Engineered PLGA nano- and micro-carriers for pulmonary delivery: challenges and promises

Authors


Francesca Ungaro, Department of Pharmaceutical and Toxicological Chemistry, University of Naples Federico II, Via D. Montesano 49, 80131 Naples, Italy. E-mail: ungaro@unina.it

Abstract

Objectives  The aim of this review is to summarize the current state-of-the-art in poly(lactic-co-glycolic acid) (PLGA) carriers for inhalation. It presents the rational of use, the potential and the recent advances in developing PLGA microparticles and nanoparticles for pulmonary delivery. The most promising particle engineering strategies are discussed, highlighting the advantages along with the major challenges for researchers working in this field.

Key findings  Biodegradable polymer carriers, such as PLGA particles, may permit effective protection and long-term delivery of the inhaled drug and, when adequately engineered, its efficient transport to the target. The carrier can be designed for inhalation on the basis of several strategies through the adequate combination of available particle technologies and excipients. In so doing, the properties of PLGA particles can be finely tuned at micro-size and nano-size level to fulfill specific therapeutic needs. This means not only to realize optimal in vitro/in vivo lung deposition of the formulation, which is still crucial, but also to control the fate of the drug in the lung after particle landing.

Summary  Although many challenges still exist, PLGA carriers may be highly beneficial and present a new scenario for patients suffering from chronic lung diseases and for pharmaceutical companies working to develop novel inhaled products.

Introduction

Drug inhalation has a long and rich history that can be traced back to the writings of many ancient civilizations, already employing the pulmonary route of administration to deliver natural remedies either locally or to the body through the lung. If in the past inhaling ‘vapours’ (Egyptians in 1500) or smoking cigarettes (Potter's cigarettes in 1802) was sufficient, with the rapidly growing popularity and sophistication of therapies, there is an increasing demand for tailor-made inhalable drug formulations able to yield the best therapeutic outcomes via the most efficient delivery to the lungs.[1]

It is well established that key features of the formulation (e.g., aerodynamic size, flow and aerosolization properties) will affect the likelihood of it being deposited in the desired region of the lung.[2] In the case of inhaled drugs required for the treatment of lung diseases, such as asthma, chronic obstructive pulmonary disease (COPD) and cystic fibrosis (CF), an effective ‘local’ delivery (i.e. centrally rather than at the periphery) of the drug is necessary.[3,4] Strategies for targeting drugs to the alveoli have become increasingly important, as many inhaled drugs are developed either to be absorbed systemically (e.g. peptides and proteins)[5] or to exert their pharmacological effect in this region of the lung (e.g. antitubercular drugs directed to alveolar macrophages).[6]

Although particle ‘landing’ at the site of interest remains crucial in determining the therapeutic efficacy of inhaled therapeutics, lung retention and the ability to overcome extracellular and cellular lung barriers is just as significant.[7] Mucociliary clearance, designed by evolution to eliminate inhaled and possibly noxious material from the airways, may considerably limit the residence time of the deposited formulation and, subsequently, the amount of drug reaching the target. Analogously, the nature and extent of the interactions of the drug with lung lining fluid (LLF), airway macrophages and lung epithelial cells will inevitably affect drug permanence in situ and duration of the effect.

To overcome these drawbacks, controlled and targeted delivery of drugs to the lungs by means of polymeric drug delivery systems (DDS) is under investigation.[1,8] As a matter of fact, DDS may offer substantial advantages over conventional dosage forms for inhalation, including long-term maintenance of drug concentrations within a desired therapeutic range, reduced doses, reduced dosing frequency, the potential for limited side effects and, last but not least, improved patient adherence to the therapy. This would represent a real benefit for those patients suffering from chronic lung diseases.

Despite encouraging premises, the development of safe and effective polymeric DDS for inhalation faces many challenges, among which the choice of the right polymer excipient[9] and its adequate engineering in the form of inhalable particles are primary.[10] In the 1990s, Edwards and colleagues developed a new type of inhalation aerosol based on poly(lactic-co-glycolic acid) (PLGA), characterized by particles of small mass density and large size, which permitted efficient and prolonged delivery of insulin into the systemic circulation.[11] This was the first attempt to apply PLGA-based carriers to lung delivery. After almost 10 years, we are experiencing the revival of the research interest in biodegradable particles for drug inhalation (Figure 1). Nonetheless, it is apparent from the current ‘manuscript boom’ that it is growing and presents a future challenge.

Figure 1.

Research studies on poly(lactic-co-glycolic acid) carriers for pulmonary delivery identified in PubMed databases from 1997 to 2011.

Excellent reviews have recently summarized the new concept of carrier-mediated lung targeting,[12,13] and the most interesting techniques to engineer particles for inhalation.[10,14,15] Nonetheless, these works only mention biodegradable PLGA particles among a huge range of technological strategies. This review focuses on PLGA carriers and their potential in the inhalation field. After a brief description of the rationale behind the choice of polymer carriers for drug inhalation, it describes the recent advances in developing PLGA microparticles and nanoparticles for pulmonary delivery. The most promising particle engineering strategies are discussed, highlighting the advantages along with the major challenges for researchers working in this field.

Why a polymer carrier for inhalation?

Nowadays, drugs are administered via the pulmonary route for two main purposes: (i) local therapies; (ii) systemic absorption. It is currently believed that drug delivery into the human lung represents the best way of treating pulmonary diseases.[2] Inhalation allows direct delivery of the drug to the site of action, with fewer systemic effects than oral therapy. Meanwhile, the lungs may be used as a portal of entry to the body, allowing systemic delivery of drugs via the airway surfaces into the bloodstream.[5] This is an attractive prospect owing to the large surface area of the alveolar region, thickness of the epithelial barrier, extensive vascularization, relatively low proteolytic activity in the alveolar space and absence of first-pass metabolism.

As can be seen in Table 1, locally acting drugs for inhalation are mainly bronchodilators, corticosteroids and antibiotics, which can reverse constriction of bronchial airways and control airway inflammation/infection. Inhaled products administered for their systemic effects mainly include fast-onset analgesics, peptides and proteins, which would otherwise need to be given by injection. In both cases, to yield the best therapeutic outcomes, the drug has to effectively deposit along the airways, remain in situ as long as possible, overcome extracellular and cellular airway barriers and, when needed, reach intracellular targets. This is the case of the emerging oligonucleotide therapies for chronic respiratory diseases, able to reach cell targets underlying disease pathogenesis.[18] Indeed, formulating biopharmaceuticals for inhalation presents new challenges to scientists. The structural complexity of proteins and nucleic acids compared with conventional drugs demands an efficient drug-delivery system that allows the intact macromolecule to gain access to the target site at the right time and for proper duration.[38]

Table 1.  Drugs marketed or undergoing preclinical/clinical studies by the pulmonary route
 Therapeutic applicationDrug classDrugInhalation device*Current development status**
  1. *DPI, dry powder inhaler; pMDI, pressurized Metered Dose Inhaler. **More details on clinical studies can be found at http://clinicaltrials.gov/.

Local delivery AsthmaShort-acting β2 agonist[16,17]Salbutamol (albuterol)Nebulizer, pMDI, DPIMarketed
FenoterolpMDI
PirbuterolNebulizer, pMDI
Long-acting β2 agonist[16,17]SalmeterolpMDI, DPIMarketed
FormeterolDPI
Anticholinergic compounds[16,17]Ipratropium bromideNebulizer, pMDIMarketed
Tiotropium bromidepMDI, DPI
Inhaled corticosteroids[16,17]Beclomethasone dipropionateNebulizer, pMDIMarketed
BudesonideNebulizer, DPI
Fluticasone propionatepMDI, DPI
COPD
Antisense oligonucleotides [18,19]EPI-2010NebulizerTerminated
AIR-645Phase II
PXSTPI-1100Preclinical
ATL-1102Preclinical
CpG oligonucleotides [18,19]AVE-7279NebulizerTerminated
QAX-935 (IMO-2134)Phase I
siRNA[18,19]ExcellairNebulizerPhase II
Cystic fibrosisAntibiotics[20–23]TobramycinNebulizer, DPIMarketed
Aztreonam lisineNebulizerMarketed
Colistimethate sodiumNebulizerPilot trials
 DPIPreclinical
Liposomal ciprofloxacin (ARD-3100)NebulizerPhase II
Levofloxacin (Aeroquin, formerly MP-376)NebulizerPhase III
Mucolytics/mucous mobilizers[20,24]Dornase alfaNebulizerMarketed
Lancovutide (Moli1901)Phase II
Antiproteases[25–27] α 1-antitrypsin (AAT)NebulizerPhase II
Respiratory DistressPulmonary surfactant[28]Phospholipids/surfactant proteinsEndo-tracheal tubeMarketed
Syndrome
Systemic delivery AnalgesiaOpioids[29,30]Fentanyl (Staccato)NebulizerPhase I
Liposomal fentanyl (AeroLEF)Phase II
DiabetesPeptides[31–37]Insulin (Affrezza)DPIPhase II
Glucagon-like peptide (MKC253)Phase I
OsteoporosisCalcitoninDPIPreclinical
Parathyroid hormone

Lung deposition results from the complex interplay of a range of factors, which relate to the size of the particles, the inhalation device, the mode of inhalation or breathing pattern and even to the lung anatomy of the individual.[2,39,40] In particular, the aerodynamic particle size, defined as the diameter of a sphere of unit density characterized by the same settling velocity in air as the particle aerosol being measured, will critically influence the deposition of the inhaled formulation in the upper airways, rather than the conducting and the alveolated airways of the lungs.[2,40,41] It is well recognized that particles below 5 µm can be distributed deep into the smaller airways and this penetration correlates well with a good clinical response to local treatment. Differently, a particle fraction with an aerodynamic diameter in the 1–2 µm range is probably the most efficient for deposition into the capillary-rich alveolar airspaces, the target for the systemic delivery of drugs. Finally, submicron particles can be exhaled, if they are not aggregated and/or if insufficient time is available for their migration to the lung walls.

A number of inhalation devices are available on the market to effectively deliver inhaled drugs to the lungs, such as nebulizers, pressurized metered dose inhalers (pMDIs) and dry powder inhalers (DPIs).[42] Since they have superior handling, metering and reliability as compared with nebulizers, pMDIs are the most commonly used inhalers to treat asthma and COPD.[43] Nonetheless, pMDIs require the use of liquid drug formulations (mainly suspensions) in polluting propellants (chlorofluorocarbon and hydrofluoroalkane are gases causing ozone depletion and greenhouse effects) and delivery efficiency strongly depends on the actuation-breath coordination of the user.[44] DPIs, on the other hand, are breath-actuated and do not cause any coordination problem and, thus, have become accepted increasingly by both patients and formulators.[45,46] In particular, DPIs are currently regarded as the device of choice for macromolecule delivery to the lungs, due to their potential to overcome solubility, bioavailability and stability issues encountered by nebulizers and pMDIs.[14]

From a technological point of view, the handling, processing and inhalation of excipient-free dried drug particles are challenging. Novel strategies have been developed and investigated for drug inhalation via DPIs, including microparticles.[8,9] Advanced products include PulmoSphere™ particles, engineered tobramycin-containing hollow porous microparticles made of dipalmitoyl-phosphatidylcholine (DPPC), the principal component of endogenous lung surfactants.[23] Technosphere™ Insulin (MannKind Corp., Valencia, USA), based on the intermolecular self-assembly of fumaryl diketopiperazine (FDKP), forming a three-dimensional highly porous sphere, is also facing the market.[31,33] Nonetheless, these systems may suffer from the short duration of their therapeutic effect (i.e. frequent administrations). In the same way, lipid-based vesicles (e.g. liposomes, lipoplexes, solid lipid nanoparticles), the most extensively investigated carriers for controlled delivery to the lung, do not allow sustained drug release and are mainly delivered by nebulization.[8,9]

In this scenario, biodegradable polymer carriers have become an increasingly attractive option for inhalation therapies (Table 2).[12,15,47] Although calling into question safety issues, which are far to be fully settled, the use of polymer carriers may allow the reduction of daily drug doses, and this could contribute to better control of local chronic infections while diminishing the rate of drug appearance in the bloodstream (i.e. nonspecific distribution to non-target tissues) and subsequent side effects. The carrier can also offer the potential to combine different drugs[48] or a drug with helper excipients (e.g. absorption promoters, mucolytic agents).[49] This can be very important in the case of complex therapeutic regimens, such as those currently required to treat or, better, to eradicate chronic pulmonary diseases. Furthermore, polymer particles can play a crucial role in improving the therapeutic index of macromolecular drugs by: (i) increasing the stability of the drug both ‘in the bottle’ and in vivo; (ii) increasing the amount of drug that reaches the site of action; (iii) prolonging the residence time of the drug in situ (i.e. sustained pharmacological effect).[50,51]

Table 2.  Main advantages of polymer carriers in pulmonary delivery
Microcarriers Protection of the entrapped drug
Sustained drug release
Reduction of daily drug doses (i.e. side effects)
Combined therapy
Dry formulation
Macrophage escape (>10 µm)
Nanocarriers Protection of the entrapped drug
Sustained drug release
Reduction of daily drug doses (i.e. side effects)
Combined therapy
Efficient transport through lung lining fluid
Cell targeting (e.g. macrophages)
Mucosal vaccination

Another important issue that can be potentially addressed by polymer carrier-based systems intended for inhalation is overcoming extracellular and cellular barriers imposed by the lungs (Figure 2).[7]

Figure 2.

Schematic representation of cellular and extra-cellular barriers imposed by the lung to inhaled drugs. The trachea and bronchi epithelia are made up of basal, ciliated, brush and goblet cells (∼60 µm-height). The same cells are present in bronchiolar epithelium but are not as tall (∼10 µm). In the alveolar region, type-I and secretor type-II cells build up the epithelium. The conducting airways are lined with airway surface liquid, a mucus gel-aqueous sol complex composed by the periciliary and the luminal mucus layers, which get thinner from the level of the trachea (up to 100 µm) to the bronchioles (∼3 µm). The alveoli are lined with alveolar subphase fluid and pulmonary surfactant, which is composed of approximately 80–85% phospholipids, 5–10%proteins and 5–10% other lipids. All the lung surfaces are supervised by resident mononuclear phagocytes (i.e. airway macrophages). They are highly represented in the alveolar lining fluid (∼80% of the bronchoalveolar lavage cellular components) and so are often referred to as alveolar macrophages.[5,30,31]

Although controversial, literature data highlight that LLFs, especially lung mucus, along with resident macrophages, represent a critical barrier to carrier-based therapies. By virtue of its peculiar size and surface properties, a polymer carrier may allow easy diffusion of a drug through the LLF,[52,53] its preferential uptake from lung epithelial cells or improved escape from locally circulating macrophages.[11,54] It has been demonstrated that the potential for access of particles to airway mucus can be tuned at size level.[53] While large nanoparticles with a mean size around 560 nm were almost completely blocked by a 220 µm-thick CF sputum layer, small nanoparticles were retarded only by a factor of 1.3 as compared with buffer.[53] It is worth noting that particle size plays a crucial role also in determining its preferential uptake by resident macrophages, which can rapidly engulf the foreign particles as a defence mechanism. A size of 1–2 µm for inhaled particles is ideal for macrophages to phagocytose (cell diameters ∼15–22 µm), with smaller (below 200 nm) size being taken up less efficiently.[5] In contrast, large particles with geometric diameters of 10–20 µm, able to penetrate deep into lungs and avoid macrophage engulfment, are considered ideal for systemic delivery of protein drugs.[11] Nonetheless, the reduction of the particle size to a nanometric range has been suggested to further enhance the potential of biodegradable polymer carriers for mucosal delivery of vaccines.[55]

Beside particle dimensions, the role played by the surface chemistry of the carrier in its interactions with, and adhesion to, lung cellular and extracellular barriers should also be considered.[52] It has now become clear that slight differences in surface properties may have significant implications in the cellular uptake and the interactions of the carrier with the biological environment. Particular attention must be paid to imparting a charge or adhesiveness to the particle. When modified on the surface with carboxyl groups, particles as small as 59 nm may be completely immobilized by human mucus.[56] On the other hand, large (200–500 nm) non-adhesive nanoparticles achieved by covalent modification of polystyrene particles with polyethylenglycol (PEG) have been demonstrated to rapidly penetrate human mucus.[57] More recently, hydrophilic poloxamer-coated biodegradable nanoparticles, displaying a negative charge, have been demonstrated to more easily diffuse across the mucus barrier leading to a higher intracellular accumulation as compared with positively charged nanoparticles modified with chitosan.[58] Thus, a thorough understanding of carrier interactions with LLFs is a condition sine qua non to achieve particles engineered to cross mucus barriers and, thus, for lung delivery.

Assuming the carrier successfully overcomes extracellular barriers, the drug has still to interact with the target cells of the respiratory tract to be effective. The interaction of the carrier with cells is critical in many applications such as mucosal vaccination or novel oligodeoxynucleotide (ODN)-based therapies. These applications require a firm control over particle–cell interactions, which are mainly dictated by size and surface properties of the delivery system. The induction of humoral immune responses at the mucosal lymphoid tissue (MALT) of the respiratory tract requires the transport of particulate antigens by M-cells and their delivery to the sub-mucosa, where dendritic cells and macrophages process and present the antigen to naïve T-cells in the adjacent mucosal lymph nodes. Similarly, at the alveoli, particulate antigens are taken up and processed by circulating antigen-presenting cells. Thus, when designing new inhalable vaccines, surface-modified polymer carriers, especially nanoparticles, may represent a paramount tool not only to protect the protein-based antigen, but also to tune its interaction with immune cells.[59,60] The challenges of nucleic acid delivery via the pulmonary route are possibly greater, since the drug has to cross the cellular membrane and gain access into the cytoplasm/nuclei, where the final targets are located.[19,61] Therefore, the main function of the carrier, along with macromolecule protection against enzymatic attack, is to facilitate ODN cellular uptake, promote its endosomal escape and release it at the final intracellular target. This can be achieved by biodegradable polymer carriers.[52,62–64]

Poly(lactic-co-glycolic acid) carriers in drug delivery and their potential for inhalation

Among synthetic biodegradable polymers, thermoplastic aliphatic poly(ester)s, PLGAs, have generated tremendous research interest due to their excellent biocompatibility as well as the possibility of tailoring their biodegradability by varying composition (lactide/glycolide ratio), molecular weight and chemical structure (i.e. capped and uncapped end-groups).[51,65] PLGAs characterized by very different in vivo life-times, ranging from three weeks to over a year, are available and approved for human use. Drug encapsulation within PLGA copolymers is regarded as a powerful means to achieve its sustained release for long time-frames and, in the case of labile drugs, effectively protect the molecule from in vivo degradation occurring at the administration site. Thus, injectable PLGA microparticles (< 250 µm – ideally < 125 µm) with different formulation design have been developed for the delivery of small drugs as well as protein therapeutics and nucleic acids,[50,51] resulting in several marketed products. On the other hand, PLGA nanoparticles (< 1 µm) represent a well-established tool to achieve targeted drug or gene delivery, with special emphasis on cancer treatment.[66]

In the attempt to discover novel strategies for the prolonged release of drugs in the lung, PLGA nanoparticles and microparticles have been studied also as carriers for inhalation. As summarized in Table 3, studies with PLGA microparticles for inhalation have been carried out mostly with the intention of developing inhalable formulations of proteins and other macromolecules.[49,67–70,72–74,84] Currently, protein encapsulation within PLGA copolymers is considered a very useful strategy to entrap and deliver the native macromolecule in a sustained manner.[50,51,85]

Table 3.  Examples of poly(lactic-co-glycolic acid) (PLGA)-based carriers tested for pulmonary delivery
 TypeDrugFormulationKey in vitro/in vivo findings
PLGA microparticles Conventional microparticlesInsulin[67,68]Microparticles/mannitol blendsProlonged hypoglycaemic effect in vivo
Minimal signs of lung toxicity
Interleukin-2[69,70]Microparticles/mannitol blendsPreservation of in vitro biological activity (T-cells)
Large porous particlesAnthocyanin[71]Dry powdersProlonged inhibition of A549 cells death induced by cigarette smoke
Persistence in mouse lung (20 days)
No signs of lung inflammation (10 days)
Deslorelin[72]Dry powdersHigh drug concentrations in plasma up to 7 days
Heparin[73]Dry powdersIncreased plasma half-life of the drug
No cytotoxicity (Calu-3)
Insulin[49]Dry powdersPreservation of protein structural integrity
In vivo alveolar deposition
Hypoglycaemic activity at low doses in diabetic rats
Prolonged hypoglycaemic effect in diabetic rats
Prostaglandin E1[74,75]Dry powdersIncreased stability in rat lung homogenates
Increased plasma half-life of the drug
No cytotoxicity (Calu-3)
No in vivo acute toxicity
Hollow porous particlesPalmytil-acylated-exendin-4[76]Dry powders In vivo deposition in the central lung
Prolonged hypoglycaemic effect in diabetic rats
No signs of lung inflammation (2 weeks)
PLGA nanoparticles Nanoparticle dispersionsCombined antitubercular therapy [77]Dispersion for nebulizerSustained plasma therapeutic drug levels
Increased plasma half-life
Enhanced drug bioavailability
Hepatitis B vaccine[78]Dispersion for nebulizerEnhanced mucosal and humoral responses to the vaccine
Efficient uptake of nanoparticles by rat alveolar macrophages
No cytotoxicity (Calu-3)
Insulin[79]Dispersion for nebulizerProlonged hypoglycaemic effect in diabetic
Trojan particles (porous nanoparticle-aggregate particles)Hepatitis B surface antigen[80]Dry powdersHigh mucosal immune response in the lungs
Rifampicin[81]Dry powdersSustained plasma therapeutic drug levels
Drug persistence in lung tissue and cells
Nano-embedded microparticlessiRNA[64]Dry powdersPreservation of siRNA structural integrity
Efficient in vitro gene silencing
No cytotoxicity (H1299 cells)
TAS-103 anticancer drug[82]Dry powdersEnhanced cytotoxicity against A549 cells
High local drug concentration
Tobramycin[83]Dry powders In vitro antimicrobial activity against P. aeruginosa
Nanoparticle composition affects in vivo deposition

Among the techniques developed for the production of PLGA microparticles, solvent evaporation/extraction methods are the most widely employed for the microencapsulation of therapeutic macromolecules.[51] The quality of the final product is related to the preservation of protein integrity along manufacturing and in vivo application.[86,87] Excellent reviews have summarized all the possible causes of protein degradation during the microencapsulation process highlighting the strategies efficient in preventing the phenomenon.[88,89] As a function of the critical steps affecting the quality of the final product, it is useful to single out a class of excipients and select, on the basis of the literature reports, those more suitable to the specific design.[65] For inhalable microparticles, the microencapsulation technique and the selected excipients should also contribute to the correct aerosolization properties of the particles. The technological paradox is that respirable particles have to be small for deposition, but also large enough to allow metering during manufacturing and delivery of the dosage form.

Thus, microparticle porosity, which is generally considered a microsphere defect to be overcome,[90] may oddly become a desired feature for inhalation.[91,92] Conversely, conventional microparticles may require the use of a inert carrier (e.g. mannitol) to become respirable.[69,70]

Beside macromolecule inhalation, the potential of PLGA particles as carriers for lung delivery of antibiotics has been also deeply investigated (Table 3).[77,81,93–97] Here, the primary goal is to increase the local concentration of the antibiotic within the macrophages, the host cells for some bacteria, such as Mycobacterium tuberculosis. Although some successful attempts have been made to deliver antibiotics to the lungs via PLGA microparticles, nanoparticles seem to better fit the purpose. When encapsulated into PLGA nanoparticles, the elimination half-life and mean residence time of the entrapped antibiotics are significantly prolonged as compared with the orally-administered parent drugs, resulting in an enhanced drug bioavailability.[77]

Despite the advantage of reduced doses and dosing frequency, the first-developed PLGA nanoparticle-based ‘liquid’ formulations pose several challenges to practical implementation. Long delivery times, low delivery efficiencies, stability of sustained-release formulations in aqueous solution and access to clean water in developing countries are only a few of the challenges highlighted by researchers in this field. Novel formulation approaches for antibiotic delivery have involved the creation of micron-scale dry powders based on PLGA nanoparticles.[81,83,97] In fact, dry powders based on self-assembled rifampicin-loaded PLGA nanoparticles (‘porous nanoparticle-aggregate particles’ or PNAPs) were developed.[81] Alternatively, mannitol microparticles containing rifampicin-loaded PLGA nanoparticles for inhalation therapy of tuberculosis were prepared in one step using a four-fluid nozzle spray drier.[97] These technologies have demonstrated a great potential also in the development of dry powders for protein delivery, with particular regard to mucosal vaccination,[80,98] and short interference RNA (siRNA) delivery.[64]

Although current literature data suggest the potential of PLGA particles for prolonged drug delivery in the lung, their use for inhalation is still embryonic and presents some shortcomings. In particular, a huge amount of work has still to be produced on clearance mechanisms and persistence in vivo that may considerably limit the benefit of sustained-release inhalation therapy. The persistence of PLGA particles in vivo has been only indirectly demonstrated by some authors, who showed the presence of the released drug in the lung for up to 20 days.[71] If this feature is desirable to prolong the therapeutic effect, its influence on the safety of respirable PLGA particles has not been elucidated yet. Anyway, in vitro cytotoxicity studies have shown that native PLGA has no manifest toxicity against healthy lung macrophages,[99] CF bronchial cells[100] or lung carcinoma cell lines.[64] Furthermore, the pro-inflammatory potential of PLGA nanoparticles has been recently investigated in vivo and results suggest that biodegradable nanoparticles are safer than non-biodegradable polystyrene particles[101] and do not induce lung tissue damage,[102] also up to 10 days.[71] Nonetheless, chronic toxicological data on PLGA particles are still poor and demand for further studies.

As polyesters, PLGAs undergo hydrolysis in vivo, forming biologically compatible moieties that should be cleared from the lung. With respect to biodegradation, among all commercially available PLGA types, those characterized by a rapid in vitro degradation are undoubtedly preferred for inhalation. The creation of synthetic PLGA derivatives more hydrophilic than native PLGAs has been attempted and novel poly(vinyl alcohol) (PVA)-based branched polyesters with PLGA side chains, characterized by faster degradation rates compared with linear PLGA of similar molecular weight, have been employed for pulmonary delivery [62,103,104].

Although few experimental studies have been performed so far to fully explore the potential and drawbacks of PLGA in pulmonary delivery, respirable dry powders based on PLGA particles appear an enticing technological approach for inhalation. However, depositing drug doses reliably into the lungs through the newest DPIs is straightforward in principle, but challenging in reality. The peculiar properties of the drug particles, such as size, shape, density and surface properties, will affect handling, processing, and inhalation of the formulation and, therefore, its likelihood of being deposited in the desired region of the lung [10,105]. In the following, a critical account of the most promising production and formulation techniques employed to engineer PLGA microparticles and nanoparticles for inhalation is provided, highlighting the goals already achieved and those to be gained.

Engineering poly(lactic-co-glycolic acid) microparticles for inhalation

Respirable PLGA microparticles have been generated, making use of several techniques, that feature partly competing, partly complementary characteristics. Great efforts have been made to improve lung deposition of the particle and overcoming intracellular and extracellular barriers imposed by lungs. Besides the PLGA polymeric platform, excipients are sometimes needed to optimize the final particle properties as a function of the preparation technique. The desirable product characteristics firstly include the correct aerosolization properties (i.e. low mass mean aerodynamic diameter (MMAD), high fine particle fraction (FPF) and emitted dose). To this end, PLGA particles have been engineered in very different ways, taking into account that the particles should also comply with the requirements of drug stability and prolonged release. (Figure 3)

Figure 3.

Typical architectures of respirable poly(lactic-co-glycolic acid) microparticles.

Among the parameters that can be adjusted to optimize the efficiency of dry powders for inhalation, particle mass density and size have recently drawn researchers' attention to limit loss of drug owing to PLGA particle aggregation in the inhaler and macrophage-mediated clearance of the particles from the lungs.[5,10] The idea is that conventional PLGA particles achieved by micronization (i.e. 1–3 µm) are both prone to aggregation and can be rapidly phagocytosed in the deep lung by alveolar macrophages. On the contrary, large porous particles (LPPs) with low mass density (< 0.4 g/cm3) and high geometric diameter (> 5 µm) can be used to enhance both particle aerosolization behaviour and residence time in the lung.[106] Thus, working on size, shape and density of the particles, the first PLGA-based LPPs for insulin delivery were achieved.[11] More recently, great efforts have been made to achieve ‘regularly shaped’ porous particles by operating selectively on size and density, thus achieving large porous ‘microspheres’. In so doing, PLGA-based LPPs have been engineered to sustain deslorelin delivery via the deep lungs,[72] allow efficient pulmonary adsorption of low-molecular-weight heparin[73] and increase insulin systemic availability.[49]

Large porous particles

From a technological standpoint, PLGA-based porous microspheres for inhalation have been typically obtained by the double emulsion (wi/o/we) technique in the presence of a porogen (Table 4).

Table 4.  Preparation techniques employed to develop porous poly(lactic-co-glycolic acid) (PLGA) microparticles for inhalation
Preparation techniquePorosigen agentDelivered drug
  1. *Patented technology that uses acoustic excitation and/or flow-limited field-injection electrostatic spraying to fabricate uniformly sized particles. **Particle porosity has been achieved by rapid evaporation of the organic solvent.

Supercritical fluids (CO2SC)Cyclodextrins[72]Deslorelin
Precision particleOils[107]Ciprofloxacin
Fabrication (PPF)*  
Single emulsion (o/w)Poloxamer (Pluronic F68 and F127) [84,108]Human Growth Hormone
Palmytil-acylated-exendin-4
Double emulsion (wi/o/we)Cyclodextrins[49,76,91,109]Insulin
Bovine Serum Albumin
Palmytil-acylated-exendin-4
Sodium chloride[76]Palmytil-acylated-exendin-4
Poly(ethylenimine)[73,74]LMW Heparin
Prostaglandin E1
Ammonium bicarbonate[92,100,110,111]Budesonide
Decoy oligonucleotide
Doxorubicin
Lysozyme
Rhodamine-dextran
Poly(vinyl alcohol)[112]Hepatitis B surface antigen
None[113]**Capreomycin sulfate

When microspheres are prepared by the wi/o/we technique, the presence of osmotically active drugs and/or excipients in the internal aqueous phase (wi), as well as their concentration, generates an osmotic pressure gradient which can strongly affect particle porosity (Figure 4a).[65] Thus, LPPs can be achieved adding osmogens in the internal aqueous phase of the double emulsion, thus causing water influx from the external to the internal aqueous phase during solvent evaporation (i.e. particle hardening). Under the influence of an osmotic gradient, the organic phase of the double emulsion acts as a semi-permeable membrane allowing the passage of water across the organic phase. In particular, high osmotic pressures in wi generate an influx of water from the external aqueous phase (we) to wi. The occurrence of this process during particle hardening results in the formation of a typically porous structure.

Figure 4.

Production of large porous particles by the double-emulsion technique. (a) Generation of an osmotic pressure gradient between wi and we. (b) Addition of an effervescent salt (e.g. ammonium bicarbonate) in wi.

In this sense, to control PLGA particle porosity and, consequently, flow and aerosolization properties of the developed dry powders, hydroxypropyl-β-cyclodextrin (HPβCD) has been recently tested as aid-excipient in insulin-loaded PLGA microparticles intended for pulmonary delivery.[49,91] The use of HPβCD in LPP production is very intriguing not only in the light of cyclodextrin's osmotic properties but also for its potential to enhance macromolecule adsorption through respiratory epithelium. Several formulations, differing in HPβCD and insulin loadings, were produced by the wi/o/we technique and their properties compared.[91] Insulin and, when added, HPβCD were solubilized in wi. The technological results show that the combination of appropriate amounts of insulin and HPβCD plays a crucial role in achieving PLGA/HPβCD/insulin LPPs with the desired bulk and aerodynamic properties (i.e. a highly porous structure, a very low density (0.1 g/ml) and a theoretical MMAD (MMADt) lower than 10 µm).[91] The good aerosolization behaviour of the developed LPPs was confirmed by in vitro aerosolization tests, showing that optimized PLGA/HPβCD/insulin particles had an experimental MMAD (MMADexp) ranging from 4.01 to 7.00 and an FPF estimated to be 26.9–89.6% at the different airflow rates tested (i.e. 30–90 l/min).[49] Confocal microscopy studies, performed after in vivo administration of labelled PLGA/HPβCD/insulin LPPs to the rat lung by means of a low-scale DPI, suggest that particles reach the alveoli and remain in situ after delivery. Finally, the therapeutic potential of optimized PLGA/HPβCD/insulin LPPs was confirmed by dose–response studies performed in both normoglycaemic and streptozotocin-induced diabetic rats.[49]In vivo data show that PLGA/HPβCD/insulin LPPs are able to reach the alveoli and release insulin, which is absorbed in its bioactive form. While insulin solutions administered via pulmonary route are unable to cause a significant hypoglycaemic effect, insulin delivered through PLGA/HPβCD/insulin LPPs at the same doses (0.5–4.0 IU/kg) significantly reduces blood glucose level as a function of the administered dose in both animal models. The developed LPPs, tested in hyperglycaemic rats in evident pathological conditions, exert a very significant and longer hypoglycaemic effect even at insulin doses as low as 0.5 IU/kg (about 0.5 mg of PLGA/HPβCD/insulin LPP per rat) as compared with an insulin solution. The duration of the effect was consistent with previous results achieved by Edwards and colleagues,[11] who showed falling glucose levels for the first 10 h after inhalation of PLGA/insulin LPPs, followed by relatively constant low glucose levels for the remainder of the 96-h period. It is worth noting that these effects were observed for PLGA/insulin LPPs at much higher insulin doses (about 9 mg of LPPs containing 20% of insulin by weight were administered per rat) as compared with PLGA/HPβCD/insulin LPPs.

HPβCD has been recently employed to produce also albumin-coated hollow porous PLGA particles for pulmonary delivery of palmityl-acylated exendin-4 (Pal-Ex4), a potent glucagon-like peptide-1 agonist with great potential for the treatment of diabetes.[76] The contemporary addition of HPβCD and sodium chloride within wi, along with controlled solvent evaporation, allowed the achievement of hollow LPPs. Plain particles were further processed with human serum albumin (HSA), which was conjugated with the carboxylate groups of poly(ethylene-alt-maleic anhydride), to achieve HSA-coated hollow LPPs based on PLGA. The developed Pal-Ex4 loaded particles demonstrated great potential for inhalation, in terms of drug loading, release rate, in vitro/in vivo aerosolization efficiency and duration of the hypoglycemic effects in vivo. On the basis of the results, the authors ascribe several benefits to HSA coating: prevention of particle aggregation, improvement of LPP aerosolization properties and, last but not least, potential for tight binding of the fatty-acid conjugated peptide, so as to control its release in vivo.[76]

Despite encouraging results, severe limitation of the osmogen-based technological approach is represented by the poor control of drug encapsulation efficiency, ascribable to mass exchanges and consequent drug loss occurring between the two phases during particle hardening. This phenomenon, as well as rapid drug release due to the macroporous structure of the system, can be particularly dramatic in the case of highly hydrophilic macromolecules, such as nucleic acids.[114] Establishment of interactions of the drug with helper excipients may limit macromolecule loss. Along this line, an alternative osmotic agent that has been tested to produce LPPs is polyethylenimine (PEI), a hydrophilic polycation that forms complexes by electrostatic interactions with different anionic macromolecules (e.g. nucleic acids, heparin).[115] A concentration-dependent increase in particle porosity was experienced after PEI addition in wi during preparation of PLGA microparticles for the delivery of oligonucleotides.[114,116] The use of PEI as core-modifying agent has been re-examined to achieve respirable PLGA LPPs of low-molecular-weight heparin[73] and, more recently, prostaglandin E1 (PGE1).[74] Despite particle porosity, LPPs showed high entrapment efficiency, ascribed to drug–polycation interactions. Furthermore, in both cases, a remarkable extension of the plasma half-life of the drug (from 6 h to 24 h) was observed.

An alternative formulation strategy to achieve PLGA-based LPPs for the prolonged release of hydrophilic macromolecules relies on the use of an effervescent agent, namely ammonium bicarbonate, which decomposes into ammonia and carbon dioxide during emulsification, forming a porous matrix as the gas products escape (i.e. gas-foamed LPPs)[110] (Figure 4b). Since pore formation depends on effervescence rather than on diffusional mass exchanges between aqueous phases, this technique allows efficient encapsulation of macromolecules in highly porous PLGA particles.[92,100,110]

The new process was firstly applied by Yang et al. for the encapsulation of two model molecules, lysozyme and doxorubicin hydrochloride.[110] The addition of an appropriate amount of ammonium bicarbonate was essential to achieve a homogeneous population of highly porous particles with optimal aerodynamic properties. Although the authors succeeded in improving drug encapsulation efficiency, poor control over release properties was achieved without further modification of formulation conditions. To further engineer the system, lipid-engineered gas-foamed LPPs were recently developed in our laboratory.[92] To control LPP release properties, two lipid aid excipients were tested. Our first choice was DPPC, the major component of human lung surfactant, which is gaining increasing research interest in the development of respirable dry powders for a number of reasons, including its biocompatibility.[9] As an alternative lipid excipient, we investigated the potential of 1,2-dioleoyl-3-trimethylammonium-propane (DOTAP), a cationic lipid extensively studied as transfection agent for nucleic acids.[117] The effect of the presence of DPPC or DOTAP upon the properties of gas-foamed LPPs containing a model hydrophilic macromolecule, rhodamine B isothiocyanate–dextran (Rhod-dex), was assessed. We found that in the case of hydrophilic macromolecules unable to interact with PLGA end-groups, such as Rhod-dex, excipient addition was essential to increase the amount of drug entrapped within LPPs, being as high as 80% only for DPPC-engineered or DOTAP-engineered gas-foamed LPPs. The Rhod-dex release profile from LPPs was also strongly affected by excipient addition in the initial formulation, with lipid-engineered LPPs allowing for a more prolonged release of Rhod-dex as compared with excipient-free LPPs. Conceiving the developed LPPs for drug inhalation, DPPC-engineered and DOTAP-engineered LPPs displayed optimal FPF and MMADexp. In vivo deposition studies performed after intratracheal administration of LPPs in rats confirmed the ability of the developed dry powders to deposit along bronchi and bronchioles. On the basis of encouraging technological results, we are currently developing respirable DPPC-engineered gas-foamed LPPs for prolonged pulmonary delivery of a decoy oligodeoxynucleotide to nuclear factor-kB (NF-κB) (dec-ODN) for the treatment of chronic lung inflammation (e.g. CF).[100] Dec-ODN LPPs represent the first attempt to delivery ODN-based therapeutics to the lung in the form of dry powders. Dec-ODN was effectively encapsulated (∼80%) within biodegradable particles with controlled porosity, confirming that the use of lipid excipients and gas foaming may result in efficient microencapsulation of a highly hydrophilic macromolecule (i.e. dec-ODN) within highly porous PLGA particles. Furthermore, a sustained release of dec-ODN from LPPs was achieved. Particle aserosol performance, as evaluated by multi-stage liquid impinger, demonstrated that the developed dec-ODN LPPs have very good flow properties, with more than 95% of the capsule content being emitted during aerosolization. The MMADexp of  < 6 µm, along with findings of previous studies, strongly supports a preferential deposition of dec-ODN LPPs in the distal airways.[63]In vitro pharmacological studies performed in CF human bronchial epithelial IB3-1 cells showed that lipopolysaccharide (LPS) challenge caused an increase of NF-κB/DNA binding activity, which was significantly inhibited by DPPC-engineered dec-ODN LPPs at 24 and 72 h.[100] This inhibitory effect on NF-κB/DNA binding activity was correlated with decreased interleukin (IL)-6 and IL-8 secretion and mRNA levels. In contrast, naked dec-ODN exhibited these effects only at 24 h. These results were supported by an enhanced structural stability of dec-ODN in cultured cells due to entrapment in LPPs.[63]

PLGA-based LPPs have been also achieved by the single emulsion technique (Table 1)[84,108]. Porous biodegradable PLGA microparticles with interconnected pores have been fabricated using poloxamers as extractable porogens (Figure 5). Pores are generated by the time difference between PLGA hardening and the extraction of the hydrophilic surfactants from the organic phase using water. Nonetheless, the drug should be loaded by adsorption onto LPPs. Thus, it is very important to favour drug interactions with PLGA. This is the case of the fatty acid-conjugated peptide Pal-Ex4, the palmitic acid moieties of which were demonstrated to improve Ex4 interactions with the hydrophobic PLGA and, consequently, to extend its in vitro release (completed after five days).[76] In turn, an extended therapeutic duration in vivo was observed, though ascribable not only to sustained drug delivery but also to the increased plasma half-life of Pal-Ex4 due to the induction of albumin binding, typical of the conjugated peptide itself.

Figure 5.

Production of large porous particles by the single-emulsion technique. Interconnected pores are generated by addition of extractable porogens in the organic phase.

Novel technologies, such as the supercritical fluid process and the patented precision particle fabrication (PPF), have been also tested for the achievement of PLGA-based LPPs (Table 1). Supercritical CO2 processing was applied to prepare deslorelin-loaded LPPs with the aim of reducing the residual solvent content of the particles, retain protein integrity and sustain its release in vivo.[72] The developed particles effectively sustained the systemic delivery of deslorelin via the deep lungs for up to seven days as compared with plain deslorelin (plasma concentrations sustained up to three days) and conventional deslorelin PLGA particles (plasma concentrations 2-fold lower at day 7). More recently, the patented PPF technology was used to create monodisperse PLGA-based LPPs for pulmonary delivery of nano-precipitated ciprofloxacin (NanoCipro dry powders) using oils as extractable porogens.[107] Notably, particle morphology strictly depended upon the nature and concentration of the extractable oil – canola oil led to PLGA particles with a porous web-like internal structure and with silicon oil hollow porous particles were achieved. The main limitation of the technique was the low ciprofloxacin encapsulation efficiency.

Despite the peculiar limitations, all the advanced particle-engineering techniques, along with the traditional methods of microencapsulation, have contributed to the increased possibility of formulating LPPs for pulmonary delivery. Notably, PLGA particles have been effectively engineered into LPPs with aerosol and release properties suitable for the prolonged delivery of drugs deep in the lungs. The challenge for researchers working in this field is that none of these techniques can really provide the solution to all types of molecules.

Engineering poly(lactic-co-glycolic acid) nanoparticles for inhalation

In the current era of nanotechnology, biodegradable nanoparticles are gaining momentum as carriers for the inhalation of drugs.[12,118] Indeed, nanoparticle delivery to the lungs is an attractive concept because it can cause retention of the carrier in the lungs (particles of a few hundred nanometers represent a tenacious resident of the lungs) accompanied by a prolonged drug release, meaning improved lung bioavailability over conventional PLGA microparticles. Furthermore, nanoparticles represent a promising tool to more simply penetrate airway barriers, which can be better overcome at nanosize level.[52] Finally, carriers of nanometric size offer potential for drug targeting to specific lung tissue and cell populations.[119]

From a technological point of view, the development of effective PLGA nanocarriers for inhalation first requires adequate engineering of the particles at the nanosize level. On this matter, PLGA nanoparticles with different features (size, morphology, zeta-potential) may be achieved by controlling the parameters specific to the production method employed.[120,121] Since the potential of nanoparticles to interact with lung tissue strongly depends on nanoparticle size and surface charge,[52] this is an important aspect to be taken into account for excipient selection and formulation design. Nonetheless, another great challenge of using nanoparticles for inhalation is the low inertia of dry powders with a mean size lower than 1 µm, which are generally exhaled upon inhalation.[5,12] Thus, the most widely employed nanoparticle-based inhalation approach relies on the nebulization of nanoparticulate aqueous dispersions.

Different strategies to engineer the system at the microsize level and achieve ‘micrometric’ nanoparticle-based dry powders have recently been explored (Figure 6).[98,122] Among them, self-assembly of nanoparticles or their embedding within an inert ‘microcarrier’ have been extensively investigated and, probably, represent the most promising technological approach.

Figure 6.

Inhalable poly(lactic-co-glycolic acid) nanoparticles. Nanoparticle-based dry powders for inhalation can be achieved by: adsorption on coarse inert carriers, nanoparticle self-assembling (Trojan particles or porous nanoparticle-aggregate particles) or embedding nanoparticles into an inert microparticle.

The ‘Trojan’ approach

Since nanoparticle drying to develop dry powders for inhalation is very difficult to achieve, given that nanoparticles aggregate excessively in the dry state, large porous carriers have been used to deliver nanoparticles deep in the lung (i.e. Trojan particles) in the form of dry powders.[123] As in Virgilio's epic poem, which tells of the Greek army entering and destroying the city of Troy staying hidden in a horse, PNAPs have been developed to act as carrier particles that release nanoparticles once delivered into the body, thereby acting as a ‘Trojan’ delivery systems for nanoparticles (Figure 7).[80,81,123] In so doing, all the advantages of nanoparticles for inhalation (e.g. long-term residence in situ, prolonged drug release, macrophage targeting) could be combined with the ease of flow, processing and aerosolization potential of LPPs.

Figure 7.

Scanning electron micrographs of spray dried rifampicin porous nanoparticle-aggregate particles (PNAPs) containing 80% nanoparticles by weight (PNAP80). A magnification of the surface of the PNAP indicates a shell of aggregated nanoparticles with intact structure (scale bars represent: 2 µm (a), 1 µm (b)).[81]

Formation of the large porous nanoparticle aggregates occurs via a spray-drying process that ensures the drying time of the sprayed droplet (i.e. time required for a droplet to dry) is sufficiently shorter than the characteristic time for redistribution of nanoparticles by diffusion within the drying droplet (i.e. time required for nanoparticles to diffuse from the edge of the droplet to its centre). The proof of principle of using Trojan particles for inhalation was given on polystyrene nanoparticles. This approach has been found useful for achieving hybrid DPPC and hyaluronic acid large porous carriers for PLGA nanoparticles releasing dexamethasone acetate.[124] Scanning electron microscopy and confocal laser scanning microscopy analysis showed that dexamethasone-loaded Trojan particles are spherical, hollow and possess an irregular surface ascribable to nanoparticles, likely held together by DPPC and hyaluronic acid.

More recently, rifampicin-loaded PLGA nanoparticles have been dispersed throughout a matrix of an inert excipient, l-leucine, a hydrophobic amino acid demonstrated to improve powder dispersibility and aerosolization properties.[81] Rifampicin-loaded PNAPs possessed properties suitable for efficient deposition in the lungs. In vitro release showed an initial burst of rifampicin, with the remainder available for release beyond 8 h. Analogously, systemic levels of rifampicin were detected for 6–8 h after delivery of PNAPs to guinea-pigs by insufflation. A prolonged permanence of rifampicin both in bronco-alveolar lavage (BAL) and lung tissue was observed as compared with porous particles containing free rifampicin.

A recent paper described dry powders intended for vaccine delivery were based on PNAP technology.[80] In particular, PLGA/PEG nanoparticles composed of a PLGA core and a PEG hydrophilic shell, containing recombinant hepatitis B surface antigen (rHBsAg), were prepared by the double emulsion method and subsequently spray-dried with l-leucine. Again, hollow and low-density particles with a rough inner surface, ascribed to intact PLGA/PEG nanoparticles, were achieved. Various PNAP formulations containing rHBsAg were administered to guinea-pigs by the pulmonary route and the immune response elicited in the systemic circulation and in the lungs assessed. The IgG titres were measured in the serum for 24 weeks after the initial immunization; whereas the titres of IgA, an indicator of mucosal immunity, were measured in BAL. Results showed that rHBsAg PNAPs have the advantage of eliciting a high mucosal immune response in the lungs without traditional adjuvants. Nonetheless, no formulation was able to elicit a systemic response comparable with that of a control antigen adsorbed on alum administered via the parenteral route.

Nano-embedded microparticles

Although direct drying of nanoparticle suspensions has demonstrated some success in the creation of dry powders composed of agglomerated PLGA nanoparticles, this technique requires great attention to preserve nanoparticle integrity upon drying and may not always offer the desired control over their flow and aerosolization properties. An alternative early-investigated approach consists of nanocomposite particles obtained by inclusion of drug-loaded PLGA nanoparticles within sugar micoparticles, used as inert carriers.[83,125–128] The nanoparticle-containing micron-sized particles, also termed nano-embedded microparticles (NEMs), are designed to release primary nanoparticles after reaching deep lung, upon the dissolution of the inert carrier in LLF (Figure 8).

Figure 8.

Proposed mechanism of nanoparticle release from nano-embedded microparticles into lung lining fluid.

Among the inert excipients available for NEM production, lactose is the first choice and the most commonly used in marketed DPIs. It has an established safety and stability profile, may be processed by different manufacturing techniques with tight controls over purity and physical properties, is easily available at different grades and is inexpensive.[9] It is conceived that the fine particles of lactose reaching the lungs are rapidly absorbed and metabolized by the intestinal epithelium and are principally excreted in urine. Furthermore, in contrast to oral administration, lactose swallowed at the levels present in inhaled preparations (up to 25 mg) is unlikely to present problems even in patients with lactose intolerance.[129]

Other safe inert excipients, such as mannitol and trehalose, have been tested as potential carriers in naoparticle-based dry powders for inhalation and have found their way into marketed products for inhalation.[9] Research interest in mannitol is increased by its ability to act as an airway rehydrating agent, inducing water flux into the bronchial lumen.[130] Indeed, the first attempts to produce NEMs for inhalation have been made with mannitol particles loaded with insulin-loaded PLGA nanoparticles by a spray-drying fluidized bed granulator.[79] This resulted in soft matrix composite granules displaying aerosolization properties superior to those of freeze-dried nanoparticles. In vivo deposition studies showed that more than 50% of the composite granules were able to reach rat bronchioles and alveoli. As a result, a prolonged pharmacological effect lasting more than 12 h was achieved as compared with insulin solutions administered either intravenously or intratracheally.

Recently, inhalable composite particles made of sugars and PLGA have been achieved by a classical spray-drying technique.[82,83,125,126] Trehalose and lactose were tested as inert carriers for rifampicin-loaded PLGA nanoparticles.[125,126] Formulation studies were useful to assess the effect of the preparation conditions – inlet temperature, size and weight ratio of primary NPs – on the properties of NEMs. Notably, the authors found that the inlet temperature of the spray-drier was crucial to allow the decomposition of NEMs into intact nanoparticles and its adjustment was necessary when the size of the primary nanoparticles or the sugar carrier were changed. Contrariwise, once achieved at optimum temperatures, the ratio by weight of nanoparticles did not affect the inhalation performance of NEMs. On the basis of encouraging preliminary data, PLGA nanoparticles were loaded with an anti-cancer drug, TAS-103, and spray-dried with threalose.[82] The cytotoxicity of the developed NEMs against A549 cells was higher than that of the free drug. Furthermore, when NEMs were administered to rats by inhalation, the cytotoxic agent reached higher concentrations in the lung than in the plasma or those achieved after intravenous administration of free drug.

In a recent work, we tried to shed light on the role played by nanoparticle composition on the interactions of NEMs with the lung environment.[83] To this purpose, we designed and developed a pulmonary delivery system for antibiotics, such as tobramycin, based on NEMs consisting of PLGA nanoparticles embedded in an inert microcarrier made of lactose. At nanosize level, helper hydrophilic polymers were used to impart the desired surface, bulk and release properties to PLGA nanoparticles prepared by a modified emulsion-solvent diffusion technique.[83] Results showed that PVA and chitosan are essential to optimize the size and modulate the surface properties of tobramycin-loaded PLGA nanoparticles, whereas the use of alginate allows efficient tobramycin entrapment within nanoparticles and its release up to one month. Optimized formulations displayed good in vitro antimicrobial activity against Pseudomonas aeruginosa planktonic cells. Nonetheless, since nanoparticles will be embedded in vivo in the lung mucus, which will prevent them from reaching the target, selected nanoparticle formulations were subjected to further in vitro studies to assess their ability to interact with mucus environment. The rough determination of nanoparticle–mucin interactions by the mucin-particle method suggest that chitosan-engineered nanoparticles have a higher tendency to interact with mucin as compared with PVA-engineered ones. The strength of interaction was confirmed by the ζ-potential of the mucin–nanoparticle dispersion, indicating a strong adsorption of the polymer onto the surface of chitosan-engineered nanoparticles. This phenomenon will likely affect nanoparticle diffusion through mucus. Fluorescence-assisted transport studies performed on a artificial mucus model,[131] showed that nanoparticle transport through mucus is facilitated when mucin shields the nanoparticle surface charge (i.e. chitosan-engineered nanoparticles). Notably, in vivo biodistribution studies of NEMs engineered by spray-drying showed that PVA-engineered alginate/PLGA nanoparticles reached the deep lung, while chitosan-engineered nanoparticles were found in great amounts in the upper airways, lining lung epithelial surfaces. Thus, PLGA nanoparticle composition appears to play a crucial role in determining not only the technological features of nanoparticles but also, once processed in the form of NEMs, their in vitro/in vivo deposition pattern.

A crucial issue pertaining to NEMs is related to the drug-to-powder ratio, which still needs to be improved to achieve safe and effective NEMs for in vivo inhalation of antibiotics.[83] Actually, the addition of an inert ingredient in elevated ranges decreases the amount of PLGA nanoparticles and, in turn, the dose of active drug administered. Nonetheless, this is not a challenge for drugs that are therapeutically active at very low doses, such as biotech molecules (i.e. proteins rather than oligonucleotides). On this matter, it should be underlined that spray-drying has been demonstrated to be an excellent technique also for engineering of siRNA-loaded PLGA nanoparticles.[64,127] Mannitol NEMs with aerosol properties suited for deep lung delivery have been engineered without destroying the biological activity of sensitive molecules.

In the light of current literature data it can be concluded that PNAPs and NEMs are very promising systems for sustained delivery of conventional and biotech drugs in the lungs. In perspective, they may represent important tools for mucosal vaccination as well as for efficient treatment of local lung diseases, which demand nanosized carriers.

Conclusions

Over the last 20 years, drug delivery to the respiratory tract has become of choice for the treatment of asthma, COPD and respiratory infections. The main objective is to confine the therapeutic agent to the airways so as to maximize drug concentration at the site of action and, in so doing, maximize drug efficacy. In addition, the lung is under investigation as a portal of entry to the body, permitting delivery of drugs, especially peptides and proteins, via the airway surfaces into the bloodstream. Literature review shows that advanced particle engineering techniques have contributed to the increased possibility of formulating nano-carriers and micro-carriers for both local and systemic pulmonary delivery of conventional drugs and biotech macromolecules, such as emerging inhaled proteins and oligonucleotides.

Although carrier-free formulations for inhalation are somewhat advisable due to their low toxicological profile, biodegradable and biocompatible polymer particles may permit efficient protection and long-term delivery of the inhaled drug. On this matter, PLGA particles for inhalation have been produced on the basis of numerous strategies that use a combination of available particle technologies and excipients to develop the desired dosage form. Particle properties can be finely tuned at micro-size and nano-size level to fulfill specific therapeutic needs. The challenge to carrier development is that none of the technological approaches reported in the literature is universal and it is very difficult to choose a priori the one that best fits the purpose. In vitro/in vivo studies represent a critical step before selection of the best formulations as candidates for use in humans. Although lung deposition studies are still crucial, there is an impelling need for in vitro/in vivo models to investigate what happens after the particle has landed. Special attention must be paid to the evaluation of the safety of PLGA in the lung. In particular, the shortage of chronic toxicological data on engineered PLGA particles must be balanced with studies to definitively assess the toxicological potential of such particles in the lung, before their in vivo use.

In the light of these considerations, nano-carriers and micro-carriers of PLGA seem an emerging formulation strategy in pulmonary delivery. When adequately engineered, PLGA particles may provide sustained drug delivery to the lungs, extend duration of action, reduce the therapeutic dose and reduce the adverse effects of highly toxic drugs, thus improving patient compliance. Although many challenges still exist, PLGA carriers represent a real benefit for both pharmaceutical companies working to develop novel inhaled products and patients suffering from chronic lung diseases.

Declarations

Conflict of interest

The Author(s) declare(s) that they have no conflicts of interest to disclose.

Funding

The authors wish to thank Fondazione per la Ricerca sulla Fibrosi Cistica-Onlus – Delegazione di Torino (FFC#5/2007) and –Delegazione del Lago di Garda con i GdS dell'Isola Bergamasca e di Chiasso (FFC#23/2011) for supporting their research on the development of respirable PLGA carriers for cystic fibrosis therapy.

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