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Therapeutic Delivery
Formulation considerations for dry powder inhalers
The market for inhalable dry powder medication has consistently grown over past years. Targeting the lungs has been recognized to offer several advantages compared with oral application of drugs. The successive development of inhalation products has led to advances in local treatment of different respiratory diseases, but has also demonstrated the possibility to utilize the lungs for systemic drug delivery. Since a dry powder inhalation product is always a combination of drug formulation and inhalation device, the requirements for the development of such a system may be particularly complex. Therefore, this review aims to give an overview of the necessary considerations for a successful dry powder inhaler development.
Background Inhalation therapy for the treatment of respiratory diseases such as bronchial asthma, chronic obstructive pulmonary disease or cystic fibrosis have been shown to be superior in terms of efficacy and reduction of drugrelated side effects compared with their oral administration [1–3] . For local treatments, drug particles are able to reach their therapeutic target directly through inhalation and, therefore, a rapid clinical response and higher specific drug concentration can be achieved accompanied by a minimization of systemic adverse effects. As a result, current guidelines of the Global Initiative for Asthma (GINA) [4] and Global Initiative for Chronic Obstructive Lung Disease (GOLD) [5] suggest the use of inhalation products as first-line therapeutics. The inhalation route is also attractive for systemic delivery of drugs. The lung exhibits a large surface area with usually high membrane permeability and first-pass metabolism can be excluded [6] . Consequently, new developments are aiming towards the systematic administration of peptides, antibiotics, vaccines, anti-cancer substances or other drug classes and will broaden the inhalation market considerably in the near future.
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Eike Cordts1 & Hartwig Steckel*,1 1 Department of Pharmaceutics & Biopharmaceutics, Kiel University, Gutenbergstrasse 76, 24118 Kiel, Germany *Author for correspondence: [email protected]
However, the development of an inhalation drug product is a complex task as the correct administration of a drug is dependent on multiple factors. In order to apply the requested amount effectively to the lungs, the particles need to exhibit an aerodynamic diameter smaller than 5 μm [6] . The generation of an aerosol cloud with the desired properties brings up the necessity for further formulation steps to generate the active pharmaceutical ingredient (API). Second, the type of inhalation device used has a major impact on the generated respirable fraction. Inhalation devices for pulmonary application can be categorized into four groups: nebulizers, pressurized metered-dose inhalers (pMDIs) with or without spacer, soft mist inhalers (SMIs), and dry powder inhalers (DPIs). Active nebulizers convert the drug solution or suspension into an aerosol either by a vibrating mesh, an ultrasonic transducer or by atomization (air jet nebulizer). The aerosol is then inhaled via a mouthpiece or a facemask until the solution has been fully aerosolized. As no specific coordination of the inhalation maneuver and device actuation is necessary, this type of administration is particularly suitable for small children or
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Key Terms Aerodynamic diameter: Diameter of a sphere with a unit density of 1 g/cm³, having the same terminal settling velocity in still air as the particle in consideration. Adhesive mixture: Different components adhere to each other to form units and increase overall blend stability and homogeneity.
for ventilated patients. Furthermore, multiple drug solutions may be applied at the same time [7] . On the other hand, the need for a consistent power supply and the prolonged treatment time limits the usability in everyday life. Also, an inconsistent aerosolization has been reported as a shift in droplet size occurs over the nebulization duration [8,9] . Portable pMDIs are suitable for API solutions or suspensions and have been used since the 1950s with the introduction of the Medihaler [10,11] . For this costefficient product, the API is either dispensed in a solvent that is miscible with the propellant (e.g., ethanol) or is dispersed in it to form suspensions. Additional stabilizing and/or performance-modifying excipients may be added to the formulation before it is filled into a can with appropriate metering valve. The first pMDI systems were based on chlorofluorocarbon propellants (CFCs), but have nowadays mostly been replaced by the less environmentally harmful propellants of the hydrofluoroalkane-type (HFAs). Upon actuation the metered solution or suspension is forced through a spray orifice and gets torn apart into fine droplets, leaving solid particles in the micrometer range or particle agglomerates after the rapid evaporation of propellant. Since pMDI formulations usually contain surfactants, additional coating of the particles takes place at the same time. Advantages of this system, such as low costs and the generation of a very fine particle collective, are opposed by stability issues and the high exiting velocity of the aerosol cloud, which causes massive particle deposition in the oropharynx and a low respirable fraction. Also, issues including the correct coordination of MDI actuation and simultaneous inspiration [12] , even in healthcare professionals [13] , result in an unsatisfactory MDI efficacy through reduced deposition of medication in the lungs [14] . Lung doses of 21% [15] or less were repeatedly found in various studies. With the use of spacers [16] or breath-actuated systems, the coordination issues can be overcome, however, they are not implemented in the therapy on a routine basis. Another option was introduced to the market with Boehringer Ingelheim’s Respimat® SMI. It eradicates many of the disadvantages of pMDIs by generating a more slowly moving aerosol cloud of longer duration and smaller droplet sizes out of the aqueous or ethanolic API solution. As a result, higher lung depositions
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can be obtained [17] even in patients with poor inhalation technique [15] . However, the current Respimat® is a non-reusable device, which ends up in a significant increase in cost of treatment. These days, DPIs are widely accepted to deliver diverse types of medication. They first appeared in the 1970s when Bell et al. had developed Fisons Spinhaler®, a capsule-based inhaler used to administer 20 mg of sodium cromoglycate [18] . However, interest in the use of DPIs became especially evident with the ratification of the 1987 Montreal protocol, in which the participating countries agreed to phase out CFC-propellants as they were identified to accelerate ozone layer depletion. Consequently, existing pMDI formulations had to be transferred to alternative systems. Besides replacement for pMDIs, DPIs obliterate many limitations of the previously mentioned systems. They feature a portable device, for which the dispersion energy is generated by the patient’s inhalation maneuver itself, rather than an external power supply or propellant. Therefore, today’s marketed DPIs are breath-actuated systems that demand less coordination of the inhalation maneuver and dose actuation. Most of the DPIs on the market are built as multiple use devices, either equipped with a formulation reservoir and dose counter or the ability to insert new capsules or blisters prior to inhalation. By doing so, the inhaler can easily be adapted to improve stability issues of the formulation or be adjusted to patient demands. The development of a DPI product is a very complex subject as the overall performance of the medication is influenced by a large number of factors and their interactions. This review aims to give an overview of necessary considerations for a successful dry powder inhaler development. Starting with properties of the drug itself, further process and formulation steps will be presented and be evaluated together with considerations for the correct choice of inhalation device. Powder formulation Starting with the pure drug substance, first considerations need to be made about different aspects of the final inhalation formulation. Among inhaler design and the patient’s inhalation maneuver, the drugcontaining powder formulation itself is arguably the starting point for a proper DPI development. As a consequence of the need to use drug particles below 5 μm in size, these powders are strongly cohesive and show very poor flowability. In order to achieve an accurate metering of the desired small quantities, the drugs are either formulated with additional excipients or processed to soft spheres through spheronization. The goal of both is to improve powder flowability, which is essential to allow volumetric filling of the
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Formulation considerations for dry powder inhalers
device reservoir or capsules and blister cavities, respectively. The addition of coarse lactose increases the total mass and therefore the metering volume, leading to fewer variations in dosing. Furthermore, the formation of an adhesive mixture improves the flowability significantly, as the used lactose is generally larger in size. A similar effect can be obtained through controlled agglomeration (spheronization) of the micronized drug particles. The increase in agglomerate size accompanies a decrease in interagglomerate forces and, hence, an increase in flowablity. This approach was first patented by Fisons Limited [19] , further developed by Astra Zeneca [20] , and is currently used for products in the Turbuhaler® (Astra Zeneca, UK) and Twisthaler® (Merck & Co., USA) devices. Likewise, Hartmann et al. developed a continuous process to produce comparable spheres through a vibration technique [21] .
into the lungs [27] . Target receptors are not always distributed uniformly throughout the lung. For example, Usmani et al. demonstrated in their study that variations to the particle size of inhaled albuterol lead to a shift in deposition region and patient response. [28] . Furthermore, for the treatment of infectious lungs diseases, or for a possible systemic uptake, it may be necessary to administer API particles uniformly throughout different lung regions [6] . Particles in the respirable size range can be generated via various processing methods. The most commonly used techniques will be highlighted in the following section. Micronization
The administered amount of API in inhalation products needed for an adequate treatment can range from as low as 6 μg of formoterol (e.g., Formatris® Novolizer® [Meda Manufacturing, Germany]) up to 125 mg of colistimethate in Colobreathe Turbospin® (Forest Laboratories, UK) per capsule. However, the majority of DPIs on the market are designated for the treatment of asthma or COPD and, therefore, amounts of 6–500 μg per dose are currently the most common. In recent years, there has been a consistent interest in high-dose deliveries as needed for the administration of antibiotics in tuberculosis, cystic fibrosis or other infectious lung diseases. Novartis managed to expand their treatment of nebulized tobramycin to a high-dose DPI treatment (TOBI® Podhaler®) with a single API dose of 112 mg, which is divided into four 28 mg (approximately 45 mg including excipients) capsules. In addition, another treatment option for patients with cystic fibrosis has recently been introduced to the market with the Colobreathe® Turbospin® Inhaler from Forest Laboratories, UK. The treatment with the encapsulated colistimethate sodium (125 mg twice daily) has been shown to be as effective as the administration of nebulized tobramycin [22] . Improvements in formulation and inhaler design (Twincer® [University of Groningen, The Netherlands] [23,24] , Orbital® [Pharmaxis Ltd., Australia] [25]) may simplify the administration and expand this route of administration to other drugs [26] .
Micronization of the raw API starting material via milling is arguably the most straight forward approach to obtain respirable particles in the lower micrometer size range. Size reduction can efficiently be achieved with different types of mills (e.g., fluid-energy mills, ball mills); however, air jet milling is used on a routine basis. The starting material is injected into the milling chamber, in which air or nitrogen is used to establish a typical grinding pressure of 3–10 bar. The accelerated powder gets comminuted as a result of particle-particle or particle-wall collisions with subsequent breakage until the particles are small enough to discharge to the classification chamber. Due to the rapid expansion of the introduced gas, the thermal stress is low in spite of the high energy input [29] . However, irrespective of the used milling technique, the control of critical parameters is very limited. As a result, important product characteristics, such as size, shape, morphology, surface properties, and electrostatic charge cannot be controlled in a sufficient manner. Also, thermodynamically unstable amorphous parts within the crystalline structure may be generated [30–32] and surface energies and/or attachment energies may vary strongly due to the breakage of the natural crystal surfaces. Further, recrystallization of the amourphous parts may result in particle growth and other uncontrolled changes to the powder during storage. Kubavat et al. [33,34] found that even changes in crystallization procedures for the starting materials lead to different properties of the further micronized particles. All the above mentioned parameters may become critical during the continuous development and production of a dry powder inhalation product as interparticle forces play a key role in the de-agglomeration behavior.
API
Spray drying
Particle size is a driving factor with respect to deposition site in the lungs. The smaller the aerodynamic diameter of the particles, the deeper they may penetrate
The generation of fine particles using spray drying (SD) can help to overcome some of the drawbacks of the micronization process. The technique can be used
Dose
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Review Cordts & Steckel to generate particles with a narrow size distribution out of a solution, suspension or colloidal dispersion. The obtained particles reveal a more homogenous surface energy distribution and generally show a reduced area of contact between particles, thus leading to increased flowability and dispersion efficiencies (i.e., respirable fractions) [35] . Moreover, particle properties, such as density or morphology, can be customized by changing the process and/or formulation parameters [35–37] . However, spray drying often leads to the generation of amorphous particles. Modifications to the process set-up (introduction of a secondary drying step) [38] or change of material (type of carrier) [36] may be suitable approaches to overcome these challenges. In any case, a well investigated SD process and formulation is essential to end up with a desired product and yield. As mentioned before, Novartis managed to introduce the first dry powder treatment with tobramycin to the market by making use of spray drying in order to create their formulation. In this approach, the formulation containing volatile ‘blowing agents’ is spray-dried to produce large porous particles out of an oil-in-water emulsion. This way, powder flowability and deagglomeration is improved and no additional mixing step is needed. Also, a flow rate-independent, high-dose deposition in the lungs has been demonstrated [39] . Furthermore, SD was found to be a suitable technique to produce dry powder protein formulations for inhalation [40] , and additional excipients, such as mannitol or trehalose, may be added to prevent possible degradation during processing and storage [41] . Spray freeze drying
Spray freeze drying (SFD) is a two-step process, during which the feed solution is sprayed into a freezing medium (usually liquid nitrogen). Eventually, the liquid nitrogen evaporates and the frozen particles can be lyophilized in a subsequent freeze drying step. The typical process yield is >95% [42] and the obtained particles are of light and porous nature with improved aerosolization characteristics compared with the micronized material [42,43] . Furthermore, particle properties can again be modified through adjustments of the process (atomization) or formulation set-up [44,45] . Generally, this technique can be used to create powders for various treatments, however, the time consuming and very expensive process narrows the use down to high-cost products, such as protein or antibody formulations. Furthermore, the ability to formulate liposomes [43] or encapsulated particles (e.g., in PLGA) [46] opens up a new set of formulation options for inhaled therapy. Sufficient stability of the material (protein) to shear, freezing and dehydration stresses is essential. Protein
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stability may be increased by adding sugars to the liquid feed, which can then build a glassy matrix and protect the protein during the process [47] . Supercritical fluid technology
A liquid or gaseous phase can be transferred into its supercritical state by rising the pressure and temperature above their specific critical values. The obtained supercritical fluids (SCF) present an intermediate state that combines properties of both, gaseous (low viscosity) and aqueous (density) phase. Consequently, its often advantageous ability to serve as solvent or antisolvent for several drug substances offers a promising alternative approach for the generation of particles in the respirable size range [48,49] . Practically, supercritical carbon dioxide (above 31°C, 74 bar) can be used synonymously for SCF as it is commonly used due to its low cost and non-toxicity. Particles can be generated mainly via three distinctive SCF techniques: rapid expansion of supercritical solutions (RESS), particles from gas-saturated solutions (PGSS) and supercritical anti-solvent (SAS). Since the latter method is independent on the drug substance’s SCF solubility, this process is the most commonly used and different anti-solvent variants (e.g., aerosol solvent extraction system [ASES] and solution enhanced dispersion by supercritical fluids [SEDS]) have successfully been developed [50] . Particles generated with such techniques are usually very uniform in size and exhibit a platelet-like morphology. They show improved flowability and smaller cohesive– adhesive interactions compared with jet-milled powders [51] . Hence, those powders possess enhanced dispersion, that is, increased respirable particle fractions [52] , especially at low airflow rates or low aerodynamic turbulent stresses [49,51] . The bottom-up generation via SCF precipitation has been shown to be a suitable approach for the production of low (water) soluble-drug and peptide particles. It is still subject of intense research and offers a promising alternative to SD and SFD methods. However, its complex and expensive set-up has so far limited its use on a routine basis. Recently, Allergan Inc. have failed to gain US FDA approval for their Levandex® pMDI migraine treatment, which contains dihydroergotamine particles and was meant to be the first commercially available product based on SCF-processed particles [53] . Since the decline was not associated with issues about the processing technology itself, SCF products are still likely to enter the market soon. Carrier
Since micronization of API is still the most commonly used method to obtain particles in the respirable size
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range, further formulation steps to improve flowability are necessary. The aim is always to use the adhesive forces of the micronized powder to build up agglomerates that are larger in size and, therefore, exhibit improved powder flowability. For such a process, the control over agglomerate strength, that is, the magnitude of interparticle forces, is fundamental. On the one hand, agglomerates need to be mechanically stable during storage, filling and dosing, on the other hand, the thoroughly dispersion into single particles upon inhalation is essential to achieve reasonable respirable fractions of the drug. Most of the DPIs on the market contain a powder formulation with additional carrier material. In theory, the micronized drug binds to the host crystal surfaces to form an adhesive mixture [54] and gets separated again as single particles during the inhalation maneuver. As carrier material, different sugars or sugar alcohols such as glucose monohydrate, mannitol, sorbitol, maltitol, xylitol, trehalose [55–58] and others have been tested. Nevertheless, lactose monohydrate is currently the carrier of choice for almost all current DPIs on the market [59] . The reasons for this are manifold; lactose has a well-established safety and stability profile (generally recognized as safe [GRAS] excipient), production is inexpensive, various qualities are easily available on the market, it is less hygroscopic compared with alternative sugars and the knowledge base around its physico-chemical properties is quiet extensive [60] . It presents a suitable carrier for the majority of API molecules, however, lactose may not be applicable for formulations containing substances that are incompatible with its reducing sugar function (especially peptides). A typical carrier size ranges between 50 and 200 μm in diameter and the payload of drug is usually between 0.1 and 4% by weight [61] . However, de-agglomeration efficiency of the mixture is dependent on a large number of factors and their interactions. Extensive research has been performed to understand the influence of particle size, particle morphology, physico-chemical properties (e.g., pseudopolymorphs [62] and amorphous contents), dose [61,63] and environmental conditions on the de-agglomeration behavior, but the very complex relationships are still not fully understood. Among others, one strongly influencing factor is the addition of a third component to the binary API/carrier mixtures. There are numerous publications about the increased respirable API fraction with the addition of a ternary fine (mostly micronized lactose) component. Supposed theories for the enhancement are the saturation of ‘active sites’ on the carrier surface with excipient fines (regions of increased adhesive forces) [64,65] , and the formation of drug/fine excipient agglomerates with beneficial overall dispersion properties [66,67] . It needs
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to be considered that the supplementation of the mixture with ternary fines is accompanied by an increased complexity of variables. Thus, further considerations about the type of ternary component, fines content [68] , size and morphology of the fine particles are necessary [69] . For completeness, additional hypotheses about the mechanistic impact of addition of ternary fines to an adhesive powder mixture should be mentioned, such as the ‘buffer theory’ [70] or ‘fluidization theory’ [71] ; however, they were only applied to a limited number of formulations. For a new DPI development, a literature research should be performed carefully and thoroughly. Much of the data may not be directly applicable to the carrier formulation, as small variations in the formulation and the use of different inhalation devices may change the outcome significantly [68,72] . Blending
In general, the primary goal of a mixing process is the homogeneous distribution of drug particles within the excipient matrix in order to achieve acceptable content uniformity by volumetric dosing or filling. More precisely, for an adhesive inhalation mixture, agglomerates of cohesive drug particles (and excipients fines) need to be broken up and be distributed evenly across the carrier surface. Once adhered to their host crystals, the segregation tendency of an adhesive mixture is comparably low. Primarily, two types of mixers can be utilized to produce such blends: tumbling blenders and high shear mixers. The extent of energy input of the two mixers during blending varies significantly and also so does the magnitude of the driving diffusive, convective and shear mixing effects. Consequently, appropriate mixing speeds and process times for the production of homogeneous blends also differ significantly. However, mixing is a dynamic process. Particles may be redistributed across the carrier surfaces, become part of newly formed agglomerates, or may even act as carriers themselves [73] over the time of the mixing process. Consequently, the mixing efficiency is affected by multiple factors, such as particle size, morphology, density and other physico-chemical properties, and needs to be evaluated carefully for each new drug or component. Perceiving blending as a dynamic process leads to the further understanding that despite a homogenous distribution, redistribution effects of the particles may also be of importance for the correct choice of process parameters. Thus, homogeneity, in other words uniformity, should not be looked at exclusively for an inhalation product. In particular, high shear mixing exerts a high mechanical stress on the larger carrier particles and may
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Review Cordts & Steckel cause breakage or attrition or even changes to the solid state of the drug (amorphization). In addition, prolonged mixing leads to a constant redistribution of particles and may hinder the detachment of drug particles from the host crystal surfaces (‘press-on forces’) [74] . Again, the effects of changes in the mixing protocol may or may not be of overall importance to the DPI as the impact of other formulation variables (especially the inhalation device) may be dominant in the final formulation [75] . In summary, no specific recommendation on the type of mixer can be given at present. Short process times may be advantageous for high shear mixers compared with low shear tumble mixers, however, research about the influence of differences in energy input on inhalation formulations has just moved into focus in recent years [73,75] . Combination products
Both the GINA guidelines for the treatment of patients with asthma and the GOLD guidelines for COPD patients mention the use of long-acting β2-receptor agonists (LABA) in addition to inhaled corticosteroids (ICS) for patients whose treatment with ICS alone is insufficient. This way, two fundamental processes of the diseases can be addressed: bronchoconstriction and inflammation. It was found that the co-administration of the two types of drugs in one formulation resulted not only in increased patient convenience, but also showed improved β2-adrenoceptor and glucocorticoid receptor function, as well as a reduction in inflammation and, potentially, remodeling in the airways and lung due to synergistic effects [76] . For example, the co-administration of fluticasoneproprionate (ICS) and salmeterolxinafoate (LABA) was found to be superior in terms of clinical efficacy to the separate treatment with the two drugs [77] . In a Cochrane meta-analysis of 48 studies including 15,155 patients [78] the authors concluded that the combination therapy is modestly more effective in reducing
the risk of exacerbations requiring oral corticosteroids than a higher dose of inhaled ICS alone. Combination therapy also led to a greater improvement in lung function, symptoms and use of rescue β2-agonists than with a higher dose of inhaled corticosteroids. For children under the age of 12, combination therapy did not contribute to a beneficial treatment, but rather showed a trend towards an increase in adverse effects. Combination products are gaining popularity. An excerpt of marketed products can be found in Table 1. Filling
As most of the DPIs on the market are used for the treatment of bronchial asthma or COPD, the mass of a single dose usually ranges between 5 and 25 mg of a carrier blend. Most commonly, the powder blend is metered into hard gelatin or hydroxypropylmethylcellulose (HPMC) capsules that need to be inserted into the inhalation device and punctured by the patient right before inhalation. Alternatively, powder can be metered into blister strips (single or multi-dose blister inhaler), or larger amounts can be filled into the powder reservoir of a device itself. With the exception of reservoir devices, filling is equivalent to the dosing step. Volumetric dosing is mostly used for inhalation powders, and different approaches can be followed to consistently meter the claimed mass within the specifications (3–5% RSD) [79] . The majority of powder formulations reveal poor to moderate flow properties as a result of their fines content, which results in further challenges for the filling equipment. A well characterized powder (flow properties, poured and tapped densities, compressibility, geometric properties, hygroscopicity, electrostatic charging and adhesive forces) is absolutely essential in order to repeatedly achieve exact dosing by volume. In addition to methods listed in pharmaceutical pharmacopoeias, advanced analysis may be beneficial to understand dynamic and bulk
Table 1. Currently available dry powder combination products.
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Active pharmaceutical ingredient combination
Formulation type
Marketed product
Salmeterolxinafoate + fluticasoneprorionate
Adhesive mixture
Seretide® Diskus® Multi-dose blister AirFluSal® Multi-dose blister Single-dose blister Forspiro® Rolenium® Elpenhaler®
Formoterolfumarate dihydrate + budesonide
Soft spheres
Symbicort® Turbuhaler®
Multi-dose reservoir
Formoterolfumaratede hydrate + beclometasone dipropionate
Adhesive mixture + magnesium stearate
Foster® NEXThaler®
Multi-dose reservoir
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Device type
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properties [80] such as shear cell and powder rheology measurements [71] . Filling of capsules and blisters can generally be approached with two techniques – either a metered part of the bulk powder is transferred into an empty capsule or blister (‘dosing’), or the blister cavity acts as the metering chamber itself and is completely filled (‘filling’). For the latter, the empty cavities are dipped into a well-conditioned powder stock upside down, followed by taking off any excess material and are finally rotated back before sealing. Thus, this process can only be used for powder with appropriate interparticle forces (cohesive powders); otherwise the powder would not remain in the reversed cavities during the filling step. A second technique is the filling of blister cavities with the help of low pressure and membrane filters. The powder is forced through a capillary by the low pressure until all excess air in the cavity has been replaced by the formulation (Figure 1A) . With this approach, sealing faces do not get in contact with the powder formulation. Hence, additional cleaning steps are not required prior to the final blister sealing. More commonly, capsules contain only small quantities of the formulation. Different techniques can be used to transfer these small amounts into the capsule. With a ‘dosator and pin system’ the dosator dips into a loosely packed powder bed and is filled through its open end. A pin inside the dosator compresses the powder, which is subsequently transferred to and ejected into the capsule shell [81] . The machine can be equipped with a vacuum system for powders that fail to remain within the dosator tube during handling (Figure 1C) . Alternatively, a vacuum drum filling system (OMNIDOSE® [Harro Höfliger, Germany]) [82] can be used and is a promising option for the filling of highly cohesive material. The drum with metering cavity sucks in the powder at the 12 o’clock position and rotates to the 6 o’clock position to release the dose with positive pressure into the waiting capsule. During this rotation, excess powder on the drum surface gets removed with a scraper. Finally, the cavity gets cleaned again by another air jet when rotating back to the filling position (Figure 1B) . Equipment based on other filling technologies, such as ‘rotating dosing plate and tamping pin’ [81] , the newly developed ‘vibrating capillary approach’ [83,84] , and others, are also available. Device selection A diversity of DPI devices is currently commercially available [85] . They can be roughly categorized into four types: single-dose devices (capsule or blister based), multi-dose blister, multi-dose reservoir systems, and disposable devices for single use.
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All of them are breath-actuated or passive devices meaning that the energy needed for powder dispersion is generated through the patient’s inspiratory airflow. The first decision for a type of device is usually driven by drug and formulation considerations. However, even if formulation aspects and device design are closely monitored and controlled, the correct handling by the patient is essential for a successful administration of drug. Consequently, an inhalation device should be easy to use and give direct feedback of the successful inhalation maneuver to the patient. Many inhalation devices are under development; some of them are claimed as active devices that include additional features (compressed air, vibrators, electric motors) to generate the aerosol cloud making it independent from the patient’s inspiratory maneuver. To date, only the Exubera® (Pfizer, USA) [86] active inhaler for administration of spray-dried insulin had been introduced to the market in 2006, but it was withdrawn again in 2007 due to a lack of consumer acceptance for the product and low insulin bioavailability [87] . Other active devices are still under development and may be beneficial for systemic treatments due to increased dispersion efficiency of particularly small particles (
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Therapeutic Delivery
Formulation considerations for dry powder inhalers
The market for inhalable dry powder medication has consistently grown over past years. Targeting the lungs has been recognized to offer several advantages compared with oral application of drugs. The successive development of inhalation products has led to advances in local treatment of different respiratory diseases, but has also demonstrated the possibility to utilize the lungs for systemic drug delivery. Since a dry powder inhalation product is always a combination of drug formulation and inhalation device, the requirements for the development of such a system may be particularly complex. Therefore, this review aims to give an overview of the necessary considerations for a successful dry powder inhaler development.
Background Inhalation therapy for the treatment of respiratory diseases such as bronchial asthma, chronic obstructive pulmonary disease or cystic fibrosis have been shown to be superior in terms of efficacy and reduction of drugrelated side effects compared with their oral administration [1–3] . For local treatments, drug particles are able to reach their therapeutic target directly through inhalation and, therefore, a rapid clinical response and higher specific drug concentration can be achieved accompanied by a minimization of systemic adverse effects. As a result, current guidelines of the Global Initiative for Asthma (GINA) [4] and Global Initiative for Chronic Obstructive Lung Disease (GOLD) [5] suggest the use of inhalation products as first-line therapeutics. The inhalation route is also attractive for systemic delivery of drugs. The lung exhibits a large surface area with usually high membrane permeability and first-pass metabolism can be excluded [6] . Consequently, new developments are aiming towards the systematic administration of peptides, antibiotics, vaccines, anti-cancer substances or other drug classes and will broaden the inhalation market considerably in the near future.
10.4155/TDE.14.35 © 2014 Future Science Ltd
Eike Cordts1 & Hartwig Steckel*,1 1 Department of Pharmaceutics & Biopharmaceutics, Kiel University, Gutenbergstrasse 76, 24118 Kiel, Germany *Author for correspondence: [email protected]
However, the development of an inhalation drug product is a complex task as the correct administration of a drug is dependent on multiple factors. In order to apply the requested amount effectively to the lungs, the particles need to exhibit an aerodynamic diameter smaller than 5 μm [6] . The generation of an aerosol cloud with the desired properties brings up the necessity for further formulation steps to generate the active pharmaceutical ingredient (API). Second, the type of inhalation device used has a major impact on the generated respirable fraction. Inhalation devices for pulmonary application can be categorized into four groups: nebulizers, pressurized metered-dose inhalers (pMDIs) with or without spacer, soft mist inhalers (SMIs), and dry powder inhalers (DPIs). Active nebulizers convert the drug solution or suspension into an aerosol either by a vibrating mesh, an ultrasonic transducer or by atomization (air jet nebulizer). The aerosol is then inhaled via a mouthpiece or a facemask until the solution has been fully aerosolized. As no specific coordination of the inhalation maneuver and device actuation is necessary, this type of administration is particularly suitable for small children or
Therapeutic Delivery (2014) 5(6), 675–689
part of
ISSN 2041-5990
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Key Terms Aerodynamic diameter: Diameter of a sphere with a unit density of 1 g/cm³, having the same terminal settling velocity in still air as the particle in consideration. Adhesive mixture: Different components adhere to each other to form units and increase overall blend stability and homogeneity.
for ventilated patients. Furthermore, multiple drug solutions may be applied at the same time [7] . On the other hand, the need for a consistent power supply and the prolonged treatment time limits the usability in everyday life. Also, an inconsistent aerosolization has been reported as a shift in droplet size occurs over the nebulization duration [8,9] . Portable pMDIs are suitable for API solutions or suspensions and have been used since the 1950s with the introduction of the Medihaler [10,11] . For this costefficient product, the API is either dispensed in a solvent that is miscible with the propellant (e.g., ethanol) or is dispersed in it to form suspensions. Additional stabilizing and/or performance-modifying excipients may be added to the formulation before it is filled into a can with appropriate metering valve. The first pMDI systems were based on chlorofluorocarbon propellants (CFCs), but have nowadays mostly been replaced by the less environmentally harmful propellants of the hydrofluoroalkane-type (HFAs). Upon actuation the metered solution or suspension is forced through a spray orifice and gets torn apart into fine droplets, leaving solid particles in the micrometer range or particle agglomerates after the rapid evaporation of propellant. Since pMDI formulations usually contain surfactants, additional coating of the particles takes place at the same time. Advantages of this system, such as low costs and the generation of a very fine particle collective, are opposed by stability issues and the high exiting velocity of the aerosol cloud, which causes massive particle deposition in the oropharynx and a low respirable fraction. Also, issues including the correct coordination of MDI actuation and simultaneous inspiration [12] , even in healthcare professionals [13] , result in an unsatisfactory MDI efficacy through reduced deposition of medication in the lungs [14] . Lung doses of 21% [15] or less were repeatedly found in various studies. With the use of spacers [16] or breath-actuated systems, the coordination issues can be overcome, however, they are not implemented in the therapy on a routine basis. Another option was introduced to the market with Boehringer Ingelheim’s Respimat® SMI. It eradicates many of the disadvantages of pMDIs by generating a more slowly moving aerosol cloud of longer duration and smaller droplet sizes out of the aqueous or ethanolic API solution. As a result, higher lung depositions
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can be obtained [17] even in patients with poor inhalation technique [15] . However, the current Respimat® is a non-reusable device, which ends up in a significant increase in cost of treatment. These days, DPIs are widely accepted to deliver diverse types of medication. They first appeared in the 1970s when Bell et al. had developed Fisons Spinhaler®, a capsule-based inhaler used to administer 20 mg of sodium cromoglycate [18] . However, interest in the use of DPIs became especially evident with the ratification of the 1987 Montreal protocol, in which the participating countries agreed to phase out CFC-propellants as they were identified to accelerate ozone layer depletion. Consequently, existing pMDI formulations had to be transferred to alternative systems. Besides replacement for pMDIs, DPIs obliterate many limitations of the previously mentioned systems. They feature a portable device, for which the dispersion energy is generated by the patient’s inhalation maneuver itself, rather than an external power supply or propellant. Therefore, today’s marketed DPIs are breath-actuated systems that demand less coordination of the inhalation maneuver and dose actuation. Most of the DPIs on the market are built as multiple use devices, either equipped with a formulation reservoir and dose counter or the ability to insert new capsules or blisters prior to inhalation. By doing so, the inhaler can easily be adapted to improve stability issues of the formulation or be adjusted to patient demands. The development of a DPI product is a very complex subject as the overall performance of the medication is influenced by a large number of factors and their interactions. This review aims to give an overview of necessary considerations for a successful dry powder inhaler development. Starting with properties of the drug itself, further process and formulation steps will be presented and be evaluated together with considerations for the correct choice of inhalation device. Powder formulation Starting with the pure drug substance, first considerations need to be made about different aspects of the final inhalation formulation. Among inhaler design and the patient’s inhalation maneuver, the drugcontaining powder formulation itself is arguably the starting point for a proper DPI development. As a consequence of the need to use drug particles below 5 μm in size, these powders are strongly cohesive and show very poor flowability. In order to achieve an accurate metering of the desired small quantities, the drugs are either formulated with additional excipients or processed to soft spheres through spheronization. The goal of both is to improve powder flowability, which is essential to allow volumetric filling of the
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device reservoir or capsules and blister cavities, respectively. The addition of coarse lactose increases the total mass and therefore the metering volume, leading to fewer variations in dosing. Furthermore, the formation of an adhesive mixture improves the flowability significantly, as the used lactose is generally larger in size. A similar effect can be obtained through controlled agglomeration (spheronization) of the micronized drug particles. The increase in agglomerate size accompanies a decrease in interagglomerate forces and, hence, an increase in flowablity. This approach was first patented by Fisons Limited [19] , further developed by Astra Zeneca [20] , and is currently used for products in the Turbuhaler® (Astra Zeneca, UK) and Twisthaler® (Merck & Co., USA) devices. Likewise, Hartmann et al. developed a continuous process to produce comparable spheres through a vibration technique [21] .
into the lungs [27] . Target receptors are not always distributed uniformly throughout the lung. For example, Usmani et al. demonstrated in their study that variations to the particle size of inhaled albuterol lead to a shift in deposition region and patient response. [28] . Furthermore, for the treatment of infectious lungs diseases, or for a possible systemic uptake, it may be necessary to administer API particles uniformly throughout different lung regions [6] . Particles in the respirable size range can be generated via various processing methods. The most commonly used techniques will be highlighted in the following section. Micronization
The administered amount of API in inhalation products needed for an adequate treatment can range from as low as 6 μg of formoterol (e.g., Formatris® Novolizer® [Meda Manufacturing, Germany]) up to 125 mg of colistimethate in Colobreathe Turbospin® (Forest Laboratories, UK) per capsule. However, the majority of DPIs on the market are designated for the treatment of asthma or COPD and, therefore, amounts of 6–500 μg per dose are currently the most common. In recent years, there has been a consistent interest in high-dose deliveries as needed for the administration of antibiotics in tuberculosis, cystic fibrosis or other infectious lung diseases. Novartis managed to expand their treatment of nebulized tobramycin to a high-dose DPI treatment (TOBI® Podhaler®) with a single API dose of 112 mg, which is divided into four 28 mg (approximately 45 mg including excipients) capsules. In addition, another treatment option for patients with cystic fibrosis has recently been introduced to the market with the Colobreathe® Turbospin® Inhaler from Forest Laboratories, UK. The treatment with the encapsulated colistimethate sodium (125 mg twice daily) has been shown to be as effective as the administration of nebulized tobramycin [22] . Improvements in formulation and inhaler design (Twincer® [University of Groningen, The Netherlands] [23,24] , Orbital® [Pharmaxis Ltd., Australia] [25]) may simplify the administration and expand this route of administration to other drugs [26] .
Micronization of the raw API starting material via milling is arguably the most straight forward approach to obtain respirable particles in the lower micrometer size range. Size reduction can efficiently be achieved with different types of mills (e.g., fluid-energy mills, ball mills); however, air jet milling is used on a routine basis. The starting material is injected into the milling chamber, in which air or nitrogen is used to establish a typical grinding pressure of 3–10 bar. The accelerated powder gets comminuted as a result of particle-particle or particle-wall collisions with subsequent breakage until the particles are small enough to discharge to the classification chamber. Due to the rapid expansion of the introduced gas, the thermal stress is low in spite of the high energy input [29] . However, irrespective of the used milling technique, the control of critical parameters is very limited. As a result, important product characteristics, such as size, shape, morphology, surface properties, and electrostatic charge cannot be controlled in a sufficient manner. Also, thermodynamically unstable amorphous parts within the crystalline structure may be generated [30–32] and surface energies and/or attachment energies may vary strongly due to the breakage of the natural crystal surfaces. Further, recrystallization of the amourphous parts may result in particle growth and other uncontrolled changes to the powder during storage. Kubavat et al. [33,34] found that even changes in crystallization procedures for the starting materials lead to different properties of the further micronized particles. All the above mentioned parameters may become critical during the continuous development and production of a dry powder inhalation product as interparticle forces play a key role in the de-agglomeration behavior.
API
Spray drying
Particle size is a driving factor with respect to deposition site in the lungs. The smaller the aerodynamic diameter of the particles, the deeper they may penetrate
The generation of fine particles using spray drying (SD) can help to overcome some of the drawbacks of the micronization process. The technique can be used
Dose
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Review Cordts & Steckel to generate particles with a narrow size distribution out of a solution, suspension or colloidal dispersion. The obtained particles reveal a more homogenous surface energy distribution and generally show a reduced area of contact between particles, thus leading to increased flowability and dispersion efficiencies (i.e., respirable fractions) [35] . Moreover, particle properties, such as density or morphology, can be customized by changing the process and/or formulation parameters [35–37] . However, spray drying often leads to the generation of amorphous particles. Modifications to the process set-up (introduction of a secondary drying step) [38] or change of material (type of carrier) [36] may be suitable approaches to overcome these challenges. In any case, a well investigated SD process and formulation is essential to end up with a desired product and yield. As mentioned before, Novartis managed to introduce the first dry powder treatment with tobramycin to the market by making use of spray drying in order to create their formulation. In this approach, the formulation containing volatile ‘blowing agents’ is spray-dried to produce large porous particles out of an oil-in-water emulsion. This way, powder flowability and deagglomeration is improved and no additional mixing step is needed. Also, a flow rate-independent, high-dose deposition in the lungs has been demonstrated [39] . Furthermore, SD was found to be a suitable technique to produce dry powder protein formulations for inhalation [40] , and additional excipients, such as mannitol or trehalose, may be added to prevent possible degradation during processing and storage [41] . Spray freeze drying
Spray freeze drying (SFD) is a two-step process, during which the feed solution is sprayed into a freezing medium (usually liquid nitrogen). Eventually, the liquid nitrogen evaporates and the frozen particles can be lyophilized in a subsequent freeze drying step. The typical process yield is >95% [42] and the obtained particles are of light and porous nature with improved aerosolization characteristics compared with the micronized material [42,43] . Furthermore, particle properties can again be modified through adjustments of the process (atomization) or formulation set-up [44,45] . Generally, this technique can be used to create powders for various treatments, however, the time consuming and very expensive process narrows the use down to high-cost products, such as protein or antibody formulations. Furthermore, the ability to formulate liposomes [43] or encapsulated particles (e.g., in PLGA) [46] opens up a new set of formulation options for inhaled therapy. Sufficient stability of the material (protein) to shear, freezing and dehydration stresses is essential. Protein
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stability may be increased by adding sugars to the liquid feed, which can then build a glassy matrix and protect the protein during the process [47] . Supercritical fluid technology
A liquid or gaseous phase can be transferred into its supercritical state by rising the pressure and temperature above their specific critical values. The obtained supercritical fluids (SCF) present an intermediate state that combines properties of both, gaseous (low viscosity) and aqueous (density) phase. Consequently, its often advantageous ability to serve as solvent or antisolvent for several drug substances offers a promising alternative approach for the generation of particles in the respirable size range [48,49] . Practically, supercritical carbon dioxide (above 31°C, 74 bar) can be used synonymously for SCF as it is commonly used due to its low cost and non-toxicity. Particles can be generated mainly via three distinctive SCF techniques: rapid expansion of supercritical solutions (RESS), particles from gas-saturated solutions (PGSS) and supercritical anti-solvent (SAS). Since the latter method is independent on the drug substance’s SCF solubility, this process is the most commonly used and different anti-solvent variants (e.g., aerosol solvent extraction system [ASES] and solution enhanced dispersion by supercritical fluids [SEDS]) have successfully been developed [50] . Particles generated with such techniques are usually very uniform in size and exhibit a platelet-like morphology. They show improved flowability and smaller cohesive– adhesive interactions compared with jet-milled powders [51] . Hence, those powders possess enhanced dispersion, that is, increased respirable particle fractions [52] , especially at low airflow rates or low aerodynamic turbulent stresses [49,51] . The bottom-up generation via SCF precipitation has been shown to be a suitable approach for the production of low (water) soluble-drug and peptide particles. It is still subject of intense research and offers a promising alternative to SD and SFD methods. However, its complex and expensive set-up has so far limited its use on a routine basis. Recently, Allergan Inc. have failed to gain US FDA approval for their Levandex® pMDI migraine treatment, which contains dihydroergotamine particles and was meant to be the first commercially available product based on SCF-processed particles [53] . Since the decline was not associated with issues about the processing technology itself, SCF products are still likely to enter the market soon. Carrier
Since micronization of API is still the most commonly used method to obtain particles in the respirable size
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range, further formulation steps to improve flowability are necessary. The aim is always to use the adhesive forces of the micronized powder to build up agglomerates that are larger in size and, therefore, exhibit improved powder flowability. For such a process, the control over agglomerate strength, that is, the magnitude of interparticle forces, is fundamental. On the one hand, agglomerates need to be mechanically stable during storage, filling and dosing, on the other hand, the thoroughly dispersion into single particles upon inhalation is essential to achieve reasonable respirable fractions of the drug. Most of the DPIs on the market contain a powder formulation with additional carrier material. In theory, the micronized drug binds to the host crystal surfaces to form an adhesive mixture [54] and gets separated again as single particles during the inhalation maneuver. As carrier material, different sugars or sugar alcohols such as glucose monohydrate, mannitol, sorbitol, maltitol, xylitol, trehalose [55–58] and others have been tested. Nevertheless, lactose monohydrate is currently the carrier of choice for almost all current DPIs on the market [59] . The reasons for this are manifold; lactose has a well-established safety and stability profile (generally recognized as safe [GRAS] excipient), production is inexpensive, various qualities are easily available on the market, it is less hygroscopic compared with alternative sugars and the knowledge base around its physico-chemical properties is quiet extensive [60] . It presents a suitable carrier for the majority of API molecules, however, lactose may not be applicable for formulations containing substances that are incompatible with its reducing sugar function (especially peptides). A typical carrier size ranges between 50 and 200 μm in diameter and the payload of drug is usually between 0.1 and 4% by weight [61] . However, de-agglomeration efficiency of the mixture is dependent on a large number of factors and their interactions. Extensive research has been performed to understand the influence of particle size, particle morphology, physico-chemical properties (e.g., pseudopolymorphs [62] and amorphous contents), dose [61,63] and environmental conditions on the de-agglomeration behavior, but the very complex relationships are still not fully understood. Among others, one strongly influencing factor is the addition of a third component to the binary API/carrier mixtures. There are numerous publications about the increased respirable API fraction with the addition of a ternary fine (mostly micronized lactose) component. Supposed theories for the enhancement are the saturation of ‘active sites’ on the carrier surface with excipient fines (regions of increased adhesive forces) [64,65] , and the formation of drug/fine excipient agglomerates with beneficial overall dispersion properties [66,67] . It needs
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to be considered that the supplementation of the mixture with ternary fines is accompanied by an increased complexity of variables. Thus, further considerations about the type of ternary component, fines content [68] , size and morphology of the fine particles are necessary [69] . For completeness, additional hypotheses about the mechanistic impact of addition of ternary fines to an adhesive powder mixture should be mentioned, such as the ‘buffer theory’ [70] or ‘fluidization theory’ [71] ; however, they were only applied to a limited number of formulations. For a new DPI development, a literature research should be performed carefully and thoroughly. Much of the data may not be directly applicable to the carrier formulation, as small variations in the formulation and the use of different inhalation devices may change the outcome significantly [68,72] . Blending
In general, the primary goal of a mixing process is the homogeneous distribution of drug particles within the excipient matrix in order to achieve acceptable content uniformity by volumetric dosing or filling. More precisely, for an adhesive inhalation mixture, agglomerates of cohesive drug particles (and excipients fines) need to be broken up and be distributed evenly across the carrier surface. Once adhered to their host crystals, the segregation tendency of an adhesive mixture is comparably low. Primarily, two types of mixers can be utilized to produce such blends: tumbling blenders and high shear mixers. The extent of energy input of the two mixers during blending varies significantly and also so does the magnitude of the driving diffusive, convective and shear mixing effects. Consequently, appropriate mixing speeds and process times for the production of homogeneous blends also differ significantly. However, mixing is a dynamic process. Particles may be redistributed across the carrier surfaces, become part of newly formed agglomerates, or may even act as carriers themselves [73] over the time of the mixing process. Consequently, the mixing efficiency is affected by multiple factors, such as particle size, morphology, density and other physico-chemical properties, and needs to be evaluated carefully for each new drug or component. Perceiving blending as a dynamic process leads to the further understanding that despite a homogenous distribution, redistribution effects of the particles may also be of importance for the correct choice of process parameters. Thus, homogeneity, in other words uniformity, should not be looked at exclusively for an inhalation product. In particular, high shear mixing exerts a high mechanical stress on the larger carrier particles and may
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Review Cordts & Steckel cause breakage or attrition or even changes to the solid state of the drug (amorphization). In addition, prolonged mixing leads to a constant redistribution of particles and may hinder the detachment of drug particles from the host crystal surfaces (‘press-on forces’) [74] . Again, the effects of changes in the mixing protocol may or may not be of overall importance to the DPI as the impact of other formulation variables (especially the inhalation device) may be dominant in the final formulation [75] . In summary, no specific recommendation on the type of mixer can be given at present. Short process times may be advantageous for high shear mixers compared with low shear tumble mixers, however, research about the influence of differences in energy input on inhalation formulations has just moved into focus in recent years [73,75] . Combination products
Both the GINA guidelines for the treatment of patients with asthma and the GOLD guidelines for COPD patients mention the use of long-acting β2-receptor agonists (LABA) in addition to inhaled corticosteroids (ICS) for patients whose treatment with ICS alone is insufficient. This way, two fundamental processes of the diseases can be addressed: bronchoconstriction and inflammation. It was found that the co-administration of the two types of drugs in one formulation resulted not only in increased patient convenience, but also showed improved β2-adrenoceptor and glucocorticoid receptor function, as well as a reduction in inflammation and, potentially, remodeling in the airways and lung due to synergistic effects [76] . For example, the co-administration of fluticasoneproprionate (ICS) and salmeterolxinafoate (LABA) was found to be superior in terms of clinical efficacy to the separate treatment with the two drugs [77] . In a Cochrane meta-analysis of 48 studies including 15,155 patients [78] the authors concluded that the combination therapy is modestly more effective in reducing
the risk of exacerbations requiring oral corticosteroids than a higher dose of inhaled ICS alone. Combination therapy also led to a greater improvement in lung function, symptoms and use of rescue β2-agonists than with a higher dose of inhaled corticosteroids. For children under the age of 12, combination therapy did not contribute to a beneficial treatment, but rather showed a trend towards an increase in adverse effects. Combination products are gaining popularity. An excerpt of marketed products can be found in Table 1. Filling
As most of the DPIs on the market are used for the treatment of bronchial asthma or COPD, the mass of a single dose usually ranges between 5 and 25 mg of a carrier blend. Most commonly, the powder blend is metered into hard gelatin or hydroxypropylmethylcellulose (HPMC) capsules that need to be inserted into the inhalation device and punctured by the patient right before inhalation. Alternatively, powder can be metered into blister strips (single or multi-dose blister inhaler), or larger amounts can be filled into the powder reservoir of a device itself. With the exception of reservoir devices, filling is equivalent to the dosing step. Volumetric dosing is mostly used for inhalation powders, and different approaches can be followed to consistently meter the claimed mass within the specifications (3–5% RSD) [79] . The majority of powder formulations reveal poor to moderate flow properties as a result of their fines content, which results in further challenges for the filling equipment. A well characterized powder (flow properties, poured and tapped densities, compressibility, geometric properties, hygroscopicity, electrostatic charging and adhesive forces) is absolutely essential in order to repeatedly achieve exact dosing by volume. In addition to methods listed in pharmaceutical pharmacopoeias, advanced analysis may be beneficial to understand dynamic and bulk
Table 1. Currently available dry powder combination products.
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Active pharmaceutical ingredient combination
Formulation type
Marketed product
Salmeterolxinafoate + fluticasoneprorionate
Adhesive mixture
Seretide® Diskus® Multi-dose blister AirFluSal® Multi-dose blister Single-dose blister Forspiro® Rolenium® Elpenhaler®
Formoterolfumarate dihydrate + budesonide
Soft spheres
Symbicort® Turbuhaler®
Multi-dose reservoir
Formoterolfumaratede hydrate + beclometasone dipropionate
Adhesive mixture + magnesium stearate
Foster® NEXThaler®
Multi-dose reservoir
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Device type
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properties [80] such as shear cell and powder rheology measurements [71] . Filling of capsules and blisters can generally be approached with two techniques – either a metered part of the bulk powder is transferred into an empty capsule or blister (‘dosing’), or the blister cavity acts as the metering chamber itself and is completely filled (‘filling’). For the latter, the empty cavities are dipped into a well-conditioned powder stock upside down, followed by taking off any excess material and are finally rotated back before sealing. Thus, this process can only be used for powder with appropriate interparticle forces (cohesive powders); otherwise the powder would not remain in the reversed cavities during the filling step. A second technique is the filling of blister cavities with the help of low pressure and membrane filters. The powder is forced through a capillary by the low pressure until all excess air in the cavity has been replaced by the formulation (Figure 1A) . With this approach, sealing faces do not get in contact with the powder formulation. Hence, additional cleaning steps are not required prior to the final blister sealing. More commonly, capsules contain only small quantities of the formulation. Different techniques can be used to transfer these small amounts into the capsule. With a ‘dosator and pin system’ the dosator dips into a loosely packed powder bed and is filled through its open end. A pin inside the dosator compresses the powder, which is subsequently transferred to and ejected into the capsule shell [81] . The machine can be equipped with a vacuum system for powders that fail to remain within the dosator tube during handling (Figure 1C) . Alternatively, a vacuum drum filling system (OMNIDOSE® [Harro Höfliger, Germany]) [82] can be used and is a promising option for the filling of highly cohesive material. The drum with metering cavity sucks in the powder at the 12 o’clock position and rotates to the 6 o’clock position to release the dose with positive pressure into the waiting capsule. During this rotation, excess powder on the drum surface gets removed with a scraper. Finally, the cavity gets cleaned again by another air jet when rotating back to the filling position (Figure 1B) . Equipment based on other filling technologies, such as ‘rotating dosing plate and tamping pin’ [81] , the newly developed ‘vibrating capillary approach’ [83,84] , and others, are also available. Device selection A diversity of DPI devices is currently commercially available [85] . They can be roughly categorized into four types: single-dose devices (capsule or blister based), multi-dose blister, multi-dose reservoir systems, and disposable devices for single use.
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All of them are breath-actuated or passive devices meaning that the energy needed for powder dispersion is generated through the patient’s inspiratory airflow. The first decision for a type of device is usually driven by drug and formulation considerations. However, even if formulation aspects and device design are closely monitored and controlled, the correct handling by the patient is essential for a successful administration of drug. Consequently, an inhalation device should be easy to use and give direct feedback of the successful inhalation maneuver to the patient. Many inhalation devices are under development; some of them are claimed as active devices that include additional features (compressed air, vibrators, electric motors) to generate the aerosol cloud making it independent from the patient’s inspiratory maneuver. To date, only the Exubera® (Pfizer, USA) [86] active inhaler for administration of spray-dried insulin had been introduced to the market in 2006, but it was withdrawn again in 2007 due to a lack of consumer acceptance for the product and low insulin bioavailability [87] . Other active devices are still under development and may be beneficial for systemic treatments due to increased dispersion efficiency of particularly small particles (
