31 Pulmonary drug delivery Kevin Taylor CHAPTER CONTENTS Inhaled drug delivery 473 Lung anatomy 473 Inhalation aerosols and the importance of size distribution 474 The influence of environmental humidity on particle size 474 Particle deposition in the airways 475 Gravitational sedimentation 475 Inertial impaction 475 Brownian diffusion 475 Other methods of deposition 475 Breathing patterns 476 Clearance of inhaled particles and drug absorption 476 Formulating and delivering therapeutic inhalation aerosols 476 Metered-dose inhalers 476 Containers 476 Propellents 476 Metering valve 478 Formulating metered-dose inhalers 478 Filling metered-dose inhaler canisters 479 Advantages and disadvantages of metered-dose inhalers 479 Spacers and breath-actuated metered-dose inhalers 479 Dry powder inhalers 480 Formulating dry powder inhalers 480 Unit-dose devices with drug in hard gelatin capsules 480 Multidose devices with drug in foil blisters 481 Multidose devices with drug preloaded in inhaler 481 Non-breath actuated devices 482 Nebulizers 482 Jet nebulizers 483 Ultrasonic nebulizers 483 Formulating nebulizer fluids 484 Physicochemicat properties of nebulizer fluids 484 Temperature effects during nebulization 484 Duration of nebulization and 'dead volume' 484 Variability between nebulizers 485 Methods of aerosol size analysis 485 Cascade impactors and impingers 485 References 487 Bibliography 487 INHALED DRUGDELIVERY Drugs are generally delivered to the respiratory tract for the treatment or prophylaxis of airways diseases, such as bronchial asthma and cystic fibrosis. The administration of a drug at its site of action can result in a rapid onset of activity, which may be highly desir- able, for instance when delivering bronchodilating drugs for the treatment of asthma. Additionally, smaller doses can be administered locally compared to delivery by the oral or parenteral routes, thereby reducing the potential incidence of adverse systemic effects and reducing drug costs. The pulmonary route is also useful where a drug is poorly absorbed orally, e.g. sodium cromoglycate, or where it is rapidly metabolized orally, e.g. isoprenaline. The avoidance of first-pass metabolism in the liver may also be advantageous, although the lung itself has some metabolic capability. The lung may be used as a route for delivering drugs having systemic activity, because of its large surface area, the abundance of capillaries and the thinness of the air-blood barrier. This has been exploited in the treatment of migraine with ergota- mine, and studies have demonstrated the potential for delivering proteins and peptides such as insulin and growth hormone via the airways. Lung anatomy The lung is the organ of external respiration, in which oxygen and carbon dioxide are exchanged between blood and inhaled air. The structure of the airways also prevents the entry of and promotes efficient removal of airborne foreign particles, including microorganisms. The respiratory tract can be considered as comprising conducting (central) regions (trachea, bronchi, bronchioles, terminal and respiratory bronchioles) and respiratory (peripheral) regions 473 A-PDF Split DEMO : Purchase from www.A-PDF.com to remove the watermark A-PDF Merger DEMO : Purchase from www.A-PDF.com to remove the watermark
Transcript
31Pulmonary drug delivery
Kevin Taylor
CHAPTER CONTENTS
Inhaled drug delivery 473Lung anatomy 473Inhalation aerosols and the importance of size
distribution 474The influence of environmental humidity on
particle size 474Particle deposition in the airways 475
fluids 484Temperature effects during nebulization 484Duration of nebulization and 'dead
volume' 484Variability between nebulizers 485
Methods of aerosol size analysis 485Cascade impactors and impingers 485
References 487
Bibliography 487
INHALED DRUG DELIVERY
Drugs are generally delivered to the respiratory tractfor the treatment or prophylaxis of airways diseases,such as bronchial asthma and cystic fibrosis. Theadministration of a drug at its site of action can resultin a rapid onset of activity, which may be highly desir-able, for instance when delivering bronchodilatingdrugs for the treatment of asthma. Additionally,smaller doses can be administered locally comparedto delivery by the oral or parenteral routes, therebyreducing the potential incidence of adverse systemiceffects and reducing drug costs. The pulmonary routeis also useful where a drug is poorly absorbed orally,e.g. sodium cromoglycate, or where it is rapidlymetabolized orally, e.g. isoprenaline. The avoidanceof first-pass metabolism in the liver may also beadvantageous, although the lung itself has somemetabolic capability.
The lung may be used as a route for deliveringdrugs having systemic activity, because of its largesurface area, the abundance of capillaries and thethinness of the air-blood barrier. This has beenexploited in the treatment of migraine with ergota-mine, and studies have demonstrated the potentialfor delivering proteins and peptides such as insulinand growth hormone via the airways.
Lung anatomyThe lung is the organ of external respiration, inwhich oxygen and carbon dioxide are exchangedbetween blood and inhaled air. The structure of theairways also prevents the entry of and promotesefficient removal of airborne foreign particles,including microorganisms.
The respiratory tract can be considered ascomprising conducting (central) regions (trachea,bronchi, bronchioles, terminal and respiratorybronchioles) and respiratory (peripheral) regions
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(respiratory bronchioles and alveolar regions),although there is no clear demarcation betweenthem (Fig. 31.1). The upper respiratory tract com-prises the nose, throat, pharynx and larynx; the lowertract comprises the trachea, bronchi, bronchiolesand the alveolar regions. Simplistically, the airwayscan be described by a symmetrical model in whicheach airway divides into two equivalent branches orgenerations. In fact, the trachea (generation 0)branches into two main bronchi (generation 1), ofwhich the right bronchus is wider and leaves thetrachea at a smaller angle than the left, and hence ismore likely to receive inhaled material. Furtherbranching of the airways ultimately results in termi-nal bronchioles. These divide to produce respiratorybronchioles, which connect with alveolar ductsleading to the alveolar sacs (generation 23). Thesecontain approximately 2-6 x 108 alveoli, producing asurface area of about 70-80 m2 in an adult male.
The conducting airways are lined with ciliatedepithelial cells. Insoluble particles deposited on theairways walls in this region are trapped by the mucusand swept upwards from the lungs by the beatingcilia and swallowed.
Inhalation aerosols and the importanceof size distributionTo deliver a drug into the airways it must be pre-sented as an aerosol. In pharmacy this is denned as
Fig. 31.1 Schematic representation of the human airways.(Reproduced with permission from Wilson and Washington1989.)
a two-phase system of solid particles or liquiddroplets dispersed in air or other gaseous phase,having sufficiently small size to display considerablestability as a suspension.
The deposition of a drug/aerosol in the airways isdependent on four factors: the physicochemicalproperties of the drug, the formulation, the delivery/liberating device and the patient (breathing patternsand clinical status).
The most fundamentally important physicalproperty of an aerosol for inhalation is its size. Theparticle size of an aerosol is usually standardized bycalculation of its aerodynamic diameter, da, whichis the physical diameter of a unit density spherewhich settles through air with a velocity equal to theparticle in question. Therapeutic aerosols areheterodispersed (polydispersed), and the distributionof sizes is generally represented by the geometricstandard deviation (GSD or crg), when size is log-normally distributed.
For approximately spherical particles
where dp is physical diameter, p is particle densityand p0 is unit density, i.e. 1 g/cm3.
When dp is the mass median diameter (MMD), da
is termed the mass median aerodynamic diameter(MMAD).
The influence of environmental humidity onparticle size
As a particle enters the respiratory tract, the changefrom ambient to high relative humidity (approxi-mately 99%) results in condensation of water on tothe particle surface, which continues until thevapour pressure of the water equals that of the sur-rounding atmosphere. For water-insoluble materialsthis results in a negligibly thin film of water;however, with water-soluble materials a solution isformed on the particle surface. As the vapour pres-sure of the solution is lower than that of pure solventat the same temperature, water will continue to con-dense until an equilibrium between vapour pressuresis reached, i.e. the particle will increase in size. Thefinal equilibrium diameter reached is constrained bythe Kelvin effect, i.e., the vapour pressure of adroplet solution is higher than that for a planarsurface, and is a function of the particle's originaldiameter (Pritchard, 1987). Hygroscopic growth willaffect the deposition of particles, resulting in deposi-tion higher in the respiratory tract than would havebeen predicted from measurements of their initialsize.
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Particle deposition in the airwaysThe efficacy of a clinical aerosol is dependent on itsability to penetrate the respiratory tract. To penetrateto the peripheral (respirable) regions, aerosols requirea size less than about 5 or 6 /mi, with less than 2 /mibeing preferable for alveolar deposition. Literaturevalues for 'respirable' size vary and must be consid-ered alongside the environmental changes in sizedescribed above and the heterodispersed nature ofinhalation aerosol size distributions. Larger particlesor droplets are deposited in the upper respiratorytract and are rapidly cleared from the lung by themucociliary action, with the effect that drug becomesavailable for systemic absorption and may potentiallycause adverse effects. Steroid aerosols of sufficientlylarge size may deposit in the mouth and throat, withthe potential to cause oral candidiasis. The size ofaerosolized drug may be especially important in thetreatment of certain conditions where penetration tothe peripheral airways is particularly desirable, forinstance the treatment and prophylaxis of the alveo-lar infection Pneumocystis carinii pneumonia.
There are three main mechanisms responsible forparticulate deposition in the lung: gravitational sedi-mentation, impaction and diffusion.
Gravitational sedimentation
From Stokes' law, particles settling under gravity willattain a constant terminal settling velocity, Ut:
where p is particle density, g is the gravitational con-stant, d is particle diameter and 17 is air viscosity.
Thus, gravitational sedimentation of an inhaledparticle is dependent on its size and density, inaddition to its residence time in the airways. Sedi-mentation is an important deposition mechanism forparticles in the size range 0.5-3 /mi, in the smallairways and alveoli, for particles that have escapeddeposition by impaction.
Inertial impaction
Where a bifurcation occurs in the respiratory tract,the airstream changes direction and particles withinthe airstream, having sufficiently high momentum,will impact on the airways' walls rather than followthe changing airstream. This deposition mechanismis particularly important for large particles having adiameter greater than 5 /mi, and particularly greaterthan 10 /mi, and is common in the upper airways,
being the principal mechanism for deposition in thenose, mouth, pharynx and larynx and the large con-ducting airways. With the continuous branching ofthe conducting airways, the velocity of the airstreamdecreases and impaction becomes a less importantmechanism for deposition.
The probability of impaction is proportional to:
where 6 is the change in airways direction, U isairstream velocity and r is the airway's radius.
Brownian diffusion
Collision and bombardment of small particles bymolecules in the respiratory tract produce Brownianmotion. The resultant movement of particles fromhigh to low concentrations causes them to movefrom the aerosol cloud to the airways walls. The rateof diffusion is inversely proportional to the particlesize, and thus diffusion is the predominant mecha-nism for particles smaller than 0.5 /mi.
Other methods of deposition
Although impaction, sedimentation and diffusion arethe most important mechanisms for drug depositionin the respiratory tract, other mechanisms may occur.These include interception., whereby particles havingextreme shapes, such as fibres, physically catch on tothe airways' walls as they pass through the respiratorytract, and electrostatic attraction, whereby an elec-trostatic charge on a particle induces an opposingcharge on the walls of the respiratory tract, resultingin attraction between particle and walls.
Different deposition mechanisms are important fordifferently sized particles. Those greater than 5 /miwill deposit predominantly by inertial impaction inthe upper airways. Particles sized between 1 and5 /mi deposit predominantly by gravitationalsedimentation in the lower airways, especially duringslow, deep breathing, and particles less than 1 /mideposit by Brownian diffusion in the stagnant air ofthe lower airways. Particles of approximately 0.5 /miare inefficiently deposited, being too large for effec-tive deposition by Brownian diffusion and too smallfor effective impaction or sedimentation, and they areoften immediately exhaled. This size of minimumdeposition should thus be considered during formu-lation, although for the reasons of environmentalhumidity discussed previously, the equilibrium diam-eter in the airways may be significantly larger than theoriginal particle size in the formulation.
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Patient-dependent factors, such as breathing pat-terns and lung physiology, also affect particle depo-sition. For instance, the larger the inhaled volumethe greater the peripheral distribution of particles inthe lung, whereas increasing inhalation flow rateenhances deposition in the larger airways by inertialimpaction. Breath-holding after inhalation enhancesthe deposition of particles by sedimentation and dif-fusion. Optimal aerosol deposition occurs with slow,deep inhalations to total lung capacity, followed bybreath-holding prior to exhalation. It should benoted that changes in the airways resulting fromdisease states, for instance airways' obstruction, mayaffect the deposition profile of an inhaled aerosol.
Metered-dose inhalersMetered-dose inhalers (MDIs), introduced in themid-1950s, are the most commonly used inhalationdrug delivery devices. In MDIs, drug is either dis-solved or suspended in a liquid propellant mixturetogether with other excipients, including surfactants,and presented in a pressurized canister fitted with ametering valve (Fig. 31.2). A predetermined dose isreleased as a spray on actuation of the meteringvalve. When released from the canister the formula-tion undergoes volume expansion in the passagewithin the valve and forms a mixture of gas andliquid before discharge from the orifice. The high-speed gas flow helps to break up the liquid into a finespray of droplets.
Clearance of inhaled particles and drugabsorptionParticles deposited in the ciliated conducting airwaysare cleared within 24 hours and ultimately swal-lowed. Insoluble particles penetrating to the alveolarregions, and which are not solubilized in situ, areremoved more slowly. Alveolar macrophages engulfsuch particles and may then migrate to the bottom ofthe mucociliary escalator, or alternatively particlesmay be removed via the lymphatics.
Hydrophobic compounds are usually absorbed at arate dependent on their oil/water partition coefficients,whereas hydrophilic materials are poorly absorbedthrough membrane pores at rates inversely propor-tional to molecular size. Thus, the airways' membrane,like the gastrointestinal tract, is preferably permeableto the unionized form of a drug. Some drugs, such assodium cromoglycate, are partly absorbed by a sat-urable active transport mechanism. The rate of drugabsorption, and consequently drug action, can beinfluenced by the formulation. Rapid drug action cangenerally be achieved using solutions or powders ofaqueous soluble salts, whereas slower or prolongedabsorption may be achieved using suspension formu-lations, powders of less soluble salts, or novel drugdelivery systems such as liposomes and microspheres.
FORMULATING AND DELIVERINGTHERAPEUTIC INHALATION AEROSOLS
There are currently three main types of aerosol-generating device for use in inhaled drug therapy:metered-dose inhalers, dry powder inhalers andnebulizers.
Containers
Pharmaceutical aerosols may be packaged in tin-plated steel, plastic-coated glass or aluminium con-tainers. In practice, MDIs are generally presented inaluminium canisters, produced by extrusion to giveseamless containers with a capacity of 10-30 mL.Aluminium is relatively inert and may be useduncoated where there is no chemical instabilitybetween container and contents. Alternatively, alu-minium containers with an internal coating of achemically resistant organic material, such as anepoxy resin, can be used.
Propellants
The propellants used in MDI formulations areliquefied gases, traditionally chlorofluorocarbons(CFCs) and increasingly hydrofluoroalkanes(HFAs). At room temperature and pressure these aregases, but they are readily liquefied by decreasing
Fig. 31.2 The metered-dose inhaler. (Reproduced withpermission from Moren 1981.)
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temperature or increasing pressure. The head spaceof the aerosol is filled with propellant vapour,producing the saturation vapour pressure at thattemperature. On spraying, medicament and propel-lant are expelled and the head volume increases. Tore-establish the equilibrium, more propellant evapo-rates and so a constant pressure system with consis-tent spray characteristics is produced. The CFCscurrently employed in MDI formulations aretrichlorofluoromethane (CFC-11), dichlorodifluoro-methane (CFG-12) and dichlorotetrafluoroethane(CFG-114). Formulations generally compriseblends of CFC-11 and CFC-12 or CFC-11, CFC-12, and CFC-114 (Table 31.1.), together with a sur-factant such as a sorbitan ester, oleic acid or lecithin,which acts as a suspending agent and lubricates thevalve.
CFCs and HFAs are numbered using a universalsystem. The first digit is the number of carbon atomsminus 1 (omitted if zero), the second is the numberof hydrogen atoms plus 1, and the third is thenumber of fluorine atoms. Chlorine fills any remain-ing valencies, given the total number of atomsrequired to saturate the compound. If asymmetry ispossible, this is designated by a letter. The symmet-rical isomer is assigned the number described above;of the asymmetrical isomers, that designated theletter a is the most symmetrical, b the next mostsymmetrical, and so on. The CFCs are perfectlymiscible with each other and suitable blends give auseful intermediate vapour pressure, usually about450 kPa. The vapour pressure of the mixture of pro-pellants is given by Raoult's law, i.e. the vapour pres-sure of a mixed system is equal to the sum of themole fraction of each component multiplied by itsvapour pressure:
where P is the total vapour pressure of the systemand pa and pb are the partial vapour pressures of thecomponents, a and b:
where xa and xb are the mole fractions and £° and pgare the partial vapour pressures of components a andb, respectively.
The reaction of CFCs with the ozone in the earth'sstratosphere, which absorbs ultraviolet radiation at300 nm, and their contribution to global warming isa major environmental concern. CFCs pass to thestratosphere, where in the presence of UV they liber-ate chlorine, which reacts with ozone. The depletionof stratospheric ozone results in increased exposureto the UV-B part of the UV spectrum, resulting in anumber of adverse effects, in particular an increasedincidence of skin cancer. The Montreal Protocol of1987 was a global ban on the production of the fiveworst ozone-depleting CFCs by the year 2000. Thiswas amended in 1992, so that production of CFCs indeveloped countries was phased out by 1 January1996. In the European Union, all ozone-depletingCFCs had been banned by the end of 1995.Pharmaceutical aerosols are currently exempted, butthis exemption is reviewed annually. In householdand cosmetic aerosols CFCs have been replaced byhydrocarbons such as propane and butane.Alternatively, compressed gases such as nitrogendioxide, nitrogen and carbon dioxide may be used,for instance in food products. However, compressedgases do not maintain a constant pressure within thecanister throughout its use, as the internal pressure isinversely proportionate to the head volume, and soproduct performance changes with age. For reasonsof toxicity and inflammability, hydrocarbons are notconsidered appropriate alternatives to CFCs forinhalation products, and so non-ozone depletingalternatives to CFCs are being developed.
Propellants HFA-134a (trifluoromonofluoro-ethane) and HFA-227 (heptafluoropropane) arenon-ozone depleting, non-flammable HFAs, alsocalled hydrofluorocarbons (HFCs), which have beenwidely investigated as alternatives to CFC-12(Table 31.2). However, these gases contribute toglobal warming and further replacements will nodoubt be required in the future.
HFA-134a and HFA-227 have some physicalproperties, including density, which are similar to
Table 31.1 Formulae and physicochemical properties of chlorofluorocarbons (CFCs) used in MDI formulations
Number
11
12
114
Formula
CCI3F
CCI2F2
C2CI2F4
Boiling point (°C)
23.7
-29.8
3.6
Vapour pressure (kPa at 20°C)
89 (0.89 bar)
568 (5.68 bar)
183 (1.83 bar)
Density (g/mL at 20°C)
1.49
1.33
1.47
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Table 31.2 Formulae and physicochemical properties of hydrofluoroalkanes (HFAs) used in MDI formulations
Number
134a
227
Formula
Cf2' 4(12
C3F7H
Boiling point (°C)
-26.5
-17.3
Vapour pressure (kPa at 20°C)
660 (6.6 bar)
398 (3.98 bar)
Density (g/mL at 20°C)
1.23
1.41
those of CFC-12 and, to a lesser extent, CFC-114.However, they present major formulation problems:in particular they are poor solvents for the surfac-tants commonly used in MDI formulation and noalternative to CFC-11 is currently available. Ethanolis approved for use in formulations containing HFAsto allow dissolution of surfactants, and is included inmarketed non-CFC MDI products. However,ethanol has low volatility and may consequentlyincrease the droplet size of the emitted aerosols.
Metering valve
The metering valve of an MDI permits the repro-ducible delivery of small volumes (25-100 fjJL) ofproduct. Unlike the non-metering continuous-sprayvalves of conventional pressurized aerosols, themetering valve in MDIs are used in the inverted posi-tion (Fig. 31.3). Depression of the valve stem allowsthe contents of the metering chamber to be dis-charged through the orifice in the valve stem andmade available to the patient. After actuation, themetering chamber refills with liquid from the bulkand is ready to dispense the next dose. A corollary ofthis is that the MDI needs to be primed, i.e. themetering chamber filled, prior to the first use by a
Fig. 31.3 The metering valve. (Reproduced with permissionfrom Moren 1981.)
patient. MDI valves are complex in design and mustprotect the product from the environment, while alsoprotecting against product loss during repeated use.The introduction of HFA propellants with differentsolvent properties has necessitated the developmentof new valve elastomers. The valve stem fits into theactuator, which is made of polyethylene or poly-propylene. The dimensions of the orifice in the actu-ator plays a crucial role, along with the propellantvapour pressure, in determining the shape and speedof the emitted aerosol plume.
Formulating metered-dose inhalers
Pressurized aerosols may be formulated as eithersolutions or suspensions of drug in the liquefiedpropellant. Solution preparations are two-phasesystems. However, the propellants are poor solventsfor most drugs. Cosolvents such as ethanol or iso-propanol may be used, although their low volatilityretards propellant evaporation. In practice, pressur-ized inhaler formulations have, until recently, beenalmost exclusively suspensions. These three-phasesystems are harder to formulate and all the problemsof conventional suspension formulation, such ascaking, agglomeration, particle growth etc. must beconsidered. Careful consideration must be given tothe particle size of the solid (usually micronized tobetween 2 and 5 /mi), valve clogging, moisturecontent, the solubility of active compound in propel-lant (a salt may be desirable), the relative density ofpropellant and drug, and the use of surfactants assuspending agents, e.g. lecithin, oleic acid and sorbi-tan trioleate (usually included at concentrationsbetween 0.1 and 2.0% w/w). These surfactants arevery poorly soluble (« 0.02% w/w) in HFAs, andso either ethanol must be used as a cosolvent oralternative surfactants such as fluorinated polymersmust be developed (Byron et al 1994). Recentlysolution formulations of beclomethasone dipropi-onate have been marketed. Evaporation of HFA pro-pellant following actuation of these formulationsresults in smaller particle sizes than with conven-tional suspension formulations of the same drug,with consequent changes in its pulmonary distribu-tion and bioavailability.
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Canisters are filled by liquefying the propellant atreduced temperature or elevated pressure.
In cold filling, active compound, excipients andpropellant are chilled and filled at about -30°C.Additional propellant is then added at the same tem-perature and the canister sealed with the valve. Inpressure filling, a drug/propellant (CFC-11) con-centrate is produced and filled at effectively roomtemperature and pressure (in fact, usually slightlychilled to below 20°C). The valve is crimped on tothe canister and additional propellant (e.g. CFG-12)is filled at elevated pressure through the valve, in aprocess known as gassing. Pressure filling is mostfrequently employed for inhalation aerosols.However, no ozone-sparing replacement propellanthas the properties (high boiling point: 23.7°C) ofCFC-11, which is a major problem for the pharma-ceutical industry.
Once filled, the canisters are leak tested by placingthem in a water bath at elevated temperature, usually50-60°C. Following storage to allow equilibration ofthe formulation and valve components, the contain-ers are weighed to check for further leakage, prior tospray testing and insertion into actuators.
Advantages and disadvantages of metered-doseinhalers
The major advantages of MDIs are their portability,low cost and disposability. Many doses (up to 200)are stored in the small canister and dose delivery isreproducible. The inert conditions created by thepropellant vapour, together with the hermeticallysealed container, protects drugs from oxidativedegradation and microbiological contamination.However, MDIs have disadvantages. They areinefficient at drug delivery. On actuation, the firstpropellant droplets exit at a high velocity, which mayexceed 30 m/s. Consequently, much of the drug islost through impaction of these droplets in theoropharyngeal areas. The mean emitted droplet sizetypically exceeds 40 ^t,m, and propellants may notevaporate sufficiently rapidly for their size to decreaseto that suitable for deep lung deposition. Vaporizationof the droplets is hindered by the low volatility ofCFC-11, which is present in concentrations of atleast 25% in most CFC-based formulations.Evaporation, such that the aerodynamic diameter ofthe particles is close to that of the original micronizeddrug, may not occur until 5 seconds after actuation.
An additional problem with MDIs, which isbeyond the control of the formulator and manufac-
turer, is their incorrect use by patients. Reportedproblems include:
• Failure to remove the protective cap covering themouthpiece, the inhaler being used inverted;
• Failure to shake the canister;• Failure to inhale slowly and deeply;• Inadequate breath-holding;• Poor inhalation/actuation synchronization.
Correct use by patients is vital for effective drugdeposition and action. Ideally, the MDI should beactuated during the course of slow, deep inhalation,followed by a period of breath-holding. Manypatients find this difficult, especially children and theelderly. The misuse of MDIs through poor inhala-tion/actuation coordination can be significantlyreduced with appropriate instruction and coun-selling. However, it should be noted that even usingthe correct inhalation technique only 10-20% of thestated emitted dose is delivered to the site of action.
Spacers and breath-actuated metered-doseinhalers
Some of the disadvantages of MDIs, namely inhala-tion/actuation coordination and the prematuredeposition of large propellant droplets high in theairways, can be overcome by using extension devicesor 'spacers' positioned between the MDI and thepatient (Fig. 31.4). The dose from an MDI is dis-charged directly into the reservoir prior to inhala-tion. This reduces the initial droplet velocity, permitsefficient propellant evaporation and removes theneed for actuation/inhalation coordination. Thedisadvantage of spacers is that they may be cumber-some, e.g. Fisonair (Rhone-Poulenc Rorer),Nebuhaler (AstraZeneca), and Volumatic (GlaxoSmithKline). Alternatively, extension tubes may bebuilt into the design of the MDI itself, e.g. Syncroner(Rhone-Poulenc Rorer) and Spacer Inhalers(AstraZeneca). The Autohaler (3M) is an MDI with
Fig. 31.4 The Nebuhaler spacer device, fitted with a face maskfor use by a child. Courtesy of AstraZeneca.
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an inspiratory demand valve. This breath-actuateddevice overcomes the coordination problems of aconventional MDI without adding bulk to thedevice. However, a substantial inspiratory flow rate isrequired for its operation.
Dry powder inhalersIn dry powder inhaler (DPI) systems, drug isinhaled as cloud of fine particles. The drug is eitherpreloaded in an inhalation device or filled into hardgelatin capsules or foil blister discs which areloaded into a device prior to use. DPIs have severaladvantages over MDIs. DPI formulations are pro-pellant free and do not contain any excipient, otherthan a carrier (see below) - which is almost invari-ably lactose. They are breath actuated, avoiding theproblems of inhalation/actuation coordinationencountered with MDIs, and consequently they areparticularly useful for young children. DPIs canalso deliver larger drug doses than MDIs, which arelimited by the volume of the metering valve and themaximum suspension concentration that can beemployed without causing valve clogging. However,DPIs have several disadvantages. Liberation ofpowders from the device and the deaggregation ofparticles are limited by the patient's ability toinhale, which in the case of respiratory disease maybe impaired. An increase in turbulent air flowcreated by an increase in inhaled air velocityincreases the deaggregation of the emerging parti-cles, but also increases the potential for inertialimpaction in the upper airways and throat, and so acompromise has to be found. Further, DPIs areexposed to ambient atmospheric conditions, whichmay reduce formulation stability. For instance, ele-vated humidity may cause powders to clump.Finally, DPIs are generally less efficient at drugdelivery than MDIs, such that twice the dose isusually required for delivery from a DPI than fromthe equivalent MDI (Melchor et al 1993).
Formulating dry powder inhalers
To produce particles of a suitable size (preferablyless than 5 /nm), drug powders for use in inhalationsystems are usually micronized. The high-energypowders produced have poor flow propertiesbecause of their static, cohesive and adhesive nature.The flowability of a powder is affected by physicalproperties, including particle size and shape, density,surface roughness, hardness, moisture content andbulk density.
To improve their flow properties, poorly flowingdrug particles are generally mixed with larger
'carrier' particles (usually 30-60 yum) of an inertexcipient, usually lactose. This not only improvesliberation of the drug from the inhalation device byimproving powder flow, but also improves the uni-formity of capsule or device filling. Once liberatedfrom the device, the turbulent air flow generatedwithin the inhalation device should be sufficient forthe deaggregation of the drug/carrier aggregates. Thelarger carrier particles impact in the throat, whereassmaller drug particles are carried in the inhaled airdeeper into the respiratory tract.
The success of DPI formulations depends on theadhesion of drug and carrier during mixing andfilling of devices or hard gelatin capsules, followed bythe ability of the drug to desorb from the carrierduring inhalation such that free drug is available topenetrate to the peripheral airways. Adhesion anddesorption will depend on the morphology of theparticle surfaces and surface energies, which may beinfluenced by the chemical nature of the materialsinvolved and the nature of powder processing.The performance of DPI systems is thus stronglydependent on formulation factors, and also onthe construction of the delivery device and theinhalation technique.
Unit-dose devices with drug in hard gelatincapsules
The first DPI device developed was the Spinhaler(Rhone-Poulenc Rorer) for the delivery of sodiumcromoglycate (Fig. 31.5). Each dose, contained in ahard gelatin capsule, is loaded individually into thedevice. The capsule, placed in a loose-fitting rotor, ispierced by two metal needles an either side of thecapsule. Inhaled air flow though the device causes a
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Fig. 31.5 The Spinhaler. (Modified from Bell et al 1971, withthe permission of the American Pharmaceutical Association.)
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turbovibratory air pattern as the rotor rotatesrapidly, resulting in the powder being dispersed tothe capsule walls and out through the perforationsinto the air. A minimum air flow rate of 35-40 L/minthrough the device is required to produce adequatevibrations by the rotor. The occurrence of lactoseintolerance and local irritation, coughing and bron-choconstriction caused by the inhalation of largeamounts of lactose has led to the development of anaggregated, carrier-free sodium cromoglycatecapsule formulation for use in the Spinhaler.
Another unit-dose DPI is the Rotahaler (GlaxoSmithKline), which is a simple two-piece device(Fig. 31.6). The gelatin capsule is inserted into anorifice at the rear of the device and when the twosections are rotated a fin on the inner barrel pulls thetwo halves of the capsule apart. During inhalation,the freed half of the capsule spins, dispersing itscontents, which are inhaled through the mouthpiece.The resistance to air flow is lower than that of theSpinhaler and therefore a lower inspiratory velocityis required.
Other hard gelatin capsule-based devices, workingon similar principles, are available for the deliveryof drug/carrier mixes. These include the Aerohaler(Boehringer Ingelheim) and the Cyclohaler (DuPont).
Multidose devices with drug in foil blisters
The main disadvantage of hard gelatin capsule-based devices, namely the individual loading of eachdose, was overcome with the development of theDiskhaler (Glaxo SmithKline). In this system, drugis mixed with a coarse lactose carrier and filled intoan aluminium foil blister disc which is loaded, by thepatient, into the device on a support wheel(Fig. 31.7). Each disc contains four or eight doses ofdrug and the blisters are pierced with a needle as aresult of mechanical leverage on the lid. Air flowthrough the blister causes the powder to disperse as
Fig. 31.6 The Rotahaler (Modified with permission fromKjellman 1981.)
Fig. 31.7 The Diskhaler (Reproduced with permission fromSumby etal 1993.)
the patient inhales through the mouthpiece. The foilblisters are numbered, so that the patient knows thenumber of doses remaining.
Multidose devices with drug preloaded in inhaler
The evolution of the Diskhaler led to the productionof the Accuhaler or Diskus Inhaler (GlaxoSmithKline), in which drug/carrier mix is preloadedinto the device in foil-covered blister pocketscontaining 60 doses (Fig. 31.8).The foil lid is peeledoff the drug-containing pockets as each dose isadvanced, with the blisters and lids being wound upseparately within the device, which is discarded atthe end of operation. As each dose is packaged sepa-rately and only momentarily exposed to ambientconditions prior to inhalation, the Diskhaler andAccuhaler are relatively insensitive to humidity com-pared to hard gelatin capsule-based systems.
An alternative approach is a reservoir type ofdevice, in which a dose is accurately measured anddelivered from a drug reservoir. In the ClickhalerDPI (Innovata Biomed), a drug blend is stored in areservoir. Metering cups are filled by gravity fromthis reservoir and delivered to an inhalation passage,from which it is inhaled. The device is capable ofholding up to 200 doses and incorporates a dosecounter, which informs patients when the device,which is discarded after use, is nearly empty.
The Turbohaler (AstraZeneca), has overcome theneed for both a carrier and the loading of individualdoses (Fig. 31.9). The device contains a largenumber of doses (up to 200) of undiluted, looselyaggregated micronized drug, which is stored in areservoir from which it flows on to a rotating disc in
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Fig. 31.8 The Accuhaler/Diskus Inhaler, showing (a) a schematic diagram and (b) a cross-sectional representation of the device.(Reproduced with permission from Prime et al 1996.)
Fig. 31.9 The Turbohaler. (Reproduced with permission fromWetterlin 1987.)
the dosing unit. The fine holes in the disc are filledand the excess drug is removed by scrapers. As therotating disc is turned, by moving a turning gripback and forth, one metered dose is presented to theinhalation channel, and this is inhaled by the patient,
with the turbulent air flow created within the devicebreaking up any drug aggregates. A dose indicator isincorporated. The Turbohaler requires a higherinspiratory effort than the Diskhaler, owing to itshigher internal resistance, and is more sensitive tohumidity if not closed quickly after each use.
Non-breath actuated devices
Devices are currently under development whichreduce or eliminate the reliance on the patient'sinspiratory effort to disperse the drug (Rubsamen1997). Such inspiratory effort may be affected by thepatient's age and/or clinical condition. For instance,a device that uses a battery-powered impeller todeaggregate the drug powder is being developed.The device is breath actuated, but deaggregation isindependent of the patient's inspiratory flow rate.Inhale Therapeutic Systems have produced a devicein which compressed air is used to disperse drugfrom a unit-dose package into a large holdingchamber, from which it is inhaled by the patient.
Nebulizers
Nebulizers deliver relatively large volumes of drugsolutions and suspensions and are frequently used fordrugs that cannot be conveniently formulated intoMDIs or DPIs, or where the therapeutic dose is toolarge for delivery with these alternative systems.Nebulizers also have the advantage over metered-
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dose and dry powder systems in that drug may beinhaled during normal tidal breathing through amouthpiece or face-mask, and thus they are useful forpatients such as children, the elderly and patientswith arthritis, who experience difficulties with MDIs.
There are two categories of commercially availablenebulizer: jet and ultrasonic.
Jet nebulizers
Jet nebulizers (also called air-jet or air-blast nebuliz-ers) use compressed gas (air or oxygen) from a com-pressed gas cylinder, hospital air-line or electricalcompressor to convert a liquid (usually an aqueoussolution) into a spray. The jet of high-velocity gas ispassed either tangentially or coaxially through anarrow Venturi nozzle, typically 0.3-0.7 mm indiameter. An area of negative pressure, where the airjet emerges, causes liquid to be drawn up a feed tubefrom a fluid reservoir by the Bernoulli effect(Fig. 31.10). Liquid emerges as fine filaments, whichcollapse into droplets owing to surface tension. Aproportion of the resultant (primary) aerosol
Fig. 31.10 Schematic diagram of a jet nebulizer. Compressedgas passes through a Venturi nozzle, where an area of negativepressure is created. Liquid is drawn up a feed tube and isfragmented into droplets. Large droplets impact on baffles (b),and small droplets are carried away in the inhaled airstream.(Reproduced with permission from Newman 1989.)
leaves the nebulizer directly; the remaining, large,non-respirable droplets impact on baffles or thewalls of the nebulizer chamber and are recycled intothe reservoir fluid.
Nebulizers are operated continuously, andbecause the inspiratory phase of breathing consti-tutes approximately one-third of the breathing cyclea large proportion of the emitted aerosol is notinhaled but is released into the environment. Open-vent nebulizers, incorporating inhalation and exhala-tion valves, e.g. the Pari LC nebulizer (Pari) haverecently been developed in which the patient's ownbreath boosts nebulizer performance, with aerosolproduction matching the patient's tidal volume andgreatly enhancing drug delivery. On exhalation, theaerosol being produced is generated only from thecompressor gas source, thereby minimizing drugwastage.
The rate of gas flow driving atomization is themajor determinant of the aerosol droplet size andrate of drug delivery for jet nebulizers: for instance,there may be up to a 50% reduction in the massmedian aerodynamic diameter (MMAD, see below)when the flow rate is increased from 4 to 8 L/min,with a linear increase in the proportion of dropletsless than 5 /zm (Clay et al 1983).
Ultrasonic nebulizers
In ultrasonic nebulizers the energy necessary toatomize liquids comes from a piezoelectric crystalvibrating at high frequency. At sufficiently high ultra-sonic intensities a fountain of liquid is formed in thenebulizer chamber. Large droplets are emitted fromthe apex and a 'fog' of small droplets is emittedfrom the lower part (Fig. 31.11). Some models havea fan to blow the respirable droplets out of thedevice, whereas in others the aerosol only becomesavailable to the patient during inhalation.
Fig. 31.11 Schematic diagram of an ultrasonic nebulizer.(Reproduced with permission from Atkins et al 1992.)
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Nebulizer fluids are formulated in water, occasion-ally with the addition of cosolvents such as ethanolor propylene glycol, and with the addition ofsurfactants for suspension formulations. Becausehypoosmotic and hyperosmotic solutions may causebronchoconstriction, as may high hydrogen ionconcentrations, iso-osmotic solutions of pH greaterthan 5 are usually employed (Snell 1990). Stabilizerssuch as antioxidants and preservatives may also beincluded, although these may also cause bron-chospasm, and for this reason sulphites in particularare generally avoided as antioxidants in such formu-lations. Although chemically preserved multidosepreparations are commercially available, nebulizerformulations are generally presented as sterile,isotonic unit doses (usually 1-2.5 mL) without apreservative.
Whilst most nebulizer formulations are solutions,suspensions of micronized drug are also available fordelivery from nebulizers. In general suspensions arepoorly delivered from ultrasonic nebulizers, whereaswith jet nebulizers the efficiency of drug deliveryincreases as the size of suspended drug is decreased,with little or no release of particles when they exceedthe droplet size of the nebulized aerosol.
As the formulation of fluids for delivery by nebuliz-ers is relatively simple, these devices are frequently thefirst to be employed when investigating the delivery ofnew entities to the human lung. Recently, they havebeen used for the delivery of peptides and liposomes.In general, ultrasonic nebulizers have not been suc-cessful for delivering either peptides or liposomes,because of denaturation resulting from the elevatedtemperatures produced. Consequently, ultrasonicnebulizers are expressly excluded for the delivery ofrecombinant human deoxyribonuclease in the man-agement of cystic fibrosis. Jet nebulizers have beensuccessfully used to deliver some peptide and lipo-some formulations, although the shearing forces thatoccur in the nebulizer may produce time-dependentdamage to some materials (Niven and Brain 1994).
Physicochemical properties of nebulizer fluids
The viscosity and surface tension of a liquid beingnebulized may affect the output of nebulizers, asenergy is required to overcome viscous forces and tocreate a new surface. However, the size-selectivity ofthe nebulizer design and dimensions, with more than99% of the primary aerosol mass being recycled intothe reservoir liquid, means that changes in the sizedistribution of the primary aerosol resulting fromchanges in the properties of the solution being
atomized may not always be reflected in the sizedistribution of the emitted aerosol. In general, thesize of aerosol droplets is inversely proportional toviscosity for jet nebulizers and directly proportionalto viscosity for ultrasonic nebulisers (McCallion et al1995), with more viscous solutions requiring longerto nebulize to dryness and leaving larger residualvolumes in the nebulizer following atomization.Surface tension effects are more complex, butusually a decrease in surface tension is associatedwith a reduction in mean aerosol size.
Temperature effects during nebulization
The aerosol output from a jet nebulizer comprisesdrug solution and solvent vapour, which saturatesthe outgoing air. This causes solute concentration toincrease with time and results in a rapid decrease inthe temperature of the liquid being nebulized byapproximately 10-15°C.
This temperature decrease may be importantclinically, as some asthma sufferers experience bron-choconstriction on inhalation of cold solutions.Further, the cooling effect within the reservoir fluidwill reduce drug solubility and result in increasedliquid surface tension and viscosity. Precipitation isuncommon with bronchodilators, which have highaqueous solubility, but problems may arise with lesssoluble drugs. In such instances the use of an ultra-sonic nebulizer may be appropriate, as the operationof such devices increases solution temperature byapproximately 10-15°C.
Duration of nebulization and 'dead volume'
Clinically, liquids may be nebulized for a specifiedperiod of time, or more commonly, they may benebulized to 'dryness', which may be interpreted assputtering time, which is the time when air isdrawn up the feed tube and nebulization becomeserratic, although agitation of the nebulizer permitstreatment to be continued; clinical time, which isthe time at which therapy is ceased followingsputtering; or total time, which is the time at whichthe production of aerosol ceases.
Regardless of the duration of nebulization, not allthe fluid in the nebuliser can be atomized. Someliquid, usually about 1 mL, remains as the 'dead' or'residual' volume, associated with the baffles, inter-nal structures and walls of the nebulizer. The pro-portion of drug retained as 'dead' volume is moremarked for smaller fill volumes, hence for a 2 mL fillvolume, approximately 50% of fluid will remainassociated with the nebulizer and be unavailable fordelivery to the patient. This reduces to approxi-
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mately 25% with a 4 mL fill volume, although thereis a commensurate increase in the time necessary tonebulize to dryness.
Variability between nebulizers
Many different models of nebulizer and compressorare commercially available, and the size of aerosolsproduced and the dose delivered can vary enor-mously. For instance, in a study of 18 different com-mercially available jet nebulizers, operated accordingto the manufacturers' guidelines, aerosols were pro-duced with MMADs ranging from 0.9 to 7.2 /xm(Waldrep et al 1994). Variability may not only existbetween different nebulizers but also between indi-vidual nebulizers of the same type, and repeated useof a single nebulizer may cause variability due tobaffle wear and non-uniformity of assembly.Nebulizers, unlike the DPI and MDI devices, are notmanufactured by the producers of nebulizer solu-tions and suspension. The choice of nebulizeremployed for their delivery is thus usually beyondthe influence of the pharmaceutical manufacturer.
METHODS OF AEROSOL SIZEANALYSIS
The regional distribution of aerosols in the airwayscan be measured directly using gamma scintigraphy,by radiolabelling droplets or particles, usually withthe short half-life gamma emitter technetium-99m(99mTc). However, more commonly in vitro mea-surements of aerosol size are used to predict clinicalperformance. The principal methods that have beenemployed for size characterization of aerosols aremicroscopy, laser diffraction and cascade impaction.
Optical methods of measuring the physical size ofdeposited aerosols using microscopy are laboriousand do not give an indication of their likely deposi-tion within the humid airways while being carried inan airstream. With methods of analysis based on laserFraunhofer diffraction, aerosolized droplets or parti-cles are sized as they traverse a laser beam to give avolume median diameter. Again the aerodynamicproperties of an aerosol are not being measured. Inaddition, spraying droplets into a beam exposes themto ambient conditions of temperature and humidity,which may result in solvent evaporation.
Cascade impactors and impingers
Cascade impactors comprise a series of progressivelyfiner jets and collection plates, allowing fractionationof aerosols according to their MMAD as the aerosol
is drawn through the device at a known flow rate.Traditional cascade impactors are constructed frommetal. The most commonly used comprises eightstages, with metal collection plates followed by aterminal filter. Multistage liquid impingers, workingon the same principle, are constructed from glass orglass and metal and have three, four or five stages,with wet sintered glass collection plates followed bya terminal filter. Large dense particles will deposithigher in the impactor, whereas smaller, less denseparticles will follow the air flow and only depositwhen they have been given sufficient momentum asthey are accelerated through the finer jets lower inthe impactor (Fig. 31.12). The first stage of theimpactor is usually preceded by a 90° bend of metalor glass to mimic the human throat. The cut-offdiameters for each stage at a particular air-flow ratecan be determined using monodisperse aerosols orcalculated using calibration curves. When determin-ing the size of an aerosol, cumulative percentageundersize plots of deposited aerosol on eachstage are plotted against the cut-off diameter for thatstage to allow calculation of the MMAD.
The five-stage liquid impinger (MSLI)(Fig. 31.13), with an appropriate induction port andmouthpiece adapter, is used to determine the aerody-namic size of DPIs (USP and EP), MDIs and nebu-lizers (EP). The MSLI may be operated at a flow ratebetween 30 and 100 L/min. At 60 L/min (i.e. 1 L/s)the effective cut-off diameters of stages 1, 2, 3 and 4are 13.0, 6.8, 3.1 and 1.7 //,m, respectively. The fifthstage comprises an integral filter which captures par-ticles smaller than 1.7 /xm.When testing DPIs to USPrequirements, an airflow rate (Q) calculated toproduce a pressure drop of 4.0 kPa over the inhaler isemployed. If this exceeds 100 L/min, then 100 L/minis used. The cut-off diameters of each stage at flowrate (Q) can be calculated from:
Fig. 31.12 Illustration of aerodynamic particle size separationby an impactor stage. (Reproduced with permission fromJaegfeldt et al 1987.)
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Fig. 31.13 The multistage liquid impinger. Courtesy of AstraZeneca.
where D50'Q is the cut-off diameter at the flow rate Qand n refers to the nominal cut-off valuesdetermined when Qn is 60 L/min (values givenabove).
The use of cascade impaction methods todetermine the size of aerosols has a number of dis-advantages. The high flow rates employed (typically28.3-60 L/min) result in rapid solvent evaporation,and droplets may be re-entrained in the airstreamwhereas particles may 'bounce off metal collectionplates, although this latter effect may be reduced bycoating the collection surface, for instance with asilicone fluid or glycerol. These effects can result in asignificant decrease in the measured aerosol size.Also, these measuring devices are operated at a con-stant air-flow rate. However, the dispersion of drypowder formulations and the deposition profile ofinhaled aerosols will very considerably with flowrate. To overcome the limitations of measurement ata single flow rate the Electronic Lung (TheTechnology Partnership) has been developed, whichuses a computer-controlled piston to draw airthrough the inhaler and into an impaction sizer,
following a predetermined inhalation profile(Brindley et al 1994).
Cascade impactor methods are invasive, laboriousand time-consuming, but necessary to derive infor-mation about median aerosol size and the polydisper-sity of the aerosol. To ensure that inhalation productsare likely to be clinically effective, quality control mea-surements usually involve measurement of theemitted dose and the 'fine particle fraction', (that frac-tion of the emitted dose less than a stated size, often 5or 6.4 /mi), which are combined to give a 'useful' or'respirable' dose or mass (Ganderton, 1995). Forroutine analysis, a simplified two-stage (twin)impinger is frequently employed (Fig. 31.14). Aerosolcollected in the throat and the upper stage (stage 1) isconsidered 'non-respirable', whereas that collected inthe lower stage (stage 2) is considered 'respirable'. Forthis glass device, the cut-off diameter for stage 2 is 6.4^im, i.e. aerosols collected in this stage have an aero-dynamic diameter less than 6.4 /xm and are for thismeasurement technique considered 'respirable'. Thistwin impinger is included in the BP as a method fordetermining the emitted dose from MDIs and DPIs.
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Fig. 31.14 The two-stage impinger. (Reproduced withpermission from Hallworth and Westmoreland 1987.)
REFERENCES
Atkins, PJ, Barker, N.P., Mathisen, D. (1992) The designand development of inhalation drug delivery systems.In: Pharmaceutical Inhalation Aerosol Technology.(Ed. A.S.J. Hickey) Marcel Dekker, New York,pp!55-185.
Bell, J.H., Hartley, P.S. and Cox, J.S.G. (1971) Drypowder aerosols I: a new powder inhalation device. J.Pharm Sci., 60, 1559-1564.
Brindley, A., Sumby, B.S., Smith, IJ. (1994) Thecharacterisation of inhalation devices by an inhalationsimulator: the electronic lung. J Aerosol Med., 7,197-200.
Byron, P.R., Miller, N.C., Blondino, F.E., Visich, J.E.,Ward,G.H. (1994). Some aspects of alternative propellantsolvency. Respir. Drug Del. IV, 231-242.
Clay, M.M., Pavia, D., Newman, S.P., Lennard-Jones,T,Clarke, S.W. (1983) Assessment of jet nebulisers for lungaerosol therapy. Lancet, ii, 592-594.
Ganderton, D. (1995) In-vitro testing of inhalation products.Drug Delivery to the Lungs VI. The Aerosol Society, Bristol,pp 85-88.
Hallworth, G.W., Westmoreland, D.G. (1987) The twinimpinger: a simple device for assessing the delivery ofdrugs from metered dose pressurised aerosol inhalers.J. Pharm. Pharmacol, 39, 966-972.
Jaegfeldt, H., Andersson, J.A.R.,Trofast, E., Wetterlin,K.I.L. (1987) Particle size distribution from differentmodifications of Turbohaler. In: A New Concept InInhalation Therapy. (Eds. S.P. Newman, F. Moren andG.K. Crompton). Medicom Europe, Bussum, pp 90-99.
Kjellman, N-I.M. (1981) Letter to the Editor Allergy,36, 437-438.
McCallion, O.N.M.,Taylor, K.M.G., Thomas, M., TaylorAJ. (1995) Nebulization of fluids of differentphysicochemical properties with air-jet and ultrasonicnebulizers. Pharm. Res., 12, 1682-1688.
Melchor, R., Biddiscombe, M.F., Mak, V.H.F., Short, M.D.,Spiro, S.G. (1993) Lung deposition patterns of directlylabelled salbutamol in normal subjects and in patients withreversible airflow obstruction. Thorax, 48, 506—511.
Moren, F. (1981) Pressurised aerosols for oral inhalation. Int.J. Pharm., 8, 1-10.
Niven, R.W. and Brain, J.D. (1994) Some functional aspectsof air-jet nebulizers. Int. J. Pharm., 104, 73-85.
Prime, D., Slater, A.L., Haywood, P.A., Smith, IJ. (1996)Assessing dose delivery from the Flixotide Diskus Inhaler -a multidose powder inhaler. Pharm. Tech. Europe, 8 (3),23-34.
Pritchard, J.N. (1987) Particle growth in the airways and theinfluence of airflow. In: A New Concept In InhalationTherapy. (Eds. S.P. Newman, F. Moren and G.K.Crompton). Medicom Europe, Bussum, pp 3-24.
Rubsamen, R. (1997) Novel aerosol peptide drug deliverysystems. In: Inhalation Delivery of Therapeutic Peptides andProteins. (Eds. A.L. Adjei and P.K. Gupta). Marcel Dekker,New York, pp 703-731.
Sumby, B.S., Churcher, K.M., Smith, I J., Grant, A.C.,Truman, K.G., Marriott, R.J., Booth, S J. (1993) Dosereliability of the Serevent Diskhaler system. Pharm. Tech.Int., 6 (5), 20-27.
Waldrep, J.C., Keyhani, K., Black, M., Knight, V. (1994)Operating characteristics of 18 different continuous flowjet nebulizers with beclomethasone dipropionate liposomeaerosol. Chest, 105, 106-110.
Wetterlin, K.I.L. (1987) Design and function of Turbohaler.In: A New Concept In Inhalation Therapy. (Eds. S.P.Newman, F. Moren and G.K. Crompton). MedicomEurope, Bussum, pp 85-89.
Wilson, C.G. and Washington N. (1989) PhysiologicalPharmaceutics: Biological Barriers to Drug Absorption. EllisHorwood, Chichester, pi57.
BIBLIOGRAPHY
Adjei, A.L. and Gupta, P.K. (eds) (1997) Inhalation Deliveryof Therapeutic Peptides and Proteins. Marcel Dekker,New York.
Farr, S.J, Kellaway, I.W. and Taylor, G. (1990). Drug deliveryto the respiratory tract. In: Routes of Drug Administration.(Eds. A.T. Florence and E. Salole). Wright, London, pp48-77.
McCallion, O.N.M. and Taylor, K.M.G. (1999) Ultrasonicnebulizers. In: Encyclopaedia of Pharmaceutical Technology.(Eds J. Swarbrick and J.C. Boylan). Marcel Dekker,New York, pp339-352.
McCallion, O.N.M.,Taylor K.M.G., Bridges, PA.,Thomas,M., Taylor AJ. (1996) Jet nebulisers for pulmonary drugdelivery. Int. J. Pharm., 130, 1-11.
Moren, F. (1981) Pressurised aerosols for oral inhalation.Int. J. Pharm., 8, 1-10.
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Timsina, M.P., Martin G.P., Marriott, C., Ganderton,D.,Yianneskis, M. (1994) Drug delivery to the respiratorytract using dry powder inhalers. Int. J. Pharm., 101, 1-13.
Wilson, C.G. and Washington N. (1989) PhysiologicalPharmaceutics: Biological Barriers to Drug Absorption. EllisHorwood, Chichester.
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