The development of an inhalant therapy that is efficacious and safe depends not only on a pharmacologically active molecule, but also on a well-designed delivery system and formulation. It is the optimization of the whole system (drug, drug formulation and device) that is necessary for the successful development of inhalation therapies, both new and old, for the treatment of local and systemic diseases. Drug-device combinations must aerosolize the drug in the appropriate particle size distribution and concentration to ensure optimal deposition and dose in the desired region of the lung. In Part II of this review, the influence that the type of aerosol delivery system and drug formulation has on a drug's therapeutic effectiveness will be discussed.
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Pulmonary delivery devices
The lung has served as a route of drug administration for thousands of years. The origin of inhaled therapies can be traced back 4000 years ago to India, where people smoked the leaves of the Atropa belladonna plant to suppress cough. In the 19th and early 20th centuries, asthmatics smoked asthma cigarettes that contained stramonium powder mixed with tobacco to treat the symptoms of their disease. The development of modern inhalation devices can be divided into three different categories (Figure 1), the refinement of the nebulizer and the evolution of two types of compact portable devices, the metered-dose inhaler (MDI) and the dry powder inhaler (DPI). The advantages and disadvantages of each system will be discussed below and are summarized in Table 1. More detailed reviews of inhalation technology have been previously published [1-6].
Figure 1
Figure 1
Evolution of pulmonary delivery devices
Table 1
Table 1
Advantages and disadvantages of inhalation devices
Nebulizers
Nebulizers have been used for many years to treat asthma and other respiratory diseases. There are two basic types of nebulizer, jet and ultrasonic nebulizers. The jet nebulizer functions by the Bernoulli principle by which compressed gas (air or oxygen) passes through a narrow orifice creating an area of low pressure at the outlet of the adjacent liquid feed tube. This results in drug solution being drawn up from the fluid reservoir and shattered into droplets in the gas stream. The ultrasonic nebulizer uses a piezoelectric crystal vibrating at a high frequency (usually 1-3 MHz) to generate a fountain of liquid in the nebulizer chamber; the higher the frequency, the smaller the droplets produced.
Constant output jet nebulizers can aerosolize most drug solutions and provide large doses with very little patient co-ordination or skill. Treatments using these nebulizers can be time-consuming but are also inefficient, with large amounts of drug wastage (50% loss with continuously operated nebulizers [7]). While these disposable nebulizers are inexpensive, the compressors supplying the air or oxygen are not. Most of the prescribed drug never reaches the lung with nebulization. The majority of the drug is either retained within the nebulizer (referred to as dead volume) or released into the environment during expiration. On average, only 10% of the dose placed in the nebulizer is actually deposited in the lungs [7].
The physical properties of drug formulations may have an effect on nebulization rates and particle size. The viscosity, ionic strength, osmolarity, pH and surface tension may prevent the nebulization of some formulations. If the pH is too low, or if the solution is hyper- or hypo-osmolar, the aerosol may induce bronchoconstriction, coughing and irritation of the lung mucosa [8, 9]. As well, high drug concentrations may decrease the drug output with some nebulizers; colomycin at concentrations >75 mg ml−1 foams in all nebulizers, especially ultrasonic ones, making aerosolization of the drug very inefficient if not impossible [8].
Advances in technology have led to the recent development of novel nebulizers that reduce drug wastage and improve delivery efficiency. Enhanced delivery designs (Pari LC Star, Pari, Germany) increase aerosol output by directing auxiliary air, entrained during inspiration, through the nebulizer, causing more of the generated aerosol to be swept out of the nebulizer and available for inhalation. Drug wastage during exhalation is reduced to the amount of aerosol produced by the jet air flow rate that exceeds the storage volume of the nebulizer. Adaptive aerosol delivery (Halolite; Medic-Aid, Bognor Regis, UK) monitors a patient's breathing pattern in the first three breaths and then targets the aerosol delivery into the first 50% of each inhalation. This ensures that the aerosol is delivered to the patient during inspiration only, thereby eliminating drug loss during expiration that occurs with continuous output nebulizers. The results of a radiolabelled deposition study showed that 60% of the emitted dose from the Halolite device is deposited in the lungs, 37% in the oropharynx and stomach and only 3% is lost to the environment [10]. A number of metered-dose liquid inhalers, including AERx (Aradigm, Hayward, CA, USA), AeroDose (AeroGen, Sunnyvale, CA, USA) and Respimat (Boehringer Ingelheim, Ingelheim, Germany), have been developed that produce a fine aerosol in the respirable range by forcing the drug solution through an array of nozzles with 30-75% of the emitted dose being deposited in the lungs [6, 11]. These new technologies have been discussed extensively in recent reviews [5, 6].
Metered-dose inhalers
The MDI was a revolutionary invention that overcame the problems of the hand-bulb nebulizer, as the first portable outpatient inhalation device and is the most widely used aerosol delivery device today. The MDI emits a drug aerosol driven by propellants, such as chlorofluorocarbons (CFC) and more recently, hydrofluoroalkanes (HFAs) through a nozzle at high velocity (> 30 m s−1). MDIs deliver only a small fraction of the drug dose to the lung. Typically, only 10-20% of the emitted dose is deposited in the lung [3]. The high velocity and large particle size of the spray causes approximately 50-80% of the drug aerosol to impact in the oropharyngeal region [12]. Hand-mouth discoordination is another obstacle in the optimal use of the MDI. Crompton and colleagues [13] found 51% of patients experienced problems co-ordinating actuation of the device with inhalation, 24% of patients halted inspiration upon firing the aerosol into the mouth, and 12% inspired through the nose instead of the mouth when the aerosol was actuated into the mouth.
The delivery efficiency of an MDI depends on a patient's breathing pattern, inspiratory flow rate (IFR) and hand-mouth co-ordination. The studies by Bennett et al.[14] and Dolovich et al.[15] demonstrated that for any particle size between 1 and 5 µm mass median aerodynamic diameter (MMAD), deposition was more dependent on IFR than any other variable. Increases in IFR result in decreases in total lung dose deposition and penetration into the peripheral airways. Fast inhalations (> 60 l min−1) result in a reduced peripheral deposition because the aerosol is more readily deposited by inertial impaction in the conducting airway and oropharyngeal regions. When aerosols are inhaled slowly, deposition by gravitational sedimentation in peripheral regions of the lung is enhanced [16]. Peripheral deposition has also been shown to increase with an increase in tidal volume and a decrease in respiratory frequency. As the inhaled volume is increased, aerosols are able to penetrate more peripherally into the lungs [17]. A period of breath holding on completion of inhalation enables particles which penetrate the periphery to be deposited in that region, instead of being exhaled during the expiratory phase. Thus, the optimal conditions for inhaling MDI aerosols are from a starting volume equivalent to the functional residual capacity, actuation of the device at the start of inhalation, IFR of <60 l min−1 followed by a 10-s breath-hold at the end of inspiration [16, 18].
An assortment of different spacer tubes, valved holding chambers and mouthpiece extensions have been developed to eliminate co-ordination requirements, reduce the 'cold Freon®' effect (when the below-freezing spray temperature causes some patients to stop inhaling) and reduce the amount of drug deposited in the oropharynx by decreasing the particle size distribution and slowing the aerosol's velocity. The aerosol from a MDI and holding chamber is finer than that with the MDI alone, with an approximate 25% decrease in the MMAD compared with the original aerosol [19, 20]. This finer aerosol is more uniformly distributed in the normal lung, with increased delivery to the peripheral airway. However, in patients with airway obstructions, the addition of a holding chamber to the MDI does not change the distribution of the aerosol [21, 22].
Breath-actuated MDIs have also been developed to eliminate co-ordination difficulties by firing in response to the patient's inspiratory effect. In patients with poor MDI technique, the breath-actuated pressurized inhaler, Autohaler™ (3M Pharmaceuticals, Minnesota, USA), increased lung deposition from 7.2% (with a conventional MDI) to 20.8% of the dose [23]. However, breath-actuated MDIs do not help patients who stop inhaling at the moment of actuation, nor do they improve lung deposition in those patients with good MDI technique. In addition, the oropharyngeal dose remains the same as for the MDI device. Patients preferred using the Autohaler™ to the MDI even though clinical outcomes were the same [24].
Because of their deleterious effect on the ozone layer, CFCs have been banned by the United Nations as per the Montreal Protocol [25]. HFAs, specifically HFA-134a and HFA-227, have been identified as suitable substitutes for CFC propellants. HFA does not contain chlorine and thus has no ozone-depleting potential. HFA and CFC propellants possess different physical and chemical properties. The spray temperature in HFA-MDIs is above freezing compared with CFC-containing MDIs, which may diminish the 'cold Freon®' effect caused by the CFC propellant. The jetting force and velocity of the spray are also reduced which may decrease the oropharyngeal dose [26]. Surfactants have been used in CFC-based MDIs to achieve the desired physical stability of suspensions, prevent aggregation of drug particles and to lubricate the metering valve. Surfactant solubility is reduced in HFA formulations. With some HFA formulations, ethanol has been used to dissolve the surfactant in the new propellants. Some steroids, such as beclomethasone dipropionate (BDP), are soluble in ethanol forming a solution rather than a suspension, as in the CFC-based formulation. These HFA-drug solutions can produce aerosols with a superior fine particle fraction and can achieve much greater deposition into the periphery of the lung [27]. HFA-BDP solution (QVAR; 3M Pharmaceuticals, St Paul, MN, USA) is an extra-fine solution aerosol with a MMAD of 1.1 µm with 97% of the dose deposited below the larynx within the respirable range (< 4.7 µm). In comparison, CFC-BDP suspension (Beclovent; GlaxoSmithKline, Middlesex, UK) has a MMAD of 3.2 µm with 77% of the dose deposited below the larynx in the respirable range [26]. In vivo deposition studies in subjects with asthma found 53% of the emitted dose of QVAR deposited in the lung, 3.2-fold greater than the 16.2% of the emitted dose of Beclovent [27, 28]. Efficacy and safety studies have demonstrated that the small particle size of the HFA-BDP solution should be prescribed at half the dose of the CFC-BDP formulation to achieve a similar therapeutic effect without additional side-effects [29].
Dry powder inhalers
DPIs were designed to eliminate the co-ordination difficulties associated with the MDI. Renewed interest in this delivery system has emanated from the urgency to eliminate CFC-containing MDIs [25]. There is a wide range of DPI devices on the market, from single-dose devices loaded by the patient (e.g. Aerolizer, Rotahaler) to multiunit dose devices provided in a blister pack (e.g. Diskhaler), multiple unit doses sealed in blisters on a strip which moves through the inhaler (e.g. Diskus) or reservoir-type (bulk powder) systems (e.g. Turbuhaler).
Lung deposition varies among the different DPIs. Approximately 12-40% of the emitted dose is delivered to the lungs with 20-25% of the drug being retained within the device [4, 6, 30]. Poor drug deposition with DPIs can be attributed to inefficient deaggregation of the fine drug particles from coarser carrier lactose particles or drug pellets. Slow IFR, high humidity and rapid, large changes in temperature are known to effect drug deaggregation and hence the efficiency of pulmonary drug delivery with DPIs [31, 32]. With most DPIs, drug delivery to the lungs is augmented by fast inhalation. For example, Borgstrom and colleagues [33] demonstrated that increasing the IFR from 35 l min−1 to 60 l min−1 through the Turbuhaler™ increased the total lung dose of terbutaline from 14.8% of nominal dose to 27.7%. This is in contrast to the MDI, which requires slow inhalation and breath holding to enhance lung deposition of the drug. The Spiros DPI, however, deposits significantly more in the lung when IFRs <30 l min−1 are used [34].
With DPIs, the drug aerosol is created by directing air through loose powder. Most particles from DPIs are too large to penetrate into the lungs due to large powder agglomerates or the presence of large carrier particles (e.g. lactose). Thus, dispersion of the powder into respirable particles depends on the creation of turbulent air flow in the powder container. The turbulent airstream causes the aggregates to break up into particles small enough to be carried into the lower airways and also to separate carrier from drug [35]. Each DPI has a different air flow resistance that governs the required inspiratory effort [36, 37]. The higher the resistance of the device, the more difficult it is to generate an inspiratory flow great enough to achieve the maximum dose from the inhaler [38-40]. However, deposition in the lung tends to be increased when using high-resistance inhalers (Figure 2) [19, 40-45].
Figure 2
Figure 2
Lung deposition measured by several different investigators from a variety of dry powder inhalers (DPIs) vs. the specific resistance of the DPI. The increase in deposition seen with the higher resistance devices may, in part, be a function of the degree (more ...)
Recent developments in DPI technology have focused on eliminating these problems. Active DPIs are being investigated that reduce the importance of a patient's inspiratory effort. By adding either a battery-driven propeller that aids in the dispersion of the powder (Spiros; Dura Pharmaceuticals, San Diego, CA, USA) or using compressed air to aerosolize the powder and converting it into a standing cloud in a holding chamber, the generation of a respirable aerosol becomes independent of a patient's inspiratory effort (Inhance Pulmonary Delivery System, Inhale Therapeutic Systems, San Carlos, CA, USA). However, DPIs in use today are breath actuated and are dependent on a patient's IFR of 30-130 l min−1 to achieve an aerosol within the respirable range [19].
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Bioequivalence of inhaled medications
Although the number of different classes of inhaled medications is small, there is an increasing number of products within each class. With the need to eliminate the use of CFC inhalers, new HFA formulations MDIs and DPI formulations are being introduced. Any one drug may have a number of different formulations and be packaged with various delivery devices. Each inhalational device, whether a nebulizer, DPI or MDI, generates its drug aerosol differently and thus, the particle size, respirable dose, lung deposition and distribution will also differ. Consequently, the same drug at the same nominal dose delivered from different devices or in different formulations may not be bioequivalent [46, 47].
The importance of matching a drug with the appropriate nebulization system is often ignored. There is considerable variation in performance among different nebulizer designs and not all brands of nebulizer are ideal for all drug therapies [48-50]. The drug output and particle size vary according to type and brand of nebulizer, initial drug volume, residual volume and flow rate of compressed gas [49, 51-53]. Many manufacturers of drug formulations for nebulization do not routinely recommend a specific nebulizer or nebulizer and compressor combination to be used with their medication. Consequently, a wide variety of nebulization systems may be used to aerosolize the drug. This can lead to variability in the dose delivered and the effectiveness of the therapy. Newman et al.[54] evaluated four brands of jet nebulizers at four different compressed gas flow rates for use with gentamicin. They found a 10-fold difference between the most and least efficient delivery systems, suggesting patients might be underdosing when using inefficient nebulizer systems. In another study evaluating the performance of 12 nebulizers for use in the treatment of cystic fibrosis (CF), the investigators reported a wide variation in efficiencies among nebulizers, with the percentage of efficiently aerosolized drug ranging from 30% to <5% of initial dose [55]. Johnson and colleagues [56] demonstrated that the bronchodilator response to salbutamol was significantly greater when nebulized using a nebulizer that produced an aerosol with a MMD of 3.3 µm compared with one that generated a larger size aerosol (7.7 µm). A possible explanation for this is that most of the 7.7 µm particle aerosol is deposited in the oropharyngeal region and larger airways and therefore does not reach the small airways, where the majority of β2 receptors are located, in sufficient quantity to induce an equivalent bronchodilator response to the finer particle aerosol. Physicians may need to adapt a prescription to the performance of the nebulizer available to their patient or determine the most efficient nebulizer/compressor system to ensure optimal therapeutic effectiveness of nebulized medications.
The reformulation of MDIs with HFAs and development of other alternative inhalation products, such as dry powder formulations, require manufacturers to demonstrate that these new products are bioequivalent to their CFC-based counterparts. For orally or intravenously administered drugs, bioavailability data are used to substantiate the claim of bioequivalence. However, unless the sensitivity of the drug assay is high, bioavailability studies to establish bioequivalence for inhaled drugs may be inappropriate [57, 58]. Concentrations in the blood and urine are low and generally minimally detectable after aerosol delivery. Unless charcoal is used to block gastrointestinal absorption, it can be difficult to differentiate what proportion was absorbed from the lung and what was absorbed from the oropharyngeal or gastrointestinal region.
HFA formulations of salbutamol, salmeterol and fluticasone propionate have been shown to be as effective and safe as existing CFC formulations at equivalent dosages, as demonstrated in Figure 3[59-61]. In contrast, HFA-based BDP MDI (QVAR; 3M Pharmaceuticals) has the same emitted dose as the CFC-BDP MDI, but has a greater fine particle fraction and lung deposition [27]. In a direct comparative study in moderately severe asthmatics, Davies and colleagues showed that a 12-week treatment with QVAR 800 µg daily resulted in an equivalent change in pulmonary function to 1500 µg of beclomethasone CFC-BDP MDI [62]. Thus, the HFA formulation of BDP provides equivalent control over asthma to the CFC formulation at half the daily dose, with a similar safety profile and a more uniform distribution in the lung.
Figure 3
Figure 3
Comparison of the bronchoprotective effects of three doses of chlorofluorocarbon (CFC) and hydrofluoroalkane (HFA)-based salbutamol in patients with mild asthma. Dose-response curves of salbutamol plotted against PC20 methacholine Proventil-HFA (more ...)
Guidance and recommendations for establishing equivalence testing of inhalers have been developed in Canada [63], USA [64, 65] and the UK [66]. There are four principal methods available for determining bioequivalence of different inhalational devices. (i) In vitro particle size distribution measurements. This evaluation has been favoured by regulatory authorities. However, there are several disadvantages of this methodology. Measurements are done under ambient conditions that are not representative of the humid environment of the lung. Drug aerosols have been shown to increase in size when exposed to humidity and therefore deposition would be altered. As well, the role of lung defence in clearing the drug from the lung is removed in these bench studies and therefore the predictions based on particle measurements may not correlate with the clinical efficacy and safety in patients [66]. (ii) Radioaerosol drug deposition studies, which assess the pattern of deposition of the aerosol in the lung and quantify its distribution. These studies should be done in patients with lung disease instead of healthy volunteers, as deposition patterns can change dramatically in the presence of airway obstructions. With direct radiolabelling of the drug, pharmacokinetic parameters could be measured. However, to date this has not been studied. (iii) Pharmacokinetic studies, which have limited value since the doses administered are small and difficult to measure in serum as the drug concentrations are often undetectable and do not necessary correlate with the dose delivered to the lung. Pharmacokinetic studies are usually conducted in healthy volunteers. As stated above, lung disease can not only alter deposition, but also have an effect on drug clearance from the lungs via mucociliary clearance and systemic absorption. (iv) Comparative pharmacodynamic, clinical efficacy and safety studies are the most reliable measure of beneficial and adverse effects of drugs [66]. However, many published comparative studies have been deficient in their study design and have had insufficient sample size to claim bioequivalence between inhalational devices [67]. In properly designed equivalence studies, the null hypothesis is not equivalence but inequivalence. Rejecting this hypothesis leads to the proper interpretation of both treatments being equivalent [68]. Another flaw, especially with bronchodilators, is that the clinical effects are often measured using doses on the plateau of the dose-response curve. As a result, no difference may be measured between inhalers and the conclusion of bioequivalence could be inaccurate [69].
The various government agencies have made recommendations as to how bronchodilator studies should be conducted in patients with asthma using a range of doses to establish a dose-response curve [63, 66]. The bioequivalence of clinical efficacy should be reported as relative potency (RP) and defined as a 90% confidence interval of the RP between 0.67 and 1.50 [58]. However, recommendations for bioequivalence of clinical efficacy of inhaled corticosteroids have not been made. The difficulties and lack of consensus in providing appropriate methodologies for comparing the clinical efficacy of different inhaled corticosteroids have been reviewed in a recent workshop [70]. Standards for bioequivalence evaluation of inhalational therapies need further development and implementation, especially as we move from treating local respiratory disease to treating systemic diseases with inhalable forms of medications.
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Inhaled drug formulations
Drug formulation plays an important role in producing an effective inhalable medication. Not only is it important to have a drug that is pharmacologically active, it must be efficiently delivered into the lungs, to the appropriate site of action and remain in the lungs until the desired pharmacological effect occurs. A drug designed to treat a systemic disease, such as insulin for diabetes, must be deposited in the lung periphery to ensure maximum systemic bioavailability. For gene therapy or antibiotic treatment in CF, prolonged residence in the lungs of the drug may be required to obtain the optimal therapeutic effect. Thus, a formulation that is retained in the lungs for the desired length of time and avoids the clearance mechanisms of the lung may be necessary. Research into dry powder formulations has been an area of growth in recent years and will be the focus of this section. Formulating dry powders for inhalation involves either micronization via jet milling, precipitation, freeze-drying or spray-drying using various excipients, such as lipids and polymers, or carrier systems like lactose. Each one offers its own unique advantages as well as disadvantages, depending on the therapeutic agent being formulated.
Lactose carrier systems
When drug particles are in the size range required for lung deposition, the surface electric forces associated with the particles exceed the gravitational force acting upon them, resulting in the development of cohesive powders with poor flow [71]. To overcome this problem, the drug is blended with a coarse carrier system (30-100 µm), such as lactose. At present, marketed dry powder inhalers contain either the drug alone or mixed with a bulk carrier, usually lactose (α-lactose monohydrate). Lactose is one of three sugars (the others being glucose and mannitol) allowed as carriers by the Food and Drug Administration.
Lactose has an established safety profile and improves the flow properties of the formulation necessary for reproducible filling and promoting dosing accuracy. The drug particles become bound by physical forces to active sites on the surface of the carrier particles (Figure 4) [72]. The mixture provides a degree of resistance to segregation caused by the interaction between lactose and drug molecules, but also allows for deaggregation of the drug from the carrier during drug delivery. Any drug still adhering to the lactose after the aerosol has been generated will be deposited in the oropharyngeal region, which has the potential to cause local side-effects, as in the case of inhaled corticosteroids.
Figure 4
Figure 4
Electron micrographs of a 2% albuterol sulphate-lactose powder blend. (a) Tomahawk-shaped lactose particle with drug on its surface. (b) The higher magnification shows individual drug particles (elongated crystals) on the lactose. (Reprinted with (more ...)
At present, the respirable fraction from most carrier-based DPI formulations is relatively low, only 10% of the total dose being delivered to the lower airways. To improve its delivery efficiency, the effect of increasing the surface roughness of lactose particles has been investigated. Kawashima and colleagues [73] showed that increasing the surface roughness does increase the emitted dose because the lactose can carry more drug particles, but it decreases the respirable fraction. Mixing fine lactose particles (approximately 5 µm) with coarse lactose does improve deaggregation, as well as increasing the fine particle fraction of the formulation. The drug particles interact with both the fine and coarse lactose particles. The liberation of the drug from the surface of fine lactose can occur more readily than from the coarser particles, since the smaller particles have a lower degree of surface roughness [74, 75]. Thus, the respirable fraction of the formulation is increased in the presence of the fine lactose particles, while the flowability of the powder is maintained by the coarse lactose particles. Another formulation approach is the modification of the carrier's surface properties or 'corrasion'[76, 77]. This is accomplished by adding a ternary component, such as L-leucine, to the formulation. It is hypothesized that the ternary agent decreases the strength of the interaction between the drug and carrier system by occupying high-energy binding sites that would otherwise be occupied by the drug particle. The lower energy binding sites on the carrier particles will bind the drug particles, thereby increasing the dispersibility of the drug, and thus the respirable fraction of the formulation.
Liposomes
Sustained release from a therapeutic aerosol can prolong the residence of an administered drug in the airways or alveolar region, minimize the risk of adverse effects by decreasing its systemic absorption rate, and increase patient compliance by reducing dosing frequency. A sustained-release formulation must avoid the clearance mechanisms of the lung, the mucociliary escalator of the conducting airways and macrophages in the alveolar region.
Liposomes, as a pulmonary drug delivery vehicle, have been studied for years and used as a means of delivering phospholipids to the alveolar surface for treatment of neonatal respiratory distress syndrome. More recently, they have been investigated as a vehicle for sustained-release therapy in the treatment of lung disease, gene therapy and as a method of delivering therapeutic agents to the alveolar surface for the treatment of systemic diseases.
Liposomes can be produced from a variety of phospholipids from natural sources, that are endogenous to the lungs and that carry either no charge or net negative or positive charges [78, 79]. They consist of an aqueous volume entrapped by one or more bilayers of natural or synthetic lipids with or without cholesterol. They exist either as a single lipid bilayer sphere (unilamellar) or as a multilayer vesicle (multilamellar) (Figure 5) [80]. They are capable of encapsulating a variety of drugs with a wide range of lipophilicities. Lipophilic drugs, such as corticosteroids, intercalate in the lipid bilayers, while others interact with the bilayer interface. Hydrophilic drugs, such as sodium cromoglycate, partition in the aqueous spaces (Figure 5). Drugs with intermediate solubility are poorly retained in liposomes. However, when the drugs are weak acids or weak bases, the pH and chemical composition of the internal aqueous compartments can be manipulated so that the drug concentrates in the interior of the liposomes, resulting in a high degree of retention [81].
Figure 5
Figure 5
Types of liposomes
Liposomes are very versatile drug carriers. They can be formed from a variety of lipids leading to a wide range of physiochemical properties that can alter the trapping efficiencies and release rates of drug. The physiochemical properties, such as liposome size, bilayer fluidity, surface charge, as well as the method of preparation, affect their in vivo behaviour. The vesicle size and number of bilayers are critical parameters in determining the circulating half-life and extent of drug encapsulation [82]. Small liposomes (≤ 0.1 µm) are opsonized less rapidly and to a lesser extent than large liposomes (> 0.1 µm) and therefore have a longer circulating half-life. Small liposomes also have a slower release rate. Niven et al.[83] demonstrated that 80 min after nebulization of the liposomal preparation, large multilamellar vesicles lost 77% of their content while vesicles with a diameter of 0.2 µm lost only 8%. The preferred size range for clinical applications has been suggested to be 50-200 nm in diameter. Liposomes of this size would avoid phagocytosis by macrophages and still trap useful drug loads [81].
Bilayer fluidity also influences the behaviour of liposomes. Lipids have a characteristic phase transition temperature (Tc). The Tc depends on the length and saturation of the fatty acid chains and can vary from −20 °C to 90 °C. The lipids exist in different physical states above and below this temperature. Below the Tc, the lipids are in a rigid, well-ordered arrangement (gel phase), and above the Tc in a liquid-crystalline state (fluid phase). The presence of high Tc lipids (Tc > 37 °C) makes the liposome bilayer less fluid at the physiological temperature and thus, less leaky. The fluidity of the bilayer appears also to influence the interaction of liposomes, with macrophages with high Tc lipids having a lower uptake [82]. Incorporation of cholesterol into the lipid bilayer affects the fluidity. At high concentrations (> 30 molar percentage), cholesterol can eliminate the phase transition, making the liposome more stable and less leaky [78, 84, 85].
The type and density of charge on the liposome surface are also important parameters. A negative charge decreases liposome aggregation and increases encapsulation efficiency; however, it increases liposome-cell interactions and charged liposomes may be cleared faster than neutral liposomes. Cationic liposomes have been studied extensively as a nonviral delivery vehicle for gene therapy, including the delivery of CFTR gene for CF [79, 86, 87]. The negatively charged genetic material is not encapsulated by the liposomes but complexes with cationic lipids by electrostatic interactions (Figure 6). Unlike anionic liposomes, cationic liposomes deliver their contents to cells by fusion with the cell membrane.
Figure 6
Figure 6
Atomic force microscopy (AFM) images of free plasmid DNA, free cationic liposomes (using a cholesterol derivative) and their complexes. The complex formation induces the changes in the shape and size of the liposomes as shown in (c) and (d). (a) Free (more ...)
Liposomes, like other inhaled particles reaching the alveoli, are cleared by macrophages. Unlike other inhaled particles, the fate of liposomes can have a similar fate to endogenous lipids. The processing, uptake and recycling of liposomal phospholipids occurs through the same mechanism as endogenous surfactant via the alveolar type II cells [78]. A search for a liposomal formulation that would evade the recognition and uptake of the immune system and prolong its residence led to the development of liposomes with a polymer surface coating, such as polyethylene glycol (PEG) (Figure 5). These formulations are known as sterically stabilized liposomes or Stealth™ liposomes (Sequus Pharmaceuticals, Menlo Park, CA, USA). The hydrophilic polymer coating attracts water to the liposome surface, preventing the association and binding of opsonins to the liposome, thereby inhibiting the body's molecular recognition processes of labelling the molecule as foreign for subsequent uptake and removal by macrophages, and subsequently extending its circulation time [81, 88, 89].
Liposomal formulations can also take advantage of opsonization by macrophages when treating intracellular respiratory infections, such as Francisella tularenis, which reside and multiply in macrophages. In a study of mice infected with F. tularenis, mice treated with liposome-encapsulated ciprofloxacin survived 15 days post infection compared with a 100% mortality rate in the group treated with free ciprofloxacin by day 9 post infection [90]. Since liposomes are naturally taken up by macrophages, the efficient intracellular delivery of the liposomes may account for the superior efficacy of the liposomal ciprofloxacin formulation.
Pulmonary delivery of liposomal formulations of antibiotics to treat respiratory infections also appears promising. Liposome-encapsulated tobramycin, gentamicin and ciprofloxacin have been investigated in several animal studies [90-92]. The studies have demonstrated sustained-release properties, as well as an increased susceptibility of bacteria to the liposome-encapsulated antibiotic. A prolonged pulmonary retention time was achieved with a liposomal formulation of tobramycin compared with free drug after intratracheal instillation into the lungs of rats infected with Pseudomonas aeruginosa. Sixteen hours post administration, liposomal tobramycin was still present in the lungs, compared with only 15 min with the free tobramycin formulation [92]. An increase in efficacy was also noted with a dry powder formulation of the liposome-encapsulated tobramycin. One, 3 and 6 h after intratracheal instillation into the lungs of a similar rat model of P. aeruginosa infection, there was no significant difference between drug groups. However, 16 h after treatment, the colony-forming unit (CFU) counts in the rats treated with the liposome formulation continued to decline from 1.4 - 106 CFU lung−1 at baseline, to 4.3 - 105 CFU lung−1 compared with an increase to 1.3 - 108 CFU lung−1 in the free tobramycin group [91].
Liposomal preparations of various antiasthma drugs, including salbutamol, sodium cromoglycate, terbutaline and corticosteroids, have been studied. Animal studies have demonstrated sustained-release profiles and reduced systemic effects [80]. A limited number of small clinical studies have reported similar findings. A pharmacokinetic study in five healthy adult volunteers comparing nebulization of a liposomal formulation of sodium cromoglycate with free drug found that the liposome-encapsulated formulation showed a prolonged lung retention [93]. Nebulization of two radiolabelled liposomal formulations of BDP to 11 healthy volunteers demonstrated good lung penetration and slow clearance of the radiolabel, suggesting that the liposomal formulation may act as a local sustained-release reservoir [94].
Most corticosteroids are lipophilic and should easily be incorporated in a liposome. This, however, has not turned out to be true for all corticosteroids. When diluted in a larger volume, whether a nebulizer chamber or even in the lung, corticosteroid liposomes tend to lose their contents in a biphasic pattern determined by their partition coefficient. For example, triamcinolone acetonide liposomes are stable in a lipid dispersion, but when diluted in the large aqueous volume of the lung the steroid is rapidly released from the liposomes and absorbed systemically. The formulation provided neither sustained release nor lung targeting. In contrast, a hydrophilic pro-drug triamcinolone acetonide phosphate encapsulated in liposomes produced a sustained-release profile, nearly doubling the duration of lung glucocorticoid receptor occupancy relative to the free drug [80].
With the versatility of liposomes as a drug carrier, in the future we may see liposomes playing a prominent role in pulmonary delivery for gene therapy, sustained-release preparations and for targeting specific cells to treat intracellular infections and local tumour cells.
Large porous particles
A new type of aerosol formulation is the large porous hollow particles, called Pulmospheres™ (Figure 7) [95]. They have low particle densities, excellent dispersibility and can be used in both MDI and DPI delivery systems [96]. These particles can be prepared using polymeric or nonpolymeric excipients, by solvent evaporation and spray-drying techniques [97]. Pulmospheres™ are made of phosphatidylcholine, the primary component of human lung surfactant. They are prepared in a two-step process. In the first step, an oil-in-water emulsion is prepared using oils, such as perflubron or perfluoroctyl ethane [96]. The oil phase serves as a 'blowing agent' during the spray drying step, retarding shrinkage of droplets while simultaneously creating pores in the particle surface. The second step in the preparation is the spray-drying of the emulsion.
Figure 7
Figure 7
Figure 7
Confocal microscopy images of large porous particles (pulmospheres). (A) Poly(lactic acid-co-glycolic acid) (PLGA) particle with diameter of 8.5 µm and density equal to 0.1 g cm−3. Double emulsion solvent evaporation techniques were used (more ...)
Traditional therapeutic powders consist of particles with a mass density of 1 ± 0.5 g cm−3 and mean geometric diameters of <5 µm to maximize peripheral deposition. Pulmospheres™ are lighter and larger than the typical dry powder particles with a mass density of approximately 0.4 g cm−3 and geometric diameter of >5 µm. By virtue of their hollow and porous characteristics, Pulmospheres™ give rise to smaller aerodynamic diameters than their geometric diameter. Because of their large size and low mass density, the particles can aerosolize more efficiently (less aggregation) than smaller nonporous particles, resulting in higher respirable fractions of the formulation. Cromolyn Pulmospheres™ have a respirable fraction (< 5 µm) of 68% compared with 24% with micronized cromolyn particles (Intal) [98]. An increase in the respirable fraction results in an increase in peripheral deposition and thus a lower dose requirement to achieve the same therapeutic effect. A decrease in oropharyngeal deposition of Pulmospheres™ also reduces the risk of local side-effects.
The large size of Pulmospheres™ allows them to remain in the alveolar region longer than their nonporous counterparts by avoiding phagocytic clearance. After intratracheal administration into rats, only 8% and 12.5% of macrophages contain Pulmospheres™ particles immediately and 48 h after inhalation, respectively, compared with 30% and 39% of macrophages containing nonporous particles during a similar time interval [95].
It has also been shown that Pulmospheres™ can increase systemic bioavailability of certain drugs. Edwards and colleagues [95] have also demonstrated an increase in systemic bioavailability of insulin and testosterone using this technology, making Pulmospheres™ attractive for systemic inhalation therapies, as well as for sustained-release therapies, with dosing every 1-2 days while avoiding local side-effects.
Biodegradable polymers
In addition to liposomes, biodegradable polymer microspheres are currently being studied as sustained-release pulmonary drug carriers. Polymers such as poly(lactic acid) (PLA) used in medical applications such as sutures, orthopaedic implants and medical dressings, and poly(glycolic acid) have been investigated. However, PLA may not be suited for pulmonary delivery because of their long biological half-life with dosing occurring once every few weeks. An in vitro release experiment demonstrated a sustained-release of BDP entrapped in PLA microspheres occurring over 6 days, and over 8 days for nedocromil sodium PLA microspheres [99]. Oligolactic acid, an oligomer of lactic acid, has a shorter biological half-life than PLA and therefore may be better suited for pulmonary drug delivery. Mucoadhesive polymer, hydroxypropyl cellulose (HPC), has been shown to prolong the pharmacokinetic and pharmacodynamic profile of crystalline BDP by avoiding mucociliary clearance [100]. Following powder aerosol administration to guinea pigs, cBDP/HPC microspheres were retained in the lungs longer than cBDP alone, with 86% of BDP remaining at 180 min compared with <20% with cBDP alone. Using BDP's inhibitory effect on airway eosinophil accumulation as a measure of efficacy, cBDP/HPC microspheres maintained drug action for 24 h compared with 1-6 h with cBDP. Although a limited amount of research has been published in this area, the sustained-release profiles achieved with corticosteroids appear promising. However, the toxicity of this type of formulation has not yet been established for pulmonary delivery.
Conclusions
As more efficient pulmonary delivery devices and sophisticated formulations become available, physicians and health professions will have a choice of a wide variety of device and formulation combinations that will target specific cells or regions of the lung, avoid the lung's clearance mechanisms and be retained within the lung for longer periods. The more efficient, user-friendly delivery devices may allow for smaller total deliverable doses, decrease unwanted side-effects and increase clinical effectiveness and patient compliance.
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