Advanced Biomaterials and Biodevicess

Nanoparticles: Scope in Drug Delivery 491 Liposomes are small spherical vesicles formed of amphiphilic lipids enclosing an aqueous core. They are mainly carrier systems for hydro- philic drugs [13, 14]. The first formulation of liposomes was prepared in 1986 by the Christian Dior Laboratory in collaboration with Pasteur Institute. The amphiphilic nature of liposomes, the ease of surface modifi- cation and good biocompatibility make them good drug delivery systems. Liposomes are composed of mixture of lipids such as phosphatidylcholine, cholesterol, diacetylphosphate-o-steroylamylopectin, monosialoganglio- side, distearylphoshatidylethanolaminepoly-(ethylene glycol) for targeted delivery of anti-TB drug to the lung [15]. The drug encapsulated lipo- somes significantly help in reduction of the bacterial count in spleen and liver as compared to free drug [16]. Several different kinds of liposomes are being widely employed for cancer treatment and for delivering cer- tain vaccines. During cancer treatment they encapsulate drugs, protecting healthy cells from their toxicity and prevent their concentration in tis- sues such as patient’s kidney, liver, pancreas showing details in Figure 14.2 (previous page). They also reduce side effects related to chemotherapy such as hair loss and nausea [17]. Polymeric nanoparticles have also emerged as a prom- ising tool for targeted and controlled drug delivery. In case of polymeric nanoparticles the drug is encapsulated or entrapped in polymeric core and depending upon method of preparation they are called nanospheres or nanocapsules [18]. Polymeric nanoparticles include those made from chitosan, gelatin, poly-(lactide-co-glycolide), copolymer, polylactic acid, poly-(cyanoacrylate), poly-(methylacrylate). These nanoparticles can be functionalized into other types of nanoparticles to change and improve their biodistribution properties. The surface property mainly plays an impor- tant role in drug targeting. When they come in direct contact with cellular membranes their surface properties help in determining the mechanism of internalization and intracellular localization. Biodegradable polymeric nanoparticles such as PGA and PLGA can be formulated to encapsulate several classes of therapeutic drugs but not limited to low molecular weight compounds. They represent as an alternative to liposomes with much improved drug delivery to target sites and with lesser side effects. Recently stimuli responsive polymeric nanoparticles have been demonstrated that response to an external or internal stimulus such as pH, redox, magnetic field and light [19]. Pandey et al, developed sustained release RIF, INH and PYZ loaded poly-(lactide-co-glycolide) [PLG] nanoparticles for oral deliv- ery in mice. As a result the drugs were detected in the plasma for up to 4 days for RIF and 9 days for INH and PYZ; whereas in the tissues were detected till 9 to 11 days after single oral administration of nanoparticles

492 Advanced Biomaterials and Biodevices whereas free drugs were cleared within 12–24h from the plasma. Five oral doses were sufficient for completely bacterial clearance whereas free drugs took 46 doses to get the same results [20]. Injectable PLGA (poly lactic- co-glycolic acid) nanoparticles were also administered subcutaneously in a murine model. A single subcutaneous dose of PLGA nanoparticles main- tains drug levels in plasma, lungs, and spleen for <1 month and bacterial count remain almost undetectable in these organs [21]. Niosomes are similar to liposomes and are mainly composed of non- ionic surfactant with or without incorporation of lipids. Recently nio- somes were prepared by reverse phase evaporation method and given a charge with a charge-inducing agent, dicetyl phosphate. Drug entrap- ment efficiency was determined by spectrophotometer. In vitro drug release and cellular uptake studies was also carried out on macrophage J774A. As a result the cellular uptake of drug loaded by macrophage cells was as high as 61.8% a level which is capable for effective treatment of tuberculosis [22]. Micelles are sub microscopic aggregates (20–80 nm) of surfactant mol- ecules resulting in liquid colloid [8]. PLA (poly lactic acid) modified chi- tosan oligomer micelles release 35% drug release within 10h followed by more sustained drug release till 5 days suggesting the role of micelles as drug carrier with reduced side effects. Alginate nanoparticles are also used as drug delivery for tuberculosis treatment. Zahroor et al used ionotropic gelation method for preparing alginate nanoparticles (235nm) of anti-TB drug. Oral administration of drug to mice resulted in detection of level free drugs in tissues till next day. Whereas in plasma alginate nanoparticles were detected up to 7 days for ETB (ethambutol), 9 days for RIF, 11 days for INH and 15 days for PYZ in tissues [24]. Gold Nanoparticles, Mesoporous Silica Nanoparticles, Quantum Dots are also employed for detection, imag- ing and treatment of various diseases. Gold nanoparticles preparation involves the chemical reduction of gold salts in aqueous, organic, or mixed solvent system. Under these condi- tions gold surface are extremely reactive as a result aggregation occurs. To reduce aggregation, gold nanoparticles were reduced in the presence of a stabilizer which binds to the surface and remove aggregation via cross link- ing and charge properties. Gold nanoparticles are also used for biological imaging and sensing. A simple gold nanoparticle probe assay was done to detect Mycobacterium tuberculosis and its complex in clinical specimen. An assay using gold nanoparticles derivatized with thiol modified oligo- nucleotides was carried out. The gold nanoparticle probes GP-1/GP-2 for IS6110 and GP-3/GP-4 for Rv3618 were designed to hybridize with target DNAs of MTB and MTBC strains. Then the efficiency of gold nanoparticle

Nanoparticles: Scope in Drug Delivery 493 probes assay was evaluated directly, detecting both MTB and MTBC from 600 clinical sputum specimens [25]. Liposomes and solid lipid nanoparticles because of their poor chemical stability and degradation by the serum resulted in decrease in the drug delivery to the target cells and increasing the potential systemic toxicity [26]. Due to some of the disadvantages related to the liposomes and solid lipid nanoparticles, there was a discovery of mesoporous silica nanopar- ticles. Mobil first discovered the mesoporous silica based nanomaterials MCM-41 in 1992. They are 100nm sized silica nanoparticles with 2nm sized pores. These pores run parallel through mesoporous sized nanopar- ticles and form hexagonal pattern [27]. These biocompatible solid mesopo- rous silica nanoparticles framework of hexagonal pattern provides it more intrinsic stability as compared to the liposomes, copolymers or polymeric nanoparticles drug delivery platforms. Some of the in vivo experiments resulted in more enhanced blood stability of mesoporous silica nanopar- ticles as compared to liposomes and polymeric nanoparticles and have favourable biodegradability, biocompatibility and excretion properties [28, 29, 30]. 14.3 Tuberculosis Targeting Nanoparticles Tuberculosis is a ubiquitous and highly contagious chronic granulomatous bacterial infection. It is one of the leading causes of mortality and mor- bidity worldwide. In the year 1993, World Health Organization (WHO) declared tuberculosis as a global emergency. Tuberculosis is caused by M. tuberculosis, which infects about one third of the World’s population and causes approximately 9 million new cases of active tuberculosis and 1.7 million deaths annually [31]. Most of the cases occurred in South East Asia (55%) and the African regions (30%). The five countries with largest num- bers of the cases include India, China, South Africa, Nigeria and Indonesia. Out of the 9 million new cases of TB, 15% were HIV positive, 78% of these HIV positive cases were in the Africa and 13% of the cases were in South East Asia regions [8]. It serves as second most common cause of death after HIV/AIDS. Treatment of active TB is becoming more and more complex due to the emergence of MDR-XDR strains. The identification of multidrug-resistant (MDR) strains, makes mycobacteria resistant to at least Rifampicin (RIF) and Isoniazid (INH) (two first line anti-TB drugs) and extensively drug resistant (XDR) strains defined as MDR strains with additional resistance to the fluoroquinolone. MDR-TB has become one of the major obstacles to effective global TB control. To solve these problems

494 Advanced Biomaterials and Biodevices of drug resistance WHO (World’s Health Organization) implemented DOTS (Directly Observed Treatment, short course), program which was not successful in solving the patients non-compliance [32]. As Figure 14.3 below discloses, most patients fail to adhere to multidrug therapy for lon- ger periods of time. Tubercle bacilli are slender, acid fast, non-motile gram-positive bacilli. These bacteria remain viable in the air for longer period of time as a result of which they get inhaled by the lungs and engulfed by alveolar macro- phages (white blood cells) where they start replicating within 2–3 weeks [33]. If these bacteria is not completely destroyed, they remain dormant for several days and may reactivate years later. The cell wall has high lipid content which allows the bacteria to survive within macrophages. Because of the impermeable nature of the cell wall, drugs are only partially effective and organisms develop resistance [4]. The bacilli can spread from the site of infection in the lung through the lymphatic or blood to the other parts of the body. Extra pulmonary TB of the pleura, bone, genitourinary sys- tem, skin or peritoneum, meninges occurs in approximately 15% of the TB patients [8]. Risk of development of active disease varies according to time according to infection, age, host immunity, however life time risk of death for a newly infected child has been estimated 10% [34–36]. Available TB treatment involves daily administration of four oral antibiotics for a period Mycobacterium tuberculosis from air get engulfed by alveolar macrophages Free Infected macrophages entered Mycobacterium tuberculosis into blood vessel released from granuloma center Life Cycle of Mycobacterium tuberculosis Formation of caseum Differentiation of infected (cells are in dormant phase) macrophages into foamy macrophages Formation of fibrous cuff made up of extracellular matrix around foamy macrophages Figure 14.3 Lifecycle of Mycobacterium tuberculosis.

Nanoparticles: Scope in Drug Delivery 495 of six months or more. The daily administration of antibiotics has side effects like nephrotoxicity and ototoxicity and longer duration of treatment has resulted in low patience adherence [37]. To improve the treatment and reduction in death cases due to tuberculosis, various chemotherapeutic strategies against TB include :- (1)-Derivatization of the existing Anti- tubercular drugs (ATDs) into more potent compounds. (2)-Screening of the compounds which are active against replicating as well as latent bacilli. (3)-Identification of novel drug targets and designing of appropriate inhib- itors. (4)-Targeting of host-pathogen related processes which are essential for the survival in the human diseases [38]. The currently used vaccine for the tuberculosis treatment is Bacillus- Calmette-Guerin (BCG) which provides extremely limited protection. Recently, several Mycobacterium tuberculosis antigens have been identi- fied for potential use as vaccine antigens, including three protein Antigen 85 complex (Ag85A, Ag85B and Ag85C), the surface exposed lipoproteins PstS (PstS-1, PstS-2 and PstS-3) and early secretary antigenic target pro- tein ESAT-6 [39–41]. Cationic liposomes entrapped antigenic tuberculosis ESAT-6 protein, complexed with TLR agonists was evaluated as a prophylac- tic vaccine system by Zaks research group [42]. These control measures for TB such as BCG vaccination and chemoprophylaxis has proven to be unsat- isfactory, so there come anti-TB drugs the only option for TB treatment. The goal of treatment with drugs is to cure without relapse, to prevent death, to stop transmission, and to prevent the emergence of drug resistance [43]. As suggested by WHO, treatment of TB and drug resistant cases required multi- drug therapy, comprising (a): an initial intensive phase of rifampicin (RIF), isoniazid (INH), pyrazinamide (PYZ), ethambutol (ETB) for 2 months; and (b): a continuation phase of RIF and INH for further 4 months, either daily or 3 times per week has to be administered [44]. (Figure 14.4) 14.3.1 Action of anti-TB drugs Rifampicin (RIF)-Inhibition of bacterial RNA synthesis by binding to the β subunit of bacterial DNA dependent-RNA polymerase, as a result it inhibits RNA synthesis. It is one of the most effective anti-tuberculosis agents and is bactericidal for extra and intracellular bacteria [45, 46]. Isoniazid (INH) - It is most active drug for treatment of TB caused by susceptible strains. It is a prodrug activated by katG, which exerts its lethal effects by inhibiting the synthesis of mycolic acids, an essential component of mycobacterial cell walls through formation of covalent complex with an acyl carrier protein (AcpM) and KasA, a beta-ketoacyl carrier protein synthetase [45, 47].

496 Advanced Biomaterials and Biodevices Cell wall and cytoplasmic membrane Mycolic acid INH Arabinogalactan ETB Cytoplasm Short chain fatty acid PYZ precursors RNA polymerase RIF (beta subunit) Systematic diagram for the site of action of principle anti-TB drugs Figure 14.4 Site of action of principal anti-tuberculosis drugs against M. tuberculosis H37Rv. Pyrazinamide (PYZ) - It gets converted into pyrazanoic acid, which lowers the pH of the surroundings of M. tuberculosis and thus organism unable to grow. It is also antimetabolite of nictoniamide and interferes with the synthesis of NAD, thus inhibiting synthesis of short chain fatty acid precursors [45, 47]. Ethambutol (ETH) - It inhibits mycobacterial arabinosyl transferases involved in the polymerization of D-arabinofuranose to arbinoglycan, an essential cell wall component of mycobacterium [45, 47]. INH, ETB along with streptomycin helps in eradicating most of the rapidly replicating bacilli in first 2 weeks of treatment. RIF and PYZ are two drugs which play an important role in sterilisation of lesions by eradi- cating organisms, and are crucial for 6 months treatment regimens. RIF is responsible for killing non-replicating organisms and high sterilising effect of PYZ act on semi-dormant bacilli and not affected by any other TB agents, in sites hostile to the penetration and action of the other drugs [48, 49]. Despite of having availability of highly effective treatments of TB, cure rates remain low. Patients have to consume large amount of drugs, the major cause of patient’s non-compliance. These short course regimens results in decrease of the therapeutic potential of patient resulted in escala- tion in the mortality rate and increased risk of developing acquired drug

Nanoparticles: Scope in Drug Delivery 497 Encapsulation Nanoparticle Drug Nanoparticle with drug Nanoparticle get Degradation of nanoparticle to endocytosed by the release drug cell Uptake of nanoparticle through intracellular membrane; localisation and internalisation Figure 14.5 Systemic uptake of nanoparticle through intracellular membrane, localization and internalisation. resistance [50, 51, 52]. Resistance of M. tuberculosis to anti-TB agents is a worldwide problem in both immunocompetent and HIV-infected popula- tions [53, 54]. Though many new antibiotics have come into existence but treatment of such intracellular pathogens still remains a problem as the infection remain localized within phagocytic cells and most of the antibi- otic are highly active in vitro, so they do not actively pass through cellular membrane and therefore it’s difficult to achieve the relatively high concen- tration of the drugs within the infected cells [55, 56]. To solve this problem of intracellular chemotherapy there is a need to design such a carrier system for antibiotics that could efficiently endocy- tosed by phagocytic cells and once inside the cells should prolong release of antibiotics so that the number of doses frequency and drug toxicity can be reduced as in Figure 14.5 (above). All these problems associated with the chemotherapy led to the inves- tigation of drug carriers for treating intracellular pathogens such as anti- biotics loaded into liposomes, microspheres, polymeric nanoparticles, and nanoplexes [57, 58]. Present efforts are being in progress in improving treatment of diseases by shortening time period of treatment or using new carrier based drug delivery strategies in addition to alternative adminis- tration routes, which have important role in improving anti-tubercular chemotherapy efficacy, thus enhancing patient’s compliance, and reduc- ing dosing frequency. Nanotechnology can be defined as formation of the

498 Advanced Biomaterials and Biodevices submicron colloidal particles, which has become a great advancement in the drug delivery [9]. Modification of new drugs as a nanoparticle-based delivery system is feasible, cost-effective, and readily available alternative to chemotherapy. In turn, nanoparticle carrier based drug delivery system hold a significant importance in the reduction of drug resistance TB cases. This system of drug delivery through nanotechnology enhances the effec- tiveness of approved drugs and extends their applicability by overcoming technological limitations, such as low availability, resistance, cellular and anatomical barriers, among others [59]. Different nanoparticulate strategies are being developed to specifi- cally deliver chemotherapeutic compounds to target disease sites. These nanoparticles increase drugs therapeutic index by localizing its pharmaco- logical activity on its target site or organ of action. The particle size between 50 to 200 nm is desired for maximum drug localization upon administra- tion by inhalation [60]. In case of lungs, particles deposition takes place by inertial impaction, sedimentation and diffusion [61, 62]. Large particles (>5μm) get deposited by impaction in the extra-thoracic cavities, particles (1–5μm) deposited deeper in the lungs by inertial impaction and sedimen- tation while very small particles (<1μm) are taken up by diffusion and the most of the particles which remain suspended are exhaled out [63]. Recently, the microencapsulation of pharmaceutical substances in bio- degradable polymers is an emerging technology. Many natural and syn- thetic carriers are used in drug delivery systems. Natural carriers mainly include lipids (liposomes and solid lipid nanoparticles), alginic acid, gela- tin, dextrins etc whereas synthetic carriers include poly-(DL-lactide-co glycolide) (PLG), polylactic acid (PLA), polymethyl acrylates, polyanhy- drides, carbomer etc. Carriers not only help in designing different delivery system but also provide flexibility for selecting the route of drug delivery system [64]. PLG (poly-DL-lactide-co-glycolic acid), copolymer of lactic acid and glycolic acid is completely biodegradable, biocompatible and has role in medical procedures as well as used for encapsulating antibiotics, antigens, peptides in order to develop sustained-release delivery system [65]. Various nanoparticulate systems include polymeric nanoparticles, lipid nanoparticles, nanosuspensions, nanoemulsions etc can be used to treat several types of parasitic infections [66–70]. Dutt and Khuller have entrapped anti-TB drugs such as INH and RIF in PLG polymers. When these nanoparticles were injected subcutaneously as a single dose in mice, the microparticles with a diameter range of 11.75μm to 71.95μm resulted in sustained release of drugs over 6–7 weeks [71]. Sheogkar explored nanoparticulate systems for the anti-TB therapy and briefly described three drugs under clinical trials [72]. Sosnik et al. reviewed the

Nanoparticles: Scope in Drug Delivery 499 development of nano-based drug delivery systems for encapsulation and release of antitiberculosis drugs [73]. Pandey et al. reported the formula- tion of three anti-TB drugs i.e. RIF, INH, and PYZ encapsulated in PLG nanoparticles. In M. tuberculosis infected mice, after oral administration of drug loaded nanoparticles at every 10th day, resulted in no detection of the tubercle bacilli in the tissues after 5 oral doses of treatment. These oral nano based drug delivery of anti-TB drugs has resulted in reduction of dosing frequency and better management of tuberculosis [20]. Prabakaran developed an osmotically regulated capsular multi-drug oral delivery sys- tem made of asymmetric membrane coating and dense semipermeable membrane coating-capsular systems for the controlled administration of RIF and INH for the treatment of TB [74]. Various anti-TB drugs have been formulated in dry microparticles for pulmonary delivery of drugs. These microparticles provided promising strategy in targeting those TB sites, which can be directly administered to the lungs with reduced sys- temic side effects. According to the experiment three formulations of PLG (each encap- sulating rifampicin) were developed i.e. non-porous (based on their drug release behaviour), hardened (based on the use of polyvinyl alcohol as a hardening agent) and porous. Out of three, the hardened PLG microcap- sules which showed 12–14% encapsulation of rifampicin and sustained drug release for 42 days in all the organs can be used as controlled drug delivery [75]. Nanoparticle with drug released in blood capillary Drug release on degradation Endolytic vesicle Lysozyme Blood capillary Release of nanoparticle encapsulated drug into the infected macrophage Figure 14.6 Release of nanoparticles encapsulated drug into infected alveolar macrophages of human.

500 Advanced Biomaterials and Biodevices Liposomes are widely used drug carriers for macrophage-specific antibacterial drug delivery (see Figure 14.6 above). In one of the experi- ment, streptomycin loaded liposomes were intravenously injected into the infected mice led to the decrease of the mycobacterium count in the spleen but not in the lungs but prolonged mouse survival and reduced drug toxicity was observed as compared to the free drugs [76]. According to the Klemens et al, gentamicin loaded liposomes were evaluated for the antibacterial activity in M. avium infected mouse model. It signifi- cantly reduced the bacterial count in spleen as well as liver compared to free drug [77]. Deol and Khuller developed stealth liposomes for the targeted delivery of anti-TB drugs to the lungs. Liposomes composed of phosphatidylcholine, cholesterol, dicetylphosphate-o-steroylamylopectin and monosialogangliosides / distearylphosphatidylethanolaminepoly (ethylene glycol) 2000. After intravenous administration in healthy and tuberculosis infected mice, increase in accumulation from 5.1% for con- ventional liposomes to 31% for Poly ethylene glycolated liposomal sys- tems after 30 min was observed. Drug uptake levels in the lungs increased to approximately 40% for the poly ethylene glycolated nanocarriers when administered to pretreated infected animals after 30 min 3[78]. Ahmed et al. developed various nanoemulsions of RIF (47 and 115nm) using GRAS listed excipients (US-FDA). As a result the entrapment effi- ciency was 99% with excellent stability over 3 months with slight increase in particle size and initial burst drug release of 40–70% after 2h [79]. Anisimova et al. encapsulated RIF, INH and streptomycin within poly-(n- butylcyanoacrylate) (PBCA) and poly-(isobutylcyanoacrylate) (PIBCA) nanoparticles and their accumulation in the human blood monocytes was tested in vitro [80]. Econazole and moxifloxacin loaded PLG nanopar- ticles were prepared by the multiple emulsion and solvent evaporation technique. As a result the drug levels in lungs, liver and spleen lasted till 6 days as compared to the pure drugs which were cleared within 12–24 h. In M. tuberculosis infected mice, only 8 doses of polymeric nanopar- ticles were sufficient to suppress bacterial clearance as compared to the pure drug which requires 56 doses daily of moxifloxacin and 112 doses of econazole twice a day. Further to improve treatment there was addition of third drug for tubercular chemotherapy [81]. Encapsulation of RIF, INH, and PYZ in alginate microspheres and oral administration to guinea pigs, maintain the drug concentration in plasma for 4–5 days and in the organs for 7–9 days. Weekly treatment of alginate microspheres resulted in the complete bacterial clearance in the organs of infected guinea pigs after 8 oral doses daily as compared to the admin- istration of free drugs [82]. Alginate-chitosan coated microcapsules were

Nanoparticles: Scope in Drug Delivery 501 developed as oral sustained delivery carriers for the anti-TB drugs. These microparticles were developed using ionotropic/external gelation method. Microparticles were formulated into three different forms containing rifampicin, isoniazide and pyrazinamide in the ratio of 1:2:2 (drug: sodium alginate: chitosan). These prepared microcapsules were then evaluated by SEM analysis, size analysis, spherecity, drug content, encapsulation effi- ciency, swelling studies and mucoadhesion which was then compared to the pure drug. In vitro release studies were carried out and the amount of drug released was analysed and was compared with the free drug [83]. Sung tested PA-824 (an alternative anti-TB candidate) which resulted in sustained release of the drug and maintained its level in the lungs for 32h [84]. PLG nanoparticles encapsulating anti-TB drugs such as PYZ, RIF, INH and ETB remained in the circulation for 72h. As a result, PLG encap- sulated INH was found to be higher than its MIC value (0.1mg/ml) [85]. A single subcutaneous dose of PLG encapsulated nanoparticles maintained drug level in the plasma, lungs, and spleen concentrations for more than 1 month and led to undetectable bacterial counts in the different organs [86]. Stearic acid encapsulating RIF, INH and PZA drugs nanoparticles after a single oral administration, the therapeutic concentration was maintained in the plasma for 8 days and in the organs for 10 days whereas free drugs get cleared within 1–2 days [87]. The nebulization of the drug loaded PLG nanoparticles with anti-TB drugs RIF, INH, and PZA was detected in plasma after 6h and therapeutic concentrations were detected until day 6 for RIF and day 8 for both INH, PZA. Nebulization of the nanoparticles to the M. tuberculosis-infected guinea pigs at every 10th day, no detection of the tubercle bacilli in the lungs was observed after only 5 doses of treatment, whereas 46 daily doses of orally administered drug required obtaining an equivalent therapeutic benefit [88]. In tracheal administration of Ofloxacin-loaded hyaluronan particles resulted in 50% lower serum bioavailability with respect to the intrave- nous or oral ofloxacin. This observation led to the conclusion that inhaled nanoparticles reduce systemic side effects, but it also suggested that extra pulmonary TB cannot be treated only by the inhaled therapies [89]. Saraogi et al prepared mannosylated gelatin nanoparticles for the selective delivery of INH to the alveolar macrophages and concluded that these nanopar- ticles can be a potential carrier for the safer and efficient management of TB through targeted drug delivery [90]. Ohasi et al designed RIF loaded biodegradable PLGA nanoparticles which were incorporated into the mannitol microspheres in single step by means of a four-fluid nozzle spray drier. As a result due to mannitol, the in

502 Advanced Biomaterials and Biodevices vivo uptake of the drug by alveolar macrophages in rat lungs was improved as compared to the RIF containing PLGA [91]. Chitosan has also been used for the development of effective drug deliv- ery systems because of its unique physiochemical and biological proper- ties. The primary hydroxyl and amine groups located on the backbone control its physical properties. The small size of chitosan nanoparticles favours intravenous administration of drugs to target sites [92]. They are widely used as drug delivery systems for low molecular drugs, peptides and genes [93, 94]. Machida et al evaluated the potential of lactosaminated N-succinyl- chitosan (Lac-Suc) synthesized by the reductive amination between the N-succinyl-chitosan and lactose in the presence of sodium cyanoborohy- dride, as a liver specific drug carrier [95]. Recently, Liu et al prepared the polyion complex micelles (PIC micelles) based on methoxypoly-(ethylene glycol) (PEG)-graft-chitosan and lactose-conjugated PEG-graft-chitosan for targeted delivery of diammonium glycyrrhizinate (DG) to liver [96]. Chitosan nanoparticles with PEG have gain attention because of its potential in the therapeutic applications [97]. PEGylation mainly increases physical stability and prolongs their circulation time in blood. Recently, the effect of PEG conjugation on PTX loaded N-octyl-sulfate chitosan nanoparticles were investigated by Qui et al. They found that these con- jugated particles were less phagocytised as compared to the unconjugated nanoparticles by the reticuloendothelial system. They have also been investigated as carriers for the delivery of different types of small molecules drugs such as paciltaxel, camptothecin, methotrexate and all trans-retionic acid (ATRA) [98–102]. Dendrimers are well defined, regularly hyper branched and 3D archi- tecture having relatively low molecular weight, polydispersity and high adjustable functionality. The unique structure is responsible for the encap- sulation and delivery of anti-TB agents. Kumar et al developed manno- sylated polypropylimine dendrimeric nanocarriers for the delivery of RIF to macrophages. The RIF loaded dendrimers in alveolar macrophages in lungs of rat showed an increase in the intracellular concentration of the antibiotic [103]. The therapeutic drugs with polymeric nanoparticles and solid lipid nanoparticles, results in more sustained drug release and the ability to tar- get specific cells and organs. Delivery of the drugs to the lungs has to face many challenges such as formulation instability due to particle-particle interactions and poor delivery efficiency due to the exhalation of low iner- tia nanoparticles. Concerning all these problems led to the invention of novel methods of formulating nanoparticles into the form of micron scale

Nanoparticles: Scope in Drug Delivery 503 dry powders. Through these nanoparticles the lungs can be targeted for the drug delivery to specific lung cells such as alveolar macrophages for the treatment of tuberculosis [104]. Significant advances in medical aerosol development began in the 1950s with a focus on the delivering of the asthma drugs directly to the lungs, the target organ, thereby resulting in reduction in the amount of systemic drug and their adverse effects [105]. Since mid-1800s nebulizers have been used for delivering solutions of the drugs to target site, by generation of aerosol droplets from liquid with the help methods such as ultrasonic or air jet technology. Recently, nanoscale aerosol vaccines have also been developed which perfuse more throughout the respiratory tract and also increases the amount of drug reaching to the target alveoli. Garcia ontreas et al. have recently synthesized a particle system both micrometer and nanometer dimensions for aerosolized delivery of the attenuated tuberculosis vaccine, BCG. The aerosol delivery of BCG encapsulated nanomicroparticles in guinea pigs increased their resistance to tuberculosis infection, and gen- eration of better immune protection than a standard parenteral BCG for- mulation [106]. The nanoparticulate delivery of aerosolized IFN-gamma through the pulmonary route has been more efficient new adjunct treat- ment for tuberculosis [107]. Para-aminosalicyclic acid (PAS) is a tuberculostatic agent recently being formulated into large porous particles for direct delivery into the lungs via inhalation. These particles possess some optimized physical properties for deposition through respiratory tract; the drug was loaded with 95% by weight over 4 weeks at elevated temperatures. PAS concentrations were measured in the plasma, lung lining fluid and homogenized whole lung tissue. As a result the PAS get cleared within 3h from the lung lining fluid and plasma. The above experiment led to the conclusion that the inhala- tion delivery of PAS help in the reduction of total dose delivered [108]. Dry powder inhalation systems have also come into existence with a potential of storing the drug in a dry state, which confers long term stability and sterility. The first type of dry powder inhalation system utilized the patients breathing system for dispersing and delivering of the milled micron sized particles. The large geometric sized particles improved the dispersion properties and higher lung deposition of the delivered dose (upto 59%) [109]. Vyas formulated aerosolised liposomes with the incorporation of RIF through a cast-film method. The liposomes coated with the alveolar macrophage-specific ligands resulted in the more accumulation in alveolar macrophages, thus maintaining high concentrations of RIF in the lungs, even after 24h [110].

504 Advanced Biomaterials and Biodevices Recently, encapsulation, characterization and in vitro release of anti-TB drug (RIF) using chitosan-polyethylene glycol nanoparticles were devel- oped. Chitosan polyethylene glycol 600 (PEG) nanoparticles were pre- pared by ionic gelation technology. The preparation of these nanoparticles is based on the interaction between positively charged chitosan solution and negatively charged TPP solution. When PEG binds with chitosan- rifampicin it changes the character and the surface of the nanoparticles and slightly increases its particle size, as well as encapsulation of drugs also increased. PEG bind with CS-RIF has resulted in more prolonged retention as compared with non-coated CS-RIF. Various parameters and method- ologies such as loading capacity, encapsulation efficiency, SEM (Scanning electron microscopy), FITR (Fourier transmission infra-red microscopy) and in vitro release of drugs have been utilised for characterization of nanoparticles [111]. Recently, there was targeted intracellular delivery of anti-TB to Mycobacterium infected macrophages through functionalized mesopo- rous silica nanoparticles. Mesoporous silica nanoparticle (MSNP) drug delivery system was coated with polyethyleneimine (PEI) to release rifam- picin. These mesoporous silica nanoparticles get internalized by human macrophages and delivered to the lysosomes and to the acidified endo- somes as a result intracellular release of high concentrations of antituber- culosis drugs occurred. Coated MSNP have shown much greater loading capacity than uncoated MSNP. The amount of RIF loaded on the PEI coated nanoparticles was determined after elution by spectrophotometer at the wavelength 475nm against RIF standards [112]. Microspheres are spheri- cal, free flowing particles ranging in average particle size 1–50 microns. Currently, the potential of microspheres as carriers for target drug delivery systems has been exploring. Microspheres are prepared from the different methods such as protein gelation technique, sonication technique, solvent evaporation technique, spray and freeze technique, polymerization tech- nique and solvent extraction method [113]. Oral nanoparticles are also being used for the delivery of the antitu- berculosis drugs. In an experiment, orally administered poly-lactide- co-glycolide (PLG), a synthetic polymer nanoparticle encapsulating antituberculosis drugs such as RIF, INH and PZA was developed for cere- bral drug delivery in murine model. These nanoparticles were prepared by multiple emulsion and solvent evaporation technique and then they were administered orally to mice for their biodistribution, pharmokinetic and chemotherapeutic studies. As a result of this experiment, a single oral dose to mice maintains sustained drug levels for 5–8 days in the plasma and for 9 days in the brain. As well as in M. tuberculosis H37Rv infected mice, five

Nanoparticles: Scope in Drug Delivery 505 oral doses of these nanoparticles formulation was administered every 10th day which in turn resulted in absence of tubercle bacilli in the meninges, on the basis of CFU and histopathology [114]. Niosomes can also be used as an alternative to liposomes for specific tar- get drug delivery. They are basically thermodynamically stable liposomes like vesicles produced by the hydration of cholesterol and charge induc- ing components such as charged phospholipids (e.g. Dicetylphosphate and stearyl amine) and non ionic surfactants (e.g. monoalkyl or dialkyl polyoxyethylene ether) [115]. Micron sized RIF loaded niosomes contain- ing Span-85 as surfactants were prepared by the Jain and Vyas [116]. As a result upto 44% of the drug was localized in the lungs by adjusting the size of the carrier. The same group of the scientist further extended their studies to investigate the biodistribution of niosomes of smaller sizes with the different sorbitan esters (Span 20, 40, 60, 80 and 85) and cholesterol in 50:50 mol fraction ratios [117]. In vitro studies showed 80% maximal and 52% minimal levels for Span-20 and Span-85 based systems. All these studies led to the conclusion that more lipophilic the surfactant, slower was the drug release in the aqueous medium. Nanoparticles of size 250 nm were used for delivery of the anti-TB drugs such as isoniazid, rifampicin and streptomycin. The accumulation of these drugs was checked in human monocytes as well as thieir antimi- crobial activity against M. tuberculosis residing in the human-monocyte derived macrophages. The result was that the intracellular concentration of the free INH was equal to or slightly higher than that of the extracellular fluid [118]. 14.4 Cancer & Tumor Targeting Nanoparticles Nanoparticles are very well suited materials for targeted tumor delivery because of their ability to circulate in the bloodstream for relatively longer period of time as well as their ability to accumulate in the tumor spaces. Some of the in vivo and in vitro experiments have shown nanoparticles to be fruitful for the tumor treatment. There is a growing evidence which sug- gests that many nanoparticles accumulated at the tumor site is indepen- dent of the presence or absence of the targeting ligand. In many studies, nanodelivery systems with target ligands have shown better performance than non-targeted system (Figure 14.7, below). These entire conclusions led to the improvement in the performance of the nanoparticles asso- ciation with target cell membranes and target cell internalization. EPR (enhanced permeability and retention) effect is one of the dominating

506 Advanced Biomaterials and Biodevices Receptor Free drugs Nanocarrier Through Blood capillary enhanced permeation effect Ligand Cancer cell Increase in intracellular drug concentration Nanoparticles targeting Cancer cells Figure 14.7 Nanoparticles target cancer cells through ligand receptor binding. mechanisms for localization of tumour site [119]. Active tumour targeting involved the use of molecules which specifically interact with the physi- ological target whereas the passive targeting involves the use of natural properties and processes in the tissues for localization of the delivery agent at a desired target site. Passive tumour targeting is mainly by EPR effect. Particles within the range of 100–200 nm have been shown to accumulate in tumour by EPR effect. Wang et al studied the biodistribution of targeted nanoparticles com- posed of heparin-folate-paclitaxel conjugates loaded with paclitaxel and compared it with the non-targeted nanoparticles of heparin-paclitaxel loaded with paclitaxel, as a result, this targeted system loaded with the drug reduces the tumour volume over nanoparticle and paclitaxel con- trol in the KB-3–1 human nasopharyngeal carcinoma-xenograft bearing mouse model [120]. Another most common targeting ligand such as trans- ferrin (Tf) has been conjugated to a variety of targeted delivery systems for targeting over expressive Tf receptors which are common in many can- cerous cells. They basically improve the drug delivery [121]. Liposomes possessing an anti-HER2 monoclonal antibody (Mab) composed of phos- phatidylcholine and polyethylene glycol (PEG) modified distearoyl phos- phatidylethanolamine with an average particle diameter of 100 nm has resulted in improve antitumor efficacy of doxorubicin over the various control formulations [122]. Kirpotin et al. further evaluated the biodistri- bution and uptake of these targeted liposomes containing the anti-HER2 Mab and compared it with that of non -targeted liposomes [123]. In 1975,

Nanoparticles: Scope in Drug Delivery 507 Ringsdorf first proposed the concept of polymer-drug conjugates for deliv- ery of the hydrophobic smaller drugs to their target site of action [124]. This polymer-drug conjugates composed of a water-soluble polymer and is chemically conjugated to the drug via a biodegradable spacer. This spacer basically cleaved at the target site by hydrolysis or enzymatic degradation. They can be accumulated at the tumour site by the EPR effects, followed by release of the drug through the cleavage of spacer [125]. In recent years, chitosan anticancer drug conjugates have also been investigated. Doxorubicin conjugated glycol chitosan (DOX-GC) with cis-aconityl spacer was synthesized by the chemical attachment of N-cis- aconityl DOX to GC using carbodiimide. DOX-GC conjugates contain- ing 2–5wt% DOX form self-assembled nanoparticles in aqueous condition but if DOX is present in higher concentration i.e. above 5.5wt% in the nanoparticles, it will get precipitated due to increase in hydrophobicity. The loading contents of DOX in the nanoparticles increased upto 38.9wt%. The release rate of DOX from the nanoparticles is dependent on the pH of the media because the cis-aconityl spacer is cleavable at low pH [126]. When these DOX-GC conjugated nanoparticles were administered into the mice, they get preferentially accumulated in the tumour tissue thus describing EPR effect. A variety of hydrophobic drugs can be loaded into the chitosan nanoparticles; the loading efficiency depends on the physio- chemical characteristics and preparation methods used. For the cancer therapy, a hydrophilic 5-fluorouracil was loaded into the chitosan nanoparticles (250–300 nm in diameter) using water- in-oil emulsion method, followed by the chemical cross linking of the chitosan in the presence of glutaraldehydes [127]. Some of the drug loaded solid nanoparticles could release the drugs at the rate which is in turn dependent on the type of hydrophobic moiety, degree of substitution and the physio- chemical properties of the drugs. These chitosan based nanoparticles have been reported to be selectively accumulated at the tumour site, primarily owing to the EPR effect. As a result these drug loaded nanoparticles have shown better therapeutic efficacy than the free drug in vivo. Magnetic targeting is also an attractive physical targeting technique, generating a substantial attention for the drug delivery applications. The therapeutic agents to be delivered are either immobilized on the surface or they are encapsulated into the magnetic micro or nanoparticle carriers. Upon intravenous administration, they get concentrated at the tumour site using an external high-gradient magnetic field [128]. After the accumula- tion of the magnetic carrier at the target tumour site in vivo, drugs are released from the magnetic carrier and then effectively taken up by the tumour cells. The efficiency of the carrier accumulation mainly depends

508 Advanced Biomaterials and Biodevices on various parameters such as intensity of the magnetic field, rate of blood flow and the surface characteristics of carriers. Gallo et al developed magnetic chitosan microspheres containing oxantrazole (MCM-OX), an anti cancer drug, for the treatment of brain tumour [129]. He monitored the level of OX in the brain after administer- ing intra-arterial injections of MCM-OX to male Fischer 344 rats under the magnetic field of 6000G for 30 min. As a result, a 100 fold increase in OX concentrations in the brain after administration of MCM-OX was observed as compared to the OX in solution. In a similar study, Chen et al prepared chitosan-bound magnetic nanoparticles loaded with epirubi- cin, an anthracycline drug used for cancer chemotherapy. The magnetic nanoparticles were stable at pH 3–7 and approximately 80% of the drug was released after 150–300 min in a biological buffer. In vitro experiment showed the efficiency of anticancer property of drug loaded nanoparticles as compared with that of free drugs [130]. Magnetic field enhances the cellular efficiency uptake of the mesopo- rous silica nanoparticles. The internalization of M-MSNs (magnetic mes- oporous nanoparticles) by A549 cancer cells could be enhanced by the magnetic field. The endocytosis studies indicate that the M-MSNs inter- nalization by A549 cells is mainly energy dependent pathway, namely clathrin-induced endocytosis. With the help of magnetic field, anticancer drug loaded M-MSNs induced cancer cell growth inhibition. Delivery of hydrophilic and hydrophobic drugs through magnetic mesoporus silica nanoparticles (MMSNs) inhibits cancer cells growth [131]. Recently, a report on the development of drug delivery system for pho- tosensitive delivery of an anticancer drug campothtecin along with cyto- toxic cadmium sulphide from a magnetic drug nanocarrier was developed. During this experiment, core–shell nanoparticles consisting of magnetic iron-oxide-cores and mesoporous silica shells were synthesized with a high surface area (859 m2 g−1) and hexagonal packing of mesopores (2.6 nm) in diameter. The mesopores were loaded with an anticancer drug camp- tothecin and the entrances of the mesopores were blocked with 2-nitro- 5-mercaptobenzyl alcohol functionalized CdS nanoparticles through a photo cleavable carbamate linkage. Camptothecin release from the mag- netic delivery system was measured by the fluorescence spectroscopy upon irradiation by the UV light. As a result, treatment of cancer cells with these drugs lead to the decrease in viability of the cells because of the activity of capping of CdS nanoparticles. The capping of Cds nanoparticles and loaded camptothecin exert an anticancer activity [132]. Transferrin conjugated paclitaxel biodegradable nanoparticle were also tested in the murine model for treatment of prostate cancer. It was

Nanoparticles: Scope in Drug Delivery 509 hypothesized that nanoparticle conjugated to transferrin ligand enhances the therapeutic efficiency of the drug. Nanoparticles of 220 nm diameter loaded with 5.4% w/w paclitaxel drug under in vitro condition exhibited sustained release of 60% encapsulated drug in 60 days. The anti-prolifera- tive activity of NPs was then determined in human prostate cancer cell line (PC3) and their effect on tumour inhibition was observed in the murine model of prostate cancer. Animals which received a single-dose of intra- tumoral injection of transferrin conjugated nanoparticles resulted in com- plete tumor regression and greater survival rate [133]. Lymphatic drainage plays an important role in uptake of particulates in respiratory system and also being associated with the spreading of lung cancer through metastasis development. Recently, solid lipid nanoparticles (SLN) have used as carriers of anti-tumoral drugs. Nanoparticles of about 200 nm diameter, radiolabelled with 99m Tc using the lipophilic chelator D,L-hexamethylpropyleneamine oxime (HMPAO) were developed for the pulmonary uptake. The biodistribution studies were also carried out following aerosolisation and administration of a 99mTc-HMPAO-SLN suspension to a group of adult male Wistar rats. As a result, 60 min dynamic image followed by the static image were col- lected at 30 min intervals for up to 4h post inhalation which has shown the significant uptake of the radiolabelled SLN into the lymphatic system after inhalation and thus controlling spreading of lung cancer [134]. Recently, nanoparticle-aptamer bioconjugate, a new approach for treat- ment of prostate cancer was developed. In this nucleic acid ligands (aptam- ers) are well suited for therapeutic targeting of the encapsulated drugs and controlled release of polymer particles in cells or tissues. A bioconjugate was synthesized, mainly composed of controlled release polymer nanopar- ticles and aptamers. Its efficacy was examined for targeted delivery to pros- tate  cancer  cells. Nanoparticles composed of poly-(lactic acid) blocked polyethylene glycol (PEG) copolymer with a terminal carboxylic acid func- tional group (PLA-PEG-COOH) and encapsulated rhodamine-labeled dextran (as a model  drug) within PLA-PEG-COOH was synthesized. Nanoparticle-aptamer bioconjugates with RNA aptamers which bind to the prostate specific membrane antigen overexpressed on prostate acinar epi- thelial cells. As a result, these bioconjugates can be efficiently targeted and taken up by the prostate LNCaP epithelial cells. This was the first report on targeted drug delivery through nanoparticle aptamer bioconjugates [135]. Recently, Heparin is used as a carrier for cancer targeting and imaging. Heparin-anticancer drug conjugates shows higher anticancer activity than free drug. The conjugated heparin (heparin-deoxycholate sodium) retained its ability to bind with angiogenic factors, thus resulting in a

510 Advanced Biomaterials and Biodevices significant decrease in endothelial tubular formation. Secondly, targeting ligands conjugated with that of heparin derivatives have been used for the receptor mediated delivery of anticancer drug. Heparin-folic acid- retinoic acid (HFR) bioconjugates are being used for treatment of cancer cells [136]. Recently, a pH responsive drug carrier is developed, based on chondrio- tin sulfate functionalized mesostructured silica nanoparticles (NMChS- MSNs) i.e., the amidation between NMChS macromer and amino group functionalized MSNs. These nanoparticles were spherical in shape with a mean diameter of about 74nm. A well known anti-cancer drug, Doxorubicin hydrochloride (DOX) was loaded into the channels of NMChS-MSNs through the electrostatic interactions between drug and matrix. As a result, drug release rate was pH dependent and it increases with the decrease in pH. In vitro cytotoxicity test also proves that these nanoparticles are highly biocompatible and can be used as a drug carrier. The above experiment proves that these chondriotin sulfate mesostructured functionalized silica nanoparticles are good platforms for pH dependent controlled drug deliv- ery systems for cancer therapy [137]. Pancreatic cancer is a highly lethal disease with a 5-year survival rate less than 5% due to the lack of an early diagnosis method and effective therapy. Recently, a multifunctional nanoimmunoliposome with high loading of ultra small super paramagnetic iron oxides (USPIOs) and doxorubicin (DOX) was prepared by transient binding and reverse-phase evaporation method and then conjugated with anti-mesothelin mono- clonal antibody by post-insertion method to target anti-mesothelin- overexpressed pancreatic cancer cells. In vivo studies have shown that, comparing with FD (free DOX) and PLDU, M-PLDU possessed higher inhibitory effect on tumour growth and the tissue distribution assay fur- ther proved that M-PLDUs get selectively accumulated in the tumour xenograft [138]. A novel magnetic nanoparticle for drug carrier has been recently dis- covered for the enhanced cancer chemotherapy. Magnetic nanoparti- cles loaded with antitumor drugs in presence of external magnetic field resulted in improvement in cancer treatment. In this experiment, DOX- PGMNPs nanoparticles were synthesized and cytotoxicity was assessed in vitro. Along with this, intravenous administration of DOX-PGMNPs to H22 hepatoma cell tumour bearing mice, biodistribution of DOX was also measured in different tissues [139]. One of the nanoparticle poly-(D,L-lactide-co-glycolide)/montmo- rillonite were decorated by Trastuzumab, an epidermal growth fac- tor receptor-2(HER-2) antibody and was being used for targeted breast

Nanoparticles: Scope in Drug Delivery 511 cancer chemotherapy with paclitaxel as a model anticancer drug. These nanoparticles were prepared by solvent extraction/evaporation method. Internalization of the coumarin-6-loaded PLGA/MMT NPs with or with- out the antibody decoration by both of Caco-2 colon adeno carcinoma cells and SK-BR-3 breast cancer cells were visualized by confocal laser scanning microscopy and quantitatively analyzed, which shows that the antibody decoration achieved significantly higher cellular uptake of the NPs to treat breast cancer [140]. Development of functionalized super paramagnetic iron oxide nanopar- ticles were being reported to interact with human cancer cells. The capacity of interaction of the cells with these nanoparticles as well as cytotoxicity was evaluated in human melanoma cells. Out of the four formulations of nanoparticles, only the polyvinyl alcohol super magnetic iron oxide nanoparticles interact with cells and cytotoxicity was negative in human melanoma cells [141]. 14.5 Conclusion Nanotechnology is globally emerging technology for targeted drug deliv- ery. Due to some disadvantages related to the unconventional strategies, such as larger amount of drug consumption by patients, more time in recovering lead to the more use of different kinds of nanoparticles as drug carriers. They basically act by crossing the biological barriers and target- ing the site. Nanoparticles such as solid lipid nanoparticles, polymeric nanoparticles, liposomes, mesoporous silica nanoparticles are being used for treatment of various types of diseases. Several carriers based drug deliv- ery system incorporating anti-TB drugs have been developed for target site action. Nanoparticles encapsulating anticancer drugs such as doxorubicin have been also developed for treatment of various types of cancers such as prostate cancer, breast cancer, pancreatic cancer and lung cancer. All these developments in drug delivery reports have resulted in significant merits such as improved drug bioavailability, reducing dosing frequency, versatil- ity of routes of administration and long term stability which become the basis of better management of diseases. Recently, aerosol vaccines are being in progress for drug delivery. Besides having so many advantages, some of the toxicological issues related to understanding the fate of nanocarri- ers, polymeric constituents in the body as well as elimination of residual should be deleted. harmful residual organic solvents have to be resolved. Future research on vectorized delivery system of drug has been focussed for large amount of drug delivery and for better results.

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15 Smart Polypeptide Nanocarriers for Malignancy Therapeutics Jianxun Ding, Di Li, Xiuli Zhuang and Xuesi Chen* Key Laboratory of Polymer Ecomaterials, Changchun Institute of Applied Chemistry, Chinese Academy of Sciences, Changchun, People’s Republic of China Abstract Nowadays, malignancy (i.e., cancer) has been one of the major global causes of morbidity and mortality. In the past few decades, the attractive potentiality of anticancer nanomedicines has led to explosive research of polymeric nanocar- riers, which exhibit targeting, modifiable and controlled drug delivery capabili- ties. Based on the previous works and in view of the physiological repellents of drug delivery in vivo, some fruitful efforts have recently focused on the design and investigation of smart drug delivery systems with stimuli-responsive polypeptides as matrices. Polypeptides, consisting of repeated amino acid units, are unique biocompatible and biodegradable synthetic polymers with mimicking structures of natural proteins. In addition, various merits, such as their dexterous stimuli- responsiveness to external environmental changes, make polypeptides become promising materials for intelligent antitumor drug delivery. In this chapter, we have reviewed the recent advances in stimuli-responsive polypeptide nanocarri- ers, especially pH and/or reduction-responsive ones, for malignancy therapeutics. Keywords: Malignancy therapies, nanocarriers, polypeptides, stimuli-responsive 15.1 Introduction Malignancy (also called cancer) has been one of the most serious world- wide diseases that threaten human health and even life [1–3]. Despite the remarkable developments of various antitumor drugs including *Corresponding author: [email protected] Ashutosh Tiwari and Anis N. Nordin (eds.) Advanced Biomaterials and Biodevices, (523–546) 2014 © Scrivener Publishing LLC 523

524 Advanced Biomaterials and Biodevices doxorubicin (DOX), paclitaxel, chlorambucil, platinum drugs, et al. [4–7], the clinical effects are disappointing because of some serious impediments, such as severe side effects to normal tissues and low efficacy towards mul- tidrug-resistant tumor cells. In order to overcome the aforementioned adverse reactions, various nanocarriers, that is, micelles, vesicles, nanogels and so on, have been explosively exploited as nanovehicles for antitumor drug delivery [8–12]. Among all of the aforementioned nanovehicles, the smart ones that sharply respond to the external environmental stimuli including pH, reduction, temperature, enzyme and light, as well as electromagnetic field, have received extensive attention for their unique advantages in the field of controlled transportation of antitumor drugs [13–16]. The “on- demand” release of payloads from the intelligent delivery systems can be triggered by the stimuli of microenvironments in the pathological sites, resulting in an enhanced inhibition efficacy to tumor cellular proliferation with reduced side effects [17, 18]. Given the specific chemical microenvi- ronments in tumor tissues and cells, the nanocarriers responding to pH and/or reduction exhibit great potential in smart antitumor drug delivery. In detail, the pH values in late endosome (pH ~ 5.0–6.0) and lysosome (pH ~ 4.5–5.0) (Table 15.1) are remarkably lower than the extracellular pH (pH ~ 6.5–7.4) [19–22]. Additionally, the reduction-responsive link- ers (e.g., disulfide and diselenide bonds) are stable at normal physiological condition, but degrade to free thiols or selenols in the presence of gluta- thione (GSH) or reductases [2, 3, 23]. The intracellular concentration of GSH (~ 0.5–10.0 mM) is significantly higher than that of outside cells Table 15.1 The pH values in normal and tumor tissues and cellular compartments [19–22]. Tissue or cellular compartment pH Normal tissue 7.4 Tumor tissue 6.5 – 7.2 Circulatory system 7.35 – 7.45 Early endosome 6.0 – 6.5 Late endosome 5.0 – 6.0 Lysosome 4.5 – 5.0

Smart Polypeptide Nanocarriers for Malignancy Therapeutics 525 (~ 2.0–40.0 μM) [2, 3]. Furthermore, the GSH concentration in some can- cer cells (e.g., A549 cells, a human lung adenocarcinoma cell line) has been reported to be several times higher than that in normal cells [3]. Apart from the stimuli-responsive properties, excellent biocompatibility and appropriate biodegradability are the other requirements for ideal nano- carriers in practical applications [24]. Polypeptides, which are poly(amino acid)s linked by amide bonds, are some of the most potentially biocompat- ible and biodegradable synthetic polymers with mimicking structures of natural proteins, and have been widely used in various biomedical realms including drug and gene delivery, biosensors and diagnostics, etc. [25, 26]. Compared with the conventional biomaterials, polypeptides may form stable secondary structures, such as α-helix and β-sheet, attributed to the cooperative hydrogen-bonding [27]. Furthermore, the self-assemblies of polypeptides exhibit smart response to different external environmental stimuli, especially pH and reduction [17, 28]. Therefore, the smart poly- peptide nanocarriers have attracted increasing attention for their great potential in controlled antitumor drug delivery. This chapter focuses on the recent progress in stimuli-responsive polypeptide nanovehicles, par- ticularly pH and/or reduction-responsive ones, which have been exploited as nanocarriers for antitumor drugs. 15.2 Smart Polypeptide Nanovehicles for Antitumor Drug Delivery 15.2.1 Polypeptide Micelles The polymeric micelles are nanosized colloids that are spontaneously self- assembled from amphiphilic copolymers in aqueous environment driven by microphase separation [29–32]. The micelles exhibit a hydrophobic core commonly composed of hydrophobic segments (e.g., polypeptides), which are shielded by a nonfouling hydrophilic shell (e.g., poly(ethylene glycol) (PEG)) [33–36]. The hydrophobic cores serve as the sustained release reservoirs of bioactive molecules, especially antitumor drugs, while the hydrophilic shells improve the stabilities and compatibilities of micelles in sophisticated in vivo circulatory systems [33, 35, 36]. In addition, the appropriate nanoscales of micelles (~10–200 nm) endow them with posi- tive targeting function ascribed to their enhanced permeation and reten- tion effect [17, 24, 37, 38]. As a result, polymeric micelles have attracted significant attention as drug delivery platforms for malignancy thera- peutics. Among them, the micelles originating from stimuli-responsive

526 Advanced Biomaterials and Biodevices polypeptides are some of the most promising members benefited from their excellent biocompatibility and biodegradability, regular secondary confirmations, and stimuli-responsive assembly and disassembly proper- ties [21, 35, 36, 39–41]. Herein, the pH and/or reduction-responsive poly- peptide micelles as smart antitumor drug carriers are summarized. 15.2.1.1 pH-Responsive Polypeptide Micelles Due to their tumor tissular (pH ~ 6.5–7.2) and intracellular acidic microen- vironments (pH ~ 4.5–6.5) (Table 15.1), polypeptide micelles consisting of pH-responsive polypeptide segments or acid-labile polypeptide-drug con- jugates show great potential in the field of intelligent antitumor drug deliv- ery [22]. Chen and coworkers have exploited methoxy poly(ethylene glycol)-block-poly(L-glutamic acid) (mPEG-b-PLGA) as pH-responsive carriers of cis-diamminedichloroplatinum (cisplatin, CDDP) or DOX for non-small cell lung cancer (NSCLC) treatment (Figure 15.1) [41, 42]. The A Nucleus DOX-NH3+ DOX-NH3+ mPEG-b-PLGA Self-assembly Endosome Ionic-complex formation Endocytosis DOX 100 Accumulative DOX Release (%)a C 1400 PBS Tumor Volume (mm3)b 1200 DOX (2 mg kg–1) B c 1000 DOX (4 mg kg–1) mPEG-b-PLGA-DOX (2 mg kg–1) 80 d 800 mPEG-b-PLGA-DOX (4 mg kg–1) 60 40 600 400 20 200 0 0 0 10 20 30 40 50 60 0 5 10 15 20 Time (h) Time (Day) Figure 15.1 (A) Schematic illustration of drug loading, endocytosis and intracellular drug release of the pH-responsive amphiphilic complex based on mPEG-b-PLGA and DOX. (B) Time and pH-dependent DOX release profiles of the DOX-loaded micellar nanoparticles in (a) phosphate-buffered saline (PBS) at pH 5.5 with 10% (v/v) fetal bovine serum (FBS), (b) PBS at pH 5.5, (c) PBS at pH 7.4 with 10% FBS and (d) PBS at pH 7.4. (C) In vivo antitumor efficacy of various DOX formulations in the A549 tumor-bearing mouse model. ***p < 0.001 versus PBS group. Reproduced with permission from [41].

Smart Polypeptide Nanocarriers for Malignancy Therapeutics 527 CDDP or DOX was combined with LGA unit through respective chelation or electrostatic attraction, and its release from the drug-loaded micellar nanoparticle was accelerated at the acidic pH, mimicking the endosomal/ lysosomal microenvironment (Figure 15.1B). In an in vivo assay towards A549-bearing female BALB/c nude mice, the CDDP-loaded mPEG-b- PLG micelle showed approximative antitumor efficacy but significantly lower body weight loss in comparison with free CDDP [42]. Preferably, mPEG-b-PLGA–DOX micellar nanoparticle exhibited both low toxicity and higher antitumor efficacy against human NSCLC xenografted tumor model (Figure 15.1C) [41]. Additionally, Chen’s group also developed the graft and block copolymers consisting of PLGA as pH-responsive DOX carriers for their therapeutic effect against hepatic carcinoma [21, 43]. DOX was loaded into micelles through nanoprecipitation, and in vitro DOX release was accelerated as either the length of PLGA segment or pH decreased. Remarkably, the galactosylated DOX-loaded block copoly- mer exhibited improved antitumor efficacy against HepG2 cells (a human hepatoma cell line); conversely, the OEGylated one in vitro and in vivo exhibited great potential as a highly specific drug delivery platform for enhanced chemotherapy efficacy in human hepatoma [21]. Bae and coworkers have developed the DOX-loaded poly(ethylene glycol)-block-poly(L-histidine) (PEG-b-PLH) micellar platforms [44, 45]. The in vitro internalization of DOX toward A2780 cells (a human ovarian cancer cell line) cultured with the DOX-loaded micelle was improved more than five times by decreasing the pH from 7.4 to 6.8, which resulted in an enhanced in vitro cellular antiproliferative activity at pH 6.8 [45]. After being administrated to A2780 xenografted nude mice, the pH-responsive DOX-loaded micelle showed significantly enhanced antitumor efficacy and markedly increased half-life time com- pared to free DOX. In addition, various antitumor drugs (i.e., DOX) have been covalently conjugated in the polypeptides via acid-labile linkers to fabricate pH- responsive micelle-like drug delivery systems. Kataoka and coworkers have developed a PEG–poly(L-aspartate-hydrozone-doxorubicin) (PEG- b-P(LA-H-DOX)) diblock copolymer, in which DOX was conjugated to the LA unit via an acid-cleavable hydrazone bond with a substitution ratio of 67.6 mol% [46, 47]. The copolymer formed into micelle (~ 65 nm) with the DOX-conjugated polypeptide segment as a core and PEG block as a shell. DOX was stably conjugated within the micelles at a pH more than 6.5, whereas the release of DOX was triggered at acidic pH. The DOX- conjugated micelle was trapped in endocytic compartments, and exhib- ited enhanced in vitro and in vivo antitumor efficacy toward SBC-3 cells (a

528 Advanced Biomaterials and Biodevices human small-cell lung cancer cell line) and lower in vivo toxicity compared to free DOX [46, 47]. Gu’s group also conjugated DOX to the dendritic PLGA containing biotin via hydrazone bonds [48]. The dendritic PLGA–DOX conjugate self-assembled into spherical micelle with an average diameter of ~ 50 nm. The DOX-conjugated micelle showed dramatically accelerated DOX release at pH 5.0, and enhanced intracellular uptake and cellular prolifera- tion inhibition toward HeLa cells (a human cervical carcinoma cell line). Zhuang et al. have prepared the micelle based on tumor-acidity-sensitive PEG–poly(L-lysine) (PLL)–DOX conjugate with acid-labile amide bond for the proliferation inhibition of HeLa cells with similar activity of free DOX [49]. 15.2.1.2 Reduction-Responsive Polypeptide Micelles The polypeptide micelles containing reduction-responsive bonds (e.g., disulfide linker) also exhibited great potential in targeting intracellular antitumor drug delivery, benefiting from the dramatic variation of GSH concentration outside and inside the cells [2]. In 2011, Li and coworkers exploited two kinds of reduction-responsive micelles originating from the disulfide-linked PEG–polypeptide copolymers (i.e., poly(ε-benzyloxy- carbonyl-L-lysine) (PZLL) and poly(rac-leucine)) to achieve intracellular delivery of DOX [50, 51]. The micelles would undergo a fast dissociation of PEG shelter when these were incubated in a reductive environment (10.0 mM DL-dithiothreitol (DTT) or GSH. The reduction-responsive micelles had good biocompatibility and high drug loading efficiency (DLE) for DOX. The DOX-loaded micelles exhibited enhanced intra- cellular DOX release and antiproliferative activities toward MCF-7 (a human breast cancer cell line) or HepG2 cells with elevated intracellular GSH concentration. In Chen’s group, the mPEG–PZLL diblock copolymers were also syn- thesized with a more facile approach [2]. The copolymers self-assembled into micelles in PBS at pH 7.4 through both direct dissolution and dialysis methods. As shown in Figure 15.2A, DOX was loaded into the micelles through nanoprecipitation with a DLE of about 30 wt%. The in vitro DOX release from all DOX-loaded micelles was accelerated in PBS with 10.0 mM GSH, mimicking intracellular reductive conditions (Figure 15.2B). The DOX-loaded micelles showed improved cellular internalizations (Figure 15.2C) and higher proliferation inhibitions towards glutathione monoester (GSH-OEt) pretreated HeLa and HepG2 cells as compared to unpretreated or buthionine sulfoximine (BSO) pretreated cells.

Smart Polypeptide Nanocarriers for Malignancy Therapeutics 529 A Self-assembly PBS, pH 7.4 Endocytosis DOX-loaded mPEG-S2-PZLL micelle GSH B a Intracellular DOX release with a b reductive cleavage of disulfide linkage c 90 c d C b 60 100 30 50 Cumulative DOX Release (%) Counts 0 0 20 40 60 010–1 100 101 Time (h) Flurescence Intensity Figure 15.2 (A) Schematic illustration of DOX loading and targeting intracellular release. (B) DOX release from (a and b) The DOX-loaded mPEG113-S2-PZLL18 micelles and (c and d) mPEG113-S2-PZLL35 micelles at pH 7.4 (a and c) without GSH and with (b and d) 10.0 mM GSH in PBS at 37 °C. (C) Flow cytometric profiles of HeLa cells incubated with the DOX-loaded mPEG113-S2-PZLL18 micelles for 2 h: (a) cells without pretreatment; (b) cells pretreated with 0.5 mM BSO; (c) cells pretreated with 10.0 mM GSH-OEt. Reproduced with permission from [2]. 15.2.2 Polypeptide Vesicles Polymeric vesicles (also known as polymersomes) are one kind of nanoscale or microscale hollow reservoir-like spheres composed of hydro- phobic interlayer, hydrophilic external and internal shells [55–59]. In con- trast to the polymeric micelles, polymeric vesicles can both encapsulate hydrophilic therapeutic molecules in their aqueous cavities and integrate hydrophobic bioactive agents within the hydrophobic membrane [57, 58]. Polymeric vesicles have great potential as versatile carriers ascribed to their colloidal stability, tunable membrane thickness and permeability,

530 Advanced Biomaterials and Biodevices and ability to encapsulate or integrate a broad range of drugs. The rela- tively long blood circulation times can be achieved by the introduction of a hydrophilic nonfouling surface layer (e.g., PEG) [57]. Additionally, diverse functionalities and stimuli-responsiveness can be incorporated into the structures of vesicles to fabricate the intelligent platforms for controlled delivery of bioactive agents, particularly antitumor drugs [58]. Of these, the vesicles based on stimuli-responsive polypeptides (e.g., pH- responsive ones) have been developed as emerging systems for antitumor drug delivery. Recently, the pH-responsive vesicle from poly(trimethylene carbonate)-block-P(L-glutamic acid) (PTMC-b-PLGA) was developed by Lecommandoux and coworkers for smart DOX delivery (Figure 15.3A, B and C). The vesicle was comprised of a PTMC interlayer as well as outer and inner PLGA shells, and was prepared by either a direct dissolution or solvent displacement (nanoprecipitation) method [52]. An ionizable antitumor drug (i.e., DOX) was encapsulated into the PTMC-b-PLGA vesicle at a pH of either 7.4 or 10.5 [54]. The distribution of DOX in the vesicle was significantly influenced by the loading pH, which can prob- ably be attributed to the ionization or non-ionization of DOX at pH below or above its pKa (i.e., 8.25), respectively [3, 60]. When loaded at pH 7.4, positively charged DOX was partially absorbed in the PLGA shell, whereas the neutral DOX was encapsulated inside the PTMC layer when loaded at pH 10.5. As depicted in Figure 15.3D, in vitro DOX released from the drug loaded vesicle was pH and thermo-dependent. An obviously faster release was observed at pH 5.5 in comparison with that at pH 7.4, which resulted from the improvement of the hydrophilicity of DOX, and reduc- tion of electrostatic interaction between PLGA and DOX at pH 5.5 [21, 54]. Furthermore, the faster DOX release was obtained as the temperature increased from 25 to 45°C, perhaps due to the enhanced permeability and mobility of the vesicle membrane. Additionally, the above vesicular system has also been exploited for the codelivery of DOX and superparamagnetic iron oxide nanoparticle (USPIO, i.e., γ-Fe2O3) for magneto-chemotherapy and magnetic resonance (MR) imaging [61, 62]. 15.2.3 Polypeptide Nanogels Polymeric nanogels are the swellable nanoscale crosslinked particles that are prepared through emulsion polymerization, precipitation polymerization, photolithographic and micromolding techniques, radical heterogeneous polymerization, supramolecular assembly and shell or core crosslinking of polymeric micelles, etc. [63–65]. The nanogels exhibited great prospects

Smart Polypeptide Nanocarriers for Malignancy Therapeutics 531 A PTMC-NH2 BLG PTMC-b-PBLG B THF/MeOH H2/Pd PTMC-b-PLGA PTMC-b-PLGA D 60 C 50 Cumulative DOX Release (% initial load) 40 30 20 20° C, pH 7.4 10 37° C, pH 5.5 37° C, pH 7.4 45° C, pH 7.4 5° C, pH 7.4 200 nm 0 0 4 8 12 16 20 24 Time (hour) Figure 15.3 (A) Synthesis of PTMC-b-PLGA. (B) Schematic representation of vesicle obtained from the self-assembly of PTMC-b-PLGA in water. The green part (deep color) represents PTMC and the blue part (light color) expresses PLGA. (C) Transmission electron microscopy (TEM) microimage of the DOX-loaded vesicle prepared at pH 10.5. (D) Cumulative release of DOX loaded within PTMC-b-PLGA vesicle (Rh = 60 nm) in various pH and temperature conditions. The DOX-loaded vesicles were prepared at pH 10.5. Reproduced with permission from [52–54]. as potential drug delivery platforms, which benefited from their tunable chemical and three-dimensional physical structures, high stabilities, excel- lent drug loading capabilities, and responsiveness to environmental fac- tors, such as pH, redox, temperature and ionic strength [3, 66]. Among them, the stimuli-responsive nanogels with nonfouling biocompatible hydrophilic shells have attracted much attention ascribed to their excellent dispersions in aqueous environment and stabilities in the complex in vivo circulatory system [23, 67, 68]. For malignancy therapy, various antitumor drugs can be loaded into the restrictively stable nanogels and released as

532 Advanced Biomaterials and Biodevices needed, triggered by the microenvironmental stimuli at lesion sites [3, 23]. In the past few years, polypeptide nanogels, which respond to pH or reduc- tion stimuli, have been exploited for smart antitumor drug delivery. 15.2.3.1 pH-Responsive Polypeptide Nanogels As in the abovementioned micellar and vesicular systems, pH-respon- sive polypeptide nanogels that undergo swelling-shrinking transitions in response to pH change, particularly at the acidic tumor tissular or intracel- lular endosomal pH, are meaningful as pH-triggered drug delivery nano- vehicles [66]. PEG-b-PLA nanogels were prepared by crosslinking the PLA segments with 1,6-hexanediamine and N,N’-diisopropylcarbodiimide as crosslinker and coupling agent, respectively [70, 71]. The nanogels exhib- ited a pH-dependent swelling-shrinking transition. As the pH increased from 4.0 to 9.0, the sizes of nanogels increased from less than 20 to above 40 nm, ascribed to the swelling of polypeptide cores that resulted from the gradual ionizations of LA residues at pH above their pKa (∼ 3.9) [13]. The DOX-loaded nanogels were fabricated by mixing DOX with nanogels in deionized water. It was observed that DOX was rapidly released from the nanogels at both pH 5.0 and 7.4, while a slightly faster drug release was observed at pH 5.0 than that at pH 7.4. The swollen state of nanogels at 7.4 facilitated the drug diffusion, and the increased solubility of DOX and reduced interaction between LA unit and DOX resulted in the faster release of DOX at acidic pH. As shown in Figure 15.4A, the reversible or irreversible nanogel was prepared by crosslinking PEG–PLA–poly(L-phenylalanine) (PLP) block copolymer micelle with an acid-cleavable ketal-containing crosslinker (K-nanogel) or an insensitive one (nanogel) [69]. DOX was loaded into nanogels through nanoprecipitation and subsequent shell crosslinking via condensation reaction. The DOX-loaded nanogel containing acid-uncleav- able linker (DOX–nanogel) exhibited a similar drug release pattern at pH 5.0 and 7.4. In contrast, drug release from the DOX-loaded nanogel with acid-liable ketal bond (DOX–K-nanogel) was accelerated at pH 5.0 com- pared with that at pH 7.4. Herein, an acid-catalyzed hydrolysis of the ketal bond in nanogel was proposed to be in charge of the improved DOX release at acidic condition (Figure 15.4B). After incubation with MCF-7 cells for 1 or 5 h, DOX–K-nanogel resulted in an enhanced intracellular DOX flu- orescence intensity compared to DOX–nanogel, indicating an improved DOX release from the former nanogel in acidic endosome or lysosome. In Chen’s group, a series of pH-responsive polypeptide nanogels based on PEG–poly(L-glutamic acid-co-γ-cinnamyl-L-glutamate)

Smart Polypeptide Nanocarriers for Malignancy Therapeutics 533 A Self-assembly Doxorubicin PEG45–PLA8–PLP19 Ketal cross-linker Endosome (pH~5.0) DOX release Nucleus (a) 1 min CSuperimpose DOX LysoTracker B (a) (b) (c) (d) (e) (b) 1 h 2.5 2.0 1.5 1.0 δ (ppm) (c) 3 h 2.5 2.0 1.5 1.0 δ (ppm) 2.5 2.0 1.5 1.0 δ (ppm) Figure 15.4 (A) Illustration of the DOX-loaded nanogels with a pH-labile ketal crosslinker and intracellular release of DOX triggered by the endosomal pH. (B) Proton nuclear magnetic resonance (1H NMR) spectra of K-nanogel in deuterium oxide at pH 5.0 at various incubation periods and the postulated corresponding structures of ketal crosslinked shell domains. Arrows indicate the gradual disappearance of ketal peaks. (C) Confocal laser scanning microscopy (CLSM) microimages of MCF-7 cells incubated with (a) free DOX for 1 h, (b) DOX–K-nanogel for 1 h, (c) DOX–K-nanogel for 5 h, (d) DOX–nanogel for 1 h and (e) DOX–nanogel for 5 h (equivalent 5 μg mL-1 DOX; green fluorescence is associated with LysoTracker and the red fluorescence is expressed by free DOX and released DOX). Scale bar = 20 μm. Reproduced with permission from [69]. (P(LGA-co-CLG)) block copolymers were prepared by the initiator-free photocrosslinking [64]. As pH increased from 4.0 to 7.4, there was a sig- nificant increase in the hydrodynamic radii (Rhs) of nanogels attributed to the swelling of polypeptide cores induced by the gradual ionizations of LGA residues. Drug-loaded nanogels were prepared through the mix- ture of PEG–P(LGA-co-CLG) copolymers with a model drug, rifampin, and subsequent in situ photocrosslinking. The rifampin-loaded nanogels exhibited a faster drug release at pH 7.4 compared to that at pH 4.0 due to the swelling of the nanogels at neutral or alkaline condition.

534 Advanced Biomaterials and Biodevices 15.2.3.2 Reduction-Responsive Polypeptide Nanogels As the pH-responsive polypeptide nanogels, the reduction-responsive polypeptide nanogels are another potential intracellular drug delivery sys- tem that benefit from the obvious different GSH level outside and inside cells. Jing and coworkers have developed the reversible nanogel through the shell crosslinking of poly(L-cysteine)-b-poly(L-lactide) micelle [72]. The diameter of nanogel was reduction-dependent, which increased slightly from 41.7 to 55.1 nm after the addition of DTT, while it decreased to 47.1 nm after the removal of DTT. Rifampin, as a hydrophobic model drug, was loaded into nanogel via the self-assembly of copolymer and drug, and followed aerial oxidation with DLE at 17.5 wt%. In comparison to its parent micelle, the rifampin-loaded nanogel with a larger diameter (∼ 65 nm) showed faster drug release in the presence of DTT ascribed to the decrosslinking of nanogels induced by the DTT-induced cleavage of disulfide bond. More meaningfully, the disulfide-core-crosslinked nanogels based on polypeptides have been independently developed by a facile one-step ring- opening polymerization (ROP) technique by the groups of Chen and Yan [73, 74]. In two separate studies, a difunctional N-carboxyanhydride (i.e., L-cystine NCA (LC NCA)) was synthesized. The reduction-responsive nanogels were fabricated through the copolymerization of LC NCA indi- vidually or together with other NCA monomers (e.g., L-phenylalanine (LP) NCA and γ-benzyl-L-glutamate (BLG) NCA). As a model antitumor drug, DOX was loaded into mPEG–P(LC-co-LP) nanogels through the diffusion approach [73]. More than 90 wt% of loaded DOX was released from nano- gels at 93.5 h in the presence of 10.0 mM GSH, while only less than 20 wt% of loaded DOX was released in PBS without GSH. Additionally, a faster DOX release was observed from the drug-loaded nanogels with lower con- tent of LC or overall polypeptides. Furthermore, the DOX-loaded nanogel exhibited enhanced intracellular drug release and antiproliferative activ- ity toward GSH-pretreated HeLa cells. Similarly, a GSH-mediated indo- metacin release from a PEG–P(LC-co-BLG) nanogel was confirmed by Yan and coworkers [74]. In addition, the near-infrared cyanine dye modi- fied reduction-responsive polypeptide nanogels were further exploited as directly imaged intracellular antitumor drug carrier in Yan’s group [75, 76]. With convenient fabrication, excellent stability in the circulation systems, accelerated intracellular drug release and potential modification, the above polypeptide nanogels hold great prospects for chemotherapy in malig- nancy therapeutics.

Smart Polypeptide Nanocarriers for Malignancy Therapeutics 535 A BADS mPEG-NH2 PLG NCA mPEG-b-PPLG Crosslinking Micellization bads B 100 a C 150 ac b b 80 c 100 d 60 DOX Release (%) Counts 40 50 20 0 0 100 101 102 103 0 10 20 30 40 50 60 70 80 10–1 Fluorescence Intensity Time (h) Figure 15.5 (A) Preparation procedure of reduction-responsive nanogel. (B) In vitro release of DOX from (a) the DOX-loaded micelle, (b) nanogel without GSH, (c) nanogel with 5.0 mM GSH and (d) nanogel with 10.0 mM GSH in PBS at pH 7.4, 37°C. (C) Flow cytometric profiles of HeLa cells incubated with the DOX-loaded nanogel for 2 h: (a) cells without pretreatment; (b) cells pretreated with 0.5 mM BSO; (c) cells pretreated with 10.0 mM GSH. Reproduced with permission from [77]. Moreover, Chen and coworkers have prepared the reduction-responsive polypeptide nanogel through core crosslinking poly(ethylene glycol)- block-poly(γ-propargyl-L-glutamate) (PEG-b-PPLG) micelle by Cu(I)- catalyzed Huisgens cycloaddition “click” reaction with bis(2-azidoethyl) disulfide (BADS) (Figure 15.5A) [77]. The DOX-loaded nanogel displayed accelerated in vitro drug release in a reductive condition (Figure 15.5B) and exhibited enhanced intracellular DOX release (Figure 15.5C) and in vitro cellular proliferation inhibition toward HeLa cells with the pretreat- ment of GSH. 15.2.3.3 pH and Reduction Dual-Responsive Polypeptide Nanogels Of all the intelligent nanogels, multiple stimuli-responsive nanogels, especially those responding to both pH and reduction, exhibited great promise for targeting intracellular transportation of antitumor drugs.

536 Advanced Biomaterials and Biodevices A  typical instance is the interlayer-crosslinked polypeptide nanogel that has been developed in Shuai’s group [78]. As shown in Figure 15.6A, the pH and reduction dual-responsive nanogel based on poly(ethylene gly- col)-block-poly(N-(2-mercaptoethyl) L-aspartamide)-block-poly(N-(2- (diisopropylamino)ethyl) L-aspartamide)) (PEG-b-PMELA-b-PDIPLA) was prepared through the disulfide-crosslinking of PMELA interlayer at pH 10.0. The nanogel showed pH and reduction dual-responsive properties A B (a) (b) pH10 High H2O pH5.0 DTT O2 Dialysis HP-ICM (59.4 nm) pH7.4 (c) pH10 ICM (47.7 nm) D (a) C PEG-b-PMELA-b-PDIPLA DOX Nuclei DOX Merged Bright Field low Free (b) 1200 HP-ICM DOX PEG-PCL micelle 1000 Free DOX Tumor Volume/mm3800 PBS 600 Administration 400 PEG-PCL 200 Micelle HP-ICM 50 5 10 15 20 25 30 40 Time/d 30 20 10 0 0 Figure 15.6 (A) Formation and structural transitions of the dual-responsive nanogel (n = 45, m = 15 and k = 14; determined by 1H NMR). (B) TEM microimages of the nanoassemblies at (a) pH 7.4, (b) 5.0, (c) 7.4 with 10.0 mM DTT and (d) 5.0 with 10.0 mM DTT. The nanogel shown in (a) was decorated with Au, and other samples were stained with uranyl acetate. (C) Intracellular DOX release and migration into nuclei observed by CLSM. Nuclei were stained with Hoechst 33342 (blue). (D) (a) In vivo DOX fluorescence images showing passive tumor accumulation of the DOX-loaded nanogel after tail-vein injection into nude mice bearing the Bel-7402 xenograft (dose: 5.0 mg DOX per kg body weight); (b) Tumor growth inhibition in nude mice bearing the Bel-7402 tumor after tail-vein injection of different formulations (n = 20; dose: equivalent 5.0 mg DOX per kg body weight per injection for DOX or the DOX-loaded nanoparticles). *p < 0.01 versus PEG-b-PCL micelle. Reproduced with permission from [78].

Smart Polypeptide Nanocarriers for Malignancy Therapeutics 537 attributed to the presence of both the tertiary amino group in the core and disulfide bond in the interlayer. The nanogel took a solid spherical at pH 7.4 (Figure 15.6B-a), whereas it showed an apparent nanocage structure at pH 5.0 without DTT ascribed to the interlayer constrained complete dissolution of the inner PDIPLA segment (Figure 6B-b). As shown in Figure 6B-c and d, after the addition of 10.0 mM DTT, the nanogel evolved into swollen micelle at pH 7.4, while it disassembled at pH 5.0 due to the cleavage of disulfide bond and the protonation of PDIPLA block. DOX, a model antitumor drug, was encapsulated into nanogels through mixing the triblock copolymer and DOX, and subsequently crosslinking the inter- layer through bubbling of an oxygen flow. The DOX release from drug- loaded nanogel was significantly accelerated by either reducing the pH to 5.0 or adding 10.0 mM DTT. In addition, the DOX-loaded nanogel exhib- ited a fast accumulation of DOX in the nuclei of Bel-7402 cells (a human hepatoma cell line), while the distribution of DOX was mainly in the cyto- plasm for the DOX-loaded poly(ethylene glycol)-block-poly(ε-caprolac- tone) (PEG-b-PCL) micelle (Figure 15.6C). In addition, the DOX-loaded pH and reduction dual-responsive nanogel exhibited remarkable aggrega- tion in tumor and the most outstanding antitumor efficacy compared with both free DOX and the DOX-loaded PEG-b-PCL micelle. Zhuang, Chen and coworkers have developed the pH and reduction dual-responsive polypeptide nanogels through the one-step ROP of LC NCA and BLG NCA or ZLL NCA, and subsequent removal of protecting groups [3]. DOX was loaded into nanogels through the diffusion technique, and a high DLE was obtained for mPEG–P(LC-co-LGA) nanogels due to the electrostatic interaction between LGA unit and DOX. The in vitro DOX release from the DOX-loaded nanogels could be accelerated in mimicking intracellular reductive (10.0 mM GSH) and/or acidic conditions (pH 5.5). The DOX-loaded nanogels exhibit faster DOX release behavior in GSH- OEt pretreated HeLa cells than that in unpretreated and BSO pretreated cells. In addition, higher cellular proliferation inhibition efficacy was achieved toward GSH-OEt pretreated HeLa and HepG2 cells compared to unpretreated and BSO pretreated cells. In addition, the pH and reduc- tion dual-responsive mPEG–P(LC-co-LL) nanogel was also employed to deliver various negatively charged DOX derivatives including acid-insen- sitive succinyl-DOX (SAD), acid-sensitive cis-aconityl-DOX (CAD) and 2,3-dimethylmaleyl-DOX (DAD) [67]. In vitro drug release could be accel- erated in the intracellular acidic and/or reductive conditions. Drug-loaded nanogels exhibit a faster drug release behavior in GSH-OEt pretreated HeLa cells than in unpretreated or BSO pretreated cells. Furthermore, the CAD- and DAD-loaded nanogels exhibited a higher anti-proliferative

538 Advanced Biomaterials and Biodevices activity than the DOX-loaded nanogels after incubation for 72 h, while free CAD and DAD showed lower cytotoxicities toward both HeLa and HepG2 cells in contrast to free DOX for the duration of test. Furthermore, in Chen’s group, the diselenide bond was employed to crosslink methoxy poly(ethylene glycol)-block-poly(L-glutamic acid-co-γ-2-chloroethyl-L- glutamate) micelle, yielding a pH and dual-redox responsive nanogel [23]. DOX was loaded into nanogel through the diffusion approach, and a faster release was observed in the presence of GSH (only 0.5 mmol). Deng, Zhong and coworkers have exploited the pH and reduction dual-responsive nanogels through crosslinking lipoic acid (LA) and cis- 1,2-cyclohexanedicarboxylic acid (CCA) decorated PEG-b-PLL (PEG-b- P(LL-g-LA/CCA)) micelles in the presence of a catalytic amount of DTT in PBS at pH 7.4 [79]. DOX was loaded into nanogels through nanopre- cipitation and subsequent core crosslinking. The release of DOX from drug-loaded nanogels was improved at endosomal pH (i.e., 5.0) or under a reductive condition containing 10.0 mM GSH, possibly triggered by the cleavage of acid-labile amide bond of CCA or disulfide bond, respectively. The pH and reduction dual-responsive nanogels were biocompatible, whereas the DOX-loaded nanogels caused pronounced cytotoxic effects toward HeLa and HepG2 cells. 15.2.4 Other Smart Polypeptide Nanovehicles In addition to the abovementioned mainstream systems, there are some other interesting smart polypeptide nanovehicles with various morphol- ogies (e.g., nanocapsules) and stimuli-responsivenesses (e.g., thermo- responsiveness) for intelligent antitumor drug delivery. For example, Huang and coworkers have prepared polypeptide nanocapsule through layer-by-layer self-assembly of PLGA and platinum(IV) conjugated PLL (PLL-g-Pt(IV)) on colloidal silica templates, followed by removal of the templates [80]. The release rate and amount of platinum from nanocap- sule were increased in acidic and/or reductive conditions. In addition, the PLGA/PLL-g-Pt(IV) nanocapsule displayed enhanced antitumor efficacy against CT-26 cells (a murine colon carcinoma cell line) in comparison with free CDDP. Moreover, a series of thermo-responsive micelles based on “hairy-rod” polypeptides, which were efficiently synthesized by graft- ing of azide-terminated copolymers of 2-(2-methoxyethoxy)ethyl meth- acrylate (MEO2MA) or 2-(2-(2-methoxyethoxy)ethoxy)ethyl methacrylate (MEO3MA) (i.e., N3-PMEOiMA) onto PPLG through a “click” reaction, were exploited by Chen’s group [39]. DOX was efficiently loaded into

Smart Polypeptide Nanocarriers for Malignancy Therapeutics 539 micelles through nanoprecipitation, and DOX release was accelerated by a decrease of temperature and/or pH. The DOX-loaded micelles could be internalized into HeLa cells and efficiently release the payload. More significantly, the micelles demonstrated good cytocompatibility, while the DOX-loaded micelles could effectively inhibit the cellular proliferation. 15.3 Conclusions and Perspectives In this chapter, a summary of the recent developments of stimuli-respon- sive polypeptide nanocarriers for smart antitumor drug delivery was presented. Various advantages including excellent biocompatibility, appro- priate biodegradability, effective drug loading, “on-demand” drug delivery triggered by stimuli in tumor tissular and/or intracellular microenviron- ments (e.g., pH and/or reduction), as well as improved therapeutic effi- cacy in vitro and in vivo, have been achieved through the dexterous design and accurate preparation of intelligent nanovehicles with polypeptides as matrices. Particularly, the pH and/or reduction dual-responsive micelles, vesicles and nanogels have been intensively developed for targeting intra- cellular transmission of antitumor drugs, resulting in enhanced chemo- therapy efficacies and reduced side effects. However, most of the studies described in this chapter are currently in the proof of concept stage. It should be affirmed that the above-presented systems have revealed the design rules for intelligent nanocarriers and demonstrated the feasibility of improving chemotherapy efficacies through controlled delivery with smart polypeptide nanovehicles. Over time, the application of stimuli-responsive polypeptide nanocarriers for smart drug delivery probably will become a clinical reality benefiting from their excellent biocompatible matrices and improved antiproliferative activities in vitro and in vivo. References 1. Z.Y. Xiao, C.W. Ji, J.J. Shi, E.M. Pridgen, J. Frieder, J. Wu, and O.C. Farokhzad, DNA Self-assembly of targeted near-infrared-responsive gold nanopar- ticles for cancer thermo-chemotherapy, Angewandte Chemie, International Edition, Vol. 51, Iss. 47, pp. 11853–11857, 2012. 2. J.X. Ding, J.J. Chen, D. Li, C.S. Xiao, J.C. Zhang, C.L. He, X.L. Zhuang, and X.S. Chen, Biocompatible reduction-responsive polypeptide micelles as nanocarriers for enhanced chemotherapy efficacy in vitro, Journal of Materials Chemistry B, Vol. 1, Iss. 1, pp. 69–81, 2013.

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