Cascade Impactors for the Size Characterization of Aerosols From Medical Inhalers- Their Uses and...

JOURNAL OF AEROSOL MEDICINEVolume 16, Number 4, 2003© Mary Ann Liebert, Inc.Pp. 341–377

Cascade Impactors for the Size Characterization of Aerosols from Medical Inhalers:

Their Uses and Limitations

JOLYON P. MITCHELL, Ph.D., and MARK W. NAGEL, H.B.Sc.

ABSTRACT

Cascade impactors, including the multi-stage liquid impinger, are by far the most widely en-countered means for the in vitro determination of the particle size distribution of aerosolsfrom medical inhalers, both in product development, batch release and in applications withadd-on devices. This is because they directly measure aerodynamic size, which is the mostrelevant parameter to describe particle transport within the respiratory tract. At the same time,it is possible to quantify the mass of active pharmaceutical ingredient in different size rangesindependent of other non-physiologically active components of the formulation. We beginby providing an overview of the operating principles of impactors and then highlight the var-ious configurations and adaptations that have been adopted to characterize the various classesof inhaler. We continue by examining the limitations of the cascade impaction method, inparticular looking at potential sources of measurement bias and discussing both appropriateand inappropriate uses of impactor-generated data. We also present a synopsis of current de-velopments, including the Next Generation Pharmaceutical Impactor, and automation of cas-cade impactors for routine inhaler performance measurements.

Key words: impactor, impinger, inertial fractionator, medical inhaler, particle size analyzer

341

INTRODUCTION

CASCADE IMPACTORS (CIs), including the multi-stage liquid impinger (MSLI), are the most

commonly encountered group of instruments forin vitro size-analyzing aerosols produced by med-ical inhalers. They are the equipment of choice inboth U.S. and European Pharmacopeias,1,2 andare also recommended in current ‘guidance’ doc-uments for industry published by the corre-

sponding regulatory authorities.3,4 The exceptionis the size-characterization of aqueous nasalspray-pumps, since these inhalers generatedroplets that are in general larger than the mea-surement range capability of CIs.

At a fundamental level, a CI should not be con-sidered an in vitro lung simulator, as it operatesat constant flow rate, in contrast with the contin-ually varying flow rate associated with thebreathing cycle. Nevertheless, this class of in-

Trudell Medical International, London, Ontario, Canada.

Invited Paper

strument, if used correctly, can provide particlesize information, based on aerodynamic diame-ter (dae), which may be indicative of the likely de-position of active pharmaceutical ingredient(API) in the respiratory tract. dae takes into ac-count the influences of both particle density andshape and is related to particle physical size (dp)measured by microscopy through the equation:

ÏCae dae 5 dp 3 41/2

(1)

where the Cunningham slip correction factors(Cae and Cp) are close to unity when dp . 1.0 mm,and the dynamic shape factor (x) is unity forspherical particles.5 If the model of Rudolph etal.,6 which assumes oral tidal breathing, is cho-sen as an example, depending on inhalation flowrate, particles with dae larger than approximately6–8 mm deposit mainly in the oropharyngeal re-gion, central (bronchial) airway deposition peakswith particles having dae of 7–9 mm, and periph-eral (alveolar) deposition in the lung reaches amaximum with particles having dae of 2–4 mm.7

Since CIs that operate at ambient pressure are ef-ficient size-separators in the range of 0.3–12 mmin aerodynamic diameter,5 they are well suitedfor characterizing almost all oral inhaler-pro-duced aerosols. Although particles with dae finerthan 0.3 mm can be analyzed with adequate res-olution by CIs having some stages operated atsub-atmospheric pressure (so-called “low-pres-sure impactors”),8 such ultrafine particles are oflittle interest in the context of inhaler characteri-zation, because they contain almost no mass ofAPI, and furthermore they are likely to be exhaledif they do not deposit either by phoretic or diffu-sion-based mechanisms.7 Such CIs will thereforenot be discussed further. Of greater significancefrom the standpoint of inhaler testing is the in-fluence of gravitational sedimentation on the mo-tion of larger particles, which imposes the uppersize limit for CI measurements in the region of20–25 mm in aerodynamic diameter.9 The in-creasing effect of gravity on the inertial size-sep-aration of larger particles has so far proved to bea barrier to impactor development, and in conse-quence, commercially available CIs have littlesize-discriminating capability for particles largerthan this limit.9 They are therefore unsuitable tocharacterize aqueous droplet sprays produced bymechanical pump-driven nasal drug delivery de-vices,10 other than to quantify the mass fraction,typically defined as being ,10 mm aerodynamicdiameter, that is likely to penetrate beyond the

rpCp}xr0

nasal cavity.11 Elutriators that operate on theprinciple of sedimentation have been developedfor size separating larger particles,12 but these de-vices have not so far been widely applied to thetesting of either nasal or oral inhalers.

OPERATING PRINCIPLES

A typical CI comprises several stages, each ofwhich functions as a size-separator or fractiona-tor of the incoming aerosol in a gas stream mov-ing at constant velocity (flow rate). In concept, asingle stage impactor comprises a jet or nozzleplate containing one or more circular or slot-shaped orifices located a fixed distance from acollection surface that is usually horizontal (Fig.1). The stage functions by classifying incomingparticles of various sizes on the basis of their dif-fering inertia, the magnitude of which reflects theresistance to a change in direction of the laminarflow streamlines.5 As the incoming flow passesthrough the nozzle plate, the streamlines divergeon approach to the collection surface, whereas thefinite inertia of the particles causes them to crossthe streamlines. The dimensionless Stokes num-ber (St), which is the ratio of the stopping dis-tance of a particle to a characteristic dimension,in this case the nozzle diameter, W (or averagediameter, for a multi-orifice stage), describes theprocess, defining a critical particle size that willreach the collection surface for a particular stagegeometry.5 The theory underlying impactor func-tion has been developed over the past 25 yearsby solving the Navier-Stokes equations definingthe gas flow field in the absence of particles, andthen using Newton’s equation of motion to modelthe passage of different sized particles throughvarious stage geometries.13–15 Marple et al.9 haverecently summarized the current status of im-pactor theory, so that only the essentials are given

MITCHELL AND NAGEL342

FIG. 1. Cross-section through a single stage impactor il-lustrating principle of operation.

here. These concepts have been applied to the de-sign of the Next Generation Pharmaceutical Im-pactor (NGI, MSP Corp., Shoreview, MN).16

For a single nozzle (jet) impactor, St is relatedto W through the expression:

St 5 5 (2)

in a self-consistent set of units, based on eitherparticle physical or aerodynamic diameter.

Equation 2 predicts that the particle collectionefficiency (E) of an ideal impactor stage, ex-pressed as a percentage, will increase in a step-wise manner between limits of zero to 100%. Inpractice, for a well-designed stage, E is a monot-onic sigmoidal function of St or dae that increasessteeply from E of <0% to .95%, reaching its max-imum steepness when E is 50% (Fig. 2). At thislocation, defined as the cut size (d50):

ÏC50 d50 5 3 41/2

ÏSt50 (3)

or in terms of volumetric flow rate (Q):

ÏC50d50 5 3 41/2

ÏSt50 (4)9pmnW3}4r0CaeQ

9hW}r0CaeU

r0Caed 2aeU}}

9 mWrpCpd2pU}9 mW

for a multi-orifice stage comprising n circularnozzles.

It is possible to take into account the shape ofthe actual collection efficiency curve of the stagein the analysis of impactor data,17 but this re-finement is rarely done for measurements of in-haler performance. Instead, the assumption ismade that the mass of particles larger than d50(the size corresponding to E50), that penetrate thestage, is exactly compensated by the mass asso-ciated with particles finer than this size, that arecollected. Thus, the cut size can be defined as asingle valued constant for a given stage at a fixedflow rate. Particles with of dae of $ d50 are as-sumed to be fully collected, whereas all particlesfiner than d50 are deemed to penetrate the stage.

The so-called “sharpness-of-cut” of the stage canbe defined as the geometric standard deviation (GS-Dstage) of the efficiency curve by analogy with theproperties of the log-normal distribution function:

GSDstage 5 !§ (5)

GSDstage for a well-defined stage is ideally lessthan 1.218 (a GSDstage of unity corresponds to the

d84.1}d15.9

CASCADE IMPACTORS AND SIZE OF AEROSOLS FROM INHALERS 343

FIG. 2. Impactor collection efficiency curve, showing parameters characterizing stage performance.

ideal size-separator). However, with many exist-ing designs of CI, GSDstage values are in excess ofthis limit, particularly for those stages that size-fractionate particles larger than about 5 mm inaerodynamic diameter where gravitational set-tling contributes significantly to the size-separa-tion process.19

Marple and Liu13 and Rader and Marple,15

have identified that the value of ÏSt at E50(ÏSt50) should be close to 0.49 for well-designedround-nozzle impactors, where differences inparticle inertia dominate the size separationprocess. However, at least two other parametersalso appear to be important for effective size-sep-aration. The ratio of nozzle-to-collection surfacedistance/nozzle diameter (S/W) describes thegeometry of the stage. ÏSt50 (and hence d50) isunaffected by small variations in S, if S/W . 1.0.At the same time, the dimensionless flow Reynoldsnumber (Ref), defined as:

Ref 5 (6)

should be in the range of 500–3000 to optimizeGSDstage. Further unpublished investigationshave indicated that S/W can be between 1 and10 for effective size-separation to occur, allowingconsiderable latitude in stage design.20 There istherefore some freedom in establishing the pre-cise location of the collection surface beneath thenozzle-plate with many impactor designs. Moreimportantly from a user’s standpoint, the pres-ence of a coating to improve particle collectionbehavior is unlikely to affect stage performance.The ratio of nozzle throat length (T) to W can alsoinfluence size separation efficiency,13 decreasingÏSt50 with increasing T/W. However, the effectis likely to be small with commercially availableimpactors, where T/W is typically ,10.20 Eventhough these criteria provide the basis for a soundimpactor design, it is notable that S/W is ,1.0 forboth stages 0 and 1, and Ref is below 500 (110 ,Ref , 394) with stages 0 to 6 of the widely usedAndersen 8-stage CI ([ACI] ThermoAndersenInc., Smyrna, GA) at the design flow rate of 28.3L/min (1 ACFM).21

Cross-flow induced by air exiting the nozzlesnear the center of the nozzle plate and flowingoutward past other air jets located near the pe-riphery of the nozzle cluster, can prevent the airjets near the edge of the cluster from reaching theimpaction plate with multi-nozzle designs.22 Un-der these circumstances, particles that would oth-

raUW}m

erwise be collected are instead captured by thecross-flow and transferred beyond the collectionsurface. Increased inter-stage losses as well asbias towards smaller sizes can therefore resultwith multi-stage impactors.20 The dimensionlesscross-flow parameter (Xc), is defined as:

Xc 5 (7)

Xc should be ,1.2 to avoid cross-flow relatedproblems, and this criterion is achieved for all ofthe commonly encountered CIs used with inhalertesting, with the exception of stage 2 of the ACI,where Xc is 1.2.21

Several individual impaction stages are con-nected together in a CI, most commonly in a ver-tical stack, but it is notable that the NGI has allof its stages located adjacent to each other in thehorizontal plane, primarily for ease of use insemi- or fully-automated operation.16,18 Since thepurpose of a CI is to fractionate the incomingaerosol into progressively finer particle sizes, be-ginning with the coarsest particles, U is increasedfrom one stage to the next, primarily by reducingthe nozzle diameter. The number of nozzles perstage as well as the number of stages within theCI can also be adjusted to optimize size resolu-tion. However, in practice, inserting more thanfive stages per decade of particle size is counter-productive, because the stages immediately be-fore a given stage will interfere with efficient par-ticle collection, due to the non-ideal nature oftheir collection efficiency curves. This restrictionlimited the number of stages of the NGI with cutsizes in the range of 0.5–5 mm in aerodynamic di-ameter to five.16

DATA INTERPRETATION FROM CIs

Fundamentally, CIs determine particle aerody-namic mass-weighted size distributions5,9 fromwhich several parameters can be derived in or-der to quantify inhaler performance. The MMAD,as the measure of central tendency of the size dis-tribution, is the most often reported value. Berget al.23 have suggested that, for the purpose ofstandardizing inhaler testing, the MMAD shouldbe based on that portion of the dose entering theCI, rather than the total emitted dose ex inhaleractuator. This approach is valid if the intention isto use the size distribution data to estimate lungdeposition. However, the determination ofMMAD should include the components of the

nW}4Dc


dose depositing in the induction port (and pre-separator if used), when the purpose of the mea-surements is to evaluate add-on devices for in-halers that are intended to remove the portion ofthe dose likely to deposit in the oropharynx.24

In addition to determining MMAD, it is oftenhelpful to have a single parameter that quantifiesthe spread of the size distribution. If the portionof the dose entering the impactor is considered,the size distribution can often be approximated toa log-normal function (particularly with solutionpressurized metered-dose inhaler (pMDI)–basedformulations,25,26 and the spread of the distribu-tion, defined by the geometric standard deviation(GSD), can be estimated as:

GSD 5 !§ (8)

where d15.9 and d84.1 are the sizes correspondingto the mass-percentile values of 15.9% and 84.1%,respectively, for the cumulative size distribution.A degree of caution has to be exercised in the useof GSD values, since there is no fundamental rea-son why inhalers should generate aerosols havinglog-normal particle size distributions. For in-stance, the coarse ballistic fraction results in a sec-ond mode in the size distribution of pMDI-gen-erated aerosols when the entire dose from isconsidered.23 Deviations from log-normality havealso been observed with size distribution databased only on mass entering the CI for at least onepMDI-generated formulation, in this instance be-ing associated with a large number of drug freesurfactant particles.27 Bi-modality has been re-ported when considering the entire dose leavingthe inhaler with certain carrier-based DPI-gener-ated aerosols,28 rendering the use of GSD as a de-scriptor inappropriate unless some restriction isplaced on the size range being considered.23

In an effort to minimize the amount of data re-quired for the description of inhaler-basedaerosols, Thiel,29 like Berg et al.,23 proposed sep-arating the dose captured in the induction portand pre-separator (if used), from the “non-ballis-tic fraction” (NBF) entering the impactor. MMADand GSD were both used to define the portion ofthe dose entering the CI, as has already been de-scribed. The NBF became the fraction of the totalemitted dose ex inhaler mouthpiece that com-prises the area under the log-normal curve de-fined by the MMAD and GSD. Thiel29 then usedvalues of MMAD, GSD and NBF, together withthe inhaled volume, inhalation rate and breath-hold, and entered into the Stahlhofen-Rudolph

d84.1}d15.9

model of lung deposition,6 to show that these pa-rameters have predictive value for lung deposi-tion for representative examples of both pMDIsand DPIs, supported by g-scintigraphic evidencefrom four separate clinical studies.

In the context of judging similarity of CI-mea-sured particle size distributions from different in-halers, a working group at the Product QualityResearch Institute (PQRI) has begun exploringthe use of statistical methods based on modifiedx2 analyses of CI-stage data to develop a robustmetric for their comparisons.30 PQRI is a collab-orative process involving representatives fromthe pharmaceutical industry, academia and theFDA, and the focus of this activity is concernedwith comparing inhalers from manufacturers ofinnovator and generic formulations for regula-tory purposes. Their methodology is only ap-plicable where large databases from CI measure-ments are available, but it may offer a means tocompare test and reference products with a met-ric that does not depend on particular producttype or particle sizing equipment. However,Clark and Kadrichu31 also investigated the ap-plicability of both the f2 similarity factor and x2

statistic in a comparison of ACI- and MSLI-gen-erated size distribution data with several refer-ence distributions having MMAD values in therange of 1–8 mm with a constant GSD of 1.2. Bysubsequently computing lung deposition data forthe reference and test distributions using theStahlhofen-Rudolf model,6 they showed that nei-ther statistic responds to changes in size distrib-ution in a meaningful way in terms of lung de-position. As an alternative methodology, theydeveloped the concept of theoretical depositionfraction (TDF) of the dose, obtained by multiply-ing the mass-percentage of API collected on eachCI stage by the lung deposition probability basedon the mean diameter for that stage. They vali-dated this approach by demonstrating a sensitivecorrelation between TDF and differences in stagedeposition patterns in ACI and MSLI data be-tween laboratories that had participated in around-robin inter-laboratory study using a com-mercially available albuterol pMDI inhaler.32 Inthe light of these findings, further work is neededto establish the applicability of appropriate sta-tistical methodologies for comparing CI-gener-ated size distribution data from other inhalers.

For many purposes, it is unnecessary to go tosuch lengths to compare data from CI-testing,and it is sufficient to determine a sub-fraction ofthe dose entering the CI that is appropriate for


the therapeutic class of the API (fine particle frac-tion [FPF]). The European Pharmacopeia speci-fies the upper size limit of 5 mm in aerodynamicdiameter for FPF,2 which very often necessitatesinterpolation of the cumulative size distributiondata to arrive at the required value. In contrast,the U.S. Pharmacopeia does not specify precisesize limits for FPF, allowing the tester to choserange limits appropriate for the formulation/in-haler being tested.1 This approach recognizes thatformulations in different therapeutic classes mayhave optimum efficacy in different size ranges. Inpractice, the upper size limit for FPF is set at ei-ther 5.8 or 4.7 mm, corresponding to two cut-pointsizes of the widely used ACI at 28.3 L/min toavoid interpolation of the size distribution data.The compendial procedures only define the up-per size limit, and therefore an underlying as-sumption is made that the entire content of thedose collected in the CI smaller than this sizecomprises fine particles likely to deposit in thelungs. This supposition may be appropriate forproducts where almost all of the dose emittedfrom the inhaler is contained in particles in thesize range from 1 to 5 mm in aerodynamic diam-eter. However, there is some concern about thetreatment of some newer solution-based pMDI-delivered formulations, in which as much as 40%of the mass that enters the CI can be contained inparticles with dae , 1 mm. A significant portionof such extra-fine particles may be exhaled with-out depositing in the lung.7 In response to thisconcern, the reporting of performance data basedon two sub-fractions (fine component havingdae , 4.7 mm [FPF4.7 mm] and extra-fine compo-nent with dae , 1.1 mm [EPF1.1 mm]) has been pro-posed in a recently developed Canadian Standardfor spacers and holding chambers.24 The litera-ture contains numerous other examples of simi-lar sub-fraction definitions, for instance those ofDolovich26 (three sub-fractions corresponding toparticles with dae , 1, 3, and 5 mm). However, itshould be noted that FPF defined in this way stillincludes the entire CI contents beyond the ap-propriate stage defining the upper limit, ratherthan adopting a more logical approach based onlikelihood of lung deposition, by defining FPF interms of a range between appropriately definedsize limits. Some formulators therefore group CIstages to represent that portion of the dose mostlikely to reach the target receptors, for exampleconsidering the mass collected on stages 3–5 ofan ACI operated at 28.3 L/min, and thereby rep-

resenting particles 1.1–4.7 mm in aerodynamic di-ameter.33

The use of the term lung targetable fraction(LTF) suggested by Newhouse34 in the context ofinterpreting in vitro data is interesting, since thisdefinition is based in part on the outcome from asmall clinical investigation, rather than from at-tempts at predicting lung deposition based onmodels. This study involved the delivery of apMDI-based bronchodilator via a holding cham-ber to 20 patients with mild-to-moderate andmoderate-to-severe asthma. Newhouse34 foundthat particles with dae , 2 mm were most relevantto predict the likely response in terms of thewidely used metrics forced expiratory volume in1 sec and maximum expiratory flow of 25–75% ofvital capacity, following deposition at the recep-tors in the lungs. A limit for LTF at finer particlesizes was not reported, and the applicability ofthis definition to other therapeutic classes of in-haled medication and patient categories (in-fant/child/adult) is unknown. However, thisstudy established the principle of establishing alink between sub-fraction limits and physiologi-cal response, which is a desirable goal to attainwhere it is possible to do so.

More than particle size information can be ex-tracted from CI-based measurements. However,there are concerns about the value of such met-rics to quantify measurement quality as well asinhaler performance. Perhaps the most con-tentious issue is the CI-derived material (mass)balance. In the compendial methods,1,2 the APImaterial balance from the inhaler actuator/mouthpiece, induction port and CI is required tovalidate that the equipment operated correctlyand that the inhaler functioned when actuated.This additional piece of data can therefore pro-vide important information about the quality ofthe measurement, but it is relatively imprecisecompared with the direct measurement of dose(total emitted dose ex actuator) by the proceduresfor determining inhaler content uniformity. A CI-generated material balance is composed of sev-eral individual measurements in which sourcesof variability (e.g., sample preparation, analyticalprecision) accumulate, compared with the singlemeasurement of API mass collecting on the filterand interior surfaces of the dosage unit samplingapparatus. Current FDA Guidances to Industryfor inhaler product quality control3,35 indicatethat the mass of API recovered in the CI and ac-cessories should be within 615% of the label


claim dose on a per actuation/spray basis. Theselimits are comparable with those recommendedfor dose content uniformity, based on the meanfor each of the beginning and end determinations.Poochikian and Bertha36 have expressed the viewthat, whereas the material balance by itself doesnot provide information on the size distributionof the emitted dose from the inhaler, in conjunc-tion with dose content uniformity testing, it doesgive additional reassurance that both the CI func-tioned properly and that the analytical method-ology was performed correctly. Furthermore, ifthe CI-generated material balance is inconsistentwith the equivalent data from dose content uni-formity testing, this outcome should signal a po-tential problem with the CI system. Although itis generally accepted that the CI-material balancecannot be thought of as a system suitability cri-terion, because of the absence of reference stan-dards against which to establish CI accuracy andreproducibility,36 there are widespread concernswithin the pharmaceutical industry that the pro-posed limits are too restrictive, supported by theoutcome of a retrospective analysis of an exten-sive database including both pMDI- and DPI-based formulations.37 The relative imprecision ofthe CI-material balance, together with the per-ception of significant producer-risk that a failureto achieve the specified range might result in abatch failure, when the cause was a fault with theCI methodology, has resulted in a comprehensivereappraisal of the causes of CI-material balancefailures and the appropriateness of this metric asa batch release specification.38 This work is beingundertaken at PQRI, where a working group hasdeveloped an analysis of the various causes offailures for both material balance and particle sizedistribution measurements by the CI method.39

This consensus about what constitutes so-called“Good Cascade Impactor Practice” is part of aconcerted effort to develop the most appropriateuse of the CI-generated material balance in thecontext of inhaler product testing. In this way, di-agnosis of the cause of a material balance failureshould be able to be undertaken in a logical man-ner.

Since total emitted dose from the inhaler (TED)is also obtainable from a CI determination, thisparameter is often used in the presentation of invitro performance data. The additional informa-tion is particularly useful for the assessment ofadd-on devices with pMDIs,40,41 where it may beimportant to have an indication of the magnitude

of reduction of the coarse component broughtabout by the spacer or holding chamber. Outsideof the context of inhaler quality control testing, itcan be argued that TED may better be measuredwith the inhaler attached to a breathing simula-tor, than by sampling at constant flow rate into aCI. This supposition is certainly true when test-ing holding chambers at low flow rates, wherethe behavior of inhalation and exhalation valvescan only be properly evaluated by simulating thefull respiratory cycle.42 Whether CI or breathingsimulator alone is used to establish TED, the fineparticle dose (FPD) based on an appropriate up-per size limit can be calculated from CI-baseddata in accordance with:

FPD 5 TED 3 FPF (9)

However, it should be noted that if TED is ob-tained from CI measurements, then both FPD andTED are also subject to loss of precision from thesame cause as has been described for the mater-ial balance. Grouping the mass of API collectedon several stages to represent a therapeuticallysignificant size range33 does not eliminate thisloss of precision, unless the recovered API fromthe stages concerned is combined at assay. Underthese circumstances, variability may be reduced,arising from individual stages on which the massof API collected is close to the limit of detectionby the assay procedure.

CIs IN USE WITH INHALER TESTING

Although there are many different types of CIin use for aerosol characterization, the vast ma-jority of inhaler testing is undertaken with the im-pactors that are listed in the compendia (Table 1),and the reader is referred to these publicationsfor detailed information about their design. Thestage characteristics including values of Ref, S/Wand d50 for the ACI at 28.3 L/min, 150/160-seriesMarple Miller impactors (MMI) at 30 and 60L/min, and the MSLI at 60 L/min, correspond-ing to flow rate-CI combinations widely used forinhaler characterization, have been summarizedby Marple et al.21 Well-designed comparisons ofdifferent CIs in current use with a range of in-haler types are few, and in this context, the studyby Olsson et al.43 is significant. They comparedsize distribution data for four different formula-tions (3 DPIs and 1 pMDI) measured by ACI, 160


series MMI and five-stage MSLI, each operated at60 L/min, observing that the MMI and MSLI gavewell correlated and similar size distribution datafor all the inhalers. However, the outcome of anexamination of an extensive database of ACI- andMSLI-measured data from the budesonide Tur-buhaler DPI (n 5 40 measurements) reported inthe same paper, indicated that ACI-based sizedistributions were consistently narrower (lesspolydisperse (GSD < 1.4)) than those obtainedwith the MSLI (GSD < 2.0).

Nevertheless, MMADs measured by both CIswere quite similar, being in the range of 2.3–2.8mm (the ACI-based values clustering towards theupper end of the range with the MSLI values atthe lower end). Subba Rao et al.44 also observeda similar finding in a comparison of size distrib-utions following 10 actuations from a pMDI-de-livered peptide suspension formulation mea-sured by ACI and 150 series MMI at 28.3 and 30L/min, respectively. In their case, the GSDs were1.8 and 1.95 using the ACI and MMI, respectively;however, the MMAD determined by MMI (4.0mm) was significantly larger than the corre-sponding ACI-measured MMAD (2.9 mm). Theyattributed these differences to bias caused byhigher inter-stage (wall) losses of larger particlesin the ACI, which resulted in an over-estimationof the proportion of finer particles with this CI.Larger losses were also observed with the ACI byOlsson et al.,43 but not quantified in terms of par-ticle size. Although Subba Rao et al.44 concludedthat, for batch control purposes, absolute equiv-alency of these CIs was not critical, such a con-clusion would seem to be less apt, if the intentionbehind the particle size measurement is ulti-mately to estimate lung deposition. Perhaps theadvice of Olsson et al.,43 that caution should beexercised when comparing results from differentCIs, is the most appropriate stance to take, given

the variety of different purposes underlying thesemeasurements.

Besides the standard CI configurations, thereare some important variants that are used to en-compass the entire range of flow rates likely tobe encountered with inhaler testing. It should benoted that the ACI was developed as a room-airbacteriological sampler, operating at a fixed flowrate of 28.3 L/min.45 This impactor in its standardform is therefore unsuitable for use at flow ratessignificantly in excess of this value, as is requiredfor the testing of many DPIs. Recognizing thislimitation, Nichols et al.46 undertook modifica-tions that enable the ACI to be used at 60 L/minby the removal of stage 7 and insertion of a newhigher flow rate stage 21. Stage 21 has largernozzles than stage 0 to permit the cut size to bemaintained close to 9.0 mm in aerodynamic di-ameter at the higher flow rate. Stage 7 is removed,as its cut size at 60 L/min would be too fine tobe of much use with inhalers. The stage calibra-tion data presented with the modified ACI indi-cated comparable size selectivity to the standardACI for stages 0–4 (GSD values 1.25–1.4).46 How-ever, although the sharpness of cut for stage 21is comparable with that for stage 0, there appearsto be considerable overlap of the two collectionefficiency curves, suggesting that the new stagemay be interfering with the collection of particlesthat ought to reach stage 0. More recently, a fur-ther similar development of the ACI has takenplace, replacing stage 6 by stage 22. This modi-fication has enabled the impactor to be used atflow rates up to 90 L/min,47 which is particularlyapplicable for testing low resistance DPIs. How-ever, the pre-separator cannot be used with eitherof these modifications to the ACI, since its cutsize, which is close to 9.0 mm in aerodynamic di-ameter at 28.3 L/min,48 is significantly finer thanthe cut size of the first impaction stage at the


TABLE 1. COMPENDIAL CIS

Impactor US Pharmacopeiaa European Pharmacopeiab

Andersen 8-stage (ACI)—no pre-separator Apparatus 1 for pMDIs Apparatus DMarple-Miller series 160 (MMI) Apparatus 2 for DPIs —Andersen 8-stage (ACI)—pre-separator Apparatus 3 for DPIs Apparatus DMulti-Stage Liquid Impinger (MSLI) Apparatus 4 for DPIs Apparatus CNext Generation Pharmaceutical Impactor (NCI) Apparatus 5 for DPIs Apparatus E

Apparatus 6 for pMDIs

aProposals for the NCI.bApparatuses A and B of the European Pharmacopeia are the two-stage glass and metal twin impingers, respec-

tively. The status of these devices is currently under review.

higher flow rates. Stage cut sizes at 60 and 90L/min based on measured values for both mod-ified versions of the ACI47 are summarized inTable 2, together with the nominal values sup-plied by the manufacturer for the standard con-figuration operated at 28.3 L/min.

All configurations of the ACI make use of cir-cular-shaped flat metal or glass discs (plates) tocollect particles. A filter circle supported on the(inverted) plate of each stage can also be used asa collection substrate, but this option is seldomused for inhaler testing on account of increasedcomplexity for API recovery. There is also evi-dence that filter substrates modify significantlythe stage collection characteristics of the ACI asa result of flow penetrating the porous filter ma-trix.49,50 The simple geometry associated with arigid metal or glass plate makes recovery of theAPI a comparatively easy process compared withcollection surfaces that have depth.

The MMI series of five-stage impactors are cur-rently available in three sizes, all based on theoriginal work of Miller51 and Marple et al.,52 en-abling measurements to be made at flow rates of4.9–90 L/min with cut sizes that are all locatedwithin the useful range for inhaler testing (Table3). Stage collection efficiency curves for all versions of this impactor expressed either interms of dae or ÏSt are steep, and associated withGSDstage values that are close to or below 1.2.51–53

These impactors employ collection cups as parti-cle collectors, with the purpose of improving pro-ductivity. However, API recovery can be moredifficult than with the simpler geometry of col-

lection plates with recovery procedures requiringmore than contact with solvent to dissolve the col-lected particles. The model 160 (high flow) MMIis the standard configuration intended for use at60–90 L/min. The model 150 MMI has half thenumber of nozzles per stage compared with themodel 160 MMI, for use as an alternative to theACI for pMDI-characterization at 30–60 L/min.A low flow rate version (model 150P) was devel-oped to permit pMDIs with add-on devices in-tended for low flow patients to be tested at moreappropriate conditions of use (4.9 and 12L/min).53 Although these impactors have notbeen widely adopted, internal losses for model150 and 160 MMIs reported by Marple et al.,52

were no more than 5% of the incoming aerosol atworst case (4 , dae , 6 mm), decreasing to ,1%for finer particle sizes and ,2% for larger parti-cles. These measurements were based on calibra-tion with monodisperse droplets. The model 160MMI has subsequently been reported as havinginternal losses at 60 L/min with at least one DPI(Bricanyl Turbuhaler)54 that were comparablewith those indicated by Marple et al.,52 providedthat precautions were taken to eliminate particlebounce and re-entrainment by coating the collec-tion surfaces with a tacky surface (silicone oil).Losses within the low flow MMI have also beenreported as being less than 5% of the material bal-ance from two types of pMDI-generated formu-lations.53

The relative lack of interference between suc-cessive stages within all versions of the MMI51–53

compared with that evident particularly with theupper stages of the ACI48,55 may account for thedifferences observed between model 160 MMI-


TABLE 2. STAGE CUT SIZES (mM) FOR THE

ACI AT SELECTED FLOW RATES

Flow rate (L/min)a

Stage 28.3 60 90

22 Not used Not used 8.00b21 Not used 8.60 6.50b0 9.0 6.50 5.20b1 5.8 4.40 3.50b2 4.7 3.20 2.60b3 3.3 1.90 1.70b4 2.1 1.20 1.00b5 1.1 0.55 0.22b6 0.7 0.26 Not used7 0.4 Not used Not used

aData at 60 and 90 L/min from ref. 47 (nominal valuesfrom the manufacturer for 28.3 L/min).

bCut size reported by Thermo Andersen for this stageis 0.43 mm.

TABLE 3. STAGE CUT SIZES (mM) FOR

VARIOUS MODELS OF THE MMI

MMI model and flow rate (L/min)

150P 150P 150 160 160Stage 4.9 12.0 30 60 90

1 10.0 10.0 10.0 10.0 8.12 7.2 4.7 5.0 5.0 4.03 4.7 3.1 2.5 2.5 2.04 3.1 2.0 1.25 1.25 1.05 0.77 0.44 0.63 0.63 0.5

Data values at flow rates other than 90 L/min are basedon calibration data.46,47 Data at 90 L/min were calculated,based on the behavior of inertial collectors as a functionof flow rate.

and ACI-measured size distribution data fromfour different albuterol pMDIs reported byLeBelle et al.56 This group observed that althoughGSD values were comparable between the twotypes of CI, MMI-measured MMADs were sys-tematically finer, by an average of 18%.

The aerodynamic characteristics of the NGIwere developed from similar concepts used in theMarple-Miller series of impactors, in particularthe use of collection cups rather than flat platesand nozzle configurations to minimize the impactof cross-flow.16 Calibration data obtained with aso-called ‘archival’ NGI, having critical parame-ters (principally nozzle diameters) close to themean values specified in the design have con-firmed that the sharpness of cut for most stagesis close to the design goal of 1.2 between 30 and100 L/min.57 The micro-orifice collector (MOC)that is a substitute for the normal back-up filtermay allow significant penetration of extra-fineparticles with dae , 0.5 mm,16 but it can either bereplaced or backed-up by a conventional filterwhen analyzing the few formulations containingan appreciable portion of the dose associated withsuch particles.

The MSLI has undergone several changes dur-ing its evolution from a three-stage single nozzledevice developed by May.58 The original versionused for inhaler testing contained four stages in-cluding back-up filter.59 This version was recentlyaugmented to five stages by the addition of alower seven-nozzle stage,60 with cut sizes deter-mined by calibration with monodisperse particlesat 60 L/min to be 13.0, 6.8, 3.1, and 1.7 mm.60 TheMSLI has the advantage of not having apprecia-ble inter-stage losses. Thus, only about 1% of DPI-generated API was claimed as being lost to theinternal surfaces of the nozzles (jet tubes),60 sincethe recovery solvent for the API is also used towash down the walls of each stage. The advan-tage of having fewer stages compared with theACI, from the standpoint of reducing the time permeasurement, is offset by the lack of size resolu-tion in the critical range of 0.5–5.0 mm in aerody-namic diameter. Furthermore the sharpness ofcut (GSDstage) for MSLI stages (in the range of1.4–1.5, estimated from the reported calibrationdata at 60 L/min60) is poorer than that typicallyachieved with either the ACI or MMI. Marple etal.21 suggested from theoretical analysis thatstage D50 values for the MSLI are likely to be quitedependent on small changes in nozzle-to-collec-tion surface distance (S), given the higher than

desirable values of Ref at 60 L/min (3.3–10.3 3103). This prediction is borne out by the outcomeof an experimental assessment of sensitivity tochanges in collection plate surface level below theedge of the holder for stages 2 and 3.60 However,the presence or absence of liquid in the stagesdoes not appear to be an important factor gov-erning stage cut sizes. The purposes of this liquidare to wet the collection surfaces and dissolve thedeposited particles.58

Although not encountered as widely as the CIslisted in the compendia, several other impactorsare occasionally encountered in the characteriza-tion of inhaler-produced aerosols, and three ofthe more frequently encountered instruments areworthy of further mention.

The Delron or Battelle impactor, like the ACI,had its origins as a sampler associated with in-dustrial hygiene, but has since been used to char-acterize inhaler-based aerosols.28,61 The DCI-6 isa six-stage vertical stack design with glass col-lection plates. It incorporates a back-up filter andis operated at 12.5 L/min by means of a criticalorifice. The impactor is reported as having cutsizes of 11.2, 5.5, 3.3, 2.0, 0.9, and 0.5 mm.28 Stagecollection efficiency data based on the calibrationof a prototype impactor62 indicate that GSDstagevalues are better than 1.2 for all of the stages, andtherefore comparable with the performance of thevarious MMI designs.

The 10-stage quartz crystal impactor (QCM,California Measurements Inc., Sierra Madre, CA)has been used with some success as a means forrapid screening of pMDI-based formulations.63

Although quick to use compared with conven-tional CIs, once the reference and measurementsensors for each stage have been properly grease-coated, its major drawback is the lack of speci-ficity for API in the presence of solid or liquid ex-cipients in the formulation. The operating flowrate (0.24 L/min) is very low by comparison withother CIs, making it difficult to interface this im-pactor with conventional induction port designs.Nevertheless, Tzou and Elvecrog64 reported com-parable MMAD values for formulations that con-tain little or no surfactant by comparison withACI-based data. The GSDstage values for moststages of the QCM impactor are comparable withthose of the MMI, although the sharpness of cutfor stage 2 is poorer than this limit, resulting inappreciable overlap with the next stage.65 TheQCM has also been shown in a study withmonodisperse particles carrying between 3500


and 4000 electronic charges to be susceptible tobias caused by electrostatic charge accumulationon the insulated stage collection surfaces, in-hibiting the further collection of incoming parti-cles.65 It may therefore be necessary to take pre-cautions to minimize this effect when using thisCI to measure the charged aerosols that are en-countered with inhalers of almost all classes.66–68

However, in the context of electrostatic charge-related effects in impactors in general, it is worthnoting that Dunbar and Hickey69 did not observesignificant changes to ACI-measured droplet sizedistributions from a jet nebulizer in experimentsin which their CI stages were either electricallygrounded or isolated by their inter-stage “O”-ring seals (normal condition). This finding wouldsuggest that accumulation of droplet-suppliedelectrostatic charge within the mostly aluminumACI is unimportant. However, further work iswarranted to establish if their finding is of gen-eral applicability to other types of inhaler-pro-duced aerosols and CI types.

In recent years, it has become recognized thatnebulizer-produced droplets should ideally besize characterized non-invasively by light scat-tering techniques, such as phase-Doppler analy-sis69 or laser diffractometry.70,71 Currently avail-able light scattering methods, however, do notdetermine the mass of API and may thereforelead to misleading results, particularly with sus-pension formulations where droplets may beformed without containing any drug-bearing par-ticles.72 Current guidance for the testing of neb-ulizers published by the FDA therefore requirethat at least one of the sizing methods for nebu-lizer characterization be a CI-based method.73

Dolovich70 noted that typically the volume(mass) median droplet diameter reported fromlaser diffractometry measurements for nebulizersis 1–2 mm greater than that obtained by CI analy-sis, attributing the cause primarily to dropletevaporation in the impactor, a process which isconsidered in more detail later. Dennis et al.74

proposed that aerosols should ideally be sampledat low flow rates (,15 L/min) to avoid bias dueto evaporation by the entrainment of excessiveamounts of drier ambient air. Even when this pre-caution is taken, heat transfer from conventionalCIs that have high thermal mass is likely to biasdroplet size measurements unless care is taken toensure that the aerosol and CI temperatures aresimilar.75,76 The Marple Model 290 Personal CI77

(Thermo Andersen Inc., Smyrna, Georgia, USA)

was therefore proposed in a new CEN standard78

as the impactor of choice for the size characteri-zation of these aerosols, as it is a compact eight-stage device having low thermal mass comparedwith impactors such as the ACI. The use of ab-sorbent glass fiber filter substrates increases itscapacity to collect droplets without overloading.This miniaturized CI uses slot, rather than circu-lar nozzles for stages 1–6, so that their sharpnessof cut is slightly reduced (GSD values of moststages are in the range of 1.3–1.4). However, thepresence of fewer stages compared with the ACIreduces overlap except for the first three stages,where some starvation is apparent from pub-lished calibration data.77 The impactor was de-veloped for workplace personal exposure moni-toring of particulates, and like the ACI, hassubsequently been adapted for inhaler testing. Assuch, there are some concerns that are emergingwith increasing user experience in the pharma-ceutical industry.76 In particular, the acquisitionand fitting of the specialized filters is difficult,and filter fibers can contaminate API work-up forassay. The analytical sensitivity of the techniqueis quite low, given the fact that only just over 13%of the aerosol leaving the nebulizer can be sam-pled in accordance with the methodology givenin the CEN standard. The sampling configurationis a straightforward 22-mm internal diameter“T”-connector arrangement, but its particle size-flow rate sampling characteristics have not beenestablished. Finally, the use of sodium fluoridesolution as a surrogate for drug formulations,whilst fulfilling the intent of providing baselinedata with which to compare different nebulizerbrands,79 is unlikely to be predictive of the behavior of suspension formulations wheredroplets can be formed that do not contain API.72

The low flow impactor technique has not been in-cluded in either the European or US compendiaat the present time, although the European Res-piratory Society has endorsed the test methodol-ogy in the CEN standard as a means of provid-ing clinicians with reliable data for the widevariety of nebulizers that are in the market-place.80 The total flow rate of 15 L/min was cho-sen to match the mid-point flow of the sinusoidalpattern proposed for breathing simulator-basedtesting of these devices.79 However, there is someflexibility in the choice of this value, since themethodology in the CEN standard is claimed tobe adaptable to virtually any breathing pattern.79

Recently, Jauernig et al.,76 based on the out-


come of preliminary experimental studies, sug-gested that the NGI operated at 15 L/min maybe adaptable for characterizing nebulizer-pro-duced aerosols following an adapted CEN pro-tocol, since this CI can sample the entire aerosol.They noted that the simplest approach would beto utilize the existing design, calibrating thearchival NGI at this flow rate to establish the cor-rect stage cut sizes. However, as an alternativestrategy, they showed that by plugging 50% ofthe nozzles for each stage, they could achievecomparable droplet size distribution data from amodified e-Flow™ nebulizer, operating the NGIat 15 L/min. They used the same stage cut sizesas those proposed for the unmodified NGI at 30L/min (note that they reduced the area of the sin-gle nozzle for stage 1 by 50%). Although theseconfigurations show promise for making reliableCI-based measurements of nebulizer-produceddroplets, either option will require extensive val-idation studies, given that this impactor is beingused in an application for which it was not specif-ically intended.

CI AUTOMATION

CI measurements are labor intensive, so thatefforts have been under way in recent years toautomate their use, particularly for inhaler de-velopment and batch release testing, where alarge number of similar measurements must beperformed on a regular basis.

If CIs listed in the compendia are considered,both the ACI and MSLI are problematical to ser-vice using robotics, since they are both verticalstacks, requiring individual stages to be accessedwithout disturbing the others, and the collectionplates tightly fit their holders (typically ,0.1 mmgap). Despite these difficulties, Smith81 describeda means of overcoming the handling problem bymodifying the outside of the impactor withgrooves and location balls, enabling the robot tolocate and manipulate the individual componentsaccurately. In the same article, he presented com-parable size distribution data from an (unspeci-fied) pMDI obtained by the automated ACI withresults acquired manually, indicating that 30–40inhalers could be tested in a 24-h period by theautomated system. The NGI by virtue of its hor-izontal stage layout, including the use of theMOC in place of a filter, lends itself more read-ily to automation. Although Smith has since pro-

posed an outline scheme for undertaking thistask,82 a design for a fully automated system hasyet to appear. A group of automation expertsfrom member companies involved with the de-velopment of the NGI is also currently workingtowards the same goal.83

Fully automated CIs are highly expensive, anestimate for the automated ACI being close toUS$1M.81 As a means of controlling equipmentcost, Miller et al.84 recently proposed simplifyingthe liquid handling procedures involved withAPI recovery from CIs by making the primarymotion of the process in the liquid streams re-quired for API dissolution and cleaning (so-called“service-head” approach), and applying the con-cept in relation to the NGI. Such a methodologyavoids the need to move the components of theCI physically, resulting in a greatly simplifiedand therefore less costly approach to automation.However, for most users, lower cost aids thatspeed up the process of making CI measurementsin a semi-automated way, for instance by apply-ing stage coatings in a reproducible manner or byassisting with API recovery from individualstages, are likely to be more attractive than fullyautomated systems. Initial findings with one suchsystem indicated both excellent API recovery (ca.99% of label claim) together with greatly reducedcycle times compared with manual ACI opera-tion.85

ADAPTATIONS TO CIs FOR USE WITHINHALERS: INDUCTION PORTS

Some form of inlet is required to ensure thatthe aerosol produced by the inhaler is sampled ina reproducible manner. From a practical per-spective, most inhalers deliver their medicationin the horizontal plane and entries to CIs are inthe vertical downwards direction. The solution tothe problem is the induction port, which alsoserves the purpose of mimicking to a greater orlesser extent depending on its design, the humanoropharyngeal region. There are many designs ofinduction port (seven examples are given in Fig.3), reflecting differing viewpoints on how inhaleraerosols should be sampled,86 but by far the mostcommonly encountered is the metal right-anglebend described in both U.S. and European andPharmacopeias.1,2 In its role as a model of the en-trance to the respiratory tract, the induction portcollects almost all of the fast moving and so-called


“ballistic” component of pMDI-produced aero-sols formed by flash evaporation of the propel-lant, and therefore likely to deposit in theoropharynx.87 It also serves to remove larger andoften aggregated particles generated by mostDPIs upon inhalation. The USP/EP design is in-tended to provide a common benchmark to com-pare different formulations by standardizing itscritical dimensions (internal diameter and unob-structed path length). Unfortunately, unlike animpactor stage, its collection efficiency is not eas-ily determined, since flow is non-laminar at flowrates encountered typically for inhaler testing,and particles in the ballistic component have ve-locities greater than that of the surrounding air-flow when entering the induction port. In an at-tempt to develop a better understanding of theprocesses that influence particle collection in theinduction port, Stein and Gabrio88 investigatedquantitatively the deposition of solution-basedpMDI HFA formulations at various locationswithin the USP/EP design as a function of flow

rate within the exceptionally wide range of 5–90L/min. They discovered interestingly that iner-tial impaction is not the dominant particle depo-sition process; since turbulent deposition is moreimportant, especially at the higher flow rateswhere the turbulent intensity of the flow movingthrough the induction port is at its greatest.

Van Oort and Downey89 have argued that, inthe case of pMDIs, “blow-back” of aerosol to-wards the inhaler can be exacerbated by the smallsize of the USP/EP design. In support, Van Oortand Truman90 have since shown that depositionof pMDI-produced albuterol within this induc-tion port, expressed as a percentage of the labelclaim dose, decreases in a linear relationship withincreasing length of the entry. A configurationbased on the larger volume Twin Impinger inlet,manufactured out of either metal or glass wasproposed as an alternative to the USP/EP de-sign,89 based on experimental studies indicatingthat the proportion of the emitted dose containedin fine particles with dae , 5 mm entering the im-


FIG. 3. Various designs of induction port.86 (Reprinted from J. Aerosol Med. 11(S1), Dolovich, M., and R. Rhem. Im-pact of oropharyngeal deposition on inhaled dose, 112–115, Copyright 1998, Used with permission of Mary AnnLiebert Inc.)

pactor was increased. Earlier work by this groupwith pMDI-produced aerosols had indicated atrend towards increased fine particle transmis-sion to the impactor with increasing inductionport internal volume up to 5 L in capacity, withmost improvement evident between 250 cm3 and1 L.91 The underlying mechanism is believed tobe the increased time for solvent/propellantevaporation to be completed within the largervolume induction ports. The 1L glass entry portdeveloped recently by Sequeira et al.92 thereforeprobably represents a near optimum condition interms of allowing the aerosol plume from thepMDI sufficient room to expand, and thereby en-hancing fine particle delivery to the impactor. Intheir justification for such a large induction port,these authors claimed that the additional dosecollected in the impactor affords greater sensitiv-ity when following subtle changes in particle sizedistribution from one batch to another. This sizeof induction port should be viewed as an aid toenhance product quality control, rather than anattempt to mimic the geometry of the upper res-piratory tract as a means of predicting lung de-position, given the relatively small capacity of thehuman oropharyngeal cavity.

The SRI induction port developed by Williamsand Witham,93 and originally proposed by By-ron94 as an alternative to the USP/EP inductionport, was also an attempt at an improved designof CI entry, primarily for quality testing low re-sistance DPIs at flow rates up to 120 L/min. Thedesign intent was to enable an ACI to be oper-ated at its standard flow rate (28.3 L/min) by in-troducing a flow of clean make-up air in an ex-tended vertical section, followed by an isokineticsampling arrangement to allow flow to enter theimpactor without biasing the size distribution.Although the underlying principles appeared tobe sound, concerns were subsequently expressedabout the ability to sample the bolus of aerosolreleased from a DPI on actuation isokineticallyand the inherent variability associated with theturbulent mixing process involving the make-upair.95 The Williams and Witham inlet has not be-come adopted as a standard induction port, sincecurrent compendial methods for DPI testing1,2 re-quires that a fixed volume (4.0 6 0.2 L) of air besampled from the inhaler at a flow rate deter-mined by the resistance of the inhaler. The addi-tional dead volume introduced by this design isundesirable when the volume of air that can besampled is restricted, since it affects the charac-

teristic flow rate-time curve for the inhaler. Themove to variable flow rate testing, at least forDPIs, was spurred on by the development of im-pactors that were designed from the outset to beoperated at a wide range of flow rates.51,52 In ad-dition, the current method avoids concerns aboutsampling, since the entire dose emitted from theinhaler is collected.

The testing of breath-actuated inhalers poses aproblem with conventional induction port de-signs, since the act of connecting the inhaler tothe inlet may trigger the breath-actuation feature.In response, Brouet et al.96 recently proposed amodified USP/EP induction port that incorpo-rates a fast (5–40 msec) switching solenoid valvelocated in the horizontal section. In use, the in-haler is connected to the entry of the inductionport with the valve initially switched to permitair to enter the impactor via a by-pass port. Flowis diverted to enter the induction port via the in-haler at the appropriate time.

In addition to the challenges of breath-actuatedinhaler testing, the evaluation of add-on devicesfor pMDIs, particularly HCs, also creates diffi-culty when simulating poor patient coordination,including mistiming of the inhalation maneuverand breath-holding following inhaler actuation.In an attempt to address this concern, Mitchelland Nagel42 reported the use of an electro-me-chanically operated shutter that attaches to the in-duction port entry. This apparatus can be used tosimulate delay times from fractions of a secondto several seconds in duration between actuationof an inhaler and the onset of sampling. It avoidsmodifications to the induction port itself, andadds less than 5 mL of dead volume to the sam-pling arrangement. In principle, the delay appa-ratus could be used with almost any inductionport, although it has only been evaluated with theUSP/EP design to date.

In the context of a meeting of experts whosepurpose was the development of an industry-wide consensus towards more meaningful labo-ratory measurements for evaluating inhaler-pro-duced aerosols, Dolovich and Rhem86 noted thatin vitro sizing with the various induction portsthat are available may introduce inconsistencieswith in vivo data from various laboratories. Theytherefore proposed the adoption of an inlet de-sign that preferably most closely resembles theaerosol collection properties of the human mouth,pharynx and trachea. The USP/EP inductionport, although widely adopted, does not meet this


criterion, as its internal geometry is too simple.Nevertheless, it could be argued that for a phar-maceutical manufacturer, the acquisition of con-sistent in vitro data from one batch to the next ofa given formulation with a given design of in-duction port is more important than achievingcomparability of data obtained with different de-signs. Under these circumstances, the choice ofinduction port (whether USP/EP or another typehaving simplified geometry) should be made tooptimize the sensitivity of the CI method for theformulation being evaluated, as proposed by Se-queira et al.92 On the other hand, a number ofgroups have in the last few years recognized theover-riding importance of establishing reliable invivo correlations with laboratory-generated parti-cle size distribution measurements in the contextof evaluating likely clinical performance.97 Thisunderstanding has resulted in the creation of sev-eral induction port geometries based on both ca-daver casts and MRI imaging of live patients, inorder to model as closely as possible the clinicalsituation.98,99 The effect of induction port choiceon the dose sampled by the CI can be significant.For instance, Berg99 observed that the fine parti-cle dose (dae , 5 mm) from a budesonide pMDIdelivered by a large volume holding chamber(HC) to an ACI operated at 28.3 L/min with aglass Twin Impinger induction port was compa-rable at 35% label claim dose (200 mg) to that fromthe pMDI alone. However, when an inductionport whose entry profile was based on theanatomical model of an adult throat replaced thisinlet, fine particle dose ex HC doubled from 20%to 40% of the label claim dose compared with thatfrom the pMDI alone. This difference was morepronounced (8%, pMDI alone; 25%, pMDI 1 HC)when the induction port was changed to onebased on an anatomical child throat. At about thesame time, Olsson et al.,100 using the charcoalblock technique, compared the lung deposition ofalbuterol in healthy adult volunteers with CI-measured fine particle dose (dae , 5 mm) for threedry powder inhalers (DPIs) and a pMDI. A bet-ter in vivo/in vitro correlation was achieved withdata obtained using an adult throat replica com-pared with a glass Twin Impinger induction port,the latter significantly over-predicting the lungdeposition that was actually achieved. Their useof Brij-35 (polyoxyethylene 23 lauryl ether) sur-factant (0.75 g) in glycerol (25 g) applied inethanolic solution (7 mL) as a coating to thereplica throat is an interesting extension of the

idea of more closely modeling the in vivo situa-tion by simulating the wet mucosa of the upperrespiratory tract. More recent studies by Berg etal.101 with induction ports modeling five differ-ent human oropharyngeal anatomies, includingadult with tongue “up” or “down,” mouth openor mostly closed, as well as that of a ,3.5-year-old child, have demonstrated as much as a five-fold difference in fine particle delivery for the de-livery of CFC- and HFA-formulated fluticasonepropionate and budesonide. These differenceswere attributed to the varying geometry of the in-lets and were also found to be formulation inde-pendent. Srichana et al.102 reported an investiga-tion into the applicability of replacing the TIinduction port with a cast of an adult male throathaving wetted interior surfaces, coming to theconclusion that the cast had higher retentionwhen sampling lactose carrier-based DPI-gener-ated aerosols at 60 L/min. Given what has al-ready been presented, it is unsurprising that Mas-soud et al.98 observed significantly greaterretention of DPI-generated aerosols sampled bysmall compared with large size oropharyngealmodels made from MRI scans of patients, repre-senting extreme dimensions likely to be encoun-tered in clinical practice.

Alternatives to an induction port as CI entryare possible in the context of nebulizer testing,since these inhalers do not generate ballisticdroplets and there is no need to simulate bolusdelivery of medication, as is the case with DPItesting. For instance, the use of a 22-mm diame-ter “T”-connector constructed from componentsused in mechanical ventilator circuitry has beenproposed to enable flow from a jet nebulizer tobe sampled by a CI operating at a fixed flow rate,whilst the air flow leaving the nebulizer is variedto simulate tidal breathing.103 Although not developed as an induction port through a for-malized design process, this readily availablearrangement having reproducible internal geom-etry may be particularly useful for the testing ofbreath-enhanced or breath-actuated nebulizers,where breath-simulation is essential to get the airentrainment and/or breath-actuation feature tofunction as they would under patient use. Likemost induction port designs, the aerodynamiccharacteristics of this “T”-connector inlet havenot yet been evaluated either by computationalfluid dynamics or experimentally with monodis-perse particles. Furthermore, the CI did not sam-ple the entire aerosol produced by the nebulizer.


Nevertheless, indirect evidence of its suitabilityis given in an article in this issue,104 whichdemonstrates a correlation between lung deposi-tion based on g-scintigraphy and CI-measuredFPF,2.5 mm. These in vitro measurements weremade sampling jet nebulizer-produced dropletsat a very low flow rate (1 L/min) into a 10-stageQCM CI. Deposition within the “T”-connector isbelieved to represent the high inertia and there-fore coarse particle component of the nebulizedaerosol that is most likely to be deposited in theupper respiratory tract.

The challenge of testing oral inhalers in a real-istic manner, but at the same time preserving sim-plicity in technique has been addressed in a dif-ferent way by Miller and Purrington.105 Theydescribed the comparative in vitro evaluation ofthe metal induction port for the Marple-Miller(MMI) series of impactors with a cast adult throatreplica, utilizing the aerosol penetration charac-teristics of a variety of undisclosed CFC- andHFA-formulated solution and suspensionpMDIs. The portion of the emitted dose that en-tered the CI via the MMI induction port was ob-served to follow closely the equivalent measureof aerosol penetration through the replica throat,sampling at both 30 and 60 L/min. In an addi-tional comparison with the MMI and USP/EP in-duction ports, they discovered significant differ-ences with some formulations, especially at thelower flow rate, with greater penetration of theemitted dose from the inhaler generally, but notalways apparent with the USP/EP inductionport. Although such data support the use of theMMI induction port as a surrogate at least foradult throat geometry when evaluating pMDI-produced aerosols, further studies of this sort areneeded to establish the comparability of this in-let, or a similar shaped but smaller induction portwith anatomically correct child throat models.

In passing, it is self-evident that the use of aninduction port would be inappropriate for CImeasurements to characterize inhalers used withdevices that instill the aerosol directly to the re-gion of the carina via an endotracheal tube (ET),bypassing the oropharyngeal region. In this con-text, Mitchell et al.106 described the use of a stan-dard ACI to evaluate the performance of a hold-ing chamber fitted with an 8.0-mm internaldiameter ET, intended for mechanically venti-lated adult patients. The aerosol leaving the ETwas sampled on-axis directly into the impactorinlet cone, enabling the size distribution of the

API to be determined at the point of delivery tothe lower respiratory tract. A further refinementwas the incorporation of humidification and heat-ing of the air supply passing through the hold-ing chamber to the impactor to simulate moreclosely the conditions in a mechanical ventilatorcircuit.

PRE-SEPARATORS

A pre-separator is often required when sam-pling aerosols produced from DPIs, since theseformulations in many instances contain the APIattached to the surface of much larger glucose orlactose carrier particles. The shear forces gener-ated by inhalation detach some, but not all of theAPI particles from the carrier material, and theformer are sufficiently fine to penetrate beyondthe upper respiratory tract into the lungs.28 BothU.S. and European compendia1,2 refer to the useof a pre-separator for DPI-based particle size dis-tribution measurements with the ACI, recom-mending that its interior surfaces be coated eitherwith a tacky agent in the same way as the col-lection surfaces of the CI stages, or with up to 10-mL of a suitable solvent to eliminate particlebounce and re-entrainment. Ideally, the pre-sep-arator should not starve the first stage of the im-pactor of particles. However, the pre-separatorused with the ACI has been shown by calibrationto have its cut size close to 9 mm in aerodynamicdiameter, almost identical with that of the firststage (stage 0) at 28.3 L/min.48 Its sharpness of cut is relatively poor at this flow rate by comparison to equivalent values for the CI stages(GSDpre-sep is about 1.548), due primarily to theinfluence of gravity. The pre-separator thereforestarves the second as well as the first stage of theimpactor. Sethuraman and Hickey107 have stud-ied the size-separating performance of the ACIpre-separator at 60 L/min using computationalfluid dynamics, supported by experimental workin which polydisperse fluorescent particles in therange of 45–125 mm were sampled by this im-pactor. Their modeling revealed the existence oflow velocity locations, located particularly abovethe entry orifices, where particles can depositaway from their intended sites, thereby increas-ing wall losses. Interference in the flow betweenadjacent nozzles was cited as a cause of its poorsize-selectivity. Their experiments indicated thatcoating the internal surfaces with silicone oil (un-


specified viscosity) improved the size-selectivity(sharpness of cut) of the pre-separator, reducingthe carry-over of large particles to the impactor.Unfortunately, this group did not calibrate theirpre-separator using monodisperse particles, sothat GSDpre-sep cannot not be established. How-ever, apart from improvements in size-selectivitybrought about by coating the interior surfaces,GSDpre-sep would be expected to fall below 1.5with increasing flow rate, given the increaseddominance of inertia in the size-separationprocess at higher Ref. Hence the overlap betweenits collection efficiency curve and those of the up-permost impactor stages would be expected to bediminished at flow rates 5 60 L/min, comparedwith that observed at 28.3 L/min.

Recognizing the limitations of the ACI pre-sep-arator, the consortium developing the NGI de-signed its pre-separator to function as a singlecomponent, but with two distinct steps in thesize-separation process.108 The incoming aerosolis first passed through a so-called “scalper” im-pingement stage that removes the coarsest parti-cles. The remaining aerosol then passes immedi-ately through a more conventional impactionstage before leaving the pre-separator. Calibra-tion with monodisperse particles has demon-strated that its GSDpre-sep is close to 1.3 at 30, 60,and 100 L/min,57 which is comparable with val-ues of GSDstage for well-designed impactors. Itsmeasured d50 values (10.0, 12.7, and 14.9 mm at100, 60, and 30 L/min, respectively) are suffi-ciently separated from the corresponding valuesfor stage 1 (6.07, 8.29, and 11.4 mm) that starva-tion of the first stage of the impactor should notoccur to a significant extent.

ADAPTATIONS OF CIs WITHBREATHING SIMULATORS

Although CIs must operate at constant flowrate for the stage cut sizes to remain stable dur-ing a measurement, there have been several at-tempts to link these instruments directly withbreathing simulators in order to arrive at aerosoltransport conditions that more closely mirror theclinical situation.

In the case of DPI testing by the compendialprocedure,1,2 the inhalation portion of a singlebreathing cycle is simulated by opening a two-way solenoid valve located between the CI andpump for a pre-determined time. The flow en-

tering the CI can take a short but significant timeto reach the nominal flow rate for the test, de-pending on the magnitude of the dead volume inthe system, during which the stage cut sizesrapidly decrease to their final stable values. It isnormal, however, to treat the measurement ashaving been undertaken at the nominal flow ratethroughout the entire sampling period, so thatfixed cut sizes, such as those given in Tables 2and 3, can be used to determine the size distri-bution data. In the case of pMDI testing, it is nor-mal to couple the inhaler mouthpiece directly tothe induction port of the CI with the flow set atthe nominal value, actuating at the appropriatetime. Under these circumstances, the assumptionof stable cut sizes throughout the measurementis valid, provided that flow rate control is ade-quate (typically 65% of nominal1,2).

The testing of HCs poses some additional chal-lenges, since these devices typically have inhala-tion and exhalation valves whose opening andclosing characteristics are flow rate dependent.42

Attempts thus far to interface CIs directly tobreathing simulators for testing these deviceshave proved somewhat problematic. Fink andDhand109 described an arrangement in which aQCM impactor samples aerosol emitted from anHC into a 1-L plenum that is also connected to aventilator circuit. Although simple in concept,this configuration has the drawback that the flowvelocity profile at the inlet to the CI is constantlychanging as a result of the imposed breathing pat-tern. Sampling of polydisperse aerosols typicallyproduced by inhalers should ideally be isokinetic(i.e., the air velocity at the inlet should match thatof the air flow from which the sample is beingtaken) to eliminate size-related bias.5 The systemsdeveloped by the group at the University of Al-berta110–112 and by Foss and Keppel113 are moresophisticated, since they both attempted to sepa-rate the continually varying flow generated bythe breathing simulator and fed to the HC fromthe constant flow path required by the CI. In theFinlay and Zuberbuhler model110–112 (Fig. 4), anACI operating at 28.3 L/min was allowed to sam-ple the aerosol emitted from the HC fitted to areplica face. The signal from a stepper motor-con-trolled piston-driven breathing simulator wasused to actuate a two-way solenoid valve, push-ing sufficient flow of air into the system duringthe inhalation portion of each breathing cycle toenable the impactor to operate normally and al-low the system to ‘inhale’ from the holding cham-


ber. At exhalation, the solenoid valve closed, sothat the piston pushed flow back towards the CIand HC, correctly simulating the exhalation flowprofile at the replica face. At the same time theCI was able to continue to draw a constant flowwithout evacuating the system. This arrangementwas used to study the behavior of several HCs atlow flow rates of ,10 L/min, corresponding touse by small children.110–112 However, in somecircumstances, particularly when simulating useat low flow rates, it may be possible for pressurepulsations arising from the operation of the fast-acting solenoid valve in this arrangement to“pop” open an otherwise stuck closed inhalationvalve with at least one type of HC. This sugges-tion is based on the outcome of testing of HCselsewhere (Vent-170, SpaceChamber).114 Bothwere shown to have a tendency for valve adhe-sion preventing delivery of medication at lowflow rates, when tested using a breathing simu-lator without CI, and employing filter collectionto capture the total emitted dose at low flow rates(tidal volume , 100 mL). The system developedby Foss and Keppel113 both avoided pressure pul-sations and generated a sinusoidal breathing pat-tern, which could also be readily adapted to gen-erate other waveforms. In their arrangement, asupply of pressurized air was fed continuouslyto a “Y”-shaped connector, the other arm of

which was connected to a breathing simulator(Fig. 5). The combined flow entered a “T”-con-nector, one arm of which was coupled to the HCand the other arm that was connected to the CI.Aerosol from the HC passed through a USP/EPinduction port at variable flow rate, simulatingthe action of inhalation. After passing the straightsection of the “T”-connector, the aerosol streamattained constant velocity under the influence ofthe vacuum pump that is connected to the CI.This arrangement requires careful flow control toavoid losing aerosol in transit from the holdingchamber to the impactor, and is also limited toinhalation flow rates that are less than the flowrate required by the CI. However, the timing ofthe breathing cycle could be adjusted to simulateactuation of the pMDI both in and out of phasewith the onset of inhalation to simulate realisticuse poorly coordinated patients, for which HCsare frequently prescribed.

The Electronic Lung™ developed by Glaxo-SmithKline Plc was an attempt to allow a moreaccurate simulation of DPI performance duringpatient use than that which can be achieved withdirect constant flow rate sampling methods. Inthis apparatus, aerosol from a DPI was first “in-haled” into a 6 L capacity vessel,115 by with-drawing a piston through a chamber connectedto the chamber via a sidearm at its base (Fig. 6).


FIG. 4. Breathing simulator–cascade impactor system developed by Finlay,110 and Finlay and Zuberbuhler.111,112(Reprinted from Int. J. Pharm. 168, Finlay, W.H., Inertial sizing of aerosol inhaled during pediatric tidal breathingfrom an MDI with attached holding chamber, 147–152, Copyright 1988, with permission from Elsevier.)

A feedback mechanism ensured that the actualpressure drop across the DPI corresponded to thedemanded pressure drop associated with the in-halation maneuver being simulated, and the pis-ton movement adjusted accordingly. The aerosolwithin the chamber was therefore claimed to berepresentative of that inhaled in normal use. Fol-lowing inhalation, a solenoid valve was opened

so that the aerosol could be sampled at 28.3L/min via a standard ACI connected to the baseof the chamber for sufficient time to ensure com-plete emptying of the chamber with two completeair changes. Although this system has the ad-vantage that patient-derived inhalation flow rate-time profiles can be used to simulate actual DPIuse, it is vulnerable to size-related sampling bias,


FIG. 6. Electronic lung.115,116 (Reprinted from J. Aerosol Sci. 29(8), Burnell, P.K.P., A. Malton, K. Reavill, and M.H.E.Ball. Design, validation and initial testing of the Electronic Lung™ Device, 1011–1025, Copyright 1998, with permis-sion from Elsevier.)

FIG. 5. Breathing simulator–cascade impactor system developed by Foss and Keppel.113 (Reprinted from Respir.Care. 44(12), Foss, S.A., and J.W. Keppel. In vitro testing of MDI spacers: a technique for measuring respirable doseoutput with actuation in-phase or out-of-phase with inhalation, 1474–1485, Copyright 1999, with permission fromDaedelus Enterprises Inc.)

principally caused by gravitational sedimenta-tion within the chamber. However, validation experiments using polydisperse particles of com-parable size range to those produced by com-mercially available DPIs, indicated that particlesfiner than about 5.8 mm in aerodynamic diame-ter are sampled with high efficiency (.80%).116

Although it is well suited for DPI evaluations, theElectronic Lung is not readily applicable to thetesting of HCs, where it may be necessary to sim-ulate exhalation through the device in the contextof evaluating performance with an uncoordi-nated patient.

In the context of DPI testing, Finlay and Gehm-lich117 recently adapted the Finlay-Zuberbuhlersystem to permit low to moderate resistance DPIs(Spiros and Ventodisk [Diskhaler]) to be tested ata wider range of inhalation flow profiles simu-lating actual patient use (average inhalation flowrates varied from 21 to 162 L/min depending oninhaler type). At the same time, their system per-mitted a CI to sample the aerosol via an anatom-ically correct oropharyngeal cast at constant flowrate (Fig. 7). They were able to demonstrate theimportant finding that for these inhalers at least,CI-derived size distributions measured directly at

constant flow rate (via the oropharyngeal model)were comparable with those obtained using both“fast” and “slow” breathing patterns with theirbreathing simulator-CI system, as long as theflow rate was carefully chosen to be comparablewith those obtained in the simulations. In orderto encompass the wide range of flow rates en-countered with these DPIs, it is worth noting thatFinlay and Gehmlich used a two-stage virtual im-pactor to permit sampling via a CI at 30 L/minfrom the air stream leaving the inhaler through aconnecting tube to which make-up air could beadded to achieve a total flow rate of 285 L/min.This system probably represents the state-of-theart sampling arrangement for obtaining particlesize data from DPIs being operated under realis-tic conditions of use, but it remains to be evalu-ated with higher resistance inhalers.

Given the outcome of the Gehmlich and Finlaystudy with DPIs,117 as well as evidence that, forat least one press-and-breathe pMDI-based for-mulation (HFA-fluticasone propionate), flow ratevariations between 28.3 and 60 L/min had an in-significant impact on CI-measured size distribu-tion data,118 a case might be made for simplify-ing in vitro testing under realistic breathing


FIG. 7. High flow rate virtual impactor–CI system developed by Finlay and Gehmlich.117 (Reprinted from Int. J.Pharm. 210, Finlay, W.H., and M.G. Gehmlich, M.G. Inertial sizing of aerosol inhaled from two dry powder inhalerswith realistic breath patterns versus constant flow rates, 83–95, Copyright 2000, with permission from Elsevier.)

conditions to the assessment of TED by filter col-lection. A CI could then be used as an adjunctprocedure, testing at constant flow rates judi-ciously chosen to be appropriate for the intendedpatient user groups. However, it is recognizedthat further research is needed to ascertain howapplicable such a simplified approach would beto the testing of the wide variety of inhalers thatare available. On a precautionary note, the datafrom Smith et al.119 would indicate that althoughsuch a simplification might be justifiable withconventional press-and-breathe and breath-actu-ated pMDIs, it may not be appropriate with cer-tain DPI-delivered aerosols.

USE OF CIs WITH INHALERS FORNASAL DRUG DELIVERY

Nasal drug delivery is rapidly evolving into apopular route for administering both topical andsystemic pharmaceutical agents, resulting in re-cently issued FDA draft Guidances for Industrycovering appropriate in vitro assessments35 andbioavailability/bioequivalence issues.120 In bothdocuments, even though the API is not directlyquantified, light scattering techniques such aslaser diffractometry121,122 are recommended foraqueous nasal products using mechanicalpumps,120 since this class of inhaler typically pro-duce droplets with dp in the 20 to 200 mmrange,122 which is well beyond the region inwhich inertial size-separation methods are effec-tive.5,9 Despite this limitation, CIs are also pro-posed for the characterization of API containedin fine droplets, but mainly to provide assurancethat the mass of drug in droplets with dae of #10mm is comparable between products.123 There isinterest in quantifying this so-called “fine” com-ponent, which is typically ,5% of the dose, sincesuch droplets can pass beyond the nasal cavityleading to undesired deposition elsewhere in therespiratory tract.124 However, this type of mea-surement has low sensitivity for reasons alreadydiscussed.125

Suman et al.11 used a nasal cavity model madefrom an inverted twin-necked round-bottomflask, and improved measurement precision byusing a shortened ACI (pre-separator and im-paction stages 0, 1, and 2 and filter only) therebyincreasing analytical sensitivity for the API. Theyactuated the nasal spray upwards into the spherevia one neck of the flask, sampling vertically

downwards from the other neck located at itsbase at 28.3 L/min. They were able to demon-strate comparable fine component delivery fromtwo types of commercially available aqueousnasal spray pump devices, in that 97.8 6 2.1%and 99.8 6 3.8% of the dose from nicotine solu-tion sprays delivered by either device was con-tained in particles with dae of $9 mm. More re-cently, Guo et al.126 extended this technique toillustrate how the volume of the glass entry port,which was varied from 300 mL to 5 L in capac-ity, affected the magnitude of the fine component(,3.8% of the dose ex actuator under all condi-tions) collected by their short stack ACI. Al-though no effect was evident at 30 L/min, re-ducing the volume of the sphere slightlydecreased the magnitude of the fine componentof the dose. As might be expected, the fine com-ponent collected in the CI increased slightly whenthe flow rate was raised to 60 L/min. The authorsnoted that these data contrast with the outcomeof an earlier in vivo study by one of the co-au-thors,127 in which no lung penetration was ob-served with a similar nasal spray, so the conclu-sion was drawn that an anatomically accuratenasal cavity model would probably result in evenlower fine component with their in vitro mea-surements with these oversized entry ports.

Doub and Adams128 recently reported an ex-amination of the fine component of aqueous be-clomethasone dipropionate (BDP) nasal spraypump aerosols using an ACI also equipped witheither a 2- or 5-L spherical glass induction port.Their short stack CI, which was constructed withcomponents normally used at 90 L/min (pre-sep-arator, stages 22, 21, 5, and filter), but operatedat 28.3 L/min, was slightly more sensitive thanthe configuration used in the studies by Sumanet al.11 and Guo et al.,126 since the cut size forstage 22 is 13.6 mm, compared with 9.0 mm forstage 0 at this flow rate, so that more of the dosecould enter the CI. However, Doub and Adams128

acknowledged that the absolute mass of API peractuation entering the CI was still a very smallfraction of the dose ex actuator (0.5–1%). With theadvent of the NGI, there may be an opportunityto re-evaluate the precision of this measurement,given the good sharpness of cut for its pre-sepa-rator (cut size 5 12.7 mm at 60 L/min).57 Thiscomponent might conceivably be inverted andused as a substitute entry port. Furthermore, theNGI can be made to function effectively with areduced number of operating stages, by inserting


deep collection cups in place of the normal cupsfor those stages where particle collection is notrequired, thereby optimizing the analytical sen-sitivity.83

The CI is a more appropriate technique for size-characterizing the aerosols produced from pro-pellant-driven nasal metered-dose inhalers (so-called n-pMDIs), since these aerosols comprisefiner droplets produced as a result of the energyavailable from propellant expansion when the in-haler is actuated. However, as with all inhaler test-ing, selection of the most suitable CI configura-tion is determined by both the geometry of theinduction port (in this instance to accommodateupward actuation of the inhaler), as well as thecut sizes for the impactor stages at the chosen flowrate. Yu et al.61 first reported the use of a six-stageDelron CI equipped with 5-L glass chamber entrymade from a round-bottomed flask to simulate thenasal cavity that collected droplets with dae . 16mm from a mechanical pump nasal spray. Morerecently, Ostrander et al.129 described the use of amore realistically shaped glass induction port(without turbinate structures) that fitted above aTwin-Impinger induction port having its entry(mouth) stopped to simulate a closed mouth. Thisarrangement therefore simulated both the humannasal and oral cavities as an entry to their CI. Le-ung et al.130 questioned the value of CI-based mea-surements for this class of inhaler if the entry tothe CI only mimics a simplified geometry of thenasal passageways. Their ACI-based data, ob-tained at 28.3 L/min with glass and silicone rub-ber induction ports shaped to mimic human nasalanatomy, showed the presence of a significantnon-ballistic fraction having MMAD values of2.6–4.2 mm, depending upon the nature of the pro-pellant. However, good nasal deposition was alsoachieved in their nasal cavities, particularly theone manufactured from silicone rubber contain-ing turbinate-like structures within its confines.Given the complexity of the nasal passageways,in particular the importance of turbulence as ameans of depositing fine particles, which is notwell reproduced by conventional induction portdesign, they concluded that in vivo nasal deposi-tion from n-pMDIs is likely to be much higherthan that indicated using a smooth-surfaced glassinduction port that can only represent the overallshape of the nasal cavity. Further work is neededto resolve this issue, which will hopefully lead tothe development of appropriate induction port-CIconfigurations for use with n-pMDIs.

OTHER CONCERNS RELATINGSPECIFICALLY TO CI OPERATIONINTERNAL (INTER-STAGE) LOSSES

Internal losses of particles on CI surfaces otherthan those intended for collection (so-called wall,or inter-stage losses) are related to the issue ofparticle bounce, since particles with tacky sur-faces are more susceptible by virtue of their in-creased ability to adhere to all types of surfacesin which they come into contact.9 An upper limitof 5% of the total delivered drug mass per actu-ation from the inhaler has been defined as a CIsystem suitability requirement in the U.S. Phar-macopeia,1 to limit the impact of internal losseson measurement accuracy. However, a recom-mendation is also made that, in the event that thelosses exceed this limit, the procedure should beperformed in such a way as to include wall lossesin the assessment of API. In practice, it is not pos-sible to apportion such losses from stage to stagewithin the CI, because they lack size classifica-tion. The alternative option to utilize a differenttype of CI is therefore to be preferred. Inter-stagelosses are also particle size dependent, particu-larly when dae . 10 mm, as illustrated by cali-bration studies in which different sizes of uni-form particles were sampled by ACIs operated at28.3 L/min.48,55 It follows that the pre-separatoris the most vulnerable component of the CI tosuch losses, but little has been done to quantifythem for the commonly encountered CI designs.This is probably because for most inhaler testing,the pre-separator is not considered as an addi-tional impaction stage, but as a pre-classifier ofparticles whose inertia is greater than that asso-ciated with the cut size of the first impactionstage, and therefore unlikely to reach the lung.Under this assumption, losses are only importantif they are related to API that cannot be recov-ered or are associated with fine particles thatshould have entered the impactor. In this context,Sethuraman and Hickey107 identified low-veloc-ity regions above and after exiting the exit ori-fices in an ACI pre-separator, in which increasedinternal losses would be expected to occur awayfrom the intended collection area on the floor be-low the entrance orifice, and which might there-fore not be recovered. Unfortunately, the magni-tude of these losses could not be quantified bytheir experimental technique that involved poly-disperse DPI-generated aerosols, and that wasmore focused on the use of a coating to eliminate


particle deaggregation. Data from an investiga-tion using monodisperse particles of appropriatesizes would therefore be valuable in determiningthe magnitude of this non-ideal collection be-havior as a function of particle size as well as flowrate. However, on a cautionary note, the surfaceproperties of the calibration particles will affecttheir adhesion, so it may be impossible to sepa-rate unambiguously true internal losses due toflow irregularities from those caused by particlebounce and relocation within the pre-separator.

Although inter-stage deposition with CIs has adirect impact on the material recovery of API,data quantifying such losses are very limited inconnection with inhaler testing. In an attempt todevelop an empirical relationship linking suchlosses with particle size for one particular CI,Marple et al.21 expressed ACI-based internal lossdata obtained at 28.3 L/min,48 in terms of an em-pirical relationship:

L(dp) 5 0.06(exp{0.1dp}) (10)

in which L represents the sum of all inter-stagelosses within the CI. These losses exceeded 10%of the total mass entering the impactor for parti-cles larger than 5 mm in aerodynamic diameter.More detailed analysis of their distribution as afunction of particle size demonstrated that theyclustered about the stage at which the particlesshould have been collected.48 Internal losseswithin a 150-series MMI operated at either 30 or60 L/min were lower than those determined forthe ACI, being ,5% for particles with dae in therange of 1–19 mm, and typically in the range of1–2%, based on data obtained with liquid droplet(fluorescent-tagged oleic acid) aerosols.52 Thelower losses for the MMI were attributed as muchto the choice of calibration aerosol (increased re-tention of solid, but moist and therefore tackymethylene blue particles on internal surfaces inthe ACI calibration), as to impactor design. How-ever, if this is the case, it follows that the magni-tude of internal losses within the CI, like those referred to for the pre-separator, will be formu-lation/inhaler dependent, and this assertion issupported by the outcome of a comparative as-sessment of four types of pMDI- and DPI-gener-ated aerosols by Olsson et al.43 They reported thattotal inter-stage losses in an ACI were 1.6% of theinhaler-delivered dose for a budesonide pMDI,2.4% for an albuterol DPI, 3.4% for a budesonideDPI and 3.5% for an ipratropium bromide DPI.

Equivalent losses within a MSLI were lower(,1.2%), but followed the same pattern. They alsofound that inter-stage losses in a model 160 MMIwere similar to those in the ACI, with the excep-tion of the albuterol DPI, where losses were sig-nificantly higher, at 9.6%. These anomalouslyhigh losses were attributed to overloading of thegrease-coated surfaces of the upper stages in theMMI with excessive amounts of carrier lactose,resulting in over-sized particles cascading furtherdown the CI, implying the need for a high-ca-pacity pre-separator with this CI for this type offormulation. Apart from the isolated measure-ment just considered, these internal losses wereall comfortably below the U.S. Pharmacopeia sys-tem suitability limit, so that variations betweenformulations/inhaler types were small in termsof the magnitude of the label claim dose. How-ever, the authors cautioned that a detailed reviewof the underlying stage data revealed large rela-tive differences between the distributions oflosses between the combinations that were tested,counseling the need for validation with respect toeach application (product-CI methodology).

Internal loss data for production NGIs have yetto be published, but measurements made withprototype instruments indicate that they amountto ,5% of the inhaler-delivered dose from bothpMDIs131 and DPIs,132 provided that the pre-sep-arator is used with certain high unit dose pMDI-delivered formulations, and is appropriatelycoated with an agent to eliminate particle bounce.

OPERATOR CONTROLLABLEVARIABLES AFFECTING

PSD ACCURACY

It is generally agreed that the CI measurementprocess is complex and labor intensive, but at thepresent time it provides the only way to deter-mine inhaler particle size distributions by quan-tifying the mass of API separately from othercomponents in the formulation.

The PQRI working group developing method-ology for identifying the causes of material bal-ance failures have addressed in some detail thenumerous operator-controllable variables that in-fluence both the accuracy of this parameter aswell as that of the size distribution measure-ment.39 Although issues such as choice of solventand technique for API recovery are product de-pendent and therefore outside the scope of this


review, the choice whether or not to use a col-lection surface coating, as well as air leakage andflow rate control, are technique dependent, andare therefore discussed further.

COLLECTION SURFACE PREPARATION

Coating of CI stage surfaces with an agent toimprove particle adhesion is well understood tobe necessary to prevent bounce and re-entrain-ment of solid particles,26,49,50,133 with the excep-tion of the MSLI, where the liquid contained ineach impaction stage fulfils this purpose, therebyavoiding the need for further treatment, otherthan to rinse all interior surfaces, including thewalls and ceiling of each stage, diligently to re-cover the API quantitatively.60 Particle bounce bi-ases the measured size distribution data to finersizes, and in the case of DPI-testing, some formof coating with a viscous, compliant agent, suchas silicone oil as one of many examples, is almostessential.94,134 This precaution is necessary be-cause of the high probability of elastic impact be-tween the solid particles and the otherwise stiffcollection surface, associated with a large coeffi-cient of restitution.135 Coating of the interior sur-faces of the ACI pre-separator has also been rec-ommended as a means of reducing the break-upof large aggregate particles at impact from someDPI-based formulations. The pre-separator,therefore, collects aggregates that would not nor-mally fragment upon inhalation, appropriately,and bias towards finer particle sizes in the sizedistribution measurements is avoided.107

Many pMDI-based formulations contain sur-factant, which is compliant, and therefore morelikely to result in inelastic collisions of drug-con-taining particles on impact with CI collection sur-faces, thereby increasing the likelihood of effec-tive adhesion. However, there is evidence thatstage coating may be necessary, particularly forsingle- or two-actuation measurements of pMDIperformance simulating guidance given in thepatient insert,136 despite the fact that these in-halers are generally tested at lower flow ratesthan DPIs. For instance, Nasr et al.137 observed aso-called “loading effect,” whereby with un-coated surfaces in either an ACI or 150-seriesMMI operated at 28.3 and 30 L/min, respectively,one or four doses of pMDI-delivered albuterolparticles were subject to significant bounce,whereas this phenomenon was not observed (to

a significant extent) when 10 actuations were de-livered. By contrast, good particle adhesion wasobserved with either impactor when the stage col-lection surfaces were coated with glycerin or sil-icone fluid. Nasr and Allgire138 and El-Araud etal.139 have also reported similar findings. Thesefindings support the advice to treat coating ofstages as an issue to be considered as part of themethod validation process on a formulation/in-haler/CI basis.

LEAKAGE AND FLOW RATE CONTROL

The elimination of leakage and achievement ofstable flow rate control for CI-based measure-ments are both factors that influence measure-ment accuracy, but the identification and removalof leaks has received scant attention in the liter-ature. Since in almost all configurations, the air-flow through the CI system carries the entire sam-ple of particulate emitted from the inhaler, thisstream must be directed in a controlled pathway.Leaks within the CI, for example caused by de-fective inter-stage seals in the case of the ACI, willprovide an additional route for the in-flow of am-bient air. Such leaks, including those at the con-nections between the inhaler and induction portas well as those associated with a pre-separator(if used) are not easily detected, and thereforemust be guarded against by appropriate cleaningand maintenance.140 Apart from the possibility ofincreasing local inter-stage deposition by deflect-ing the airflow from its proper pathway, leakswill reduce the upstream airflow, if flow rate ismonitored downstream of the CI, decreasing flowvelocity at the nozzles of the affected stages. Con-versely, if flow rate is set correctly at the entry tothe CI, leakages will increase the actual flowthrough stages that are downstream. Either situ-ation will result in unpredictable changes to thecut sizes of the affected stages. Although leakagesmay have an ill-defined but possibly significantimpact on the accuracy of the particle size distri-bution measurements, the material balance is un-likely to be affected, since there is no opportunityfor material to escape from the system.39 Since thepressure differential within the CI and the sur-rounding ambient atmosphere is greatest at thestages nearest to the vacuum source, it followsthat air leakage due to sealing defects are mostlikely to occur at these locations. At least onemanufacturer has therefore developed a test fa-


cility in order that a user can verify that each stageof their CI is leak-tight before use.83

The size-discriminating ability of CIs is depen-dent on flow and hence particle velocity throughthe nozzles of each stage (equation 3). It thereforefollows that the volumetric flow rate at the CI in-let is the parameter that must be established cor-rectly.5 For testing pMDIs and nebulizers, thisprocess can be relatively easily accomplished bythe use of a soap-film flow-meter that has negli-gible resistance to flow, located upstream of theCI,140 since there is no requirement to establish apressure drop at the inhaler interface, as is thecase with DPI testing.1,2 A calibrated dry gas me-ter is likely to be more suitable for flow rates inexcess of 30 L/min.141 However, if other types offlow-meter are used, the measured flow rateshould be corrected to the conditions at the in-duction port entry.39

From the analysis of CI operating principles, itis evident that stage cut sizes will only be fixedif the volumetric flow rate through the CI is keptconstant during the entire aerosol sampling pro-cedure. This criterion is readily achieved whenevaluating either pMDIs or nebulizers, as there isno need to combine the process of sampling bythe CI with operation of these classes of inhaler.For pMDI testing, particularly when evaluatingthe impact of delayed inhalation through an HC,the practice of actuating the inhaler followed byapplying the vacuum to permit the CI to sam-ple142 should be avoided, as the significant inter-nal airspace (dead volume) within the CI system,including induction port, pre-separator (if used),and pipe-work delays the rise in flow rate to thefinal value. During this period, which can be ofthe order of 1 sec, and which takes place whenmost of the aerosol has yet to be sampled, thestage cut-points are undefined, ultimately lead-ing to bias that is difficult to quantify in the mea-sured size distribution data.42

Flow rate setting to evaluate DPI performancecan be difficult to accomplish when undertakenfollowing compendial procedures, since it is nec-essary first to attach the DPI to the induction portof the CI, and adjust the flow rate until the pres-sure drop across the inhaler (measured at the in-duction port entry) is 4 kPa, chosen as beingbroadly representative of pressure drops pro-duced by patients using DPIs.143 The DPI is thenreplaced by a flow-meter that is capable of pro-viding the volumetric flow rate either directly orthrough an appropriate pressure correction. Pro-

viding that critical (i.e., sonic) flow is maintainedat the regulating valve located between the CIand vacuum source, the assumption is made thatexchanging the DPI for the flow-meter will notaffect the flow rate at the CI inlet. Cox et al.144

identified two concerns with this methodology.Firstly, all currently available flow measuring de-vices introduce some disturbance into the flowbeing measured, making it difficult to select asuitable reference location to define exit flow rate(equivalent to CI entry flow rate). Although thecritical orifice prevents pressure pulsations fromthe vacuum source being propagated upstreamto the CI and inhaler, the mass flow rate throughthe system is affected by changes to air pressureand density upstream of the orifice. The authorsconceded that the effect of such changes on CI in-let volumetric flow rate is negligible if:

(1) The pressure drop across the critical orifice ismuch more than twice the upstream pressuredrop across inhaler and CI system.

(2) The pressure drop across both DPI and flow-meter are small compared with that across theCI (both likely conditions for most low andmedium resistance DPI testing).

They concluded by proposing that the com-pendial method be followed as far as setting thepressure drop at the induction port to 4 kPa withthe inhaler attached to the induction port and es-tablishing critical flow at the regulation valvedownstream of the CI. The procedure was thenmodified by replacing the DPI with a flow-meterand second regulation valve, both located up-stream of the induction port. This additional reg-ulating valve was adjusted to restore the pressuredrop at the impactor inlet to 4 kPa, thereby giv-ing the same inlet flow rate to the CI as with theinhaler in place.

Olsson and Asking145 more recently showed,by inserting variable flow resistances in place ofan inhaler, that the compendial method which re-lies on achieving critical flow downstream of theCI to make the volumetric flow rate at its entryinsensitive to changes in pressure drop at the in-let, is valid to a pressure drop of at least 12 kPa.This limit encompasses the range likely to be ex-perienced in the testing of the highest resistanceDPIs.146 Olsson and Asking145 further demon-strated that under these conditions, the volumet-ric flow rate downstream of the variable resis-tance is constant (varying mass flow rate), when


critical flow is maintained at the regulating valve.They therefore rejected the claim by Cox et al.144

that the pharmacopeial method is flawed, on thegrounds that they drew their conclusion withoutmeasuring the flow rate with the DPI in place, as-suming that the flow rate entering the system isindependent of its resistance (constant massflow). However, Olsson and Asking145 also notedthat, if the compendial method is used un-amended, a practical problem arises, since manyflow-meters are calibrated for the entering andnot the exiting flow that is required to gauge in-let flow rate to the CI correctly. To resolve thisconcern, they proposed that the relevant flow rate(Qout) simply be calculated from the indicatedflow rate (Qin) and the pressure drop across theflow-meter (DPflow-meter), applying Boyle’s law

correction for isothermal expansion of an idealgas:

Qout 5 3 4 (11)

where Patm is atmospheric pressure.In the case of DPI testing following compen-

dial procedures,1,2 inhaler operation and aerosolformation are intimately linked with the inhala-tion maneuver simulated by sampling throughthe CI system, and dead volume is therefore ofimportance. A large dead volume is normallyconsidered undesirable, as sampling from theDPI only takes place for the required duration towithdraw a fixed volume (4 L) of air from the in-haler mouthpiece. Flow is initiated for the re-quired time by opening a solenoid valve located

QinPatm}}}(Patm 2 DPflow-meter)


FIG. 8. Two methods developed by de Boer et al.147 to vary flow rate rise time when testing DPIs. (Reprinted fromInt. J. Pharm. 153, de Boer, A.H., G.K. Bolhuis, D. Gjaltema, and P. Hagerdoorn. Inhalation characteristics and theireffects on in vitro drug delivery from dry powder inhalers: Part 3—The effect of flow resistance increase rate (FIR) onthe in vitro drug release from the Pulmicort 200 Turbuhaler, 67–77, Copyright 1997, with permission from Elsevier.)

downstream of the CI. The dead volumes associ-ated with the ACI, MSLI and NGI with inductionport (contributing 85 mL) amount to approxi-mately 450 mL, 1.2 L, and 1.0 L, respectively. If apre-separator is used, the additional dead volumeof the ACI and NGI increases to approximately150 mL and 1.0 L, respectively (the pre-separatorfor the MSLI is stage 1 of this CI).

De Boer et al.147 exploited CI system dead vol-ume to study flow rate rise behavior when test-ing DPIs. They significantly increased the deadvolume associated with their MSLI-based systemby inserting a 1-L vessel between the MSLI andregulating valve (Fig. 8a). The added volume re-sulted in a reduced flow rate increase (FIR) up topeak (constant) flow rate when the inhaler wasattached and flow initiated through the system.A capillary inserted in the exit tube from the ves-sel was used to fine tune the flow rate-time pro-file. Increased FIRs were achieved by creating anunder-pressure in the CI, operated without in-creased dead volume, before connecting and ac-tuating the inhaler (Fig. 8b). The underlying pur-pose of these modifications was the desire tomore closely understand how a high-resistanceDPI (Turbuhaler) performs in clinical use, givenprevious experience with a small group ofhealthy volunteers.148

Also recognizing the importance of flow raterise as a function of time when testing DPIs, Cha-van and Dalby149 simulated different flow rateramps that were linear with respect to time, rang-

ing from 100 msec to 3 sec in duration, by regu-lating the air-flow fed via the DPI using a com-puter-controlled proportionating valve. Their CIoperated at constant flow rate greater than themaximum flow rate achieved via the DPI, fol-lowing the compendial procedure, with make-upair coming from an inlet that by-passed the in-haler (Fig. 9). Using this approach, they were ableto correlate increases in FPF from a Rotahaler DPIwith decrease in ramp duration (fastest “inhala-tion”), attributing the effect to increased particlede-aggregation and/or the capture of larger ag-gregates in crevices or regions of low flow withinthe inhaler at longer ramp durations. Interest-ingly, they observed an insignificant difference inFPF when 2 L was sampled, rather than the 4-Linhaled volume recommended in the compendialmethod. Further work is needed, with DPIs hav-ing a wide range of resistances, to evaluate thistechnique more thoroughly.

The establishment of the correct CI stage cut-point sizes at flow rates other than those forwhich the impactor has been calibrated is also ofconcern to those testing DPIs by the compendialprocedure, as the resistance of the inhaler deter-mines the flow rate at the CI inlet.1,2 This is ap-propriate, given the evidence that testing at flowrates appropriate to those achieved by patients isnecessary for meaningful comparisons of the per-formance of this class of inhaler.146,150 For theideal inertial separator, it can be shown by ap-plication of equations 3 or 4, that the stage cut-


FIG. 9. Method developed by Chavan and Dalby149 for simulating different flow rate ramps for DPI testing by CI.(Reprinted from AAPS PharmSci. 4(2), Chavan, V., and R. Dalby. Novel system to investigate the effects of inhaledvolume and rates of rise in simulated inspiratory air flow on fine particle output from a dry powder inhaler, article6, Copyright 2002, used with permission by the American Association of Pharmaceutical Scientists.)

point size (d50,1) at flow rate (Q1) is related to thecut-point size (d50,ref) at a reference flow rate(Qref) where calibration data are available, in ac-cordance with:

d50,1 5 d50,ref 3 41/2

(12)

which is the basis of the calculation provided inthe current compendial methods, where d50,ref isfixed at 60 L/min.1,2 This process enables the CIto sample the entire aerosol from the DPI, avoid-ing concerns about size-related bias, if an alter-native such as isokinetic sampling of the aerosolby the CI was to be chosen,93 so that the impactorcan be operated only at its design/calibrationflow rate. However, equation 12 is only strictlyvalid when inertial forces dominate the particleseparation process so that St is constant.13–15 Thisis not the case when gravity has a significant ef-fect on the size-separation process, as occurs withcomponents of CIs in which particles with dae .10 mm are being size-separated.151 Under thesecircumstances, it may be more appropriate to fitcalibration data, if acquired at several flow rateswithin the range of operation of the CI, by apower law expression of the form:

d50,1 5 Y3 4x

(13)Qref}Q1

Qref}Q1

where Qref is a chosen reference condition (gen-erally 60 L/min for DPI testing), and the para-meters Y and x are chosen to fit the calibrationdata. This more general approach was thereforeadopted to predict stage cut-point sizes for theNGI.57 Although not based on a model of the un-derlying physics of impactor operation, the tech-nique is practical, and can also be applied to cor-rect for other non-ideal behavior, such as theeffect of the slip correction term describing par-ticle motion, which can be significant with stagesthat separate particles finer than 0.5 mm.

Although most investigators utilize the com-pendial method for evaluating DPIs, the study byWeuthen et al.152 is of significance, in that it em-ployed an alternative strategy making it possibleto characterize inhaler performance at a range ofdifferent flow rates from 28.3 to 80 L/min, whilstoperating an ACI at a fixed flow rate of 28.3L/min. In their arrangement, the excess flow wasdiverted away from the CI via a “Y”-connectorlocated between the inhaler and induction port,passing through a by-pass line where the volu-metric flow rate was measured, to re-join the flowexiting the CI (Fig. 10). A three-way valve locatedbetween the CI and vacuum source was used todivert flow from the system so that the inhalercould be attached before making each measure-ment. Once the valve was switched so that vac-


FIG. 10. CI sampling arrangement for DPI testing developed by Weuthen et al.152 (Reprinted from J. Aerosol Med.15(3), Weuthen, T., S. Roeder, P. Brand, B. Müllinger, and G. Scheuch. In vitro testing of two formoterol dry powderinhalers at different flow rates, 297–303, Copyright 2002, used with permission of Mary Ann Liebert Inc.)

uum was applied to the DPI, the entire flow actedon the device within 0.25 sec. In addition to de-termining the particle size distribution with theCI, the total mass of API emitted from the inhalercould also be quantified by diverting the flow ratedownstream of the bypass line to an absolute fil-ter. Although simple in concept, this system ap-pears to lack the advantages of flow stability im-posed by the critical orifice, and the particlesize-dependent sampling characteristics of the“Y”-piece interposed between the inhaler and in-duction port have not been established. Justify-ing the latter limitation, Weuthen et al.152 notedthat particles that impacted in either the “Y”-piece or induction port are not supposed to reachthe lungs, basing the value of their technique onits ability to quantify FPD , 5.8 mm in aerody-namic diameter accurately.

CI CALIBRATION OR STAGE MENSURATION?

In common with other measurement tools con-cerned with inhaler performance assessments, CIqualification involves a means of providing as-surance that performance is to the design speci-fication upon receipt and that it has not degradedwith use. Such checks are usually provided bycalibration using traceable methods. In the caseof a CI, this procedure involves determining eachstage collection efficiency-particle size curve toestablish its cut size, and this process is most ac-curately accomplished using monodisperse par-ticles having a range of different sizes in the ap-propriate range.153 Although unambiguous,calibration with challenge particles of this natureis exceedingly time-consuming, and requiresmeticulous attention to the methodology to be ac-curate. It is also dependent upon the accuracy ofthe calibration particle detecting equipment. In-formation about internal losses cannot be ob-tained by methods involving particle countingupstream and downstream of the stage being cal-ibrated, but requires the generation of aerosolscomprised of particles containing a chemicalmarker that can be quantified by appropriate an-alytical methods.21

For a well-designed CI, the only structure sig-nificantly influencing cut size is nozzle diameter,or mean nozzle diameter for a multi-jet stage(equation 4). The U.S. Pharmacopeia,1 as an al-ternative to calibration, therefore specifies as part

of system suitability testing for CIs, that the noz-zle diameters be measured (stage mensuration).However, no advice is given on the frequency forundertaking this process, other than it should beperformed on a regular basis, presumably be-cause the requirement will be formulation andAPI-recovery method dependent. Nichols,154 inan assessment of standard and modified ACIs foruse at high flow rates, recommended that a stan-dardized methodology for stage mensuration bedeveloped for CIs, in association with nozzlespecifications that include both mean nozzle di-ameter and a limit for the range of variation (63SDs). This approach was taken with the manu-facturing specification for the NGI,16 in which anarrow range is specified for nozzle diameters,related to machining tolerances that are specificto each stage.155 Automated optical inspectionwas proposed for stages whose nozzles are finerthan 2.5 mm with the use of “go/no-go” pin-gauges as a more precise technique to inspectstages with larger nozzles. These nozzle diame-ter specifications are associated with the calibra-tion of a representative “archival” impactor,whose range of nozzle diameters for each stagewas purposely chosen to be close to the mid-pointof the specification for manufacture.57 In princi-ple, the user need only confirm after stage men-suration on a regular basis, that the nozzle sizesare still within specification to have confidence inthe CI performance, continuing to use the genericarchival calibration to define the stage cut sizes.However, recent experience with aluminum ACIshas shown that care is required with the use of“go/no go” gauges to check stages made fromsofter metal or where corrosion may have oc-curred.156 In addition to taking the precaution tolimit corrosion by rapid removal of liquid waterafter cleaning, optical inspection methods may bepreferred for these materials,157 since these ex-amination methods are non-invasive.

Although the approach developed for the NGIcould in theory be applied to other CIs, until re-cently there was almost no information availablefrom CI manufacturers to enable standards to beset for stage mensuration. In consequence, usersset their own specifications without knowingwhat effect (if any) deviations outside of theselimits might have on stage d50 values, often rely-ing on generic calibration data, such as that sup-plied by the manufacturer of the ACI. In 1997,Stein and Olson158 highlighted that the nozzle di-ameters on several stages of 14 aluminum ACIs


did not correspond to the manufacturer’s speci-fications. As a result, cut sizes for individual‘identical’ stages were calculated to vary by asmuch as 0.45 mm. Irregular shaped nozzlesformed during manufacture were also identifiedby microscopy. They then went on to calculatethe effect of such variations on those stages wherethe bulk of the particles would have collected, as-suming a model aerosol having MMAD and GSDvalues of 2.40 mm and 1.70, respectively, wassampled at 28.3 L/min. The mass collected as apercentage of the total sampled mass was esti-mated to vary from 5.3% to 22.8% for stage 3,25.3% to 39.2% for stage 4, 29.0% to 38.9% forstage 5, and 5.0% to 10.8% for stage 6. These aresubstantial variations in the context of inhalertesting, indicating that a more rigorous approachwas required for both CI manufacturing specifi-cations and in methods for validating individualCIs. More recently, Stein159 observed from test-ing undertaken with three different pMDI-pro-duced formulations, that even with ACIs whosenozzles were within the manufacturer’s specifi-cation, large differences in mass of API collectedcould occur on a stage-by-stage basis. For in-stance, the mass of API from a solution-basedpMDI formulation that collected on stage 6 ofACIs operated at 28.3 L/min ranged from 28.7%to 41.3% of the mass entering the impactor. Thisfinding indicates that it may be impractical to an-alyze data from this particular CI on an individ-ual stage basis, even when stage mensuration hasconfirmed that the nozzle dimensions are withinspecification. In response to these concerns, ver-sions of the ACI manufactured in stainless steelwith and without gold-plating have been manu-factured in an attempt to improve nozzle manu-facturing tolerances as well as to enhance resis-tance to changes in use caused by wear and/orcorrosion. CI manufacturers have also becomeaware of the need to publish nozzle diameterspecifications and some now offer stage mensura-tion services. Such improvements appear to benecessary, to judge from a recent comparison ofaluminum and stainless steel ACIs from two dif-ferent manufacturers reported by Shelton et al.,160

in which a solution-formulated pMDI was chosenin order to minimize inhaler-based variability. Mi-croscopically visible imperfections in the alu-minum impactor nozzles, including non-sphericaloutline, partial obstruction with debris, ridging ofmetal at the orifice exit due to the manufacturingprocess, were associated with small, but signifi-

cantly greater inter-impactor variability in size dis-tribution measurements than equivalent data ob-tained with CIs manufactured from stainless steel.

CONCLUSION

The CI technique is challenging and requiresmeticulous attention to detail to make accurateand precise measurements consistently, irrespec-tive of inhaler type. This review has highlightedseveral aspects of CI design and operation thathave an influence on the accuracy of these mea-surements, but the user has to decide which ofthese issues should be addressed, based on theapplication being undertaken. There are two dis-tinct uses for CI-based measurements: (1) the as-sessment of inhaler quality and (2) the estimationof deposition behavior of the resulting aerosol inthe respiratory tract. It would be both an addi-tional burden and divergent from current com-pendial and regulatory practices to interface a CIwith a breathing simulator for the purpose of in-haler batch release testing. Although the ideal re-sult is to establish accurately an absolute measureof the size distribution, a more critical goal inproduct performance testing is to optimize andmaintain consistent measurement capability frombatch to batch. Under these circumstances, it isdesirable to keep the measurement system as sim-ple as possible. However, the use of the CI in con-junction with a simulator capable of at least re-producing an inhalation maneuver, or better theentire breathing cycle, would be a more appro-priate approach to take, if the measurements werebeing undertaken to understand the likely desti-nation of the inhaler-produced particles in therespiratory tract. A further refinement would bethe use of an anatomically correct induction port.Whatever the purpose of the measurement, closeattention to details that directly affect CI perfor-mance—such as collection substrate type andcoating, leakage, and flow rate control—shouldalways be a priority to maximize the value of themeasurements obtained using this technique.

NOMENCLATURE

Cp 5 Cunningham slip correction of particle dpCae 5 Cunningham slip correction of particle daeC50 5 Cunningham slip correction factor of a par-

ticle of size d50


d16 5 size corresponding to the 16th percentile ofthe particle size distribution or 16th percentileof the stage collection efficiency curve

d84 5 size corresponding to the 84th percentile ofthe particle size distribution or 84th percentileof the stage collection efficiency curve

d50 5 stage cut sizedac 5 aerodynamic particle diameterdp 5 particle physical diameterDc 5 diameter of the overall cluster of nozzles in

a multi-nozzle impactor stageE 5 collection efficiency of an impactor stageE50 5 collection efficiency corresponding to the

cut size of an impactor stageFPD 5 fine particle dose per actuation from the

inhalerFPF 5 fine particle fractionGSD 5 geometric standard deviation of the par-

ticle size distributionGSDstage/pre-sep 5 geometric standard deviation of

the stage or pre-separator collection efficiencycurve (sharpness-of-cut)

n 5 number of nozzles per impactor stagePatm 5 atmospheric pressureDP 5 pressure drop across a flow-meterQ 5 total volumetric flow rate through an im-

pactor stageQ1 5 flow rate for an impactor measurementQin 5 volumetric flow rate at the entry to a flow-

meterQout 5 volumetric flow rate at the flow-meter exitQref 5 reference flow rateRef 5 flow Reynolds numberS 5 nozzle to collection surface distanceSt 5 Stokes numberSt50 5 Stokes number corresponding to d50

T 5 nozzle throat lengthTED 5 total emitted dose per actuation from the

inhalerU 5 average air velocity at nozzle exit of im-

paction stageW 5 nozzle diameterx 5 adjustable power term to fit impactor cali-

bration dataXc 5 cross-flow parameterY 5 adjustable coefficient to fit impactor calibra-

tion datam 5 air viscosityra 5 air densityrp 5 particle densityr0 5 unit density (i.e., 1 g/cm3)x 5 dynamic shape factor (1.00 for a spherical par-

ticle)

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Received on January 31, 2003; in final form, March 27, 2003

Reviewed by: Leon Gradon, Ph.D.

Address reprint requests to:Jolyon P. Mitchell, Ph.D.

Trudell Medical International725 Third St.

London, Ontario, Canada N5V 5G4

E-mail: [email protected]


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Cascade Impactors for the Size Characterization of Aerosols From Medical Inhalers- Their Uses and...

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