invited review
`
`Recent advances in pulmonary drug delivery
`using large, porous inhaled particles
`
`DAVID A. EDWARDS,1 ABDELAZIZ BEN-JEBRIA,1 AND ROBERT LANGER2
`1Department of Chemical Engineering, The Pennsylvania State University,
`University Park, Pennsylvania 16802; and 2Department of Chemical Engineering,
`Massachusetts Institute of Technology, Cambridge, Massachusetts 02139
`
`Edwards, David A., Abdelaziz Ben-Jebria, and Robert Langer.
`Recent advances in pulmonary drug delivery using large, porous inhaled
`particles. J. Appl. Physiol. 84(2): 379–385, 1998.—The ability to deliver
`proteins and peptides to the systemic circulation by inhalation has
`contributed to a rise in the number of inhalation therapies under
`investigation. For most of these therapies, aerosols are designed to
`comprise small spherical droplets or particles of mass density near 1
`g/cm3 and mean geometric diameter between ⬃1 and 3 µm, suitable for
`particle penetration into the airways or lung periphery. Studies per-
`formed primarily with liquid aerosols have shown that these characteris-
`tics of inhaled aerosols lead to optimal therapeutic effect, both for local
`and systemic therapeutic delivery. Inefficient drug delivery can still arise,
`owing to excessive particle aggregation in an inhaler, deposition in the
`mouth and throat, and overly rapid particle removal from the lungs by
`mucocilliary or phagocytic clearance mechanisms. To address these
`problems, particle surface chemistry and surface roughness are tradition-
`ally manipulated. Recent data indicate that major improvements in
`aerosol particle performance may also be achieved by lowering particle
`mass density and increasing particle size, since large, porous particles
`display less tendency to agglomerate than (conventional) small and
`nonporous particles. Also, large, porous particles inhaled into the lungs
`can potentially release therapeutic substances for long periods of time by
`escaping phagocytic clearance from the lung periphery, thus enabling
`therapeutic action for periods ranging from hours to many days.
`inhalation therapies; respiratory illness; aerosol particles
`
`DRUG DELIVERY TO THE LUNGS by inhalation has attracted
`tremendous scientific and biomedical interest in re-
`cent years. This trend accompanies a rise in respira-
`tory illnesses, dramatized by an increase in asthma
`population in the United States of 46% between 1982
`and 1993 (Centers for Disease Control and Prevention,
`Atlanta, GA). Of at least equal impact has been the
`development of many new biotherapeutics (primarily
`peptide and protein drugs) that in most cases can be
`delivered to humans only by intravenous injection,
`often with low patient compliance. To avoid needles,
`noninvasive delivery strategies have been extensively
`explored. Among these, inhalation delivery has proven
`especially attractive, since the epithelium of the human
`lungs is highly permeable and easily accessed by an
`inhaled dose (7, 15, 23, 24, 26, 30).
`While promising, inhalation therapy is not yet opti-
`mized (22). In many cases, the loss of efficiency or
`reproducibility that this lack of optimization entails
`can preclude inhalation as a practical noninvasive
`
`human therapy. Presently, this is true for many bio-
`therapeutics currently injected intravenously,
`like
`growth hormone, glucagon, or ␣1-antitrypsin, each of
`which could possibly be delivered to humans by inhala-
`tion were the efficiency of inhalation therapy greater.
`Losses of inhaled therapeutic can be attributed to a
`variety of factors; for example, inhaled aerosol particles
`must possess a very narrow range of ‘‘aerodynamic
`diameters’’ (related to a particle’s geometric diameter
`and mass density) to pass through the filter of the
`mouth and throat. Even if properly designed and
`produced, aerosol particles may be propelled with too
`high a velocity and consequently deposited in the mouth
`and throat by inertia. Once in the lungs, particles must
`release the therapeutic substance at a desired rate and,
`in some cases, escape the lungs’ natural clearance
`mechanisms until their therapeutic payload has been
`delivered.
`To meet these challenges, new inhaler devices have
`been, and continue to be, developed. These fall in the
`
`http://www.jap.org
`
`8750-7587/98 $5.00 Copyright r 1998 the American Physiological Society
`
`379
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`categories of metered-dose inhalers, dry-powder inhal-
`ers, and nebulizers (14, 21, 32). When combined with
`optimized aerosol formulations, these new inhalers
`promise to significantly expand the use of inhalation
`therapy in humans (20).
`Among the factors that can be adjusted to optimize
`the efficiency of aerosol formulations, particle chemis-
`try and surface morphology (manipulated to reduce
`particle-particle aggregation or hygroscopicity) and par-
`ticle solubility (altered to influence the rate of therapeu-
`tic release) are rather well documented (10, 11). Less
`well documented is the potential for major improve-
`ments in aerosolization efficiency by diminishing aero-
`sol particle mass density and increasing particle size,
`as done in recent studies for select formulations (6, 35).
`The delivery of large, porous particles to the lungs may
`also permit exceptionally long-acting therapeutic deliv-
`ery following inhalation, as further described in this
`review.
`
`THE AERODYNAMIC DIAMETER WINDOW
`
`To understand the rationale behind therapeutic aero-
`sol particle design, it is helpful to briefly review the
`concept of aerodynamic diameter and its relation to the
`pattern of particle deposition in the lungs. Aerody-
`namic diameter is the geometric diameter a particle
`appears to possess on the basis of its in-flight speed,
`were it assumed to be spherical and to possess a mass
`density of 1 g/cm3; stated differently, the geometric
`diameter of a spherical particle possessing unit mass
`density (1 g/cm3) is equivalent to its aerodynamic
`diameter. Because many naturally occurring particles
`possess a mass density near this value and because
`sphericity is a tendency of nature based on surface
`energetic considerations, such a ‘‘base-case’’ particle
`has proven useful for discussing the sites and extent of
`aerosol particle deposition in the lungs as a function of
`particle size.
`A more quantitative idea of aerodynamic diameter
`can be gathered by imagining a spherical particle
`falling under gravity through air; so long as the charac-
`teristic particle size is substantially larger than the
`mean free path of the surrounding air molecules, it can
`be shown that the particle will settle with a velocity (v)
`
`v ⫽
`
`mg
`3␲µd
`
`where m is the particle mass, g is the gravitational
`constant, µ is the viscosity of air, and d is the particle
`diameter. Expressed in terms of particle mass density
`(␳), this gives
`
`
`
`v ⫽ 1 g18µ2 ␳d2
`
`(1)
`
`Equation 1 shows that spherical particles will fall
`under gravity with a velocity that is proportional to
`their mass density ␳ and the square of their geometric
`diameter (d2). If we consider the heuristic definition of
`aerodynamic diameter provided above, it is possible to
`
`rewrite Eq. 1 in terms of the particle’s aerodynamic
`diameter, as
`
`v ⫽ 1␳a g18µ2 da
`where ␳a ⫽ 1 g/cm3, and where we have defined the
`aerodynamic diameter (da) by the relationship
`
`
`
`2
`
`(2)
`
`2 ⫽ ␳d2
`␳ada
`
`(3a)
`
`or
`
`da ⫽冑 ␳
`
`d
`
`(3b)
`
`␳a
`Equation 2 shows that a spherical particle of any mass
`density ␳ will settle with a velocity that depends only on
`its aerodynamic diameter da, i.e., is dependent on size
`and mass density through the specific relation of Eq. 3.
`Modifications to Eq. 3 arise for nonspherical particles
`(10); these modifications can include more than a single
`aerodynamic coefficient, particularly in the case of
`nonisotropic particles (e.g., cylinders), wherein the
`particles translate with a preferred orientation.
`Because gravitational settling constitutes one of the
`principal mechanisms of aerosol particle deposition in
`the lungs, the concept of aerodynamic diameter be-
`comes a useful intrinsic particle property with which to
`discuss a particle’s expected lung deposition perfor-
`mance following inhalation. Moreover, as shown by
`Landahl (17) in his landmark study of aerosol particle
`deposition in the lungs, the second principal mecha-
`nism of particle deposition in the lungs, inertial impac-
`tion, also depends uniquely on da (see Fig. 1, repro-
`duced from Ref. 9). Aerodynamic diameter has thus
`been used for several decades to quantify an aerosol
`particle’s inherent propensity to deposit in the lungs,
`essentially independently of its shape, mass density,
`and (in principle) geometric size.
`Numerous experimental (2, 5, 12, 19) and theoretical
`(3, 9, 13) studies have demonstrated that particles of
`mean aerodynamic diameter of 1–3 µm deposit mini-
`mally in the mouth and throat and maximally in the
`lung’s parenchymal (i.e., alveolar or ‘‘deep-lung’’) re-
`gion. Tracheobronchial deposition, generally not de-
`sired for an inhalation therapy, is maximized for aero-
`dynamic diameter between ⬃8 and 10 µm. Particles
`possessing an aerodynamic diameter smaller than ⬃1 µm
`(although greater than several hundred nanometers)
`are mostly exhaled, and particles larger than ⬃10 µm
`have little chance of making it beyond the mouth.
`The aerodynamic diameter window of 1–3 µm has
`also proven optimal for inhalation drug delivery. By
`using liquid aerosols (whence aerodynamic and geomet-
`ric diameters coincide) Clay et al. (4) found that optimal
`bronchodilation in human asthmatics occurred after
`inhalation of 1.8-µm-diameter (terbutaline) droplets,
`relative to droplets of 4.6 or 10.3 µm in diameter (all
`other factors held constant). Johnson et al. (16) and
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`381
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`µm), natural fog (2–80 µm), coal dust (3–30 µm), and
`metal fumes (0.01–100 µm).
`
`THE BIOAVAILABILITY QUESTION
`
`To further clarify the optimal characteristics sought in
`an ‘‘ideal’’therapeutic aerosol formulation, we review one of
`the major driving forces for the expanding use of inhaled
`therapies; namely, to replace injectables with an easy-to-
`use, affordable, noninvasive delivery technology.
`Intravenous (iv) injection, the primary route of pro-
`tein therapeutic administration, is painful, and patient
`compliance is often low as a consequence; several
`excellent reviews have documented the case for pulmo-
`nary administration as a practically viable alternative
`(22, 30, 34). Relative to oral administration, pulmonary
`delivery exposes the administered drug to a less harsh
`environment, and the permeability of the alveolar
`epithelium is high. Compared with the sites of transder-
`mal, nasal, vaginal, buccal, or ocular delivery, the exposed
`surface area of the lungs is extremely large, estimated to
`approximate the area of a tennis court (22).
`The efficiency of drug delivery by a noninvasive route
`can be expressed in terms of bioavailability (B) defined as
`
`B ⫽1 AUCAUCiv2 ⫻1doseiv
`
`dose2 ⫻ 100
`
`(4)
`
`whereAUC represents the integrated area under the curve
`of systemic plasma concentrations of administered drug vs.
`time, given a particular dose (dose); and AUCiv is the
`comparable area obtained after iv injection of the same
`drug with dose (doseiv). Because inhalation of drugs
`involves losses of drug in the inhaler as well as in the
`mouth and throat, researchers define bioavailability
`either in terms of the actual dose administered or the
`dose delivered to the pulmonary system, the latter
`being a kind of optimal intrinsic value and the former
`having more direct relevance to an actual therapy.
`Depending on the dosage form, a relative bioavailabil-
`ity may also be defined in terms of subcutaneous
`(rather than iv) injection, i.e.
`
`Bsc ⫽ 1 AUCAUCsc2 ⫻ 1dosesc
`
`dose 2 ⫻ 100
`
`(5)
`
`where sc denotes subcutaneous administration. As
`measures of bioavailability, values of Bsc overestimate
`the ‘‘absolute’’ value of B. Table 1 lists the bioavailabili-
`ties of several different proteins following inhalation,
`with absolute definitions of B.
`Together with the question of appropriate pharmaco-
`kinetics, the question of bioavailability is central to the
`viability of a systemic inhalation therapy. Given a
`required dose to be administered to the systemic circu-
`lation, as well as the acceptable dose variability and
`upper limit on the amount of mass that can be inhaled
`into the lungs at any one time without unwanted
`side-effects such as coughing (e.g., ⬃20 mg), a practical
`lower limit is placed on B, below which inhalation is an
`unacceptable alternative to iv injection. This value may
`be, for example, 1, 10, or 30%, depending on the drug
`
`Fig. 1. Relative deposition in a human lung model after inhalation of
`a monodisperse aerosol as a function of aerosol particle aerodynamic
`diameter, for particles with a mean aerodynamic diameter of 1–10
`µm. Contributions from the 3 predominant mechanisms of deposition
`are listed; these include gravitational sedimentation, inertial impac-
`tion, and Brownian diffusion. The latter mechanism of deposition
`becomes predominant for particles of mean aerodynamic diameter
`less than ⬃0.5 mm (9), hence it does not substantially contribute to
`therapeutic particle-deposition patterns in lungs. Other deposition
`mechanisms, including electrostatic precipitation (e.g., in an inhaler)
`or interception, are not considered and may be assumed either to be
`minor contributors to overall particle deposition or to play an
`important role in limited circumstances, such as those involving
`particles with significant surface charge. [From Gerrity et al. (9).]
`
`Ruffin et al. (28), respectively, showed that liquid
`aerosols comprising 3.3-µm (salbutamol) and 1.5-µm
`(isoproterenol) droplets also produced greater broncho-
`dilation in human asthma patients than did droplets
`substantially larger than 3 µm. Also, Zanen et al. (37,
`38) in two separate studies found that 2.8-µm aerosols
`of salbutamol and ipratropium bromide produce greater
`bronchodilation than droplets of larger or smaller size
`(although little difference was seen between the 1.5-
`and 2.8-µm ipratropium bromide aerosols). The aerody-
`namic diameter window of 1–3 µm has also been shown
`to be optimal for systemic delivery in dogs after inhala-
`tion of leuprolide acetate (liquid) aerosols (1).
`Given that most published studies of particle size
`effect on therapeutic efficacy have used approximately
`spherical aerosol particles of mass density near unity (1
`g/cm3) and considering that manipulation of aerody-
`namic diameter by particle mass density has tended to
`be viewed as hard to achieve (10, 11), the aerodynamic
`diameter window of 1–3 µm has become associated in
`many published research articles with an intrinsic
`geometric diameter window in the same range (1–3
`µm) (1, 30). This partly explains why therapeutic
`aerosol particles are primarily designed today with
`geometric diameters in the range of 1–3 µm, even
`though aerosol particles of much larger size (and lower
`mass density) can be encountered in the environment
`and often enter into the lungs. This latter point is
`reflected in the following list of airborne particles (note
`that increasing particle size generally coincides with
`diminishing mass density, in accordance with the rela-
`tionship indicated by Eq. 3) (25, 33): pollen (10–100
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`Table 1. Bioavailabilities and absorption times
`of peptides and proteins after administration
`to the lung
`
`its purpose is to control the pharmacokinetic profile or
`to protect a protein over the time period of its absorp-
`tion, controlled release from an aerosolized particle into
`the lungs over periods ranging from minutes to, per-
`haps, many days may be desirable.
`Barriers that must be surmounted to achieve high
`bioavailability and proper pharmacokinetic profile natu-
`rally differ with drug but generally include 1) loss of
`drug owing to filtration of aerosolized dose in the
`inhaler, mouth, or throat; and 2) clearance of drug from
`the lungs before drug action. The fact that particles of
`mean diameter of 1–3 µm are both prone to aggregation
`and can be rapidly phagocytosed in the deep lungs by
`alveolar macrophages makes these barriers intrinsi-
`cally problematic for conventional aerosol particles. For
`example, by using conventional aerosol particles of 1- to
`3-µm size, current inhalers deliver ⬃10–20% of inhaled
`drug to the lungs (32). Clearance of drug from the lungs
`after deposition can include enzymatic degradation,
`mucocilliary clearance, and phagocytosis. The first of
`these can be important for molecules like vasoactive
`intestinal polypeptide (34). Mucocilliary clearance is
`especially important in the upper and central airways
`(22), but phagocytosis of slowly soluble particles may be
`the most significant mechanism of clearance from the
`deep lungs (6); Gehr et al. (8), for example, report that
`approximately one-third of polystyrene particles are
`immediately phagocytosed after inhalation into the
`lungs of hamsters. This rapid clearance acts against
`the potential benefits of an inhaled particle with con-
`trolled-release features, rendering sustained release of
`drugs into or from the lungs hard to achieve with
`current inhalation aerosol particles.
`
`ADVANTAGE OF LARGE AND POROUS
`AEROSOL PARTICLES
`
`A new type of inhalation aerosol has recently been
`identified that may help to address limitations of the
`current inhalation therapies (6). This aerosol is formed
`by particles possessing low mass density and, conse-
`quently, large size, such that the particles’ mean aerody-
`namic diameter fits into the window of 1–3 µm. The
`advantage of large size and low mass density is twofold:
`first, increased particle size results in decreased ten-
`dency to aggregate; hence, in combination with low
`mass density, this leads to more efficient aerosolization
`in a given air field; second, since phagocytosis of particles
`by macrophages diminishes with increasing particle size
`beyond ⬃2–3 µm (27, 31), very large particles deposited in
`the pulmonary region may escape clearance by alveolar
`macrophages and, therefore, permit drug release for longer
`periods of time and more efficiently.
`To clarify the potential for large and porous aerosols
`to increase systemic bioavailability as well as to pro-
`vide sustained-release capability in the lungs, Edwards
`et al. (6) encapsulated insulin into a biodegradable
`copolymer commonly used in biodegradable sutures
`and in controlled-release, implantable or injectable
`depot systems (18). Poly(lactic-co-glycolic) acid polymer
`particles were prepared with encapsulated insulin in
`two forms: a small, nonporous (conventional) aerosol
`particle and a large, porous (novel) aerosol particle of
`
`Compound
`(mol wt)
`
`Leuprolide (1,209) Human
`Calcitonin (3,416) Rat
`Glucagon (3,481) Rat
`PTH-34 (4,278)
`Rat
`Insulin (5,786)
`Human
`Growth hormone
`(22,000)
`AIA (52,000)
`IgG (150,000)
`
`Rat
`Sheep
`Rat
`
`Absorption
`Time
`tmax ⬃1–2 h
`17
`tmax ⬃15 min
`32
`Undetermined 33
`tmax ⬃15 min
`33
`tmax ⬃15 min
`34
`tmax ⬃1–4 h
`9–10%
`48-h study
`⬍5%
`1.5–1.8% 192-h study
`
`Species Bioavailability
`
`Reference
`
`17%
`17%
`⬍1%
`40%
`7–16%
`
`35
`36
`37
`
`PTH, phenylthiohydantoin; AIA, ␣1-antitrypsin; tmax, maximal
`time of absorption.
`
`and therapy, with the understanding that an increase
`in B value above this theoretical minimum will always
`tend to benefit the therapy by decreasing cost and likely
`increasing reproducibility. In the case of insulin, widely
`studied for its role in the chronic treatment of diabetes,
`reproducibility must be especially high to avoid hypogly-
`cemic shock. Sufficient bioavailability and reproducibil-
`ity have, however, been observed in human insulin
`trials (29) to justify hopes that an inhalation insulin
`therapy may soon reach the commercial US market.
`Calcitonin, interferons, parathyroid hormone, and leu-
`prolide are also in human clinical trials for systemic
`action after inhalation. On the other hand, proteins
`such as glucagon (used to treat hypoglycemic coma),
`␣1-antitrypsin (used for the treatment of emphysema),
`or growth hormone (used for the treatment of growth
`deficiency), while attractive molecules for inhalation,
`have not yet reached the human trial stage.
`A second key element to inhalation therapy concerns
`pharmacokinetic profile following administration. Lack-
`ing control over delivery by a long-lived, controlled-
`release particle, the systemic delivery rate of drugs
`from the lungs depends strongly on molecular size and
`structure, as indicated by the varied plasma absorption
`times listed in Table 1. A drug’s intrinsic absorption
`profile may, however, not perfectly suit the required
`therapy, as is the case for sustained release of insulin
`for diabetic patients (29). A sustained-release inhaled
`therapeutic particle might serve to change the pharma-
`cokinetic plasma profile, making it more suitable to the
`particular therapy. Sustained release from an inhaled
`therapeutic particle may also help to improve the
`bioavailability of very large proteins, such as IgG
`(Table 1), that absorb extremely slowly into the sys-
`temic circulation from the lungs. To the extent that the
`low bioavailability of slowly absorbing macromolecules
`can be attributed to long residence time in the alveolar
`fluid before absorption, during which time activity is
`lost by enzymatic degradation or aggregation, a con-
`trolled-release form of the drugs might again be benefi-
`cial. By controlling the release of a drug from an
`inhaled particle such that it approximates the rate of
`absorption into the bloodstream, the overall bioavailabil-
`ity of the drug might be increased without necessarily
`altering the pharmacokinetic profile. Hence, whether
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`383
`
`similar da (2.15 µm). Relative bioavailability of the
`conventional aerosol particle after aerosolization as a
`dry powder into the lungs of rats was 12%, and the
`release time was ⬃4 h. Figure 2 shows that insulin
`release into the systemic circulation from the large and
`porous aerosol particle lasted ⬃96 h, with similar
`inhalation and subcutaneous release profiles. The rela-
`tive bioavailability (cf. Eq. 5) of the large and porous
`insulin particle was 87.5%, about seven times greater
`than that of the small and nonporous insulin particle.
`One reason for the relatively high bioavailability of
`the inhaled large-particle insulin is more efficient
`delivery of drug to the lung as a consequence of less
`powder aggregation (6). An increase in the size of
`aerosol particles results in a reduced fractional surface
`
`Fig. 2. Systemic concentrations after administration of porous thera-
`peutic particles. A: serum insulin concentration after inhalation and
`subcutaneous injection of 9 mg of large, porous insulin particles. No
`insulin particles were administered to nontreated controls. Porous
`particles contained insulin (20.0 wt%) and 50:50 poly(lactic-co-
`glycolic) acid (80.0 wt%). Their mean geometric and mean aerody-
`namic diameters were, respectively, 6.8 and 2.15 µm. Mean and SE
`values are based on n ⫽ 3 subjects. B: serum glucose concentrations
`after inhalation of 9-mg large, porous insulin particles. No insulin
`particles were administered to nontreated controls. Mean and SE
`values are based on n ⫽ 3 subjects. [Reprinted with permission from
`D. A. Edwards, J. Hanes, G. Caponnetti, J. Hrkach, A. Ben-Jebria,
`M.-L. Eskew, J. Mintzes, D. Deaver, N. Lotan, and R. Langer. Large
`porous particles for pulmonary drug delivery. Science 276: 1868–
`1871, 1997, Copyright 1997 American Association for the Advance-
`ment of Science.]
`
`area (or likelihood) of particle-particle contact in a dry
`powder (or liquid suspension) and thus in less tendency
`to aggregate. This diminished aggregation means that
`less energy is required to aerosolize particles or that
`particles are more efficiently aerosolized with a given
`energy of aerosolization.
`However, perhaps the predominant cause of the
`sustained insulin delivery is the role of large particle
`size in discouraging phagocytosis. In a parallel study
`(6), the lungs of rats were lavaged both immediately
`after inhalation and 48 h after inhalation, with an aim
`to determine the location of the two types of insulin
`particles in the alveolar fluid. Similar numbers of
`porous and nonporous particles were inspired into the
`rat lungs by administering identical masses of porous
`and nonporous particles into the trachea.1 Immediately
`after inhalation, 30 ⫾ 3% of macrophages contained
`insulin particles in the case of rat lungs into which
`small and nonporous particles had been inspired, com-
`pared with 8 ⫾ 2% of macrophages containing large and
`porous particles. After 48 h, 17.5 ⫾ 1.5% of macro-
`phages contained three or more small and nonporous
`particles compared with 4 ⫾ 1% of macrophages contain-
`ing three or more large and porous particles. Figure 3B
`shows two alveolar macrophages 48 h after inhalation
`of the small and nonporous insulin particles. Each
`macrophage contains numerous particles. Figure 3C
`shows several alveolar macrophages 48 h after inhala-
`tion of the large and porous insulin particles. The
`macrophages surround a large, porous particle with-
`out, however, engulfing it; other particles can be seen
`along the periphery of the macrophages. The lower
`extent of particle phagocytosis in the case of the large
`and porous therapeutic particles accompanies a rela-
`tively low inflammatory response as well. These results
`support the findings of in vitro experimental studies, cited
`above (27, 31), that show a trend of diminished particle
`phagocytosis with increasing particle size beyond ⬃3 mm.
`The potential of large, porous particles for inhalation
`of a variety of drugs is presently being explored. In
`addition to insulin and testosterone (6), promising
`results in animals have recently been obtained with
`long-lasting formulations of estradiol (35), for hormone
`therapy, and albuterol (35), for asthma. In the estradiol
`study, particles were produced by spray drying, compris-
`ing estradiol and combinations of lung surfactant (dipal-
`mitoylphosphatidylcholine), human albumin, and lac-
`tose. Large, porous particles with a mean geometric
`diameter of 10 µm and bulk (tapped) density 0.09 g/cm3
`
`1 Due to the relatively inefficient aerosolization properties of the
`nonporous particles, based on in vitro and in vivo measurements
`using particles of similar size and porosity (6), approximately twice
`the porous powder mass is estimated to have entered the lungs
`relative to nonporous powder mass. At the same time, the larger size
`of the porous particles resulted in fewer particles per powder mass;
`this can be seen as follows: if 1 and 2 denote porous and nonporous
`particles, respectively; N is the number of particles; d is the particle
`diameter; and ␳ is the particle mass density, the porous powder mass
`can be represented as N1(1/6)␲␳1d1
`3 and the nonporous powder mass
`as N2(1/6)␲␳2d2
`3. Accounting for the lesser amount of aerosolized
`nonporous mass entering the lungs, and noting Eq. 3a, gives the
`relation N1/N2 ⬃ 2d2/d1. Using d1 ⫽ 6.8 µm, and d2 ⫽ 4.4 µm (6), the
`ratio of porous to nonporous particle numbers entering the lungs is
`near unity (N1/N2 ⬃ 1.3).
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`Fig. 3. Pulmonary macrophages (large circular
`objects with dark nuclei) removed from lungs. A:
`immediately after inhalation of (conventional)
`small and nonporous particles; B: 48 h after
`inhalation of small and nonporous particles; and
`C: 48 h after inhalation of large, porous particles.
`Inhaled particles (appearing like white crystals
`on photographs) largely surround macrophages
`in A, are primarily engulfed inside two macro-
`phages in B, and remain (notably, the largest
`particle near center of photograph) outside sev-
`eral macrophages in C.
`
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`385
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`were prepared, as were nonporous particles possessing
`a mean geometric diameter of 3 µm and bulk (tapped)
`density 0.48 g/cm3. After being inspired as an aerosol
`into the lungs of rats, the large, porous particles
`produced elevated systemic levels of estradiol for 5
`days, whereas the small nonporous particles produced
`elevated systemic estradiol levels for only 1 day (35).
`Relative bioavailability in the case of the large, porous
`estradiol particles was ⬃87%. In the albuterol study,
`large, porous particles prepared, again, by spray drying
`with albuterol and a combination of dipalmitoylphos-
`phatidylcholine, albumin, and lactose resulted in sus-
`tained bronchodilation in guinea pigs for ⬃1 day, at
`relatively low inhaled albuterol doses, compared with
`several hours in the case of small, non-sustained-
`release aerosol particles.
`Whereas considerable work remains to clarify the poten-
`tial bioavailability and efficiency gains that can be achieved
`after inhalation of large, porous aerosol formulations in
`humans, the results to date suggest that such particles
`may play a role in the development and optimization of
`new inhalation therapies in the future.
`Address for reprint requests: D. A. Edwards, Advanced Inhalation
`Research, Inc., 840 Memorial Drive, Cambridge, MA 02139.
`
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