Sustained-release from nanocarriers: a review.

COREL-07202; No of Pages 17 Journal of Controlled Release xxx (2014) xxx–xxx

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Sustained-release from nanocarriers: a review Jayaganesh V. Natarajan, Chandra Nugraha, Xu Wen Ng, Subbu Venkatraman ⁎ Nanyang Technological University, School of Materials Science and Engineering, Blk N4.1, Nanyang Avenue, Singapore 639798

a r t i c l e

i n f o

Article history: Received 25 February 2014 Accepted 17 May 2014 Available online xxxx Keywords: Nanocarriers Drug delivery Targeting Sustained drug release Localized drug release

a b s t r a c t Nanocarriers have been explored for delivering drugs and other bioactive molecules for well over 35 years. Since the introduction of Doxil®, a nanoliposomal delivery system for the cancer drug doxorubicin, several products have been approved worldwide. The majority of these products focus on cancer chemotherapy, and utilize the size advantage of nanocarriers to obtain a favourable distribution of the drug carrier in the human body. In general, such carriers do not sustain drug release over more than a few days at best. In this review, we explore the reasons for this, and present an overview of successful research that is capable of generating sustained-release products in non-cancer applications. A variety of nanocarriers have been studied, and their advantages and shortcomings are highlighted in this review. The achievement of sustained release of bioactive molecules opens new doors in nanotherapeutics. © 2014 Elsevier B.V. All rights reserved.

1. Introduction A drug delivery system (DDS) can be defined as a device or a formulation that is capable of introducing a therapeutic substance into the body in a manner that enhances its safety and efficacy over the two “standard” methods of drug administration: oral tablets and intravenous (IV) injections. The improved efficacy can be due to greater localization of drug, enhanced bioavailability of the drug or sustained duration of action. Of these, sustained release delivery systems have been intensely studied for many decades and have enjoyed some success in the pharmaceutical arena. A drug delivery system that is capable of achieving a prolonged therapeutic effect by slow release of the therapeutic substance over an extended duration (days or months) after administration of a single dose is termed as a sustained release delivery system. There are several advantages of sustained release dosage forms which include 1) Lesser frequency of administration 2) Reduced side effects 3) Stable drug absorption levels in blood and plasma 4) Better patient compliance. Accordingly, physico-chemical (size, dosage size, solubility, partition etc) and biological factors (half-life, absorption, distribution, metabolism etc) need to be accounted for the optimized design of sustained release dosage forms. Also, suitable in vitro release methods should be identified for evaluation of release from various dosage forms including nanoparticles. A major challenge in the evaluation of release from nanoparticles is the lack of a universal method for quantifying release in vitro. This has hindered the interpretation of in vitro release data and thus the ⁎ Corresponding author at: Nanyang Technological University, School of Materials Science and Engineering, Blk N4.1 #02-05, Nanyang Avenue, Singapore 639798. Tel.: +65 6790 4259; fax: +65 6790 0921. E-mail address: [email protected] (S. Venkatraman).

comparison between different classes of nanocarriers such as liposomes, micelles and polymeric nanoparticles. Some of these methods reported in the literature include microdialysis [1], fractional dialysis [2], reverse dialysis [3], sample and separate methods [4]. Different experimental techniques that are used to evaluate release of drugs from nanoparticles are reviewed elsewhere [5,6]. We will highlight the main differences or challenges for drug release studies involving nanocarriers as opposed to their micron-sized counterparts. In general, when nanocarriers incorporating drug are incubated in buffer receiver fluid, it is difficult to separate the released drug from the encapsulated drug, unlike the case of microparticulate suspensions which may be centrifuged or even filtered. Thus, dialysis is by far the preferred method for nano-sized carrier systems. In this method, drug-incorporated nanoparticles are “suspended” in a buffer inside a dialysis tube with a certain molecular weight cut-off (MWCO). The dialysis tube is separated from the release medium and assayed for drug release at various time intervals [6]. The choice of the membrane MWCO is critical here, as the choice will dictate whether release of drug from nanocarriers or membrane diffusion is the rate-limiting step [7]. For the dialysis membrane, various modifications have been proposed, such as using a more standardized surface area of the membrane [7], or by dispersing the nanoparticles into a cell covered by a dialysis membrane, and rotating of said cell at certain speed [8]. In these methods, they were able to distinguish certain release profiles which otherwise were indistinguishable through normal dialysis membrane. Apart from dialysis, another commonly considered method is by separation of the sample through centrifugation. This is a commonly used procedure for evaluating drug release from microparticles. Apart from the shared problem with centrifugation with dialysis (whether sink condition criterion is fulfilled), there is an additional concern of whether particles can be re-dispersed properly into the buffer. This is

http://dx.doi.org/10.1016/j.jconrel.2014.05.029 0168-3659/© 2014 Elsevier B.V. All rights reserved.

Please cite this article as: J.V. Natarajan, et al., Sustained-release from nanocarriers: a review, J. Control. Release (2014), http://dx.doi.org/10.1016/ j.jconrel.2014.05.029

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J.V. Natarajan et al. / Journal of Controlled Release xxx (2014) xxx–xxx

more pertinent due to the smaller size of nanoparticles that makes it more difficult to both centrifuge non-destructively and to re-disperse the pellet. The use of centrifugation, therefore, is not recommended for nanoparticles which are unable to withstand high centrifugal force. More recently, efforts are being made to determine the drug amounts directly without resorting to the need for separation of nanoparticles from the drug; such methods include differential pulse polarography [9], voltammetric techniques [10], or using a Langmuir balance [11]. However, there are other associated limitations with these methods, such as in the type of drug to be quantified, potential signal interference with the nanocarrier, or simply the additional dedicated equipment involved. Under the broad umbrella of Nanomedicine which includes the use of nanotechnology for: (a) delivering drugs and genes; (b) for enhancing in vivo imaging; (c) for enhancing the sensitivity of detection of biomarkers in in vitro diagnostics; and (d) for functionalizing biomaterial surfaces to enhance the performance of implants, nanoparticulate drug/protein/gene delivery remains the major component in terms of research papers published, patents granted, and as well as approved products [12]. Rather surprisingly, however, none of the approved products using nanoparticulate carriers can lay claim to sustained release as one of the principal attributes. Admittedly, it is more difficult to achieve prolonged release of small or high molar mass drugs from nanocarriers compared to their micron-sized counterparts, where a number of products have been approved. It is the purpose of this review to examine critically why this situation exists, and whether we can expect more sustained-release nanomedicinal products in the future. The major focus of the research and development of nanocarriers is to exploit their enhanced cellular penetration [13,14]. The second reason for using nanoparticles is to increase blood circulation lifetimes, which facilitates passive targeting of tissues [15,16]. Although nanoparticles have been explored extensively for active targeting of tissues, this approach has not been clinically successful to date [14,17]. The other use of nanoparticulates is to improve solubility of highly hydrophobic drugs such as paclitaxel while retaining injectability of the formulation [18,19]. On the other hand, there are microparticulate products that are classified as sustained-release: for example, Lupron Depot™ is a microparticulate formulation of leuprolide acetate in biodegradable poly (lactide-co-glycolide) particles, used as a subcutaneous injection for benign prostatic hyperplasia [20]. Here the particle sizes average about 50–100 microns, and sustained release is achieved over several weeks. Another example is Vivitrol®, a once-a-month sub-cutaneous injection of PLGA microspheres for treatment of alcohol dependence. In fact there are about 8 microsphere carrier formulations that are approved based on their sustained-release characteristics, to treat conditions ranging from dwarfism to acromegaly and schizophrenia [21]. Most of them use PLGA (a copolymer of poly (L-Lactide) and poly (glycollide)). In general, sustained delivery is warranted for many chronic conditions to achieve better patient compliance and safety. Currently, sustained release is achieved with some success by using microspheres injected intramuscularly (IM) or subcutaneously (SC). On the contrary, microparticles do not last long enough in the bloodstream for passivelytargeted chemotherapeutic particles when injected intravenously and their penetration into cells (delivery of plasmid DNA or siRNA) is almost negligible compared to nanocarriers. It is also a corollary that it is more difficult to achieve sustained release from nanoparticles than from microparticles, for 2 reasons: (a) Drug loading is generally limited and inefficient for nanocarriers (b) Release is faster, and there is usually a pronounced burst release for nanocarriers

It is possible that the difficulty of achieving sustained release from nanocarriers is the reason behind the lack of approved nanomedicine products in that category, but it is also likely that there were no compelling advantages for nanocarrier use in some of the above applications. For example, it is unlikely that sub-cutaneous injections of nanoencapsulated leuprolide acetate offers any advantages in bioavailability or pharmacokinetics over microspheres, because the likely path to the systemic circulation for the peptide is via molecular diffusion from the particles to the blood vessels in the dermis. All other factors being equal, it is likely that nanospheres will release leuprolide acetate faster than microspheres and hence may even prove disadvantageous for this application. Thus this review will purport to highlight first of all how far nanocarriers (Fig. 1 shows a pictorial representation of different types of nanocarriers) have been successful in achieving sustained release, and where their nanosize is a definite advantage for efficacy of action of the encapsulated bioactive molecule. 2. Liposomal carriers A number of nanoparticles have been evaluated over the years to improve loading and sustained delivery of drugs, including liposomal nanoparticles, polymer nanoparticles and nanosuspensions. Among the nanoparticles studied, liposomes have been the most successful drug delivery carriers [22]. In general, the research on liposomes as nanocarriers was focussed on the passive targeting (also called the Enhanced Penetration & Retention effect or EPR) for tumour shrinkage. For this concept to succeed, there are generally three requirements of the particles (Fig. 2): (a) Their blood lifetimes must be substantial enough for passive targeting to happen; this translates to about 24–50 hours halflife in blood (b) The sizes must be small enough to facilitate fenestration through the leaky blood vessels around the tumour site; empirically this requirement is diameter b 200 nm (c) Clearly, the release duration of the drug from the circulating nanocarriers must be longer than the half-life; otherwise free drug will be released and cleared quickly from the bloodstream To satisfy requirement (a), the most successful approach appeared to be to “PEG-ylate” the particles, i.e., cover them with a corona of PEG molecules. Indeed, the first successful nanocarrier for drug delivery was Doxil® which used nano-sized liposomes (average size 90 nm). Poly(ethylene glycol) molecules (molar mass ~2000 Da ) are projected on to the surface of the liposomes by mixing the PEGylated lipids with the main lipids that formed the liposome [23]. The chemotherapeutic drug, doxorubicin, was loaded into the aqueous core of the liposome at a concentration exceeding saturation. This high loading (~12.5% by weight of the lipids) was achieved through a special technique called active loading using an ammonium sulphate gradient [25]. The doxorubicin that is undissolved exists in the core as crystals. It is this feature that delays the release of the drug beyond the circulation halftime, as drug dissolution becomes rate-limiting. This combination of properties (PEGylation, active loaded crystallized drug and small size) enabled some selectivity of action towards tumor tissue, thus reducing side effects of the drug (schematic shown in Fig. 3). The biggest challenge in the development of Doxil® was certainly the loading of the drug in amounts that could be therapeutically meaningful, and released slowly enough. Doxil® was approved by the US FDA in 1995 [23]. Since then it has been used to treat AIDS-related Kaposi’s sarcoma, ovarian cancer and multiple myeloma in combination with bortezomib. The use of the nanocarrier does reduce some side effects of doxorubicin, such as cardio-toxicity but does not eliminate it. Nevertheless, the carrier formulation is preferred to naked doxorubicin in chemotherapy. Since the introduction of Doxil® several other liposomal products have been developed and approved for human use [26]. They can be



Fig. 1. Pictorial representation of different nanocarrier types and their functionalities. Targeting moieties, surface modifiers (PEG molecules) and incorporated agents (drugs and contrast agents) are shown.

classified based on the size of the liposomes and/or lipid complex as summarized in Table 1. In general, none of these products were approved on the basis of their duration of action. Most of the cancer-related products were approved on the basis of reduced side effects due to their passive targeting capabilities. Other liposomal anticancer products, such as DaunoXome and Myocet were primarily used to reduce toxicity in comparison to free doxorubicin and not to sustained release of encapsulated drug [40]. Another difficulty in liposomal product development is the stability of the drug-encapsulating liposomal formulation. Generally, lyophilization is used to ensure the stability of the active agent and the lipid, as otherwise there are either aggregation problems or drug leakage issues or both. Ambisome and Visudyne are two such examples [26]. On the other hand, reconstitution with a diluent/buffer could result in both variable size ranges and hence unpredictable drug release after injection, thus leading to variable therapeutic outcomes. This is particularly critical for sustained-release products where the reproducibility of the overall release profiles is essential to product performance.

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clinical studies. Such molecules leak out from the bilayer, hence any product that is developed will have to be lyophilized and reconstituted prior to use, rather than stored as a suspension. This is the problem that our group has recently overcome, in the development of a long-acting nanoliposomal drug for glaucoma [41]. This has been accomplished by the use of partition control of release rather than diffusion control. In the concept of partition control, it is the physico-chemical interaction of the drug molecule (Fig. 5) that controls the efflux of the drug (latanoprost) from the bilayer. Polar and nonpolar interactions may be utilized to advantage in this regard. For hydrophilic entities other than cancer drugs, there have been attempts to incorporate these in liposomes and to control their efflux. Starting with doxorubicin, which was remote-loaded into the aqueous core of the liposomes using special techniques, a comparison study of how the phospholipid transition temperature (Tg–f; this is the gel to fluid transition) affects release has been reported [42]. Several phospholipids were evaluated with Tg–f ranging from −20 to +55 °C. As seen in Table 2 below, the release half-life is directly dependant on the Tg–f, with a half-time of release of 120 hours being attained when the Tg–f is 55 °C (with added cholesterol, for HSPC). Such attempts enabled the release of doxorubicin to be slower than the circulation half-life, which was the main goal for passive-targeting systems. Any further slowing down of release is not desirable for these systems, because the drug needs to be delivered to tumour tissue soon after accumulation. Thus for systems similar to doxorubicin, the maximum release duration is about 4–5 days. In a study by Bhardwaj et al. [43], hydrophobic drug, dexamethasone, was loaded into liposomes prepared from lipids that differed in their phase transition temperatures (DMPC, C14:0, Tm ~ 23 °C; DPPC, C16:0, Tm ~ 41 °C; and DSPC, C18:0, Tm ~ 55 °C) incorporated with cholesterol. The authors reported that sustained in vitro release of dexamethasone at 37 °C (about 42 days) could be achieved using micronsized vesicles (MLVs) that were prepared using a high phase transition temperature lipid, DSPC (Tm ~ 55 °C). On the contrary, release of the drug from nano-sized vesicles of similar composition could not be sustained beyond 5 days, due to relatively higher burst. The slower release of the drug in case of MLVs was attributed due to slower diffusion of the drug across multiple bilayers of the vesicles, when compared to single bilayer vesicles (LUVs). Also, a general trend of slower release with increase in chain lengths of the lipids (DMPC N DPPC N DSPC) was shown suggesting that differences in the physicochemical properties of the drug and phase transition of lipids could be used to control release rates for both hydrophilic and hydrophobic molecules.

2.1. Challenges for sustained release from liposomes The main challenge in developing sustained release liposomal products, however, lies in the difficulty of using diffusion control alone for long-term delivery. For nanoliposomes, the bilayer membrane dimensions are of the order of a few nanometers (5–10 nm) so that the diffusion path length is of the same order. Unless the lipid bilayer is rigid (in the sense of having a high transition temperature) the diffusion through this layer is relatively facile for drugs encapsulated in the aqueous core (hydrophilic drugs). Attempts to make this layer more rigid have been somewhat successful using cholesterol (Fig. 4). Nevertheless, the bilayer exists essentially only in either a fluid state (above its “melting” temperature) or in a gel-like state (below the transition temperature). Even the gel state is not as resistant to diffusion of small molecules as a polymeric matrix is, when the polymer is below its glass transition temperature (Tg). The gel state is more akin to the rubbery state of polymeric matrices, and hence diffusivity in gel matrices is about 2–3 orders of magnitude higher than through a poly (lactide-co-glycolide) matrix, whose Tg is above 37 °C. For lipophilic molecules inside liposomes, the diffusion path length is even shorter; hence the difficulty of sustained release is even greater than for hydrophilic drugs contained in the core. In fact, there are no reports at all of slow lipophilic drug release from liposomes, even in pre-

Fig. 2. Schematic showing both passive and active targeting by nanocarriers. In order to reach tumours, the carriers must circulate in blood for at least 24 – 30 hours, during which time they should release very little of the encapsulated drug. This is true of both active and passive targeting nanocarriers. Reprinted (adapted) from Dahnier et al. [24] with permission from Elsevier.


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Fig. 3. Schematic depiction of how the Doxil system works: crystallized drug in the core and the presence of PEG groups on the surface are shown; the PEG layer is responsible for long blood lifetimes while the crystallization allows for slow release of drug. Reprinted (adapted) from Barenholtz [23] with permission from Elsevier.

2.2. Hydrophilic drugs in liposomes For (hydrophilic) drugs loaded into the core of liposomes, the optimal control of release is achieved through drug crystallization in the core. For example, anthracyclines such as doxorubicin, duanorubicin and idarubicin were loaded in LUVs at a high D/L ratio (0.3, w/w). The release was slowest for doxorubicin due to the formation of drug precipitates (stacked together as dense precipitates) in the core with a relatively lower solubility and lower affinity of the drug to the bilayer [44]. Also, the dissolution of the precipitate into soluble doxorubicin was found to be the rate limiting step for doxorubicin release [45]. Ciprofloxacin, on the other hand, had a relatively faster release. One study reported that ciprofloxacin formed dense precipitates inside the core of liposomes similar to doxorubicin; however, faster release was attributed due to liposome rupture by drug crystals [46]. In another study, the nature of the drug was evaluated by 1H-NMR. Their studies concluded that ciprofloxacin unlike doxorubicin was not precipitated although the drug concentration inside the core exceeds its solubility by several orders of magnitude; the observed faster release is attributed to the lack of crystallization [47]. A similar explanation is offered for controlled release of encapsulated vincristine by Johnston et al. [48]. The drug release rate of bilayerpermeable species is correlated to the state and nature of the precipitate in the core of liposomes. At low D/L ratios, the drug exists as a soluble form, while with increase in D/L ratios; the drug exists as a precipitate with relatively lower proportion of soluble drug. The authors reported an increase in circulation half-lives of drug and slower release with increase in D/L ratios, suggesting that the dissolution rate of the precipitate into the soluble form of the drug was the rate limiting step at higher D/L ratios. A similar trend was also observed for liposomal formulations of doxorubicin that showed slower release rates and increased circulation half-lives with increase in D/L ratios [49]. Attempts to prolong the release of hydrophilic entities from nanoliposomes have been only partially successful. For example, an antibody (bevacizumab) originally approved for management of gastrointestinal tumours, has also been used off-label for halting angiogenesis in age-related macular degeneration. To reduce the frequency of intravitreal injections, sustained-release formulations have been evaluated. One such evaluation is reported by Abrishami et al., using eggPC and DPPC liposomes [50]. Extruded liposomes were rehydrated with bevacizumab solution, with reported efficiencies of about 15 to 45% depending upon whether the bevacizumab was added before or after freeze-drying. Cholesterol (1:1) was also used to make the membrane rigid, although it is not clear if the bevacizumab locates on the surface or in the core of the liposome in this case.

Nevertheless, injection of liposomal and non-liposomal bevacizumab does show a difference in retention of the antibody in rabbit vitreous humour over 42 days. At 42 days, the amount of bevacizumab detected in vitreous humour was about 4-fold higher for the liposomal formulation. This study indicates the persistence of the antibody in the vitreous humour, presumably attributable to slower clearance of the liposomes, and some degree of sustained release of the antibody from the liposomes. More importantly, the study shows the incorporation of therapeutically comparable doses of bevacizumab inside liposomes (at least at levels of ~ 30% encapsulated and ~ 70% free bevacizumab). Nevertheless, in this instance, attainment of about 40 days of release duration is at least partly due to slow clearance of the bioactive molecule itself from the posterior eye segment. There have been similar attempts to prolong the activity of antiangiogenesis factors in vitreous humour using liposomes, but no liposomal formulation incorporating an antibody has been translated to the clinic. A major reason for this is the fact that liposomes have been found to cause blurred vision after injection into the posterior eye segment [51]. The tendency for liposomes to aggregate is possibly enhanced in the viscous posterior segment, and dispersion is hindered. Attempts to reduce this have been reported using turbidity measurements in vitro [52] but the relevance of such measurements to the in vivo situation is questionable. This limitation, however, is probably applicable to all nanoparticulate systems for intra-vitreal administration. To take another example, vaso-intestinal peptide (VIP) has been successfully encapsulated into multilamellar vesicles [53] approximately 1 micron in size. The leakage was negligible for the encapsulated VIP, for 14 days at 4 °C, but no information is given regarding release characteristics at ambient conditions; presumably it is longer than the systemic half-life. When injected into rats, the mean arterial pressure was lowered substantially compared to the naked VIP, and this effect is maintained for several hours with a single injection. It is not clear where the VIP resides, as the method of preparation is via reverse evaporation; most likely it is associated with the surface of the vesicles rather than in the core. Loading of hydrophilic bioactive molecules into liposomes remain a challenge. Insulin has been a favourite target for liposomal encapsulation, but its loading efficiency via reverse evaporation is only 2–8%; using a pH gradient and positively-charged vesicles improves this efficiency somewhat [54]. Thus insufficient loading of insulin has been the main reason for limiting its use in sustained delivery nanoparticle systems. 2.3. Lipophilic drugs in liposomes On the other hand, lipophilic entities have been more successfully incorporated, and their release sustained. A good example is dexamethasone, an anti-inflammatory steroid used for various indications, including restenosis in stents. Using eggPC liposomes, Kallinteri et al. were able to load dexamethasone using 3 different approaches [55]. Loading efficiency was high (~ 78%) when the dehydrate-rehydrate method (DRV) was employed; the release was also sustained longer when coated onto polymer-covered metallic stents for endoprosthesis applications, as seen below in Fig. 6. In all 3 cases, the loading is in the bilayer, and the maximum duration achieved is of the order of 3 days. Liu et al. [56] developed a novel hydrophobic anticancer agent ML220 (2-(5-bromo-1H-indol-3-yl)1H-phenan-thro-[9, 10-d] imidazole and studied the release and anticancer properties in a colon HT29 carcinoma mice model. The authors were able to load the drug in very high concentrations (~ 29% by mole ratio) by using a mixture of egg phosphatidylcholine (EggPC) and polyethylene conjugated distearoyl-phosphatidyl ethanolamine (DSPE-PEG200) nanoliposomes. The in vitro release was sustained for over two weeks (50% drug release at the end of two weeks) with a minimal burst. This delayed release also led to lower cytotoxicity in two


No Product name Therapy

Drug

Particle diameter Type (Zavg)

Lipid composition

Drug form

Liposome – Nano sized products 1 Ambisome Fungal

Amphotericin B

b100 nm

LUV

Lyophilized powder To improve stability and reduce nephrotoxicity [27,28]

2 3

DaunoXome Doxil

Cancer Cancer

Daunorubicin Doxorubicin

~45 nm ~100 nm

SUV LUV

4

Lipo-Dox

Cancer

Doxorubicin

~180 nm

LUV

HSPC/Cholesterol/DSPG, D/L wt ratio = 0.16 DSPC/Cholesterol D/L wt ratio = 0.054 HSPC/Cholesterol/DSPEmPEG 2000 D/L wt ratio = 0.125 DSPC/Cholesterol/DSPEmPEG 2000

5 6 7

Myocet Visudyne Epaxal

Cancer Ocular Vaccine

Doxorubicin ~190 nm Verteporfin b100 nm Inactivated hepatitis A virus ~150 nm (strain RG-SB)

LUV LUV LUV (Virosomes)

EggPC/Cholesterol D/L wt ratio = 0.25 EggPG/DMPC D/L wt ratio = 0.13 DOPC/DOPE

8

Inflexal V

Influenza

Inactivated hemaglutinine of influenza virus strains A and B

~150 nm

LUV (Virosomes)

DOPC/DOPE

10–20 μm

MVL (DepoFoam) DOPC/DPPG/Cholesterol/Triolein, D/L wt ratio = 0.813 MVL (DepoFoam) DOPC/DPPG/Cholesterol/Triolein/ Tricalrylin, D/L wt ratio = 1.13

Liposome – Micron sized products 9 Depocyt Neoplastic and lymphomatous Cytarabine meningitis 10 DepoDur Pain management Morphine sulfate Lipid complex – Nano and Micron sized products 11 Abelcet Fungal 12 Amphotec Fungal

Amphotericin B Amphotericin B

10–20 μm

b5 μm ~130 nm

Disc like Ribbon like

Purpose

Suspension Suspension

To reduce toxicity To increase blood circulation/control release/ drug targeting Suspension To increase blood circulation/control release/ drug targeting Lyophilized powder To improve stability due to non-pegylation Lyophilized powder To increase solubilization and stability Suspension To induce immunity to variety of antigens without adverse effects associated with other adjuvants Suspension To induce immunity to variety of antigens without adverse effects associated with other adjuvants

Reference

[29] [23,30] [31] [30] [32,33] [34]

[35]

Suspension

To control release

[36]

Suspension

To control release

[26]

DMPC/DMPG, D/L wt ratio = 1.02 Suspension To improve drug loading Cholesteryl sulfate, D/L wt ratio = 1.89 Lyophilized powder To improve drug loading and stability



Table 1 List of liposomes/lipid based drug approved products in market.

[37] [37,38]

Reprinted (adapted) and modified from Natarajan et al. [39] with permission from American Chemical Society.

5

6


Hydrophilic head

Cholesterol

Aqueous Solution

Hydrophobic tails

Fig. 4. Pictorial depiction of a liposomal molecule and how the cholesterol moiety inserts into a liposome.

human cancer cell line SKOV-3 and MCF-7, suggesting the effectiveness of the drug loading in liposomal nanocarriers. Although, sustained release was achieved, mechanisms to explain this kind of release behaviour were not elucidated. In fact the only reported long-term release products using liposome (prior to our work with liposomal latanoprost) are DepoCyt®, and DepoDur®, both of which are based on the DepoFoam technology. Depocyt sustains the release of an anti-neoplastic agent (cytarabine) and is administered intra-thecally, whereas DepoDur is an epidural injection of morphine for sustained anaesthetic effect. The DepoFoam vehicle is called a Multi-Vesicle Liposome (MVL) and is composed of vesicular compartments within a membrane bilayer [57]. The size of MLV can range from 20 to 100 microns, so strictly speaking the sustained release is achieved only with a micron-sized liposomal particle, in contrast to liposomal latanoprost, where the average particle size is 100 nm. In summary, there have been very few nanoliposomal products that have been currently approved on the basis of sustained release benefits. Part of the reason is the difficulty of loading sufficient amount of bioactives in either the core or the bilayer without disruption of the self-assembled structure; the other main reason is the relative fluidity of the bilayer membrane which allows facile diffusion. Moreover, unless the aggregation issue is resolved for intra-vitreal injections of liposomal particles, they are unlikely to be used in the posterior cavity of the eye. 3. Solid lipid particles Solid lipid nanoparticles (SLNs) were developed as an alternative class of colloidal drug carriers in early 90’s for drug delivery applications. SLNs have shown considerable promise in the last decade for delivery of drugs (hydrophobic and hydrophilic) and proteins/peptides. SLNs were first described by Muller [58] and Gasco [59] in early nineties where the liquid lipid phase in oil-in-water emulsion (O/W) is replaced with a solid lipid phase. There are different methods available to prepare SLN as outlined in Pardeshi et al. [60]. High pressure homogenization and micro-emulsion based techniques have been routinely used to prepare SLNs.

Fig. 5. Schematic of the nanomedicine system for glaucoma, showing the stereographic insertion of the drug into the lipid bilayer. Reprinted (adapted) from Natarajan et al. [39] with permission from American Chemical Society.

Two different methods of drug/protein solubilization in the lipid matrix have been reported. Hot homogenization involves the use of an emulsified drug loaded lipid matrix (temperatures above the melting point of lipid) with the solution of surfactant (at similar temperatures) using a high speed stirrer. The primary emulsion is then subjected to high pressure homogenization; this leads to an inherent increase in the temperature of the lipid melt due to increased pressure and repeated cycles to obtain particle size in the sub-micron size range. Finally, the hot molten emulsion is cooled to room temperature to produce a solidified drug/protein dispersed lipid melt yielding SLN (re-partitioning of the drug from the aqueous phase to lipid phase, Fig. 7). There are several drawbacks using this method such as; (1) accelerated degradation of thermo-labile drugs and proteins (2) increased partitioning of the drug into the aqueous phase which eventually leads to drug leakage/expulsion on release and on storage (3) polymorphic transition of the lipids leading to instability of SLN. To overcome these drawbacks, cold homogenization technique was developed. In this technique, excess temperature is circumvented by solidifying the drug-lipid mixture emulsion in liquid nitrogen or dry ice to ensure homogenous dispersion of the drug particles in the matrix. The particles are then grounded to fine particles in the sub-micron range before dispersing it in a cold surfactant solution. This technique overcomes the limitation of melting of the lipid during homogenization and as well minimizes the partition of drugs into the aqueous phase leading to better stability on release and on storage. Another strategy of SLN preparation involves the use of microemulsions. The lipid usually fatty acids and/or triglycerides are melted and then a mixture of water, co-surfactant and the surfactant are added to the lipid melt at the same temperature under stirring to form a micro-emulsion. This is immediately followed by dispersion of the mixture in a cold aqueous phase (around 4 °C) under mixing conditions to precipitate the lipids in sub-micron size ranges to yield SLNs. Surfactants can include a wide range of bile salts, lecithins and some alcohols such as butanol [61]. 3.1. Drug/protein loading into SLN SLN possess similar advantages to that of liposomes and biodegradable nanoparticles. SLN have shown considerable promise as drug delivery carriers for controlled and sustained drug delivery in the last decade. SLN have been predominantly used to load lipophilic drugs and peptides/proteins with a relatively low success for loading and release of hydrophilic molecules. This could be attributed due to relatively lower volume of the aqueous phase for loading and surface localization of hydrophilic molecules leading to relatively higher burst and hence not suited for sustained delivery. One of the important parameters for sustained drug release over few weeks is its loading capacity and is dependant on various factors such as solubility, structural properties of drug and lipid melt, polymorphic state of the lipid and miscibility of the drug with the lipid melt [60,61]. A large number of lipophilic drugs have been evaluated for loading into solid lipid microparticles


J.V. Natarajan et al. / Journal of Controlled Release xxx (2014) xxx–xxx Table 2 Leakage half-lives of Doxorubicin (DXR) from liposomes. Lipid Composition

Caelyx (HSPC:CHOL:mPEG) HSPC:CHOL:mPEG DMPC:CHOL:mPEG POPC:CHOL:mPEG DOPC:CHOL:mPEG

Phase Transition Temperature of phosphatidylcholine (° C) 55 55 23 −2 −20

Ratio (mol)

n

t1/2 (h)

55:40:5

5

118.4 ± 18.8

2:1:0.1 2:1:0.1 2:1:0.1 2:1:0.1

5 4 2 2

91.8 ± 11.2 23.0 ± 6.4 14.6, 11.9 14.9, 10.2

Reprinted (adapted) from Charrois et al. [42] with permission from Elsevier.

and solid lipid nanoparticles. Loading capacities between 1–20% of drugs and macromolecules have been routinely achieved. Some studies have also reported higher loading capacities of about 25–50%, suggesting that SLN could be a useful drug delivery carrier for sustained release applications [61]. There are different models available for drug loading and release as described by Muller and colleagues (shown in Fig. 8). Three different models have been proposed based on the preparation strategies namely (i) drug-enriched core (ii) drug-enriched shell (iii) solid solution. In drug enriched core, the drug is precipitated inside the lipid matrix before the lipid could recrystallize during preparation. This usually happens when the drug is supersaturated during cooling of the lipid melt and eventual drug crystallization inside the lipid as a core with a lipid shell surrounding it. Drug-enriched shell is usually obtained during hot homogenization preparation when the drug partitions itself into the aqueous phase due to increased temperature inside the homogenizer. A solid solution of SLN is prepared by using cold homogenization technique and without any further addition of aqueous surfactants The main challenge with SLN is the drug leakage/expulsion on storage/release due to regular crystal lattice arrangements of the solid lipid (from a less ordered α-form to more stable β-form) and hence limited space availability for drug distribution [61]. To overcome this drawback, nanostructured lipid complex (NLC) was developed. NLC uses mixtures of lipid/triglycerides that differ in the structural properties leading to an imperfect crystal lattice that improves both loading (more area available for loading) and less drug leakage/expulsion during storage and controlled release [62–64]. With this as the background, let us now focus on the studies that have shown sustained release of bioactives with solid lipid nanoparticles and microparticles. 3.2. Drug delivery using solid lipid nanoparticles

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drugs loaded into SLN showed considerable promise in terms of loading and stability, there are very few studies that showed sustained release capabilities beyond three days. Majority of the studies reported, have shown significant burst release due to surface adsorption of the drug in the shell and eventually, loss of drug within a shorter timeframe as one of the major limiting factors. Mehnert et al. [65] studied the effect of loading (1, 5 and 10%) and release of three model drugs tetacaine, etomidate and prednisolone with different processing parameters such as size, surfactant concentration, temperature and on the nature of the lipid matrix from SLN. The etomidate and tetracaine drug were released within an hour with a relatively large burst while prednisolone release was sustained over a period of 5 weeks in vitro (Fig. 9). The possible explanation for faster release for etomidate and tetracaine was attributed due to larger surface area of nanoparticles and major drug distribution in the shell (drug-enriched shell model prepared by hot homogenization technique) leading to faster diffusion of the drug from the surface. On the contrary, prednisolone release was slower due to the differences in the melting points and nature of the lipids used (leading to increased drug-lipid interaction) and slower diffusion of the drug from the core to the external aqueous phase (drug-enriched core model) for release due to preparation using cold homogenization technique. This study clearly suggest that choosing appropriate drugs that have higher melting point than the lipid matrix and prepared by using cold homogenization technique could be used to sustain release of drugs for more than a month in vitro and how much of this release in vitro could be correlated with in vivo release is not available and needs to be evaluated.

3.3. Protein/peptide delivery using solid lipid micro and nano particles More recently, efforts to deliver large macromolecules using solid lipid microparticles (SLM) and solid lipid nanoparticles (SLN) have attracted much attention. A significant advantage includes improved protein loading by using milder preparation conditions (i.e. without using harsh organic solvents) and enhanced protein stability on storage and on release. A review by Almeida et al. [66] outlines the research focused on loading, release kinetics and stability of protein/peptide delivery using solid lipid particles. Much attention was also directed at the formulation aspects of protein encapsulation and stability of the protein/peptides; however very few studies have shown promise for sustained release applications. Many different classes of model proteins (bovine serum albumin [67,68], lysozyme [69]), antigens (hepatitis B [67]) and therapeutic peptides (calcitonin [70], cyclosporine A [71],

A large number of hydrophobic drugs have been loaded into SLN and drugs that have been loaded into SLN are outlined in the review article reported by Muller et al. [61]. Although, a significant proportion of the

Fig. 6. Dexamethasone release from 3 different liposomal preparations. Reprinted (adapted) from Kallinteri et al. [55] with permission from Elsevier.

Fig. 7. Partitioning effects of drug during preparation using hot homogenization technique. (i) Partitioning of the drug from the lipid phase to aqueous phase during increased temperature (ii) Re-partitioning of the drug to the lipid phase during cooling of the produced o/w nanoemulsions. Reprinted (adapted) from Muller et al. [61] with permission from Elsevier.


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evaluated, suggesting that IFN-α loaded SLN could be a viable formulation for controlled release in veterinary therapeutics. Although, slower release can be obtained using solid lipid nanoparticles, a major obstacle still remains is burst (almost 20–40%) for majority of macromolecules evaluated till date. 4. Biodegradable (solid) nanoparticles

Fig. 8. Drug incorporation models: solid solution; drug-enriched shell; and drug-enriched core. Reprinted (adapted) from Muller et al. [61] with permission from Elsevier.

insulin [72,73], and somatostatin [74]) have been evaluated for loading and release using solid lipid particles. There is always a challenge associated with the loading of hydrophilic bioactives due to the lipophilic matrix of SLN. This challenge was overcome in a study reported by Hu et al. [75] where a model peptide (gonadorelin) was incorporated into solid lipid nanoparticles (Zavg ~ 420 nm) using a solvent diffusion method in an aqueous system for oral administration. An incorporation efficiency of ~70% was achieved using this preparation strategy. The release was tested in simulative gastrointestinal fluid (SGF) and simulative intestinal fluid (SIF). A biphasic release pattern of burst followed by slow release (zero order) was observed in both release conditions. The only difference between the two release conditions was the burst, which was nearly twice in acidic conditions (SGF) when compared to neutral conditions (SIF). Almost 80% of the bioactive was released in a span of two weeks. Also, biological activity of the peptide (100%) was maintained on storage and release suggesting that SLN could be developed for sustained release applications for other kinds of macromolecules. In another study, Yuan et al. [76] compared the loading and release of a model hydrophilic peptide, Leuprolide (LR), from SLN using two different preparation strategies. In one strategy, LR was loaded using the solvent diffusion method in an aqueous system and was compared against the combination of hydrophobic ion pairing (HIP) and oil in oil (O/O) emulsion-evaporation methods. The loading efficiencies of LR and LR-sodium stearate (LR-SA-Na) was increased by ~ 1.5–2 fold by using O/O emulsion evaporation. The release of LR and LR-SA-Na was slow and sustained upto 4 days in vitro using (O/O) emulsion method; while the release was faster (within two days) using solvent diffusion system. Li et al. [77] developed a controlled release formulation of yak interferon alpha (IFN-α) for veterinary therapeutic application. In this study, IFN-α was loaded into SLN by double emulsion solvent evaporation method (w/o/w) and an encapsulation efficiency of ~83% was observed with a loading capacity of ~1.75%. A release pattern of burst followed by slow release was achieved over a period of 16 days in vitro and the released IFN-α also showed antiviral activity for the entire period

Fig. 9. In vitro release of prednisolone from cholesterol SLN and from Compritol SLN over a period of 5 weeks. Reprinted from Mehnert et al. [65] with permission from Elsevier.

These are the class of nanoparticles that are solid at 37 °C and composed mainly of polymers. These polymers generally have glass transitions temperatures (Tg) or melting points above 37 °C. PLGA is the most versatile polymer that has been evaluated extensively till date with relative success. This is partly due to the ability to modify and fine tune the physicochemical properties of the polymer such as the ratio of the lactide and glycolide in the copolymer, molecular weights etc. For example, changing the ratio from 50% of L-lactide to 85% Llactide can change the Tg between 45 °C and 70 °C. The polymers and the co-polymers are fully biodegradable and the degradation products are generally safe for human use. The degradation rates could also be manipulated by varying the lactide to glycolide ratio. In this section we will focus on PLGA nanoparticles as applied to cancer therapy. In solid nanoparticles, drug is either dissolved or dispersed in a polymer matrix and processed to a nanoparticulate form (other than by selfassembly). This is different from nanocapsules or otherwise core-shell nanoparticles, which consists of two different types of polymers or with lipids as a core and polymer as a shell. The drug could be dissolved either in core or in the shell depending on the types of polymers (hydrophobic/hydrophilic) or lipids used. Of the two, it appears that nanocapsules are difficult to fabricate and generally core-shell particles have micron sized dimensions. Hence, most of the discussion in the following sections relates to more of matrix types polymeric nanoparticles. In comparison to micelles and liposomes, polymeric nanoparticles are easier to manufacture with improved stability and better control over release of the drug; but the downside is the relatively difficulty in covalently linking hydrophilic moieties such as PEG or ligands to the particle surface. Also, polymeric microparticles have been in use for relatively longer periods of time, but nanoparticulate ones are rare. To our knowledge, there are no nanosystems based on solid nanoparticles that have been approved by the Food and Drug Administration (FDA). 4.1. Fabrication techniques In this section, various strategies used for the preparation of biodegradable polymeric nanoparticles will be briefly summarized for studies that have shown sustained release [78]. A number of articles [79,80] detailing the general synthesis process have been reviewed elsewhere for PLGA [81], polylactic acid (PLA) [82] and chitosan nanoparticles [83]. In general, different strategies are used to load bioactives either in the core of the nanoparticles or by adsorption onto the surface of the nanoparticles. For bioactives that are loaded inside the core are generally classified as nanocapsules and for bioactives that are adsorbed onto the surface of the nanoparticles are classified as nanospheres. Polymeric nanoparticles have been synthesized by various methods depending on the properties of the encapsulant and its intended use for delivery [78,84]. For example, generally PLGA nanoparticles have been prepared by either one of these techniques (i) solvent emulsion-evaporation [85], (ii) emulsion-diffusion [86], (iii) nanoprecipitation [87] (iv) interfacial deposition methods [84]. In solvent evaporation method, polymers are dissolved separately in a highly volatile organic solvent (chloroform, acetone, dichloromethane etc) and added to an aqueous phase containing an emulsifier/stabilizer (usually polyvinyl alcohol) and sonicated to prepare nanoparticles. In emulsion-diffusion method, PLGA polymers are dissolved in an organic solvent (methyl ethyl ketone, ethyl acetate etc) and then mixed with an aqueous phase containing the stabilizer and emulsified using a homogenizer. In interfacial deposition, polymers



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spite of several years of research after the approval of first PLGA microsphere formulation (Zoladex) in 1989. The reasons for this are two-fold: (i) Stability issues associated with aggregation over time, and lyophilization and reconstitution does not always produce particles in the same size range; and (ii) increased drug loading is possible, however release is retarded due to diffusional barriers. This retardation is not useful for passive targeting, as slow release might lead to sub-therapeutic levels of drug with minimal cytotoxic effects in tumours. In a typical example [102], PLGA nanoparticles were decorated with PEG by mixing with PLGA-PEG block copolymer in an organic solvent containing the drug (paclitaxel). Subsequently, the particles are either prepared by emulsion techniques or by nanoprecipitation to produce PLGA-PEG nanoparticles in the range of 110–190 nm by diameter. These particle release paclitaxel partly via a burst followed by slow release, see Fig. 10. The burst release is unhelpful in vivo, if considerable amounts are released prior to extravasation into tumour tissue. The authors do show improved tumour volume reduction for the PLGA nanoparticles with PTX, compared to paclitaxel by itself; no comparison was made to Doxil® in the study.

are dissolved in the water miscible organic solvent with the aqueous phase and nanoparticles are prepared by centrifugation of the mixture. Nanoprecipitation method is more commonly accepted and reproducible method of preparation of PLGA nanoparticles. In this method, polymer is dissolved in acetone and added drop-wise into a continuously stirring aqueous phase containing an emulsifier/stabilizer with solvent evaporated under reduced pressure. A number of drugs have been loaded into PLGA nanoparticles and have shown considerable promise for sustained release. In a study by Mu et al. [88], PLGA nanoparticles have been synthesized using solvent evaporation technique using vitamin E TPGS as an emulsifier and loaded with taxol. A slow and sustained release of the drug upto 20 days in vitro was achieved. Diffusion of the drug with polymer swelling and erosion were the reasons behind slow and sustained release of the drug. A similar solvent evaporation technique was followed to load 2-aminochrome into PLGA nanoparticles surface modified with didodecyldimethylammonium bromide (DMAB). A slow and sustained release of the drug up to two weeks was reported by Labhasetwar et al. [89]. In another study by Derakhshandeh et al. [90], 9-Nitrocamptothecin (9-NC) was loaded into PLGA nanoparticles by nanoprecipitation method. A slow and sustained release of the drug upto a week was reported and the release mechanism was mediated by diffusion. A more advanced preparation technique involves the use of a mixed solvent system which is then emulsified into an aqueous medium containing the surfactant. This technique is termed as spontaneous emulsification solvent diffusion method (SEDS) [91]. The polymer phase consists of a water-miscible solvent and a low molecular weight PVA surfactant is used to minimize agglomeration of the particles. A particle size range of 100–200 nm is achieved and remains stable even after freeze-drying. Such nanoparticles have been predominantly used for loading of hydrophobic drugs [92]; while proteins [93] and even plasmid deoxyribonucleic acid (DNA) [94] have also been evaluated. As mentioned earlier, such nanoparticles may again be surface modified: but it requires special techniques to do so. PLGA nanoparticles can also be decorated on the surface with PEG and PEG-like molecules including the use of a surfactant molecule, PEG-PLA-PEG [95]. PEG grafting to already formed nanoparticles is seldom reported, while “chemically-bound” PEG nanoparticles rely instead on copolymerization with PLA [96] or PLGA [97] or even poly cyanoacrylate [98]. PLA nanoparticles have been prepared using solvent diffusion, solvent displacement and salting out method. In a study, Savoxepine was loaded into PLA nanoparticles that were surface modified with PEG by salting out method. This method involved the addition of salting out agents such as calcium chloride or magnesium chloride to separate the water miscible solvent from aqueous solution. This method is mostly suited for loading of proteins. A slow and sustained release of the drug was reported over one week with sustained plasma levels after intramuscular and intravenous injection [99,100]. Poly-ε-caprolactone (PCL) polymeric nanoparticles have also been studied for sustained release and prepared using methods similar to PLGA. In a study by Prabu et al. [101], vinblastine was loaded into PCL nanoparticles by emulsion technique. A slow and sustained release of the drug up to 20 days was reported with efficient uptake of the nanoparticles with the drug in a breast cancer cell line (MCF-7). From these studies, we can conclude that by carefully choosing the drug and the preparation strategy, more products based on polymeric nanoparticles can be expected to be successful for sustained release applications in future.

Due to lack of success of PLGA nanoparticles in passive targeting, most of the research work has concentrated on an alternative “active” targeting approach. In this approach, ligands (small molecular weight compounds, protein fragments) are covalently linked to the surface of nanoparticles that will specifically bind to receptors or molecules that are unique to cancer cells, carrying with it the cytotoxic payload. Of the ligands used, folate has been studied extensively due to the overexpression of the folate receptors in a variety of tumor cells, particularly in epithelial cancers such as breast cancer, nasopharyngeal cancer and endometrial cancer. In this strategy, folate ligands are covalently attached to the PEG molecules with a spacer. An example of this approach is to be found in Liang et al. [103]. Liang et al. synthesized a new terpolymer which consists of blocks of PLAGA, PEG and folate (Fig. 11). Folate ligands are not completely attached leaving behind free PEGs on the surface that leads to some “stealthiness”. An emulsion technique was used to prepare nanoparticles and the size was about 220 nm in diameter with a surface charge of − 8 mV with the added folate. Folate conjugated particles showed enhanced internalization of particles in human carcinoma cell lines, when compared to control PLGA-PEG particles (no folate) suggesting that internalization was mediated by folate receptor binding to cancer cells. Paclitaxel loaded nanoparticles injected via the tail vein in a mouse bearing endometrial cancer cells, showed greater tumour volume reduction for folate targeted PLGA nanoparticles. The drug release profile here appears more acceptable than in the case of the PLGA blended with PLGA-PEG or PCL-PEG (compare Fig. 12). Although, numerous studies have shown the superior animal data with targeted nanoparticles, none have been translated to have beneficial outcome in humans. It does appear that combined benefits provided by “stealthiness” and targeting in the design of nanoparticles does not provide any advantage to longer circulation half-lives and enhanced targeting/penetration in cancer therapy. So, the best solution appears to be local administration of the nanoparticles rather than intravenous administration.

4.2. Passive targeting of PLGA nanoparticles

4.4. Cardiovascular applications

In addition to liposome and micelles, core-shell nanoparticles and solid lipid nanoparticles have been studied for targeting. In this section, we briefly summarize on studies that have traditionally used PLGA nanoparticles for passive and active targeting of bioactives. PLGA nanoparticles have not been successful for passive targeting to tumours, in

Apart from the mentioned applications, nano-size becomes vital when delivering into cells. These include the delivery of low molecular weight drugs (for anti-restenosis or anti-thrombotic effects) or other bioactives such as siRNA or plasmid DNA. Here the long circulation lifetime is no longer a predominant requirement, as the mode of

4.3. Active targeting of PLGA nanoparticles


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administration is localized to the site of action, by various means, including the use of infusors. One important use of this concept is treating thrombus formation and restenosis following angioplasty. Several infusors, such as Dispatch® and Infiltrator®, have been approved by the FDA for catheterbased drug delivery. The Dispatch® device, for example, was given a 501 K approval in 1996 for delivery of anti-thrombotic agents such as heparin and urokinase following angioplasty. Since then the device has been used in a variety of preclinical and clinical studies. The interesting feature of this indwelling catheter is that it is designed to deliver drug intra-arterially for up to 4 hours without blocking blood flow. This makes it an ideal delivery mechanism for various bioactive molecules from drugs to proteins to plasmid DNA. Since the infused drug usually is in a liquid (either suspended or dissolved), the delivery of nanocrystalline drugs and particles has evident advantages including ease of delivery, without the agglomeration effects or rapid clearance that usually accompanies intravenous administration. It appears that sustained and localized delivery is possible, but surface modification of particles has emerged as a critical factor for cellular uptake. Song et al. have developed an ex vivo model for measuring arterial uptake, using an explanted canine femoral artery perfused with the drug solution for 30 seconds at 37 °C [104]. Labhasetwar and co-workers investigated surface modified PLGA nanoparticles tailored from chemical and physical means [89]. They found that the formulation with cationic detergent (didodecyldimethylammonium bromide, DMAB) adsorbed at 5% onto the PLGA particle yielded the highest arterial uptake, which was approximately 40 times that of the unmodified PLGA particle. Other cationic adsorbents studied also increased arterial uptake and the sizes of these particles were reported to be about 100 nm. Some of these findings were translated into positive results in a rat study, to confirm deep-tissue delivery of dexamethasone incorporated in PLGA particles [105]. An artery injury model was used in this study, with “local” infusion of the drug-containing nanoparticles, over a 3minute period in the carotid artery. The local infusion is accomplished by closing off the arterial segment and creating a closed arterial space where the drug suspension is infused. The study demonstrated that these nanoparticles (without any surface modification) were able to penetrate the luminal, the medial and the adventitial layers, with sufficient dexamethasone being delivered. It is claimed that the infused drug inside the nanoparticles is able to reduce stenosis by 31% compared to intraperitoneally delivered dexamethasone, but the caveat is that this is a non-conventional animal model for restenosis, and such studies need to be extended to the standard porcine models. To date, there is no report of an approved delivery system for restenosis using this approach. We suspect that the concept has been superseded by the emergence of the drug-eluting stents, since stenting is widely accepted for arterial stenosis to overcome not just stenosis but also to accommodate vessel recoil which cannot be done via the infusor concept. The localized delivery of two anti-proliferative drugs, rapamycin [106] and paclitaxel [107] has received considerable success in overcoming approximately 20–30% restenosis caused by the use of bare metal stents. Although other problems have surfaced with the use of drug-eluting stents, notably late-stage thrombosis, these are being addressed with the use of fully biodegradable stents [108] and another concept, drug-eluting balloons. However, it is possible that the use of nanoparticles in localized delivery will see other applications, including siRNA and plasmid DNA delivery, which is beyond the scope of this review, but are nevertheless critical applications of nanocarriers with the capacity for sustained and localized release.

Fig. 10. Cumulative paclitaxel release from PLGA nanoparticles blended with PLGA-PEG and PCL-PEG copolymers. Note the high initial burst followed by slow release. It is likely that the initial burst releases substantial paclitaxel amounts into the blood in vivo before extravasation. Reprinted (adapted) from Danhier et al. [102] with permission from Elsevier.

they form colloidal polyelectrolyte complexes or multilayers of polyelectrolyte when complexed with an underlying layer of the opposite charge, the process being known as polyelectrolyte layer-by-layer (LBL) assembly and the resulting materials are termed polyelectrolyte multilayers (PEM). In recent decades, PEMs have been explored for drug delivery. Studied as thin films on substrates and shells on particles, recent work has showed that they are responsive to various stimuli [109], useful for functionalization of nanoparticles [110] and cell-targeting purposes [111], as well as being capable of being built into composite architectures [112], all done in aqueous solution. Due to the interest in this system, in 2013 alone there were 17 review papers indexed by the ISI on polyelectrolyte applications in drug delivery. Despite the huge interest in LBL technology, the number of works that achieved sustained release with LBL assembly, however, remains low. Within the last 5 years, very few PEMs on nanoparticles achieved sustained release for more than 3 days. We have tabulated the literature survey in Table 3. The platform in general is divided into polyelectrolyte complex, LBL assembly, or nanocapsules. A polyelectrolyte complex is a self-assembled structure formed by a simple admixture of 2 oppositely charged polyelectrolyte solutions, whereas LBL assembly requires build-up of the PEM shell on a core particle by either repeated dispersion in each polyelectrolyte solution or alternating saturation of the solution by each polyelectrolyte. When the core particle content is dissolved, it converts into nanocapsules. A polyelectrolyte complex, therefore, has no underlying core, and the resulting complex itself is the nanoparticle of interest. LBL assembly, on the other hand, has cores that may vary, such as PLGA nanoparticles, liposomes, oil emulsion, or mesoporous silica nanoparticles, and the assembly itself may consist of differing polyelectrolytes, or terminated by a distinct polyelectrolyte layer for certain desired surface properties. To-date, platforms that exhibit successful sustained release with the use of polyelectrolyte are PLGA nanoparticles [120–123], liposomes [113–115], oil emulsion [117–119], mesoporous silica nanoparticles [116], single component nanocapsules [124,125], as well as polyelectrolyte complex [126–133]. These exhibit sustained release for at least 3 days, and have diameters below 500 nm. 5.1. Nanoparticle-LBL assembly

5. Polyelectrolyte nanoparticles Polyelectrolytes are charged polymer chains with the capability of forming an electrostatic complex with other polyelectrolytes of the opposite charge. When complexed with other polyelectrolyte in solution,

For PLGA nanoparticles, several low molecular weight (MW) drugs/ model drugs have been employed for sustained release: fluorescent probes such as fluorescein, rhodamine 6G, carboxyfluorescein, and anti-cancer drugs such as docetaxel and doxorubicin [120,123]. One



Fig. 11. Synthesis scheme for PLGA particles incorporating both active and passive targeting capabilities, using the folate molecule as a targeting moiety. Reprinted (adapted) from Liang et al. [103] with permission from Elsevier.

notable study done in mice, more specifically, shows that PLGA-core PEMs are capable of both sustained release and imaging (theranostics) of both the cargo and the carrier [123]. The cargo was a model drug cardiogreen for doxorubicin. The LBL assembly here has a two-fold function: for particle fluorescence trafficking and to determine the biodistribution. The former was done through a near-IR-labelled polyL-Lysine (PLL) as a counter poly-ion of dextran sulfate (DES) in the PEM, whereas the latter was accomplished by adding a terminal layer such as hyaluronic acid (HA) or alginate (ALG), reducing significantly the accumulation of the carrier particles, and therefore the drug, in mouse liver and spleen. The PEM, although primarily used primarily for biodistribution purposes, also benefits the release kinetics by retarding the release. However, in this study, there was no report of

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the odd/even effect of the number of layers on the release, contrary to another reported work on similarly low MW drug [120]. Our lab has contributed to studies of this platform by establishing that different polyelectrolyte types can not only affect the release retardation by PEM, but the amount of release suppressed increases with particle size, indicating higher efficacy of PEM as diffusional barrier in nano-sized particles than micron-sized ones [121]. This was done with BSA as the delivery cargo, and could achieve 30-day sustained release. With FITC-dextran (4 kDa), the release duration was much shorter, 4 days [122]. A similar strategy of incorporating agents in the core particle and in the PEM was also attempted on liposomes [115]. The drug was doxorubicin, encapsulated in the liposome, and the additional cargo siRNA in the PEM was also released, forming co-delivery of drugs for antitumor therapy. Although siRNA was released from the multilayer with poly-L-arginine, the release of doxorubicin was still sustained by the additional diffusional PEM barrier, retarding the release from 100% to 50% within 3 days. The result was a significantly reduced tumor volume, compared to free doxorubicin and free siRNA. Higher MW model cargo have also been encapsulated in liposomes with LBL assembly, particularly BSA [113] and glucose [114]. For BSA, the release was much suppressed, from 100% to 30% over 30 days using 5 alternating layers of chitosan (CHI) and alginate (ALG), whereas glucose release was slowed from 37% in 6 days to 0.5% with 2 layers of chitosan/DES. Another platform that has exhibited sustained release with LBL assembly is an oil core. Bazylinska et al. [117–119] managed to retard the release from oil emulsion, from 80% at 14 days to 20%, using an LBL assembly. Unlike other core materials, however, stabilization of the oil cores with dicephalic-type surfactant is necessary for the anchoring of the PEM. The retardation of release with LBL assembly is not always achieved, particularly when loading of the particles is done after LBL assembly. This is exemplified by a study of mesoporous nanoparticles (MSN) releasing doxorubicin [116]. The release increases with the number of layers up to 20; however, the release of doxorubicin is not complete and remains at 20% of the total drug loading, indicating that the drug that is released is from a closed PEM shell, rather than from the core MSN. The release period however, could be sustained to potentially more than 5 days, and pH-sensitive due to the closing/opening of the PEM shell at different pH values. In addition, it is claimed that the PEM also helps in reducing uptake by macrophages. A final category of PEMs is that of single component nanocapsules [124,125]. This utilizes a single polyelectrolyte layer, such as CHI, PLL, or polyglutamic acid (PGA), and cross-linking of the shell, followed by further dissolution of the core particle, producing capsules which are able to sustain the release of incorporated drug: for example it has been shown that ibuprofen release can be sustained over for 3 days from chitosan nanocapsules [124]. When the template has a solid core but with a mesoporous periphery, a very thick shell could be produced in the order of 30 nm. This is in excess of a typically 2 nm single polyelectrolyte layer thickness [125]. Encapsulating doxorubicin, such cross-linked PGA capsules were able to completely suppress the release of doxorubicin, and shows similar efficacy in reduced human colorectal tumor cell viability to free doxorubicin, but with expected reduced systemic toxicity. 5.2. Polyelectrolyte complex nanoparticles

Fig. 12. Paclitaxel release profile from PLGA-PEG nanoparticles. Reprinted (adapted) from Liang et al. [103] with permission from Elsevier.

The concept of sustained delivery with polyelectrolyte complexes have been directed towards diverse applications. For instance in wound healing, a PEI/DES complex was formed that encapsulates FGF10 [126]. These complex particles were reported to yield higher proliferation of human umbilical vein endothelial cells (HUVEC) than free FGF-10, with a sustained delivery of 11 days. For bone tissue regeneration, CHI/chondroitin sulfate complex could release BSA for 1 month with a 40% of burst release, followed by slower rate of release [128].


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Concurrently, higher osteogenic differentiation than the control was demonstrated. For cancer treatment, the bacterial exopolysaccharide, mauran, could be complexed with CHI, delivering 5-fluorouracil for 12 days and showed a more gradual cytotoxic effect compared to free drug [131]. The architecture may also be more complex. Self-assembled particles may be formed through one polyelectrolyte, such as the case of a hydrophobically modified chitosan, which is then complexed with alginate polyelectrolyte [130]. In such a case the cluster formed would be multiple chitosan particles entangled with alginate chains, depending on the stoichiometry. Carrying BSA, this complex could sustain the release for up to 30 days. The initial constituting polyelectrolyte could be modified for specific purposes that bestow similar features on the resulting complex, such as with the creation of amphiphilic polyelectrolyte consisting of a carbohydrate backbone with hydrophobic poly-L-lactide and anionic succinyl group, complexed with the cationic protein etanercept [132]. Above the clouding temperature of the polyelectrolyte, the complex exhibits hydrophobic interaction at body temperature, increasing stability of the particles in the body. Another example of this is by the use of antibacterial, mucoadhesive, water-soluble chitosan derivative as the complexing polyelectrolyte with the drug. The resulting complex could release diclofenac sodium, a model NSAID, for 3 days [127].

6. Non-spherical nanoparticles and carbon nanotubes (CNT) 6.1. Non-spherical nanoparticles Although this review is predominantly focused on lipid and polymer based nanospheres, it is important to note that the geometrical effects of particle can also directly impact drug delivery [134]. Although the precise role of particle shape has not been fully elucidated, some studies have identified degradation as the dominant factor, where water penetration and degradation do not occur homogenously unlike in spherical particles. Apart from providing variable release, the use of non-spherical nanoparticles may confer additional benefits such as targeting ability and internalization. The red blood cell (RBC) is a good example to demonstrate the importance of shape and flexibility; in the spleen, only spherical particles less than 200 nm can pass through however RBC of diameter 10 μm can also pass through [135]. Hence, in this section, we would like to highlight the usefulness of non-spherical nanoparticles in drug delivery. For example, it has been found that the shapes of particles minimize phagocytosis [136], thus leading to longer blood lifetimes. In the study of ellipsoid particles, Champion et al. [136] found that macrophages initially attached to the pointed end of particles took only a few minutes to internalize the particle, while macrophages that were attached to the flat regions took over 12 hours to internalize the entire particle. Compared to spherical particles that may be internalized rapidly from any point due to its symmetry, these ellipsoid particles may have an advantage. However, it should be noted that particle volume is also a critical factor in phagocytosis and further studies are required to elucidate the importance of shape and size in phagocytosis. Abidian and co-workers exploited conducting polymer nanotubes for precise triggered release of drugs [137]. The control of release of dexamethasone (DEX) was effected using external electrical stimulation on poly (3,4-ethylenedioxythiophene) (PEDOT) nanotubes, as shown in Fig. 13G. In vitro, they were able to achieve several complex release profiles of dexamethasone using an applied electrical field. The mechanism lies in the contraction and expansion of the nanotubes during electric stimulation. This mechanical shrinkage causes DEX to be eluted from the ends or from the fine cracks within the nanotubes. This concept could be used to release drugs at desired point in time using electrical stimulation of the nanotube. The authors reported that this approach could be useful for bio-applications in stimulation of neurite outgrowth

and guidance for neural tissue regeneration, but its use in vivo presents many challenges, similar to iontophoretic delivery of drugs. 6.2. Carbon nanotubes (CNT) Although carbon nanotubes (CNT) are non-polymeric nanocarriers, we would like to briefly highlight some progress made in the use of CNTs in drug delivery. Despite a number of publications on the use of CNT as drug carriers, the majority of the studies are focused only on drug targeting and very few studies have shown sustained release capabilities. Liu and co-workers presented the use of single-walled carbon nanotubes (SWNT) to deliver paclitaxel (PTX) in mice for anticancer therapy [138]. The PTX release from SWNT-PTX was performed in PBS and mouse serum for 48 hours. At the 48 hour time-point, approximately 20% of PTX was released into PBS while almost twice the amount was released in mouse serum at 37 °C. Although in-vitro release was only studied for 48 hours, the authors have shown more sustained effects in vivo (Fig. 14). Tumor inoculated mice were used and efficacy of the formulation was evaluated based on reduction of the tumor size. The authors presented the tumor size as a tumor growth inhibition (TGI) index, where a higher value equates to more inhibition and better effectiveness, relative to untreated mice. At the 22-day timepoint, the TGI for pure PTX alone was 27.7% while TGI for SWNT-PTX was 59.4% and significant reduction in tumor growth was obtained. The sustained inhibition of tumor growth could be attributed due to the increase in circulation half-lives of PTX in blood from 18.8 min ± 1.5 to 81.4 ± 7.4 min in SWNT-PTX. The authors also evaluated the biodistribution of PTX compared to SWNT-PTX in tumor and other organs. They found significantly higher PTX uptake in tumor 2 hours after injection, where SWNT-PTX group exhibited 10-fold and 6-fold higher uptake than the PTX formulation 2 h and 24 h after injection respectively; demonstrating higher drug delivery efficacy in SWNT-PTX. Also, they reported high tumor-to-normal tissue PTX uptake ratio, thus making SWNT-PTX more favourable for suppression of tumor at lower doses. In another work by Abidian et al., loading of doxorubicin (DOX) into the sidewall surface area of PEGylated SWNTs via noncovalent binding was demonstrated [139].This work showed that SWNT can provide active drug release based on pH effects between DOX and SWNTs. At neutral or alkaline, minimal release of DOX was observed, while at acidic pH 5.5, SWNTs-DOX released about 50% of DOX in 2 days. There is a wealth of information regarding the use of CNTs in drug delivery; however experimental evidence suggests conflicting results on the toxicity of CNTs at cellular levels and in animals. This could be attributed to lack of standardized methods available to evaluate nanotoxicity. The impact of different parameters such as size, shape, length, agglomeration, delivery routes etc should be systematically evaluated for a thorough understanding of the key factors that can contribute to CNT toxicity [140]. If the specific safety and toxicity concerns of CNTs can be identified and addressed, then products based on CNTs for use in humans could be a reality in the future. 7. Summary and prognosis In comparison to approved microsphere formulations, the number of nanocarrier-based therapeutic products is still low, and is mostly confined to the area of passive targeting of tumour tissue. Tumor-targeted delivery does not require sustained release of drug beyond a few days. In general, it is more difficult to achieve sustained delivery over weeks using nanocarriers than with micron-sized carriers. With the demonstration of sustained (long-term) release from nanoliposomes and from other nanoparticles, it is possible now to extend the applications to other disease conditions. In particular, there is demonstrated superiority for use of sustained-release nanocarriers in therapy of ocular


Ref.

Delivery platform

Applications

Cargo

Polyelectrolyte

Size (nm)

Delivery period (days)

In vitro release setup

[113] [114]

Liposome-LBL assembly Liposome-LBL assembly

Biomacromolecule delivery Hollow nanocapsules

CHI/ALG CHI/DES or DNA

380 100–170

30 suppressed

Centrifugation, unspecified buffer Dialysis, unspecified buffer

[115]

Liposome-LBL assembly

poly-L-arginine/siRNA and final HA layer

120

3

Dialysis, PBS

[116] [117] [118]

MSN-LBL assembly Oil core-LBL assembly Oil core-LBL assembly

Doxorubicin IR-786, IR-780, Oil Red O IR-786

PAH/PSS PDADMAC/PSS PSS/PDADMAC or PLL

260 70–210 100

4 8 8

Dialysis, acetate buffer/phosphate buffer Dialysis, PBS Dialysis, PBS

[119] [120]

Oil core-LBL assembly PLGA-LBL assembly

Multiple therapeutics, treatment of triple-negative breast cancer Improving MSN-based carrier systems Innovative drug nanocarrier DDS that combine biocompatibility and stimuli-responsiveness Innovative drug nanocarrier Sustained delivery in bloodstream

BSA 1-hydroxy pyrene-3,6,8-trisulfonic acid (HPTS), alendronate, and glucose, Doxorubicin, SiRNA

λ-carrageenan/PDADMAC PEI/PAA

100 280

9 30

Dialysis, PBS Centrifugation, PBS

[121] [122] [123]

PLGA-LBL assembly PLGA-LBL assembly PLGA-LBL assembly

Release retardation Surface functionalization Theranostics

IR-786 Rhodamine 6G or Fluorescein or 5(6)-Carboxyfluorescein BSA FITC-dextran Doxorubicin or Cardiogreen

300 300 180–250

30 4 3

Centrifugation, PBS Centrifugation, NaCl Dialysis, PBS

[124]

Novel nanocapsule fabrication method Novel nanocapsule platform

Ibuprofen

190–250

3

Dialysis, PBS

Doxorubicin

PLL or PGA

270–420

Wound healing Mucoadhesive polymeric platform

FGF-10 Diclofenac sodium

170–300 40–90

Centrifugation, pH 7 PBS Dialysis, PBS

[128] [129] [130]

Polyelectrolyte complex Polyelectrolyte complex Polyelectrolyte complex

Bone tissue engineering Oral protein/peptide administration Sustained delivery in bloodstream

BSA or Platelet lysates Superoxide dismutase BSA

250 110–290 350

28 3 30

Centrifugation, phosphate buffer Dialysis, unspecified buffer Centrifugation, PBS

[131] [132]

Polyelectrolyte complex Polyelectrolyte complex

Cancer treatment Rheumatoid arthritis treatment

5FU Etanercept

PLL or CHI or PEI/DES Water-soluble CHI derivative/diclofenac sodium CHI/chondroitin sulphate Superoxide dismutase/DES-cholic acid Deoxycholic acid hydrophobically modified chitosan coated with sodium alginate Mauran/CHI Succinylated pullulan-g-oligo(L-lactide)/ etanercept

suppressed without degradation 11 3

Unspecified method, hydrolase in pH 5.8 PBS

[126] [127]

Single component nanocapsules Single component nanocapsules Polyelectrolyte complex Polyelectrolyte complex

PAH/PSS, PLL/DES PAH/PSS, PLL/DES PLGA coated with PLL/DES, terminated with DES, HA, or ALG CHI

30–200 250–450

10–12 Bioactivity up to 18 d

Dialysis, phosphate buffer/PBS Unspecified, PBS

[125]


PAH: polyallyamine hydrochloride; PSS: polysulfonate sodium; PDADMAC: polydiallyldimethylammonium chloride; PEI: polyethyleneimine; PAA: polyacrylic acid.

13


Table 3 Sustained delivery platforms with PEM and polyelectrolyte complex.

14


Fig. 13. G: Cumulative mass release of dexamethasone from: PLGA nanoscale fibers (black squares). Cumulative mass release of dexamethasone from: PLGA nanoscale fibers (black squares), PEDOT-coated PLGA nanoscale fibers (red circles) without electrical stimulation, and PEDOT-coated PLGA nanoscale fibers with electrical stimulation of 1 V applied at the five specific times indicated by the circled data points (blue triangles). Reprinted (Adapted) from Abidian et al. [137] with permission from Wiley online library.

diseases such as glaucoma; it is anticipated that other diseases will now be treatable with nanocarriers, provided the benefits over micro carriers is demonstrated unequivocally. The benefits specifically include improvements in patient compliance, cost reduction and a new method

of treatment of the disease. It is our belief that where sustained release is needed concurrently with intra-cellular delivery of the drug cargo, nanocarriers will play a pivotal role: one such area is the delivery of siRNA. Acknowledgements The authors would like to acknowledge the support from School of Materials Science and Engineering, Nanyang Technological University, as well as the NTU-Northwestern Nanomedicine Institute. References

Fig. 14. Nanotube PTX delivery suppresses tumor growth of 4 T1 breast cancer mice model. Tumor growth curves of 4 T1 tumor-bearing mice that received different treatments indicated. The same PTX dose (5 mg/kg) was injected (on days 0, 6, 12, and 18, marked by arrows) for Taxol, PEG-PTX, DSEP-PEG-PTX, and SWNT-PTX. Reprinted (Adapted) from Liu et al. [138] with permission from AACR.

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Controlled Drug Release from Pharmaceutical Nanocarriers.

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Polymer nanocarriers for dentin adhesion.

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Food antimicrobials nanocarriers.

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Controlled co-delivery nanocarriers based on mixed micelles formed from cyclodextrin-conjugated and cross-linked copolymers.

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Nanocarriers for cancer-targeted drug delivery.

siRNA Delivery by Stimuli-Sensitive Nanocarriers.

Mucus barrier-triggered disassembly of siRNA nanocarriers.

Photoresponsive polymer nanocarriers with multifunctional cargo.

Carbohydrate-Based Nanocarriers Exhibiting Specific Cell Targeting with Minimum Influence from the Protein Corona.

Ligation strategies for targeting liposomal nanocarriers.

Genetically engineered nanocarriers for drug delivery.

Lipid-Based Nanocarriers for RNA Delivery.

Buckysomes: New Nanocarriers for Anticancer Drugs.

Polymer-based nanocarriers for vaginal drug delivery.