Influence of environmental humidity on particle size

As a particle enters the respiratory tract, the change from ambient to high relative humidity (approximately 99%) results in condensation of water on to the particle surface, which continues until the vapour pressure of the water equals that of the surrounding atmosphere. For water-insoluble materials, this results in a negligibly thin film of water; however, with water-soluble materials a solution is formed on the particle surface. As the vapour pressure of the solution is lower than that of pure solvent at the same temperature, water will continue to condense until equilibrium between vapour pressures is reached, i.e. the particle will increase in size. The final equilibrium diameter is constrained by the Kelvin effect, i.e. the vapour pressure of a droplet solution is higher than that for a planar surface, and is a function of the particle’s original diameter. Hygroscopic growth will affect the deposition of particles, resulting in deposition higher in the respiratory tract than would have been predicted from measurements of their initial size.

Inertial impaction

The air stream changes direction in the throat, or where a bifurcation occurs in the respiratory tract. Particles within the air stream, having sufficiently high momentum, will impact on the airways’ walls rather than following the changing air stream. This deposition mechanism is particularly important for large particles having a diameter greater than 5 µm, and particularly greater than 10 µm, and is common in the upper airways, being the principal mechanism for deposition in the nose, mouth, pharynx and larynx and the large conducting airways. With the continuous branching of the conducting airways, the velocity of the air stream decreases and impaction becomes a less important mechanism for deposition.

The probability of impaction is proportional to:

(37.2)

where θ is the change in airways direction, V is air stream velocity and r is the airway’s radius. V_t is the terminal settling velocity (see Eqn 37.3).

Gravitational sedimentation

From Stokes’ Law, particles settling under gravity will attain a constant terminal settling velocity, V_t:

(37.3)

where ρ is particle density, g is the gravitational constant, d is particle diameter and η is air viscosity.

Thus, gravitational sedimentation of an inhaled particle is dependent on its size and density, in addition to its residence time in the airways. Sedimentation is an important deposition mechanism for particles in the size range 0.5–3 µm, in the small airways and alveoli, for particles that have escaped deposition by impaction.

Brownian diffusion

Collision and bombardment of small particles by molecules in the respiratory tract produce Brownian motion. The resultant movement of particles from high to low concentrations causes them to move from the aerosol cloud to the airways’ walls. Diffusion is inversely proportional to particle size. It is the predominant mechanism for particles smaller than 0.5 µm, with the rate of diffusion, given by the Stokes-Einstein equation:

(37.4)

where D is the diffusion coefficient, k_B is the Boltzmann’s constant, T is the absolute temperature, η is viscosity and d is particle diameter.

Other mechanisms of deposition

Although impaction, sedimentation and diffusion are the most important mechanisms for drug deposition in the respiratory tract, other mechanisms may occur. These include interception, whereby particles having extreme shapes, such as fibres, physically catch on to the airways’ walls as they pass through the respiratory tract, and electrostatic attraction, whereby an electrostatic charge on a particle induces an opposing charge on the walls of the respiratory tract, resulting in attraction between particle and walls.

Effect of particle size on deposition mechanism

Different deposition mechanisms are important for different sized particles. Those greater than 5 µm will deposit predominantly by inertial impaction in the upper airways. Particles sized between 1 and 5 µm deposit predominantly by gravitational sedimentation in the lower airways, especially during slow, deep breathing and particles less than 1 µm deposit by Brownian diffusion in the stagnant air of the lower airways. Particles of approximately 0.5 µm are inefficiently deposited, being too large for effective deposition by Brownian diffusion and too small for effective impaction or sedimentation, and they are often quickly exhaled. This size of minimum deposition should thus be considered during formulation, although for the reasons of environmental humidity discussed previously, the equilibrium diameter in the airways may be significantly larger than the original particle size in the formulation.

Breathing patterns

Patient-dependent factors, such as breathing patterns, lung physiology and the presence of pulmonary disease also affect particle deposition. For instance, the larger the inhaled volume, the greater the peripheral distribution of particles in the lung, whilst increasing inhalation flow rate enhances deposition in the larger airways by inertial impaction. Breath-holding after inhalation increases the deposition of particles by sedimentation and diffusion. Optimal aerosol deposition occurs with slow, deep inhalations to total lung capacity, followed by breath-holding prior to exhalation. It should be noted that changes in the airways resulting from disease states, for instance airways’ obstruction, may affect the deposition profile of an inhaled aerosol.

Clearance of inhaled particles and drug absorption

Particles deposited in the ciliated conducting airways are cleared by mucociliary clearance within 24 hours and are ultimately swallowed. The composition of mucus and the process of mucociliary clearance are discussed in Chapter 38. Insoluble particles penetrating to the alveolar regions, and which are not solubilized in situ, are removed more slowly. Alveolar macrophages engulf such particles and may then migrate to the bottom of the mucociliary escalator, or alternatively may be removed via the lymphatics. The clearance of particle-loaded macrophages occurs over a period of days or weeks.

Hydrophobic compounds are usually absorbed at a rate dependent on their oil/water partition coefficients, whereas hydrophilic materials are poorly absorbed through membrane pores at rates inversely proportional to molecular size. Thus, the airways’ membrane, like the gastrointestinal tract, is preferably permeable to the unionized form of a drug. Some drugs, such as sodium cromoglicate, are partly absorbed by a saturable active transport mechanism, whilst large macromolecules may be absorbed by transcytosis. The rate of drug absorption, and consequently drug action, can be influenced by the formulation. Rapid drug action can generally be achieved using solutions or powders of aqueous soluble salts, whereas slower or prolonged absorption may be achieved using suspension formulations, powders of less soluble salts or novel drug delivery systems such as liposomes and microspheres.

Containers

Pharmaceutical aerosols may be packaged in tin-plated steel, plastic-coated glass or aluminium containers. In practice, pMDIs are generally presented in aluminium canisters, produced by extrusion to give seamless containers with a capacity of 10–30 mL. Aluminium is relatively inert and may be used uncoated where there is no chemical instability between container and contents. Alternatively, aluminium containers with an internal coating of a chemically resistant organic material, such as an epoxy resin or polytetrafluoroethylene (PTFE), can be used.

Propellants

The propellants used in pMDI formulations are liquefied gases, traditionally chlorofluorocarbons (CFCs) which are now largely replaced by hydrofluoroalkanes (HFAs). At room temperature and pressure, these are gases but they are readily liquefied by decreasing temperature or increasing pressure. The head space of the aerosol canister is filled with propellant vapour, producing the saturation vapour pressure at that temperature. On spraying, medicament and propellant are expelled and the head volume increases. To reestablish the equilibrium, more propellant evaporates and so a constant pressure system with consistent spray characteristics is produced. The CFCs currently employed in pMDI formulations are trichlorofluoromethane (CFC-11), dichlorodifluoromethane (CFC-12) and dichlorotetrafluoroethane (CFC-114). Formulations generally comprise blends of CFC-11 and CFC-12 or CFC-11, CFC-12 and CFC-114 (Table 37.1), together with a surfactant such as a sorbitan ester, oleic acid or lecithin, which acts as a suspending agent and lubricates the valve.

CFCs and HFAs are numbered using a universal system. The first digit is the number of carbon atoms minus 1 (omitted if zero), the second is the number of hydrogen atoms plus 1, and the third is the number of fluorine atoms. Chlorine fills any remaining valencies, given the total number of atoms required to saturate the compound. If asymmetry is possible, this is designated by a letter. The symmetrical isomer is assigned the number described above; of the asymmetrical isomers, that designated the letter a is the most symmetrical, b the next most symmetrical, and so on. The CFCs are perfectly miscible with each other and suitable blends give a useful intermediate vapour pressure, usually about 450 kPa. The vapour pressure of the mixture of propellants is given by Raoult’s Law, i.e. the vapour pressure of a mixed system is equal to the sum of the mole fraction of each component multiplied by its vapour pressure:

(37.5)

where P is the total vapour pressure of the system and p_a and p_b are the partial vapour pressures of the components, a and b:

(37.6)

(37.7)

where x_a and x_b are the mole fractions and and are the partial vapour pressures of components a and b, respectively.

The reaction of CFCs with the ozone in the earth’s stratosphere, which absorbs ultraviolet radiation at 300 nm, and their contribution to global warming are major environmental concerns. CFCs pass to the stratosphere, where in the presence of UV they liberate chlorine, which reacts with ozone. The depletion of stratospheric ozone results in increased exposure to the UV-B part of the UV spectrum, resulting in a number of adverse effects, in particular an increased incidence of skin cancer. The Montreal Protocol of 1987 was a global ban on the production of the five worst ozone-depleting CFCs by the year 2000. This was amended in 1992, so that production of CFCs in developed countries was phased out by January 1996. In the European Union and USA, all ozone-depleting CFCs were banned by the end of 1995, except for certain limited uses. The inclusion of CFCs in pMDIs currently has an ‘essential use exemption’, which will remain until medically acceptable non-ozone depleting alternatives to the remaining CFC-based pMDIs are available. This exemption is reviewed regularly and very few CFC-based pMDIs are now marketed. In household and cosmetic aerosols, CFCs have been replaced by hydrocarbons, such as propane and butane. Alternatively, non-toxic compressed gases such as nitrogen dioxide, nitrogen and carbon dioxide may be used, for instance in food products. However, compressed gases do not maintain a constant pressure within the canister throughout its use, as the internal pressure is inversely proportionate to the head volume, and so product performance changes with usage. For reasons of toxicity and inflammability, hydrocarbons are not considered appropriate alternatives to CFCs for inhalation products and so non-ozone depleting alternatives to CFCs have been developed.

Propellants HFA-134a (trifluoromonofluoroethane) and HFA-227 (heptafluoropropane) are non-ozone depleting, non-flammable HFAs, also called hydrofluorocarbons (HFCs), which are now used as alternatives to CFC-12 (see Table 37.1). However, these gases contribute to global warming and further replacements may be sought in the future.

HFA-134a and HFA-227 have some physical properties, including density, which are similar to those of CFC-12 and, to a lesser extent, CFC-114. However, they have presented major formulation challenges; in particular, they are poor solvents for the surfactants commonly used in pMDI formulation and no alternative to CFC-11 is currently available. Ethanol is approved for use in formulations containing HFAs to allow dissolution of surfactants, and is included in marketed HFA pMDI products. However, ethanol has low volatility and its inclusion may consequently increase the droplet size of the emitted aerosols.

Metering valve

The metering valve of a pMDI permits the reproducible delivery of small volumes (25–100 µL) of product. Compared to the non-metering continuous-spray valve of conventional pressurized aerosols, the metering valve in pMDIs is used in the inverted position (Fig. 37.3). Depression of the valve stem allows the contents of the metering chamber to be discharged through the orifice in the valve stem and made available to the patient. After actuation, the metering chamber refills with liquid from the bulk and is ready to dispense the next dose. A corollary of this is that the pMDI needs to be primed, i.e. the metering chamber must be filled, prior to the first use by a patient. pMDI valves are complex in design and must protect the product from the environment, while also protecting against product loss during repeated use. The introduction of HFA propellants with different solvent properties has necessitated the development of new valve elastomers. The valve stem fits into the actuator, which is made of polyethylene or polypropylene. The dimensions of the orifice in the actuator play a crucial role, along with the propellant vapour pressure, in determining the shape and speed of the emitted aerosol plume.

Formulating pressurized metered-dose inhalers

Pressurized aerosols may be formulated as either solutions or suspensions of drug in the liquefied propellant. Solution preparations are two-phase systems. However, the propellants are poor solvents for most drugs. Cosolvents such as ethanol or isopropanol may be used, although their low volatility retards propellant evaporation. In practice, pressurized inhaler formulations have traditionally been almost exclusively suspensions. These three-phase systems are harder to formulate and all the problems of conventional suspension formulation, such as caking, agglomeration, particle growth, etc., need to be considered. Careful consideration must be given to the particle size of the solid (usually micronized to between 2 and 5 µm), valve clogging, moisture content, the solubility of active pharmaceutical ingredient in propellant (a salt may be desirable), the relative densities of propellant and drug, and the use of surfactants as suspending agents, e.g. lecithin, oleic acid and sorbitan trioleate (usually included at concentrations between 0.1 and 2.0% w/w). These surfactants are very poorly soluble (<0.02% w/w) in HFAs and so ethanol is usually employed as a cosolvent, though alternative surfactants are being developed. Solution formulations of some drugs, such as beclometasone dipropionate are now available. Evaporation of HFA propellant following actuation of these formulations results in smaller particle sizes than with conventional suspension formulations of the same drug, with consequent changes in its pulmonary distribution and bioavailability. The dose may be adjusted accordingly or the volatility of the product modified by the addition of a less volatile component, such as glycerol.

Filling pressurized metered-dose inhaler canisters

Canisters are filled either by liquefying the propellant by reducing its temperature (cold filling) or by filling the vapour at elevated pressure (pressure filling).

In cold filling, active compound, excipients and propellant are chilled and filled at about −60 °C. Additional propellant is then added at the same temperature and the canister sealed with the valve. In pressure filling, a drug/propellant (CFC-11) concentrate is produced and filled at effectively room temperature and pressure (in fact, usually slightly chilled to below 20 °C). The valve is crimped on to the canister and additional propellant (e.g. CFC-12) is filled at elevated pressure through the valve, in a process known as gassing. Pressure filling is most frequently employed for inhalation aerosols. However, no HFAs have the properties (high boiling point: 23.7 °C) of CFC-11 so a single-stage pressure filling process has been developed for HFA-based formulations whereby a concentrated solution or suspension of drug in propellant, under pressure, is filled into canisters through the valve, followed by addition of further propellant.

Once filled, the canisters are leak tested by placing them in a water bath at elevated temperature, usually 50–60 °C. Following storage to allow equilibration of the formulation and valve components, the containers are weighed to check for further leakage, prior to spray testing and insertion into actuators.

Advantages and disadvantages of pressurized metered-dose inhalers

The major advantages of pMDIs are their portability, low cost and disposability. Many doses (up to 200) are stored in the small canister and dose delivery is reproducible. The inert conditions created by the propellant vapour, together with the hermetically sealed container, protect drugs from oxidative degradation and microbiological contamination. However, pMDIs have disadvantages. They are inefficient at drug delivery. On actuation, the first propellant droplets exit at a high velocity which may exceed 30 m/s. Consequently, much of the drug is lost through impaction of these droplets in the oropharyngeal areas. The mean emitted droplet size typically exceeds 40 µm, and propellants may not evaporate sufficiently rapidly for their size to decrease to that suitable for deep lung deposition. Vaporization of the droplets is hindered by the low volatility of CFC-11, which is present in concentrations of at least 25% in most CFC-based formulations. Evaporation, such that the aerodynamic diameter of the particles is close to that of the original micronized drug, may not occur until 5 seconds after actuation.

An additional problem with pMDIs, which is beyond the control of the formulator and manufacturer, is their incorrect use by patients. Reported problems include:

Correct use by patients is vital for effective drug deposition and therapeutic action. Ideally, the pMDI should be actuated during the course of a slow, deep inhalation, followed by a period of breath-holding. Many patients find this difficult, especially children and the elderly. The misuse of pMDIs through poor inhalation/actuation coordination can be significantly reduced with appropriate instruction and counselling. However, it should be noted that even using the correct inhalation technique, only 10–20% of the stated emitted dose may be delivered to the site of action.

Spacers and breath-actuated metered-dose inhalers

Some of the disadvantages of pMDIs, namely inhalation/actuation coordination and the premature deposition of large droplets high in the airways, can be overcome by using extension devices or ‘spacers’ positioned between the pMDI and the patient (Fig. 37.4), and thus these are frequently employed with a pMDI, for administering aerosol medications to young children (see also Chapter 43). The dose from a pMDI is discharged directly into the reservoir prior to inhalation. This reduces the initial droplet velocity, large droplets may be removed by impaction, efficient propellant evaporation occurs and the need for actuation/inhalation coordination is removed. The disadvantage of traditional spacers, though effective, is that they may be cumbersome because of their large volume, e.g. Volumatic^® (GlaxoSmithKline), though smaller, medium-volume spacers are now available, e.g. AeroChamber Plus^® (GlaxoSmithKline). Alternatively, extension tubes may be built into the design of the pMDI itself as an extended mouthpiece, e.g. Syncroner^® (Sanofi-Aventis) and Spacer Inhalers^® (AstraZeneca). Breath-actuated pMDIs do not release drug until inspiration occurs. In the Autohaler^® (3M), an inspiratory demand valve triggers a spring mechanism to release drug, whilst in the Easi-breathe^® (Teva), a vacuum in the device is released on inspiration to trigger the actuation. Breath-actuated devices overcome the coordination problems of conventional pMDIs and are easy to use without adding bulk to the device.

Formulating dry powder inhalers

To produce particles of a suitable size (preferably less than 5 µm), drug powders for use in inhalation systems are usually micronized. Alternatives are spray drying, spray freeze drying and supercritical fluid technology. The high-energy powders produced by micronization have poor flow properties because of their static, cohesive and adhesive nature. The flowability of a powder is affected by physical properties, including particle size and shape, density, surface roughness, hardness, moisture content and bulk density.

To improve their flow properties, poorly flowing drug particles are generally mixed with larger ‘carrier’ particles (median size usually 30–150 µm) of an inert excipient, usually lactose (α-lactose monohydrate). Drug and carrier particles are mixed to produce an ordered mix in which the small drug particles attach to the surface of the larger carrier particles. This not only improves liberation of the drug from the inhalation device by improving powder flow, but also improves the uniformity of capsule or device filling. Once liberated from the device, the turbulent air flow generated within the inhalation device should be sufficient for the deaggregation of the drug/carrier aggregates. The larger carrier particles impact in the throat, whereas smaller drug particles are carried in the inhaled air deeper into the respiratory tract.

The success of DPI formulations depends on the adhesion of drug and carrier during mixing and filling of devices or hard gelatin capsules, followed by the ability of the drug to detach from the carrier during inhalation, such that free drug is available to penetrate to the peripheral airways. Adhesion and detachment will depend on the morphology of the particle surfaces and surface energies, which may be influenced by the chemical nature of the materials involved and the nature of powder processing. Additionally, a ternary mix may be employed. Thus, fine particle size lactose can be added to conventional carrier lactose, to occupy the high energy sites on the larger carrier particles. Only low energy sites remain for drug-carrier interaction, enhancing the detachment of drug particles during inhalation of the formulation. Likewise, materials such as leucine and magnesium stearate may be included in formulations to modify the adhesion properties between drug and carrier particles. The interactions between micronized drugs, fine carrier particles and large carrier particles in DPI formulations are discussed in Chapter 8.

The performance of DPI systems is thus strongly dependent on formulation factors. Other factors affecting aerosol delivery and deposition include the design of the delivery device and a patient’s inhalation technique.

Unit-dose devices with drug in hard gelatin capsules

The first DPI device developed was the Spinhaler^® (Rhône-Poulenc Rorer) for the delivery of sodium cromoglicate. Each dose, contained in a hard gelatin capsule, was placed individually into the device, in a loose-fitting rotor. The capsule was pierced by two metal needles on either side of the capsule, and inhaled air flow though the device caused a turbovibratory air pattern as the rotor rotated rapidly, resulting in the powder being dispersed to the capsule walls and out through the perforations into the inspired air. Although the tendency nowadays is to develop multiple-dosing devices, other hard gelatin capsule-based devices, working on similar principles are still available, e.g. the HandiHaler^® (Boehringer Ingelheim/Pfizer) and the Aerolizer/Cyclohaler^® (Novartis/Teva, Fig. 37.5).

Multidose devices with drug in foil blisters

The main disadvantage of hard gelatin capsule-based devices, namely the individual loading of each dose, has been overcome with the development of the Diskhaler^® (GlaxoSmithKline). In this system, drug is mixed with a coarse lactose carrier and filled into an aluminium foil blister disc which is loaded, by the patient, into the device on a support wheel (Fig. 37.6). Each disc contains four or eight doses of drug and the blisters are pierced with a needle as a result of mechanical leverage of the lid. Air flow through the blister causes the powder to disperse as the patient inhales through the mouthpiece. The foil blisters are numbered, so that the patient knows the number of doses remaining.

Multidose devices with drug preloaded in inhaler

The evolution of the Diskhaler led to the Accuhaler^® or Diskus^® Inhaler (GlaxoSmithKline), in which drug/carrier mix is preloaded into the device in foil-covered blister pockets containing 60 doses (Fig. 37.7). The foil lid is peeled off the drug-containing pockets as each dose is advanced, with the blisters and lids being wound up separately within the device, which is discarded at the end of operation. As each dose is packaged separately and only momentarily exposed to ambient conditions prior to inhalation, the Diskhaler and Accuhaler are relatively insensitive to humidity compared to hard gelatin capsule-based systems.

An alternative approach is a reservoir type of device, in which a dose is accurately measured and delivered from a drug reservoir. In the Clickhaler® DPI (Innovata Biomed), a drug–lactose blend is stored in a reservoir. Metering cups are filled by gravity from this reservoir and delivered to an inhalation passage, from which the dose is inhaled. The device is shaken before use. The device is capable of holding up to 200 doses and incorporates a dose counter which indicates the number of metered doses and informs patients when the device, which is discarded after use, is nearly empty. After the final dose has been dispensed, the push button locks to prevent further use.

The Turbohaler® (AstraZeneca) has overcome the need for both a carrier and the loading of individual doses (Fig. 37.8). The device contains a large number of doses (up to 200) of undiluted, loosely aggregated micronized drug, which is stored in a reservoir from which it flows on to a rotating disc in the dosing unit. The fine holes in the disc are filled and excess drug is removed by scrapers. As the rotating disc is turned, by moving a turning grip back and forth, one metered dose is presented to the inhalation channel and this is inhaled by the patient, with the turbulent air flow created within the device breaking up any drug aggregates. A dose indicator is incorporated. The Turbohaler requires a higher inspiratory effort than the Diskhaler, owing to its higher internal resistance, and is more sensitive to humidity if not closed quickly after each use.

Breath-assisted devices

Several devices have been developed which reduce or eliminate the reliance on the patient’s inspiratory effort to disperse the drug. This is advantageous as inspiratory effort may be affected by the patient’s age and/or clinical condition. For instance, Nektar Therapeutics produced a device for the delivery of insulin as a fine powder, in which compressed air was used to disperse drug from a unit-dose package into a large holding chamber, from which it was inhaled by the patient. The device was marketed briefly by Pfizer for the delivery of insulin by inhalation, but was withdrawn for many reasons including; cost, the cumbersome design of the delivery device and the need for injections of insulin to supplement inhaled drug. Another novel device, the Spiros® DPI (Dura) (Fig. 37.9), is a breath-actuated, effort-assisted device which uses a battery-powered impeller to deaggregate and aerosolize the drug powder, which is held in a rotating cassette. The device functions relatively independently of the patient’s inspiratory flow rate and can be used by those patients who can only achieve low inspiratory flow rates.

Jet nebulizers

Jet nebulizers (also called air-jet or air-blast nebulizers) use compressed gas (air or oxygen) from a compressed gas cylinder, hospital air-line or electrical compressor to convert a liquid (usually an aqueous solution) into a spray. The jet of high-velocity gas is passed either tangentially or coaxially through a narrow Venturi nozzle, typically 0.3–0.7 mm in diameter. An area of negative pressure, where the air jet emerges, causes liquid to be drawn up a feed tube from a fluid reservoir by the Bernoulli effect (Fig. 37.10). Liquid emerges as fine filaments, which collapse into droplets as a result of surface tension. A proportion of the resultant (primary) aerosol leaves the nebulizer directly; the remaining large, non-respirable droplets impact on baffles or the walls of the nebulizer chamber and are recycled into the reservoir fluid.

Nebulizers are operated continuously and because the inspiratory phase of breathing constitutes approximately one-third of the breathing cycle, a large proportion of the emitted aerosol is not inhaled but is released into the environment. Open-vent nebulizers, incorporating inhalation and exhalation valves, e.g. the Pari LC® nebulizer (Pari), have been developed in which the patient’s own breath boosts nebulizer performance, with aerosol production matching the patient’s tidal volume and greatly enhancing drug delivery. On exhalation, the aerosol being produced is generated only from the compressor gas source, thereby minimizing drug wastage.

The rate of gas flow driving atomization is the major determinant of the aerosol droplet size and rate of drug delivery for jet nebulizers; for instance, there may be up to a 50% reduction in the mass median aerodynamic diameter when the flow rate is increased from 4 to 8 L/min, with a linear increase in the proportion of droplets less than 5 µm.

Ultrasonic nebulizers

In ultrasonic nebulizers the energy necessary to atomize liquids comes from a piezoelectric crystal vibrating at high frequency. At sufficiently high ultrasonic intensities, a fountain of liquid is formed in the nebulizer chamber. Large droplets are emitted from the apex and a ‘fog’ of small droplets is emitted from the lower part (Fig. 37.11). Some models have a fan to blow the respirable droplets out of the device, whereas in others the aerosol only becomes available to the patient during inhalation.

Mesh nebulizers

Recently, mesh nebulizers have come commercially available, in which aerosols are generated by passing liquids through a vibrating mesh or plate with multiple apertures. The energy of vibration comes from a vibrating piezoelectric crystal attached to a horn transducer which transmits vibrations to a perforated plate with up to 6000 tapered holes (Fig. 37.12) or from an aerosol generator comprising a domed aperture plate, with up to 10 000 tapered holes and a vibrational element which contracts and expands to generate the aerosol. These devices generate aerosols with a high fine particle fraction, deliver fluids very rapidly and have very small residual volumes (see below) compared to jet and ultrasonic nebulizers. A recent development, has been the ‘Adaptive Aerosol Delivery’ (AAD) system employed in the I-neb AAD® mesh nebulizer (Philips Respironics), which analyse the patient’s breathing pattern and emits aerosol only during inhalation, thus eliminating wastage during exhalation.

Formulating nebulizer fluids

Nebulizer fluids are formulated in water, occasionally with the addition of a cosolvent such as ethanol, and with the addition of surfactants for suspension formulations. Because hypoosmotic and hyperosmotic solutions may cause bronchoconstriction, as may high hydrogen ion concentrations, isoosmotic solutions of pH greater than 5 are usually employed. Stabilizers such as antioxidants and preservatives may also be included, although these may also cause bronchospasm, and for this reason sulfites in particular are generally avoided as antioxidants in such formulations. Although chemically preserved multidose preparations are commercially available, nebulizer formulations are generally presented as sterile, isotonic unit doses (usually 1–2.5 mL) without a preservative.

Whilst most nebulizer formulations are solutions, suspensions of micronized drug are also available for delivery from nebulizers. In general, suspensions are poorly delivered from ultrasonic nebulizers, whilst mesh nebulizers are considered suitable for delivering suspensions. With jet nebulizers the efficiency of drug delivery increases as the size of suspended drug is decreased, with little or no delivery of particles when they exceed the droplet size of the nebulized aerosol.

As the formulation of fluids for delivery by nebulizers is relatively simple, these devices are frequently the first to be employed when investigating the delivery of new entities to the human lung. Recently, they have been used for the delivery of biopharmaceuticals and drug delivery systems, such as liposomes. In general, ultrasonic nebulizers have not been successful for delivering either biopharmaceuticals or liposomes, because of denaturation resulting from the elevated temperatures produced. Consequently, ultrasonic nebulizers are expressly excluded for the delivery of recombinant human deoxyribonuclease in the management of cystic fibrosis. Mesh and jet nebulizers have been successfully used to deliver some peptides, nucleic acids and liposome formulations, although the shearing forces that occur in jet nebulizers may produce time-dependent damage to some materials.

Physicochemical properties of nebulizer fluids

The viscosity and surface tension of a liquid being nebulized may affect the output of nebulizers, as energy is required to overcome viscous forces and to create a new surface. However, the size selectivity of the nebulizer design and dimensions, with more than 99% of the primary aerosol mass being recycled into the reservoir liquid, means that changes in the size distribution of the primary aerosol resulting from changes in the properties of the solution being atomized may not be reflected in the size distribution of the emitted aerosol. In general, the size of aerosol droplets is inversely proportional to viscosity for jet and mesh nebulizers and directly proportional to viscosity for ultrasonic nebulizers, with more viscous solutions requiring longer to nebulize to dryness and leaving larger residual volumes in the nebulizer following atomization. Surface tension effects are more complex, but usually a decrease in surface tension is associated with a reduction in mean aerosol size.

Temperature effects during nebulization

The aerosol output from a jet nebulizer comprises drug solution and solvent vapour, which saturates the outgoing air. This causes solute concentration to increase with time and results in a rapid decrease in the temperature of the liquid being nebulized by approximately 10–15 °C. This temperature decrease may be important clinically, as some asthma sufferers experience bronchoconstriction on inhalation of cold solutions. Further, the cooling effect within the reservoir fluid will reduce drug solubility and result in increased liquid surface tension and viscosity. Precipitation is uncommon with bronchodilators, which have high aqueous solubility, but problems may arise with less soluble drugs. In such instances the use of an ultrasonic nebulizer may be appropriate, as the operation of such devices may increase solution temperature by up to 10–15 °C. As indicated above, this temperature increase may have detrimental effects on heat-sensitive materials intended for nebulization. Mesh nebulizers have a negligible effect on fluid temperature.

Duration of nebulization and residual volume

Clinically, liquids may be nebulized for a specified period of time or, more commonly, they may be nebulized to ‘dryness’, which may be interpreted as sputtering time, which is the time when air is drawn up the feed tube and nebulization becomes erratic, although agitation of the nebulizer permits treatment to be continued; clinical time, which is the time at which therapy is ceased following sputtering; or total time, which is the time at which the production of aerosol ceases.

Regardless of the duration of nebulization, not all the fluid in the nebulizer can be atomized. Some liquid, usually about 1 mL for jet and ultrasonic nebulizers, remains as the ‘dead’ or ‘residual’ volume, associated with the baffles, internal structures and walls of the nebulizer. The proportion of drug retained as residual volume is more marked for smaller fill volumes; hence for a 2 mL fill volume, approximately 50% of fluid will remain associated with the nebulizer and be unavailable for delivery to the patient. This reduces to approximately 25% with a 4 mL fill volume, although there is a commensurate increase in the time necessary to nebulize to dryness. For this reason small-volume nebulizer fluids may be diluted to a larger volume by addition of a suitable diluent, usually sterile 0.9% sodium chloride solution. Vibrating-mesh nebulizers have a much smaller residual volume than jet or ultrasonic devices, which may mean that the nominal dose, or dose volume used with a mesh nebulizer should be reduced compared to that employed with a conventional nebulizer.

Variability between nebulizers

Many different models of nebulizer and compressor are commercially available, and the size of aerosols produced and the dose delivered can vary enormously. Variability may not only exist between different nebulizers but also between individual nebulizers of the same type, and repeated use of a single nebulizer may cause variability due to baffle wear and non-uniformity of assembly. Nebulizers, unlike the DPI and pMDI devices, are not manufactured by the producers of nebulizer solutions and suspension. The choice of nebulizer employed for their delivery is thus usually beyond the influence of the pharmaceutical manufacturer.