The early medicinal inhalation products deposited to a large extent in the oropharyngeal cavity (80-95% of the inhaled dose), with the majority of the dose swallowed into the gastrointestinal tract [10-18]. This was true even in normal volunteers, despite the fact that nominally the aerodynamic size was often in the range that, in principle, should deliver at least 60% of the dose to the tracheobronchial and pulmonary regions. The specific reasons for this discrepancy are several, and some need to be discussed with reference to a particular means of aerosol delivery system (i.e., the combination of the formulation and the generation device). However, there are two fundamental differences between the aerosols used in the studies described in the previous section and those administered in clinical practice . Therapeutic aerosols are polydisperse (heterodisperse), and their size usually changes after generation, in contrast to the stable monodisperse aerosol studies reviewed earlier.
Because the therapeutic aerosols frequently have a log-normal type of size distribution, polydispersity usually implies that there is a long tail of particles with large aerodynamic diameters. These big particles contain a significant fraction of the therapeutic dose. Several theoretical calculations indicate that there should be very significant differences in the regional deposition of aerosols with the degree of polydispersity found in conventional therapeutic aerosols compared to the deposition of monodisperse aerosols [50,57-61]. A reduction in alveolar deposition from about 60% of the inhaled dose to less than 30% was calculated for a polydisperse aerosol with a mass median aerodynamic diameter of about 3 mm and a degree of polydispersity characterized by the geometric standard deviation of 3.5 compared to a monodisperse aerosol with the same mass median aerodynamic diameter . These more recent findings are in sharp contrast to an earlier, widely quoted occupational hygiene report .
The instability of the particle size of therapeutic aerosols is exhibited in several ways. Perhaps the most thoroughly studied effects are those associated with the generation of drug particles from the propellant-driven metered-dose inhalers. These have usually been formulated as suspensions of drug particles. On release from the metered-dose inhaler, the particles are surrounded by a large volume of propellant and other excipients. The effective aerodynamic size of the drug is, therefore, initially much greater than the intended size (i.e., the size of primary drug particles) . This, together with the high velocity of the aerosol released from the metered-dose inhaler, is predominantly responsible for the high proportion of the dose that deposits at the back of the patient's mouth. If sufficient time is given between generation and entry into the mouth (e.g., by using a holding chamber or a spacer), the propellant may evaporate. It is possible, however, that the resulting particles will be markedly greater than the primary drug particles. This is because they may be associated with involatile excipients and also with other drug particles that were in the same droplet of propellant. Whichever pattern develops with a particular device or formulation, the changing aerodynamic properties of the whole system of the dispersed phase, and not the static properties of the naked drug particles, determine the deposition [64,65]. Recent advances in the design and formulation of metered-dose inhalers using hydrofluoroalkane (HFA) propellants resulted in a product in which a corticosteriod, beclomethasone dipropionate, is dissolved rather than suspended. This metered-dose inhaler (MDI) generates more favorable droplets and evaporation dynamics, leading to significantly smaller particle size and resulting in higher lung deposition and lower loss in the oropharynx than those obtained with the older, chlorofluorocarbon suspensions .
The second reason for the difference between the deposition of stable aerosols and many therapeutic aerosols is hygroscopic growth and evaporation. For drugs to be effective, they must show an appreciable aqueous solubility. Solid drug particles may pick up water vapor and dissolve, and droplets of drug solutions can exchange water with the environment to equalize vapor pressure. The relative humidity profile in the respiratory tract depends on the ambient conditions and the breathing pattern .
The relative humidity eventually reaches almost 100% in alveoli . The drug particles, or droplets of solution, therefore, will tend to exchange water with the surrounding atmosphere, to equilibrate as isotonic solutions at 37°C. Of course, they may deposit, or be exhaled, before the equilibrium is reached if the rate of water transfer is slow. This instability of particle size has been thought for a long time to be a significant factor causing major differences between the deposition patterns of nonhygroscopic and hygroscopic aerosols (note that at very high relative humidity, almost all reasonably soluble substances become hygroscopic). Hygroscopic growth and evaporation were studied in theoretical and in vitro models [68 -72]. The impact of this size instability on the distribution of nonisotonic aqueous aerosols in the human respiratory tract was investigated by gamma scintigraphy [73,74]. A computational model of aerosol deposition that can also deal with aerosol hygroscopicity showed good agreement with the human gamma scintigraphic data .
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