimally invasive cancer therapy that utilizes the thermal energy generated by MNPs under an alternating magnetic field (AMF) in order to kill tumor through apoptosis-mediated cell death. SPIONs-loaded nanocapsules of thermoresponsive poly(organophosphazene) hydrogels have been recently synthesized for multiple MHT and simultaneous MRI [54]. Animal tests showed that the temperature of glioblastoma injected with SPIONs increased up to 45°C after multiple MHT, affecting the tumor growth. The thermal energy

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delayed the initial tumor growth within the first 5 days, but one or two cycles of MHT were not sufficient to stop tumor development (Fig. 1.4C). On the contrary, the tumor was completely eradicated after four MHT cycles, without inducing damages to the surrounding healthy tissue (Fig. 1.4C).

1.4. Future Perspectives As discussed in this chapter, significant advances have been achieved in the design and synthesis of smart nanomaterials that exert a therapeutic action by responding to a stimulus. In order to improve the efficiency of those systems when exposed to challenging microenvironments within the human body, scientific attention is now devoted toward materials that possess multiple functions and that are multiresponsive [7]. Smart biohybrid NPs that simultaneously or sequentially respond to combinations of temperature, pH, and redox stimuli have been developed for triggering drug release in intracellular and reductive environments [57]. Polymeric NPs that are sensitive to near-infrared light, pH, and ROS have shown their efficacy against lung cancer by exploiting the synergic effect of photothermal therapy and chemotherapy [58]. The final aim of those research efforts is to create autoregulating therapeutic agents that make the most of both physiological cues and external localized sources of energy.

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