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Targeted Drug Delivery with Polymers and Magnetic Nanoparticles: Covalent and Noncovalent Approaches, Release Control, and Clinical Studies Karel Ulbrich,†,§ Kateřina Holá,‡,§ Vladimir Šubr,† Aristides Bakandritsos,‡ Jiří Tuček,‡ and Radek Zbořil*,‡ †

Institute of Macromolecular Chemistry, The Czech Academy of Sciences, v.v.i., Heyrovsky Square 2, 162 06 Prague 6, Czech Republic ‡ Regional Centre of Advanced Technologies and Materials, Department of Physical Chemistry, Faculty of Science, Palacky University, 17 Listopadu 1192/12, 771 46 Olomouc, Czech Republic ABSTRACT: Targeted delivery combined with controlled drug release has a pivotal role in the future of personalized medicine. This review covers the principles, advantages, and drawbacks of passive and active targeting based on various polymer and magnetic iron oxide nanoparticle carriers with drug attached by both covalent and noncovalent pathways. Attention is devoted to the tailored conjugation of targeting ligands (e.g., enzymes, antibodies, peptides) to drug carrier systems. Similarly, the approaches toward controlled drug release are discussed. Various polymer−drug conjugates based, for example, on polyethylene glycol (PEG), N-(2-hydroxypropyl)methacrylamide (HPMA), polymeric micelles, and nanoparticle carriers are explored with respect to absorption, distribution, metabolism, and excretion (ADME scheme) of administrated drug. Design and structure of superparamagnetic iron oxide nanoparticles (SPION) and condensed magnetic clusters are classified according to the mechanism of noncovalent drug loading involving hydrophobic and electrostatic interactions, coordination chemistry, and encapsulation in porous materials. Principles of covalent conjugation of drugs with SPIONs including thermo- and pHdegradable bonds, amide linkage, redox-cleavable bonds, and enzymatically-cleavable bonds are also thoroughly described. Finally, results of clinical trials obtained with polymeric and magnetic carriers are analyzed highlighting the potential advantages and future directions in targeted anticancer therapy.

CONTENTS 1. Introduction: Drug Targeting and Its Significance in Tumor Therapy 2. Principles of Tumor Targeting 2.1. Passive Accumulation of Macromolecular Systems in Solid Tumors 2.2. Active Drug Targeting 2.2.1. Important Targets and Targeting Ligands for Tumor Drug Delivery Systems 2.2.2. Conjugation of Targeting Ligands to Drug Carrier Systems 2.2.3. Other Systems with Active Drug Targeting 3. Synthetic Water-Soluble Polymer−Drug Conjugates for Cancer Therapy (Polymer Prodrugs) 3.1. Conjugates of Poly(ethylene glycol) (PEG) 3.1.1. PEG Conjugates with Proteins/Glycoproteins 3.1.2. PEG Conjugates with Low Molecular Weight Anticancer Drugs 3.2. Copolymers of N-(2-Hydroxypropyl)methacrylamide

© 2016 American Chemical Society

3.2.1. HPMA Copolymer−Drug Conjugates for Passive Drug Delivery 3.2.2. Synthesis and Activity of Actively Targeted HPMA Copolymer−Drug Conjugates 3.2.3. Drug-Free P(HPMA) Conjugates for Cancer Treatment 3.3. Other Polymer−Drug Conjugates 3.3.1. Polymer−Drug Conjugates with Biodegradable Polymer Chain 3.3.2. Poly(amino acids) 3.3.3. Acrylate and Vinyl Polymers and Copolymers 3.3.4. Poly(amido amine)s 3.3.5. Polyacetals 3.3.6. Poly(2-oxazolines) 4. Polymeric Micelles 4.1. Preparation of Micelles and Loading of Drugs

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DOI: 10.1021/acs.chemrev.5b00589 Chem. Rev. 2016, 116, 5338−5431

Chemical Reviews 4.2. Amphiphilic PEG-Based Copolymers as Micellar Drug Carriers 4.2.1. PEG-b-poly(amino acid) Micelles as DDS 4.2.2. PEG−Polyester Amphiphilic Copolymers and Micelles 4.2.3. PEG−Polyester Micelles as DDS 4.2.4. Polyethers (Poloxamers) 4.3. Polymer Micelles Prepared from Other Amphiphilic Block Copolymers 4.4. Thermoresponsive Micellar Systems 5. Polymer Nanoparticles 5.1. Polymer Nanoparticles for Drug Delivery 5.2. Methods for Preparing Polymer Nanoparticles 5.3. Nanoparticle DDS in Clinic and Clinical Studies 6. Magnetic Iron Oxide Nanoparticles 6.1. Advances in the Use of SPIONs for Drug Delivery 6.2. Design and Structure of Magnetically Engineered Drug Delivery Systems 6.2.1. Synthetic Approaches 6.3. Noncovalent Conjugation of Drugs with SPIONs 6.3.1. Single Magnetic Nanocrystals and Soft Clusters Thereof 6.3.2. Condensed Magnetic Clusters 6.3.3. Magnetic Particles in Matrices 6.3.4. Magnetoliposomes, Magnetopolymersomes and Magnetic Micelles 6.4. Covalent Conjugation of Drugs with SPIONs 6.4.1. Single Magnetic Nanocrystals 6.4.2. Other Nanoarchitectures with Covalent Bonds 6.5. SPIONs in the Clinic and Clinical Studies 6.5.1. SPIONs for MRI 6.5.2. SPIONs for Treatment 7. Summary and Future Outlook Author Information Corresponding Author Author Contributions Notes Biographies Acknowledgments Abbreviations References

Review

threat with the risk of causing a pandemic or when drugs become ineffective against specific viruses or bacteria due to the evolution of resistance mechanisms. Cancer is a leading cause of death in developed countries (e.g., almost 600 000 cancer deaths estimated in 2016 in the United States).1 The term “cancer” refers to uncontrolled cellular growth and multiplication resulting from a cell phenotype that produces growth signals and is insensitive to anti-growth signals. Such cells have unlimited replicative potential, evade apoptosis, induce angiogenesis, and stimulate invasion and metastasis.2 There are many types of cancer with few typical or common characteristics, so its treatment is very challenging. The efficacy of conventional chemotherapy is reduced by the nonspecific distribution and rapid clearance of many anti-cancer agents, drug resistance at the tumor and cellular level, low efficiency, and the often significant toxicity of existing anti-cancer agents when administered at higher doses. Consequently, in recent decades vast efforts have been devoted to understanding the molecular and cellular mechanisms of this disease as well as to the design of drugs for its treatment. This has prompted the exploration of new nanomedicines (drug carriers) that can overcome the major drawbacks of conventional cancer treatments by attacking disease-specific mechanisms and properties. Many kinds of drug carrier have been developed to date, including water-soluble high molecular weight (HMW) polymer carriers, polymeric nanoparticles, polymeric micelles, dendrimers, liposomes, viral nanoparticles, carbon-based systems (e.g., carbon nanotubes and graphene oxide), magnetic nanoparticles (e.g., iron oxides), and silica and gold nanoparticles (Figure 1). Recent advances in polymer chemistry and the development of novel polymerization techniques have enabled the synthesis of polymers with well-defined structures, narrow molecular weight distributions, and tunable properties.3−5 Similarly, developments in nanomaterial chemistry have

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1. INTRODUCTION: DRUG TARGETING AND ITS SIGNIFICANCE IN TUMOR THERAPY The development of effective treatments for diseases has been a major goal of the human race for over 2000 years. During this period we have greatly improved our understanding of the human body and the functions of its various cooperating components (organs, bones, muscles, blood vessels, nerves, and so on) as well as the living cells that constitute their building blocks. This understanding has led to the development of a great range of medications, both natural and synthetic, to fight against infectious diseases, inflammatory illnesses, systematic (physical or mental) disorders and/or damage, and malfunctions of the body’s components. The demand for new and more effective medicaments is particularly strong during the emergence of new diseases and/or disorders that pose a global

Figure 1. Overview of polymer, nanoparticle, and other drug delivery systems with many possibilities of drug targeting and its attachment. 5339


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suited to covalent drug conjugation (e.g., polymer−drug conjugates or single magnetic nanocrystals) and some to noncovalent approaches (e.g., polymer nanoparticles or magnetic nanoclusters). Broadly speaking, these approaches are applicable to diverse kinds of DDS for cancer treatment including magnetic nanoparticles and nanoclusters as well as polymer−drug conjugates, polymer micelles, and polymer nanoparticles. Other polymer-based drug delivery systems such as nonmodified and polymer-modified liposomes (one of a few DDS to have been extensively clinically tested and approved for clinical use12), nanocapsules, and other nanomedicines are not covered in this review due to space limitations, although we recognize their great importance for the treatment of cancer and other conditions. This review’s aim is to comprehensively cover recent advances in the fields of polymeric and magnetic DDS with covalently or noncovalently attached drugs for cancer treatment. Section 2 discusses some general requirements that must be satisfied by any DDS intended to target tumors actively or passively, stressing the unique properties and physiological attributes of tumor regions. Section 3 deals with covalent polymer−drug conjugates including extensively explored watersoluble polymers such as poly(ethylene glycol), N-(2hydroxypropyl)methacrylamide, and others. The most important results obtained with such polymer-bound prodrugs in in vivo studies and clinical trials are also discussed. Section 4 describes polymeric micelle-based DDS with covalently attached drugs as well as polymeric micelles with encapsulated medicine payloads, while section 5 is devoted to polymeric nanoparticles bearing drugs that are attached only via noncovalent interactions. Section 6 discusses iron oxide nanoparticle-based DDS using both the covalent and the noncovalent methods of drug loading. Clinical results that have been obtained with the relevant DDS are emphasized in each section. Finally, the Summary and Future Outlook section summarizes the potential advantages of DDS in cancer treatment.

produced nanoparticle carriers with narrow size distributions and controllable physicochemical properties that can be exploited for various purposes such as monitoring the treatment’s effects or enhancing its efficiency. Together with advances in cellular and molecular biology, these developments offer opportunities to create sophisticated selectively targeted nanomedicines based on carriers and carrier conjugates with various biologically active molecules (BAM) such as drugs, genes, enzymes and other proteins, or nucleotides. Such agents could potentially form the basis of very specific, safe, and efficient cancer treatments. Another important benefit of using polymers and nanoparticles as drug carriers derives from their ability to increase the water solubility of hydrophobic drugs, extend the circulation of drugs in the blood, and suppress or eliminate fast renal excretion. Together, the use of drug carriers dramatically increases organ or cell-specific drug accumulation6,7 and opens up the possibility of controlled activation (i.e., release) of the delivered drug where the therapeutic effect is required, e.g., in tumors and tumor cells. Selective activation in this way could prevent the drug’s toxicity from affecting normal tissues and cells, mitigating or eliminating any harmful side effects it might otherwise have. This review discusses recent achievements in the development and characterization of drug delivery systems (DDS) based on polymers and magnetic nanoparticles. In particular, we focus on DDS that exploit interactions between ligands and specific receptors localized on cell or organelle membranes to facilitate site-specific drug delivery. In vivo experiments have shown that such DDS have great potential in human medicine, therapy, and/or the diagnosis of cancer diseases. Considerable attention is paid to chemical aspects of the design and synthesis of DDS that facilitate the dissolution and delivery of hydrophobic anticancer drugs that are otherwise insoluble in water and bodily fluids. In particular, we discuss covalent and noncovalent approaches for attaching drugs to nanocarriers and the profound ways in which such attachment can change the drugs’ pharmacodynamic behavior. Any DDS, whether polymeric or based on magnetic iron oxide nanoparticles (MIONs), must satisfy a number of common criteria.8 Specifically, a DDS must (i) avoid nonspecific interactions with the body or the induction of adverse reactions and avoid capture by cells of the reticulo-endothelial system (RES), (ii) facilitate the BAM’s transport to the site of action (organ, tissue, cell, or organelle) from the site of administration in a high yield while keeping the BAM in a safe (i.e., inactive) state during transport, (iii) protect the BAM from detrimental effects during transport (e.g., enzymatic degradation or hydrolysis) in the body, (iv) release effective quantities of the active BAM in or around the target, ideally in a controlled fashion such that a desired tissue/cell concentration is achieved, (v) enable elimination of all components of the DDS from an organism after its function as a carrier has been fulfilled. There are many ways in which these requirements could be met, and each individual DDS offers its own set of specific solutions. Some of these are common to many (or even all) systems, but others are unique and depend on the details of the carrier’s structure and architecture.9−11 Some DDS are better

2. PRINCIPLES OF TUMOR TARGETING Because of the great phenotypic diversity of malignant cells and tumors there are few generally applicable methods for targeting tumors and tumor cells (or their organelles). The most important such approach is based on the fact that many tumors and vascularized solid tumors as well as some vascularized metastatic tumor nodules exhibit an enhanced permeability and retention (EPR) effect13 that can be exploited for so-called “passive targeting” of antitumor agents. This effect occurs because many solid tumors have a leaky vasculature and absent or impaired lymphatic drainage, which causes the accumulation of high molecular weight molecules (polymers) as well as small particles of diameter ∼20−500 nm within the tumor tissue. It should be noted that the exploitation of this effect for “drug targeting” is not the same thing as true specific receptordependent targeting; the macromolecules or macromolecular systems captured in tumors due to the EPR effect do not act in a hugely selective fashion. Consequently, the term “targeting” should strictly only be used in reference to “active targeting” approaches based on specific interactions between a membranebound receptor and a complementary ligand that is bound to the surface of the DDS, where it serves as a targeting moiety. The main advantage gained by exploiting the EPR effect for drug delivery to solid tumors is its universality, which enables the DDS to be used in the treatment of diverse tumors. 5340


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RES uptake as well as a suitable surface charge and adequate stability in circulation. It is important to note that many tumors have necrotic and hypoxic regions, particularly in the interior of the tumor mass, as well as living cells. These regions and the cancer stem cells are not readily accessible to DDS, which is an important point to bear in mind when designing DDS for passive tumor targeting. The scope for overcoming these challenges using polymeric and magnetic DDS and modified variants of these systems are discussed in the sections on the corresponding DDS types. The triggered release of drugs from a carrier in response to a stimulus encountered on entry into the diseased tissue (e.g., a change in pH, redox state, or temperature) is sometimes also regarded as a form of passive drug targeting. However, this review focuses on the site-specific delivery of the drug and its delivery system as a whole; the issue of specific drug release is beyond our scope here. The passive targeting process for macromolecular DDS (including soluble conjugates, nanoparticles, micelles, and liposomes) and inorganic nanoparticles is outlined in Figure 2: the DDS circulates in the bloodstream, undergoes

However, there are significant challenges associated with this approach in practice, as discussed in the next section. 2.1. Passive Accumulation of Macromolecular Systems in Solid Tumors

Because healthy liver tissue is extensively fenestrated, many nanocolloids (NCs) with certain physicochemical properties are nonspecifically accumulated in the liver and taken up by cells of the reticulo-endothelial system (RES), namely, macrophages and mononuclear phagocytes. In addition, NCs may be taken up by inflammatory tissues. These processes are fast and do not require prolonged circulation of the NCs. The extravasation of NCs and macromolecules from the circulation into the interstitial spaces of healthy tissues is generally very limited. However, the EPR effect, caused by the leaky vasculature and dysfunctional lymphatic drainage of many solid tumors, results in the efficient extravasation of longcirculating NCs from the tumor vasculature and their retention in the tumor interstitium.13,14 It should be noted that NCs will only accumulate efficiently in tumor tissues if they satisfy certain criteria. Primarily, they must be long circulating (i.e., not quickly removed by the RES) and small enough to pass through gaps between the endothelial cells of the tumor vasculature but large enough to avoid escape from the tumor tissue. In practice, this usually means that their size (or equivalently, the diameter of the macromolecule coil) should be between 20 and 500 nm.15−17 In addition, the EPR effect is sensitive to a number of other physicochemical properties of the NCs including their biocompatibility,16 hydrophilicity/hydrophobicity, and surface charge. These properties have important effects on the (nonspecific) interactions between the NCs and the cells of the RES during transport in blood and also on their extravasation and interactions with various entities inside the tumor.18 Moreover, colloidal carriers that have not been modified with hydrophilic polymers are rapidly taken up by phagocytosis and the RES. To achieve prolonged circulation and protection from opsonization, the NC surface must be modified, e.g., with dextrans or pullulans, by “pegylation”19 or by modification with other hydrophilic polymers such as poly(HPMA)20−22 to form a “stealth” system. This is not a problem for highly hydrophilic and water-soluble drug carrier systems, which are naturally long circulating.23 Accumulation in tumors via the EPR effect has been documented for various NCs including soluble macromolecules such as plasma proteins, PEGylated proteins, and soluble polymer−drug conjugates, as well as for various organic and inorganic NC systems such as polymer micelles, nanoparticles, nanocapsules, and liposomes.24 All of these NCs can be used as vehicles for the tumor-specific delivery of BAMs (drugs) by encapsulating or covalently attaching the low molecular weight (LMW) molecule to the vehicle. The vascular permeation of NCs can be enhanced by a variety of factors,17 which may be coadministered with the DDS or produced endogenously by the tumor tissue or in its vicinity. Important endogenous factors that can enhance vascular permeation include bradykinin, VEGF, TNF-α, NO, and heme oxygenase.17 In addition, the carrier’s retention in the tumor tissue can be enhanced by increasing its hydrophobicity. This can be exploited to manipulate the accumulation of the delivered drug in the tumor. Because the EPR effect is time dependent, a DDS must remain in circulation for at least several hours to be efficient. As noted above, this will only be the case if the DDS is of an appropriate size and has sufficient biocompatibility to avoid

Figure 2. Schematic representation of accumulation of macromolecules in solid tumors due to the EPR effect (red dots, LMW drug; blue line, macromolecule; nanoparticles are supposed to accumulate in the same way as macromolecules).

preferential extravasation in tumor tissues due to their leaky vasculature, and accumulates in the tumor due to the EPR effect. The drug is then activated (released) in the tumor by pH-dependent hydrolysis induced by the lower pH of the tumor tissue or inside the tumor cells by endosome-mediated hydrolysis, enzymolysis (hydrolysis by enzymes in secondary lysosomes), or redox processes occurring in the cytosol. The concentration of the released drug can be controlled by the rate of degradation of the spacer in covalently conjugated drugs or by biodegradation and dissolution of the polymer matrix. 2.2. Active Drug Targeting

All of the DDS types mentioned above can be actively targeted using specific ligands, targeting moieties attached to the carrier macromolecules or decorating the surface of NPs. Active targeting can be achieved at different levels depending on the extent of penetration; it may occur at the organ, cell, or subcellular level. It is important to note that in solid tumors, even active targeting processes begin with the passive accumulation of the DDS in the tumor tissue, so any actively targeted carrier must satisfy the basic requirements outlined for passively targeted systems: the polymer/nanoparticle should be biocompatible, and the system must be stable in circulation, 5341


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enable the transport of various chemical species across the membrane, act as enzymes, and control cell adhesion. Membrane proteins have domains that are embedded into the cell membrane (in the case of transmembrane proteins, these domains extend across the membrane’s width) as well as surface-exposed domains that are accessible to ligands and can potentially constitute tumor-specific antigens. NC-based DDS consist of a carrier, a covalently bound or entrapped drug, and a targeting ligand. A wide range of ligands can be used to bind to cell membrane receptors and promote internalization of the NC system into the cell. For example, galactose has been used to promote binding to parenchymal liver cells,31,32 mannose and fucose for Kupfer cells,7 and folic acid for folate receptorexpressing33−35 cancer cells. The most specific ligands are antibodies, preferably IgG-type antibodies with an Mw of around 150 kDa or antibody scFv fragments. Some representative tumor cell membrane receptors and the corresponding receptor-binding ligands are listed in Table 1;

long circulating, and of a size that permits efficient extravasation and accumulation in tumors. In principle, it is not possible to use a single DDS with a single active targeting mechanism against a large group of malignancies with diverse origins or a broad variety of tumor cells. However, actively targeted DDS can be very selective and effective against solid tumors as well as leukemias and other circulating tumor cells that are less sensitive to passively targeted DDS. In addition, actively targeted systems are often more effective in general than passively targeted alternatives. The design of actively targeted DDS requires detailed information on surface receptors (proteins, glycoproteins, and lipoproteins) that are more abundant on tumor cells than their healthy counterparts. It is also necessary to identify ligands that bind strongly to these receptors, such as saccharides, lectins, antibodies and antibody fragments, peptides, enzymes, or enzyme inhibitors. By conjugating the DDS with such ligands, it is possible to efficiently guide the delivery vehicle to the desired site of drug action and avoid nonspecific delivery to healthy cells and tissues.25,26 The general concept of active DDS targeting is outlined in Figure 3.

Table 1. Selected Cancer Cell-Specific Antigens and the Corresponding Ligands cancer cell membrane receptor Asialoglycoprotein receptor Kupfer cells biotin receptor (MCF7 cells) folate receptor Asialoglycoprotein receptor Integrins (αvβ5, αvβ3) MSH-R EGF-R CD20 (B cells) CD52 (lymphocytes) VEGF-A CD30 HER2/Neu transferrin receptor epidermal growth factor receptor

Figure 3. Active targeting of DDS to tumor cells and subsequent mechanisms of intracellular drug release. After the extravasation and potential accumulation of DDS in the tumor due to the EPR effect, the targeting moiety selectively binds the DDS to a specific receptor on the cancer cell membrane. The conjugate then enters the cell via pinocytosis (endocytosis), and the drug is released from the system after carrier degradation (in the case of nanoparticles), disassembly (micelles) or pH-sensitive hydrolysis in primary endosomes, redox reactions in the cytoplasm, or by enzymolysis in secondary lysosomes for systems with covalently bound drugs.

cell or receptor-specific ligand

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glucose mannose, fucose biotin (Vitamin B7)

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folic acid (Vitamin B9) galactose, galactosamine RGD, c(RGDfK) melanocyte-stimulating hormone (MSH) EGF, Erbitux, GE11 peptide Rituximab Alemtuzumab Bevacizumab Brentuximab Trastuzumab Transferrin scFv anti-RGDR

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many new cell membrane receptors and corresponding receptor-binding ligands are identified every year. The suitability of a given ligand as a targeting moiety for a DDS is not determined solely by its specificity for the desired receptor: the ligand’s cost and its availability in sufficient quantities for the synthesis and in vivo evaluation of the DDS are also important to consider when attempting to design and develop a practical drug carrier. In this context, it is noteworthy that recent developments in the large-scale production of a broad range of monoclonal antibodies (mAb) have increased their utilization as targeting moieties for site-specific therapies and diagnostics.36,37 Targeting strategies for nanomedicines designed for cancer treatment were discussed at length in a recent review.18 Established targeting strategies include passive targeting and active targeting methods based on the use of monoclonal antibodies (mAb) or specific peptides, but it is also possible to use integrin-mediated targeting, aptamer-mediated targeting, folate receptor targeting, transferrin receptor targeting, or magnetic nanoparticle drug targeting. In addition, stimulus-sensitive approaches such as acid-triggered, light-

2.2.1. Important Targets and Targeting Ligands for Tumor Drug Delivery Systems. Currently used chemotherapeutic agents have a wide range of mechanisms of action, including intercalation (doxorubicin), DNA alkylation (nitrogen mustards, cisplatin), microtubule inhibition (taxanes), and the inhibition of metabolic processes (gemcitabine)27 or angiogenesis (endostatin, TNP-470).28−30 The targets of polymeric and inorganic DDS are dictated by the mechanism of action of the drug they are used to transport. These targets may be specific for particular organs (for example, when seeking to treat lung, brain, or colon cancer), for solid tumors, or for tumor vasculature. However, in most cases the actively targeted DDS will be targeted at individual tumor cells and their organelles. The selection of an appropriate targeting moiety should be guided by detailed information on the membrane proteins of the relevant tumor cells. The cell membrane proteins perform a variety of functions: among other things, they serve as receptors for specific molecules, 5342


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Figure 4. Some representative reactions used to covalently conjugate a targeting ligand (T) with a polymer carrier.

2.2.2.1. Conjugation of Antibodies, Enzymes, and Peptides with Polymer Carriers. Antibodies (Ab) are glycoproteins that bind very specifically to well-defined antigens such as the surface-exposed domains of membrane proteins. As such, they can be very effective targeting moieties for the receptormediated targeting of DDS to tumors or tumor cells. IgG Abs (Mw ≈ 150 kDa) consist of two Fab polypeptide antigenbinding fragments and an FC fragment with a saccharide content of 3−15%. The individual peptide chains (two light and two heavy chains) are interconnected by disulfide bridges, and each IgG molecule has two antigen-binding sites situated at the amino ends of the Fab fragments. An IgG Ab has approximately 73 primary amino groups64 that could potentially be used for conjugation with polymer carriers. For example, if the polymer carrier bears carboxylic acid functional groups, these can be activated (e.g., by conversion into reactive esters via the carbodiimide method) and then aminolyzed by the Ab amino groups to form an Ab-targeted conjugate with the Ab bound via a stable amide linkage.65,66 Unfortunately, because the Ab is multivalent (as is the polymer carrier in most cases), the product is likely to be branched with a broad molecular weight distribution. Moreover, the Ab binding site contains fairly reactive primary amino groups that may participate in the aminolytic reaction; if this happens it will significantly reduce

triggered, ultrasound-mediated, and enzyme-mediated drug release can also be regarded as targeting strategies.7 The following sections describe the most important strategies and methods for conjugating targeting ligands with polymer carriers. 2.2.2. Conjugation of Targeting Ligands to Drug Carrier Systems. Strategies and reactions for the covalent conjugation of targeting ligands with polymer NC carriers must be selected to avoid impairing the functionality of the targeting moiety or reducing its ability to bind to the antigen (i.e., the cell surface receptor). Methods of conjugation include standard reactions from peptide chemistry such as aminolysis, catalytic acylation (e.g., the carbodiimide method), addition of amino or sulfhydryl groups to double bonds, and disulfide exchange reactions.62,63 “Click” chemistry can also be used to connect a polymer carrier to a ligand. Some illustrative conjugation reactions are shown in Figure 4. As an alternative to covalent conjugation, methods based on the self-assembly of oligopeptides (coiled-coil peptides) or the formation of stable complexes (e.g., biotin-avidin) can be used to decorate polymer carriers’ surfaces with targeting ligands. These methods are most useful when the ligand is a sensitive macromolecule such as an enzyme or antibody. 5343


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Figure 5. Antibody modification for binding reactions with polymer precursors. (a) Protection of the most reactive Ab amino groups with DMMA, leaving the remaining amino groups available for conjugation reactions; (b) synthesis of a monomeric Ab fragment; (c) introduction of a sulfhydryl group by mild reduction; (d) introduction of aldehyde groups by periodate oxidation; (e) introduction of sulfhydryl groups by reaction with 1-imino2-thiolane.

cycloaddition reaction whereby an unprotected peptide bearing a terminal azide group is directly attached to a polymer or inorganic drug carrier without the need for specific protection of other reactive groups. Three implementations of the click reaction for conjugation of ligands with drug carriers were recently compared:77 (1) click chemistry in aqueous and organic solvents using a Cu(I) catalyst, (2) click reactions using a catalytic ruthenium complex in DMF, and (3) metal-free click chemistry based on a dibenzocyclooctyne (DBCO) and azides. The catalyst-free final option appears best suited for the synthesis of conjugates to be used in vivo. However, each one has its own advantages and drawbacks, and care should be taken to select a method that is appropriate for both reactants (i.e., the DDS and the peptide) before they are modified and conjugated. All of the reactions described above for Abs (aminolysis, monomeric ligands, various acylation reactions, and click chemistry) can also be used to conjugate DDS with proteins or specific peptides. 2.2.2.2. Noncovalent Attachment of Targeting Moieties to Polymer Carriers. Single-chain fragment variable (scFv) antigen-binding fusion proteins derived from monoclonal antibodies (mAb) are very promising candidates for the targeting of polymer DDS. They are prepared using genetic engineering methods, and their synthesis involves the incorporation of specific spacers that enable their conjugation with polymer carriers. Surface-modified polymeric DDS bearing scFv proteins have already been designed and used for tumor cell-specific gene delivery. For example, a universal bungarotoxin/bungarotoxin-binding peptide linker78,79 was used in a gene delivery system for prostate cancer therapy.80 Here the linker was formed by simple mixing of the α-bungarotoxin peptide-modified scFv fragment of an anti-PSMA (prostatespecific membrane antigen) Ab and bungarotoxin binding

the Ab’s antigen-binding activity in most cases. Cross-linking and reductions in Ab-binding activity can be minimized by protecting the majority of the reactive amino groups (or at least the most reactive ones, especially those in the antigen-binding site), for instance, by reaction with dimethylmaleic anhydride65 (DMMA), to prevent their participation in the aminolytic conjugation reaction. The protecting groups can then be removed from the final conjugate by simple mild acidification of the solution at the end of conjugation reaction.65,67 Branching during the conjugation reaction can be avoided by using a polymer carrier with only one reactive group situated at the polymer chain end (a semitelechelic polymer). This strategy produces a star-like conjugate structure with a relatively narrow molecular weight distribution.65,66 A more specific conjugation process that cannot affect the Ab-binding site involves periodate oxidation of saccharide units in the FC part of the Ab molecule to produce aldehyde groups that can be reacted with primary amino or hydrazide groups in the carrier to produce an azomethine or hydrazone group, which can then be stabilized by mild reduction with cyanoborohydride.68−70 Alternatively, a sulfhydryl group introduced into the Ab structure by mild reduction or reaction with 2-iminothiolane can be employed for conjugation with a polymer carrier bearing vinyl, dithiopyridyl, or maleimidyl groups by addition or an exchange reaction.71 Another strategy that has been used to prepare soluble polymer drug conjugates with a monomeric Fab′ antibody fragment involves functionalization of the Fab′ fragment with a methacryloyl group followed by direct copolymerization of the macromonomeric methacryloylated Fab′ fragment with other hydrophilic monomers (Figure 5).72−74 Click chemistry is another powerful and chemoselective method for conjugating DDS with peptide targeting moieties.75,76 This method is based on an azide−alkyne 5344


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specific binding activity in vitro.83 A detailed study on selfassembly, coiled-coil heterodimer formation, and the use of such heterodimers as noncovalent spacers in the synthesis of polymer−peptide conjugates was recently published.84 In conclusion, the use of antibodies, their fragments, and other protein-based targeting moieties for the synthesis, further development, and clinical application of actively targeted DDS still needs to overcome certain hindrances preventing the targeted DDS large-scale synthesis and adoption for cancer treatment. It is not only the number of suitable, available, and approved antibodies for human treatment (increasing every year), cost of these molecules (decreasing every year), and cost of their conjugates but, as already mentioned earlier, also transition from preclinical animal tests to clinical evaluation in humans which still meets difficulties with development of proper animal model systems enabling serious prediction of activity and safety in humans. Among other drawbacks remaining to be solved a drop in binding activity to antigen during conjugation reaction, potential immunogenicity of the proteins, and their side effects should be also mentioned. Nevertheless, high selectivity, superior anticancer activity at low doses, and promising preliminary data from treatment of human patients85 are the driving forces for further development in this area. Conjugation reactions with targeting moieties and individual drugs are discussed at greater length in the sections dealing with specific DDS types. 2.2.3. Other Systems with Active Drug Targeting. More than 20 mAb are currently approved for clinical use, and several others are under development. They act through antibody-dependent cellular and complement-activation-dependent cytotoxicity and play crucial roles in the management of diverse cancers. Their conjugates with anticancer drugs or, more importantly, with radionuclides represent a promising group of immunotherapeutics and radioimmunotherapeutics for the efficient and specific treatment or diagnosis of cancer. In these therapeutics, α emitters such as 213Bi, 223Ra, 149Tb, and 230 U or β emitters 131I, 90Y, or 60Co capable of damaging DNA are combined with mAbs that have detrimental effects on cancer cells by themselves.7 Numerous clinical trials are being conducted to evaluate the safety and antineoplastic potential of antibody-targeted chemotherapeutics based on established drugs (e.g., cisplatin, oxaliplatin, taxols, or gemcitabine), preferably in combination with an anti-VEGF mAb. Trials are also being conducted with various radioimmunotherapeutics based on radioisotopes of I, Bi, In, and Lu among others. In addition, targeted treatments that combine chemotherapy and radiotherapy are being developed based on multicomponent

peptide-modified viral NPs. This resulted in the noncovalent conjugation of the scFv with the viral gene delivery vector. The same strategy was used to create an anti-PSMA scFv-targeted water-soluble DOX conjugate to treat prostate tumors in mice.81 A linker prepared by the self-assembly of coiled-coil peptides was also used for the noncovalent conjugation of a targeting scFv (derived from an M75 mAb) with a HPMA copolymer− drug carrier.82 The scFv was modified with a peptide tetramer (VAALKEK)4 without affecting its binding site, and the HPMA copolymer−DOX was modified with a complementary (VAALEKE)4 peptide tetramer (Figure 6). Self-assembly of

Figure 6. Structure of a polymer−DOX conjugate in which a GFLG spacer is used to link the polymer backbone to the conjugated drug (DOX) and to a peptide spacer that is complementary to a peptide spacer linked to an scFv protein. The two peptide spacers interact to form a coiled-coil peptide spacer (depicted in ribbon form) that conjugates the polymer to the targeting scFv.

the two components in an aqueous solution yielded the coiledcoil heterodimer spacer, which was stable and thus effectively bound the scFv to the DDS, producing a conjugate that interacted selectively with the desired receptor. The same strategy was also used to target the DOX polymer conjugate to BCL1 cells using an scFv derived from a B1 mAb as the targeting ligand. The resulting conjugate exhibited effective and

Figure 7. Basic principles of the ADEPT system. The conjugate of a target cell-surface receptor-specific Ab with a selected enzyme is administered in the first step. After the conjugate is attached on the cell surface, the prodrug is injected. The active drug is then released from the prodrug in the vicinity of the target cell by enzymolysis. 5345


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nanoparticles featuring a targeting mAb, a fluorophore for fluorescence imaging, a radionuclide, a chemotherapeutic agent, and also potentially a gene therapy agent.7 A range of nanoparticles have been tested in systems of this kind, including liposomes, albumin and dextran nanoparticles, PEG-polyester micelles, and others. These complex systems are very interesting but will require further development to overcome some problems that have been identified with their use. Antibody-directed enzyme prodrug therapy (ADEPT) represents a unique strategy based on a two-component system for the treatment of malignancies and solid tumors.86−88 The first component is a prodrug, i.e., an anticancer drug bearing an inactivating moiety; the second is tumor-specific antibody that is conjugated with an enzyme capable of releasing the active drug by cleaving the bond to the inactivating moiety. The pretargeted Ab-enzyme is allowed to accumulate in the tumor, and the prodrug is then administered, resulting in the specific targeted release of the drug within the tumor as a consequence of the enzymatic reaction (Figure 7). To be functional, the system must fulfil some basic requirements: the receptors for the Ab should be overexpressed on tumor cells but not in normal tissues, the conjugate should avoid internalization by cells, the enzyme should be specific for its prodrug substrate, and the prodrug should accumulate in the tumor, ideally by exploiting the EPR effect (which could be achieved by administering it in polymer-bound form). The Ab in this system could potentially be replaced by an Ab fragment or the corresponding scFv. Careful selection of a suitable enzyme (carboxypeptidase-A or G2, β-glucosidase, and β-lactamase may all be suitable) and substrate is essential when designing an ADEPT system. Modification of the above-mentioned strategy has resulted in the development of related systems for gene-directed enzyme prodrug therapy or virus-directed enzyme prodrug therapy, among other things. However, it seems that further development and refinement of the strategy will be required to achieve convincing performance in vivo. This category of targeted nanomedicines also arguably includes immunotoxins and immunocytokines, but a detailed description of these systems is beyond the scope of this review.

body can be easily controlled by changing the polymer’s molecular weight and architecture during polymer synthesis or by exploiting polymer chain biodegradation while the carrier is in the body. With a few exceptions, water-soluble polymer−drug conjugates (polymer prodrugs) generally consist of (i) a water-soluble polymer backbone of variable architecture that gives the conjugate biocompatibility and good solubility while ensuring its extended circulation in the body, (ii) a biologically active molecule (BAM), which is typically a LMW drug, nucleotide, or biologically active hormone or protein, and (iii) a biodegradable spacer between the BAM and polymer chain, which is usually designed to facilitate the BAM’s controlled release from the polymer carrier. In many cases there is also a fourth component: a targeting moiety that enables specific delivery of the conjugate to its target. The polymer carrier’s multivalency means that it can also be modified or decorated with additional functional molecules such as another drug (resulting in a conjugate enabling combination therapy) or radionuclide or a Gd complex or fluorescent label (to form polymer diagnostics or theranostics) among other things, as shown in Figure 8.

Figure 8. Scheme of polymer drug conjugate. The conjugate consists of polymer backbone, one or more BAM (drugs) attached by biodegradable spacer, fluorescent or radioactive label, and targeting moiety.

The discussion below deals with the design and structure of each component of the polymer carrier as well as the synthesis and physicochemical properties of the final polymer−drug conjugates. The number of polymer carriers and drug conjugates under investigation is rapidly expanding, and developments in this field have been reviewed extensively.89−96 Consequently, we focus mainly on conjugates with in vivo activity, preferably verified in preclinical and clinical trials. Two groups of macromolecules with different origins have been used during the development of drug carriers: (i) naturally occurring, often biodegradable, macromolecules such as proteins, glycoproteins, and polysaccharides and (ii) synthetic polymers. The use of natural macromolecules as drug carriers has several advantages but also disadvantages. They can be often obtained from renewable resources even on large scale, they are biodegradable and often well tolerated by the body, and in most cases their chains are uniform and well defined. On the other hand, their batch-to-batch reproducibility is rather low, they may lose biodegradability after chemical modification, their physicochemical behavior is unstable, and they may become immunogenic after conjugation with other components of the DDS.97,98 While these macromolecules represent an

3. SYNTHETIC WATER-SOLUBLE POLYMER−DRUG CONJUGATES FOR CANCER THERAPY (POLYMER PRODRUGS) DDS that combine cytotoxic chemotherapeutics with targeted water-soluble polymer drug carriers are highly promising and offer the prospect of safer cancer therapies that are more specific, efficient, and patient friendly than current options. The versatility of the polymer carrier structure and its potential for modification make it possible to tune the properties of the polymer−drug conjugate to the needs of specific treatments. Water-soluble polymers are a group of large mostly linear molecules with a long polymer chain that can accommodate multiple functional groups, which can be exploited to tailor the polymer’s properties. Moreover, the characteristics of the polymer chain may be imparted to its conjugates; consequently, the conjugates of nonsoluble drug molecules with such polymers may exhibit water solubility, resistance to degradation or undesired uptake during transport, prolonged blood circulation, and little or no toxicity or immunogenicity. In addition, the size of the polymer carrier (i.e., the diameter of the polymer coil in solution) and thus its distribution in the 5346


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Figure 9. Selected reactive groups used to conjugate PEG/mPEG chains with BAM.

with protein or drug molecules. However, rather than an individual drug carrier, PEG is often used for coating other delivery vectors (nanoparticles, viruses, liposomes), as a block copolymer for the preparation of polymer micelles, or as a spacer between a BAM and a drug carrier with a different structure.103,106,107 mPEG chains with activated end groups are used in the synthesis of some conjugates that are currently on the market or undergoing clinical trials, including conjugates with asparaginase,108 adenosine deaminase,109 interferon-α 2b or 2a,110,111 and anti-TNFα Fab112,113 or anti-VEGF114 antibodies. The modification of biologically active proteins with PEG enhances their stability, increases their circulation time, and reduces their immunogenicity and antigenicity while preserving all or most of the protein’s original activity.104 Many polymer− protein or glycoprotein conjugates have been synthesized by acylating the terminal α-amino or ε-amino groups of lysine residues in the protein molecule. However, protein conjugates with tailored properties can also be prepared by the specific modification of histidine residues;102 this approach is especially useful for the synthesis of conjugates with a hydrolytically removable PEG shield. A variety of functional groups have been used to activate mPEG for conjugation with biologically active proteins, including reactive carbonate groups such as succinimidyl succinate or carbonate, 4-nitrophenyl carbonate, and benzotriazolyl carbonate, as well as aldehyde or dichlorotriazine groups.102,115 Some important reactive groups suitable for PEG activation are shown in Figure 9. The selection of a suitable reactive group can influence the rate and yield of the conjugation reaction and also the properties and stability of the formed bond. It can also affect the protein’s surface charge. Reactions performed at higher pH (pH > 8) preferentially lead to the acylation of ε-amino groups such as those of Lys residues, whereas acylations conducted at lower pH (5.5−6.5) favor modification of α-amino groups. Some activated PEGs can slowly react with hydroxy (in Ser or Tyr residues) or imidazolyl (in His) groups to form linkages that are typically susceptible to hydrolysis,102 enabling removal of the shielding polymer under specific conditions. The covalent conjugation of proteins with PEG need not be achieved via amino groups; other groups that may be present in polypeptide chains such as thiol or carboxylic groups can also be exploited. In particular, reduced proteins and glycoproteins

interesting and significant group of drug carriers, their use for this purpose has recently been reviewed at length,99−101 so we focus here on water-soluble drug carriers based on synthetic polymer materials. In principle, biodegradable or nondegradable synthetic polymers and copolymers can be prepared with a broad variety of chemical structures and polymer chain architectures, making it possible to tailor the polymer prodrug’s properties to suit its intended purpose and anticipated mechanism of action. To be effective in vivo, a water-soluble polymer−drug carrier should be biocompatible and nonimmunogenic. It should also contain functional groups that permit the attachment of a drug and other ligands, be long circulating in the bloodstream, and be readily eliminated from the body after fulfilling its task as a carrier. The final requirement means that if nondegradable polymers are to be used as drug carriers, their molecular weight must be below the renal threshold (commonly ∼50 kDa). A variety of synthetic hydrophilic polymers have been tested as drug carriers, with nondegradable derivatives of poly(ethylene glycol), or PEG, probably being the most important. 3.1. Conjugates of Poly(ethylene glycol) (PEG)

3.1.1. PEG Conjugates with Proteins/Glycoproteins. Of the known hydrophilic synthetic polymers, PEG has been most widely used as a hydrophilic drug carrier. In particular, it has been used extensively for the delivery of biologically active proteins and glycoproteins102−105 but it is also often used for delivery of LMW drugs. PEG is a highly hydrated linear polyether consisting of −CH2CH2−O− units with a narrow distribution of molecular weights and one monomethoxy (mPEG) or two hydroxy chain end groups. The nomenclature of PEG is often rather confusing, which can make it difficult to orient oneself when reading the literature. The material known as PEG is also referred to as poly(oxy ethylene) (POE) and poly(ethylene oxide) (PEO). Strictly speaking, “PEG” should be used for polymers prepared by polycondensation of the monomer ethylene glycol by removal of water and “POE” or “PEO” for polymers prepared by polymerization of ethylene oxide. However, for the sake of simplicity we use the abbreviation PEG for all polymers with two −OH groups at the chain ends or polymers with no identified structure and mPEG for polymers with one −OH and one CH3O− (methoxy) group at the polymer chain ends. Terminal −OH groups can be activated and used for subsequent conjugation 5347


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Table 2. PEG−Protein Conjugates Approved for Clinical Use116,117 tradename

protein

PEG

indication

ref

Adagen Oncaspar Pegvisomant PEG-intron Pegasys Neulasta Mircera Somavert Cimzia

adenosine deaminase asparaginase GH antagonist interferon α2b interferon α2a GCSF epoetin beta HGH antagonist anti-THF Fab

5 kDa 5 kDa 5 kDa 12 kDa 40 kDa 20 kDa 5 kDa, conj. 60 kDa 5 kDa, conj. 42−52 kDa PEG-Fab conjugate

combined immunodeficiency disease acute lymphatic leukemia excessive growth hepatitis C hepatitis C neutropenia anemia, kidney acromegalia rheumatoid arthritis, Crohn’s disease

109, 118 119 120 121 122, 123 124 125 125 125

Figure 10. Structure of a branched PEG polymer conjugated with Ara-C. Branches are formed by Asp residues.

The importance of PEG conjugates with proteins/ glycoproteins as DDS for the treatment of various diseases is demonstrated by the diversity of such conjugates that have already been accepted in clinical practice (Table 2). Surprisingly, only two of the conjugates listed in Table 2 were designed for the treatment of cancer. 3.1.2. PEG Conjugates with Low Molecular Weight Anticancer Drugs. In addition to conjugates with proteins and glycoproteins, PEG has often been used as a carrier for LMW drugs. Single-chain polymers, multiblock polymers, and polymers of various molecular weights and architectures (linear, star, and grafted) have all served as drug carriers for the delivery of drug molecules attached to the polymer via nondegradable or, preferably, degradable linkages or spacers. In most cases, conjugates that do not release an active anticancer drug do not have convincing in vivo activity.126,127 Consequently, PEG conjugation with the drug being attached to the polymer via a biodegradable spacer seems to be a very promising general strategy for drug conjugation and delivery. One drawback of this approach is that a simple linear PEG chain has only one or two reactive groups available for drug attachment, which leads to low drug loadings, especially in the case of HMW PEGs. This limitation can be overcome by synthesizing PEG carriers with star,128 dendron-type,129 or biodegradable multiblock structures130−133 that have higher contents of reactive groups. Drug loading can also be improved by using PEGs with branches formed by (for example) the incorporation of amino acid residues such as Lys, Asp, or Glu at one or both ends of the PEG chain. This approach has been used to prepare conjugates

can be conjugated with 2-pyridyldisulfanyl PEG derivatives, yielding conjugates with a PEG-based hydrophilic shield connected through a disulfide bond that is broken in reducing environments such as that in the cytoplasm of the target cells. Maleimide, iodoacetamide, and vinyl sulfone groups can also be used to activate PEG and achieve site-specific PEGylation of biologically active proteins (i.e., PEGylation at a specific location in the protein molecule).103 An alternative approach to selective PEGylation of proteins is the enzyme-catalyzed reaction of polymers with appropriate residues in the protein structure.102,103 In recent years click chemistry has offered new possibilities for achieving specific conjugation including the “PEGylation” of proteins.75,76 The physicochemical properties and biological behavior of PEG-modified proteins can be controlled by adjusting the molecular weight of the mPEG chains and the degree of protein substitution. It is generally accepted that the use of long polymer chains or a high degree of substitution improves the protein’s resistance to degradation and reduces the incidence of undesirable interactions but also reduces antigen-binding activity in Ab conjugates and enzyme activity in PEG conjugates with enzymes while increasing the conjugates’ hydrodynamic radius and thus affecting their ability to penetrate biological membranes. The density of the polymer shield can be increased by using branched or semidendritic structures that facilitate the attachment of more PEG chains to a single binding site102 which in combination with the type of polymer attachment (site-specific, stable, or via a biodegradable bond) influences the conjugate’s final properties. 5348


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Figure 11. Structure of PEG-doxorubicin conjugates with different drug release mechanisms: (a) GFLG spacer−drug release by enzymolysis mediated by lysosomal enzymes (cathepsin B); (b) hydrazone-containing spacer−drug release by hydrolysis of hydrazone bond at pH ≈ 5.0 (< 6.5); (c) drug release by the benzyl elimination mechanism; (d) drug release by the trimethyl lock lactonization mechanism.

of cytarabine (Ara-C)134 and daunomycin.104 The structure of a previously reported branched PEG conjugate with Ara-C is shown in Figure 10. Ideally, the spacer that links the conjugated drug to the PEG chain should be stable in circulation but susceptible to chemical hydrolysis once it reaches the tumor tissue, where the pH drops from 7.4 to ∼5-6, or after uptake by cells, either by enzymolysis in secondary lysosomes or by reductive degradation in the cytoplasm. Enzymolysis mediated by lysosomal enzymes in secondary lysosomes is substrate (spacer)-specific127 and thus advantageous for controlled drug release because it can be exploited to design a spacer that is stable during transport in circulation and then selectively cleaved once the conjugate has been taken up by a cell. A variety of spacers that fulfill these requirements have been used to attach cytotoxic drugs to PEG carriers.104 In most cases, primary or secondary amino groups already present in the drug’s structure were used for conjugation with the PEGs. PEGs containing a variety of reactive groups are commercially available and can be used for direct attachment of amino-groupcontaining drugs via aminolytic or addition reactions. The reactive groups that are most widely used in conjugation reactions of this kind are reactive esters, aldehydes, tresyl or tosyl groups, epoxides, and cis-aconityl, dichlorotriazine, or thiazolidine-2-thione groups.103,135 The free active drug can be released directly by degradation of the spacer or a linkage containing either an enzymatically degradable oligopeptide or a hydrazone or trityl group that is susceptible to pH-sensitive hydrolysis. Double prodrug strategies based on biodegradable

spacers that enable two-step drug release (the initial hydrolysis of the spacer releases a prodrug, which is subsequently degraded to release the active drug) can also be effective.103,104 In addition to amides and hydrazones, ester, carbonate, and carbamate groups are also frequently used to conjugate drugs with PEGs. They are easily prepared and undergo hydrolytic degradation under physiological conditions. Importantly, the precise conditions required to induce their hydrolysis can be controlled by modifying the spacer structure in the vicinity of the conjugating group. Thiol-containing drugs can be conjugated with PEG derivatives bearing maleimide, 2pyridyldisulfanyl, iodoacetamide, or vinyl sulfone groups,103 with which they form linkages that are stable during transport but degraded in the reducing environments of cell compartments. In principle, LMW drugs can be attached to PEG carriers via a wide range of linkers, enabling reasonable control of drug release104,136,137 Some illustrative linker structures are shown in Figure 11. Most PEG−drug conjugates do not contain a targeting moiety, and their delivery to tumors is based entirely on the EPR effect (passive accumulation). The molecular weights of single-chain PEG carriers and their low functionality limit their use for anticancer drug delivery. Nevertheless, there are numerous examples of effective DDS based on PEG carriers. PEG conjugates with taxanes have arguably yielded more promising results than any other PEG−anticancer drug conjugates tested to date. PEG-conjugated camptothecin (PEG−CPT) and paclitaxel (PEG−PTX)92,104,138 have re5349


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(CRLX101), exhibited significant antitumor effects in mouse xenografts (including some that are resistant to CPT) and also showed promising results in Phase I/II clinical trials.148 Doxorubicin (DOX) is one of the most frequently studied drugs in polymer carrier-conjugated systems. Its structure features two functional groups that enable covalent attachment to a PEG carrier: the amino group and the C13 oxo group. PEGs with diverse architectures (linear and branched) with molecular weights of 5−20 kDa were conjugated with DOX using oligopeptide spacers having various structures.137 When incubated in a cathepsin B solution, these conjugates released DOX at a rate that depended on the structure of the oligopeptides. There was no direct relationship between the rate of DOX release and the conjugates’ cytotoxicity to B16F10 melanoma cells, but all of the conjugates were 10-fold less toxic than free DOX. Relative to free DOX, all of the conjugates exhibited greatly enhanced circulation times and levels of tumor targeting, with significantly lower accumulation of the anthracycline in the heart (cardiotoxicity is serious problem associated with DOX treatment in humans). A linear PEG5000GFLG-DOX conjugate proved to be most effective at treating L1210 leukemia and B16F10 melanoma in mice, significantly increasing survival times in both cases. Biodegradable DOX-conjugated PEG2000 multiblock copolymers (40 kDa) in which drug conjugation was achieved via a pH-sensitive hydrazone bond or an enzymatically degradable GFLG spacer have been synthesized with the aim of improving tumor accumulation due to the EPR effect (Figure 14). The conjugates were stable at pH 7.4 but released DOX in phosphate buffer of pH 5 or in cathepsin B solution and significantly inhibited tumor growth in colorectal carcinoma C26-bearing mice.130,132 Direct conjugation of a similar biodegradable multiblock PEG2000 DOX conjugate132,149 with a human IgG as a targeting moiety yielded an immunoconjugate that was fairly stable at pH 7.4 but released DOX readily in mildly acidic environments. This immunoconjugate reportedly exhibits significant cytotoxicity and in vivo anticancer activity in mice bearing EL4 T-cell lymphoma. Enzymatic or hydrolytic degradation of the multiblock polymer carrier generated PEG2000 fragments that were excretable by glomerular filtration. Conjugation of a polymeric DDS with an IgG antibody attached via an oxidized FC fragment or reducible disulfide bridges has also been reported.149 However, in all of these cases it is not clear whether the IgG moiety enabled active tumor targeting or merely improved the conjugates’ tumor accumulation by increasing their molecular weight and thereby enhancing the magnitude of the EPR effect.

ceived FDA approval for use in clinical studies. In the PEGCPT conjugate the drug was linked to carboxyl groups on the PEG40000-(Ala-OH)2 carrier via ester bonds (formed using the free hydroxyl group at the C20 position of the drug molecule) that facilitate hydrolytic drug release (Figure 12).

Figure 12. Structure of a clinically tested PEG-CPT conjugate.

The CPT payload of this conjugate was rather low (1.7 wt %), which is typical for PEG conjugates with only two binding sites. In the Phase I−II clinical trials, the conjugate showed good water solubility, low side toxicity (MTD 3240 mg m−2), and prolonged circulation in comparison to the parent drug. Unfortunately, only 2 of the 27 participating patients exhibited partial responses, and it seems that further development of the drug has been discontinued.139,140 A PEG−PTX prodrug has also been tested in clinical trials with patients having advanced solid tumors and lymphomas. In this conjugate, PTX is attached to the PEG using an amino acid spacer and an ester bond at the drug’s C2 position. Unfortunately, detailed results from these trials have not been published. A carrier of the same size was also used to prepare PEG(40000)-Gly-PTX conjugates141,142 that exhibited a higher therapeutic index than PTX alone and had significant antitumor activity in mice inoculated with solid HT-29 colon, A549 lung, and SKOV3 ovarian tumors. The HMW conjugate’s superior performance was attributed to controlled drug release and enhanced accumulation in the tumors due to the EPR effect. Another report described a PEG conjugate with a CPT derivative (7-ethyl-10-hydroxycamptothecin) bound to a multiarm PEG carrier.143 This conjugate had a high drug loading, good solubility, excellent anticancer activity against a panel of cancer cell lines in vitro, and enhanced activity in the MX-1 xenograft mice model relative to the commercially available CPT derivative irinotecan, which was used as a positive control. The potential structural diversity of PEG-based drug carriers is demonstrated by the polyrotaxanes, noncovalent complexes of PEG with cyclodextrins.144−146 CPT has been covalently attached to ß-cyclodextrin-modified PEG via a glycine linker (Figure 13).146,147 This conjugate, which has an Mw of ∼90 kDa

Figure 13. Schematic representation of the structure of a PEG−ß-cyclodextrin conjugate of camptothecin. PEG and cyclodextrin blocks are connected by thioether linkages formed by terminal Cys residues on the PEG blocks, and the CPT molecules are bound to the Cys residues via Gly spacers.146 5350


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using a double-prodrug strategy based on benzyl elimination.104,138 The rate of drug release depended on the spacer’s structure, and the conjugates were more effective in in vivo treatment of M109 or mammary carcinoma MX-1 tumor models than the parent drugs. PEG−DNR conjugates were also prepared using an alternative conjugation strategy based on trimethyl lock lactonization.138,150 Unfortunately, the release of DNR from these conjugates was almost completely uncontrolled. The potential for creating PEG conjugates with an enhanced EPR effect was investigated using a conjugate with a double mechanism of action based on the cytotoxicity of epirubicin and the vasodilatory effect of nitric oxide. The resulting conjugate exhibited greater anticancer activity than free epirubicin and was also less cytotoxic toward nonneoplastic cells.103,151 Studies on PEG−DOX conjugates have also demonstrated that PEG carriers can be actively targeted to angiogenic tumors. A cyclic RGD peptide was used as a targeting moiety for vascular endothelium cells and tumor cells overexpressing the αvβ3 integrin. The PEG-DOX-E-[c(RGDfK)2] conjugate152 (whose structure is shown in Figure 15) bonded to the endothelium cells and selectively accumulated in murine mammary tumors in BALB/c mice, whereas a similar conjugate with a random peptide in place of the targeting moiety was inactive. Cytarabine (Ara-C) is another drug that has been tested in PEG conjugates for cancer treatment. Various strategies have been used to prepare PEG−Ara-C prodrug conjugates.153−155 Most of them relied on acylation of the drug’s N4-amino group to form the conjugate, with cleavage of the resulting bond releasing the drug under physiological conditions.104 Unfortunately, the initial conjugates had low drug loadings that limited their anticancer activity. The loading problem was solved by replacing the PEG carrier with branched PEG−aspartic acid derivatives or aspartic acid-based dendrons, enabling the attachment of tetrameric or octameric Ara-C derivatives.134 The release of Ara-C from these conjugates was achieved by introducing spacers with benzyl elimination or trimethyl lock lactonization mechanisms to enable drug release at a rate determined by the rate of spacer hydrolysis. Adequate Ara-C loadings were achieved by covalently linking the drug to linear or branched PEG5000, PEG10000, and PEG20000156 chains such

Figure 14. Structure of a biodegradable multiblock PEG conjugate with DOX bound via a GFLG spacer with GluLysGlu linkers between PEG blocks.

A similar strategy was used to design a multiblock PEG carrier composed of PEG4000 blocks connected with bis(4hydroxybutyl) maleate units to form a multiblock polymer containing biodegradable carbonate bonds in the polymer backbone. The polymer was grafted with HS-PEG3000-GFLGDOX side chains, so the DOX “payload” was attached via an enzymatically degradable spacer.136 The relatively low antitumor activity of this conjugate was attributed to a low rate of DOX release, emphasizing the importance of selecting a suitable biodegradable spacer to achieve a rate of drug release that generates a therapeutically effective concentration. The dependence of conjugate activity on PEG carrier size was studied using PEG carriers of 20 and 70 kDa bearing DOX bound via a pH-sensitive hydrazone linker. The conjugate based on the 70 kDa carrier was approximately 20-fold less active than that based on the 20 kDa carrier in vitro, and both conjugates exhibited much lower cytotoxicity than the parent drug but much higher cytotoxicity than comparable conjugates with the DOX moiety bound via an amide bond.127 The influence of the spacer’s structure on the activity of PEG conjugates was studied using daunorubicin (DNR) and DOX conjugates in which the drugs were conjugated to the PEG

Figure 15. Structure of a PEG−DOX conjugate targeted with the antiangiogenic peptide c(RGDfK).151 5351


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that four or eight Ara-C molecules were attached to one PEG carrier with branches formed by dicarboxylic amino acid units. These conjugates exhibited increased stability to degradation, slow blood clearance in mice, and reduced cytotoxicity relative to free Ara-C. Branched PEG carriers that permit multiple drug loading have also been used to prepare conjugates of 6mercaptopurine (6-MP)157 and gemcitabine (GEM).158 In the PEG−6-MP conjugates, the thiol group of 6-MP was modified with a benzyl elimination system and used for attachment to the carrier. The conjugate prodrugs exhibited significant activity against murine leukemia and selected solid tumors in mice. In the PEG−GEM conjugates, the drug was attached via an amide bond and a divalent aminoadipic acid spacer together with folic acid as a targeting group. The affinity of the targeted conjugate for cells overexpressing the folic acid receptor was 2−3 times greater than that of nontargeted analogs. PEG carriers also have potential applications as theranostics.159 In one study, PEG was conjugated with a photosensitizing fullerene to form a polymer bearing a gadolinium complex at the PEG terminal group. When injected intravenously into tumor-bearing mice the conjugate exhibited good MRI activity and significant antitumor activity after light irradiation.160 The results discussed above clearly show that PEG is a promising hydrophilic polymer carrier for the delivery of diverse drugs. It is particularly well suited for conjugation with biologically active proteins and glycoproteins, for coating and “stealthing” nanoparticulate DDS, and for preparing drug carriers based on polymer micelles and nanoparticles. Unfortunately, PEGylation has not achieved similar levels of success in the delivery of small molecule drugs, despite the expenditure of considerable effort in this field. It thus seems that PEG-based systems will be superseded by conjugates based on multivalent polymers, preferably with biodegradable backbones and better solubility.

Table 3. List of HPMA Copolymer Conjugates Approved for Clinical Evaluation conjugate PK1 PK2 MAG-CPT PNU166945 AP5280 AP5346 a

drug DOX DOX (galactosamine) CPT PTX carbo platinate DACH platinate

phase of testing

spacer GFLG GFLG

I and II I and II

GAcapGa GFLG GFLG GGG-amino malonate

I I I I and II

ref 168, 169 32, 165, 170, 171 172−174 175 176, 177 178, 179

Acap, aminocaproyl.

conjugate molecules (which are primarily due to the polymer carrier’s polydispersity), and the tendency for such conjugates, namely, those actively targeted, to perform differently in preclinical animal models than in the human body. The latter effect is probably due to differences in animals’ and humans’ tumor cell receptors and in immune systems and appears to be especially pronounced when dealing with polymer therapeutics that have both cytotoxic and immunostimulating effects.180 Treatment of human tumor models in animals requires use of immunodeficient mice in which the immunostimulating effect of the polymer conjugates cannot be effective. Even if the immunostimulating effect as an important constituent of anticancer activity was described for HPMA copolymer conjugates, it can be expected also for other polymer conjugates. Nevertheless, many P(HPMA)-based conjugates, particularly the conjugates effectively accumulating in solid tumors by EPR, exhibit outstanding in vivo properties and significantly outperform the free parent drugs, as documented below, thus giving the chance for foreseeable application in human therapy. The most frequently used strategy for the synthesis of P(HPMA) conjugates is based on copolymerization of HPMA with monomers containing spacers that terminate in a reactive group such as a reactive ester (e.g., a 4-nitrophenoxy ester) or a thiazoline-2-thione. This reactive group then undergoes aminolysis with a primary or secondary amino group in the drug. P(HPMA) conjugates can also be prepared by copolymerization with a monomeric drug; this has been done with a DOX-terminated methacryloylated oligopeptide.65,164,181,182 Finally, conjugation can be achieved by activating the carboxylic acid groups in the P(HPMA) backbone with carbodiimides, which then undergo a binding reaction with a drug bearing a free hydroxyl group (e.g., camptothecin), a free thiol, or some other group with similar characteristics. Direct conjugation was employed in the synthesis of a P(HPMA) conjugate with DOX bound via a hydrazone bond-containing spacer;183 the structure of the resulting P(HPMA)−DOX conjugates is shown in Figure 16. Spacers or bonds formed during conjugation should be susceptible to pH-sensitive hydrolysis or specific enzymolysis that will facilitate the release of the parent drug in its original form. Details on the synthesis of the P(HPMA) conjugates can be found in recent reviews.164,184 Another approach for HPMA copolymer synthesis that has been used by Zentel’s group involves preparing a polymeric active ester precursor, poly(pentafluorophenyl methacrylate), by controlled RAFT (reversible addition−fragmentation chain-transfer) polymerization. The activated ester polymer is then subjected to aminolysis with 1-aminopropan-2-ol.185,186 This approach yields a product

3.2. Copolymers of N-(2-Hydroxypropyl)methacrylamide

Copolymers of N-(2-hydroxypropyl)methacrylamide (HPMA) rank among the most frequently studied water-soluble drug carriers. Unlike PEG carriers, which have only one or two endchain functional groups available for polymer modification, HPMA copolymers or P(HPMA)s are multivalent polymers with functional groups randomly distributed along the polymer chain. The number of functional groups in the copolymer can be varied substantially by changing the relative abundance of the functional comonomer in the polymerization feed mixture. 3.2.1. HPMA Copolymer−Drug Conjugates for Passive Drug Delivery. HPMA copolymer−drug conjugates intended as prodrugs for cancer treatment have been studied since P(HPMA)s were first suggested as drug carriers.91,92,161 Key achievements in this field were recently summarized in another article.162 A large body of information has been collected concerning their in vitro activity and mechanism of action as well as their in vivo activity, mainly in mice163,164 but also in humans.32,85,165 Surprisingly, only a few soluble drug conjugates based on P(HPMA) have entered clinical trials (Table 3). Difficulties with introduction of the polymer−drug conjugates into clinics can be documented by the fact that after more than 30 years of development only one of all the hitherto clinically evaluated polymer−drug conjugates has been introduced into regular clinical practice.166,167 The limited clinical use of polymer−drug conjugates to date, including P(HPMA), is primarily due to their rather complicated synthesis, the unsatisfactory characteristics of the 5352


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Figure 16. HPMA copolymer conjugates of DOX. (Left) Polymer conjugate with DOX bound via an enzymatically degradable GFLF spacer. (Right) Polymer conjugate with DOX bound via a pH-sensitive hydrazone bond-containing spacer.

Figure 17. Scheme of the synthesis of the HPMA copolymer functionalized by radical thiol−ene chemistry.

with a narrow Mw distribution and enables the synthesis of block copolymers and versatile targeted drug carriers for sophisticated molecular imaging and tumor immunotherapy.187,188 The synthesis of multifunctional P(HPMA)s via postpolymerization modification and sequential thiol−ene chemistry was described by Alexander’s group. 189 Allyl groups were introduced by reacting the P(HPMA) chain’s pendant hydroxyl groups with allyl isocyanate to form the corresponding N-allyl carbamates. The resulting allyl-modified HPMA copolymer was then easily functionalized by radical thiol−ene chemistry. Copolymer derivatization in this way has been performed with complex thiols bearing pendant functionalities including mercaptopropionic acid, cysteamine, mercaptoethanol BODIPY-SH, and the cell-targeting peptide RGDC.189 For a scheme of the copolymer synthesis see Figure 17.

Over the years, various anticancer drugs have also been attached to nondegradable P(HPMA) carriers via the enzymatically degradable oligopeptide (GFLG) spacer. One such conjugate called PK1 and its galactosamine-targeted analogue PK2 were the first polymer−DOX conjugates subjected to Phase I/II clinical trials.168−171,190 DOX aside, a number of other anticancer drugs have been conjugated with similar linear nondegradable P(HPMA) carriers via oligopeptide spacers, including puromycin,191,192 aminoellipticin,193 methotrexate,194 melphalan and bis(2-chloroethyl)amine,195−197 camptothecin,172 taxol,175,198,199 platinates,176 5-fluorouracil,200 derivatives of cyclosporine A,201,202 geldanamycin,203−205 an 8-aminoquinoline analogue,206 and other drugs. The biological activity and enzyme-mediated drug release of these conjugates have been tested, and some have also undergone clinical evaluation (Table 3). 5353


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transplanted with a lethal dose of cancer cells with no subsequent treatment effectively resisted this second attack of cancer, achieving good long-term survival rates (Figure 18).

In addition to DOX PK1/PK2 conjugates, a P(HPMA)− PTX conjugate has also been tested in phase I/II clinical studies. It exhibited less side toxicity than the parent drug but only induced partial remission of skin metastases in breast cancer patients.175 A P(HPMA) conjugate with CPT was also tested in Phase I/II clinical studies. In this case, a MAG P(HPMA) chain with GlyAcapGly spacers bound to the side chains was used as the drug carrier.174 Unfortunately, these trials were not continued into phase II because the conjugate had only a partial effect and caused obvious bladder toxicity. The final type of P(HPMA)-based conjugates that have been tested in Phase I/II trials against cancer are P(HPMA) platinates.176 Two conjugates were tested for adverse events and to determine their maximum tolerated doses and doselimiting toxicity: a carboplatin analog attached to the P(HPMA) chain via a GFLG spacer and a malonato 1,2diaminocyclohexyl derivative attached via the same spacer. These polymer conjugates of platinum complexes exhibited minimal toxicity compared to LMW Pt derivatives with significantly increased MTD values and prolonged plasma exposure. However, only a few patients responded to the treatment, and those responses were only partial. Although the results achieved to date with P(HPMA) conjugates in clinical trials have not matched their outstanding performance in tumor-bearing animal models, a number of sophisticated tailormade passively or actively targeted conjugates are currently under development or being tested in preclinical studies. Hopefully, some of these treatments will live up to the clear potential of these systems in cancer therapy. As mentioned above, drugs should ideally only be released from the carrier at the desired site of action. The target for anticancer drugs is the tumor tissue or even better the cancer cells themselves. The development of new biodegradable spacers containing hydrolytically degradable linkers based on hydrazone bonds, cis-aconityl moieties, acetal or ketal bonds, reductively degradable disulfide bonds,207 or other functional groups offers diverse possibilities for controlling drug release and ensuring that it occurs specifically in tumors or tumor cells due to the presence of specific enzymes, a change in the pH, or the reducing nature of the cytoplasm. A recent review listed the linear P(HPMA) drug conjugates that have been reported to date and discussed their synthesis and antitumor activity.163,164,184,208 In most cases a reasonably biodegradable spacer between the polymer and the drug was a prerequisite for in vivo antitumor activity, but some results suggested that drug release may not be essential for anticancer activity in general.209 Here, we focus exclusively on conjugates of linear HPMA copolymer carriers with cytostatic agents that have performed well during in vivo preclinical trials or that exhibit particularly noteworthy properties. A P(HPMA) DOX conjugate in which the drug was bound with a spacer containing a pH-sensitive hydrazone bond exhibited cytotoxic effects in mice and induced a specific antitumor immune response that could be transferred ̈ recipients by splenocyte transplantation.210,211 This to naive promising and structurally relatively simple polymer prodrug had an MTD that was at least three times greater than that for PK1 and showed excellent antitumor activity in vivo when administered to EL4 T-cell lymphoma-bearing mice, with 100% long-term survival achieved after a single dose. No side toxicity was observed, and importantly, the treated mice retained a degree of resistance to new attacks of the disease.212−214 Mice cured of EL4 T-cell lymphoma by treatment with the HPMA copolymer−DOX or −docetaxel conjugates and then re-

Figure 18. Survival of mice with EL4 T-cell lymphoma after treatment with HPMA copolymer conjugates of doxorubicin (dash-and-dot line (5 mice in group)) or docetaxel (dotted line (5 mice in group)) and both conjugates together (dashed line (8 mice in group)). The full line shows results for control mice re-transplanted with a lethal dose of cancer cells without drug treatment (8 mice in group). Conjugatetreated mice exhibited effective resistance to the second attack of cancer and had good long-term survival.

Similar conjugates with pirarubicin (THP) also showed substantial antitumor activity with no observable side effects in a recent study on mice bearing S-180 colon sarcoma or autochthonous breast and colon cancer.215,216 Moreover, the conjugate was safe and the treatment effective in a preliminary evaluation in humans with metastasized prostate cancer.217 Treatment of a patient with metastases in the lungs (grade IV disease) with the HPMA copolymer−pirarubicin conjugate reduced their PSMA levels to normal values and caused the clearance of the metastases (Figure 19). In another study, the potent DNA intercalator ellipticine was labeled with the radioisotope 125I (an Auger electron emitter) and conjugated to P(HPMA) via a hydrazone linker. Cancerbearing mice treated with this conjugate exhibited improved survival,218 demonstrating that P(HPMA) conjugates have potential applications also in radiotherapy and can be used to provide cancer treatment with dual mechanisms of action. After accumulation of the conjugate in solid tumor and its uptake by cancer cells the intercalating agent and the emission of Auger electrons will destroy DNA inside the cells. Polymer drug conjugates are well established as vehicles for the delivery of single therapeutic agents, but an increasingly common strategy in cancer treatments under development is to use polymer carriers to deliver multiple therapeutically active compounds simultaneously in so-called combination or multiagent therapies.219 This makes it possible to exploit synergies between different drugs’ mechanisms of action and could potentially greatly improve the quality of treatment. Combination therapies can involve treatment with two different polymer conjugates each bearing a single drug or treatment with a single conjugate carrying two different drugs bound to the same polymer chain. Conjugates of the latter sort bearing DOX and dexamethasone220 or DOX and mitomycin C221 have been synthesized and shown to exhibit improved activity and synergistic effects. The potential effectiveness of combination therapy is demonstrated by the example of a P(HPMA) conjugate containing DOX covalently conjugated via a pH-sensitive hydrazone bond and 5-fluorouracil (Fu) conjugated via the 5354


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Figure 19. Structure of the polymer conjugate with pirarubicin (left) and CT images of metastatic nodules in a prostate cancer patient (stage IV) before treatment (right, top) and 5 months after the last treatment (6 doses) with the PHPMA-THP conjugate (right, bottom). Reprinted with permission from ref 217. Copyright 2015 Springer.

oligopeptide sequence GFLG.222 The conjugate was also targeted by decoration with a galectin-3 targeting peptide. The resulting dual-drug-loaded conjugate achieved 81.6% inhibition of PC-3 tumor xenograft growth in mice and was thus more effective than the corresponding nontargeted conjugate P(HPMA)-DOX-Fu (71.2% inhibition), the singledrug-loaded conjugates P(HPMA)-DOX (63%) and P(HPMA)-Fu (32.0%), and the corresponding free drugs DOX·HCl (40.5%) and 5-Fu (14.6%). An interesting example of synergy in drug action was also reported for a biodegradable conjugate based on a P(HPMA) diblock carrier with one block bearing DOX bound via a GFLG oligopeptide spacer (which preferentially causes necrosis) and a second bearing DOX bound via a hydrazone group (which preferentially causes apoptosis).214,223 Significant in vivo antitumor activity and synergistic drug action was reported for the double conjugate bearing DOX bound by both types of linker and also for a combined treatment based on a mixture of single conjugates each bearing DOX bound by only one type of linker.224 Synergistic effects were also observed in cytotoxicity tests against MCF-7 breast cancer cells using a P(HPMA) conjugate bearing the aromatase inhibitor aminoglutethimide and DOX, both of which were bound to the carrier via oligopeptide spacers.225,226 This double conjugate thus represents an example of a polymer prodrug with synergy derived from dual activity. Biodegradable HMW P(HPMA) carriers have also been used to effectively deliver other drug combinations including gemcitabine and diaminocyclohexyl platinum227 or gemcitabine and paclitaxel.228 In keeping with the current focus on the importance of cancer stem cells in cancer treatment, a P(HPMA) conjugate bearing a PI3K/mTOR inhibitor and docetaxel was synthesized. The conjugate exhibited preferential inhibitory effects on

prostate cancer stem cells and cytotoxicity toward bulk tumor cells.229 Simultaneous effects on cancer stem cells and bulk tumor cells were also achieved using P(HPMA) conjugates bearing both cyclopamine (active against stem cells) and docetaxel (toxic toward tumor mass cells).230 These combination therapies that simultaneously attack cancer stem cells and differentiated cancer cells are mentioned to introduce one of the most important new strategies in cancer treatment. This approach could potentially significantly improve cancer therapy, although many issues remain to be addressed. For example, it is not clear whether it is better to use a single conjugate bearing multiple drugs or a combination of two conjugates each bearing a single drug. It will also be important to optimize the ratio of the two drugs, the release profile of each drug, the overall drug loading, and the timing of the drugs’ administration among other things. The ADEPT and PDEPT systems (antibody- or polymerdirected enzyme prodrug therapy) bear mentioning as specific examples of therapies that use a combination of two active molecules with a polymeric delivery system.8,231 These approaches involve treatment with a P(HPMA) prodrug and a separate poly(HPMA)−enzyme conjugate that selectively generates the cytotoxic drug from the polymer-bound precursor within the tumor tissue. Various combinations of polymer prodrugs and polymer−enzyme conjugates have been tested232 including a PK1 and polymer−cathepsin B system.233 Polymer theranostics are multifunctional polymer conjugates combining at least two types of molecules, typically a drug and a chemical species that enables in vivo visualization of the conjugate’s distribution, that are both attached to the same polymer carrier usually via specific spacers. Their conjugation may be achieved via covalent bonding (in the case of fluorescent labels), complexation with chelating molecules (in the case of radionuclides or metal ions for MRI), or entrapment 5355


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Figure 20. Structures of biodegradable HMW drug conjugates including a branched carrier structure (a), linear diblock and multiblock carriers (b), a graft polymer carrier (c), and a star polymer carrier (d).

ing the initial description of branched P(HPMA)−drug conjugates (Figure 20a),251−253 better defined grafted254 (Figure 20c), diblock,224 multiblock,255−258 (Figure 20b) and star259−263 (Figure 20d) HMW conjugates with anticancer drugs were developed. The formation of grafted systems with a relatively broad size (Mw) distribution can be avoided by incorporating hydrophobic moieties (e.g., cholesterol)264 into P(HPMA) chains that are linked to anticancer drugs such as DOX via a pH-sensitive spacer. A linear polymer of this sort with an Mw of around 30 kDa self-assembled into a HMW supramolecular structure (100−200 kDa, ∼25 nm in diameter) in aqueous solution. The conjugates with cholesterol moieties showed prolonged blood circulation, enhanced tumor accumulation, and yielded longterm survival rates of up to 100% when administered in one or two relatively small doses to mice bearing EL4 T-cell lymphomas. Elimination of the carrier can be achieved by disassembly of the system after dilution or/and after biodegradation of a spacer situated between the polymer and the cholesterol moieties.265 A similar P(HPMA) system incorporating lauryl methacrylate was designed as a drug carrier capable of penetrating the blood−brain barrier without harming it.266 Although originally developed as a carrier for neurological drugs, this system clearly has potential as a vehicle for delivering cancerostatics to the brain. Biodegradable linear diblock or multiblock P(HPMA) drug carriers have been synthesized with the aim of creating carriers with prolonged blood circulation and greater drug accumulation in solid tumors than can be achieved with the widely used linear copolymers. Disulfide spacers that are degraded reductively in the cytoplasm or GFLG spacers that are degraded enzymatically in lysosomes were positioned between the polymer blocks, enabling intracellular polymer carrier degradation and the subsequent elimination of the resulting shorter degradation fragments. The size of the polymer coil in solution controls the rate of polymer elimination by glomerular filtration rather than the polymer’s Mw per se, although the Mw is a convenient and easily calculated measure related to the size,

in the nanoparticle system. These systems are important and have opened up new opportunities in cancer treatment by making it possible to directly monitor the fates of drugs within the body and to evaluate their effectiveness in the treatment of solid tumors. However, developments in this field have been reviewed recently,234,235 so we refer interested readers to those reviews and will not discuss them further here. P(HPMA) derivatives of biologically active proteins (enzymes) represent a specific group of polymer nanomedicines. Both semitelechelic and multivalent P(HPMA)s have been used for protein modification, which is mainly achieved by aminolysis with reactive ester-bearing polymer precursors. This approach yields conjugates with high polydispersity when performed with multivalent polymers or a less dispersed star structure when performed with a semitelechelic polymer grafted onto a central protein molecule. P(HPMA) conjugates have been prepared with the model protein chymotrypsin,236 bovine seminal ribonuclease,237−239 and RNAse A;240,241 the latter conjugate exhibited remarkable antitumor activity in mice in vivo. In addition, poly(HPMA) conjugates with superoxide dismutase,242 lysosyme,243,244 trypsin,245 insulin,246−248 and acetylcholine esterase249 have been synthesized and tested for anticancer activity. A P(HPMA) conjugate with interleukin-2250 exhibited a significant in vivo immunostimulatory effect, which has implications for immunotherapy. In general, conjugation with hydrophilic polymers often reduces proteins’ toxicity,250 increases their resistance to proteolysis in circulation, alters their pharmacokinetics and pharmacodynamics, and improves their catalytic/ biological activity. When administered in combination with anticancer drugs, such conjugates may exhibit synergistic effects (as was observed for DOX and RNases) that confer significant in vivo anticancer activity in model tumor-bearing mice.239−241 3.2.1.1. HPMA Copolymer Conjugates for Improved Passive Tumor Accumulation. After it was recognized that P(HPMA)s accumulate in solid tumors due to the EPR effect with molecular weight-dependent efficiency,14,24 HMW biodegradable conjugates were designed and synthesized. Follow5356


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Figure 21. P-gp mechanism of multidrug resistance in cancer cells.

Figure 22. P(HPMA) conjugate bearing DOX (left) together with the ABC transporter inhibitor ritonavir (right) in order to overcome multidrug resistance to DOX treatment.

so its value is often used as characteristic for the elimination limit of polymers. Nanomedicines above a certain size cannot achieve effective extravasation or tumor accumulation. For P(HPMA) derivatives, this size limit appears to be around 50 nm, which corresponds to an Mw of around 600 kDa;23 conjugates with Mw values above 1000 kDa (100 nm) exhibited markedly reduced accumulation in tumors. Linear HPMA copolymers with Mw values of up to 70 kDa can be excreted in the urine, but more rigid star conjugates have a lower renal threshold of around 50 kDa.23 Controlled RAFT polymerization of HPMA using a specific GFLG oligopeptidecontaining chain-transfer agent yielded a diblock polymer carrier with a narrow molecular weight distribution that was used as a click reaction substrate to prepare a range of HMW multiblock carriers and conjugates.256,258,267 The resulting longcirculating carriers were used to deliver individual drugs (DOX,268 gemcitabine, paclitaxel,255 prostaglandin269) and also proved suitable for combination therapy. Biodegradation of the conjugates with model or lysosomal enzymes resulted in the formation of polymeric degradation products with Mw values below the renal threshold.162

In addition to linear and biodegradable graft conjugates, biodegradable star conjugates have been evaluated as carriers for tumor-specific drug delivery. At the core of these conjugates is a low-generation polyamidoamine dendrimer with terminal amino groups, which are grafted with semitelechelic P(HPMA)s bearing a covalently bound drug260 (Figure 20d). To make the star carriers biodegradable, disulfide or oligopeptidic GFLG spacers were introduced between the dendrimer end groups and the polymer grafts. 260,270 Reductive or enzymatic degradation of the conjugates or their incubation with cancer cell cultures converted them into polymer fragments with Mw values below the renal threshold. Conjugates of this sort with DOX,271 pirarubicin,216 or docetaxel199 (all bound via pHsensitive hydrazone bond-containing spacers) exhibited prolonged blood circulation, enhanced tumor accumulation, and excellent anticancer activity with good long-term survival rates in murine models. These very promising HMW carriers have a relatively narrow molecular weight distribution that can be made even narrower by using RAFT polymerization to prepare the semitelechelic polymer grafts.259 In vivo studies showed that the star conjugates had no detectable side toxicity and that 5357


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further studies will be needed to properly evaluate the potential of PDT in cancer therapy and imaging in humans. Another site-specific drug delivery system designed for oral application is a colon-targeted P(HPMA) conjugate with 9aminocamptothecin284 bound via an oligopeptide spacer with an aromatic azo group. Because the colon has a specific combination of azoreductase and proteolytic activity285 not found elsewhere in the body, the carrier selectively released the drug in this organ after oral administration. Importantly, the conjugate was stable during transport through the GI tract (i.e., in the stomach and small intestine).286 Conjugates of P(HPMA) with DOX bound via a GFLG spacer to a carrier containing one of the nonspecific human immunoglobulins Intraglobin and Endobulin have also been prepared and shown to completely heal mice with developed tumors after treatment with a single dose of the prodrug.287 Because these conjugates contain nonspecific Abs they are not (strictly) actively targeted, but they nevertheless effectively eradicated tumors in the murine model without harming the immune system; in fact, immunoprotective effects and the induction of systemic antitumor resistance were observed in some cases.288−292 The very promising results obtained with these conjugates in animal experiments were verified in a preliminary clinical assessment conducted with human patients with generalized breast cancer.290,292 The role of the nonspecific Ab in these conjugates is not entirely clear; its effect is probably partly due to an enhancement of the EPR effect and partly to immunostimulation. 3.2.2. Synthesis and Activity of Actively Targeted HPMA Copolymer−Drug Conjugates. In addition to drugs, multivalent P(HPMA)s can be conjugated to a wide variety of ligands that may serve as efficient targeting moieties or change the carrier’s properties and thereby modify its biodistribution and interactions with specific body compartments. Daunomycin- and DOX-containing HPMA-based conjugates bearing D-galactose, D-mannose, D-fucose, and melanocytestimulating hormone (MSH) ligands as early targeting moieties were evaluated in tests of receptor-mediated targeting.31,39,293−296 The results obtained showed that simple ligands can significantly influence the body distribution of polymer conjugates and promote receptor-mediated targeting in vivo (e.g., D-galactose promotes targeting to hepatocytes, Dfucose to L1210 cells, and MSH to B16F10 melanoma). Glycoproteins (antibodies) with high affinity for cell membrane antigens can serve as more specific targeting moieties. Moreover, they can contain a wide range of functional groups (primary and secondary amines, carboxyl and thiol groups, aldehydes obtained by oxidizing saccharide units) that offer many possibilities for their conjugation to polymer carriers bearing appropriate reactive groups. Reactions used to conjugate an Ab with a polymer should be efficient and high yielding and preserve the antibody’s affinity for its antigen. Simple monosaccharide or oligopeptide targeting moieties as well as more complex polyclonal and monoclonal antibodies (Ab) or Ab fragments targeting T-lymphocytes,297 transferrin receptor,298 or human colorectal carcinoma299 were attached to carriers using an aminolysis reaction in which reactive ester groups on a polymer precursor were aminolyzed by primary amino groups on the ligands. One of these Ab-targeted conjugates, a conjugate of P(HPMA) with the photosensitizer chlorin e6 (mesochlorin), was used to assess the concept of double-targeted anticancer agents. This system combined active targeting of mouse splenocytes using an anti-Thy1,2 mAb with

their anticancer activity was superior to that of equivalent linear or grafted conjugates. Multidrug resistance (MDR) is a major obstacle in tumor chemotherapy using LMW drugs. Cancer cells that are repeatedly exposed to anticancer drugs can develop resistance to the treatment via multiple mechanisms including overexpression of the transmembrane P-glycoprotein (P-gp) in the cell membrane. This ATP-dependent efflux pump expels drug molecules from the cell interior, reducing their effective intracellular concentration (Figure 21) and potentially causing the therapy to fail.272 Many attempts have been made to develop tumor therapies that can overcome MDR. These typically involve combining anticancer drugs with inhibitors of ABC transporters (MDR modulators).273 In principle, polymer nanomedicines with covalently bound drugs that enter the cell by endocytosis could potentially overcome MDR based on P-gp overexpression because they only release the bound drugs inside secondary lysosomes where they are not exposed to P-gp.162 The viability of this strategy has been demonstrated in studies on ovarian carcinoma using P(HPMA)−DOX and other conjugates.274−277 The potential of polymer conjugates with a drug bound via an enzymatically degradable spacer to overcome MDR has also been investigated using antibody-targeted conjugates278 and conjugates targeted with chlorin e6.279,280 However, MDR can be a serious problem for vehicles that release drugs extracellularly, for example, in response to the lower pH of tumor tissues (as is the case for hydrazone conjugates). The potential for using simple polymer conjugates to overcome MDR in such cases is probably limited. One way of resolving this problem may be to use conjugates bearing a drug (e.g., DOX) together with an MDR modulator. P(HPMA) conjugates of this sort bearing DOX together with one of the ABC transporter inhibitors reversin121 or ritonavir281 (Figure 22) have been synthesized and tested in vitro. While the results obtained were promising, MDR continues to represent a major challenge in cancer treatment. 3.2.1.2. Other Systems for Passive Site-Specific Drug Delivery. All of the systems mentioned above rely on the EPR effect to achieve tumor-specific delivery, which means they are primarily applicable in the treatment of vascularized solid tumors. Stimulus-responsive HPMA-based polymer systems that release or activate drugs locally in response to specific signals originating from external or internal stimuli represent alternative tools for passive site-specific drug delivery. A wide variety of stimuli can be used to induce drug release; for example, in photodynamic therapy (PDT), light irradiation serves as an external stimulus. Photosensitizing agents bound to a polymer carrier can accumulate in a tumor due to the EPR effect, and the drug is activated (leading to the formation of singlet oxygen) only after irradiation with light of the appropriate wavelength. P(HPMA) conjugates with chlorin e668,282 or protoporphyrin IX215,283 have been prepared to test this concept. HPMA copolymer conjugates of protoporphyrin IX exhibited tumor-selective accumulation and remarkable antitumor activity after local irradiation in mice bearing autochthonous breast cancer and in other implanted mouse tumors (C-26 and S-180) without apparent toxicity. Moreover, when this conjugate was used as a fluorescent nanoprobe for tumor imaging it did not induce hypersensitive responses in the skin, which are commonly observed with conventional LMW fluorescent photosensitizers. While these results are interesting, 5358


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specific local drug activation by light irradiation.68,282,300−302 The drug was shown to have a photodynamic effect on splenocyte viability and suppressed the primary antibody response of mouse splenocytes to sheep red blood cells. To determine how the nature of the link between the Ab and the polymer affected the conjugate’s final biological activity, conjugates prepared by the aminolytic approach were compared to otherwise equivalent conjugates in which the IgG Ab was linked to the polymer by oxidizing its FC region. Most P(HPMA)−drug−antibody conjugates have been synthesized by aminolysis of a polymer precursor bearing reactive p-nitrophenoxy (ONp) groups.68,297,303 In the first step of such reactions the drug (typically DOX) is attached to the carrier by aminolysis of some of ONp groups, followed by aminolysis of the remaining ONp groups with the chosen Ab to form a “classic” conjugate. Careful selection of buffers, concentrations, pH, and temperature is important for complete conjugation and preservation of the conjugated antibody’s ability to bind to its receptors.304 The polymer precursor can be prepared by direct copolymerization of HPMA with the monomeric drug and a methacryloylated ONp ester. Aminolysis of a multivalent precursor with a multivalent antibody results in the formation of HMW structures with a broad distribution of molecular weights and usually also partly reduced Ab binding activity. Despite these drawbacks the synthetic procedure is practically straightforward, and the resulting DOX conjugates exhibit appreciable anticancer activity in humans.292 The protection of some Ab amino groups with DMMA during the reaction somewhat improves the activity of the resulting conjugate65 but makes the synthesis more complicated. The synthesis was significantly improved by using polymer precursors containing thiazolidine-2-thione (TT) reactive groups, which are more stable than ONp groups in aqueous solution and thus enable the single-step simultaneous aminolysis of the polymer precursor with DOX and Ab together.287 Further improvement of anticancer activity of the Abtargeted conjugates was achieved through the synthesis of “star” conjugates with a central Ab molecule grafted onto semitelechelic P(HPMA)−DOX chains. DOX was attached to the polymer via a biodegradable GFLG spacer or a pH-sensitive hydrazone bond.305 Semitelechelic HPMA copolymers (copolymers with one specific end-chain reactive group) can be prepared by chain-transfer radical polymerization,306 by using an azo-initiator containing TT or 2-pyridyldisulfanyl groups,307 or by RAFT polymerization.259,308 The monovalency of the semitelechelic copolymers means that attachment is only possible at one point and thus prevents branching reactions with a multivalent Ab. The structures of some initiators and the corresponding chain-transfer agents used to prepare semitelechelic HPMA copolymers with defined polymer chain end functional groups are shown in Figure 23. Polymers terminating in hydroxyl, thiazolidine-2-thione, azo, alkyne, pyridyldisulfanyl, hydrazo, and amino groups can be synthesized in this way. The star conjugates are better defined than the classic conjugates (only one Ab is conjugated), and their polydispersity is lower. The star conjugates also exhibited significantly better in vivo activity in the treatment of mice bearing human colorectal carcinoma SW 620305 compared to classic conjugates. Similar conjugates bearing DOX bound via a pHsensitive hydrazone bond and a monoclonal Ab (mAb) conjugated via its oxidized FC fragment using a sulfide linkage or a reducible disulfide bridge149 were prepared by reacting a

Figure 23. Structures of azo initiators and chain-transfer agents used for the synthesis of semitelechelic HPMA copolymers by RAFT polymerization.

thiol-group-containing antibody (i.e., an Ab modified by reaction with 2-iminothiolane) with a polymer precursor bearing DOX and maleimide or 2-pyridyldisulfanyl groups to form sulfide bonds or reducible disulfide bridges, respectively.253,307,309 These procedures for the synthesis of HPMA copolymer−antibody conjugates are outlined in Figure 24. All of the conjugates exhibited significant antitumor activity in vivo with good long-term survival rates in mice bearing EL4 T-cell lymphoma, although the method used to conjugate the Ab to the polymer carrier was found to influence the conjugates’ anticancer efficiency to some extent. The targeting efficiency of mAb-targeted conjugates can be further improved by copolymerizing HPMA with a monomeric antibody fragment (MA-Fab), as demonstrated by the synthesis of chlorin e6 conjugates targeted using the Fab fragment of the OV-TL 16 Ab.73 A monomer containing a short PEG spacer exhibited greater reactivity in copolymerization than one with no spacer. Administration of these conjugates to mice bearing human ovarian carcinoma caused significant inhibition of tumor growth over an extended period of time,310 confirming the potential of this new synthetic approach. Another selective method for conjugating a polymer precursor to an Ab is based on the selective oxidation of saccharide units in the FC part of the Ab molecule. Periodate oxidation does not affect the antibody’s antigen-binding site, and the azomethine or hydrazone bonds formed by conjugation with amino or hydrazine groups on the polymer can be stabilized by cyanoborohydride reduction. (Figure 24). Chlorin e6 conjugates targeted with an anti-Thy 1,2 Ab were prepared by this method68,301 and exhibited better cytotoxic activity and suppression of antibody response than a similar “classic” conjugate prepared by aminolysis. DOX and cyclosporine A conjugates targeted with oxidized anti-Thy 1,2 Ab311 also showed better cytotoxic and immunosuppressive activity in vivo than those prepared by aminolysis, but the improvement was not sufficiently pronounced to compensate for the more complex synthesis. A major drawback that currently limits the applicability of this synthetic approach for the preparation of drug carriers is the limited number of polymer chains that become conjugated to a single mAb molecule. DOX-bearing P(HPMA)s have also been conjugated with targeting lectins such as wheat germ agglutinin (WGA) or peanut agglutinin (PNA) by two-step aminolysis. These conjugates were evaluated against human colorectal carcinoma SW 620.312 The WGA-targeted conjugates exhibited significant cytotoxicity in vitro comparable to that of an antibody-targeted 5359


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Figure 24. Methods for conjugating targeting Abs or Ab fragments to HPMA copolymer precursors: (a) synthesis of a classic conjugate; (b) synthesis of a star conjugate; (c) conjugation via a sulfide spacer; (d) conjugation via a biodegradable disulfide spacer; (e) conjugation via periodate oxidation of the FC region; (f) synthesis via direct copolymerization of monomers.

polymer analog, demonstrating the potential of lectins for targeting polymer drug carriers. HPMA conjugates targeted with WGA and PNA also reportedly showed high affinity for goblet cells and other cells of the gastrointestinal tract.313,314 The drawbacks of using monoclonal antibodies as targeting moieties for conjugates are the high cost of the antibody, the comparatively low reproducibility of the synthesis, the complicated structures that are typically obtained, and the risk of immunogenicity. These can be avoided by replacing the Ab with an oligopeptide corresponding to the active sequence responsible for the Ab’s interaction with its target antigen.

Alternatively, the recognition sequence of a protein such as fibronectin or laminin can be used for the same purpose. Oligopeptides with cell receptor specificity can be also designed by phage display.315 Synthetic oligopeptides can be produced in much larger quantities and much less expensively than mAbs. The covalent attachment of a short synthetic oligopeptide to a polymer carrier can be accomplished using standard methods of peptide, click, and polymer chemistry, all of which yield welldefined polymer carriers. In principle, the conjugates can be directed straight to the receptors of malignant cells or they can be aimed at the endothelial cells of the tumor vasculature, 5360


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which overexpress αVβ3 integrins that should ideally be targeted with peptides that have angiogenesis-inhibiting activity.316−318 Cyclic oligopeptides containing RGD (RGDfK or RGD4C) are often used to deliver radionuclides (111In, 99Tc, 90Y) to tumor vasculature for therapeutic or diagnosis purposes.319−323 The resulting conjugates adhere strongly to the targeted cancer cell receptors and perform well in vivo, demonstrating the high potential of these oligopeptides as targeting moieties for actively targeted polymeric anticancer drugs. As noted above, MSH has been used as a targeting moiety for polymer−drug conjugates. Similarly, the Epstein−Barr virus nonapeptide EDPGFFNVE has been used to promote receptor-mediated targeting of a HPMA copolymer−DOX conjugate to T- and B-cell lymphomas.324 Similar P(HPMA)− DOX conjugates targeted with the nonapeptide CPLHQRPMC exhibited efficient binding to cells of the human metastatic cancer cell line PC3MM2 and significant antiproliferative activity.325 A particularly interesting approach to cancer cell targeting is subcellular targeting of individual organelles such as the nucleus. P(HPMA)−DOX conjugates targeted with the Tat-peptide (GRKKRRQRRR) originating from the HIV-1 Tat protein326 have been synthesized and shown to be capable of subcellular drug delivery in human ovarian carcinoma cells.327−329 Folic acid, a derivative of glutamic acid, is also often used to target various DDS toward fast-dividing cells.330 P(HPMA)-methotrexate or -Pt complexes conjugated to vitamin B12 or folic acid as a targeting moiety exhibited superior tumor growth inhibition relative to equivalent untargeted conjugates.211 Because folate receptors are overexpressed on many cancer cells, targeting using folic acid has been tested with many DDS (including nanoparticles, micelles, and liposomes43,331) intended for cancer treatment or diagnosis. Unfortunately, to our knowledge, none of these systems have been successful in clinical trials. A final class of LMW targeting moieties that merits discussion is bisphosphonates such as alendronate, which can be used to target P(HPMA)s to hard tissue (bone). An alendronate-targeted copolymer332 accumulated in bone in vivo, while similar HPMA conjugates bearing the radionuclides 125 I or 111In and DOX333 bound efficiently to hydroxyapatite and released DOX in a controlled fashion at a rate that depended on the detailed structure of its biodegradable spacer. While these results are promising, further in vivo studies will be required to properly assess the potential of this approach. The results discussed in this section demonstrate that active drug targeting is effective and can be very specific, with mAbs being the most potent targeting moieties. Polymer−mAb conjugates are effective at treating animal cancer models in vivo, but their structure tends to be poorly defined, and they are costly to prepare. The use of LMW targeting moieties therefore appears to be more practical: the resulting conjugates are structurally and chemically better defined and can be synthesized in larger batches. Unfortunately, switching from animal experiments to human clinical trials is complicated for both system types because the tumor and tissue receptors in animals and humans often differ significantly. Consequently, more studies will be required to support the design of polymer conjugates suitable for human treatment and to develop a better understanding of the relationship between conjugate’s structure and their biological behavior. 3.2.3. Drug-Free P(HPMA) Conjugates for Cancer Treatment. Drug-free macromolecular therapeutics are nanomedicines that possess specific biological activity without

having a LMW anticancer drug attached to the carrier.162 Such systems represent a new paradigm in nanomedicine and act by cross-linking cell surface receptors using coiled-coil peptide interactions to induce apoptosis and cancer cell death. They are administered in two steps: the first step entails treatment with an anti-CD20 antibody- or antibody fragmenttargeted conjugate bearing a coiled-coil forming oligopeptide, and the second involves treatment with a P(HPMA) polymer conjugated with a complementary oligopeptide. Self-assembly of the peptides into antiparallel helices then results in crosslinking of the CD20 receptors on the cancer cells’ surfaces, which induces apoptosis. More detailed information on this new and interesting strategy for cancer treatment can be found in recent publications from the Kopeček group.162,334,335 3.3. Other Polymer−Drug Conjugates

A large number of polymer−drug conjugates for cancer therapy have been synthesized using polymers other than PEG or P(HPMA)s, and their biological properties have been studied in vitro and also in vivo in many cases. These conjugates do not generally differ greatly from PEG or P(HPMA) derivatives with respect to the methods used for drug attachment and drug release (enzymatic, pH-controlled hydrolysis, or reduction) or with respect to the targeting moieties with which they are used. A full description of all systems that have been reported would be outside this review’s scope; the interested reader is directed to some recent reviews of the field for more information.92,94,99,162,336,337 Here we only discuss synthetic polymer systems that have entered clinical trials or exhibit promising behavior in vivo and, in our opinion, demonstrate some structural, conceptual, or strategic features that may be useful in the development of more potent anticancer DDS. Table 4 Table 4. Polymer−Drug Conjugates Based on Carriers Other Than P(HPMA) or PEG That Have Been Entered into Clinical Trialsa product CT-2106 Xyotax (Poliglumex) MTX-HSA DE-310 Delimothecan DOX-OXD a

polymer carrier

drug

polyglutamic acid polyglutamic acid

CPT PTX

serum albumin carboxymethyl dextran and polyalcohol carboxymethyl dextran dextran

Methotrexate CPT derivate CPT analogue Doxorubicin

phase

ref

I/II III II I

338 339, 340 341 342

I I

343 344

Adapted from ref 337.

shows the structures of some soluble polymer−drug conjugates based on carriers other than PEG or P(HPMA)s that have entered clinical trials. Interestingly, all of them are based on natural biodegradable polymers and their derivatives. The only carrier in this class to have been entered into Phase III trials is polyglutamic acid, P(Glu). 3.3.1. Polymer−Drug Conjugates with Biodegradable Polymer Chain. To be safe, DDS based on polymer carriers with nondegradable backbones must have a molecular weight below the renal threshold, which is typically around 50 kDa. However, much greater molecular weights are required for efficient drug accumulation in tumors via the EPR effect. The safe elimination of the polymer carrier system from the organism after it has fulfilled its task as a drug carrier will be essential in any polymer therapeutic that is to be used in routine medical practice. As discussed above, one way of 5361


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which were bound to the carrier in a synergistic ratio. The conjugate exhibited superior antitumor efficacy and safety in the treatment of orthotopic mammary adenocarcinoma in mice compared to a combination of the free drugs or single-drugbearing polymers.362 These results represent a valuable contribution to the discussion on combination therapy using polymer carrier systems. An interesting DOX-bearing P(Glu)based conjugate actively targeted to αVβ6 integrin-expressing cells with a tetrameric RGD-containing peptide was synthesized as a branched block copolymer of five undeca(oxyethylene) units with P(Glu).363 DOX was attached to the P(Glu) block via an acid-labile hydrazone bond. Specific internalization into αVβ6 positive cells was demonstrated, but unfortunately only in vitro data on this conjugate’s antitumor activity are available. Only a few water-soluble P(Glu) conjugates with anticancer drugs have been reported to be effective at treating in vivo over the last decade. However, numerous nanoparticle and micellar systems use P(Glu) amphiphilic block and graft copolymers to form micelles, which have been used effectively as drug carriers in cancer therapy. These systems are reviewed in the section dealing with polymer micelles and nanoparticles. There are also other poly(amino acid)- and polypeptidebased biodegradable DDS, some of which have been actively targeted (using cyclic RGD)364,365 and conjugated with DOX.366 While these carriers have shown promising in vivo antitumor effects in mice, with high long-term survival rates, it seems unlikely that they will match the anticancer activity of the P(Glu) carrier-based conjugates. 3.3.3. Acrylate and Vinyl Polymers and Copolymers. Many synthetic hydrophilic vinylic polymers and copolymers have been studied as water-soluble drug carriers. These carriers are typically based on substituted acrylic−methacrylic acid derivatives or on vinyl-group-containing monomers (e.g., vinylpyrrolidone). In principle, these polymers contain a functional group that allows other ligands (e.g., drugs and targeting moieties) to be attached directly or after functional group activation. Alternatively, such groups can be incorporated into the polymer structure by copolymerization with comonomers that contain suitable reactive groups, as is the case with HPMA copolymers. Some basic polymers of this kind exhibit therapeutic activity by themselves, including poly(N-vinylpyridine-N-oxide), poly(acrylic acid), and a divinylether-maleic anhydride copolymer. The others require modification with biologically active molecules. Some representative examples of these drug carriers and their conjugates are discussed below. A conjugate of a poly(styrene-co-maleic acid) copolymer with the protein drug neocarzinostatin (SMANCS) was the first polymer drug approved for the treatment of cancers (specifically, hepatocellular carcinoma) in humans.166,367,368 In this conjugate (Figure 26), two chains of the styrene-maleic acid copolymer are conjugated via an amide bond to neocarzinostatin, a protein with anticancer activity. This conjugate is not a typical water-soluble polymer prodrug; due to its hydrophobicity, it has to be administered in Lipiodol solution. SMANCS/Lipiodol has been shown to be effective both as a diagnostic tool and for therapeutic use in solid tumors in humans. Divinyl ether maleic anhydride (DIVEMA) is a copolymer containing anhydride groups (Figure 27) that allows it to undergo direct conjugation with ligands containing primary amino groups. This alternating copolymer has been conjugated with methotrexate369 and with enzymes such as SOD370,371 or TNFα. The conjugates with SOD exhibited significant anti-

overcoming this problem is to prepare HMW polymer micelles by self-assembly of smaller amphiphilic block copolymers or by introducing biodegradable linkages into the HMW polymer carrier structure. A potentially more viable strategy is to use carriers based on biodegradable natural macromolecules such as albumin, starch, hyaluronic acid, dextran, or pullulan or a fully degradable synthetic polymer carrier. More information about one class of natural drug carriers, the polysaccharides, can be found in some recent reviews and books.6,99 Biodegradable natural carriers aside, multiple water-soluble synthetic polymers with biodegradable backbones have been developed in the last two decades, including poly(amino acids), polyesters based on malic acid,345−348 polyacetals, and poly(amido amine)s. However, only a few have yet shown appreciable promise in practice. 3.3.2. Poly(amino acids). The most widely studied biodegradable poly(amino acid)-based drug carriers are based on poly(aspartic acid) (P(Asp)) or poly(glutamic acid) (P(Glu)) derivatives. They can be prepared by ring-opening polymerization of the corresponding amino acid N-carboxy anhydrides, enabling the synthesis of polymers with controlled molecular weights and narrow distributions.349 Their conjugation with drugs350−356 provides biodegradable drug conjugates that have been tested in the treatment of cancer. Conjugates of mitomycin C and N,N-di(2-chloroethyl)-4phenylenediamine mustard bound via GFG or GFAG oligopeptide spacers to a poly[N-5-(2-hydroxyethyl)-L-glutamine] (PHEG) carrier as well as similar conjugates with PEG side groups357 showed notable in vivo antitumor activity, inducing extended survival in animal models358 and lower side toxicity than the parent drugs. P(Glu) conjugates of PTX and CPT also exhibited significant in vivo activity in mice. These conjugates entered Phase I/II trials, and the PTX conjugate (Xyotax, Paclitaxel poliglumex) subsequently entered Phase III trials.359 This last conjugate (Figure 25) has proven to be the

Figure 25. Structure of Xyotax (Paclitaxel poliglumex), a polymer conjugate of PTX bound to the biodegradable P(Glu) polymer carrier via an ester linkage.

most potent polymer drug yet identified. Its PTX payload is attached to the γ-carboxylic groups of the Glu residues in the polymer backbone via ester bonds formed by esterification of the drug’s 2′-OH group.360 The conjugate is known to undergo intracellular enzymatic degradation mediated by cathepsin B to form a diglutamyl-PTX product. In clinical trials, this conjugate significantly increased survival among patients with advanced nonsmall-cell lung and ovarian cancer, and there are high hopes that it will replicate this performance in Phase III studies.359,361 A P(Glu) carrier was also used for codelivery of DOX and PTX, 5362


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Figure 26. Representation of the first polymer drug conjugate approved for cancer treatment in humans (SMANCS).

potential for renal drug targeting.376 However, most research on P(VA) and P(VP) in the context of DDS has focused on their applications in micellar drug delivery. Among the carriers based on vinylic polymers, thermoresponsive polymer drug carrier systems are particularly interesting because of their responsiveness to external stimuli. At physiological temperature these polymers are fully water soluble, but at temperatures above their lower critical solution temperature they undergo phase separation. Such temperatures may be encountered at sites of local inflammation or hyperthermia, causing the carrier to come out of solution and form nonsoluble deposits that slowly release the covalently bound drug while the polymer slowly dissolves; this has been demonstrated in a system featuring DOX bound to an isopropyl and n-propyl methacrylamide copolymer.377 Thermoresponsive carrier systems may have specialized uses in local radiotherapy,378 but their major area of application is likely to be in thermoresponsive micellar systems for drug and gene delivery and tissue engineering.379−381 3.3.4. Poly(amido amine)s. Poly(amido amine)s are biocompatible and biodegradable synthetic polymers that were developed to facilitate the design of new biomedical materials and polymer therapeutics. They have been prepared by stepwise polyaddition of primary or secondary aliphatic amines to bis(acrylamide)s382 to form linear or branched products. Properly controlled polyaddition reactions of bis(acrylamide)s with equimolar quantities of diamine derivatives generate linear water-soluble polymeric carriers containing functional groups suitable for drug attachment.383−385 Hydrolytic degradation of the main chain of poly(amido amine)s forms small degradation products that are easily excreted, e.g., by glomerular filtration. Although poly(amido amine)s have been employed as a carrier of anticancer drugs such as mitomycin C,386 melitin,387 and platinates388−390 they are most likely to be important as dendritic polycationic polymer materials for the design of synthetic gene delivery vectors.382,391

Figure 27. Structure of polymer−drug carriers (precursors), copolymer DIVEMA (left) and N-vinylpyrrolidone-maleic anhydride copolymer (right).

inflammatory effects in rats with fibrotic livers, while the DIVEMA-TNFα conjugate372 showed significant necrotic effects on Sarcoma-180 solid tumors in mice and dramatic antitumor effects in mice bearing Meth-A solid tumors, causing complete tumor regression with no apparent side effects.372 However, there have been no further reports on the development of DIVEMA−anticancer drug conjugates, and it seems that this polymer carrier will not play a substantial role in the future development of polymer−drug conjugates. Poly(vinyl alcohol) (P(VA)) is a nonionic and nontoxic polymer that is well suited for use as a drug carrier. A conjugate of P(VA) with DOX bound to the carrier via a pH-responsive cis(trans)-aconityl spacer373 was synthesized by the reaction of the cis(trans)-aconityl derivative of DOX with P(VA) after derivatization with ethylene diamine using the carbodiimide method. The cis conjugate released DOX in mildly acidic environments much more quickly than the trans conjugate and also showed superior in vitro antitumor activity. Vinylpyrrolidone copolymers are nonionic and nontoxic polymers with antifouling properties that have been tested as drug carriers. Poly(vinylpyrrolidone-co-dimethylmaleic anhydride) (Figure 27) was developed as a pH-sensitive polymeric carrier capable of releasing a native drug with full activity in response to changes in the surrounding pH. Conjugates of this polymer with DOX374 or tumor necrosis factor TNFα375 reportedly have significant antitumor activity and offer the 5363


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3.3.5. Polyacetals. Water-soluble polyacetals prepared by the polyaddition of PEG and oligo(ethylene oxide) divinylether were recently identified as promising biodegradable carriers for anticancer drugs.392,393 These polymers are relatively stable at physiological pH (7.4) but undergo fast hydrolysis in mildly acidic environments to form small nontoxic degradation products that are removable from the organism by glomerular filtration. A terpolymer of PEG with divinyl ethers and serinol was used for conjugation with DOX,393 which was attached to the polymer with a succinic acid residue-containing spacer. The conjugates, which had M w values of ∼100 kDa and polydispersity values of 1.7−2.6, had a long half-life in blood and exhibited superior tumor accumulation compared to PK1 (a HPMA copolymer−DOX conjugate) as well as lower uptake by the liver and spleen. Unfortunately, it did not exhibit detectable in vivo anticancer activity. More successful was a conjugate of a biodegradable bioinert polyacetal (Fleximer) with CPT (XMT-1001) (for structure see Figure 28), which

with diameters of 10−100 nm. The hydrophobic blocks form the micelles’ cores and are surrounded by a hydrophilic polymer shell that stabilizes the micelles in aqueous solutions and protects them from undesired interactions with proteins in body fluids and specific cells and tissues in the living body (e.g., cells of the immune system or RES and uptake by the liver). Consequently, such micelles have the potential to be good drug carriers, especially for anticancer drugs, because their size is such that they have a strong EPR effect. The formation of a micelle-based drug carrier is outlined in Figure 29.

Figure 29. Formation of polymer micelle drug carriers. Diblock or multiblock amphiphilic copolymers self-assemble into spherical nanoparticles with hydrophobic cores and hydrophilic shells. The hydrophobic drug is typically entrapped in the micelle’s core by hydrophobic interactions, and the micelle can be targeted to a desired location by attaching an appropriate targeting moiety to the shell.

Figure 28. Structure of the conjugate XMT-1001.

was designed for the treatment of solid tumors and has entered into Phase I/II clinical studies.394 This conjugate reportedly has favorable pharmacokinetics, safety, and potential for therapeutic activity against various tumors. Its ongoing evaluation will reveal its true potential in cancer treatment. 3.3.6. Poly(2-oxazolines). Poly(2-oxazolines) are a very recently developed alternative to PEG conjugates in drug delivery. These water-soluble polymers have narrow molecular weight distributions and can be synthesized by ring-opening polymerization, making it possible to exercise complete control over their molecular weight. Their potential as drug carriers was evaluated by conjugation with selected enzymes (ribonuclease, catalase, uricase), which yielded promising results.395 Poly(2ethyl-2-oxazoline) (PEtOZ) conjugates with proteins and AraC have been prepared by using an amide bond to link the drug to the polymer.396 The resulting conjugate released Ara-C slowly and exhibited cytotoxicity comparable to that of a similar PEG conjugate. Applications of these polymers in biomedicine have been reported,397 and they will undoubtedly be tested further as anticancer drug carriers. In vivo studies will be needed to establish their true potential for drug delivery.

The micelles are formed if the concentration of the copolymers (unimers) exceeds the critical micelle concentration (CMC) and are in equilibrium with the unimers that depends on the temperature, polymer concentration, pH, and ionic strength of the solution. Consequently, they can be described as dynamic polymeric micelles. However, if the exchange of copolymers between unimers and micelles is slow compared to the experimental time scale (meaning that the polymer chains are kinetically trapped), the aggregates are called frozen micelles.398 The micelles’ stability depends strongly on the Tg of the hydrophobic polymer block forming the core of the system. Spherical micelles are formed from block copolymers with longer hydrophilic segments, while longer hydrophobic segments usually generate nonspherical rods and lamellae.399 At concentrations below the CMC, the micelles disassemble to the original unimers, whose properties are usually tailored to make them excretable by glomerular filtration. The CMC values for polymeric micelles are low, making them relatively stable at the concentrations used in DDS treatments. Polymeric micelles are designed to transport bioactive molecules, both hydrophilic and hydrophobic, to tumor tissues while avoiding the induction of immune responses and nonspecific drug distribution to normal tissues.400,401 They can be used as universal delivery systems that facilitate solubilization and enable tumor-specific delivery of hydrophobic molecules that are otherwise insoluble in water.

4. POLYMERIC MICELLES Advances in the synthesis of amphiphilic block copolymers with narrow molecular weight distribution and polymer blocks of well-defined structure have facilitated the preparation of micellar drug carrier systems by self-assembly of the copolymers in aqueous solution to generate polymeric micelles 5364


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solvent but not in the second. Nanoprecipitation is achieved by adding the polymer solution to the nonsolvent, which induces rapid desolvation of the polymer. The amphiphilic triblock copolymer poly(acrylic acid)-poly(ε-caprolactone)-poly(acrylic acid) was used to prepare kinetically stable micelles by a modified nanoprecipitation procedure, yielding core−shell micelles with diameters of 120−140 nm, a negative zeta potential, and satisfactory loadings of the hydrophobic drug (7ethyl-10-hydroxy camptothecin).417 In the dialysis method, the amphiphilic block copolymer and drug are dissolved in a water-miscible solvent and introduced into a dialysis tube. Micelles are formed upon dialysis against water. In the thin f ilm method, the amphiphilic block copolymer is dissolved in a good solvent (e.g., acetonitrile) in a roundbottom flask, the solvent is evaporated by rotary evaporation, and micelles are formed from the film by adding a saline solution before intense vortexing. The influence of different micelle preparation methods on particle size and drug loading was discussed by Tyrell et al.414

Polymeric micelle-based delivery systems can be divided into two classes depending on the method of drug loading. In systems of the first class, the drug is loaded in the micelle’s hydrophobic core via hydrophobic interactions. This approach can be used to load any hydrophobic drug regardless of its chemical structure or content of specific functional groups. In systems of the second class, the drug is conjugated with the amphiphilic polymer via a covalent bond.402 Entrapment by physical forces (e.g., hydrophobic interactions) is generally achieved during micelle formation. In contrast, covalently conjugated drugs are bound to the micelle core until the micelles have accumulated at the site of action, where they are exposed to stimuli such as changes in the ionic strength, pH, or abundance of endogenous signal peptides or enzymes. These stimuli then trigger micelle disassembly and drug release. Micelles bearing covalently conjugated drugs appear to be more stable than those with physically entrapped drugs, while the covalent bond remains intact. Moreover, their drug release patterns can be modified by varying the properties of the drugbinding linker. Covalent conjugation therefore enables environment-responsive controlled drug release.402,403 4.1. Preparation of Micelles and Loading of Drugs

4.2. Amphiphilic PEG-Based Copolymers as Micellar Drug Carriers

It has been reported that about 40% of the drugs being developed by the pharmaceutical industry are poorly watersoluble molecules.404 Water insolubility has always been a key obstacle in pharmaceutical formulations, affecting formulation stability and drug bioavailability. The solubilizing agents currently used to prepare drug formulations are often not effective enough and may even be toxic or have harmful side effects. Cremofor EL, which is used in injectable formulations of paclitaxel, cyclosporine A, and other drugs, causes hypersensitive reactions, hyperlipidemia, and neurotoxicity. Similarly, Tween 80 and sodium deoxycholate, which are used to solubilize amphotericin B, are hemolytic.405,406 Nanoparticle systems based on polymeric micelles 407−409 or liposomes410−412 represent very promising alternatives to these solubilizing agents. The self-assembly of preformed amphiphilic copolymers is the most convenient method for preparing polymeric micelles. However, the choice of micelle preparation method will depend on the polymeric material and its physicochemical properties. The final drug loading capacity of the micellar system will be determined by the micellar core’s compatibility with the drug to be encapsulated. The self-assembly of amphiphilic block copolymers in aqueous solution is driven by differences in the water affinity of the individual blocks. The most widely used methods for micelle preparation are10,11,413,414 (i) emulsion− solvent evaporation, (ii) nanoprecipitation, (iii) dialysis, and (iv) the thin film method.415 In the emulsion−solvent evaporation process, a waterimmiscible organic solution of a hydrophobic drug and the preformed amphiphilic block copolymer is added into an aqueous phase containing a surfactant under vigorous stirring to obtain nanosized organic solvent droplets that serve as templates for nanocarrier assembly. The organic solvent is subsequently removed under reduced pressure. This method has been used to prepare micelles from the amphiphilic pentablock copolymer P(LGA)-PEG-PPO-PEG-P(LGA) in which the hydrophobic drug docetaxel was encapsulated at a loading yield of around 84%.416 The nanoprecipitation technique uses two miscible solvents. The block copolymer and the drug must be soluble in the first

The polymerization technique used to prepare an amphiphilic block copolymer should be selected based on the polymer’s desired properties and intended use. Many polymerization techniques can be used for this purpose including anionic polymerization, ring-opening polymerization (ROP), atom transfer radical polymerization (ATRP), and reversible addition−fragmentation chain-transfer (RAFT)418−420 polymerization. In addition to traditional AB- and ABA-type block copolymers, ABC triblock copolymers (or block terpolymers) and copolymers of various architectures (star, core, or shell cross-linked, etc.) have been synthesized and used for micelle preparation. The hydrophilic blocks of most amphiphilic block copolymers designed for the preparation of micellar DDS consist of poly(ethylene oxide) (PEG). PEGs used to initiate the polymerization of the monomers that form the hydrophobic segments of such block copolymers typically have α-methoxy and ω-amino or hydroxyl groups. 4.2.1. PEG-b-poly(amino acid) Micelles as DDS. Although many types of amphiphilic block copolymers could potentially form micellar structures, poly(ethylene oxide)-bpoly(amino acid) (PEG-b-P(AA)) block copolymers have been most extensively studied in this context.409,421,422 N-Carboxy anhydride (NCA) derivatives of aspartic acid, glutamic acid, and lysine are generally regarded as the most appropriate precursors for the synthesis of such copolymers, yielding PEGb-poly(β-benzyl-L-aspartate) (PEG-b-P(BLA)), PEG-b-poly(γbenzyl-L-glutamate), or PEG-b-(ε-benzyloxycarbonyl-L-lysine) block copolymers, respectively. An α-methoxy ω-amino PEG is generally used as a macroinitiator for the ring-opening polymerization of these amino acids; PEG derivatives with molecular weights between 2000 and 20 000 g mol−1 have been successfully used for this purpose.423 4.2.1.1. Nontargeted PEG-b-P(AA) Micelles. PEG-b-P(AA) micelles were designed to achieve prolonged drug circulation and enhanced accumulation in solid tumors. In early studies, DOX molecules were incorporated into poly(ethylene glycol)b-poly(β-benzyl-L-aspartate) (PEG-b-P(BLA)) polymeric micelles via physical entrapment. The preparation of PEG-bP(BLA) micelles with diameters of approximately 50−70 nm 5365


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Figure 30. Structure of micellar NPs formed by conjugate of PEG-b-P(Asp) with DOX.422,425,427

HT29 human colonic cancer cell line; PTX micelles administered at a PTX-equivalent dose of 25 mg kg−1 exhibited activity comparable to that achieved by free PTX at a dose of 100 mg kg−1.429,430 Importantly, some PEG-b-P(AA) micellar DDS have been evaluated in clinical trials. Micelles based on the block copolymer poly(ethylene glycol)-b-poly(glutamic acid) (PEGb-P(Glu)) were covalently bound to the drug 7-ethyl-10hydroxy-camptothecin (SN-38; see Figure 31), an active

and DOX loadings of 15−20 wt % achieved by physical entrapment demonstrated the potential of polymeric micelles as drug carriers, but the micelles’ stability was low due to the poor DOX affinity of the P(BLA) segment.424 The hydrophobic blocks of PEG-b-P(AA)s often contain functional groups that may be derivatized to improve the properties of the core-forming blocks with respect to drug delivery. Chemical modification of the core-forming block of poly(ethylene glycol)-b-poly(α,ß-aspartic acid) (PEG-b-P(Asp)) with a hydrazine enabled the conjugation of the anticancer drug DOX to the hydrophobic segments of polymeric micelles via pH-sensitive hydrazone bonds (see Figure 30) that are stable under physiological conditions at pH 7.4 but release DOX selectively at the acidic pH values (4−6) occurring in endosomes and lysosomes. The intracellular pH thus controls the systemic, local, and subcellular distributions of DOX.422,425,426 In vivo studies showed that these micelles with DOX bound via hydrazone bonds released the free drug in the intracellular environment and had effective antitumor activity with remarkably low levels of side toxicity.422,425,427 PEG-b-P(Asp) micelles have also been investigated as carriers of physically entrapped docetaxel, paclitaxel, and cisdichlorodiammine platinum(II) (CDDPt). Docetaxel-loaded micelles based on PEG-b-poly(racemic-leucine) (P(RLeu)) block copolymers were prepared by dialysis, yielding particles whose size could be varied between 170 and 250 nm by changing the length of the hydrophobic blocks. Regardless of their size, the loaded micelles exhibited similar release behavior, with a sustained drug release over up to 72 h, and were more active than the free drug toward MCF-7 human breast cancer cells.428 Polymeric micellar particles loaded with PTX were formed in an aqueous medium by facilitating the self-association of amphiphilic block copolymers having hydrophilic PEG segments and modified polyaspartate hydrophobic segments in which one-half of the carboxylic groups were esterified with 4phenyl-1-butanol. PTX was incorporated into the inner core of the micelle system by physical entrapment through hydrophobic interactions between the drug and the specially designed hydrophobic block. The antitumor activity of the PTX-loaded micelles was evaluated in vivo on nude mice implanted with the

Figure 31. Structure of a poly(ethylene glycol)-b-poly(glutamic acid) copolymer that is covalently conjugated with 7-ethyl-10-hydroxycamptothecin (SN-38).

metabolite of irinotecan (CPT-11) that is up to 1000-fold more active than the parent drug against various cancer cell lines in vitro.431,432 The drug was conjugated to the polymer by a 1,3-diisopropylcarbodiimide-mediated condensation reaction between the carboxylic acid moieties of P(Glu) and the phenolic −OH group of SN-38. The polymer-bound SN-38 and SN-38 cleaved from the micelles by the hydrolysis of the phenolic ester linkage under physiological conditions were slowly eliminated from the plasma with half-lives of approximately 140 and 210 h, respectively. Patients with a refractory esophageal cancer and a lung carcinoid tumor exhibited objective responses to the loaded micellar DDS over treatment periods of 5 and 12 months, respectively.431,432 5366


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Platinum-based anticancer agents such as CDDPt and carboplatin have many problematic side effects when used in the clinic. A range of delivery vehicles for platinum-based therapeutics including Pt−polymer complexes, dendrimers, micelles, and microparticles have been investigated, partly with the aim of alleviating these side effects; the synthesis and properties of these systems have been reviewed extensively.433,434 Micelles containing CDDPt in the core with diameters of approximately 20 nm and narrow size distributions were prepared by the complexation of CDDPt with a PEG-bP(Asp) block copolymer in an aqueous medium. The release rate of CDDPt was inversely correlated with the chain length of the copolymer’s P(Asp) segments.435−437 Moreover, micelles based on a PEG-b-P(Asp) block copolymer conjugated to CDDPt exhibited similar in vivo activity to free CDDPt against MKN 45 human gastric cancer xenografts but caused significantly less CDDPt-induced nephrotoxicity and neurotoxicity.438,439 Poly(ethylene glycol)−poly(glutamic acid) block copolymers (PEG-b-P(Glu)) with P(Glu) blocks of different lengths were mixed with dichloro(1,2-diaminocyclohexane)platinum(II) (DACHPt) in distilled water to prepare micelles with dimensions of around 40 nm bearing an entrapped Pt complex. The micelles’ strongly hydrophobic cores made them very stable: they retained their initial size for over 240 h when stored in PBS buffer. However, after an induction period of 15 h, the entrapped DACHPt complexes were gradually released from the micellar core. In vivo biodistribution and antitumor activity experiments in CDF1 mice bearing subcutaneously inoculated murine colon adenocarcinoma C-26 showed that mice treated with the micelles exhibited 20-fold greater accumulation of DACHPt in their tumors than was achieved with free oxaliplatin. Micelles prepared from PEG-b-P(Glu) copolymers with short P(Glu) blocks exhibited low nonspecific accumulation in the liver and spleen, resulting in greater specificity for solid tumors.440,441 A total of 17 patients were enrolled in a Phase I clinical study of these CDDPt-bearing PEG-b-P(Glu) micelles, which were administered to the participants intravenously in 3 week cycles. The maximum tolerated dose was 120 mg m−2 CDDPt, at which point renal impairment and hypersensitivity reactions were observed. The recommended dose was thus set to 90 mg m−2. Stable disease was observed in seven of the patients with only very mild toxicity.442 Despite this partial success, there have been no reports on subsequent trials with these micelles or on their clinical use. 4.2.1.2. Actively Targeted PEG-b-P(AA) Micelles. The use of NP carriers to deliver pharmaceuticals into brain tumors is very challenging because of the need to cross the tight endothelium of the blood−brain barrier (BBB).443 For large molecules, this obstacle can be overcome by exploiting receptor-mediated transport. To this end, peptide-modified polymeric micelles bearing the cyclic Arg-Gly-Asp, c(RGD) or c(RGDfK) structure were developed to deliver platinum anticancer drugs through the BBB to glioblastomas (Figure 32).444 The c(RGDfK) peptide has a high affinity for the αvβ3 and αvβ5 integrins, which are overexpressed on GBM cells and the endothelial cells of angiogenic vessels serving tumors. PEG-bP(Glu) micelles with a diameter of around 30 nm and a c(RGDfK) content of 5−40% were prepared by conjugating a cysteine-ε-aminocaproic acid derivative of c(RGDfK) onto maleimide-functionalized DACHPt-PEG-b-P(Glu) micelles obtained by mixing PEG-b-P(Glu) and maleimide-containing

Figure 32. Preparation of c(RGDfK)-targeted PEG-b-P(Glu) micelles loaded with DACHPt for drug delivery to glioblastomas.444 Similar micelles targeted by the Fab’ fragment of a human anti-transferrin Ab were used to target pancreatic tumor xenografts. Reprinted with permission from ref 445. Copyright 2013 American Chemical Society.

MI-PEG-b-P(Glu). DACHPt-bearing micelles decorated with 20% c(RGD) exhibited a significantly higher tumor growth inhibitory effect in mice inoculated with U87MG tumor cells than micelles targeted with the nonspecific cyclic-Arg-Ala-Asp peptide444 at the same loading, verifying the effectiveness of this strategy for transporting micellar vehicles across the BBB. The effectiveness of Ab fragments as targeting moieties for PEG-b-P(Glu) micelles is exemplified by the successful treatment of human pancreatic cancer BxPC3 in BALB/c nu/ nu mice using similar DACHPt-PEG-b-P(Glu) micelles to those described above with a thiol-Fab’ fragment of the human anti-transferrin Ab in place of the c(RGD) targeting peptide. The Fab’-targeted DACHPt micelles exhibited greater in vitro toxicity than the corresponding nontargeted micelles, in part because of a 15-fold increase in cellular binding and more rapid cellular internalization. In vivo, the Fab’-targeted micelles significantly suppressed the growth of pancreatic tumor xenografts for more than 40 days.445 From these experiments follows that both targeting moieties, c(RGDfK)/c(RGD) and specific antibody fragments, are promising candidates for targeting of micellar DDS to selected tumors. 4.2.2. PEG−Polyester Amphiphilic Copolymers and Micelles. An alternative class of amphiphilic block copolymers can be formed by combining hydrophilic PEG block(s) with biocompatible and biodegradable hydrophobic polyester block(s)446 based on materials such as poly(lactic acid) (P(LA)),447−449 poly(glycolic acid) (P(GA)) poly(lactide-coglycolide) (P(LGA)),450 poly(ε-caprolactone)(P(CL)),451,452 or poly(δ-valerolactone).453 Ring-opening polymerization in the presence of stannous octoate (Figure 33) is a standard method for synthesizing these copolymers. This method has also been used to prepare the isotactic stereoisomeric block copolymers PEG-b-poly(L-lactide) (PEG-b-P(LLA)) and PEG5367


Chemical Reviews

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Figure 33. Scheme of the synthesis of PEG-b-polyester (PLA) copolymers.

175 mg m−2 PTX equivalents of Genexol every 3 weeks, preoperatively, unless they developed profound side effects or disease progression. After curative surgery, two additional cycles of chemotherapy were administered to patients who had shown a positive response to the neoadjuvant chemotherapy. Forty-seven of the fifty patients underwent all four cycles of designated treatment. Complete disappearance of invasive foci of the primary tumor and negative axillary lymph nodes were confirmed in eight patients (16%) postoperation. The cumulative 5-year disease-free survival rate was 70% for patients with complete remission and partial remission and 33.3% for patients with stable disease and progressive disease.459 GenexolPM also has ongoing Phase II trials against pancreatic cancer,460 ovarian cancer,461 gastric cancer,462 and nonsmall-cell lung cancer.463 Curcumin (Cur) is a very hydrophobic molecule with antiangiogenic and apoptosis-inducing effects. Its clinical applicability is limited by its poor solubility in water, which suggests that it may be a good candidate for micellar delivery. Cur-loaded biodegradable mPEG-b-P(LA) micelles with diameters of ∼30 nm exhibited promising antiangiogenic and antitumor effects, inducing apoptosis when used to treat lung metastasis in a murine colon cancer model. Notably, the Curloaded micelles had a significantly stronger inhibitory effect on both the lung metastases and the original tumors than could be achieved with an equivalent dose of free Cur. This demonstrates the potential of micellar DDS to improve the solubility of hydrophobic drugs and their ability to exert antitumor effects on metastatic tumors in vivo.464 Modification of the Cur-loaded PEG-b-P(LA) micelles with the RGD peptide, which binds strongly to αvß3 integrin-expressing cells, significantly increased the cellular uptake of the micellar system by human umbilical vein endothelial cells (HUVEC) and mouse melanoma cells (B16). In B16 tumor-bearing mice, the Cur-loaded RGD-PEG-b-P(LA) micelles inhibited tumor growth more strongly than the corresponding drug-loaded nontargeted micelles,465 demonstrating that active targeting can have important effects on the delivery of micellar systems. 4.2.3.1. Active Targeting of PEG−Polyester Micelles. Many human cancer cells express the folate receptor (FR) on their surfaces, so folic acid (FA) is very widely used as a targeting moiety for delivering DDS to cancer cells and triggering their cellular uptake by endocytosis. FA-targeted biodegradable polymeric micelles loaded with DOX have been prepared by self-assembly from a PEG-b-P(LGA) diblock copolymer in which DOX was chemically conjugated to the terminal end of the P(LGA) block and FA was conjugated to the terminus of the PEG chain. Micelles were obtained by mixing PEG-bP(LGA)-DOX and FA-PEG-b-P(LGA) diblock copolymers and then tested against KB cells from a human epidermal carcinoma xenograft cell line that were subcutaneously implanted in female athymic nude mice. The FA-targeted DOX-containing micelles were efficiently accumulated in the tumor and suppressed tumor growth more strongly than free DOX or nontargeted PEG-b-P(LGA)-DOX micelles.466

b-poly(D-lactide) (PEG-b-P(DLA)) in which P(LLA) and P(DLA) form isotactic semicrystalline hydrophobic blocks while poly(D,L-lactide) (P(DLLA)) is atactic and amorphous.454,455 Micelle biodegradability is primarily determined by the composition of the hydrophobic block and the tacticity and crystallinity of the hydrophobic polymer or copolymer. Studies on the effects of the macromolecular architecture of amphiphilic PEG-b-P(LA) copolymers on micelle formation and the micelles’ cellular uptake showed that the micelles’ interactions with biological systems depend on both the size of the PEG block and the structure of the PLA block. It was also shown that mixed micellar solutions containing both PEG-bP(LLA) and PEG-b-P(DLA) copolymers exhibited greater solution stability (i.e., had lower CMC values) than solutions of the separate copolymers. This may indicate that P(LLA) and P(DLA) blocks interact more strongly with one another than with blocks of their own chirality.455 Solution stability has important effects on micelle disassembly and drug release. 4.2.3. PEG−Polyester Micelles as DDS. An interesting passively targeted micellar DDS was prepared by encapsulating DOX noncovalently in monomethoxy PEG (mPEG)-b-P(CL) micelles. The antitumor activity of this system was evaluated in vitro and in vivo in C57BL/6 mice inoculated with B16-F10 melanoma cells. Encapsulation in the micelles increased the cytotoxicity of DOX and enhanced its cellular uptake by B16F10 cells in vitro. In vivo, the DOX-loaded micelles significantly inhibited the growth of aggressive B16-F10 tumor xenografts (achieving 75% of maximum inhibition relative to controls) and markedly prolonged the survival of the treated mice. The antitumor response was accompanied by a significant increase in tumor cell apoptosis and a noticeable reduction in cell proliferation. Moreover, less systemic toxicity was observed for the group treated with DOX-loaded mPEG-b-P(CL) than for that treated with free DOX.456 A number of other drugs have also been encapsulated in micellar DDS, notably the taxols. We will not discuss the performance of these systems at length because the in vivo results obtained with PTX or docetaxel-loaded polyester-based micelles, and nanoparticles have been reviewed comprehensively by Gaucher et al.457 We therefore restrict our discussion to the one micellar system that has been approved for clinical use to date. Genexol-PM, the first micelle-based DDS approved in Europe, was developed by the South Korean company Samyang Co. It is a PTX-containing PEG-b-P(DLLA) micelle formulation with micelle sizes of 20−50 nm. A multicenter Phase II clinical trial was conducted to evaluate its safety and efficacy in metastatic breast cancer patients. The micelles were administered to 41 patients at a dosage of 300 mg m−2 PTX equivalents by iv infusion over 3 h every 3 weeks without premedication. The overall response rate was 58.5%, with 5 complete and 19 partial responses. The median time to progression for all patients was 9.0 months. No febrile neutropenia was observed in any of the participants.458 Fifty patients with locally advanced breast cancer were then enrolled in a multicenter Phase II study. All of these patients were scheduled to receive four cycles of 60 mg m−2 of epirubicin and 5368


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antitumor effect in S-180 sarcoma-bearing mice than any other tested treatment, and histological studies showed that it induced more pronounced cytoarchitectural changes associated with tumor regression than the other treatments.415 4.2.4. Polyethers (Poloxamers). Polyethers are another class of polymers that can be used to prepare amphiphilic micelles for the dissolution and delivery of hydrophobic drugs. Most polyethers of pharmaceutical interest belong to the poloxamer family, i.e., they are copolymers of poly(ethylene oxide)-b-poly(propylene oxide)-b-poly(ethylene oxide), which is abbreviated (PEG-b-PPO-b-PEG).473 Poloxamers are synthesized by sequential alkaline-catalyzed polymerization. The initiation step of the triblock synthesis is polymerization of the PPO block, after which the ends of the PPO chain are extended with PEG chains. The lengths of each block can be varied independently, offering wide scope for tuning the properties of the final copolymer. In aqueous solutions at concentrations above the CMC (from 1 mM to 1 μM), these copolymers self-assemble into micelles with diameters that usually vary between 10 and 100 nm.474 Poloxamer block copolymers have been identified as efficient DDS that have multiple beneficial effects. Specifically, the incorporation of hydrophobic drugs into the micellar core increases their solubility and metabolic stability as well as the drug’s circulation time in the body. Poloxamers have been shown to preferentially target cancer cells due to their distinctive membrane properties compared to normal cells. Multidrug-resistant cancer cells are sensitized by their interactions with the poloxamer unimers, inhibiting the action of the proteins responsible for multidrug resistance and other drug efflux transporters expressed on cancer cells’ surface. Consequently, the micelles increase the cancer cells’ susceptibility to chemotherapeutic agents such as DOX and other cancerostatics. The unimers also inhibit drug efflux transporters in both the blood−brain barrier and the small intestine, enhancing the transport of the drugs into the brain and their oral bioavailability. Poloxamers also inhibit ATP production in multidrug-resistant cancer cells by inhibiting respiratory proteins I and IV and have been shown to enhance proto-apoptotic signaling, suppress antiapoptotic defense in MDR cells, inhibit the glutathione/glutathione S-transferase detoxification system, induce the release of cytochrome C, increase the concentration of reactive oxygen species in the cytoplasm, and abolish drug sequestration within cytoplasmic vesicles. Several reviews have been published in the last decade on their applications in drug delivery473,475 and gene delivery476 and as tools for overcoming multidrug resistance.477,478

In another study, FA-targeted mixed pH-sensitive polymeric micelles loaded with DOX were formed from two block copolymers, PEG-b-P(LLA)-FA and the poly(L-histidine) block copolymer PEG-b-P(His). The antimetastatic activity of these micelles was examined in mice inoculated with murine mammary 4T1 carcinoma, which is one of the most aggressive metastatic cancer cell lines. The mice were treated with the equivalent of 10 mg kg−1 DOX each in 4 doses with a 3-day interval between doses. The FA-targeted micelle formulation resulted in retarded tumor growth, no weight loss, no death for 4−5 weeks, and no apparent metastasis after 28 days of treatment. However, significant metastases to the lung and heart were found on day 28 in mice treated with free DOX in PBS or with nontargeted PEG-b-P(LLA)/PEG-b-P(His) micelles.467 Both examples given above document the suitability of FA as efficient targeting moiety in DDS designed for delivery of drugs to specific tumors. 4.2.3.2. PEG−Polyester Micelles in Combination Therapy. Unlike soluble drug conjugates, micellar DDS can be used to deliver hydrophobic molecules such as drugs even if the delivered molecule does not contain specific reactive functional groups amenable to covalent bond formation. Moreover, micelles could in principle be used to codeliver multiple hydrophobic drugs that are encapsulated in the hydrophobic core or one hydrophobic drug that is entrapped in the core and a hydrophilic drug that is covalently attached to the hydrophilic block of the amphiphilic copolymer. This makes micelles good candidates for the codelivery of multiple molecules having different structures and origins, which is useful in combination therapy and theranostics. Table 5 lists some PEG−polyester Table 5. Multidrug Polymeric PEG−Polyester Micelles for Concurrent Drug Delivery polymer

drugs

indication

status

ref

PEG-bP(DLLA) PEG-bP(CL)

PTX/17-AAG/ rapamycin PTX/ cyclopamine/ gossypol PTX/17-AAG/ rapamycin

A549 NSCLC and MDAMB-231breast cancers ES-2-luc and SKOV-3-luc ovarian cancers

in vivo in vivo

468, 469 470

ES-2-luc ovarian cancer

in vivo

471

DOX/PTX

A549 NSCLC, B16 mouse melanoma, and HepG2 liver cancers

in vitro

472

P(LGA)-bPEG-bP(LGA) PEG-bP(LGA)

micelles intended for codelivery of drug combinations such as PTX with a derivative of geldanamycin (17-AAG) and rapamycin, PTX with DOX, and PTX with cyclopamine and gossypol. Micellar systems significantly increase the aqueous solubility of the encapsulated drugs, often resulting in a durable antitumor response with acceptable acute toxicity and enhanced PTX killing. The potential of combination therapy is illustrated by the example of mPEG-b-P(LA) micelles. Polymer micelles with a size of 20 nm were prepared from an mPEG-b-P(LA) copolymer and used for sequential codelivery of the chemosensitizer resveratrol (RES) and PTX with the aim of achieving synergistic anticancer activity. Micelles loaded with both drugs together were prepared by the thin film method and used to treat mice bearing lung adenocarcinoma A549/T and S-180 sarcoma cells. Synergistic effects against both cancer cell types were observed, and the micellar formulation exhibited very limited toxicity toward normal human hepatic and kidney cells. Moreover, the combination therapy achieved a stronger

4.3. Polymer Micelles Prepared from Other Amphiphilic Block Copolymers

In addition to poly(amino acids) and polyesters, amphiphilic block copolymers have been synthesized using with hydrophobic core segments consisting of vinylic polymers such as poly(ethyl acrylate), poly(n-butyl acrylate), or poly(tert-butyl methacrylate). The use of such polymers in DDS has been reviewed previously.418,454,479 Similarly, while PEG is most commonly used as the hydrophilic segment, alternative nonionic biocompatible and water-soluble synthetic polymers exist including P(VP) and P(HPMA); both have been used in the synthesis of amphiphilic copolymers. The synthesis of P(VP)-b-poly(D,L-lactide) block copolymers and their use in micelle preparation was described by Luo et al.480 The P(VP) blocks interact with a variety of hydrophilic and hydrophobic pharmaceutical agents, increasing the micelles’ solubilizing 5369


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capacity.481 P(HPMA) is another hydrophilic polymer that could potentially replace PEG in micelle-forming copolymers. It is readily amenable to functionalization, enabling the attachment of various drugs and targeting moieties to the micelle’s hydrophilic shell; this should facilitate the development of new micelle-based technologies and structures. The uses of P(HPMA) and its role in the field of polymeric micelles have been reviewed;482 notably, it can be used as a micellar “stealth corona”, as a stabilizing unit in core-cross-linked micellar structures,483 or for covalent core entrapment of drugs. Micelles incorporating P(HPMA) blocks exhibit good stability and drug retention in circulation, efficient tumor accumulation, and tailored drug release at the target site.483 An interesting strategy for preparing well-defined P(HPMA)b-P(LLA) and P(HPMA)-b-P(DLLA) block copolymers was based on a combination of ring-opening polymerization of L- or D,L-lactide, conversion of the resulting polylactides into polymer chain-transfer agents, and then using them in RAFT polymerization with pentafluorophenyl methacrylate to form a second polymeric block. Transformation of the poly(pentafluorophenyl methacrylate) block by aminolysis with 1-amino-2-propanol then yielded P(HPMA)-b-P(LLA) or P(HPMA)-b-P(DLLA) block copolymers, which formed polymer micelles that were efficiently taken up by adenocarcinoma cells (HeLa).484,485 The synthetic strategy used to prepare these copolymers is based on a general reactive precursor approach and can thus be adapted to prepare diverse functional systems from a single polymer precursor, yielding block polymers suitable for use as micellar systems or soluble drug carriers. This strategy was also used in the synthesis of statistical and block copolymers of P(HPMA)b-poly(lauryl methacrylate) that self-assemble into HPMA copolymer nanoaggregates of various topologies with particle diameters that range between 100 and 200 nm depending on the molecular weight of the block copolymer.486 The use of polymer micelles to modulate immune responses represents a whole new field of research. Micellar nanoparticles of 25−30 nm were formed by self-assembly of a diblock copolymer consisting of a hydrophilic P(HPMA) coronaforming block and an amphiphilic core-forming terpolymer block comprising poly(propylacrylic acid), poly(dimethylaminoethyl methacrylate), and poly(butyl methacrylate). The incorporation of pyridyl disulfide groups into the P(HPMA) blocks enabled the reversible conjugation of thiolated ovalbumin.487 In vitro studies on these ovalbumin− polymer conjugates revealed that the micelles enhanced intracellular antigen retention and cytosolic antigen accumulation. Moreover, subcutaneous immunization of mice with the conjugates significantly enhanced their antigen-specific CD8+ T cell responses. These results demonstrate that micelles of this sort have great potential for use in the design and preparation of synthetic vaccines. A representative but relatively simple biodegradable micellar system based on an amphiphilic diblock copolymer was prepared by RAFT copolymerization of HPMA with the functional monomer 2-(2-pyridyldisulfide)ethyl methacrylate (PDSM).488 The versatility of the pyridyl disulfide groups on the P(HPMA)-b-P(PDSM) block copolymer was then exploited to covalently conjugate the copolymer with the anticancer drug DOX via a pH-sensitive hydrazone bond. The same groups were also used to stabilize the micellar assemblies by forming cross-linking internal reductively cleavable disulfide bridges. The biodegradable disulfide cross-links of the resulting micelles were reduced by cytosolic glutathione following

cellular uptake, causing the micelles to disintegrate into unimers. Poly(2-oxazoline)s (P(Oz)) are a relatively new class of polymers that have been used in the design of biomaterials and polymer therapeutics. Their tunable solubility in water can be used for preparation of nanostructures of diverse sizes and architectures bearing a wide range of chemical functionalities.489,490 Amphiphilic diblock copolymers have been synthesized based on poly(2-ethyl-2-oxazoline) P(EtOz) as a hydrophilic block and aliphatic polyesters such as poly(Llactide) P(LLA) or P(CL) as the hydrophobic blocks. Their micellar characteristics in an aqueous solution were investigated using dynamic light scattering and fluorescence techniques. The block copolymers formed micelles in the aqueous phase with CMCs in the range of 1.0−8.1 mg L−1. The mean diameters of the micelles were in the range of 108−192 nm with a narrow size distribution.491 Amphiphilic diblock copolymers with P(EtOz) derivatives as the hydrophilic block and poly(1,3trimethylene carbonate) P(TMC) as the hydrophobic block also formed micelles with CMCs in the range of 2.8−25 mg L−1.491 A third class of P(Oz)-based amphiphilic block copolymers was prepared using a one-step protocol in which 4-cyano-4-(dodecylthiocarbonothioylthio)pentyl-4-methylbenzenesulfonate (CDPS) serves as a dual initiator for RAFT polymerization and cationic ring-opening polymerization (CROP).492 Methyl (meth)acrylate, butyl (meth)acrylate, tertbutyl (meth)acrylate, and N-isopropylacrylamide were polymerized to form the hydrophobic block, while monomers of 2methyl-2-oxazoline and 2-ethyl-2-oxazoline were used to form the hydrophilic block. Drug carriers formed from polymer micelles based on P(Oz) have been studied extensively, but further investigations will be required to fully elucidate their usefulness in the development of new DDS that are effective in vivo. Sequential RAFT polymerization was used to prepare amphiphilic block copolymers in which a hydrophilic poly(2hydroxyethyl acrylate) block was combined with a hydrophobic poly(2-hydroxyethyl acrylate-co-2,2-dimethyl-1,3-dioxolane-4yl)methyl acrylate) (P(HEA)-b-P(HEA-co-DMDMA)) copolymer block. A DMDMA content above 11mol % was required for self-assembly in aqueous solution, with micelle sizes ranging from 23 to 338 nm depending on the DMDMA content. Under acidic conditions the nanoparticles decomposed into soluble unimers at a rate that was inversely proportional to the relative mass of the hydrophobic block copolymer. The in vitro cytotoxic activity of P(HEA)-b-P(HEA-co-DMDMA) micelles loaded with paclitaxel (PTX) was comparable to or better than that of the commercial PTX nanoformulations Abraxane and Genexol-PM at equivalent PTX doses.493 In addition to diblock copolymers, multiblock copolymers have also been synthesized and used to form micelles. The results obtained with systems of this sort have been reviewed on multiple occasions.494−496 Triblock copolymers containing blocks of poly(ethylene oxide), poly(glutamic acid), and poly(phenylalanine) (PEG−P(Glu)−P(Phe)) were successfully synthesized via NCA ring-opening copolymerization. The selfassembly behavior of PEG−P(Glu)90−P(Phe)25 was exploited to prepare biodegradable and biocompatible polypeptide-based polymeric micelles with a P(Phe) hydrophobic core, a crosslinked ionic P(Glu) intermediate shell layer, and a PEG corona. These cross-linked micelles of around 90 nm in diameter had narrow size distributions and underwent degradation after incubation with the proteolytic enzyme cathepsin B, which is 5370


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Table 6. Multidrug Polymeric PEG−Poly(AA) Micelles for Concurrent Drug Delivery polymer

drugs

indication

PEG-b-P(Asp-Hyd)a PEG-b-P(Asp-Hyd)a PEG-b-P(γ-benzyl-L-glutamate) + PEG-b-P(LLA) PEG-b-P(Glu)-b-P(Phe)

DOX/wortmannin DOX/17-hydroxethylamino-17-demethoxygeldanamycin DOX/etoposide, DOX/PTX PTX/cisplatin

MCF-7 breast cancer MCF-7 breast cancer CT-26 murine colorectal cancer A2780 ovarian cancer

a

status in in in in

vitro vitro vivo vivo

ref 498 499 500 497

Chemical loading (conjugation) of drugs.

Figure 34. Structure of triblock poly(2-methyl-2-oxazoline)-b-poly(tetrahydrofuran)-b-poly(2-methyl-2-oxazoline) copolymer for curcumin delivery.

commonly present in lysosomes. To test their potential as DDS, they were loaded with a combination of drugs having very different physical properties such as cisplatin (15 w/w %) and PTX (9 w/w %).497 The doubly loaded micelles exhibited synergistic cytotoxicity against human ovarian A2780 cancer cells in a xenograft model, with superior antitumor activity to that observed for micelles loaded with either drug alone or free cisplatin. Positive results were also obtained in cytotoxicity tests where DOX was incorporated into PEG−P(AA) micelles in combination with a geldanamycin derivative, wortmannin, PTX, etoposide, or PTX with cisplatin. The drugs were incorporated into the micelles by exploiting hydrophobic interactions or covalently via hydrazone bonds (Table 6). Drugs bound via a hydrazone bond were released from the micelles in a pHdependent manner, and the micelles’ cytotoxicity was comparable to that of the corresponding free drugs. Unfortunately, no in vivo results for these interesting systems have been reported. The ABA triblock terpolymer poly(2-methyl-2-oxazoline)-bpoly(tetrahydrofuran)-b-poly(2-methyl-2-oxazoline) (Figure 34), or TBCP, was synthesized by means of a cationic ringopening polymerization reaction in which tetrahydrofuran (THF) and then 2-methyl-2-oxazoline (MeOx) were added sequentially to a cationic initiator. The length of the two block types was adjusted by varying the polymerization time for each monomer.501 Self-assembly of the resulting copolymers in water generated a core−corona micellar structure. When these polymer micelles were used to entrap curcumin, the resulting conjugates’ permeation properties effectively promoted the drug’s cellular internalization in A549, HeLa, and 16HBE14o cells. Random copolymers that self-assemble into micellar DDS capable of binding covalently to drugs via functional groups in the hydrophilic shell can be formed by preparing P(HPMA) derivatives from mixtures of comonomers bearing hydrophobic dodecyl, oleic acid, or cholesterol moieties. Copolymers of this sort were conjugated to DOX via a pH-sensitive hydrazone bond-containing spacer. In aqueous media, copolymers of this class with a hydrophobic comonomer content of 1.5−5 mol % formed HMW supramolecular structures with diameters of 13− 37 nm, depending on the type and relative abundance of the hydrophobic substituent.264,502 P(HPMA)s bearing DOX and hydrophobic cholesterol derivatives (i.e., cholesterol acylated with an oxoacid, cholest-4-en-3-one, or 5-α-cholestanone), both bound by pH-sensitive hydrazone bonds (Figure 35), were also synthesized. Hydrolysis of these hydrazone bonds in the mildly acidic environment of tumor cells resulted in drug release and

Figure 35. Chemical structure of a cholesterol-containing HPMA copolymer conjugate used as a micellar delivery system for DOX. Both DOX and cholesterol are attached to the hydrophilic HPMA copolymer via pH-sensitive hydrazone bonds that control the drug’s release and the micelle’s disassembly.

dissociation of the supramolecular structure into unimers, which should facilitate the renal removal of the P(HPMA) carrier from the body after drug release.265 The drug release profiles of these systems were broadly similar to those for the related soluble P(HPMA) conjugates. EL4 T-cell lymphomabearing mice treated with these conjugates achieved good longterm survival and were fully cured at lower doses than were required to achieve comparable results with HPMA-based soluble systems. A list of micellar anticancer drugs that have received authorization for evaluation in clinical trials is presented in Table 7. Almost all of these systems are relatively simple and prepared by self-assembly of amphiphilic block copolymers containing hydrophilic PEG blocks and hydrophobic polyester (P(LA)), poly(aspartic acid), or poly(glutamic acid) blocks. The drug molecules used in these systems are mostly classical cancerostatics (paclitaxel, camptothecin and its analogue DOX, or platinum complexes). The systems that have reached the most advanced testing stages are conjugates of PTX (Genexol, 5371


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Table 7. Polymer-Based Micellar DDS That Are Currently in Clinical Trials product

polymer

drug

indication

Genexol-PM

PEG−poly(D,L-lactide)

Paclitaxel

breast cancer

Genexol-PM Genexol-PM Genexol-PM

PEG−poly(D,L-lactide) PEG−poly(D,L-lactide) PEG−poly(D,L-lactide)

Paclitaxel Paclitaxel Paclitaxel

nonsmall cell lung cancer recurrent or metastatic breast cancer pancreatic cancer bladder cancer, ureter cancer

Lipotecan

polymeric micelle

TLC388 (CPT derivate)

advanced solid tumor advanced/metastatic RCC patients

Nanoxel

mPEG−poly(D,L-lactic acid) PEG-b-poly(L-glutamic acid) mPEG-b-poly(L-glutamic acid) PEG-b-poly(α,βaspartatehydrazone) PEG-b-poly(L-glutamic acid) PEG-b-poly(L-glutamic acid) mPEG-b-poly(α,βaspartic acid) PEG-b-poly(α,β-aspartic acid) PEG-b-poly(L-glutamic acid) polymeric micelle

Paclitaxel

NC-4016 NC-6004 (Nanoplatin) NC-6300 NK-012 NK-012 NK-105 NK-911 NC-4016 Paclical a

phase marketed S. Korea II III II

ref

I

NCT01770795a NCT00876486e NCT00111904,c NCT01426126,c514 NCT00747474c

advanced breast cancer

II I

NCT01831973e NCT00915369e

Oxaliplatin

advanced cancers lymphoma

I

NCT01999491d

Cisplatin

pancreatic cancer

III

Epirubicin

solid tumors

I

NCT00910741,c NCT02043288d 503

SN-38

solid tumors, small cell lung cancer, breast cancer

II

SN-38

I

Paclitaxel

refractory solid tumors advanced solid tumors, metastatic colorectal cancer gastric cancer, breast cancer

III

NCT00951613,c NCT00951054c NCT00542958,c NCT01238939c NCT01644890b

Doxorubicin

various solid tumors

I

515

Oxaliplatin

advanced cancers lymphoma

I

NCT01999491d

Paclitaxel

epithelial ovarian cancer, primary peritoneal cancer, fallopian tube cancer

III

NCT00989131c

Clinical Trials.gov Identifier (https://clinicaltrials.gov/). bActive, not recruiting. cCompleted. dRecruiting. eUnknown.

NK-105, paclical) and cisplatin (NC-6004) with Genexol-PM having received approval for use in Europe and Korea as a treatment for breast and small cell lung cancer. NC-6004 (PEGP(Glu)) micelles loaded with carboplatin are currently undergoing Phase III clinical trials. Previous studies showed that they did not induce significant nephrotoxicity or any observable adverse reaction even at a high dose of 120 mg m−2. In Phase II trials, PEG-b-poly(aspartate) micelles loaded with PTX (NK-105) exhibited high tolerability and antitumor efficacy in patients with advanced stomach cancer, justifying the commencement of Phase III clinical trials. In addition, NC6300, a PEG-poly(aspartate hydrazide) (PEG-b-P(Asp-Hyd)) with pendant epirubicin, has entered Phase I clinical trials in Japan.503,504 This unique system in which epirubicin is covalently bound to a carrier via a pH-sensitive hydrazone linkage is stable in blood but releases the bound drug in the mildly acidic environments presented by tumors, endosomes, or the lysosomes of cancer cells.

the hydrophilicity and hydrophobicity of their constituent blocks, which can be modulated by the temperature. The second method relies on the use of multiblock copolymers in which one or two blocks consist of temperature-responsive polymers that form a hydrophilic shell or hydrophobic inner core above or below the lower critical solution temperature (LCST), respectively.510−512 The LCST is a unique characteristic of temperature-responsive polymers that depends on the polymer’s concentration. Below the LCST, the polymer is soluble due to extensive hydrogen bonding interactions with the surrounding water molecules and restricted intra- and intermolecular hydrogen bonding between polymer chains. Upon heating, hydrogen bonding with water is disrupted and intra- and inter-molecular hydrogen bonding and hydrophobic interactions dominate, causing a solubility transition.513 A list of temperature-responsive polymers with a wide range of transition temperatures has been compiled by Roy et al.516 Many thermoresponsive polymers with potential applications in biomedicine have transition temperatures between room and body temperature. For example, poly(N-isopropylacrylamide) (P(NIPAAm)) undergoes a sharp phase transition in water at around 32 °C, poly(N,N′-diethylacrylamide) has an LCST in the range of 25−35 °C, and poly(dimethylaminoethyl methacrylate) has an LCST in range of 14−50 °C.507,517 The LCST can be adjusted by copolymerization with hydrophobic or hydrophilic comonomers or end group transformations. By increasing the hydrophilicity of the copolymers, the overall hydrogen bonding ability of the macromolecules is increased, which leads to higher transition temperatures. Conversely, incorporating hydrophobic comonomers lowers the LCST.511,518−520 Two types of micellar structures containing thermosensitive poly(N-isopropylacrylamide) in the outer corona or inner core

4.4. Thermoresponsive Micellar Systems

Drug carriers based on stimuli-responsive polymers have great potential in biomedicine.505−509 Stimuli-responsive polymers are defined as polymers that undergo relatively large and abrupt physical or chemical changes in response to small external stimuli from the external environment. Temperature- and pHresponsive mechanisms have been investigated extensively because temperature and pH differences are associated with a number of disease states and related conditions, making them convenient natural stimuli for many biomedical applications. Here we focus on temperature-responsive polymeric micelles. Micelles of this kind can be prepared by two methods. The first method is based on the fact that the micellization of amphiphilic block copolymers is governed by the balance of 5372


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Figure 36. Thermoresponsive polymer micelles formed by self-assembly of a doubly thermosensitive diblock copolymer on raising the temperature above LCST1. On raising the temperature above LCST2, the formed micelles collapse, releasing the drug. This image was adapted from ref 528. Copyright 2015 American Chemical Society.

such diblock copolymers have been prepared with alternative hydrophobic core blocks consisting of materials such as copolymers of lactic and glycolic acid, poly(N-acryloyl-2pyrrolidone) and its derivatives,525 or poly(ε-caprolactone)526 among others. Individual thermally-responsive polymer blocks can also be prepared by statistical copolymerization of thermosensitive and nonthermosensitive monomers. Thus, NIPAAm was copolymerized with DMAAm to form a statistically copolymerized thermosensitive block, which was then paired with a nonthermosensitive P(DLLA) block to form the thermally responsive diblock copolymer P(NIPAAm)-co-DMAAm-b-P(DLLA). DOX-loaded micelles of this material were then prepared by dissolving it in dimethylacetamide with DOX and dialyzing the resulting solution against water. This yielded polymeric micelles with an LCST of 40 °C in phosphatebuffered saline, a diameter of 69.2 nm, and a monodisperse size distribution. The micelles released more DOX above the LCST (at 42.5 °C) than below it (at 37 °C) and showed little cytotoxic activity against bovine aorta endothelial cells at 37 °C, whereas their cytotoxicity at 42.5 °C was high. On the other hand, free DOX exhibited high cytotoxicity at both temperatures. These results suggest that such polymeric micelles could be effective for treating solid tumors because of the local hyperthermia of tumor sites: the micelles would be stable at 37 °C in the circulation but would release the trapped drug efficiently in the tumor, where the local temperature is appreciably higher.527 Another interesting thermoresponsive block copolymerbased micellar DDS for DOX delivery has been designed in which each block has a distinct LCST. The copolymer was prepared by RAFT polymerization and consisted of an A block formed by a copolymer of N-vinylcaprolactam with 3-methylN-vinylcaprolactam (LCST1 19−27 °C) and a B block formed by a copolymer of N-vinylcaprolactam with N-vinylpyrrolidone (LCST2 41−42 °C). Below LCST1, both blocks are hydrophilic and well soluble in water. If DOX is added to such a cold solution and the temperature is raised above LCST1 but kept below LCST2, the A block becomes hydrophobic. This causes spontaneous micelle formation with the entrapment of DOX in the micelle core, while the B block remains hydrophilic and forms a temperature-sensitive corona around the core. The resulting micelles are efficiently passively targeted to tumor sites. Raising the temperature above LCST2 makes the B block hydrophobic, collapsing the micellar corona and triggering aggregation and drug release (Figure 36).528

can be mentioned here. In micelles with P(NIPAAm) outer corona, NIPAAm block has been copolymerized with hydrophobic blocks, and in micelles containing P(NIPAAm) in the inner core, NIPAAm block has been copolymerized with hydrophilic blocks.521,522 Thus, P(NIPAAm)-b-P(BMA) micelles with a thermosensitive shell and hydrophobic core were prepared from a block copolymer of P(NIPAAm) and poly(butyl methacrylate) (P(BMA)). These micelles were loaded with DOX, and their cytotoxicity based on controlled temperature-dependent drug release was tested in vitro.523 Selfaggregation of the P(BMA) blocks resulted in the formation of a hydrophobic inner core that trapped DOX, while the P(NIPAAm) chains of the outer shell stabilized the micelles in a temperature-dependent manner. While the temperature was below the LCST the outer shell remained hydrophilic, so the micelles were stable and did not engage in unwanted interactions with cells, tissues, or other micelles. However, on raising the local temperature above the LCST (32.5 °C) the outer shell suddenly became hydrophobic, releasing the drug molecule. This dramatic thermoresponsive on/off switching caused a sharp increase in cytotoxicity in vitro driven by the temperature-dependent change in the structure of the micellar shell. The release of DOX was selectively accelerated upon heating through the LCST but was suppressed effectively at lower temperatures. Consequently, the micelles exhibited greater cytotoxic activity in vitro than free DOX above the LCST (at 37 °C) but lower cytotoxic activity than free DOX below the LCST (at 29 °C).523 The authors expressed a conviction that further development of such thermoresponsive micelles open important new opportunities to construct novel drug delivery systems that take advantage of localized hyperthermia. Another diblock copolymer with a thermoresponsive P(NIPAAm) corona and a hydrophobic polymer block in the core, P(NIPAAm)-b-P(DLLA), was synthesized by initiating the ring-opening polymerization of D,L-lactide with hydroxyterminated P(NIPAAm). Polymeric micelles were prepared from the block copolymer using a dialysis method. The resulting solutions exhibited reversible changes in their optical properties: they were transparent below the LCST and opaque above it. Dynamic light scattering measurements confirmed the formation of micellar structures with diameters of approximately 40 nm whose dimensions did not change detectably between 20 and 30 °C. Above the LCST, the polymer micelles aggregated reversibly; on cooling to below the LCST, the aggregates broke up and the micelles were restored.524 Many 5373


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Table 8. Polymer-Based Nanoparticles Featuring Natural or Synthetic Polymers That Are Used in Clinical Practice or Have Been Evaluated in Clinical Trials product

polymer

drug

indication

phase

metastatic breast cancer, non-small-cell lung cancer, pancreatic cancer metastatic breast cancer, prostate cancer solid tumors, bladder cancer solid tumors, lymphoma non-small-cell lung cancer, prostate cancer

FDA approved II I/II I/II I/II

solid tumors

I

551−553

I/II

Mitoxantrone

solid tumors, renal cell carcinoma, rectal cancer, nonsmall-cell lung cancer hepatocellular carcinoma

II

148, 554, NCT01612546,c NCT01380769c 555

polymeric nanoparticle

Docetaxel

solid tumors

I

556, NCT01103791e

polymeric nanoparticle polymeric nanoparticle

PTX DOX

peritoneal neoplasms hepatocellular carcinoma

I III

557, 558, NCT00666991c 559

Abraxane

albumin nanoparticle

PTX

ABI-008 ABI-009 ABI-011 BIND-014

albumin nanoparticle albumin nanoparticle albumin nanoparticle PEG−PLGA polymeric nanoparticle cyclodextrin polymeric nanoparticle

Docetaxel Rapamycin Thiocolchicine dimer Docetaxel

CALAA-01 CRLX-101 DHADPBCA PNP DocetaxelPNP Nanotax Transdrug BA-003 a

cyclodextrin nanoparticle polymeric nanoparticle

siRNA targeting ribonucleotide reductase subunit Camptothecin

ref 541−543 544, 545, NCT00531271a,d 546, NCT00635284c NCT01163071d 547−550, NCT01792479,b NCT01812746b

Clinical Trials.gov Identifier (https://clinicaltrials.gov/). bActive, not recruiting. cCompleted. dTerminated. eUnknown.

reviews.532,533 There are many systems under development, some of which have quite complex structures, but there have been only a few publications reporting convincing in vivo anticancer activity. It remains to be seen whether these systems will offer major advantages over classic micelles and other DDS intended for cancer therapy. Hundreds of papers on micellar DDS based on amphiphilic block copolymers suitable for cancer treatment (with structures of varying complexity) are published every year, but only a few have yet undergone clinical evaluation (Table 7); a very limited number are in the later stages of clinical testing, and one (Genexol-PM) has received regulatory approval for routine use in treating breast and lung cancer.12 Nevertheless, the clinical results that have been obtained are promising, suggesting that micellar therapies can improve patients’ prognosis while causing fewer and less severe side effects than conventional cancer chemotherapeutics.534 It is thus readily possible to envision the routine use of micellar DDS in the clinic, either individually or as components of combination therapies. The potential of micellar systems as vehicles for the simultaneous delivery of multiple drugs is particularly exciting because it could enable the exploitation of synergies between drugs, the combination of micellar systems with other DDS or different therapy types (e.g., immunotherapy or radiotherapy). Another attractive aspect of polymer micelles derives from their ability to overcome biological barriers, enabling the development of new therapeutic strategies as well as safe and efficient diagnostic tools.

There have also been attempts to actively target thermoresponsive micellar carriers to cancer cell surface receptors. Thus, thermoresponsive Pluronic F127−P(DLLA) copolymer micelles with an LCST of 39.2 °C were prepared and decorated with folic acid as a targeting moiety. At 37 °C, small quantities of the encapsulated anticancer drug DOX were released from the FP100 micelles; at a slightly higher temperature (40 °C), the shrinkage of the thermoresponsive segments caused rapid DOX release and instantly increased the drug’s local concentration. Cytotoxicity tests on the NIH 3T3 and human HeLa cell lines showed that the system had excellent cytocompatibility and that the FA-decorated micelles were taken up much more efficiently and had greater cytotoxicity toward folate receptor-expressing HeLa cells than was achieved with undecorated micelles. Importantly, the cytotoxicity of the DOX-loaded micelles toward the HeLa cells under hyperthermic conditions (40 °C) was much greater than at 37 °C. Therefore, these thermoresponsive micelles have great potential as drug vehicles for cancer therapy.529 Another type of folatedecorated thermoresponsive micelle based on the star-shaped amphiphilic block copolymer 4s[P(CL)-b-2s((P(NIPAAm)-coAAm)-b′-mPEG/PEG-FA)] was developed for the tumortargeted delivery and temperature-induced controlled release of hydrophobic anticancer drugs. These amphiphilic star copolymers self-assembled into spherical micelles with an average diameter of 91 nm and an LCST of around 39.7−40 °C. PTX was encapsulated into the micelles, and their cytotoxicity was evaluated in vitro. The drug-loaded micelles were relatively stable under physiological conditions but released the encapsulated drug when the temperature rose above the LCST. The cytotoxicity of the PTX−micelles was greater than that of the commercial PTX formulation Tarvexol, which was used as a control, and both the cytotoxicity and the cellular uptake of the PTX-loaded micelles increased sharply above the LCST. These results indicate that using folic acid as a micelletargeting moiety can improve the delivery efficiency and cancer specificity of hydrophobic chemotherapeutic drugs.530,531 A more extensive discussion on the potential of thermoresponsive micellar DDS in cancer therapy can be found in recent

5. POLYMER NANOPARTICLES 5.1. Polymer Nanoparticles for Drug Delivery

Nanoparticles (NPs) are a specific class of nanomedicines with particle sizes of 10−1000 nm and many potential medical applications. In anticancer therapy they can be used as depots within tissues, releasing a drug in the tumor’s vicinity of tumors at a rate controlled by the drug’s rate of diffusion from the particles or the rate of degradation of a polymer matrix. Alternatively, they can be used as injectable DDS to deliver 5374


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drugs to the liver, spleen, and bone marrow.535 Finally, if their surfaces are modified they can be used for drug delivery to solid tumors by passive accumulation (via the EPR effect) or by active receptor-mediated targeting to tumor cells. They can also be used as stimuli-responsive drug-releasing systems and as tools for radiotherapy or diagnostics. A fundamental requirement of any DDS is that it must be safely eliminated from the body once it has fulfilled its purpose. Therefore, many DDS are based on biodegradable natural or synthetic macromolecules. The first biodegradable nanoparticles to be investigated extensively were prepared from natural polymers; subsequently, synthetic poly(alkylcyanoacrylate) or poly(orthoester) systems were examined. More recently developed systems tend to be carefully designed and highly sophisticated, featuring a biodegradable synthetic polymer core with an entrapped drug or drug combination, and also often metal nanoparticles (of Au, Fe, or Ag), surrounded by a stealthing corona of hydrophilic polymers decorated with a targeting moiety of some kind. A detailed description of all the nanoparticle systems that have been reported to date (which include nanoparticles, nanocapsules, lipid particles and nanoemulsions, nanobubbles, and more) would be beyond the scope of this review; indeed, a separate review would be required to do the topic justice. Therefore, we will only mention some basic design principles of such systems and discuss those examples that are used in the clinic or are undergoing clinical trials (Table 8). For a more detailed overview we refer the reader to some recent reviews on nanomedicine and nanoparticle DDS.10,11,536,537 As shown in Table 8, only PNPs with rather simple particle structures based on biodegradable polymers (albumin, PEG/PLGA copolymers, or cyclodextrin derivatives) bearing hydrophobic drugs such as taxols, DOX, rapamycin, or CPT have been tested in clinical trials;337 the more sophisticated systems are currently in preclinical evaluation or in need of further development and more extensive in vivo studies. To improve the likelihood of successful preclinical and clinical evaluation of any nanomedicine, including PNPs, it is important to consider scientific problems such as biocompatibility, drug diffusion, mechanisms of polymer degradation and their control, and how best to prepare nanomaterials with well-defined structures and appropriate physico-chemical characteristics. However, it is also necessary to develop thoroughly validated methodologies and procedures for the synthesis of polymers and the preparation of nanoparticles, to consider regulatory barriers and acquire knowledge of relevant industrial processes, and to select and develop suitable biosystems against which to test the developed nanomedicines.125 Success thus requires truly interdisciplinary collaborations based on shared professional expertise from all disciplines involved in developing the nanoparticle system. The f irst-generation PNPs535 (Figure 37) were based on hydrophobic biodegradable polymers, and regardless of their composition and morphology, they were rapidly cleared from the bloodstream by the normal reticuloendothelial defense mechanism. As a result, they preferentially accumulated in the liver, spleen, and bone marrow. Nanoparticles of this sort, such as DOX-loaded poly(alkylcyanoacrylate) particles, are thus best suited for the delivery of drugs and treatment of liver metastasis. Studies on DOX-loaded poly(alkylcyanoacrylate) PNPs showed that they accumulated in Kupffer cells and released DOX slowly into tumor cells at a rate controlled by the particles’ biodegradation.535 Cardiotoxicity, which is a common problem in patients treated with DOX, was significantly

Figure 37. Three generations of PNPs. (A) First-generation PNPs based on hydrophobic biodegradable polymer; (B) second-generation PNPs shealthed with hydrophilic polymer coat formed by semitelechelic polymer; (C) third-generation PNPs shealthed with semitelechelic hydrophilic polymer, actively targeted; (D) third-generation PNPs shealthed with multivalent hydrophilic polymer, actively targeted.

reduced in patients treated with the PNPs. First-generation PNPs can be also used to target drugs for delivery to macrophages and other monocytes. The second-generation PNPs differ from their first-generation counterparts in that their surfaces are coated with a hydrophilic polymer (typically PEG). This stealthing allows the PNPs to avoid interaction with plasma proteins (opsonins) and significantly prolongs their plasma circulation. Hydrophilic stealthing can be achieved by coating the hydrophobic PNP’s surface with amphiphilic PEG derivatives (such as pluronics or bis-fatty acid derivatives) by adsorption, or the PNPs can be prepared by nanoprecipitation or solvent evaporation technology from amphiphilic block copolymers featuring blocks of PEG (or some other hydrophilic polymer) and hydrophobic biodegradable blocks of polymers such as the polyesters PLA or PLGA, poly(ε-caprolactone) (PCL), polyanhydrides, poly(αamino acids), or poly(alkylcyanoacrylates). Polymer-coated PNPs, depending on their size and density of coating, can be used for passive delivery of drugs to solid tumors by exploiting the EPR effect or as long-circulating drug reservoirs that slowly release drugs which would otherwise have short elimination half-lives. Long-circulating nanoparticles can be also used to deliver drugs through perturbed endothelial barriers, i.e., infected or inflamed tissues, into tumors, or through the blood−brain barrier to the brain. The third-generation PNPs are the most sophisticated nanoparticle delivery systems. In addition to a hydrophilic polymer coating they also contain targeting moieties that recognize specific cell receptors and thus mediate the specific delivery of PNPs to cancer tissues and cancer cells; they may also induce the uptake of the PNPs by the target cells. In principle, the targeting molecules used with soluble polymer carrier conjugates or polymer micelles (mAb and their fragments, hormones, folic acid, transferrin, RGD derivatives, and so on) could all be used for this purpose. The most specific targeting moieties are mAbs; their specificity has been used to target PNPs toward tumor cells expressing the human epidermal growth factor (HER2), epidermal growth factor 5375


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evaporated under reduced pressure. This is known as the oil-inwater (o/w) method. Emulsification by the water-in-oil (w/o/ w) method is an alternative option for preparing PNPs; this method, like all of those listed below, is suitable for preparing nanoparticles composed of both hydrophobic polymers and amphiphilic block copolymers. Salting-out methods involve dissolving the polymer in a watermiscible solvent (e.g., acetone), mixing the solution with water, and then adding molecules or salts with strong salting-out effects (e.g., sucrose, MgCl2, CaCl2, or Mg acetate) to induce precipitation of the polymer NPs. Alternatively, NPs can be obtained by a reverse salting-out effect induced by adding a large excess of water. Nanoprecipitation (solvent displacement) is another method for obtaining PNPs in which the polymer is dissolved in a water-miscible organic solvent (e.g., dioxane or acetone) and mixed with a nonsolvent or a mixture of nonsolvents. PNPs are then formed by the slow addition of water under moderate stirring. The resulting particles usually have a narrow size distribution and are well defined. Dialysis is a simple and effective method suitable for preparing small PNPs with a narrow size distribution. The polymer is dissolved in an organic solvent and dialyzed against a nonsolvent that is miscible with the solvent initially used to dissolve the polymer. It is essential to use a dialysis membrane with an appropriate cutoff in order to obtain PNPs with the desired characteristics. Methods based on supercritical fluid technology avoid the use of organic solvents, surfactants, and so on. The polymer is dissolved in the supercritical fluid, which is then allowed to expand rapidly across an orifice or through a capillary nozzle into ambient air or a liquid solution, leading to rapid particle formation. PNPs prepared in this way are very pure and completely devoid of organic solvents. Other methods of PNP preparation are based on techniques that are widely used in polymer synthesis such as emulsion polymerization of monomers emulsified in water or aqueous media by using efficient surfactants and emulsion stabilizers. It is essential to remove all traces of additives and residual polymer initiators at the end of the synthesis when using this approach. Various methods of emulsion polymerization (notably mini- and microemulsion polymerization) have been developed over the years; these techniques can deliver welldefined particles and make it possible to tailor the growing polymers’ characteristics during the polymerization process. However, some important problems such as the loading of drugs into the particles prepared by emulsion polymerization must be solved separately. With the exception of particles prepared from amphiphilic block copolymers, PNPs usually need some modification before being used in drug delivery. This typically involves coating their surfaces with hydrophilic polymers such as PEG or a HPMA (co)polymer. Alternatively, an amphiphilic PEG copolymer or a related derivative (e.g., a bis-oleyl, palmitoyl, or cholesteryl species) may be used, and the coating may be decorated with targeting moieties. It is difficult to generalize the methods and strategies used for surface modification of PNPs because the selection of methods is heavily dependent on the nature of the delivery system and its intended purpose. Consequently, this problem is best solved on a case-by-case basis.

(EGFR), or transferrin (TfR) receptors, or the prostate-specific membrane antigen (PSMA). For example, Trastuzumabmodified PNPs containing supermagnetic iron oxide were used to target tumors expressing HER2 in order to facilitate magnetic resonance imaging and magnetic relaxometry for diagnostic purposes and also for tumor treatment.538 Similarly, Cetuximab-targeted PNPs have been used to target EGFRoverexpressing cells and anti-CD20 Abs-PLA nanoparticles to target lymphoma cells. A notable example of PNP targeting with Ab fragments is the use of an anti-PSMA scFv to direct polymer-coated viral PNPs for gene delivery to prostate cancer cells.78 The targeting of PNPs was recently discussed at length in a very good review.538 The techniques used to conjugate functional molecules to the surface polymers of PNPs are also similar to those used with polymer-based DDS. In addition to anticancer drugs, PNPs can be used to deliver biologically active molecules such as DNA, plasmids, siRNA, or oligonucleotides. To achieve high loadings, the core polymer must be modified to accommodate the delivered molecules’ properties. For example, PEGdecorated block copolymers with hydrophobic blocks of PLA, PLGA, or poly(ε-caprolactone) have been supplemented with polycationic blocks of polyethylenimine, poly(2-aminoethyl ethylene phosphate), or poly(dimethylaminoethyl methacrylate) to enable the incorporation of plasmids or oligonucleotides in gene and siRNA delivery systems. In addition, alternative biodegradable polymers such as alginate, chitosan, gelatin, hyaluronic acid, and pullulan have been used to prepare PNP vectors for in vivo siRNA delivery. The design and preparation of such systems is discussed in much greater detail elsewhere.539 5.2. Methods for Preparing Polymer Nanoparticles

During the design and synthesis of polymer nanoparticles as carriers for drug delivery it is necessary to consider the composition of the polymer components, the size and size distribution of the nanoparticles, the shape, porosity, and architecture of the whole delivery system, and the properties of the drug(s) to be entrapped. The nanoparticles’ cores usually consist of biodegradable hydrophobic polymers or copolymers such as polyesters, polyalkylcyanoacrylates, polyanhydrides, or some natural polymer. Second-generation nanoparticles can be prepared from amphiphilic block copolymers or from firstgeneration PNPs by coating their hydrophobic surfaces with a hydrophilic polymer corona. The preparation of thirdgeneration PNPs requires additional techniques for conjugation with targeting moieties. Each synthesis clearly requires its own specific procedures and technologies. However, some basic principles are common to all systems. PNPs can be prepared from preformed polymers by means of solvent evaporation, salting out, dialysis, or supercritical fluid technology. Alternatively, they can be obtained by direct polymerization of appropriate monomers using microemulsion, miniemulsion, surfactant-free emulsion, or interfacial polymerization.540 Of course, methods used to prepare particles for biomedical applications should yield final products that are completely devoid of additives and other reactants. The principles of the most widely used methods for PNP synthesis are briefly discussed below. In solvent evaporation methods, the polymer and the drug to be encapsulated are dissolved in a volatile solvent that is immiscible with water (e.g., ethyl acetate). The solution is then emulsified by ultrasonication in water, and the solvent is 5376


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5.3. Nanoparticle DDS in Clinic and Clinical Studies

major breakthroughs in the treatment of malignant diseases; further work will clearly be required to reach the goal of more general and powerful treatments and cures.

It is not generally straightforward to obtain data from all of the clinical studies that have been conducted with a given type of medicine, especially in cases where the clinical evaluation was terminated. Nevertheless, this section reviews some key results obtained in various clinical evaluations of the PNP-based DDS listed in Table 8. Albumin nanoparticles loaded with PTX (Abraxane) have been approved by the FDA as an injectable suspension for the treatment of metastatic breast cancer, nonsmall lung cancer (NSCLC), and late-stage metastatic pancreatic cancer. In addition, some similar albumin PNPs loaded with docetaxel, rapamycin, and thiocolchicine have been or are being tested in Phase II trials. Abraxane in combination with docetaxel was used to treat 861 patients with pancreatic cancer; compared to docetaxel alone, the Abraxane plus docetaxel treatment delayed tumor growth by 1.8 months on average. Moreover, some of the side effects experienced by the Abraxane plus docetaxel group (including neutropenia, thrombocytopenia, alopecia, diarrhea, and vomiting) were less severe than those suffered by the docetaxel control group. In a separate trial, a combination of Abraxane and carboplatin was tested in 1052 patients with NSCLC. The Abraxane group exhibited a higher objective response than was observed for the PTX-treated control group, and their overall survival increased from 11.2 to 12.1 months. Abraxane with rapamycin (ABI-009) was tested in 27 patients with stage IV disease (which is characterized by advanced unresectable nonhematological malignancies) in a Phase I clinical evaluation. The MTD for Abraxane alone was determined to be 100 mg kg−1 weekly by iv administration. When administered alone, the drug was well tolerated with preliminary evidence of response and stable disease. Fair doseproportional pharmacokinetics were also reported. These results suggest that Abraxane represents an important new treatment option, especially in combination cancer therapies. PEG−PLGA nanoparticles (BIND-014) loaded with docetaxel have been tested as a second-line therapy for NSCLC and prostate cancer in stage III/IV of the disease. These PNPs were targeted to PSMA; in animal models they delivered up to 10 times more DTX to tumors than was achieved with DTX alone. BIND-014 was administered in three doses (over a 21 day cycle) to 40 patients, under which conditions it was well tolerated and displayed antitumor activity at low doses. Thirtythree evaluable positive responses were recorded; in 36% of patients the disease was stabilized, and neutropenia, anemia, and neuropathy were significantly reduced compared to treatment with the free drug. These data were considered to warrant further evaluation, which is currently in progress. PNPs consisting of a cyclodextrin-based polymer with a PEG coat and a transferrin targeting moiety (CALAA-001) have been tested as vehicles for the delivery of oncogenic molecules, namely, siRNA, against ribonucleotide reductase. Fifteen patients with solid refractory cancer were treated with four iv infusions. Some side effects (fatigue, fever/chills, constipation, nausea/vomiting) were observed, but recovery was recorded shortly after infusion. However, the study was terminated after Phase I clinical trials. The clinical trial results reported in the literature suggest that PNP-based drug delivery systems have considerable potential in cancer therapy and great scope for further improvement. Unfortunately, the results presented to date seem to represent evolutionary improvements over existing options rather than

6. MAGNETIC IRON OXIDE NANOPARTICLES One of the most extensively explored classes of nanosystems suitable for drug delivery are inorganic nanoparticles (INPs). Like polymeric nanoparticles, INPs range in size between 1 and 1000 nm. However, for drug delivery purposes they should be no bigger than 200 nm to avoid opsonization and subsequent elimination by the RES. Compared to polymer-based DDS, the advantages of using INPs for drug delivery are that they can both enhance the effect of the delivered treatment and also facilitate imaging and monitoring of the treatment’s efficacy. One class of INPs that is widely used in DDS are superparamagnetic iron oxide nanoparticles (SPIONs). They can be prepared in various sizes (which may be defined in terms of hydrodynamic size or core size), are highly biocompatible, and have a wider range of interesting and complex properties that are useful for drug delivery than alternative INPs such as carbon or silica nanoparticles. The major advantage of SPIONs as DDS derives from their magnetic behavior. This allows them to serve as contrast agents in magnetic resonance imaging (MRI), which is currently one of the most popular and widely available medical imaging techniques. It also allows them to be guided and held in a desired location by magnetic fields and to induce local heating in tumor regions by magnetic fluid hyperthermia. This can be used to trigger the release of a loaded drug or to cause cell death by temperature-induced apoptosis. These properties give SPIONs a wide range of potential applications as advanced theranostics agents (i.e., medicines that are useful for both therapy and diagnosis) and nanocarriers for drug delivery. For example, they could potentially be delivered to tumor tissues by region-specific magnetic targeting, where they would release the loaded/attached drug on demand while enabling the entire process to be monitored by MRI. However, the magnetic properties of SPIONs also present some drawbacks and challenges; notably, they increase the particles’ tendency to aggregate. Therefore, SPIONs are very commonly combined with biological or synthetic polymers to form nanostructures such as magnetic nanoclusters, SPIONs entrapped in organic stimuli-responsive matrices, or magnetic micelles among others. The polymers prevent aggregation and facilitate secondary functionalization with drugs, radionuclides, and/or compounds that can protect the carrier against recognition by the immune system (i.e., compounds that increase the carrier’s recognition time). Such polymeric coatings make SPIONs amenable to both covalent and noncovalent drug-loading strategies, giving them access to a wide range of drug release profiles and mechanisms including release induced by external stimuli or changes in physiological conditions in the vicinity of tumors. 6.1. Advances in the Use of SPIONs for Drug Delivery

Strategies that exploit the intrinsic magnetic properties of SPION-based drug carriers rely on their strong magnetic response to small applied magnetic fields. Moreover, once the size of SPIONs falls below a characteristic threshold they become superparamagnetic (25 nm for Fe3O4 and 30 nm for γFe2O3).560 It greatly facilitates their visualization by increasing their imaging contrast as well as their ability to be manipulated in space by magnetic fields and their capacity to induce local heating. Hence, the superparamagnetic particles do not have 5377


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properties and the features of the applied magnetic field (strength and gradient), it is important to consider hydrodynamic and physiological parameters such as the infusion route, blood half-life, the reversibility/strength of the drug/ carrier bond, and the tumor’s volume. As mentioned above, magnetic carriers can be used to facilitate medical imaging and to generate localized temperature increases. This means that they can serve as theranostic agents, i.e., multifunctional tools useful for both medical diagnostics and therapy. Superparamagnetic nanoparticles have a very large magnetic moment (superspin) amounting to several tens of thousands of Bohr magnetons (μB = 9.274 × 10−24 J/T). They show magnetic saturation behavior in the magnetic fields of an MRI scanner, locally perturbing the dipolar field and causing a substantial shortening of the T2 (spin−spin) relaxation time. Consequently, they can be used as negative contrast agents in T2-weighted MRI experiments. This provides a way of monitoring their movement through the body, determining whether they have reached the tumor’s location, monitoring the status of the cancer treatment, and manipulating their fate after they have fulfilled their purpose in delivery and treatment. In addition, magnetic carriers can be used to generate heat when placed in external alternating magnetic fields because the rate of change of nanoparticle magnetization is finite and lags behind the reorientation of the applied field. Consequently, some of the magnetic energy is converted into internal energy in the nanoparticle, increasing its temperature. If the nanoparticle is closely attached to a tissue, this heat will be effectively transferred, heating that tissue. Heating tumor tissues can increase drug efficiency by disrupting cancer cells’ resistance mechanisms.

remanent magnetization and can exhibit better colloidal stability. Superparamagnetism is a finite-size effect emerging when the size of a nanoparticle falls below a threshold value when it enters a single-domain state, i.e., a state when all the atomic magnetic moments inside the nanoparticle lie in one direction along the respective easy axis of magnetization established and maintained by the particle magnetic anisotropy. Thus, in a single-domain state, we define a magnetic moment of a nanoparticle, often termed as superspin, as a simple product of magnitudes of all the atomic magnetic moments inside the nanoparticle and pointing along the easy axis of particle magnetization. Thus, all atomic magnetic moments are aligned due to exchange magnetic interactions. Once in a superparamagnetic state, the superspin fluctuates between the directions favored by particle magnetic anisotropy if the thermal energy supplied to the system is sufficient to overcome the barriers imposed by the nanoparticle’s magnetic anisotropy; on the atomic level, all atomic magnetic moments fluctuate but in a cooperative way due to existing exchange magnetic interactions still holding the mutual alignment of atomic magnetic moments inside the nanoparticles. The duration of the superspin’s alignment along a given orientation is known as the relaxation time (τ = τ0 exp(KV/kBT), where τ0 is the preexponential factor (10−9−10−11), K is the magnetic anisotropy energy constant, V is the volume of a nanoparticle, kB is the Boltzmann constant, and T is the temperature). In a superparamagnetic state, the superspin can easily orient itself in the direction of an applied magnetic field, giving a strong magnetic response due to its high saturation magnetization, which significantly increases the T2 relaxation time of the protons of nearby water molecules in MRI. High saturation magnetization (and hence high magnetic susceptibility) is essential for the magnetic targeting of nanoparticles because the force, Fm, exerted by an applied magnetic field on a magnetic nanoparticle is equal to Fm = χV(∇B)2/(2μ0), where χ is the particle’s magnetic susceptibility, V is its volume, ∇B is the gradient of the external magnetic field, and μ0 is the permeability of free space. The external magnetic field will hold the nanoparticle in the tumor site if the magnetic force exceeds the hydrodynamic drag forces exerted by the blood flow. Once the drug carriers have been concentrated at the tumor site with the assistance of the external magnetic field, the drug is released by either enzymatic activity or by changes in physiological conditions (i.e., some combination of the pH, temperature, or osmolality). Early attempts to perform targeting with external magnetic fields were hindered by weak magnets or unsuitable field gradients, but progress in magnet construction techniques has made it possible to generate gradients that can be used to effectively drive magnetic carriers to places deep inside the human body such as the brain. Nanoparticles based on γ-Fe2O3 and Fe3O4 have proven to be particularly promising in these applications due to their magnetic properties; they are both ferrimagnetic, and their nanoparticles exhibit superparamagnetic behavior. Moreover, SPIONs are nontoxic, biodegradable, biocompatible, and efficiently cleared from the human body via the pathways of Fe metabolism.561,562 Alternative magnetic nanoparticles are based on nanometals (e.g., iron, cobalt, or nickel), nanoalloys, or iron-containing garnets. However, they are more toxic than iron oxides and must therefore be functionalized with other compounds to reduce their toxicity. To date it has been shown that the efficiency of magnetic targeting depends on several parameters. In addition to the carrier’s intrinsic magnetic

6.2. Design and Structure of Magnetically Engineered Drug Delivery Systems

Magnetically engineered drug delivery systems (MEDDS) encompass a variety of structural architectures and pathways of design (e.g., combination with other drug delivery systems as polymers or carbon nanostructures).563 It is well understood that the superparamagnetic iron oxide nanocrystals (SPIONs) must be appropriately integrated with (macro)molecules in a way that will produce new tailored structural properties and functions. These functions can range from binding mechanisms that keep the hybrid colloidal assemblies intact in complex biological environments to much more complex functions such as those required to create multicompartmental MEDDS564 in which drugs and SPIONs are enclosed in separate areas of a nanophase that can release its cargos selectively and conditionally, i.e., in response to selected biochemical or physical stimuli. The integration of the SPIONs and their organic counterparts at the nanoscale is almost always facilitated by noncovalent interactions (coordination or hydrogen bonds) between the (macro)molecule and the surface metal ions or hydroxyl terminal surface groups. Even in the case of hydrophobic SPIONs that are encapsulated in the hydrophobic compartments of polymeric micelles and emulsions their mixing with the polymers and subsequent encapsulation is effective only if they have already been modified and colloidally stabilized with small coordinating molecules such as oleic acid. The noncovalent nature of the resulting bonds enables facile exchange reactions with mono- or oligodentate ligands after incubation in solutions containing competitor molecules.565,566 Such exchange reactions are widely exploited to achieve the 5378


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phase transfer of hydrophobically modified nanocrystals.567 It is very important to consider the strength of these binding interactions and particularly collective strength when designing, because they give rise to collective phenomena and allow stability and other functions to be achieved by exploiting initially forces. Much contemporary research on MEDDS is thus focused on studying such interactions towards the manipulation and organization of the individual components into new nanoscale architectures with novel and desirable functions. On the other hand, bioactive substances are incorporated into MEDDS by means of both covalent and non-covalent interactions. Similarly, the strengths of these bonds and their sensitivity to the local environment control key properties of the MEDDS such as its drug loading and loading efficiency,568 pharmacokinetics, and potential for smart on-demand release. Figure 38 shows some indicative hybrid superstructures that have been prepared by manipulating interactions between soft and condensed matter, including directly functionalized SPIONs, magnetoliposomes, magnetic polymersomes, condensed clusters, and magnetomicelles. There can be no doubt that hybrid drug delivery systems designed on the basis of our steadily improving understanding of both nanoscale and biological processes have great potential for improving health and treating a wide range of maladies.569 6.2.1. Synthetic Approaches. The synthesis of SPIONs has been extensively reviewed in previous articles along with the pathways and materials used for their modification.575−579 There are many established routes for their synthesis based on physical, wet chemical, and microbial methods. However, wet chemical routes have been far more widely used than the others when preparing SPIONs for MEDDS. While the preparation of the magnetic iron oxide phases is not generally regarded as a major challenge, the development of improved synthetic routes continues to be important because minor modifications can significantly alter the size and shape of the obtained SPIONs and also the supramolecular assembly of the individual SPIONs into spherical or asymmetric supraparticles in many cases.570,580 Generally, there are four main routes for preparation of SPIONs: coprecipitation, microemulsion, hydrothermal synthesis, and thermal decomposition.581 The first mentioned is the simplest one and with good potential to be upscalable to industrial production for future clinical applications. The other ones sometimes offer particles with good monodispersity but yields may be low (microemulsion method), or procedures might be energy-demanding. Still, the coprecipitation technique produces particles with high polydispersity, which is the main drawback for its clinical use. Irrespectively of the method used, SPIONs require their hybridization with polymeric materials during or after the formation of the magnetic core. However, the polymer should be carefully chosen to fulfill the properties desirable for its application because not only the size but also the organic shell determines the interactions of SPIONs with biological environment and the immune system. It can strongly affect their biostability, biocompatibility, cellular uptake,582 or blood half-life. The surface coating should fulfill the general characteristics for successful in vivo delivery as neutral surface charge. For example, SPIONs covered with citrate and have a hydrodynamic size of 12 nm in the study by Schnorr et al. exhibited only 9 min of blood half-life in rats.583 In the comparison, the particles covered with neutrally charged dextran even with higher diameter (20 nm) can have the blood half-life up to 2 h.584 Still, even the SPIONs with very similar

Figure 38. Interactions between inorganic magnetic nanoparticles and organic matter give rise to stable colloidal nanoassemblies with structural attributes and functionalities that have paved the way for future advanced clinical therapeutic schemes. Nanomedicines can facilitate spatiotemporally controlled drug delivery and the exploitation of magnetism-based techniques for enhancing the efficiency of drug delivery and treatment, giving them clear (and often outstanding) advantages over conventional treatments. The figure presents a conceptual illustration of some representative magnetically engineered drug delivery systems based on colloidal nanoassemblies stabilized via covalent and noncovalent interactions. (A) Dense cluster of SPIONs after direct surface functionalization with polymeric ligands. Drug molecules are bound to the polymeric coating covalently or electrostatically. Reprinted with permission from ref 570. Copyright 2007 Wiley. (B) Magnetoliposome in which SPIONs and hydrophilic drugs are encapsulated in the aqueous compartment, while hydrophobic bioactives are trapped in the lipid bilayer. Reprinted with permission from ref 571. Copyright 2012 American Chemical Society. (C) Magnetic polymersome structure with hydrophobic SPIONs embedded in the hydrophobic shell and hydrophilic drugs encapsulated in the hollow aqueous core. Adapted with permission from ref 572. Copyright 2013 Elsevier. (D) Magnetomicelle formed from amphiphilic block copolymers that form a hydrophobic core of soft condensed matter that encapsulates hydrophobic SPIONs and bioactive substances. From the authors’ own TEM archive. (E) Condensed clusters coated with agarose and covalently conjugated to doxorubicin with a pH-cleavable hydrazone bond. Reprinted with permission from ref 573. Copyright 2012 Wiley. (F) Polymersomes with hydrophilic SPIONs entrapped in the aqueous compartment. The hydrophobic middle layer is formed from polyglutamate, conjugated with the hydrophobic (protonated) form of doxorubicin. The outer and inner layers are composed from PEG, with the latter being crosslinked for enhanced stability. Reprinted with permission from ref 574. Copyright 2010 American Chemical Society. (Insets) TEM images of real-world variants of the systems depicted in schematic representation form in the main figure.

size and surface charge can have different polydispersity index, grafting density, or different surface chemistry (charge distribution, hydrophilicity/hydrophobicity) that can affect protein microcorona and related pharmacokinetic characteristics.585 The complete data available for SPIONs and their pharmacokinetic properties can be found in the recent review by Aramie et al.562 Hence, the surface coating can also affect the 5379


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Figure 39. (A) Elongated nanoworm SPION clusters form multidentate interactions with the target cell surface, increasing the efficiency of targeting and drug delivery to the cytoplasm. Reprinted with permission from ref 596. Copyright 2009 Wiley. (B) MEDDS based on covalent attachment of janus-like SPIONS with asymmetric surface chemistry and a drug-bearing liposome. The application of a suitable radiofrequency sensitizes the SPION chains, triggering on-demand drug release from the liposome. Reprinted with permission from ref 588. Copyright 2012 American Chemical Society.

Figure 40. (A) TEM image of ferrimagnetic vortex domain nanorings. Reprinted with permission from ref 600. Copyright 2015 Wiley. (B) SEM images of iron oxide nanodiscs. Reprinted with permission from ref 601. Copyright 2015 Wiley.

cellular receptors compared to conventional SPIONs (Figure 39A). Although there were no significant differences between the nanoworms and spherical SPIONs with respect to BCT, the nanoworms clearly remained in the vicinity of the tumor for a much longer time. In principle, this should enable larger quantities of the drug to be delivered to the tumor. The SPIONS developed in this work were prepared by a typical protocol based on coprecipitation in the presence of a high molecular weight dextran (20 kDa). This approach exploits noncovalent hydrogen bonding interactions developed between the dextran’s hydroxyl groups and those on the SPION’s surface to direct the SPIONs’ asymmetric assembly. In an alternative synthesis of SPION nanoworms,588 the supramolecular arrangement of the nanoparticles was controlled by covalent interactions,591 utilizing principles from solid-phase peptide synthesis. This approach is particularly interesting because of its potential to create novel architectures based on magnetoliposomes in which the SPIONs are covalently attached to the liposomes’ exterior (Figure 39B). Importantly, structures of this sort exhibited dramatically enhanced drug release compared to conventional magnetoliposomes when a suitable radiofrequency was applied. The exploration of different nanoparticle shapes and sizes by tweaking the parameters of conventional synthetic routes also frequently leads to materials with unprecedented properties. Such tweaking resulted in procedures for the synthesis of SPION-based nanocubes of 30 nm that can be assembled into higher order aggregates and excellent hyperthermia mediators,597 as well as 20 nm individual SPION nanocubes.598 In another example, studies on the use of doped ferrites to form core−shell structures that engage in useful exchange interactions revealed but dramatic enhancements in specific absorption rates.599 Nanorings600 (Figure 40A) and nanodiscs601 (Figure 40B) formed during the hydrothermal

rate of biodegradation probably thanks to the different water diffusion through the organic shell.586 Certainly, other forms of magnetic nanomaterials (iron carbide for instance587) are alternatives with exciting properties related to therapeutic applications. The composition will clearly affect the properties of the resulting MEDDS including their loading of magnetic material and overall magnetization, while other physicochemical parameters may affect the form adopted by the SPIONs, as for instance their clustering into dense spheres or worm-like architectures. These properties in turn influence the MEDDS’ interactions with cells and their in vivo biodistribution588 as well as their responses to magnetic drug targeting (MDT), magnetic fluid hyperthermia (MFH) for treatment and drug release induced by exposure to alternating electromagnetic fields (AMF), and imaging-guided drug delivery.589 Some important recent developments in this field include new protocols for the controlled synthesis of SPIONs with new architectures such as condensed colloidal nanocrystal clusters (co-CNCs) and nanoworms.590−592 The use of coCNCs in MEDDS is advantageous because they make it possible to accommodate a greater magnetic mass in individual nanoassemblies than is possible with conventional SPION architectures570,593,594 without sacrificing superparamagnetism as a result of increasing the particle size. These systems consequently exhibit unusually strong responses to magnetic manipulation, making them particularly suitable for use in magnetic drug targeting applications. In addition, their high magnetic moments produce higher transverse relaxivities and thus increase the sensitivity of techniques for their detection.595 On the other hand, nanoworm architectures have been reported to be advantageous for molecular targeting (MoT)590,596 and because of their blood circulation times (BCTs).588 Their advantages in MoT have been attributed to a greater number of interactions between targeting moieties and 5380


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Table 9. Some Attributes of Magnetically Engineered Drug Delivery Systems Whose Therapeutic Efficacy Has Been Studied in Vivoa ref

Ms (emu/ghybrid)/ SPIONs diameter (nm)

drug/capturing matrix

622 623

n.r./∼10 4.5/n.r.

MTX/lauric acid DOX/surface ions of SPIONs DOX/P(AA)

624

46/4−8

589

∼40/4−11

620

n.r./3−4

625

n.r./4.5

626

15/10

627 628 629 630

n.r./12 n.r./9.5 n.r./22 2/20

631 603 632

20/8−10 n.r./5 43/5−15

633

n.r.

634

20/7

635 636

30/5 40/4

637

25/10

638

8−18/n.r.

619

n.r./6

DOX/PAMAM and PEI gel

639

n.r./9.4

640

n.r./20

588

n.r./10

DCX/aqueousliposomic DOX/aqueousliposomic DOX/aqueousliposomic

641 642 643

15/6−12 20−25/10 22/n.r.

PTX/PLGA ETX/PPO DOX/hydrocarbon

572

Targeted Drug Delivery with Polymers and ... - ACS Publications

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