CHAPTER 1: Fundamentals of Stimuli-responsive Drug and Gene Delivery Systems

The use of drug-loaded nanoparticulate pharmaceutical carriers (drug delivery systems, DDS) can overcome many of the limitations characteristic of free therapeutic entities, such as low stability, fast inactivation or degradation in vivo, non-specific toxicity, poor solubility, unfavorable pharmacokinetics, and poor biodistribution. Nanocarriers can be designed/engineered that may entrap both hydrophilic and hydrophobic drugs, increase their stability and longevity in vivo, provide controllable drug release, change drug pharmacokinetics in any required way (in other words, drug's pharmacokinetics can be replaced with the specifically designed pharmacokinetics of the pharmaceutical carrier/delivery system), minimize undesirable side effects, and allow targeting of drugs to pathological organs and tissues of interest, or even to individual cells and intracellular organelles.^1–3  Currently, there exists a broad variety of nanoparticulate carriers meeting different practical requirements, which includes liposomes, polymeric micelles, polymeric nanoparticles, niosomes, solid lipid nanoparticles, dendrimers, nanoemulsions, inorganic nanoparticles, nanoshells, nanotubes, and many others.^1,4–7 

If required, nanoparticulate DDS can be specifically engineered to selectively change certain parameters/functions (for example, enhance drug release or intracellular uptake) in response to certain intrinsic stimuli characteristics of pathological tissues or external stimuli applied from the outside of the body.^4,8–11  The intrinsic/internal stimuli that are characteristic for the pathological areas (tumors, infarcts, athrosclerotic lesions, sites of infection or areas of transplant rejection) make those areas different from normal tissues, include local changes of different intensity in pH,^12,13  temperature,^14–16  redox conditions,^17,18  and the expression of certain biologically/enzymatically active molecules.^19,20  External stimuli include magnetic field, heat, light (including laser beams), and ultrasound, see Figure 1.1.^8,11 

A better understanding of local changes in the microsurroundings of pathological areas, including tumors, has allowed for the development of novel materials and engineering principles in order to build DDS capable of specifically reacting to certain “abnormal” conditions of affected tissues.²¹  Thus, for example, many lipids and polymers can undergo certain changes in their physical–chemical properties as a response to even insignificant changes in pH or temperature, which can be used to engineer smart DDS.²²  Upon exposure to intrinsic or external stimuli, such stimuli-sensitive DDS switch on certain functions controlling drug release or the efficacy of the intracellular uptake/penetration.⁴  Drug release can be controlled by the structural changes in a microcontainer type DDS loaded with the drug (such as carrier degradation or an increase in its permeability) or by the cleavage of chemical bonds used to attach the drug to the carrier. The specific biochemistry of various affected areas helps to engineer stimuli-sensitive DDS responding to these conditions, which can allow for precise temporal and spatial drug delivery improving its effectiveness and minimizing off-target effects.^4,23,24 

We will present here an overview of stimuli-sensitive DDS. Redox potential, pH, hypoxia, and overexpression of certain catalytically-active molecules represent a group of intrinsic factors/stimuli characteristics of affected/pathological areas/tissues in the body. Magnetic field, light, and ultrasound belong to a group of external stimuli. Temperature can serve as both an intrinsic stimuli (local temperature increase characteristic of the inflamed areas) and as an external stimuli if applied from outside the body, or generated inside a certain area in the body under the action of external stimuli (such as high frequency magnetic field).

Low pH is a well-established property of tumors and inflamed and hypoxic areas. In 1930, Warburg and co-workers showed that tumor cells preferentially convert glucose and other substrates to lactic acid (pK 3.7), even under aerobic conditions, and it was thought that lactic acid was the main source of tumor acidity. However, it is now known that the decreased pH is also related to the high levels of CO₂ and the increased expression and activity of vascuolar type proton pumps.^25–27  Deficiencies in tumor perfusion, hypoxia, and metabolic abnormalities associated with uncontrolled cell growth and proliferation (tumorigenesis) also contribute to the low pH of tumors.

A pH-sensitive DDS has to protect the loaded drug when in the blood (pH 7.3–7.4) and release the drug or expose some of the special functions (such as cell penetrating ability) when in the interstitial space of tumors (pH 6.8–7.2) or in the intracellular compartments, such as cytoplasm (pH slightly above 7.0), endosomes (pH 5–6), and lysosomes (pH 4.5–5.5).^28–30  This allows for the massive drug release, which is especially important in tumor treatment, as if high concentrations of free cytotoxic drug appear intracellularly, it is possible to exceed the efflux capacity of drug transporters (such as P-gp) and kill the tumor cells efficiently. Drug release can also occur in slightly acidic extracellular fluids of tumor tissue after tumor accumulation via the enhanced permeability and retention (EPR) effect. In endosomes or lysosomes, which have an even lower pH value than extracellular fluids, pH-triggered drug release proceeds with high efficiency when assisted by hydrolysis (with enzymes such as catepsin B).^31,32 

DDS for pH-triggered drug release usually contain protonizable components (amino, imidazolyl, and carboxyl groups), which at acidic pH values can induce the DDS destabilization and drug release as a result of protonation. Thus, adriamycin-loaded poly(l-histidine)-b-PEG-containing mixed micelles are non-protonated and stable at pH 7.4, at pH 6.6 they undergo protonation of the histidine residues, which destabilizes the micelle structure and results in burst release of a loaded anticancer drug.³³  pH-sensitive polymers (containing acidic–carboxylic or sulfonic groups or basic ammonium salts groups) undergo pH-sensitive conformational changes resulting in dissociation, destabilization (such as collapsing or swelling), or changes in the partition coefficient drug/vehicle.³⁴  DDS with acid-labile chemical bonds are also used for pH-triggered drug release. Chemical bonds, such as the ortho ester bond, the acetal/ketal bond, and the hydrazone bond can break down in acidic conditions. When a drug is bound to the DDS via such bonds, it will be released upon acidification.^34,35  Doxorubicin conjugated onto hydrophobic segments of the folate-modified amphiphilic hyperbranched block-copolymer, H40-P(LA–DOX)-b-PEG–OH/FA, via the acid-labile hydrazone linkage were released rapidly at acidic pH typical of the tumor microenvironment and endocomes of tumor cells.³⁵  CO₂-generating pH-responsive DDS represents an interesting strategy enabling the intracellular drug release in lysosomes. In a study by Liu and coworkers, the bicarbonate ion was encapsulated in the liposomes composed of HSPC, Chol, and mPEG₂₀₀₀–DSPE. In an acidic medium (pH 5.0), the liposomes released CO₂ gas with the formation of carbonic acid and corresponding hydronium ions, which led to fast drug release from the liposomes. This allowed the DDS to overcome the resistance of breast cancer cells (MCF-7R) to doxorubicin.³⁶ 

An interesting approach to pH-mediated drug delivery suggests the use of charge-conversional or charge-reversal DDS,^37,38  which are negatively charged at neutral pH and convert their surface charge into a positive one in response to the lowered pH. Hybrid calcium phosphate (CaP)/charge-conversional polymer nanopreparations were loaded with siRNA against VEGF and tested against pancreatic cancer showing high gene silencing efficiency in a BxPC3 tumor model.³⁸  These DDS contained a block copolymer of PEG and the endosomolytic polyanionic poly(N′-{N′-[(N-cis-aconityl)-2-aminoethyl]-2-aminoethyl}aspartamide) (PAsp), which is stable at pH 7.4, but degrades at pH 5.5 inside endosomes or lysosomes due to cleavage of the cis-aconitic amide bonds, reverting back to the parent polycation.

A pH-sensitive micellar DDS for the delivery of plasmid DNA was engineered by the conjugation of phospholipids, such as phosphatidyl ethanolamine (PE), to low molecular weight polyethylene imine (PEI), and such PEI–PE conjugates effectively complexed DNA.³⁹  The PEI–PE/DNA complexes were mixed with low pH-degradable PEG–hydrazone–PE co-polymer and formed DDS with high transfection efficacy at acidic pH values due to PEG detachment and efficient uptake of remaining positively charged DNA complexes by cells.

Another strategy is the use of detachable functions or coatings of DDS, such as DDS with a protective polyethylene glycol (PEG) cover. Detachment of the PEG coat upon exposure to acidified pH could be very important, since the PEG coat may interfere with DDS interaction with the cell and/or retard the release of the encapsulated drug/DNA. pH-sensitive links (diorto esters, vinyl esters, cystein-cleavable lipopolymers, double esters, and hydrazones) can be used to attach PEG to the surface of DDS, which degrade fast at lowered pH eliminating the PEG cover and allowing for efficient intracellular delivery.⁴⁰ 

See some examples of pH-sensitive bonds/linkers are shown in Figure 1.2.

Glutathione (GSH) is a crucial tripeptide generated in the cell cytoplasm, which acts as the main reducing agent in cells facilitating the thiol-disulfide exchange reaction.^41,42  The intracellular concentration (2–10 mM) is especially prominent in certain regions, such as the cytosol, mitochondrion, and nucleus and is dramatically higher than in the blood (about 2–20 µM), where the disulfide exchange reactions are minimal.^43,44  The GSH/glutathione disulfide (GSSG) redox couple imparts the cells with their oxidative property and is kept under a reductive condition by NADPH and glutathione reductase. Other redox couples also contribute towards the intracellular levels of GSH, namely NADH/NAD⁺, NADPH/NADP⁺ and thioredoxinred/thioredoxinox. In addition to the cytosolic reducing conditions the endosome and the lysosome also have a redox potential that can be used to promote endosomal and subsequent lysosomal escape.⁴⁵  The redox potential in the endosome is modulated by a specific reducing enzyme, gamma interferon-inducible lysosomal thiol reductase (GILT) in the co-presence of reducing agents such as cysteine and not GSH, while the lysosome’s reductive environment is controlled by high amounts of cysteine-like thiols.^45,46  Interestingly, in a tumor mass, the GSH concentration is much higher than in normal tissues, and this difference makes GSH a promising candidate stimulus for redox-mediated drug delivery.

The reduction of the disulfide bonds in redox-sensitive DDS results in a rapid disruption of the nanocarrier, as well as an increase in the cytotoxic activity of therapeutics, first of all, in cancer.^42,46–50  The rate and efficacy of the reduction depends on many factors including the exact chemical structure of the bond. Poly(l-lysine)-based DNA-loaded DDS were developed by conjugating poly(l-lysine) through an N-terminal cysteine to a PEG chain by the disulfide bond or thioether bond. The nanoparticles containing the hindered disulfides (thioester) were not cleaved in the extra-cellular space of the two different cancer cell lines (HeLa and HuH7), while the nanoparticles with the unhindered bonds (disulfide) underwent the cleavage. The rate of dePEGylation was shown to be cell line dependent, suggesting different amounts of thiols are secreted from different cell types.⁵¹ 

Redox-responsive polyphosphate nanosized assemblies based on amphiphilic hyperbranched multiarm copolyphosphates (HPHSEP-star-PEPx) with redox responsive backbone were prepared, and the GSH-mediated intracellular drug delivery was investigated in the HeLa cells. The results suggested that DOX-loaded HPHSEP-star-PEPx micelles show higher cell proliferation inhibition against glutathione monoester pretreated HeLa cells compared to the nonpretreated ones.⁵²  A redox-responsive micellar nanopreparation assembled from the single disulfide bond-bridged block polymer of poly(ε-caprolactone) and poly(ethyl ethylene phosphate) (PCL–SS–PEEP) achieved more drug accumulation and retention in MCF-7/ADR breast cancer cells. The system rapidly released its DOX payload in response to the intracellular reductive environment and also significantly enhanced the cytotoxicity of doxorubicin to MCF-7/ADR cancer cells.⁵³ 

Cho and coworkers prepared redox-sensitive polymeric nanoparticles from a monomer containing TMBQ (trimethyl-locked benzoquinone) as the redox-sensitive group. TMBQ is a substance that forms a lactone via intramolecular cyclization by a two-electron reduction, either chemically or enzymatically. Paclitaxel was loaded into these polymeric nanoparticles and a triggered release of paclitaxel was observed by sodium dithionite-mediated reduction.⁵⁴ 

Redox-responsive DDS have been used widely for intracellular delivery of siRNA and genes. Conjugation of anti-GFP siRNA to the cationic lipid DOPE via a disulfide bond resulted in a product, which could be incorporate into PEG₂₀₀₀–PE micelles. The resultant nanosized micelles effectively protected siRNA against nuclease-mediated degradation and released siRNA in the presence of 10 mM GSH; a 30-fold higher GFP downregulation was obtained compared to free siRNA.⁵⁵  Furthermore, survivin siRNA was conjugated to phosphothioethanol (PE) via a disulfide bond and the resulting siRNA–S–S–PE conjugate was incorporated into PEG₂₀₀₀–PE micelles. A significant decrease of cell viability and a down-regulation of survivin protein levels were achieved after the treatment with survivin siRNA micelles in several cancer cell lines. The down-regulation of survivin provided a significant sensitization of the cells to paclitaxel in both sensitive and resistant cancer cell lines.⁵⁶ 

Bioreducible cationic lipids or polymers containing disulfide bonds, such as cholesterol disulfide cationic lipids,⁵⁷  branched poly(disulfide amine),⁵⁸  hyperbranched poly(amido amine) (PAMAM),^59,60  and disulfide based polyethyleneimine (SS–PEI),⁶¹  have been synthesized. Redox-sensitive micelles containing the above or similar component have been successfully used for the delivery of genes, anti-sense oligonucleotides, and siRNA.^62–64  Bioreducible, non-viral carriers were suggested for the intracellular delivery of siRNA against human telomerase reverse transcriptase.⁶¹  The authors used a biodegradable PEI of ∼800 Da, containing multiple disulfide bonds (SS–PEI), which effectively binds siRNA to form nano-sized positively charged complexes that easily destabilize and release siRNA in a reducing environment both in vitro and in vivo.

In addition to systems that respond to a reductive environment for the delivery of cargo, nanopreparations that are sensitive to hypoxic conditions (such as in certain tumor areas and in infarct zones) are also being developed. Perche et al. incorporated azobenzene as a hypoxia-responsive bioreductive linker between PEG and PEI in lipid–PEI–PEG copolymer micelles loaded with siRNA for hypoxia-specific detachment of PEG and effective uptake of remaining positively charged siRNA complexes with PEI by cells.⁶⁵ 

Some examples of redox-sensitive bonds/linkers are shown in Figure 1.3.

Many efforts have been made to understand the molecular events occurring during carcinogenesis and the signaling pathways participating in cancer progression. The interactions between the tumor cells and the tumor microenvironment play a pivotal role in this process. For accelerated growth and proliferation, cancer cells overexpress certain molecules, including proteins with enzymatic activity, which can be considered as cancer markers or local stimuli making a tumor microenvironment different from a normal one. Thus, DDDS is sensitive to the cation of locally overexpressed enzymes typical for many cancers and may represent a promising strategy for drug delivery to tumors.⁶⁶ 

Matrix metalloproteinases (MMPs), are the main mediators of the changes in the microenvironment which exist during cancer progression. MMP activity, especially MMP-2 and MMP-9, contributes to tumor growth and differentiation, the multistep processes of tumour angiogenesis, invasion and metastasis including proteolytic degradation of the extracellular matrix (ECM), alteration of the cell–cell and cell–ECM interaction, migration and angiogenesis.^67–70  The alterations in certain local enzyme expression and composition (such a MMPs) can be not only considered as biomarkers for cancer diagnosis and prognosis, but also provide an opportunity to design DDS that are capable of releasing their drug loads via enzyme-triggered mechanisms at the tumor site. Thus, paclitaxel was conjugated to the octapeptide Gly-Pro-Leu-Gly-Ile-Ala-Gly-Gln known to be selectively and efficiently hydrolyzed by the MMP2 yielding the PTX–AcG conjugate. As a result, PTX concentrations needed to arrest the cell cycle in the G2/M phase were 100-fold lower when PTX–AcG was used compated to the free drug. In vivo, tumor-bearing mice treated with PTX–AcG had more and larger areas of necrosis and fewer proliferating cells in tumor sections than mice treated with free PTX, confirming the activation of the conjugate by MMP2 at the tumor sites.⁷¹  In another study, a self-assembling, MMP2-sensitive, TATp-modified micellar nanopreparation composed of the paclitaxel prodrug (PEG₂₀₀₀–MMP2 clevable peptide–PTX) and two other polymers, TATp–PEG₁₀₀₀–phosphoethanolamine (PE) (a cell-penetrating enhancer) and PEG₁₀₀₀–PE (a nanocarrier building block) was formulated in an aqueous environment with high PTX loading. PTX was located in the hydrophobic “core” of the mixed polymeric micelle, and covered and protected by a hydrophilic PEG shell providing prolonged circulation and good tumor accumulation via the EPR effect. When inside the MM2-overexpressing tumor, the peptide link between the drug and PEG was cleaved, the cell-penetrating function become exposed and effectively brought the drug inside cancer cells resulting in high anticancer activity and low side effects.⁷² 

MMP-cleavable substrates (typically, cleavable peptides) have been used for the delivery of multiple drugs and imaging agents.^73–75  A functionalized gold nanoparticle (AuNP) for tumor imaging and drug delivery was recently reported, where DOX was conjugated to the AuNPs by a thiol–Au bond with a MMP-2-cleavable peptide substrate.⁷⁴  The nanoparticles were modified with PEG for prolonged circulation. The presence of a thiol–Au bond led to accelerated drug release once the particles were internalized via endocytosis. Injection of the AuNP into SCC-7 tumor-bearing mice caused rapid release of the drug, which resulted in efficient tumor growth inhibition.

Huang et al. engineered the MMP2-sensitive DDS combining the siRNA against VEGF and DOX in nanoparticles built from a dendrigraft poly-l-lysine with a cell-penetrating peptide (CPP) conjugated to PEG.⁷⁶  The CPP was masked using a covalently attached pH-sensitive peptide, with a MMP2-sensitive linker. In circulation, the masking peptide shielded the CPP and exposed it only upon exposure to lower pH conditions and elevated MMP2 levels to activate the CPP. These nanopreparations were tested on glioma models both in vitro and in vivo and promoted apoptosis and anti-angiogenesis simultaneously.

In another approach, the tetra-peptide sequence, GFLG (Gly-Phe-Leu-Gly), sensitive to cathepsin B, an enzyme known to be overexpressed by tumor cells, was inserted as a cleavable linker between DOX and PEGylated dendrimer (mPEGylated dendrimer–GFLG–DOX) for enzyme-triggered doxorubicin release in tumors. In vivo tumor growth inhibition was improved 2-fold with reduced side-effects compared to a standard DOX formulation at an equal dose.⁷⁷ 

Local hyperthermia is characteristic of cancer and inflammation, and the temperature difference between normal and pathological tissues may potentially serve as a trigger for the design of temperature-sensitive DDS. In addition, external heat sources can also be used locally to control the tissue temperature.⁷⁸  Different strategies can be applied for the external heating of the tissue, such as: dielectric heating by microwave irradiation, ohmic heating by electrode-applied high frequency currents, optical laser heating with fiberoptics, ultrasound application, interstitial laser photocoagulation, and even water bath heating.^79,80  When the pathology (tumor) site is heated, an increase in the endothelial pore size and blood flow takes place, resulting an increased extravasation of DDS.²⁴  Hyperthermia is also responsible for decreased DNA synthesis, induction of heat shock proteins, altered protein synthesis, disruption of microtubule organizing centers, changes in the expression of receptors and binding of growth factors, and ultimately, changes in the cellular morphology.⁸¹ 

To achieve a temperature-sensitive response, liposomes could be prepared with lipids, which have a specific gel-to-liquid phase transition temperature providing the maximal liposome membrane destabilization and drug release at temperatures of 40–42 °C within a 30–60 min period in clinical hyperthermia treatment protocols. Liposomal composition including DPPC/MSPC/DSPE-PEG₂₀₀₀ (90 : 10 : 4 mole ratio) is a good example of such a system. Myristoyl stearoyl phosphatidyl choline (MSPC) is a lysolipid which facilitates the rapid release of a drug at 40–42 °C while the PEG-lipid enhances liposomal circulation lifetimes and promotes their accumulation within tumors via the EPR effect.^82–84  One of the most commonly used and studied lipids for use in temperature-sensitive DDS is dipalmitoylphosphatidylcholine (DPPC), which has an ideal gel-to-liquid crystalline transition temperature (T_g) of 41 °C.^78,85–88 

Thermoresponsive polymeric materials have been used widely in formulating temperature-sensitive DDS, which have a sharp transition temperature, where they become either soluble or insoluble (the lower critical solution temperature, LCST). When the thermoresponsive component is combined to/with a hydrophilic polymer, the thermoresponsive block forms the hydrophobic core, and when the thermoresponsive component is combined with a hydrophobic polymer, this results in the formation of micelles with a thermoresponsive shell.⁸⁹  Poly(N-isopropylacrylamide) (PNIPAAm) derivatives are the most common thermosensitive polymers used in DDS, they transforms to a hydrophilic structure below LCST and a compact hydrophobic structure above LCST.⁹⁰  The thermo-sensitive amphiphilic block copolymer, P-(N,N-isopropylacrylamide-co-N-hydroxymethylacrylamide)-b-caprolactone [P-(NIPAAm-co-NHMAAm)-b-PCL] with a LCST of about 38 °C was used to prepare DOX-loaded micelles, which effectively inhibited tumor growth in nude mice.⁹¹ 

Thermo-sensitive comb-like copolymers of methoxy PEG (mPEG) blocks and hydrophobic polyacrylate (PA) backbones with thermo-sensitive PNIPAM graft chains (mPEG-b-PA-g-PNIPAM) have been used to prepare micelles loaded with camptothecin, which released the drug at temperatures above 40 °C.⁹²  Thermo-sensitive polyelectrolyte complex micelles suggested for delivery of 5-fluorouracil (5-FU) were composed of two biocompatible graft copolymers, chitosan-g-poly(N-isopropylacrylamide) (CS-g-PNIPAM) and carboxymethyl cellulose-g-poly(N-isopropylacrylamide) (CMC-g-PNIPAM).⁹³ 

Elastin-like polypeptides (ELPs) investigated as vehicles for thermo-responsive delivery^94–96  are genetically encoded biopolymers which show a phase transition temperature similar to the LCST observed with some polymers, wherein they are soluble at low temperatures, but can phase-separate into a gel-like phase above a critical transition temperature.⁹⁴ 

Magnetically-sensitive nanopreparations have been used widely in biomedical applications for enhancement of magnetic resonance imaging contrast, magnetic field-assisted radionuclide therapy, hyperthermia, and tissue-specific release of therapeutic agents. The most popular agent, iron oxide, is used as either maghemite (γ-Fe₂O₃) or magnetite (Fe₃O₄), which are converted in the required materials by precipitation of magnetite in a solution containing stabilizing agents (biocompatible polymers, such as dextran, PEG, PEI, polyvinyl alcohol) or by adsorption of polymers on the magnetic particle surface after their synthesis. Upon administration, the localization of such materials in the body for different purposes is controlled by an external magnetic field.^24,97–100 

When the size of iron oxide nanoparticles is reduced to less than 20 nm, they gain a single domain and become superparamagnetic. Superparamagnetic iron oxide nanoparticles (SPIONs) have iron oxides in the core and are coated with biocompatible polymers and can be functionalized with drugs, proteins or plasmids.¹⁰¹  Following the removal of the magnetic field, SPIONs lose their magnetization and become highly dispersed, which prevents their aggregation and recognition by the MPS. The are biocompatable and convenient as drug delivery and imaging systems.^102,103  DOX was incorporated into the polymeric coat of thermally cross-linked SPIONs (TCL-SPIONs) bearing carboxylic groups through electrostatic interactions between the positively charged DOX and the negatively charged carboxyls. The preparation controlled by the external magnetic field provided good tumor growth inhibition, and allowed effective tumor imaging in mice bearing lung carcinoma.¹⁰⁴  Anti-HER2/neu (HER, herceptin) antibody-modified, pH-sensitive drug-delivering magnetic nanoparticles (HER-DMNPs) loaded with DOX provided rapid drug release and allowed for the use of in vivo MRI in real time.¹⁰⁵ 

Vinblastin-loaded magnetic cationic liposomes (MCLs) significantly improved vascular uptake of MCLs in a murine melanoma model under the action of an external magnet, reducing tumor nodules in metastatic sites compared to the control group without the magnet.¹⁰⁶  Fe₃O₄/Au nanoparticles with an iron oxide core coated with a layer of gold were modified with PEG and loaded with DOX. The preparation provided DOX concentration in the liver exposed to the magnetic field much higher than in controls with no magnetic field.¹⁰⁷  Complexes of cationic lipids with plasmid DNA associated with SPIONs can be concentrated onto target cells using magnetic fields.¹⁰⁸  Transferrin-conjugated (for better targeting of glioma cells) magnetic silica poly(d,l-lactic-co-glycolic acid) (PLGA) nanoparticles co-loaded with DOX and PTX demonstrated an effective delivery across the blood–brain barrier.¹⁰⁹ 

MCLs loaded with SPIONs could be guided to the tumor site using an external magnet and their accumulation and biodistribution could be monitored using MRI, as well as biodistribution analysis. In melanoma-bearing mice two-fold greater accumulation of the administered MCLs in tumor masses was achieved by magnetic targeting.¹¹⁰ 

An interesting application of magnetic nanoparticles is generating magnetic hyperthermia. The action of the alternating magnetic field (AMF) of sufficient strength onto magnetic particles localized in the desired area (tumor) causes the magnetic moment inside these particles to oscillate, which is converted to heat through hysteresis losses or Néel relaxation and transferred to the surrounding environment with the efficacy depending on the strength and duration of exposure to the AMF.¹¹¹  At a temperature of 41–43 °C, the damage to normal cells is reversible, while tumor cells are killed.¹¹²  Cationic albumin-conjugated magnetite nanoparticles have also been suggested as a novel method for hyperthermia in cancer therapy.¹¹³ 

Ultrasound is a well-established clinical imaging modality¹¹⁴  with microbubbles used as the ultrasound contrast agents to enhance imaging resolution. In addition, ultrasound can be safely used for triggered drug delivery and release, by the localized destruction of DDS. With this in mind, ultrasound-responsive DDS accumulated in desired areas are made leaky by locally applied ultrasound to release incorporated drugs. Such release depends on the time of ultrasound application, its frequency and power density, and the type of ultrasound; pulsed or continuous wave ultrasound, as well as on drug lipophilicity.^115,116  Low frequency ultrasound (20–100 kHz) releases drug from DDS more effectively and can penetrate deeper into the tissue than high-frequency ultrasound (1–3 MHz),^117,118  although the latter does not allow for sharp focusing.¹¹⁷  For in vivo applications, low-frequency ultrasound is more appropriate for larger and deep-seated pathological areas (tumors), while high-frequency ultrasound is used for small and superficial lesions (tumors).^117,118  Ultrasound-triggered drug release was studied with Pluronic-based polymeric micelles.^119–123  Folate-modified P105 micelles loaded with DOX have been subjected to the action of low frequency ultrasound (70 kHz) resulting in drug release, which increased with the increase in the ultrasound power intensity of ultrasound.¹²³  High intensity focused ultrasound was used to trigger the drug release from poly(ethylene glycol)-b-poly[N-(2-hydroxypropyl) methacrylamide-lactate] (mPEG-b-p(HPMAm-Lac_n) micelles.¹²⁴ 

Ultrasound also enhances cell membrane permeabilization when applied in a high frequency pulse regime and allows for better penetration of DDS into tissues and cells, producing higher drug concentration, for example, in tumors.¹²⁵  The cavitation of microbubbles under the action of a ultrasonic field can cause transient, reversible cell membrane permeabilization.¹²⁶  Microbubbles may serve as drug carriers, as they perform the dual functions of target-specific delivery by responding to ultrasound and can deliver their payload intracellularly because of membrane permeabilization under the action of ultrasound.^127–131  Since larger microbubbles (2–10 µm) have short circulation times (minutes),^126,130  the use of nanobubbles was suggested,^130,132,133  or bubble combinations with another nano-DDS, such as liposomes.^134,135  Thus, antitumoral activity of liposome-encapsulated DOX (Caelyx®) was enhanced by ultrasound application in a mouse tumor model.¹³⁶  DOX-loaded liposomes were conjugated to the surface of microbubbles via a biotin–avidin linkage. Such complexes provided rapid cellular uptake, enhancement of DOX nuclear accumulation, and less drug efflux in the resistant cells treated by these complexes under the action of ultrasound, compared to just DOX-loaded liposomes with and without addition of verapamil or liposomes complexes with bubbles without applied ultrasound. The authors concluded that their DDS in combination with ultrasound was as an effective system to sensitize cells and overcome MDR.¹³⁷  PTX and siRNA co-loaded ultrasound-responsive nanobubbles have also been developed to treat hepatocellular carcinoma (HCC) by hetero-assembly of polymeric micelles and liposomes with nanobubbles. Such complexes have been given intravenously to nude mice bearing human HepG2 xerografts and an external low-frequency ultrasound was applied at the tumor site providing effective codelivery of the active component to the tumor. Antiapoptotic response in HepG2 cells was effectively suppressed by the co-delivered siRNA targeting an antiapoptosis gene (BCL-2 siRNA) during PTX chemotherapy.¹³⁸ 

DDS termed eLiposomes have also been described, which represent liposome nanodroplets encapsulating an emulsion (usually, pefluorinated hydrocarbons) capable of phase change, as well as a therapeutic agent.^139,140  Ultrasound causes the emulsion droplet to change from a liquid to a gas, which increases the volume/pressure within the liposome, rupturing it and provoking the drug release.

Photoactivation is an another option that can be used to trigger drug release from DDS. Light provides a very broad range of adjustable parameters (i.e. wavelength, duration, and intensity) and the light source can be controlled both spatially and temporaly to regulate drug release.¹⁴¹  Wavelengths in the range of 700–1000 nm, near-infrared light (NIR), are reported to be more suitable for biomedical applications than UV or visible light, because NIR deeply penetrates tissues with high spatial precision and shows less risk of damage to normal cells.^142,143  The photo-responsive groups or chromophores are usually encapsulated within, or conjugated to the surface of the DDS.¹⁴⁴  A variety of mechanisms can induce light-triggered release of cargoes from the DDS, including photo-isomerization, photo-cross-linking, photosensitization-induced oxidation, light-triggered reversible or irreversible switch in polarity, photo-decross-linking, or photo-degradation of the polymer backbone.^145,146  Azobenzene and spiropyran have been used as reversible photo-responsive groups in photoresponsive DDS. The azobenzene group undergoes light-induced, extremely rapid and fully reversible, E (trans)-Z (cis) isomerization,¹⁴⁷  while a reversible isomerization occurs between spiropyran and merocyanine, following UV and visible light exposure. By a different mechanism, NIR irradiation causes irreversible photocleavage of chromophores like 2-diazo-1,2-napthoquinone (DNQ), altering the polarity of the nanocarrier, as well as destabilizing its structure, followed by the release of the incorporated drugs.¹⁴²  Photopolymerization or photocrosslinking involve a polymerizable double bond that can be irradiated. After the photopolymerization of double bonds in the hydrophobic portion of the bilayer, the bilayer is disrupted by pore creation and allows the release of the encapsulated drug.^145,148,149 

Polymeric micelles and liposomes have been the two most widely utilized nanocarriers for the development of light-responsive vehicles.^150–152  However, other nanocarriers like dendrimers,¹⁵³  silica nanoparticles,¹⁵⁴  gold nanoparticles,¹⁵⁵  and nanogels¹⁵⁶  have also been employed. Liposomes composed of dioleoylphosphatidylethanolamine and 2-(hexadecyloxy)-cinnamic acid were reported, where photo-dimerization of the cinnamic acid residue destabilized the liposomal membrane and enabled the release of the loaded cargo.¹⁵⁷  Photo-sensitive DOX-loaded liposomes have been described, made of diplamitoyl phosphatidylcholine (DPPC), a photopolymerizable diacetylene phospholipid (1,2 bis(tricosa-10,12-diynoyl)-sn-glycero-3-phosphocholine), and DSPE–PEG₂₀₀₀,¹⁴⁸  the drug release from which was initiated by visible light. NIR light-sensitive polymeric micelles have been used for the delivery of doxorubicin.¹⁵⁸ 

Gold nanostructures have a high surface plasmon resonance that allows them to absorb NIR radiation and convert the photons into heat, which is either used to trigger the release of chemotherapeutics from NIR responsive DDS or for photothermal ablation (PTA), by increasing the temperature of the surrounding tissues.¹⁴⁵  Thus, DOX-loaded NIR light-sensitive liposomes have been prepared with hollow gold nanospheres attached to the liposome membrane.¹⁵⁹  Drug release was observed from the liposomes upon NIR irradiation resulting in increased cytotoxicity against cancer cells both in vitro and in vivo.

Photodynamic therapy (PDT) is a promising strategy for treating various types of cancer. In PDT, a photosensitizer (or a precursor) under the action of light reacts with oxygen in the tissue and generates reactive oxygen species (ROS) including singlet oxygen and free radicals that destroy cancer cells/damage DNA. This photophysical mechanism results in cellular toxicity and necrosis, apoptosis or autophagy.^160,161  The most commonly used photosensitizers are porphyrin derivatives, dyes such as phthalocyanins and napthalocyanins and chlorophyll-based derivatives. Many drawbacks of conventional photosensitizers (such as non-specific toxicity) have been overcome with the use of liposomal DDS,^162–164  as seen in the review by Lim et al.¹⁶⁵ 

Table 1.1 presents some examples of stimuli-sensitive DDS, which can destabilize and release their cargoes to provide better cell penetration under the action of various stimuli as described above.

Combination therapy, for example, the co-administration of two or more therapeutic agents acting through different mechanisms to achieve a therapeutic response better than with each of these agent individually, is widely used for the treatment of many diseases including cancer.¹⁸⁸  Such therapy has also been used to overcome multi-drug resistance in cancer, by combining the cancer cell anti-apoptotic defense mechanism of a modulator with chemotherapeutic agents.¹⁸⁹  A number of siRNA-based treatments that target genes involved in cancer cell survival mechanisms have also been evaluated in combination with conventional anti-cancer drugs.

As already mentioned, DDS have several favorable features and can play a significant role in the delivery of agents for combination therapy. Thus, for example, for chemically different drugs with different PK, DDS can impart a consistent and unified PK profile, biodistribution and stability, and deliver them simultaneously to the target site.¹⁹⁰  An important feature of DDS as described above, is the ability to provide a triggered/controlled release of payloads in response to various stimuli. Using such smart preparation, multiple payloads can be co-released/co-delivered within the same required site. Many of such combination preparations can include biological macromolecules such as siRNA, DNA, or antibodies, which can work effectively in combination with conventional anti-cancer agents for better cancer therapy.

The general principles of the engineering of stimuli-sensitive combination DDS are the same as those described and discussed above, so there is no need to repeat this. The following several tables (Tables 1.2–1.6) present some interesting and promising combination preparations based on stimuli-sensitive DDS.

Stimuli-responsive DDS demonstrate a better control over the temporal and spatial release of drugs compared to traditional DDS. Still, a careful understanding of stimuli-sensitivity of the designed DDS is required to predict/minimize off-target effects. Combining targeting moieties with the stimuli-sensitivity can further improve control over the distribution and localization of such systems and minimize side-effects. The future challenge will be to develop DDS responding to biomarkers present in very low concentration ranges, and engineer DDS with certain signal amplification features. Another trend is incorporating multiple functions into DDS allowing for simultaneous detection, diagnosis, and therapy of a disease using a single nanoparticle. Such DDS are garnering a great deal of interest from the scientific community, but are still largely the subject of academic research. Their clinical translation may be difficult, but definitely not impossible.