Core-shell nanocarriers for cancer therapy. Part I: biologically oriented design rules.

Review

Expert Opin. Drug Deliv. Downloaded from informahealthcare.com by Emory University on 06/21/14 For personal use only.

Core--shell nanocarriers for cancer therapy. Part I: biologically oriented design rules 1.

Introduction

2.

Intravenous delivery

3.

Lung delivery

4.

Oral delivery

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Other routes of administration

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Conclusions

7.

Expert opinion

Ivana d’Angelo, Claudia Conte, Agnese Miro, Fabiana Quaglia† & Francesca Ungaro †

University of Napoli Federico II, Department of Pharmacy, Laboratory of Drug Delivery, Napoli, Italy

Introduction: The application of nanotechnologies to the cancer field for therapeutic, imaging or diagnostic purposes represents an advanced and very attractive approach to overcome the main limits related to conventional chemotherapy. In particular, core--shell nanocarriers can be engineered to provide adequate features to overcome the main biological barriers encountered by free anti-cancer drugs. Areas covered: This review will try to summarise the design rules -- as dictated by biological requirements -- to take into account for proper nanocarrier design and to highlight the potential of administration routes other than intravenous. Expert opinion: Although intravenous injection remains the most investigated route of administration for ‘nanoncologicals’, research interest towards less explored administration routes allowing localised chemotherapy or delivery in close proximity to the tumour site might change the way cancer is treated in the near future. Nevertheless, an experimental set-up more biologically oriented taking into account an in-depth evaluation of stability in complex media, protein interaction, and cell interaction of novel nanoconstructs could allow their successful translation in pre-clinical and clinical settings. Keywords: administration route, biological barriers, cancer, core--shell nanocarriers, nanomedicine Expert Opin. Drug Deliv. (2014) 11(2):283-297

1.

Introduction

The term ‘nanoncology’ refers to the application of nanotechnologies to the cancer field for therapeutic, imaging as well as diagnostic purposes. The active agent is dissolved, entrapped, adsorbed, covalently attached or encapsulated in the nanocarrier giving an array of architectures and functionalities. In the case of conventional chemotherapeutics, this approach is considered beneficial in solving solubility issues and inconsistent bioavailability, based on the concept that pharmacokinetics (PK) of an anti-cancer drug can be usefully altered in the body to promote drug accumulation predominantly in pathological sites [1-5]. The therapeutic effects of some anticancer drugs could be significantly improved also if drug delivery occurs specifically to organelles inside cancer cells. Thus, an appropriate design allows nanoconstructs to improve drug efficacy (activity at lower doses) as compared with the free-drug treatment, giving in turn a wider therapeutic window and lower side effects. Furthermore, nanocarriers are also capable of addressing several drug delivery problems, which could not be effectively solved in the past. Overcoming multidrug resistance (MDR) phenomenon and penetrating cellular barriers that may limit device accessibility to intended targets, such as the blood--brain barrier, remain unsolved issues where nanodelivery technologies can play a key role. On the other hand, engineered 10.1517/17425247.2014.868881 © 2014 Informa UK, Ltd. ISSN 1742-5247, e-ISSN 1744-7593 All rights reserved: reproduction in whole or in part not permitted

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Nanocarriers represent an advanced and very attractive drug delivery approach for anti-cancer drugs to overcome the main limits related to conventional dosage forms. The efficacy of nanocarriers in cancer first depends on their ability to overcome biological barriers imposed by the administration route as dictated by shell features and overall colloidal properties. Core--shell nanocarriers can be especially engineered at surface to target a drug to the tumour site through passive mechanism (EPR) or specific recognition of protein/receptors over-expressed on tumour cells. Adequately engineered long-circulating core--shell polymeric nanocarriers with a sustained drug release, may be rediscovered beneficial in cancer treatment, also in the light of metronomic approaches. The knowledge of tumour microenvironment may be of great help in designing novel ‘smart’ stimuli-responsive nanocarriers able to activate on-demand drug release. Less explored administration routes allowing localised chemotherapy or delivery in close proximity to the tumour site, such as lung delivery, may pave the way to novel highly selective and potent nanomedicines for cancer. To proper design nanocarriers for cancer, an experimental set-up more biologically oriented on a characterisation standpoint is required, that is cell-level studies in relevant biological conditions and development of animal models resembling in vivo conditions.

This box summarises key points contained in the article.

nanocarriers have many potential benefits for diagnosing and treating metastatic cancer, including the ability to transport complex molecular cargo to the major sites of metastasis, such as lungs, liver and lymph nodes, as well as targeting to specific cell populations within these organs [6]. Another aspect covered by nanocarriers relies on the significant opportunity to apply nanotechnologies in molecular cancer imaging and diagnosis [6-8]. Tumour imaging plays a central role in clinical oncology by helping: i) to identify solid tumours, ii) to determine recurrence and iii) to monitor therapeutic responses. Imaging techniques use different contrast agents, which, once engineered or incorporated in nanoparticles (NPs), can allow efficient single or multimodal imaging in cancer detection. Indeed, the main advantage of increased image resolution is discovering lesions so small to be undetectable with conventional techniques and also monitoring disease progression, which would represent a significant improvement over the available clinical diagnostic methods. Furthermore, nanoprobes, such as quantum dots (QDs), can be used to identify a panel of biomarkers on intact cancer cells and tissue specimens, thus allowing a deeper understanding of cancer behaviour and early cancer detection [9]. Integration of an imaging element in a nanoconstruct makes it possible to follow biodistribution in the body in real time, which can 284

be useful to optimise carrier design in order to reach a specific area in the body and to elucidate toxicological concerns. Finally, imaging techniques and treatment strategies can be properly combined in the same entity giving rise to multifunctional theranostic nanomedicines conveying unique advantages for cancer treatment [10-13]. The behaviour of nanocarriers in a living body and their interaction with biological environment is driven by physiochemical features of the nanocarrier such as size, shape, deformability, surface properties, stability, dose and route of administration, all of them considered crucial for a beneficial therapeutic outcome. Thus, a huge amount of nanosystems with more or less complicated design have been developed over the past few decades and there has been intense effort towards understanding how nanocarrier properties can be properly manipulated for oncology applications. In particular, core--shell NPs represent a feasible way to attain multiple functionalities on a nanoscopic length scale. The core generally entraps the drug, releases it in a pre-programmed way while preventing its direct contact with biological environment. On the other hand, the shell provides an interface especially designed to tune particle interaction with biological environment. In this review, we describe the main barriers encountered by nanocarriers in their journey in the body after administration by well-established and emerging administration routes (summarised in Table 1). Then, we highlight the fundamental design rules dictated by each administration route (summarised in Figure 1) to build up core--shell nanocarriers on a rational basis. A detailed description of the different strategies to obtain core--shell nanocarriers for cancer, main building blocks employed, their assembling and specific features achieved can be found in the Part II [14]. 2.

Intravenous delivery

Intravenous injection remains the preferred route of administration for several anti-cancer drugs in order to reach different body compartments. Nevertheless, multiple obstacles must be overcome in order to drive an injectable nanocarrier to solid tumour lesions, once its stability in the blood has been demonstrated. The major challenges include: i) exploiting extravasation at tumour site through the enhanced permeability and retention (EPR) effect; ii) overcoming clearance in the body; iii) increasing cellular uptake into target tissue/cells/ organelles, iv) promoting endo-lysosomal escape, which is especially important for DNA and nucleic acid fragments (siRNA, oligonucleotides). In principle, nanocarrier delivery to tumour tissues can be achieved by either passive targeting, active targeting or stimuli-responsive materials. Passively targeted nanocarriers Passive targeting takes advantage of the unique properties of tumour vasculature and microenvironment [15]. In fact, 2.1

Expert Opin. Drug Deliv. (2014) 11(2)

Core--shell nanocarriers for cancer therapy. Part I: biologically-oriented design rules

Table 1. Main administration routes of core--shell NPs for cancer and corresponding key challenges. Administration route Intravenous


Pulmonary

Oral

Intratumoral

Intraperitoneal

Key challenges

Renal clearance Opsonisation (if tumour is not localised at MPS) Tumour type and status Tumour vascularisation Cellular uptake into target tissue/cells/ organelles Exhalation Deposition at bronchi, bronchioles and alveoli Macrophage escape Mucociliary escalator Airway mucus barrier Uptake into lung epithelial tissue/cells Specific targeting to cancer cells Nanoparticle accumulation and toxicity Stomach acidic condition Enzymatic activity First-pass extraction mechanism (liver cytochrome P450) P-gp efflux (MDR effect) GI mucus barrier Uptake into GI epithelial tissue/cells Distribution in tumour mass Composition and functions of the tumour extracellular matrix Tumour interstitial fluid pressure Occult tumour cells (i.e., cancer recurrence) Trafficking in the lymphatic system Systemic absorption Rapid clearance from the peritoneal cavity Spleen/liver accumulation

architectural defectiveness and high degree of vascular density generate abnormal tumour vessels, aberrant branching and blind loops of twisted shape. Blood flow behaviour, such as direction of blood flow, is also irregular or inconsistent in these vessels. When compared with normal vessels, tumour vessels are ‘leaky,’ owing to basement membrane abnormalities and a decreased number of pericytes lining the rapidly proliferating endothelial cells. The pore size of tumour vessels varies from 100 nm to almost 1 mm in diameter, depending upon the anatomic location of the tumours and the stage of tumour growth. Moreover, as the functionality of tumour lymphatic vessels is generally poor, solid tumours are characterised by decresed clearance of macromolecules and their consequent retention in tumor interstitium. The unique pathophysiologic characteristics of tumour vessels coupled with poor lymphatic drainage induce the EPR effect, which enables macromolecules to extravasate through these gaps into extravascular spaces and accumulate inside tumour tissues. After pioneering results [16], researchers have capitalised this concept for the delivery of various drugs by either building drug-polymer conjugates or encapsulating drugs within nanocarriers. Nowadays, it is evident that long-circulating nanosized carriers accumulate passively in tumours due to the EPR

effect [17] and the vast majority of nanomedicines developed for drug targeting to tumours rely on this effect. After approval of liposomal doxorubicin (Doxil), Abraxane and Genexol-PM have followed as first examples of marketed products based on polymeric nanocarriers. Several additional passively tumour-targeted nanomedicines are currently in clinical trials, and a large number of other ones are in earlyand late-stage pre-clinical development. Nevertheless, exploiting EPR effect to deliver a drug to a solid tumour is a complicated matter [18,19]. The journey of nanocarriers in the body has been recently reviewed [20]. Natural elimination processes, including both renal clearance and mononuclear phagocyte system (MPS) uptake, play a pivotal role in driving nanocarrier circulation in the body. Filtration of particles through the glomerular capillary wall is highly dependent on molecule size and is referred to as the filtration-size threshold. Molecules with a diameter of < 5 nm are typically filtered, while those with a size > 8 nm are not typically capable of glomerular filtration. The uptake of nanocarriers in MPS, also referred to as the reticuloendothelial system (RES), proceeds quickly and must be avoided in order to have an acceptable circulation time MPS, is the main natural clearance system for particles not filtered by the kidneys acting via phagocytosis. Recognition by the MPS is aided by opsonisation, which induces nanosystem phagocytosis and accumulation in the liver. Protein adsorption is driven by nanocarrier properties and has been considered a key factor to control nanocarrier biological behaviour ab initio [21,22]. Several tools are available to exactly identify composition of bound protein cloud [23]. The liver acts as a reservoir toward nanocarriers, conditioning their rapid first-phase disappearance from the blood and, in case of biodegradable systems, their second-phase release in the body under degraded and excretable form. This biodistribution can be of benefit for the chemotherapeutic treatment of MPS-localised tumours (e.g., hepatocarcinoma or hepatic metastasis arising from digestive tract or gynaecological cancers, bronchopulmonary tumours, myeloma and leukaemia). Ideally, an injectable nanocarrier has to be small enough to avoid internalisation by the MPS but large enough to avoid renal clearance (100 -- 200 nm). Recent findings highlight that variation of nanocarrier dimension in the scale length > 100 nm can heavily affect blood circulation time as in the case of filomicelles [24], whereas the role of geometry in driving in vivo biodistribution has not been clarified yet [25,26]. In order to overcome the opsonisation of nanocarriers, a number of strategies have been investigated to make them ‘invisible’ to the immune system, creating long-circulating nanocarriers, known also as ‘stealth’ systems. Hydrophilic polymers can form a cloud on nanocarrier surface, which repels opsonins and possibly adsorbs disopsonins giving decreased levels of uptake by the MPS and increased circulation time in the blood [27]. To disclose the mechanistic action of hydrophilic clouds around nanocarrier and manipulate its PK through rational design, relevant literature has been produced in the last few years [28]. Longevity in the blood can promote


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Intravenous – Hydrophilic and non-immunogenic surface is required for long-circulating NPs. Oral/pulmonary – A mucoadhesive surface prolongs in situ permanence; a muco-inert surface improves transport through mucus. Oral delivery – pH sensitive shells protect NPs toward stomach harsh environment; some surface modifications (PEGylation) can inhibit P-gp efflux. Intraperitoneal – Surface lowering peritoneal adhesion capability while increasing residence time is required. Intravenous – Targeting ligands may increase specificity toward tumor tissue/cells but cause the ‘binding site barrier’ effect. Intravenous/pulmonary – Magnetic NPs integrating SPION can be concentrated at tumor level. Oral/pulmonary – Targeting ligands may increase specificity toward tumor tissue/cells. Intratumoral – Targeting ligands interfere with NP penetration inside tumor mass. Intraperitoneal – Tumor cellspecific ligands may limit cytotoxicity toward local healthy cells, which are potentially exposed high chemoterapeutic concentrations.

Biomimetic/bioresponsive coating

Targeting agents

Core-shell NP

Size (nm)

Surface charge

Intraperitoneal – Large NPs (500 – 700 nm) are preferred to small NPs (250 nm) to avoid clearance from the peritoneal cavity and systemic absorption but can cause peritoneal adhesions. Oral – < 500 nm – Particle size lower than 500 nm is preferred to optimize cell uptake and to enhance drug bioavailability after oral administration. Pulmonary – < 500 nm – Particle size lower than 200 nm is preferred for non muco-inert NPs to diffuse through the mucus layer; NPs around 200 nm are taken up less efficiently by lung macrophages. Intravenous – 100 – 200 nm – This size is required for rigid and biominetic NPs to avoid internalization by MPS but the situation can dramatically change for deformable NPs.

Intravenous – Surface change can be used to modulate NP/cell interactions as well as interaction with blood components Oral and pulmonary delivery – Positively charged shells may cause strong NP/mucin interactions Intratumoral – Neutral shells are desirable to penetrate into tumor matrix

Figure 1. A summary of the main design rules/issues related to nanoparticles administration.

nanocarrier accumulation in solid tumours passing through their leaky vasculature (passive targeting). Particles with hydrophobic surfaces, in fact, will preferentially be taken up by the liver, followed by the spleen and lungs, while particles with long circulation times should be 100 nm or less in diameter and have a hydrophilic surface in order to reduce clearance by macrophages [29]. As a result, non-targeted nanocarriers accumulate in tumours of xenograft mice models in the range 1 -- 4% of the injected dose/g of tissue although these numbers are difficult to derive due to different post-injection time assessments and tumour type [29]. Furthermore, it is not clear whether circulation half-life or maximum tolerated dose is the most critical parameter for optimal accumulation in tumour tissue of a given drug [30]. In the case of humans, it is very hard to compare PK of different systems due to intervariability as well as cancer staging. Furthermore, a steric stabilisation of nanocarrier surface limits also aggregation between particles themselves in the blood and contributes to system stability in biological environments. 286

Nanocarriers prepared from poly(ethylene glycol) (PEG)modified polymers are certainly the most explored systems for intravenous cancer therapy. PEGylation facilitates MPS escape, significantly increases blood circulation time and benefits EPR-based targeting of drugs to tumours [31]. Besides hydrophilicity and surface density, chain flexibility is a major feature to provide prolonged drug circulation since PEG cloud should in theory prevent immune system from modelling an antibody against it. However, it has been recently demonstrated that PEGylated NPs and micelles can lose their long-circulating characteristics when administered within a certain time interval from injection of the first dose, in the same animal -- the so-called accelerated blood clearance (ABC) phenomenon [32]. This effect has been attributed to the fact that PEG in itself may be immunogenic and that the induced anti-PEG antibodies are linked to enhanced blood clearance, accumulation in the liver and reduced efficacy of the products. This unexpected phenomenon, being a potential challenge for PEGylated nanocarrier applications,




would greatly compromise the benefits of nanocarrier use for cancer diagnosis and therapy and will be worth of further investigation in the next future. In fact, the induction of the ABC phenomenon seems to be a complicated process. The following mechanism was very recently set to explain it: once PEGylated nanocarriers (first dose) reach the spleen, they bind and cross-link to surface immunoglobulins on reactive B cells in the splenic marginal zone and consequently triggers the production of anti-PEG IgM. Upon administration of the second or subsequent dose, if anti-PEG IgM still exists in the blood circulation, it binds to surface PEG, and subsequently activates the complement system, resulting in opsonisation by C3 fragments and enhanced uptake by Kupffer cells via complement receptor-mediated endocytosis [33]. Nevertheless, as underlined by the authors themselves, this mechanism can only partly explain ABC, which could not be totally reversed upon spleen removal, suggesting that other factors/ tissues are involved. Of note is the recent finding of a 22 -- 25% occurrence of anti-PEG IgM in healthy blood donors, which is a very high value if compared with the 0.2% occurrence reported 20 years ago [34]. The authors attributed this phenomenon to an improvement of the limit of detection of antibodies during the years as well as to greater exposure to PEG-containing cosmetic, pharmaceutical and processed food products. Nevertheless, it has been also highlighted that several literature indications derive from assays for anti-PEG antibodies that are flawed and lack specificity [35]. Also the biological effects induced by anti-PEG IgM antibodies lack the characteristics of a bona fide antibody reaction, which is a point that will require other research efforts [35]. A paradigm shift in cancer therapy resides in precise control over the exposure of tumour-associated cells to anti-cancer drugs, which can result in therapeutic outcome by regimens that control drug levels, dosing intervals and drug retention in tumour. Thus, strategies allowing the exposure of cells to metered dose of an anti-cancer drug over longer periods can act more effectively than the same dose as a bolus (metronomic approach) [36]. Through exerting direct and indirect effects on tumour cells and their microenvironment, metronomic scheduling can inhibit tumour angiogenesis, stimulate anti-cancer immune response and directly affect tumour cell growth. To this purpose, nanoplatforms with a sustained drug release may be highly beneficial in cancer treatment. Nevertheless, certain PEG-modified particles are also now understood to have a slower uptake into tumour cells than non-PEGylated molecules, thus generating drug gradients in the extracellular space of tumour site reaching further hypoxic regions. To do this, the release rate of the drug from the carrier at the target must be optimal (e.g., 3 -- 10%/day), because a too slow release results in insufficient concentrations of active drugs at sites of action. Release that is too rapid would lead to a high concentration of free drug in circulation but no drug accumulation in the tumour, the results thus being a considerably lower therapeutic effect and undesired systemic toxicity.

Actively targeted nanocarriers Targeting cancer cells through EPR effect is not feasible in all tumours because the degree of tumour vascularisation and porosity of tumour vessels can vary with the tumour type and status [37]. A positive correlation clearly exists between tumour growth rate and the EPR effect, rendering the clinical significance of vascular permeability effect in drug delivery much debated. Little is known about the actual effectiveness of the EPR effect in metastatic or microscopic residual tumours at late stage of development, where targeted chemotherapy is most desired. To improve drug accumulation at tumour/cancer cell level, different approaches can be followed. By simply changing surface properties (charge, shielding cloud), one can manipulate nanocarrier interaction with cell membrane [38]. Nevertheless, decoration of nanocarriers with different functional motifs is the approach most represented in the literature to build up actively targeted nanoconstructs [39,40]. Surface exposure of a ligand recognising a receptor over-expressed in cancer cells (folate, transferrin, prostate specific membrane antigen-interacting peptides or aptamers) is expected to increase drug amount inside cancer cells by activating receptor-mediated transport mechanisms [2,41]. Another approach resides in the decoration of nanocarrier surface with targeting moieties able to recognise tumour microenvironment (surface receptor on tumour blood vessels, as in the case of RGD peptides or extracellular matrix) [2,41]. Many active targeting strategies have been developed, but despite this truly remarkable potential in applications, they have met with mixed success to date and the reasons for that are nowadays deeply discussed [42-47]. An important drawback of targeted NPs is that they can paradoxically lose targeting ability in a biological environment. First, interaction with other proteins in the medium and the formation of a ‘protein corona’ can screen the targeting molecules on the surface of NPs and cause loss of specificity in targeting [48]. Another potential challenge related to molecular targeting is the so-called ‘binding-site barrier’ effect, which may confine the drug in the perivascular regions retarding drug/NP tumour penetration [49]. This effect is mainly dictated by particle size and ligand--receptor affinity. In fact, the diffusion of extravasated NPs from the perivascular into the avascular region will be inversely correlated with their molecular size, which increases upon exposure of surface ligands. On the other hand, when a high-affinity ligand diffuses into the tumour, the slow off rate mediated by its long association may preclude the docking of other ligands to the same receptor, impeding them from traversing tumours using sequential binding to and dissociation from unoccupied receptors [50,51]. Assuming targeted nanocarriers successfully overcome these drawbacks, the system has often to be taken up by targeted cancer cells to be effective. Several studies have shown that the cellular uptake of NPs can depend on many factors, including biological features. This is the case of the proven 2.2


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role of cell cycling not only on uptake rate but also on the amount of NPs internalised by cells [52]. In fact, internalised NPs are not exported by cell but are split between daughter cells when the parent cell divides. Thus, in a cell population, the dose of internalised NPs in each cell significantly varies as the cell advances throughout the cell cycle. In a critical perspective, more focused and dedicated approach to the understanding of the in situ (in vivo) interface between engineered nanomedicines and their targets is thereby needed to transfer the remarkable possibilities of nanoscale interactions in biology into therapeutics [53]. Nevertheless, some targeted nanoparticulate products, such as CALAA-01 and BIND-014, have recently entered in clinical trials. Stimuli-responsive nanocarriers The knowledge of peculiar tumour features can be useful to activate the release specifically at tumour level and has prompted toward stimuli-responsive materials, which offer an interesting alternative to passive and molecular targeting to tumour. To this respect, lower extracellular pH, mild hyperthermia or hyperthermia induced through clinical procedures can be exploited through development of nanoplatforms based on pH responsive polymers [54] or thermosensitive materials [55]. It is well established that in solid tumours, the extracellular pH can be significantly more acidic (~ 6 -- 7) than systemic pH due to poor vasculature and consequent anaerobic conditions prevailing in the malignant cells [56]. Besides, the cellular organelles also exhibit sharp pH differences in different locations; for instance, in cytosolic, endosomal and lysosomal compartments. In fact, a great number of engineered nanomedicines should deliver their drug cargo at intracellular level, eventually targeting specific organelles [38,56,57]. This is especially complicate for nucleic acid drugs that degrade under harsh lysosomal conditions. Nevertheless, systems truly targeting tumour pH are still limited [54]. This may be a result of the narrow pH window that requires a sharp transition of the system as well as variation of pH estimate values with tumour type and volume. Other factors include characteristics of vasculature in pathological states, which determine also the cellular micro-environment [58]. For example, poorly vascularised tumour tissues facilitate the formation of hypoxic cells with low oxygen partial pressure, low pH and poor nutrient level. Notably, hypoxic areas are environmentally reductive due to the presence of bioreductive enzymes, such as glutathione (7- to 10-fold higher in tumour than in normal cells), which can also act as a stimulus for nanocarriers equipped with a disulfide cross-linked moiety that is sensitive to intracellular GSH levels [59]. Physical stimuli, such as temperature, light, magnetic or electrical fields, can be also applied externally to bring about a triggered release of the active guest [56]. Thermo-responsive polymers have become particularly attractive candidates for designing physically targeted nanocarriers. Nevertheless, attention should be paid to the temperature range within which the thermo-responsive nanocarriers should release their 2.3

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cargo, which should fall between 37 and 42 C, as above this temperature protein denaturation and disruption of fine anatomical structures are possible [55]. Furthermore, the benefit of hyperthermia may be counterbalanced by a decreased blood flow, since hyperthermia-induced dilatation of pre-capillary arterioles results in decreased arteriolar--venous pressure gradient [58]. Amid targeting strategies based on the application of external stimuli, magnetic targeting are also garnering substantial attention for drug delivery applications [47,60]. After accumulation of a magnetic carrier, generally small paramagnetic iron oxide NPs (SPIONs), at the target tumour site in vivo, drugs are released through different mechanisms and effectively taken up by tumour cells. In this case, the major role of a non-magnetic polymer shell is to protect the functional core from damaging environments like oxygen, as well as to make them compatible with biological media, improving also NP stability in blood (the smaller and more hydrophilic the NP surface, the longer its plasma half-life) [61]. 3.

Lung delivery

Although dated back of 30 years, the opportunity to selectively target a drug to lungs remains a fascinating option to strongly limit ubiquitous distribution of conventional chemotherapeutics used to treat lung cancer and has been recently fostered by the fast evolution of lung delivery technologies [62,63]. Among them, nanocarrier delivery to the lungs is an attractive concept because it can cause retention of the carrier in the lungs (particles of a few hundred nanometres represent a tenacious resident of the lungs) accompanied with a prolonged drug release, meaning improved lung bioavailability. Furthermore, nanocarriers represent a promising tool to penetrate airway barriers, which can be better overcome at nanosize level [64]. Finally, carriers of nanometric size offer the potential of drug targeting to specific lung tissue and cell populations [65]. Particle engineering at the nanosize level From a technological point of view, the development of effective nanocarriers for inhalation first requires adequate engineering of the particles at the nanosize level to prolong drug residence time on mucosa and to allow penetration in the underlying epithelia [66]. Respiratory epithelia are covered by a mucus layer, which represents a physical barrier that may strongly affect drug bioavailability and targeting [67,68]. The amount, the composition and the rheological properties of mucus, its turnover, along with washout of fluids that line epithelia, are the main determinant of in situ permanence and extent of drug absorption. Actually, the delivered molecule may be entrapped within mucus and, not only prevented to reach epithelium, but also removed by multiple clearance mechanisms [64,69]. Amid them, mucociliary clearance, designed by evolution to eliminate inhaled and possibly noxious material from the airways, may considerably limit the benefit of inhalation therapy. In 3.1




case of severe lung diseases, the overproduction of a viscous and highly complex mucus may further limit the amount of drug reaching the target (cancer cells or, better, organelles inside them) [67,70]. Thus, the choice of the adequate delivery system is of great relevance, since the physical--chemical and technological properties of the carrier (e.g., size, surface charge, loading efficiency) will greatly affect its trafficking in the extracellular environment and, subsequently, that of the released material [69]. The potential for access of nanocarriers to airway mucus as well as their uptake by resident macrophages can be first tuned at size level. While large particles with a mean size around 500 nm may be blocked by mucus layer, NPs smaller than 200 nm are expected to diffuse more easily through the mucus layer [71,72]. Particles < 200 nm are also taken up less efficiently by circulating macrophages, which avidly phagocytose 1 -- 2 µm particles [73,74]. Nevertheless, slight differences on surface properties may have significant implications in the cellular uptake and the interactions of the carrier with the biological environment. Thus, the role played by the surface chemistry of the carrier on its interaction with/adhesion to lung cellular and extracellular barriers should be also considered. The first generation of nanocarriers for mucosal drug delivery was designed to interact with mucus in a non-specific manner, through the use of hydrophilic mucoadhesive polymers, either alone or incorporated in the carrier to modify its surface. Excellent mucoadhesive performance is typically observed for polymers possessing charged groups capable of forming strong hydrogen bonds with mucosal surfaces, such as chitosan, sodium alginate or hyaluronic acid [75]. Some of them (e.g., chitosan) may also enhance drug permeability through the epithelium by modifying the tight junctions between the cells. Nevertheless, particular attention must be paid on imparting a charge or adhesiveness to NPs for inhalation. When modified on the surface with carboxyl groups, also NPs as small as 59 nm may be completely immobilised by human mucus [76]. On the other hand, hydrophilic poloxamer-coated poly(lactic-co-glycolic acid) (PLGA) NPs, displaying a negative charge, have been demonstrated to diffuse more easily across the mucus barrier leading to a higher intracellular accumulation as compared to positively charged NPs modified with chitosan [77]. Notably, surface modification with mucus-interacting polymers may have a huge impact on NP biodistribution along the airways, as reported for chitosan-modified PLGA NPs [78]. More recent strategies to overcome the airway mucus barrier rely on the use of non-adhesive NPs achieved by covalent modification of polystyrene particles with PEG, the so-called mucus-penetrating particles (MPP) [64]. MPP with adequate size (200 -- 500 nm) may deeply penetrate a variety of human lung mucus secretions independently upon their viscosity (i.e., cystic fibrosis sputum, chronic rhinosinusitis mucus as well as mucus from healthy patients), whereas comparably sized uncoated particles are immobilised by the mucus

meshes [79-81]. By this way, NPs can reach the surface of epithelium, thus allowing a prolonged interaction of the nanocarrier with the targeted substrate. In the light of these numerous and often controversial literature reports, a thorough understanding of carrier interactions with airway mucus appears a condition sine qua non to engineer NPs for crossing human mucus barriers and, thus, for lung delivery. Assuming the carrier successfully gets away with the extracellular barriers, the drug has still to interact with the target cells of the respiratory tract to be effective. To this purpose, a firm control over particle--cell interactions is required through engineering of composition and size of the delivery system [65]. This can be very challenging in case of advanced anti-cancer agents, such as nucleic acid derivatives (DNA, siRNA and oligonucleotides), since the drug has to cross the cellular membrane and gain access into the cytoplasm/nuclei, where the final targets are located [69,82]. Therefore, the main function of the carrier, along with macromolecule protection against enzymatic attack, is to facilitate cellular uptake, promote its endosomal escape and release it at the final intracellular target. This can be achieved by adequately engineered nanoparticulate systems, comprising not only compatible polymers able to compact, protect and possibly sustain nucleic acid delivery in situ, but also a biomimetic shell containing agents able to facilitate endo-lysosomal escape (e.g., fusogenic peptides or lipids, endosome-destabilising polymers) [82,83]. Particle engineering at the microsize level Once engineered at nanosize level, effective pulmonary delivery of nanocarrier-based formulations demands for devices and inhalable formulations, which will play a crucial role in determining drug deposition along the airways [84,85]. At the simplest level, nanocarrier properties may be engineered to achieve stable nanosuspensions to be delivered via jet and advanced ultrasonic nebulisers, which are still considered the devices of choice for inhaled NPs. The balance of repulsive/attractive forces (electrostatic and steric, van der Waals), thermal energy (i.e., Brownian motion) and densityand size-dependent gravitational forces will control the stability, and hence the quality, of the formulation [84]. As a consequence, surface charge and composition as well as size of the particles should be properly taken into account when designing effective aerosols based on nanoparticulate systems. Conceiving this system for aerosol delivery, of relevance is the maintenance of a mass median aerodynamic diameter (MMAD) within the suitable range (0.5 -- 5 µm) for lower airway deposition upon nebulisation. To improve drug efficacy in the lungs, it may be advantageous to control aerosol deposition and target aerosols to diseased or disease-causing lung tissue and cellular structures in order to maximise drug potency and minimise side effects in unaffected tissue. With this idea in mind, targeted core--shell NPs have been also investigated for lung delivery of chemotherapeutics, showing in vitro and in vivo-specific accumulation in cancerous lung cells/tissue [86-88]. To bring aerosolised chemotherapeutics to 3.2


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an advanced level of specificity, also magnetic gradient fields can be used to direct magnetisable aerosol droplets containing SPIONs -- the so-called ‘nanomagnetosols’ -- specifically to desired regions of the lungs in mice [89-91]. Although nebulisers are often preferred in the clinical practice, a great deal of interest was recently aroused in delivering anti-cancer drugs directly to the lung surfaces via hand-held pressurised metered dose inhalers (pMDIs) or breath-actuated dry powders inhalers (DPIs). Although less frequent, the delivery of drug-loaded nanocarriers to the lung may be accomplished by non-aqueous suspensions to be used in pMDIs [92]. In this case, the formulation strategy is directed to efficiently disperse nanocarriers in hydrofluoroalkane propellants, and to generate aerosols suitable for deep lung deposition. As most suspension-based pMDIs comprise micrometric particles, the likelihood of nanosized carriers to be transported in the aerosol is very low. The greatest challenge is the low inertia of particles with a mean size lower than 1 µm, which are generally exhaled upon inhalation [73,93]. In recent years, different strategies to engineer NPs at the microsize level and also achieve ‘micrometric’ NP-based dry powders for inhalation are being explored [85,94]. In so doing, all the advantages of NPs for inhalation (e.g., long-term residence in situ; prolonged drug release; macrophage targeting) could be combined with the ease of use of DPIs [95]. Since drying of nanosuspensions to develop DPIs is very difficult to achieve, given that NP aggregate excessively in the dry state, large porous carriers (so-called ‘Trojan particles’) [96] or microparticles based on inert materials (so-called ‘nanoembedded microparticles’) [78,97] may be used to inhale NPs in the form of dry powders and release them once reaching respiratory tract. A similar approach has been employed to efficiently carry doxorubicin-loaded core--shell NPs with anti-cancer agents in the lung [98]. This less-invasive route of administration might change the way lung cancer is treated in the future. Nevertheless, one should be aware of potential health and safety risks associated with handling cytotoxic dry powders in the healthcare industry. 4.

Oral delivery

In the past decades, the oral administration of anti-cancer drugs is raising great interest in the scientific community, as suggested by the growing number of drugs commercially available and by the orally formulated agents in development. In fact, oral cancer therapy provides many attractive benefits as compared to parenteral routes, such as the low cost and the ease of administration, associated to great convenience and life quality of the patients [99,100]. Nevertheless, the systemic efficacy of orally administered anti-cancer drugs is seriously compromised by their poor oral bioavailability. Before the drug molecule reaches its final destination, it must go through the stomach, the lumen of the intestine, the mucus layer coating the intestinal epithelium and, finally, the epithelium itself. Both the highly acidic condition in the 290

stomach as well as the digestive enzymes that are secreted along the entire gastrointestinal (GI) tract can potentially degrade the drug before it moves across the epithelium. Furthermore, numerous chemotherapeutics show low permeability across GI mucosa, due to not only the limited solubility and structural instability in GI fluids, but also the affinity for intestinal and liver cytochrome P450 metabolic enzymes (first-pass extraction mechanism) and to the P-glycoprotein (P-gp) involved in MDR [99,100]. Adhesion and diffusion of the drug in the mucosal microenvironment of the GI tract represent another important challenge in oral chemotherapy. In fact, GI tract is covered by a mucus layer, which is a very complex protection systems from the pathogens and foreign agents and has been highlighted as a significant barrier for drug penetration [101]. Finally, the drug transport across GI epithelium can occur through: i) the paracellular transport, which depends from the opening of tight junctions, ii) the transcellular transport by active endocytic mechanism, such as pinocytosis and clathrin-mediated endocytosis and iii) the lymphatic mediated transport, which is achieved preferentially after absorption by M cells, relatively less protected by mucus and characterised by a high transcytotic activity [102]. Nanotechnology engineering of chemotherapeutics represents an interesting approach in order to overpass the limits related to systemic oral administration of anti-cancer drugs. In particular, core--shell nanocarriers provide the adequate size and possibly the appropriate surface modification to improve drug permeation across GI mucosa and absorption by the intestinal cells as well as to reduce the P-gp efflux and the cytochrome P450-mediated inactivation [103,104]. Improving drug permeation across GI mucosa Particle size and size distribution are considered crucial parameters to optimise drug absorption and distributions, as well as to minimise the toxicity, of the encapsulated drug [102,103]. Particles smaller than 100 nm may allow significant drug penetration across the intestinal mucosa, whilst the oral administration of particles with a diameter of 500 nm may lead to the permeation of a percentage of the administrated dose as low as 10%. Furthermore, if 1 µm particles can be especially accumulated in the Peyer’s patches, no permeation at the lymphoid tissues was observed with particles of higher size [102]. On the other hand, it has been reported that 100 nm particles may diffuse also through the submucosal layer, allowing a systemic distribution, while larger NPs predominantly localise in the epithelial lining of the tissue [105]. Although the correlation between particle size and GI mucosa barrier permeation is often unclear and involves unknown mechanisms, the basic criteria that particle size < 500 nm are preferred in order to optimise cell uptake and to enhance drug bioavailability after oral administration can be accepted [102,103]. Nevertheless, if this cut-off diameter may be useful to optimise in vitro uptake by GI epithelial cells, it can be sometimes detrimental for in vivo penetration 4.1




of the GI mucus layer, which represents a barrier for NPs permeation [106]. To increase the persistence of a nanocarrier at GI mucosa, surface engineering is required so as to cross the loosely adherent mucus layer, which is continuously removed by peristalsis and replaced, and to reach the firmly adherent layer, which is deposed on GI epithelium [106]. As described in the case of lung delivery, NP surface can be modified to enhance or reduce adhesion to mucus layer. The use of mucoadhesive polysaccharides, such as chitosan or alginate, has been widely employed to effectively engineer NPs for drug delivery to GI mucosal tissues [106]. In particular, through electrostatic interactions with mucus components, mucoadhesive NPs can penetrate across the mucus layer and increase the contact between encapsulated drug and the surface of absorptive cells, enhancing drug permeation [107]. In some cases, such as for chitosanbased NPs, improved drug permeation across GI barrier was related not only to the mucoadhesive properties but also to an opening effect on the tight junction [108].

Increasing drug resistance toward GI harsh environment and MDR effect

4.2

Taking into account the variety of conditions expressed in the GI tract, especially the variable pH value, strategies based on polymers enhancing the stability of the delivery system at low pHs or releasing drugs in a pH-dependent fashion are required. An interesting strategy to overcome nanocarrier degradation in stomach relies in the development of core--shell polymeric NPs, in which biodegradable NPs are protected by a coating based on polymers stable at low pH while dissolving at pH 5 -- 6 of small intestine, such as alginates [109]. This strategy, developed to overcome the harsh environment of the GI tract, acquire additional importance when looking for targeted drug delivery to GI-specific areas, such as colon. In fact, NPs made of widely employed biodegradable polymers, such as poly(lactic acid) (PLA), may be embedded in polysaccharide-based hydrogels, such as alginate/chitosan hydrogels, also to actively target drugs to inflamed colonic tissues [110,111]. In addition to enhancing particle stability toward acidic pH, nanocarrier engineering for oral cancer therapy should also take into account the MDR effect exerted by the P-gp, which is extensively expressed in the GI tract [99]. On this matter, NPs combining mucoadhesive properties and the ability to inhibit P-gp-mediated efflux mechanism can be successfully achieved by surface chemical modification with PEG [112]. In particular, PEG is reported to exert an inhibitory effect comparable to that obtained after the administration of verapamil, a calcium blocker and well-known inhibitor of the P-gp efflux pump employed in conventional clinical therapies [113]. Nevertheless, PEG molecular weight may influence NP mucoadhesion properties and, consequently, their permeation through GI mucus [112]. On the other hand, a mucoinert high-density coating of low molecular weight PEG, as observed for lung delivery, may allow the achievement of

MPP that readily penetrate the loosely adherent mucus layer and enter the firmly adherent one [106]. 5.

Other routes of administration

The intrinsic limitations of major administration routes for the delivery of anti-cancer drugs have prompted research interest in novel formulations allowing the use of other delivery routes aimed at localised chemotherapy (i.e., intratumoral injection) or delivery in close proximity to the tumour site. In the last case, topical delivery for hyperproliferative skin diseases, intravaginal administration for neck or cervix cancer, buccal delivery for oral cancer and, last but not least, intraperitoneal delivery of anti-neoplastic agents can be attempted. In each case, a thorough understanding of the factors that influence drug absorption from the site of administration is important in designing nanocarrier-based formulations. Although ensuring high levels of anti-cancer agents at the site of disease, the efficacy of localised intratumoral or peritumoral treatment strongly depends upon the accessibility of the delivered therapeutic agent to the tumour and nearby diseased tissue [114]. In particular, adequate diffusion to reach sites harbouring occult tumour cells is of paramount importance to prevent recurrence. The several factors that affect the rate and extent of drug absorption upon intratumoral delivery of NPs, including physiological conditions (i.e., composition and functions of the tumour extracellular matrix, tumour interstitial fluid pressure), physicochemical properties of the drug, particle size and surface charge as well as the presence on NP surface of tumour-targeting ligands, have been recently reviewed by Holback and Yeo [115]. When the cancer lesion is localised in the peritoneal cavity (uterus, ovaries or peritoneal carcinomatosis), intraperitoneal delivery of NPs may provide relatively high concentration and long half-life of the anti-cancer drug in situ. One of the major technological challenge is the residence time of the drug, since particle in the peritoneal cavity are known to be rapidly absorbed by the lymphatic circulation, as a function of their size (large is better than small) [116]. Although the ultimate fate of nanocarrier surviving the lymphatic circulation has not been fully elucidated yet, literature evidences suggest that NPs cleared from the peritoneal cavity end up in the systemic circulation after passing lymph nodes and ducts. An advanced approach to prevent premature particle clearance, without causing peritoneal adhesion (e.g., micrometric particles), relies on the use of a hydrogel as a carrier medium, such as in situ cross-linked hyaluronan hydrogels [117]. Despite the high potential for the treatment of skin cancer, an accessible barrier such as skin is very complicated to employ for local administration of anti-neoplastic agents due to its intrinsic barrier properties. Passive and active permeation enhancement methods have been widely applied to increase the cutaneous penetration of NPs, resulting in topical delivery (into skin strata) and transdermal delivery (to subcutaneous tissues and into the systemic circulation) [118,119]. In


291



perspectives, some promising systems based on biodegradable polymers, such as PLGA [120,121], PLA [122] and poly("-caprolactone) (PCL) [123] could offer a further advantage of sustained drug release in the skin. In particular, core--shell nanostructures with an external deformable PEG fringe are able to distribute through the skin deep layers, thus overcoming the subcutaneous barrier [124]. To increase residence time in situ, the incorporation of colloidal nanocarriers in proper vehicles with adequate rheology can be envisaged (i.e., Nanogel system). The gel can aid a uniform dispersion of the nanocarriers in the matrix and increase the contact time, which results in enhanced penetration of the drug payload [125] Increasing research interest has been recently devoted also to minor topical routes, such as intravaginal, which is nowadays a proven way to confer local protection against sexually transmitted diseases, often precancerous. Nevertheless, intravaginal delivery is complicated by the mucus gel barrier, the hormone cycle and the harsh mucosal environment. These challenges have been addressed by the novel approaches in drug formulation, including the use of mucoadhesive, pH- or temperature-sensitive or muco-inert NPs [126]. Again, surface modification of the nanocarrier with PEG, so as to achieve core--shell muco-evasive NPs able to deeply penetrate the mucus layer, seems to date the most effective NPs tested for this application [127]. An open question is how to administer NPs locally. In perspective, the particulate formulation can be applied as such in the form of dried powder, liquid suspension or dispersion, or as a NP-containing topical ointment, cream or gel. 6.

Conclusions

The last century has witnessed the maturation of nanomedicine as a viable alternative to conventional chemotherapy for the treatment of cancer. Despite many advances in drug delivery technologies, cancer treatment through nanosized systems releasing anti-cancer agents is still far from optimal because it is plagued by several biological drawbacks. Frequent challenges encountered in designing injectable nanocarriers include MPS-mediated rapid clearance, low stability in the biological environment, non-specific systemic distribution of the anti-tumour agent, inadequate drug concentrations reaching the tumour site. An enhanced understanding of tumour biology as well as the underlying mechanisms of tumourigenesis has boosted the discovery and development of highly specific targeted nanomedicines capable of exerting their effects on individual proteins/receptors over-expressed or aberrant within tumours (actively targeted nanocarriers) as well as in response to stimuli proper of the tumour (stimuli-responsive nanocarriers). These systems have gained considerable recognition, even if they suffer from several deficiencies, which clearly hamper their translation to the clinics. The intrinsic limitations of the intravenous route for the delivery of anticancer drugs have prompted research interest in novel formulations allowing the use of other delivery routes aimed at 292

localised chemotherapy or delivery in close proximity to the tumour site. In particular, the increasing knowledge acquired on physical and anatomical barriers imposed by the tissues covered with mucus, makes oral and lung delivery of anticancer nanomedicines always more feasible. In general, there is still much to do to increase the potential and strength of nanocarriers for cancer and help them, through a rationale biologically driven design, to find a definitive place in the pharmaceutical market. 7.

Expert opinion

‘Nanoncology’ represents an advanced and very attractive approach to overcome the main limits related to conventional dosage forms for anti-cancer drugs. To be successful, design of novel nanoconstructs should take into consideration welldefined rules to overcome limitations associated to each administration route, providing adequate size, composition and, possibly, appropriate surface modification. To date, intravenous injection remains the preferred, and probably the most widely investigated, route of administration for ‘nanoncologicals’ in order to reach different body compartments. Recent innovations, mainly focused on exploring novel surface modifications of the nanocarrier, may allow overcoming clearance in the body, increasing NP extravasation and its subsequent uptake into target tissue/ cells/organelles. However, a high level of complexity of multifunctional core--shell nanoconstructs comprising at same time novel materials, one or more drugs and/or imaging elements, may be sometimes unsuitable to draw general conclusions on nanocarrier biological behaviour. This generates a proliferation of papers but maybe does not allow us to understand the importance of a single parameter. For example, the wellestablished strategy to actively target a drug to the tumour site remains a fascinating option to increase dose fraction reaching pharmacological target but a strong debate is growing on its usefulness. This situation may have more to do with our current understanding and integration of knowledge across disciplines, than any intrinsic limitation on the vision itself. A paradigm shift in cancer therapy resides in precise control over the exposure of tumour-associated cells to anticancer drugs, which can result in therapeutic outcome by regimens that control drug levels, dosing intervals and drug distribution/retention in tumour. Thus, a naive approach, consisting in adequately engineered core--shell polymer NPs with a sustained drug release, may be rediscovered beneficial in cancer treatment, also in the light of metronomic approaches. In the future, it seems essential to carry out cell-level studies in relevant biological conditions, that is, reproducing in vitro the biological environment/fluids in which the particles will work. On the other hand, attention should be paid to animal models employed to study in vivo effects. Results achieved in etherotopic mouse models are not necessarily comparable to those achieved in an orthotopic model. In both cases, the




size and growth rate of tumours in mice, although favourable for the EPR effect, are not comparable to those of human patients, which typically develop over a number of years. Furthermore, little is known about the effectiveness of the passive/active targeting strategies in metastatic or microscopic residual tumours, where targeted chemotherapy is most desired. Increasing efforts should be directed also towards less explored administration routes allowing localised chemotherapy or delivery in close proximity to the tumour site, such as lung delivery. On this matter, the core--shell architecture of the NP may allow the achievement of ‘biomimetic’ nanoconstructs able to deeply penetrate the mucus barrier, which may increase the bioavailability of anti-cancer drugs also upon oral administration. Maybe we should rethink the way to look at the problem considering an experimental set-up to test nanomedicines in Bibliography Papers of special note have been highlighted as either of interest () or of considerable interest () to readers. 1.

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Affiliation


Ivana d’Angelo1, Claudia Conte2, Agnese Miro2, Fabiana Quaglia†2 & Francesca Ungaro2 † Author for correspondence 1 Second University of Napoli, Di.S.T.A.Bi.F, Caserta, Italy 2 University of Napoli Federico II, Department of Pharmacy, Laboratory of Drug Delivery, Via Domenico Montesano 49, 80131, Napoli, Italy Tel: +39 81 678707; Fax: +39 81 678707; E-mail: [email protected]


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Liposomal siRNA nanocarriers for cancer therapy.

Clinical trial design in dermatology: experimental design. Part I.

Nanocarriers Usage for Drug Delivery in Cancer Therapy.

Biologically oriented preparation technique (BOPT) for implant-supported fixed prostheses.

Optimum Design Rules for CMOS Hall Sensors.

Ovarian cancer, Part I: Biology.

Ethnopharmacological Approaches for Therapy of Jaundice: Part I.

Update on therapy for superficial mycoses: review article part I.

NEED ORIENTED WARD DESIGN.

Transformable Peptide Nanocarriers for Expeditious Drug Release and Effective Cancer Therapy via Cancer-Associated Fibroblast Activation.

Optimal structural design of mannosylated nanocarriers for macrophage targeting.

Topical pharmacotherapy for skin cancer: part I. Pharmacology.

Nanocarriers for microRNA delivery in cancer medicine.

Nanocarriers for cancer-targeted drug delivery.

Multifunctional metal rattle-type nanocarriers for MRI-guided photothermal cancer therapy.

Therapy of genital human papillomavirus infections. Part I: Indications for and justification of therapy.

Towards design rules for covalent nanostructures on metal surfaces.

Rationale for design of biologically reversible drug derivatives: prodrugs.

Design rules for efficient transgene expression in plants.

All-optical switching: Three rules of design.

Protein design: rules, empiricism, & nature.

Recent advances in biocompatible nanocarriers for delivery of chemotherapeutic cargoes towards cancer therapy.

Poly(HydroxyButyrate-co-HydroxyValerate) (PHBHV) Nanocarriers for Silymarin Release as Adjuvant Therapy in Colo-rectal Cancer.

Pathway-oriented concepts in adjuvant and neoadjuvant breast cancer therapy.