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JOURNAL OF AEROSOL MEDICINEVolume 20, Number 3, 2007© Mary Ann Liebert, Inc.Pp. 294–309DOI: 10.1089/jam.2006.0617

Computational Analyses of a Pressurized Metered Dose Inhaler and a New Drug–Aerosol

Targeting Methodology

CLEMENT KLEINSTREUER,1 HUAWEI SHI,2 and ZHE ZHANG3

ABSTRACT

The popular pressurized metered dose inhaler (pMDI), especially for asthma treatment, hasundergone various changes in terms of propellant use and valve design. Most significant arethe choice of hydrofluoroalkane-134a (HFA-134a) as a new propellant (rather than chloroflu-orocarbon, CFC), a smaller exit nozzle diameter and attachment of a spacer in order to reduceultimately droplet size and spray inhalation speed, both contributing to higher deposition ef-ficiencies and hence better asthma therapy. Although asthma medicine is rather inexpensive,the specter of systemic side effects triggered by inefficient pMDI performance and the in-creasing use of such devices as well as new targeted drug–aerosol delivery for various lungand other diseases make detailed performance analyses imperative. For the first time, exper-imentally validated computational fluid-particle dynamics technique has been applied to sim-ulate airflow, droplet spray transport and aerosol deposition in a pMDI attached to a humanupper airway model, considering different device propellants, nozzle diameters, and spaceruse. The results indicate that the use of HFA (replacing CFC), smaller valve orifices (0.25 mminstead of 0.5 mm) and spacers (ID � 4.2 cm) leads to best performance mainly because ofsmaller droplets generated, which penetrate more readily into the bronchial airways. Exper-imentally validated computer simulations predict that 46.6% of the inhaled droplets may reachthe lung for an HFA–pMDI and 23.2% for a CFC–pMDI, both with a nozzle-exit diameter of0.25 mm. Commonly used inhalers are nondirectional, and at best only regional drug–aerosoldeposition can be achieved. However, when inhaling expensive and aggressive medicine, orcritical lung areas have to be reached, locally targeted drug–aerosol delivery is imperative.For that reason the underlying principle of a future line of “smart inhalers” is introduced.Specifically, by generating a controlled air–particle stream, most of the inhaled drug aerosolsreach predetermined lung sites, which are associated with specific diseases and/or treatments.Using the same human upper airway model, experimentally confirmed computer predictionsof controlled particle transport from mouth to generation 3 are provided.

Key words: computer simulation, droplet spray formation, pMDI performance, droplet trans-port/deposition, human upper airways, controlled air-particle stream, targeted drug–aerosoltransport/deposition

1Department of Mechanical and Aerospace Engineering and Department of Biomedical Engineering, North Car-olina State University, Raleigh, North Carolina.

2Turbine Technology International, Inc., Rochester, New York.3Department of Mechanical and Aerospace Engineering, North Carolina State University, Raleigh, North Carolina.

INTRODUCTION

PRESSURIZED METERED DOSE INHALERS (pMDIs),dry powder inhalers and nebulizers are the

most common oral drug delivery devices. ThepMDI is a compact, easy-to-use droplet spray dis-penser for multiple-dose treatment of pulmonarydiseases, the most common being asthma. Typi-cally 200 high-pressure (Pcanister � 10 atm) actua-tions, containing 20 to 5000 �g of drug in meteredvolumes of 25 to 100 �L, can be generated andinhaled.1–3 A key component of the device is theactuator spray nozzle, which controls the atom-ization process and resultant spray formation interms of spray angle and droplet size distribu-tion. The formulation in pMDIs consists of a ther-apeutic material and a propellant as well as sur-factants and other excipients which dissuadedrug–aerosol coagulation and lubricate movingparts of the metering valve. While pressurizedchlorofluorocarbon (CFC) substances are beingphased out worldwide because released chlorineatoms deplete the earth’s ozone layer, hydroflu-oroalkane (e.g., HFA-134a and HFA-227) have be-come replacement propellants.4,5 However, dueto the different physical and chemical propertiesof CFC and HFA, both the formulation in pMDIsand their structural elements had to be modified.6

In any case, 70 to 80% of the discharged formu-lation of the discharged formulation is liquid withnozzle exit velocities in the range of 150 to 225m/sec.

The overall goal for pMDIs is to deliver mostof the medicine (e.g., BDP for asthma7) to a pa-tient’s bronchopulmonary tree and lower air-ways, that is, avoiding depositions inside the ac-tuator and in the oral airways. Numerous in vitroand in vivo tests with pMDIs have been conductedto approach that goal; however, no computa-tional fluid-particle dynamics (CFPD) analysishas been found in the open literature. Usuallycompanies employ a geometric model of the up-per airways, that is, the United State Pharma-copoeia (USP) throat and the Andersen cascadeimpactor, to estimate particle size from pMDIsand to estimate deposition pattern using gammascintigraphy (see http://www.westechinstru-ments.com). For example, Leach et al.8 comparedCFC- and HFA-driven pMDI performances. Whilemost of the drug aerosols from CFC–pMDIs re-mained in the oropharyngeal region (i.e., 80–90%), HFA–pMDIs delivered up to 60% of the in-haled medicine into the central and intermediate

lung zones. Such an increase in drug deliverywith newer HFA–pMDIs can be contributed tothe improved formulation characteristics and ac-tuator design, resulting in smaller drug aerosols.In vitro tests were carried out by a number of re-searchers.9–12 They argued that the use of a real-istic human upper airway replica is preferable tothe USP throat and compared CFC- versusHFA–pMDI performances, considering the hold-ing chamber and consistent drug dosing, the ef-fect of temperature, the influence of spacers, andthe impact of orifice diameter on spray force.

In this paper, the results for a conventional in-haler as well as the methodology for a new smartinhaler system are discussed. Specially, the airflowand microdroplet dynamics inside a pMDI modelattached to a representative human upper airwayreplica are simulated, considering different propel-lants, exit-nozzle diameters, and spacer impact. Thework relies on the contributions by Shi and Klein-streuer13 describing the dynamics of high-pressuresprays and by Kleinstreuer and Zhang14 simulatingair-particle flow in a human airway model adaptedfrom Cheng et al.15 The key question addressed hereis how much of the inhaled medicine, that is, dropletmass fraction, travels past the trachea, and hence,reaches the bronchial tree and beyond.

While pMDIs may deliver some of the inhaledmedicine to the central lung region, most treat-ments require high-percentage drug aerosol de-position at specific lung sites, such as tumors, orlung regions, for example, alveolated ducts. Forexample, Kleinstreuer and Zhang16 analyzed acritical tumor size for maximum drug–aerosol de-position under various inhalation conditions.Some drugs are so aggressive and/or expensivethat targeted aerosol delivery is imperative.Clearly, the development and testing of such anew generation of inhalers is a very active re-search area (refer to www.uspto.gov/patft/in-dex.html). Some preliminary results, based on anew “controlled air-particle stream” methodol-ogy,17 are presented (see later section). A smartinhaler system which implements the idea is dis-cussed by Kleinstreuer and Seelecke.18

THEORY

Modeling equations and system geometries

Assuming a flow rate of Q � 30 L/min withN � 3000 droplets tracked, three-dimensional air

ANALYSES OF pMDI AND DRUG–AEROSOL TARGETING 295

FIG. 1. pMDI structure and air-droplet modeling aspects.

ANALYSES OF pMDI AND DRUG–AEROSOL TARGETING 297

drag forces on the droplets. The drag force willresult in droplet deformation, as determined bythe Weber number, which is generally defined asthe ratio of the inertia force to surface tension:

We � (2)

where �g is the gas density, Vslip is the relative ve-locity (i.e., slip velocity), r is the droplet radius,and � is the coefficient of liquid surface tension.

�gV2slipr

��

inhalation was described by a low Reynolds num-ber k–� turbulence model.14,19 The spray evolu-tion was simulated with gas–liquid interactiveEnhanced Taylor Analogy Breakup (ETAB)model,13,20 while the resulting droplet trajectorieswere determined with Newton’s second law gov-erned by drag and gravity as the major forces fordispersed droplets in low density gas flow, thatis,

mp � ��gd2pCD(U

→

g � u→p)�U→

g � u→p� � G→

(1)

where x→p is the particle position, u→p is the particlevelocity, mp is the particle mass, �g is the gas den-sity, G

→is the gravitational force, CD is the drag co-

efficient, and dp is the particle diameter. A justi-fication for the use of Equation (1) is provided inKleinstreuer,17 among other sources. Consideringspherical microdroplets at low formulation massflow rates1 (�0.5 g/sec), the two phases are de-coupled in the present Eulerian (gas)-Lagrangian(liquid) modeling approach.21 Furthermore, tur-bulent airflow effects on droplet trajectories wereneglected because the focus was on aerosol trans-port and not deposition (see Matida et al.22), aswell as uniformity in operational device compar-isons.

The ETAB model, a basic theory for liquiddroplet breakup, was selected because of its ac-curacy and stability.13,20 Specifically, if a dropletis exposed to gas flow and there is a velocity dif-ference (i.e., slip velocity) between droplets andsurrounding gas, the gas phase will exert large

1�8

d2x→p��dt2

TABLE 1. ESTIMATED PMDI AND PROPELLANT PROPERTIES

CFC pMDI HFA pMDI

pMDI properties Nozzle diameter (mm) 0.5 (Cheng et al.10) 0.25 (Cheng et al.10)Spray angle (degree) 35 35Droplet exit velocity 150 (Dunbar et al.2) 150 (Dunbar et al.2)

(m/sec)Propellant used Mixture of CFC 11 and CFC 12 HFA-134a

(Dunbar et al.1) (Cheng et al.10)Propellant Formula CCI3F, CCl2F2 CH2FCF3

properties* Density (kg/m3) 1480 1311Viscosity kg/msec 4.0E-4 3.8E-4Diffusivity (m2/sec) 9.0E-6 8.0E-6Surface tension (N/m) 0.017 0.015Specific heat (kJ/kgK) 0.59 (Dunbar et al.1) 0.79Boiling point (°C) �24.0 (Dunbar et al.1) �26.0Latent heat (kJ/kg) 153.2 (Dunbar et al.1) 214.8

Other sources: http://webbook.nist.gov/chemistry/fluid/http://www.epa.gov/athens/learn2model/part-two/onsite/estdiffusion-ext.htm

FIG. 2. Comparison between the simulated and mea-sured microparticle deposition efficiency in the humanupper airway.

FIG. 3. Spray model validation. (a) Comparison between the simulated and measured spray droplet velocity dis-tribution. (b) Comparison between the simulated and measured spray penetration depth.

Thus, the Weber number connects the gas-in-duced drag force, which leads to droplet defor-mation against liquid surface tension which tendsto maintain a spherical droplet shape. Significantdeformation starts at a Weber number of unity.Above a certain value of the Weber number,droplet deformation leads to breakup. The ETABmodel is a Weber number-based, spray dropletbreakup model to predict droplet lifetime, chil-dren droplet size, and momentum. Detailedmathematical descriptions can be found else-where.13,20

Droplet evaporation was simulated with alumped parameter model23 where the evapora-tion rate depends on the droplet diameter,gas/vapor mixture density, diffusivity, heattransfer number, and presence of surfactants.Specifically, the droplet evaporation model usedin this study is based on theory reviewed byFaeth:23

m� d � �2� � dp � �� � DAB � ln(1 � BT) (3)

where m� d is the mass evaporation rate (kg/s), dpis the droplet diameter, �� is gas/vapor mixturedensity, DAB is the formulation vapor phase dif-fusivity in air, and BT is the heat transfer numberdefined as:

BT � (4)

Here, cp is the formulation specific heat and Twbis the formulation wet bulb temperature. Dropletsare assumed to be at their wet-bulb temperatureof the propellant and remain at this temperaturethroughout their life history. Tref is the referencetemperature which is calculated with the one-third rule of Sparrow and Gregg,28 that is,

Tref � Twb � (T � Twb) (5)

where T � 20°C is the ambient temperature.A representative geometric model of a pMDI

with an appropriate spacer (4.2-cm diameter and10.5-cm long) was constructed and connected toa human upper airway model14 which wasadapted from Cheng et al.15 The diameter varia-tions along the present oral airway from mouthto trachea are almost the same as the hydraulicdiameters from the human cast reported byCheng et al.15 Variations to the actual cast include

1�3

cp � (Tref � Twb)��hfg

the circular cross sections, a short mouth inletwith a diameter of 2 cm, a modified soft palate,and a strong bend.

A typical pMDI consists of a canister contain-ing the formulation (pcanister � 10 atm) which isconnected to a metering valve producing dosagesof 25 to 100 �L. The user triggers the release ofthe measured formulation through the actuator-nozzle, generating a liquid jet (vnozzle 150m/sec) which breaks up in two stages where secondary-breakup (ETAB model) determinesdroplet-spray characteristics. Air was assumed toenter through the annular space between canis-

ANALYSES OF pMDI AND DRUG–AEROSOL TARGETING 299

FIG. 4. Distributions of velocity profiles and turbulencekinetic energy in the inhaler, spacer, and human upperairway.

ter and inhaler body (see Fig. 1b). The nozzle,with an orifice diameter of 0.5 mm and 0.25 mm,is located where the center lines of the canisterand mouthpiece intersect. Roundoff of compo-nent edges was implemented to reflect a realisticdesign.

Numerical method and model validations

The numerical solutions of the governingequations for airflow, droplet formation, anddroplet trajectories, subject to appropriateboundary conditions, as well as models for airturbulence and droplet evaporation, were car-ried out with a user-enhanced commercial fi-nite-volume code CFX 10.0 (Ansys, Inc.,Canonsburg, PA). The final mesh topology (Fig.1a) using the mesh generator ICEM-CFD(ICEM-CFD Engineering Inc., Berkeley, CA)was determined via mesh refinement until gridindependence of the solutions was achieved. A

KLEINSTREUER ET AL.300

particle number-independence test was alsoconducted in order to verify that 2000 particleswere sufficient. Specifically, particle deposi-tion, including deposition efficiency, was testedto be independent of the number of particles re-leased. The flow field solutions were assumedto be converged when the dimensionless massand momentum residual ratios were �0.0001.All computations were performed on a dualXeon Intel 3G Dell workstation (CFPD Labora-tory, Department of Mechanical and AerospaceEngineering, NC State University) as well as onan IBM Linux Cluster at NCSU’s High-Perfor-mance Computing Center (Raleigh, NC). Thebasic boundary conditions were as follows (seeFig. 1):

1. Inlet: uniform velocity profile for Q � 30L/min and a turbulence intensity value of 5%for the turbulence kinetic energy.

2. Outlet: given pressure condition.

FIG. 5. Simulation result of CFC propelled pMDI droplet transport and deposition.

FIG. 6. Simulation result of CFC propelled pMDI droplet transport and deposition with spacer.

3. Walls: perfect particle absorption condition,that is, particles that touch the wall are as-sumed as deposited and will not be consid-ered further.

4. Actuator nozzle: given droplet initial diameter,velocity, and spray angle (see Table 1).

Validation of the present computer simulationmodel with experimental data sets included spher-ical microparticle deposition in the human upperairways, that is, mouth to trachea (Fig. 2), as wellas droplet-spray evolution in terms of a mean ax-ial velocity profile (Fig. 3a) and transient spraypenetration depth for two pressure levels (Fig. 3b).As expected, Figure 2 shows that the particle de-position efficiency increases rapidly for 103 IP 104, where the impaction parameter, IP � Q � d2

p,features two key quantities affecting microparticledeposition, that is, particle diameter and flow rate.It should be noted that the inhaler design is not in-cluded in the simulation results presented in Fig-

ure 2. Figure 3 confirms that the ETAB model usedfor simulating droplet formation is appropriate;specifically, at x � 600 D downstream from thenozzle exit with orifice diameter D, the axial ve-locity profile averaged for all spray dropletsmatches experimental data points reasonably well.Figure 3b depicts two important aspects in the sim-ulation of droplet spray behavior, that is, how farand how fast a spray front expands for differentback pressures.

Necessary geometric, operational, and prop-erty data for both the pMDI and propellants aresummarized in Table 1.

RESULTS AND DISCUSSION

pMDI initiated air-droplet dynamics

The airflow field in terms of the mid-planestreamlines and local velocity magnitudes is

KLEINSTREUER ET AL.302

FIG. 7. Simulation result of HFA propelled pMDI droplet transport and deposition.

FIG. 8. Simulation result of HFA propelled pMDI droplet transport and deposition with spacer.

shown in Figure 4. Air is uniformly sucked in viathe inlet ring at Q � 30 L/min, forming severalvortices inside the inhaler body and attachedspacer. Due to the sudden expansion, the airflowvelocity in the spacer (ID � 4.2 cm) is reduced,increasing coupled droplet residence time andhence evaporation rate. Air accelerates again inthe oral cavity (mouth diameter 2 cm) and be-comes fully turbulent in the trachea because ofthe constriction in the larynx. Before leaving thetrachea, the airflow has largely been relaminar-ized.

Considering first a CFC–pMDI for droplet trans-port and deposition, Figures 5 and 6 depict thesimulation results without a spacer and with aspacer, respectively. For CFC-driven inhalers, theexit nozzle diameter, which greatly influences thedroplet formation process and resulting spraycharacteristics, is typically 0.5 mm. As depicted inFigure 5a, the dense spray cone distributesdroplets almost fully in the oral cavity where mostthe droplets deposit, that is, about 90% (Fig. 5b).As a result, only a minor portion travels down thetrachea (Fig. 5a), leading potentially to a drug de-position in the lung of about 5% (Fig. 5b). Attach-ing a spacer (Fig. 6a and b) improves the situationmeasurably. First of all, the extra tube reduces boththe droplet spray speed and average droplet sizesignificantly; hence, it allows for broader dropletdispersion in the human upper airways (Fig. 6a).Although almost 40% of the released droplets de-posit in the spacer, over 50% make it into the lungregions. Experimental data also shows that byadding spacers, the deposition in the oral airwayscan be reduced significantly.9,10

Switching to HFA as a propellant in a pMDIwith dnozzle � 0.25 mm improves the potential ofdrug aerosol reaching the lung dramatically with-out or with a spacer (Figs. 7 and 8). Specifically,because of better atomization achieved by usingHFA and a smaller nozzle orifice, the smallerdroplets are better dispersed throughout the air-ways (Fig. 7a), and as a result, oral airway depo-sition (53.4%) is almost balanced by the through-put (46.6%) for a pMDI without a spacer (Fig. 7b).When a spacer is attached, hardly any dropletstouch the tube wall (Fig. 8a and b) because the0.25-mm nozzle opening generates more confinedspray characteristics. As a result, almost 75% ofthe released droplets travel to the lung.

The present simulation results are in agree-ment with both in vitro and in vivo tests (Fig. 9).Clearly, HFA–pMDIs perform much better than

CFC–pMDIs, that is, oral airway depositions are40–60% (HFA) versus 89–90% (CFC). Any dis-crepancies between experimental and theoreticalresults may stem from three major factors:

• Droplet coalescence was not included in thepresent model, that is, the axial mean dropletvelocities may be underestimated due to alower particle inertia effect.

• Assumed pMDI spray conditions, for example,droplet velocity at nozzle exit, spray angle, andliquid mass flow rate, may deviate from actual(experimental) conditions.

• Use of models, for example, ETAB, which donot fully match actual liquid-particle dynam-ics processes.

When reducing the exit-nozzle diameter for theCFC-propelled pMDI to 0.25 mm, the potential

KLEINSTREUER ET AL.304

FIG. 9. Comparison between the simulated and mea-sured pMDI droplet deposition fractions.

deposition efficiency increases from 5.2% (Fig. 5b)to 23.2% (Table 2); but, it is still half of that forcomparable HFA–pMDI (Fig. 7b and Table 2).Clearly, HFA properties such as its lower surfacetension and boiling point (Table 1) help in pro-ducing smaller droplets.

Smart inhaler air-particle dynamics

While HFA–pMDIs with spacers may cause upto 75% of released drug aerosols to pass throughthe trachea (Fig. 8b), their subsequent transportand deposition are arbitrary, that is, pMDIs asmost all inhalers are nondirectional. This is not ofgreat concern when routine asthma treatment isrequired, because the drug is inexpensive andusually no systemic side effects occur. MultiplepMDI actuations should ultimately supply suffi-cient medicine to the inflamed airways, if they arelocated in the upper bronchial tree.

However, in cases where expensive and/or ag-gressive drug aerosols have to reach a specificpoint in the lung (e.g., to coat a tumor) or a spe-cific lung region (e.g., to achieve rapid drug ab-sorption of, say, insulin), targeted drug–aerosoldelivery would be most desirable. This can be aachieved with a new “controlled air-particlestream” methodology. Predetermined lung sites,associated with a specific disease or treatment,can be targeted with an appropriate inhaled airstream with suitable drug–aerosol characteristics,in terms of density and size, as well as release po-sition.16,17 The necessary device to generate suchcontrolled air-particle streams is being patentedby NCSU.

As an example, an upper airway model con-sisting of the oral airway model and a symmet-ric, triple-bifurcation lung airway model repre-senting generations G0 (trachea) to G3 (see Fig.10), was employed to investigate the transportof inhaled drug aerosols. The detailed descrip-tion of this upper airway model can be found inZhang et al.29 The velocity field and particle dis-

tributions in the trachea of the oral airway modelwere adjusted as the inlet conditions of the bi-furcating airway segment (i.e., Generations G0to G3).19,29 The air flow and microparticle trans-port were simulated with a commercial finite-volume code CFX4 (Ansys, Inc.) and our in-house off-line F90 particle trajectory code.29 Thecomputations were conducted on an IBM p575machine with multiple POWER5 processors. Theaccurate prediction of the conditions for a con-trolled air-particle stream is a prerequisite fordesigning any “smart inhaler system.” As a first

ANALYSES OF pMDI AND DRUG–AEROSOL TARGETING 305

TABLE 2. COMPARISON OF DIFFERENT

PMDI SPRAY CONDITIONS

Deposition efficiency in Nozzle diameter the lung regions

CFC pMDI 0.5 mm 5.2%HFA pMDI 0.25 mm 46.6%CFC pMDI 0.25 mm 23.2%

FIG. 10. 3D views of the oral airway model and the bi-furcation airway model (Generations G0 to G3). B1—firstbifurcation, B21 and B22—second bifurcation, B31, B32,B33 and B34—third bifurcation (the dashed lines indicatethe segmental boundaries).

FIG. 11. Comparisons of simulated and measured particle distributions with different inlet release positions.

A

B

step, the simulated particle distributions in theoral airway model with different inlet releasepositions were compared with experimental vi-sualizations (see Fig. 11), assuming a steady in-spiratory flow rate of 8 L/min and a sphericalparticle diameter of 7 �m. One thousand parti-cles were released at the oral inlet. The experi-mental measurements, a proof of concept, wereconducted in-house using LDV (Laser DopplerVelocimetry) technique. Clearly, the simulatedparticle distributions agree well with the exper-imental visualizations. Minor discrepancies maybe attributed to: (1) slight differences in inlet re-lease positions between the simulations and ex-periments; (2) differences in visualization loca-tions; (3) differences in the geometries (e.g.,there was a transition tube in the experimentalsetup and the airway geometry becomes slightlydifferent after manufacturing); and (4) the dif-ference in air humidity, particle size, and flowrate, that is, particle distributions are sensitiveto factors (1) and (2).

The specific inlet positions of aerosols whichland on different targeted sites are determinedvia ‘’backtracking,’’ and then release-controlledair-particle streams are generated so that mostaerosols deposit in the desired lung regions, for example, inflamed left or right lower airways,

etc. Specifically, depositions of 7 �m particles(a representative size of drug aerosols) with arelatively low inhalation flow rate (i.e., 8L/min) are minor in the upper airways due tonegligible inertial impaction (DForal � 1.47%,and DFG0-3 � 0.76%). Thus, most of them canenter the deeper lung regions. In general, par-ticles released from the left and right sides ofthe circular mouth entrance enter the left andright lungs, respectively. However, the inlet po-sitions of particles leaving different portions ofG3 vary irregularly due to the effects of sec-ondary flows.

The feasibility of an inhaler based on the newdrug-aerosol targeting methodology may be en-hanced if the particles can be released from orbitpoints with one or more critical radii. In this case,a tube-aligned nozzle with adjustable particle-outlet diameter can rotate on a predetermined or-bital path. For example, the targeted regions forrelease positions 1 to 4 shown in Figure 12 are lo-cated in the lower airways after generation G3(see Fig. 13 and Table 3). About 3000 particleswere released from each position at the oral in-let. Clearly, particles leaving from G31 and G34may enter side and central parts of the left lung,

ANALYSES OF pMDI AND DRUG–AEROSOL TARGETING 307

FIG. 12. Schematic of particles released with a criticalradius.

FIG. 13. Distributions of particles leaving different tubesof generation G3 with different given release positions atthe mouth inlet (Qin � 8 L/min and dp � 7 �m; dry air).

respectively, while they transport into side andcentral portions of the right lung after exitingfrom G35 and G38 (see Fig. 10). As summarizedin Table 3, with the controlled inlets the captureefficiency of particles in the targeted areas can in-crease from about 10% to 60–100%. Distributionsof particles entering targeted (outlet) airways aredepicted in Figure 13. Some particle dispersionoccurs for Inlet positions 2 to 4 because of the in-fluence of secondary flows and lack of humidity.In contrast, inlet position 1 happens to be locatedin a larger particle release area reaching the G31airway outlet. In any case, dispersion could begreatly avoided when using drug aerosols withmildly attractive surface properties via surfactantcoating. Furthermore, the relatively high humid-ity environment in lung airways should reducemicroparticle dispersion.

Clearly, the development of smart inhalersbased on the new methodology may encounterseveral challenges, such as intersubject variabil-ity, device robustness, and nonuniform drug–aerosol characteristics.

ACKNOWLEDGMENTS

This effort was sponsored by the Air Force Of-fice of Scientific Research, Air Force MaterialCommand, USAF, under Grant Number FA9550-04-1-0422 (Dr. Walt Kozumbo, Program Man-ager) and NIH Grant 8R21EB006717-02. The U.S.Government is authorized to reproduce and distribute reprints for governmental purposes

notwithstanding any copyright notation thereon.The use of both CFX software from ANSYS Inc.(Canonsburg, PA) and the IBM p690 at the HighPerformance Computing Center at North Car-olina State University (Raleigh, NC) as well as thesupply of the experimental observations (see Fig.11) by Prof. W.L. Roberts and his MS student C.R.Rojas (MAE Dept., NCSU), are gratefully ac-knowledged as well. The views and conclusionscontained herein are those of the authors, andshould not be interpreted as necessarily repre-senting the official policies or endorsements, ei-ther expressed or implied, of the Air Force Officeof Scientific Research.

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TABLE 3. REGIONAL PERCENTAGE OF INHALED AEROSOLS (%)

Controlled inlets

PositionsNormal Targeted 1 2 3 4

Region inlet area G31 outlet G34 outlet G35 outlet G38 outlet

Deposition in the 2.23 0.22 2.09 0.08oral and G0-3

Exit G31 12.36 — — —Exit G32 14.29 — — —Exit G33 12.48 35.63 — —Exit G34 8.93 64.15 — —Exit G35 13.23 — 89.51 0.81Exit G36 14.75 — 8.38 —Exit G37 11.85 — 0.02 —Exit G38 9.88 — — 99.11

Qin � 8 L/min, dp � 7 �m.

0100————————

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Received on November 28, 2006in final form January 28, 2007

Reviewed by:Edgar Matida, Ph.D.

Jeffry Schroeter, Ph.D.

Address reprint requests to:Dr. Clement Kleinstreuer

Department of Mechanical and Aerospace Engineering

and Department of Biomedical EngineeringNorth Carolina State University

Raleigh, NC 27695-7910

E-mail: ck@eos.ncsu.edu

ANALYSES OF pMDI AND DRUG–AEROSOL TARGETING 309

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