University of Groningen

Optimisation of dry powder inhalationBoer, Anne Haaije de

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Citation for published version (APA):Boer, A. H. D. (2005). Optimisation of dry powder inhalation: The application of air classifier and laserdiffraction technology for the generation and characterisation of aerosols from adhesive mixtures. [S.n.].

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Chapter 1

Introduction to pulmonary drug delivery

Part of this introduction has been published in Expert Opin. Drug Deliv. 1 (2004) 67-86. Note: for convenience of the reader, the registration mark ® has not been indicated in this chapter.

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Chapter 1: contents 1. General introduction 1.1. Pulmonary drug delivery

1.1.1. Historical perspective 1.1.2. Pulmonary delivery applications 1.1.3. Stimuli for development of (new) inhalation technologies and devices 1.1.4. Stimuli for development of (new) inhalation formulations

1.2. The target area for drug deposition 1.2.1. Receptor and absorption areas 1.2.2. Inspiratory air flow manoeuvres 1.2.3. Fluid dynamics in the respiratory tract 1.2.4. Particle dynamics in the respiratory tract 1.2.5. Particle characterisation of inhalation drugs 1.2.6. Modelling of aerosol particle deposition

1.2.7. The desired aerodynamic size distribution of inhalation drugs 2. Medical aerosol generators (inhalation devices)

2.1. Nebulisers 2.2. Metered dose inhalers (mdi’s) 2.3. Dry powder inhalers (dpi’s)

2.3.1. Basic design and functional parts of a dpi 2.3.2. The powder formulation

2.3.2.1. Particle-particle interactions 2.3.2.2. Adhesive mixtures: historical review

2.3.2.3. Adhesive mixtures: the role of carrier properties 2.3.2.4. Spherical pellets 2.3.3. Powder de-agglomeration during inhalation

2.3.3.1. De-agglomeration principles 2.3.4. Balancing steps in dry powder inhalation 2.3.5. The dpi as an external resistance to air flow 2.3.6. The use of a dpi: variables and interactions 2.3.7. Dpi performance

2.3.8. Recent trends in dry powder inhalation 3. Aerosol characterisation 3.1. Relevance and principles of aerosol characterisation

3.2. Cascade impactor analysis (cia) 3.3. Laser diffraction analysis (lda)

4. The aim of this thesis 5. References

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1. General introduction 1.1. Pulmonary drug delivery 1.1.1. Historical perspective

The respiratory tract is one of the oldest routes for the administration of drugs. The history of inhalation therapy covers a period of more than 4000 years (Grossman, 1994; O’Callaghan et al., 2002). Smoking of the leaves of the Atropa belladonna plant in India, approximately 2000 BC, is frequently mentioned as the first reference to pulmonary drug delivery. More than a millennium later, the ancient Greeks recognised the benefit of the volatile gases from the pine forests of Libya in the treatment of tuberculosis of the lungs, whereas ancient Egyptians inhaled the vapours that were released when plants of the Hyosycamus muticus were exposed to heat (Miller, 1973).

Inhalation therapy for the treatment of various chest infections and tuberculosis was (re)discovered in Europe (UK) in the 17th and 19th century (O’Callaghan et al., 2002). Aromatic plants, balsams and sulfur were smoked in a pipe alone, or after mixing with tobacco. The Datura (stramonium) tobacco mixture in the Potter’s asthma cigarette is one of the best known examples in this respect. In the beginning of the nineteenth century, inhaled vapours became replaced by wet aerosols. The first atomiser (‘apparatus for the pulverisation of liquids’) was constructed in 1829, followed by a portable device in 1860. More than a century later, in 1938, the first hand held bulb nebuliser became popular (Grossman, 1994).

In the second half of the 20th century, inhaler technology diverged rapidly along three different pathways. The widespread introduction of electric pumps enabled more continuous air flows through jet nebulisers which, in combination with the use of baffles, reduced the width of the size distribution of the aerosol substantially. Portable jet nebulisers evolved in a great variety of devices and, in addition to that, ultrasonic nebulisers were introduced in the 1960s. The first metered dose inhaler (mdi), utilizing CFC (chorofluorocarbon) propellants, became available in 1956 (Medihaler, Riker) and nearly fifteen years later (1970), the first dry powder inhaler (dpi: Fisons Spinhaler) reached the market (Ashurst et al., 2000). All three types of aerosol generation devices (nebulisers, metered dose inhalers and dry powder inhalers) will be discussed more in detail in Paragraph 2. 1.1.2. Pulmonary drug delivery applications and advances The (potential) advantages of the pulmonary route in the treatment of local disorders and diseases of the lungs, like for instance asthma, chronic obstructive pulmonary disease (COPD), cystic fibrosis (CF) and lung cancer, have been described extensively (e.g. Morén, 1987; Brown et al., 2001). They include a rapid onset of therapeutic action, a strongly reduced dose compared to the oral or other systemic routes and consequently, reduced systemic side effects. The high safety margin of inhaled corticosteroids (compared to the oral route) has encouraged widespread use of this type of drug in patients with moderate to severe asthma, and corticosteroid therapy has become the cornerstone of asthma management since effective inhalation systems for this type of drug became available (Pedersen and O’Byrne, 1997). Other actives for local therapy are bronchodilators and antibiotics. There is a growing interest in drug delivery via the human respiratory tract also in order to achieve systemic effects (Corkery, K., 2000; Schultz, 2002). Systemic administration of drugs through the lungs is particularly interesting for substances that suffer from metabolic breakdown in the gastro-intestinal tract. Also large molecules, such as peptides and proteins, that can not pass the absorbing membrane after oral administration are in focus for pulmonary delivery (Skyler et al., 2001; Sakagami, 2002). Many examples of systemic delivery through the lungs have recently been mentioned, as for instance summarised by de Boer et al., 2001b. Systemic delivery through the respiratory tract has become possible in the past two decades, as a result

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of strongly improved inhalation technology, but the mechanisms of adsorption in the lungs for large molecules are still poorly understood (Patton, 1996). The pulmonary route may also become a successful alternative for the parenteral route in case of vaccination, as suggested by studies with vaccines against measels (LiCalsi et al., 1999; Dilraj et al., 2000) and influenza (Jemski and Walker, 1976). 1.1.3. Stimuli for development of (new) inhalation technologies and devices As explained in Paragraph 1.1.1, a first large scale diversification of inhalation devices occurred mid-20th century, following the introduction of electrical pumps for jet nebulisers. The introduction of isoprenaline in approximately the same period (1951) is considered as another major stimulus to inhalation therapy (O’Callaghan et al., 2002). However, the increasing growth rates for asthma and COPD have stressed most of all the developments of new inhalation drugs and therapies in the second half of the 20th century. They were the driving force for large numbers of new device developments, making use of technologies from various disciplines. Examples are the metered dose inhalers (1956), ultrasonic nebulisers (1960s) and dry powder inhalers (1970s) mentioned already in Paragraph 1.1.1. Regulatory requirements have an increasing influence on inhaler design. A major impact on the design of particularly metered dose inhalers (Newman 1990), but also a stimulus to the development of new dry powder inhalers (Prime et al., 1997), was given by the Montreal protocol in 1987, calling for signatory countries to phase out the production of CFC propellants (by January 1, 1996) in order to stop depletion of the ozone layer.

Recent developments in inhaler technology focus on principles that increase the efficiency of dose delivery, as well as the accuracy and reproducibility of drug deposition in the target area, and improve patient compliance (Dolovich, 1999). Stimulation of such developments does not exclusively come from (lung) physicians, but also from patients asking for cost reduction, ease of operation and a high degree of mobility. Breath activated aerosol generation is only one example, as applied for instance in the AERx membrane nebuliser (Aradigm), the HaloLite nebuliser with adaptive aerosol delivery (AAD) and the QVAR Autohaler (3M) metered dose inhaler. The aerosol is not delivered until a certain threshold value for the inspiratory flow rate through the device is exceeded. This principle of operation, achieved with various means, synchronises inhalation and firing of a dose. It can be further optimised by controlling the duration of the aerosolisation process, in order to achieve maximal transport of the aerosol by the inhaled air stream to the site of deposition. A positive effect on dose deposition efficacy and reproducibility is also obtained from controlled inhalation manoeuvres (Brand et al., 2000). Compliance may be increased by reducing the effort and duration of inhalation, for instance by replacing nebuliser therapy with dpi therapy (e.g. Le Brun et al., 2002; de Boer et al., 2002a). Finally, certain events or new diseases may influence the acceptance of, or change the preference for certain types of inhalation devices, as has recently (2003) been shown in Canada. The SARS outbreak in Toronto has nearly put a ban on nebuliser therapy in this country and suddenly increased the interest in dry powder inhalation. On the other hand, the BSE problems in the UK, have made the North American authorities (unwarrantably) reluctant to accept lactose as excipient in inhalation formulations. 1.1.4. Stimuli for development of (new) inhalation formulations

Great formulation efforts have recently been put in the substitution of CFC-containing drug formulations for mdi’s by hydrofluoroalkane (HFA) containing formulations (Vervaet and Byron, 1999). However most research is focussed on improving dry powder formulations, particularly the so-called adhesive mixtures in which micronised drug particles (with the desired size distribution) are attached to the surface of much larger carrier or host crystals with the objective to increase the dose reproducibility by dilution and by improving the flow

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properties of the drug. The attention of a great number of research groups is focussed on studying and controlling the parameters that determine the adhesive force between the drug and carrier particles, such as the carrier surface and bulk properties and the mixing process. This is necessary because the fine particle fractions achieved with most currently marketed dry powder inhalers is frequently only 20 to 30% of the nominal dose. Adhesive mixtures for inhalation are the subject of the Paragraphs 2.3.2.2 and 2.3.2.3 and the Chapters 5-8 of this thesis. The expectation that systemic action can be achieved from drugs delivered to the lungs, has extended and broadened the efforts to achieve controlled and reproducible peripheral lung deposition particularly in the last two decades (Patton, 1996). As discussed in Paragraph 1.1.2, this refers particularly to drugs with poor to no bioavailability when administered via the oral route. Such drugs may consist of large molecules (e.g. peptides and proteins) for which special formulation processes are required, particularly when they are administered in the dry state (Johnson, 1997). This includes for instance stabilisation with sugar glass (e.g. inulin) technology. The preparation of large porous particles has been proposed as a means to avoid rapid particle removal from the deep lungs by phagocytic clearance (Edwards, et al., 1998). This is necessary when such particles need to have a long residence time in the lungs in the case of slow drug release. Poly(lactic acid-co-glycolic acid) or poly(lactic acid-co-lysine-graft-lysine) have been mentioned as excipients for drug encapsulation (Edwards, et al., 1997). Other particle engineering techniques for the production of so-called ‘smart particles’ are surveyed in Paragraph 2.3.8. Monodisperse aerosols that are recommended to improve the precision of airway targeting (e.g. Zanen et al., 1994; Brand et al., 2000) can not (yet) be manufactured in the dry state on a production scale with the current state of the technique. Techniques for the production of monosized wet aerosols are mentioned in Paragraph 2.1. 1.2. The target area for drug deposition 1.2.1. Receptor and absorption areas

Target areas in the lungs vary with the type of disease, and therefore, with the type of drug. Although the precise location of specific receptor and absorption areas, or sites of infection, are not the subject of this study, a brief summary is given in order to understand the relevance of the aerodynamic size distribution of the drug to the penetration potential into the tracheobronchial and pulmonary compartments as well as the deposition efficiency on the site of action. Of particular interest to this study are the anti-inflammatory (budesonide) and bronchodilating drugs (salbutamol and formoterol), since they are the drugs applied in the marketed Novolizer (Viatris), which device makes use of the air classifier technology described in the Chapters 4, 9 and 10. But also relevant are the antibiotics in CF (tobramycin and colistin), for which different dry powder (test)inhalers were developed on the basis of air classifier technology (de Boer et al., 2002a; Le Brun et al., 2002).

Asthma is frequently associated with lymphocytes and eosinophil cells (Frew, 1996; Wever, 1996; Carroll et al., 1997). Saetta et al. (1998) reported the highest numbers of eosinophils in the walls of non-respiratory bronchioles (with diameters < 2 mm), but Synek et al. (1997) found an increased infiltration of (EG1 and EG2) positive eosinophil cells particularly in the larger airways (diameter > 2 mm). This has led to discussions about whether central or peripheral inflammation is more important (Hamid, 1997; Poutler, 1997). Both viewpoints are in disagreement with the findings of Carroll et al. (1997), who described that an increased recruitment of eosonophils with increased asthma severity occurs in all airway size groups. According to Chrystyn (1997), inflammation in asthma is present throughout the lungs. Pedersen and O’Byrne (1997) described that the molecular action of inhalation steroids occurs at intercellular glucocorticoid receptors which can be found in most

12

cell types. The latter three viewpoints may explain why distribution of inhaled steroids throughout the airways is often recommended (Leach et al., 1998), although it has also been assumed that the small airways comprise the optimal site for this type of drug in asthma treatment (Jackson, 1995). Bronchodilators (β2-agonists) in asthma and COPD, such as salbutamol and formoterol, interact with β2-adrenoceptors, which are also located on a variety of cells, including smooth muscle and epithelial cells. The concentration of β-receptors throughout the lungs is not the same. Places of highest density differ between studies, and vary from central lung (Howarth, 1997) to small airways (Chrystyn, 1997) and the alveoli (Barnes et al., 1982). Moreover, starting in the bronchioles, there is a gradual decrease in the ratio of β2 to β1-receptor type with decreasing airway diameter (Carstairs et al., 1985), whereas β2-agonist deposition in the alveoli results in reduced bronchodilatation because of the lack of smooth muscle in this region (Chrystyn, 1997). In contrast, muscarine receptors occur almost exclusively in the larger (proximal) airways (Howarth, 1997; Barnes et al., 1983).

The pulmonary complications in CF begin in the small peripheral airways and progress to the development of widespread bronchiectasis, which is most marked in the upper lobes (Geddes, 1997; Coates et al., 1997). This explains why the bronchial lumen (Ramsey, 1996; Van Devanter and Montgomery, 1998; Pai and Nahata, 2001; Sermet-Gaudelius et al., 2002) and the smaller bronchioles (Touw et al., 1995; Geddes, 1997) have both been considered as the target area for inhaled antibiotics in CF. More recently, it has been described that the inflammatory process is much more severe in the peripheral than in the central airways (Tiddens, 2002), which indicates the need for substantial peripheral drug deposition.

Macromolecular drugs for systemic action have to be delivered to the alveolar region, where the absorption surface area reaches values up to 125 m2 and barriers for drug passage to the systemic circulation are only between 0.2 and 0.5 µm (Gorman and Hall, 1973; Patton, 1996; Kim and Folinsbee, 1997). For delivery of vaccines through the lungs, targeting may be less critical however. LiCalsi et al. (1999) described that the CD46 receptors for the measles vaccine are present in nearly all cells. 1.2.2. Inspiratory air flow manoeuvres

The anatomy and physiology of the respiratory system has extensively been described elsewhere (e.g. Gorman and Hall, 1973). The pulmonary ventilation parameters that are relevant to inhalation therapy have been reviewed by de Koning (2001). Inhalation occurs when the chest volume is expanded. This expansion creates a pressure difference (dP) between the lungs and the atmosphere, which is the driving force for an air flow (Φ) into the lungs. The phases of exhalation are less interesting for pulmonary drug delivery, except for standard nebuliser therapy, as will be explained in Paragraph 2.1. Correlations between inspiratory and expiratory lung function parameters, that have been recommended to predict the attainable peak inspiratory flow rate (PIF) through an external air flow resistance (e.g. Brown, et al., 1995; Meijer et al., 1996) from the data obtained with standard lung function tests (e.g. PEF: peak expiratory flow rate) are neither of relevance to this thesis. During non-restricted inhalation (yielding so-called ‘control values’), PIF’s of 333 l/min for male and 214 l/min for female healthy volunteers are possible (Timsina et al., 1993). Peak flow rates at maximal effort during non-restricted inhalation of patients suffering from lung diseases may be (but are not necessarily) lower. Brown et al. (1991) measured PIF values of only 156 l/min during exacerbations of asthmatics. Sarinas et al. (1998) compared PIF values from pulmonary function testing of adult healthy volunteers with COPD and CF patients in three different degrees of severity: mild, moderate and severe, classified upon %FEV1 (forced expiratory volume in 1 s as percent of predicted). They found substantial decreases in PIF

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with increasing disease severity for both CF and COPD, PIF ranging from 248 l/min (mild; %FEV1 = 72) to 163 l/min (severe; %FEV1 = 36) for CF, and from 409 l/min (mild; %FEV1 = 69) to 239 l/min (severe; %FEV1 = 40) for COPD. Their mean value obtained for healthy subjects (331 l/min; %FEV1 = 98) confirms however, that mild COPD patients of approximately the same age may have a better inspiratory performance which could be the result of frequent inhaler use, resulting in respiratory muscle training (de Koning, 2001).

Age (Quanjer et al., 1993) also influences the inhalation performance. Holmann et al. (1989) compared the maximal inspiratory performance of different groups of children (5-25 year) and found PIF values of 150 l/min for healthy subjects versus (on average) 126 l/min for asthmatics with different degree of severity and 120 l/min for CF patients. However, the major determinant for PIF is the external air flow resistance (R), as shown in studies of de Boer et al. (1996) and de Koning (2001). This is also true for the flow increase rate (FIR) towards peak flow and the total inhalation time (t), which are the most important flow parameters for the operation of dry powder inhalers. Dpi’s normally have higher air flow resistances than nebulisers and metered dose inhalers (Richards and Saunders, 1993; Olsson and Asking, 1994), which is sometimes considered problematic, as they are breath operated devices (Paragraph 2.3.5) which rely on the kinetic energy of the inspiratory air flow through the inhaler.

It strongly depends on the working principle of the inhaler, how the patient has to be instructed to breathe or to inhale. Instructions for the use of metered dose inhalers are generally described in terms of calm and deep inhalation, whereas nebulisers are often used under normal breathing conditions. Denyer et al. (1997) measured peak flow rates of (15) adult healthy subjects during normal breathing between 25 and 48 l/min (with a mean of 35 l/min). For (15) asthmatics, (28) COPD and (8) CF patients, the ranges (means) were respectively 23-90 (56), 17-71 (37) and 29-89 (54) l/min, showing that a lung disease does not necessarily deteriorate normal breathing flows. This, at least for CF, seems in agreement with Tiddens (2002) who concluded that patients may have normal lung functions in spite of the existence of areas in the lungs with end-stage disease. But this conclusion may not be extended to inhalation at maximal effort through an external resistance (such as a dpi), as will be discussed more in detail in Paragraph 2.3.5.

For a typical inspiratory flow curve (Φ as function of the inhalation time) through a dry powder inhaler, the increase in flow rate towards the peak value is often steeper than the decrease towards zero flow (e.g. Brunell et al., 1996; Hawksworth et al., 2000). Average flow rate during inhalation through a dpi is between approximately 70 and 75% of the peak value, both at maximal and at comfortable inhalation (de Boer et al., 1996), although a considerable difference may exist between the inhalation manoeuvre at maximal and comfortable inspiratory effort (Clark and Hollingworth, 1993). Different studies have shown that the PIF’s through low resistance dpi’s at maximum effort may reach values of more than 150 l/min (Sumby et al., 1992; Sarinas et al., 1998; de Koning 2001), even for asthmatic children aged under 10 years ( Pedersen, 1986; Nielsen et al., 1997). On the other hand, a recent study with standardised inspiratory flow manoeuvres from an aerosol chamber includes a flow rate of only 6 l/min (Brand et al., 2000). This extends the total range of flow rates that has to be considered for particle deposition in the respiratory tract to a span of 6 to (at least) 150 l/min. 1.2.3. Fluid dynamics in the respiratory tract

One of the most frequently cited lung models to explain the principles of aerosol delivery to the respiratory tract is that of Weibel (eg. Hinds, 1982; Morrow and Yu, 1985; Gerrity, 1990). The model distinguishes 23 subsequent branchings of the airways, starting at the trachea (generation 0) with a diameter of 18 mm (for adults) which decreases to 0.41 mm in the alveolar sac (generation 23). As a result of the bifurcation of each airway, the number

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of airways increases from 1 (trachea) to 223 (8.388.608 alveolar sacs). This has the consequence that the cross section for air flow increases exponentially (Fig. 1.1) after an initial small decrease from trachea to lobar bronchus, in spite of a decreasing diameter (by a factor 44). Due to their relatively small total cross sectional area, the upper and central airways account for approx. 90% of total airway resistance (Netter et al., 1980).

Figure 1.1. Cross section for air flow (left) and air velocity at three different inspiratory flow rates (right) as function of airway generation.

Weibel’s model describes the airways as circular channels (in cross section). From the

total cross section, the air velocity in each airway generation can be estimated as function of the inspiratory flow rate (Fig. 1.1). The model also provides channel lengths, which enables calculation of the cumulative lung volume (5 l for adults). However, the model lacks morphological description of the alveoli, whereas it does not take account of individual branching angles, or angles of inclination of individual airways to the horizontal (Gerrity, 1990). Therefore, it is a simplified presentation of reality (Netter et al., 1980). Figure 1.2. Reynolds number as function of airway number at three different inspiratory flow rates.

The relevance of the Reynolds number to fluid mechanics and particle deposition in the human lungs has been described by others (e.g. de Jongh, 1995). As a result of the

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exponentially decreasing air velocity from the lobar bronchus towards the alveoli, and the decreasing airway diameter in the same direction, the Reynolds number decreases rapidly from more than 4000 (at flow rates of 60 l/min and higher) in the trachea (turbulent flow) to values far below 2000 (laminar flow) already at generation 4 (Fig. 1.2). Local turbulences at the bifurcations and at constrictions occur also in the smaller airways however, in spite of the fact that the average angle of bifurcation is 37 degrees, which is the angle with least disturbance of the streamlines of the air. At the bifurcations, also local regions of back flow are present during inhalation, which may contribute to the aerosol deposition (de Jongh, 1995). 1.2.4. Particle dynamics in the respiratory tract Different divisions between airway regions have been made on the basis of functionality, diameter, air flow resistance, behaviour of airborne particles in these airways, etc. For instance, upper plus central (> 2 mm) and peripheral (< 2 mm) airways have been distinguished as separate regions (Netter et al., 1980; Brown et al., 2001). Division has also been made into conducting (generation 0-11), transitional (12-16) and respiratory (17-23) airways (e.g. Martonen, 1993), or into tracheobronchial (generations 0-16) and alveolar (pulmonary) regions (generations 17-23: Hinds, 1982). Inhaled aerosol particles travel with the air through the branching system of airways, in which the air is warmed and humidified, until these particles are captured by collision with an airway or alveolar wall, or exhaled again. Figure 1.3. Forces acting on an airborne particle in a bent airway duct

Basically, three different types of forces act on airborne particles passing a bend in an airway duct: the force of gravity (FG); the drag or resistance force of the air (FD) and a stopping or inertial force (FS), as shown in Fig. 1.3. The force of gravity is constantly acting on all particles in a field of gravity. The size of the force (FG = m.g) is proportional to the mass (m), and thus to the third power of the particle diameter (d), but independent of the particle velocity (g, is the acceleration of gravity). The drag force (FD) acting on aerosol particles is given by the Stokes equation (because frictional forces are much larger than inertial forces), which contains the particle diameter to the first powder and particle velocity relative to the velocity of the air (UPA):

FD = 3.π.η.UPA.d.χ/CC [1.1]

X

Y

FS: stopping force

FD: drag forceFG: force of gravity

X

Y

X

Y

FS: stopping force

FD: drag forceFG: force of gravity

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where η is the viscosity of the air

χ is the dynamic shape factor of the particle, and CC is the Cunningham correction factor for slip flow

Different expressions can be given to the inertial or stopping force (FS), which all

relate to particle momentum (product of particle mass (m) and particle velocity (U)). Frequently, also the stopping distance (S), or inertial range, is used, which is the distance necessary to reduce the initial particle velocity (U0) to zero by the action of the resistance force (FD). The stopping distance is the product of particle mobility (B), defined as particle velocity (U) per unit resistance force (FD), and particle momentum (in terms of initial velocity: m.U0).

S = B.m.U0 = (U/FD).(π/6.d3.ρP.U0) [1.2]

where ρP is the particle density. After substitution of FD by 3.π.η.d.U.χ/CC and re-

arrangement of terms, equation 1.2 can be written as: S = (d2.ρP.U0.CC)/(18.η.χ) [1.2a]

A large stopping distance (high particle mass, high velocity or both) relative to the

drag force tends to bring the particle in contact with the airway wall by inertial deposition (in the bend at a bifurcation). A high drag force may conduct the particle into the next airway generation. Unlike in a cascade impactor, it is difficult to predict whether a particle with known velocity and mass will really impact on the airway wall in a certain bifurcation however, simply because the streamlines of the air can not be assessed properly and the particle may have different positions in the airway duct when entering a bifurcation. Besides, there exists uncertainty about the precise shape and cross section of the airway duct, the angle of bifurcation, the presence, size and shape of local flow constrictions, the effect of back flows, etc. A simplification of the situation can be made by considering only the throat (as the port of entry to the respiratory tract), which has a nearly 90 degrees bend and average radius (R) of about 20 mm. Using the centrifugal force (FC) acting on a particle in curved motion, instead of the stopping distance:

FC = m.UT

2/R [1.3] where UT is the tangential velocity (equals particle velocity), it can be calculated that

the approximate particle diameter for which the ratio of FC to FD equals 1 (when particles are inhaled at 60 l/min) is 5 micron (for particles with a density of 1500 kg/m3). This means that particles > 5 µm have a high collision probability in the throat, whereas particles < 5 µm may be transported into the trachea. On the basis of equation 1.3, it can be concluded that the FC increases with increasing particle density, particle diameter (both increasing particle mass) and velocity, resulting in an increase of the ratio of FC to FD and thus, in a greater chance of being captured by inertial deposition, which actually decreases the cut-off point of the throat.

The forces discussed above relate to different deposition mechanisms. The action of

the force of gravity results in a stationary particle settling velocity (UTS), which could lead to making contact with the wall of an airway duct by sedimentation. A high stopping force (relative to the drag force) may result in capturing of the particle by inertial impaction. In the upper airways, where air (and particle) velocity is still high (Paragraph 1.2.3), inertial

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impaction is dominant. As the air velocity decreases (with increasing airway number), the ratio of settling velocity to air transport velocity increases and sedimentation becomes more important. In the deep lung, into which no large particles can enter and where the air practically stands still, diffusion by Brownian motion gains importance for particles in the submicron range. The orders of magnitude for the particle velocities by air transport, sedimentation and diffusion in different airway generations are given in Figure 1.4 for spherical particles (three different diameters) with a density of 1500 kg/m3, inhaled at 60 l/min. The Brownian motion is represented by a linear velocity, which in fact is incorrect, as this motion randomly changes in all directions.

Figure 1.4. Particle velocity in different airway generations of the human respiratory tract, as the result of air transport, sedimentation and diffusion. Calculations are made on the basis of the Weibel model.

Figure 1.4 makes clear that only for relatively large particles (10 µm), the terminal settling (sedimentation) velocity can exceed the air transport velocity. This occurs only in the respiratory zone however, where such large particles do not enter, because they have already been removed from the air stream by inertial impaction in the upper airway tract. On the basis of the Weibel model, it can be calculated that the distance to travel for a particle from generation 15 to 23 in a typical human lung is only 0.861 x 10-2 m (Martonen et al., 1995). By making use of the velocities for air transport in Fig. 1.4, it can also be computed that it takes only 1.3 s for a particle to travel this distance, or 2.6 s to return to generation 15 at a constant flow rate of 60 l/min without delay in the alveoli. For the same airway generations, the diameter decreases from 0.65 x 10-3 (generation 15) to 0.41 x 10-3 m (generation 23). Meaning that the mean distance to travel for a particle from the central axis of the airway duct to its bottom wall is only 0.27 x 10-3 m. At the particle velocities for sedimentation and diffusion given in Fig. 1.4, it takes for a 1 micron particle 5.0 s by sedimentation, respectively 45 s by diffusion to complete this distance. For comparison, a 0.1 micron particle requires 9.0 s by diffusion (assuming linear movement). All times are longer than the 2.6 s required for transport from generation 15 to 23, suggesting that only particles that are already near the (bottom) wall of an airway duct do have a chance of being captured by sedimentation (or diffusion).

The calculations given above only indicate the orders of magnitude, because the flow rate (air transport velocity) varies during inhalation (Paragraph 1.2.2), and the real residence time will be somewhat longer (e.g. due to a short period of breath hold between inhalation and

0,000001

0,00001

0,0001

0,001

0,01

0,1

1

10

0 5 10 15 20 25airway generation

velo

city

(m/s

)

air transportsediment 10 micsediment 1 micsediment 0.1 micdiffusion 1 micdiffusion 0.1 mic

18

exhalation), which increases the residence time for small particles in the peripheral lung. On the other hand, local turbulences (e.g. from back flows) may effectively remove small particles from a position near the wall of an airway duct, whereas diffusion does not result in a linear particle movement. As a result, sedimentation and diffusion have a rather poor collection efficiency, as will be discussed in Paragraph 1.2.6. The simplified calculations suggest that the deposition efficiency by diffusion and sedimentation can be increased by reducing the air transport velocity (inspiratory flow rate) or by a breath hold between inhalation and expiration. Both increase the residence time in the alveolar region. An increase in the air transport velocity also changes the ratio of the stopping force to the drag force, or to the force of gravity respectively. This shifts the cut-off of aerosol particles of all sizes to airway generations with lower numbers (airways with larger diameters). The many attempts to model the airway deposition of aerosol particles are based on similar considerations as given above (e.g. Gerrity, 1990; Martonen, 1993; Martonen and Katz, 1993; Clark and Egan, 1994; Schulz, 1998). Such computations yield valuable information about the preferable aerodynamic diameter of aerosol particles, which is the subject of Paragraph 1.2.7. 1.2.5. Particle characterisation of inhalation aerosols The preferable size distribution of an aerosol for inhalation is strongly related to the site of deposition and the inhalation manoeuvre, as can be concluded from the previous paragraphs. Because particles of different shape and density from different aerosol generation principles (wet and dry) are used in inhalation, the term aerodynamic diameter has been introduced. By definition, the aerodynamic diameter (dA) of a particle is the diameter of a sphere with unit density (ρP = 1), having the same terminal settling velocity (UTS) in still air as the particle in consideration (Hinds, 1982). Other parameters used to describe irregular particles, like the Stokes diameter (dS), the equivalent sphere diameter (dE) and the dynamic shape factor (χ), standardise for shape only. The Stokes diameter is the diameter of a sphere with the same density and same settling velocity as the irregular particle; the equivalent sphere diameter represents the diameter of a sphere having the same volume, and the dynamic shape factor is the ratio of the resistance force acting on a non-spherical particle to the resistance force acting on a sphere having the same volume and velocity. By expressing the terminal settling velocity (UTS) in terms of these different parameters, correlations can be found, as will be explained more in detail in Chapter 2. The aerodynamic diameter enables prediction of the aerodynamic behaviour of particles of different shape and density in terms of stopping distances and settling velocities, and applies (consequently) for inertial deposition and sedimentation.

Aerosol particles do not only vary in shape and density, but also in size. Only few aerosol generators are capable of producing monodisperse particles (Paragraph 2.1): all other systems generate aerosols within the approximate size range between 0.1 and 10 µm (carrier particles in adhesive dry powder formulations excluded). Methods to describe size distributions are well known from literature (e.g. Allen, 1975; Gerrity, 1990). Mostly, particle size distributions are log-normal, for which type of distributions the geometric mean diameter (gmd) and the geometric standard deviation (GSD) are frequently used as the characteristic parameters. Aerosols from dry powder inhalers are not log-normal however, because they are nearly always a mixture of primary and secondary particles. Agglomerates in the aerosol are the reason for a tail-off at the side of the larger diameters. Therefore, the mass median diameter (mmd: the diameter above which 50% of the aerosol mass is contained) is a better parameter, although the size fraction for which mmd is calculated should be defined also. For inhalation drugs, the mass median aerodynamic diameter (mmad) is the most frequently used parameter. When laser diffraction technique is used as sizing principle, not a mass median diameter, but a volume median diameter (vmd) is obtained (calculation on the basis of

19

spherical particles). Only when all particles have the same density, which mostly is not true for aerosols from dpi’s because of the mixture of primary particles and agglomerates, the (cumulative) volume distribution curve as function of the diameter is the same as the mass distribution curve (and vmd equals mmd). Such a volume distribution curve from laser diffraction analysis is described with X50 for the median diameter and X10 plus X90 to define the span of the range (ISO 13320-1, 1999). Laser diffraction analysis neither yields aerodynamic diameters. The consequences of that will be discussed in the Chapters 2 and 3.

1.2.6. Modelling of aerosol particle deposition

The preferable range for the aerodynamic size distribution of an inhalation drug can be estimated either from mathematical models predicting lung deposition, or from in vivo deposition (scintigraphic) studies. Also clinical effect studies and pharmacokinetic methods are used. Computations are often on the basis of stochastic frameworks, which make use of deposition probability equations for inertial impaction, sedimentation and diffusion, based on the parameters that are relevant to deposition (Paragraphs 1.2.3 and 1.2.4). In the calculations made by Gerrity (1990), the expression for inertial impaction contains the Stokes number (characterising the collection efficiency: Hinds, 1982), which is a function of the particle diameter, particle velocity and the airway radius. The expression for inertial impaction also contains the angle of bifurcation between mother and daughter airway. The expression for sedimentation includes the terminal settling velocity of the particle (which is a function of particle diameter, shape and density), the residence time of the particle in the airway, the inclination angle of the airway to the horizontal and the airway radius. And the diffusion probability equation contains the residence time in the airway, next to the particle diameter, airway radius and the Cunningham slip flow correction factor.

The numerical deposition probability values obtained from such computations are arguable, because of the many assumptions and simplifications, but the influence of relevant parameters and the effects of changes therein can well be studied. For instance, Gerrity (1990) calculated that the fractional deposition of 5 micron particles at a low inspiratory flow rate of 30 l/min has two peaks: one by inertial impaction in airway generation 4, which has a high air velocity (Fig. 1.5), and one in generation 18 as the result of sedimentation. Increasing the particle size under the same inspiratory conditions, decreases the sedimentation probability around generation 18 and increases the inertial deposition probability around generation 4. Decreasing the particle size has the opposite effect for the inertial impaction. Sedimentation however (around generation 18), seems to have a maximum for particles between 3 and 5 and decreases also with decreasing particle diameter as a result of decreasing settling velocity. The effect of increasing the inspiratory flow rate from only 12 to 60 l/min is quite dramatic for 5 micron particles. At the lower flow rate, the deposition probability for sedimentation in the respiratory region with a peak in generation 18 is nearly five times as high as the probability for inertial impaction in the conducting zone, having only a minor peak in generation 4. Already at 40 l/min, both probability peaks are of the same size and at 60 l/min the deposition by inertial impaction around airway 4 is twice as high as that by sedimentation in airway 18. This strong shift in deposition towards higher airways with increasing flow rate is the basis for the expectation that an increasing fpf with increasing flow rate (as from a dpi) results in a more flow rate independent therapy than a constant fpf at all flow rates. This expectation is strongly supported by Clark and Egan (1994) who computed the shift in deposition in different lung compartments for particles in the size range between 0.5 and 12 µm as a result of increasing the flow rate from 7.5 to 120 l/min. Such calculations also suggest that 5 micron particles can not penetrate the respiratory zone effectively at flow rates higher than 40 l/min.

20

As may be expected from the relevance of the residence time in the airway to the deposition probability expressions for sedimentation and diffusion, there is a great effect of breath hold on the deposition in the respiratory zone. Martonen and Katz (1993) concluded that lung deposition is even more sensitive to particle residence time than to inspiratory flow rate. They also showed that the inhaled volume has a relatively minor effect on the deposition of 5 micron particles throughout the whole respiratory tract. There is only a faint shift from inertial impaction in the upper tract to peripheral deposition when the inhaled volume is increased from 0.5 to 1.5 l (at 40 l/min), which can be explained by increasing the residence time. Such computations should preferably be extended to the whole range of exerted inspiratory flow parameters, which for the inspiratory flow rate has a span ranging from 6 to more than 150 l/min (Paragraph 1.2.2).

The fractional deposition probabilities can be used to predict the regional and total deposition efficiencies in the respiratory tract as function of particle diameter. Generally, deposition efficiencies presented in literature all show the same trends. Total lung deposition fraction has its maximum value (of 1) for particles larger than 10 micron. It decreases with decreasing particle diameter to a minimum of 0.2 (or even less) for particles in the size range between 0.5 and 1 µm which are deposited mainly by sedimentation. When the particle diameter is further decreased to values < 0.5 µm, an increase is computed again mainly as the result of diffusion (according to some models nearly up to 1 again, e.g. Martonen and Katz, 1993; Schulz, 1998). Results of regional deposition computations at relatively low inspiratory flow rates (between 7.5 and 45 l/min) suggest that tracheobronchial deposition (in the airway generations 0-16) decreases from 1 to practically 0 when the particle diameter is decreased from 10 to 1 micron (Martonen and Katz, 1993). Calculated pulmonary deposition fractions (in airway generations 17-23) show bimodal curves. They increase from zero for 10 µm particles to a peak value of 0.4 to 0.6 for 4 µm particles (under the same inspiratory conditions). A further reduction in particle size results first in a decrease of the pulmonary deposition fraction to less than 0.2 (for 0.5 µm particles) followed by an increase again for particles smaller than 0.5 micron. The peak value in the size range around 0.02 µm can be even higher (0.6 to 0.8) than that around 4 micron and contributes to the previously mentioned increase in total lung deposition fraction for submicron particles. For particles < 0.01 µm, pronounced filtering in the upstream tracheobronchial compartment is described, for which no good explanation is given however. The computed data presented by Martonen and Katz (1993) seem to be in fairly good agreement with experimental data (studies performed by others), except for particles larger than 10 µm or those smaller than 0.1 micron. 1.2.7. The desired aerodynamic size distribution of inhalation drugs

Computations made by for instance Gerrity (1990) and Martonen and Katz (1993) indicate that effective pulmonary deposition (airway generations 17-23) at relatively low flow rates requires particles in a narrow size range around either 4 or 0.1 micron. Because submicron particles are difficult to produce, particularly with dry powder inhalers, particles in the size range between 1 and 5 micron are generally the objective. The optimum within this range can not simply be derived from the deposition calculations however. Even when the precise location of receptor areas and the effect of the inspiratory flow manoeuvre on deposition behaviour are taken into account, many other parameters may influence this optimum. For instance, also lung anatomy and morphology aspects, such as the non-uniformity of ventilation distribution as the result of airway obstructions (Martonen, 1993), or the reduced airway dimensions of infants have to be considered. Brown et al. (2001) compared published data regarding the relation between ventilation distribution and particle deposition and concluded that the results are inconsistent. Both increasing and decreasing particle deposition in poorly ventilated parts of the lung have been reported, for particles in

21

the size range between 0.5 and 1.2 µm. Increased deposition was attributed to an increased residence time. Other studies yielded no association between ventilation distribution and particle deposition, apparently due to inappropriate measuring techniques. Brown et al. (2001), using iron oxide particles with a mass median aerodynamic diameter of 5 micron, reported different associations for healthy subjects and CF patients. Within CF patients a substantial particle deposition in the tracheobronchial airways (generations 0-16) of poorly ventilated regions was observed, but within the pulmonary region (generations 17-23) better ventilated regions exhibited higher particle deposition. For healthy subjects, tracheobronchial deposition appeared to increase dramatically with ventilation. The results seem logical considering the previously discussed deposition mechanisms and the effect of ventilatory parameters thereon. In obstructed airways, diameters are reduced and air velocities are increased, which contribute to a higher probability of inertial impaction in the tracheobronchial lumen. Effective removal of particles in the tracheobronchial tree reduces the number of particles entering the pulmonary region, and by that the particle deposition in this (pulmonary) region. In healthy lungs, an increase in the ventilation increases the air velocity and by that, the inertial impaction.

The mass median diameters for drugs in dry powder formulations are frequently selected between 1 and 2 micron on the basis of the expectation that not all drug particles are released as primary entities from the dpi. Moreover, drug particle size in dry powder formulations affects the efficacy with which particles can be detached from carrier crystals during inhalation (Chapter 4). By slightly increasing the primary drug particle size even within this narrow range, the detached mass fraction of drug can be increased substantially. This is particularly relevant to inhalers with poor de-agglomeration efficiency yielding low fine particle fractions (fpf’s) (Dickhoff et al., 2002). To compensate for the minor shift in higher airway deposition when the particle diameter is increased, a lower inspiratory flow rate could be applied. It has furthermore been recognized that effective deposition on the site of highest receptor density, does not always give best clinical effect. The already mentioned reduction in bronchodilatation with increasing deposition in the alveoli as the result of lack of smooth muscle in this region could be an example for this (Chrystyn, 1997). Additionally, alveolar deposition may also result in systemic adverse drug reactions (e.g. Weda et al., 2004). Therefore, the aerodynamic diameter particularly for dry powder inhalers has to be adjusted to:

a. the precise target area, taking account of the patient’s lung morphology and anatomy, b. the presumed inhalation manoeuvre, c. the efficiency of powder de-agglomeration during inhalation, and d. avoid side effects.

Most experimental data seem to support the results from deposition modelling. But

neglecting the effect of one (or more) of the parameters a-d mentioned above (or additional effects) may lead to false conclusions about the preferable aerodynamic diameter. For instance, Zanen et al. (1994) observed that monodisperse particles of 2.8 µm (aerodynamic diameter) for salbutamol provide a greater improvement in lung function than either 1.5 or 5 micron particles in the same dose. The observation was made for an inhalation manoeuvre having a peak flow rate between 40 and 60 l/min and a breath hold period of 10 s. Considering the previously discussed effect of the inhalation manoeuvre on the deposition behaviour (e.g. Gerrity, 1990), it could well be that without breath hold, the 5 micron particles would have given greatest improvement in lung function at this relatively low flow rate (reduced sedimentation for the three micron particles). On the other hand, the 1.5 µm particles might have been optimal at flow rates higher than 60 l/min with breath hold (because of increased inertial impaction for the 3 and 5 µm particles in the upper airways). Leach et al

22

(1998) concluded that half the beclomethasone dipropionate (bdp) dose from a HFA mdi is necessary to obtain the same efficacy as that from a CFC mdi. They explained this by a difference in mmad, being 3.5 to 4 µm for the CFC versus 1.1 µm for the HFA device (Leach, 1998). As a result, the fraction of drug < 4.7 micron collected in an Andersen cascade impactor was 60% for the HFA mdi versus 30% for the CFC inhaler. The conclusion is quite remarkable, since the calculated lung depositions differed by a factor 10 between both devices (55-60% throughout the airways from the HFA mdi versus 4-7% exclusively in the central airways for the CFC mdi) whereas steroid receptors are located throughout the lungs (as confirmed by the authors). They explained the discrepancy with reasons related to the dose-response curve, which therefore is yet another parameter to consider when evaluating the effect of the aerodynamic diameter. This conclusion is supported by Weda et al. (2004) who found that an increase in the salbutamol fpf (at the same inspiratory flow manoeuvre) does not improve the lung function, but it does increase the adverse side effects (decrease of serum potassium level). Ram et al. (2001) reviewed 84 different studies in which clinical effectiveness of β2-agonists was investigated. They found no evidence for differences in effectiveness between the different types of inhalers used in these studies, in spite of differences in lung deposition efficiencies, which are partly related to differences in aerodynamic diameter for the drug. The conclusions of Leach et al. may also have been shaded by uncontrolled aspects, such as the differences in plume velocity, as reported by Gabrio et al. (1999). Not to mention that conclusions about lung deposition of drug particles on the basis of Technetium-99 radionuclide distribution are rather arguable, as the penetration index (ratio of peripheral to central lung deposition) may be different for the radionuclide and the drug, whereas the fpf may be changed by the labelling technique (Bondesson et al., 2002).

Summarising, it may be expected that the optimal aerodynamic diameter for dry

powder inhalers, that are operated at inspiratory peak flow rates between 30 and 150 l/min, lies somewhere in the range between 1 and 5 µm. It depends on the type of drug (related to a specific target area), the inhaler resistance (defining the range of attainable flow rates) and its powder de-agglomeration efficiency, as well as on the possible side effects and patient’s lung morphology where to put the mass median diameter within this range and what size distribution to strive for. Distribution throughout the lungs, as has been recommended for inhaled corticosteroids, may benefit from a somewhat wider distribution, whereas targeting at the pulmonary region is better obtained with aerosols having a small mmad and low GSD. If the distribution is too wide, and the target area is primarily in the central and small airways (as for β2-agonists), adverse systemic side effects may be obtained from particles that are deposited in the throat or the alveolar region, from where they can enter more easily the systemic circulation. A better targeting at a particular deposition site and reducing adverse side effects are two arguments to use monodisperse instead of polydisperse aerosols (e.g. Zanen et al., 1994; Brand et al., 2000; Brown et al., 2001). However, as may be clear from the deposition calculations, this makes only sense when the inspiratory manoeuvre can be controlled within quite narrow limits for the flow rate and the inhaled volume, as reported by Brand et al. (2000). And even then, the preferable particle size may differ considerable for different patient groups and age categories. Without breath control, or in case of extreme pulmonary morphology and/or anatomy (like in small children), deposition may substantially be outside the target area. Therefore, a somewhat wider size range would give increased assurance of therapeutic efficacy. Also the deposition efficiency of large porous particles may be reduced compared with solid particles having the same aerodynamic diameter. Because of their increased geometric diameter, diffusion deposition may be much lower.

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2. Medical aerosol generators (inhalation devices) The task of a medical aerosol generator (inhaler; inhalation device) is to produce an

aerosol with the desired aerodynamic particle size distribution from a single drug dose under inspiratory circumstances that enable deposition of a high fraction of that dose in de target area of the lungs. Classification of medical aerosol generators is upon working principle, and includes nebulisers, metered dose inhalers and dry powder inhalers. Nebulisers and metered dose inhalers are not the subject of this thesis. For that reason only brief reviews of these types of medical aerosol generators are given. 2.1. Nebulisers Classic nebuliser therapy is with jet or ultrasonic nebulisers. Their principles of operation have been described elsewhere (e.g. Newman, 1989; Hess, 2000; Le Brun et al., 2000). A strong disadvantage of the classic type of nebuliser is the continuous aerosol production which conflicts with the alternating flow direction during the inhalation manoeuvre. A large portion of the aerosol produced during phases of exhalation is wasted to the environment (Coates et al., 1998), or returned to the reservoir by rain out (Ho et al., 2001). Denyer et al. (1997) measured the ratios of inhalation to exhalation time for different patient groups, and found mean values greater than 1 for all groups, the ratio being highest for asthmatics (1.64) and lowest for healthy subjects (1.29). Partly as a result of the wastage, and partly because of a less favourable size distribution, the effectiveness in lung deposition of most classic nebulisers is confined to less than 10% (Newman 1989; Coates et al., 1998). A larger fraction of the produced aerosol from jet nebulisers is administered to the lungs with so-called open vent systems (e.g. Sidestream, Medic-Aid), which continuously draw air into the nebuliser cup to increase the aerosol output rate (O’Callaghan and Barry, 1997). Breath assisted open vents (e.g. Pari LC Plus, Pari) open only during the phase of inhalation (Ho et al., 2001). Vented nebulisers are not as effective as systems that actually stop the aerosol production during exhalation however. Such breath actuated nebulisers (BAN) exist in two different types: mechanically intermitting devices (e.g. AeroEclipse, Trudell Medical) and the electronically operating systems (e.g. HaloLite, Medic-Aid; Denyer and Nikander, 1998). The first type of nebuliser has a slight overpressure in the nebuliser cup during exhalation that depresses the drug solution in the capillary towards the nozzle head. The HaloLite makes use of so-called adaptive aerosol delivery (AAD): on the basis of recorded breath patterns the jet nebuliser is only switched on during the first 50% of the inhalation cycle. Breath assisted open vents increase the portion of the aerosol that is inspired (compared to non-vented systems) to 65%, whereas AAD systems have an efficiency of nearly 100% in this respect. Many different parameters influence the size distribution in the aerosol from jet and ultrasonic nebulisers. They partly relate to the properties of the drug solution, like (changes in) temperature, concentration, surface tension and viscosity (Steckel and Eskander, 2003). Particularly the surface tension appears to be of great influence (McCallion et al., 1996). But also device parameters (e.g. jet flow; Hess, 2000; Standaert et al., 2001) and the inspiratory flow rate can affect the droplet size distribution from a jet nebuliser, although this depends on the jet flow too (de Boer et al., 2003a). All these different parameters that influence the aerosol properties plea for use of validated combinations of nebuliser, compressor and drug solution. Further means to improve lung deposition from nebulisers may come from controlled inhalation, as described by Brand et al. (2000). Even greater improvements, particularly in regard of precision in drug targeting, are expected from monodisperse aerosols. Only three aerosol generation principles are (theoretically) able to generate droplets that are all exactly of the same size: electrospray (Marijnissen, 2002), Rayleigh break-up (Wissink, 2003) and the spinning top generator (Cheah and Davies, 1984). Particularly the principle of

24

Rayleigh break-up has good perspectives, because the whole operating system can be simple, cheap and small. Development of this principle has become possible thanks to developments in membrane (chip) technology, which have also been the basis of the development of so-called vibrating membranes, as applied in the e-flow (Pari, based on TouchSpray technology, ODEM), Omron NE-series (Omron) and AeroNeb (AeroGen) (eg. Dolovich, 1999; Conway, 2002). Such membranes can produce aerosols with much more uniform droplet sizes than conventional jet and ultrasonic nebulisers, although bimodal size distributions may be obtained as a result of interference between the Rayleigh break-up mechanism of the jet and the superimposed break-up by high the frequency vibration of the membrane. An advantage is that the output rates of membrane nebulisers are much higher than those of classic jet and ultrasonic devices. A final principle, though not new as such (Perry and Chilton, 1973), to be mentioned is that of impinging jets, as applied by Boehringer in the Respimat (Newman et al., 1998). 2.2. Metered dose inhalers (mdi’s) The working principle of metered dose inhalers and their performance have been described extensively by others (e.g. Newman, 1983). The first generation mdi’s used chlorofluorocarbon (CFC) propellants which had the disadvantage of producing cold plumes with high initial droplet velocities (Gabrio et al., 1999). This resulted in the so-called ‘cold-freon effect’ (Newman and Clarke, 1993) and substantial oropharyngeal deposition of the drug, frequently > 90% of the dose (Leach et al., 1998). Poor stability of drug suspensions is another error source in mdi therapy (Jackson, 1995). Furthermore, patients need to have a good co-ordination between inhalation and dose release to use an mdi effectively (e.g. Arppe et al., 1994). Different solutions have been found to overcome these problems, like the introduction of extension tubes and spacers (Bisgaard et al., 2002), and breath actuated dose release, as for instance with the (3M) Autohaler (Fergusson et al., 1991) and (IVAX) Easi-Breathe. Also mdi’s with reduced aerosol velocity (< 2 m/s, versus 30-50 m/s for a conventional mdi) have been proposed (Newman and Clarke, 1993). Spacers rather decrease the throat deposition than having great potential for increasing lung deposition. Substantial losses in a spacer may occur, due to impaction, sedimentation and electrostatic attraction (Bisgaard et al., 2002). A reduction of the electrostatic losses has been obtained with a metal spacer (Bisgaard et al., 1995). However, the most important argument against the use of ‘classic’ mdi’s is the contribution of CFC’s to the depletion of the ozone layer (Newman, 1990). This has resulted in a phase out of the production of CFC’s by January 1, 1996, as agreed to in the Montreal protocol in 1987 (also Copenhagen Amendment, 1992: Partridge and Woodcock, 2002). Hydrofluoroalkanes (HFA’s) 134a and 227 appear to be good substitutes and are safe in use (e.g. Thompson et al., 1998), although reformulation has been necessary in a number of cases (Vervaet and Byron, 1999). HFA mdi’s have a lower plume temperature and particle exit velocity (Gabrio et al., 1999) and frequently finer particles in the aerosol, which increase the lung deposition (e.g. Leach, 1998). Current developments in mdi therapy focus particularly on different add-on devices, such as dose counters, training devices, monitoring devices and lever systems to ease firing of a dose. 2.3. Dry powder inhalers (dpi’s)

Dry powder inhalers produce aerosols in which the drug is in the solid state. This has certain advantages regarding drug stability, ease of handling and the range of possible dose weights to be administered. Because the drug is processed into a powder formulation, decomposition, separation (flocculation and sedimentation) and microbiological

25

contamination hazards have been eliminated, or at least strongly been reduced compared to wet formulations. In most cases, the inspiratory air flow through the dpi delivers the energy for aerosol generation, meaning that dpi’s can be small, relatively cheap and portable. In contrast with nebulisers, high doses can be inhaled within one single breath manoeuvre, whereas preparation of drug solutions and cleaning (sterilisation) of the device after use are not necessary. It has often been considered as a disadvantage that some dpi’s have a breath controlled fine particle output (Paragraph 2.3.7). It may be questioned however, whether this is really a drawback, considering the shift in deposition area with a change in the inspiratory flow rate, which effect can be partly compensated by a change in the fine particle fraction (Chapter 10).

Although several patents for dry powder inhalation are known from the beginning of the 20th century (e.g. Di Christina, 1927; Priestley, 1952), the first portable inhaler on the market (Fisons Spinhaler) was not introduced until 1970 (Ashurst et al., 2000). Early dry powder inhalers were all single-dose devices with hard gelatin capsules as dose containers, like the Glaxo Rotahaler (Pover et al., 1982) and the ISF inhaler (Salvatore, 1976). A new dimension to dry powder inhalation was given by the multi-dose Turbuhaler (Wetterlin, 1988), which challenged many companies to copy this concept. The concept has a multi-dose powder reservoir with a metering system for the isolation of single doses from the supply chamber, that has to be operated by the patient. Another unique feature of the Turbuhaler was the omission of excipient in the drug formulation (spherical pellets), although recent developments with the Turbuhaler include that micronised lactose is used to standardise the measured dose of the drug-lactose mixture to the same weight for different drugs (and drug combinations). In the mid-1990s a new trend was started with the introduction of multiple unit dose inhalers like the Glaxo Diskhaler and Diskus (Sumby et al., 1993; Brindley et al., 1995) and Dura Spiros (Dura, 1997), having a large number of unit doses in blisters on a disk or strip. On the basis of these historical developments, dpi’s are often classified into single-dose, multi-dose or multiple-unit-dose (e.g. Srichana et al., 1998), although many other, and even more relevant differences in respect of performance exist, such as the type of formulation, the inhaler’s resistance to air flow, the type (and size) of the de-agglomeration forces generated during inhalation. Also a classification into breath operated dpi’s and devices relying on auxiliary energy would be more logical.

2.3.1. Basic design and functional parts of a dpi

From design viewpoint, a classic dry powder inhaler consists of certain primary and secondary functional parts. In spite of the differences mentioned in the previous paragraph, the primary parts are the same for all types of currently marketed devices (Fig. 1.5). They determine the performance of the inhaler with respect to fine particle fraction and lung deposition, and include a powder formulation with the drug in the correct aerodynamic particle size distribution, a dose mechanism, containing (or measuring) a single drug dose, a powder de-agglomeration principle, redispensing the powder into the inhaled air stream, and the inhaler’s mouthpiece. Obviously, the parts are assembled into a suitable housing which includes the necessary (mechanical) means to get access to a single pre-measured dose or to operate the dose measuring system. For each of these primary parts, different solutions and designs are possible, which contribute to the variety of currently available dry powder inhalers. The secondary elements are for safety, ease of handling, signalling to the patient, moisture protection of the drug formulation, etc.

26

Figure 1.5. Primary functional design parts of a classic dry powder inhaler.

Due to their size distribution (approximately 1-5 micron: see Paragraph 1.2.7), inhalation drugs are extremely co- and adhesive. They tend to stick together (agglomerate) and to (inhaler, capsule or blister) walls with which they make contact, primarily by Van der Waals forces. Doses are given in a wide range, varying between only a few micrograms (e.g. formoterol fumarate) to several tens, or even hundreds of milligrams (e.g. colistin), and particularly the lower doses can not be measured in a reproducible way without using diluent excipients. The excipient can either be micronised to (approximately) the same size distribution as the drug, or consist of larger crystals (or agglomerates) that act as a carrier for the drug.

Figure 1.6. Scanning electron micrographs of an adhesive mixture (left), a nucleus agglomerate (middle) and spherical pellets (right).

Three different types of formulation exists (adhesive mixtures, nucleus agglomerates and spherical pellets), as depicted in Fig. 1.6. Adhesive mixtures (containing large carrier crystals) and spherical pellets are discussed more in detail in Paragraph 2.3.2. Pellet size and carrier size distribution are selected to obtain good flow properties, which is a requisite for the

mouthpiece

de-agglomeration principle

powder formulation

dose (measuring) system

or preloaded dosecompartments and canisters (e.g. capsules)

mouthpiece

de-agglomeration principle

powder formulation

dose (measuring) system

or preloaded dosecompartments and canisters (e.g. capsules)

27

desired improvement in processing and dose measuring. When the drug is distributed in multiparticulate layers around carrier particles, nucleus agglomerates are formed (Sarlikiotis and de Boer, 2001). This type of formulation is not frequently used in dry powder inhalers however, and for that reason it will not be discussed in this thesis. More recently, solid dispersions of drugs with sugars like trehalose and inulin are prepared for inhalation, particularly when these sugars can maintain (in the dry state) the ternary and quaternary structures of large molecular drugs (e.g. peptides and proteins, which depend for their biological activity on these structures), presumably by water replacement (e.g. Chan et al., 2004). Also recently, various particle engineering processes are applied to improve drug particle dispersion during inhalation. Some examples will be mentioned in Paragraph 2.3.8.

From the viewpoint of lung deposition, carrier formulations are to be preferred. Micronised lactose particles in spherical pellets generally have the same size distribution as the drug and are able to enter the deep lung, which (under circumstances) may be an extra burden to the patient. This, in spite of the fact that it has been reported that the inhalation of a high dose of 100 mg inhalation lactose as dry powder (in carrier quality as used in the Diskhaler formulation) does not necessarily result in a fall in FEV1 in asthmatic subjects (Thoren et al., 2001). Srichana et al. (1998) investigated the lactose deposition in a twin impinger for three different formulations with different amounts of lactose, having different degrees of fineness, and concluded that the amount of carrier lactose deposited on the second stage (fraction < 6.4 µm) may be of the same order of magnitude (400 µg) as the drug dose (Rotacaps formulation). Their findings are supported by Karhu et al. (2000) who measured only 2.5 to 3% lung deposition from different adhesive powder mixtures, containing different qualities of 99m Tc-labelled lactose carrier. Carrier lactose is either deposited in the throat, which is unwanted because of possible adverse local side effects from (still) adhering drug particles, or in the mouth, depending on the inhaler’s discharge flow pattern. Mouth deposition is frequently mentioned as an advantage, because it confirms dose delivery, but can not be sensed by everyone in amounts smaller than 10 mg (Higham et al., 1995) and it may also annoy the patient.

As already mentioned in Paragraph 2.3, single-dose containers (blisters or capsules) with pre-weighed quantities of inhalation powder and multi-dose reservoir systems with metering principles are both applied as dose systems for dpi’s. Many different capsule and blister piercing or opening mechanisms have been developed to gain access to the powder during inhalation (Fig. 1.7A). In a majority of the principles, discharge of the capsule (blister) and powder dispersion into the inspiratory air stream occur simultaneously. Only some recently developed capsule and blister inhalers apply additional powder de-agglomeration means to increase the fine particle fraction, e.g. Aventis Eclipse and Dura (Elan) Spiros (Han et al., 2002). Multi-dose reservoirs may have different designs too, as shown in Fig. 1.7B for a selection of examples. Mostly, as for the Turbuhaler and Novolizer in Fig. 1.7B, the isolation of individual doses is by volumetric measurement. Small holes in a rotating disk (Turbuhaler), or a cavity in a slide (Novolizer) are filled with powder by gravity directly from the powder reservoir (Novolizer) or by scraper mechanisms from special chambers (Turbuhaler) respectively. In some special cases, forced metering is applied, e.g. by conducting compressed air through the powder bed, as for the (IVAX) Airmax (Keating and Faulds, 2002). Before inhalation, the filled compartment is transported to a position where it becomes in connection with a powder channel. During inhalation, the inspiratory air flow entrains the powder from the compartment towards the de-agglomeration principle. An exception to this general concept is the Jethaler (Ratiopharm, Fig. 1.7B) which has a ring compact of the drug-excipient mixture, from which small amounts are grated with a scraper disk during inhalation (Newman et al., 2002). The concept is the same as that of the Aventis Ultrahaler (Pitcairn et al., 1997), and appears to have a poor dose reproducibility (de Boer et al., 2004).

28

A

B Figure 1.7. Different capsule emptying principles (A) and different multi-dose reservoirs with metering systems (B) applied in dry powder inhalers.

De-agglomeration principles are meant to break-up spherical pellets into primary drug entities, or to detach drug particles from the carrier crystals in adhesive mixtures (or nucleus agglomerates) during inhalation. The objective is to generate an aerosol with drug particles in the approximate aerodynamic size range between 1 and 5 micron which can enter the target area for deposition. Many different principles of powder dispersion have found application in dry powder inhalers, as will be discussed in Paragraph 2.3.3.1.

The mouthpiece of the inhaler (Fig. 1.5) may not seem to be a primary design element, but it can be used to add certain functionalities to the inhaler. The mouthpiece may have bypass channels to control the inhaler resistance to air flow, and the bypass flows can be arranged in such a way, that they constitute a co-axial sheath of clean air around the aerosol cloud. This reduces the deposition of drug in the mouth from back flows, as is explained in Chapter 9 of this thesis. Mouthpiece design is also relevant to the shape of the released aerosol cloud. A strongly diverging cloud increases the mouth deposition, although it can be favourable for adhesive mixtures when deposition of the carrier crystals can be directed to the mouth instead of the throat (to reduce local side effects: Chapter 9).

Secondary inhaler parts are of lower relevance to this thesis. They may include means to prevent double dosing or inhalation without having activated a dose (safety functions). They can also indicate the status of the inhaler (dose primed), confirm correct inhalation to the patient or indicate the number of doses taken or left (signalling functions).

Turbuhaler Novolizer JethalerTurbuhaler Novolizer Jethaler

FlowCaps

Rotahaler

HandiHaler

ISF inhaler (eg. Aerolizer)

FlowCaps

Rotahaler

HandiHaler

ISF inhaler (eg. Aerolizer)

29

Multi-dose reservoir inhalers may have desiccant compartments to give moisture protection to the drug formulation. And many inhalers have protective hoods or caps against pollution either of the inhaler with dust, or the patient’s clothes or handbag with residual drug formulation. A few representatives from the groups of single-dose (capsule), multi-dose reservoir and multiple-unit-dose (blister) inhalers are shown in Fig. 1.8. Figure 1.8A. Examples of capsule inhalers: Inhalator Ingelheim (left) and ISF-inhaler (right); the latter is also known as (Novartis) Aerolizer and (Ciba Geigy) Foradil inhaler.

Figure 1.8B. Examples of multi-dose (reservoir) inhalers: (AstraZeneca) Turbuhaler (left), (Viatris) Novolizer (middle) and (Ratiopharm) Jethaler (right).

30

Figure 1.8C. Examples of blister inhalers: (GSK) Diskhaler (left) and (GSK) Diskus (right). 2.3.2. The powder formulation

Of the powder formulations shown in Fig. 1.6, adhesive mixtures and spherical pellets are the most frequently used ones in dry powder inhalers. They will be discussed in the paragraphs 2.3.2.2 to 2.3.2.4. 2.3.2.1. Particle-particle interactions

The nature and strength of the interparticulate forces acting within a cohesive or adhesive powder determine the ease of relocating individual particles within the mixture during the mixing process, the intensity of segregation during handling of the powder, as well as the degree of de-agglomeration during inhalation. If solid bridge formation (chemical interaction) is excluded, likely bonding mechanisms are either due to moisture, to electrostatic charging or to molecular (atomic) interaction forces (Harnby, 1983). The relative magnitude of these bonding forces may differ quite considerably. Coulombic (electrostatic) forces on a molecular basis are by far the strongest of all physical interaction forces, sometimes exceeding the magnitude of covalent bonds (Hinds, 1982; Myers, 1991). Capillary forces are second in order of magnitude. Their mechanism of interaction and differences in appearance as function of the amount of water in relation to the powder mass are well known from granulation studies. Capillary forces are tenfold higher than molecular forces in the presence of absorbed liquid layers (with zero interparticulate distance) and hundredfold higher than molecular forces calculated at a distance of 3.3 nm (Lloyd, 1983). The rather relatively weak molecular forces are generally grouped as Van der Waals forces. They include different types of atomic and molecular interactions with different characteristics, theoretical bases and limitations. Two types are based on relatively straightforward electrostatic principles and are closely related to Coulombic forces (dipole-dipole and dipole-induced dipole interactions), whereas a third type has its origin in quantum mechanics (Myers, 1991). This third type is the most fundamental and universal type of force, generally the weakest, and yet often the most important contributor to total Van der Waals attraction. This force is known as the London-Van der Waals force, or London dispersion force. The dipole-dipole type of force is also termed orientational or Keesome-Van der Waals force, whereas the dipole-induced dipole

31

force is referred to as induced or Debye-Van der Waals force. Hydrogen bonding is often considered as a fourth type of Van der Waals attraction (Myers, 1991).

In order to describe the particle-particle interactions in a powder, the complex relationships for interactions between individual molecules have to be converted into interactions between large groups of molecules. In adhesive mixtures for inhalation, three different situations can be distinguished. Micronised drug particles are either attached to each other or to carrier crystals, or to both. Interactions between carrier particles are less relevant, because of their size, which normally excludes attachment of carrier particles to each other (because the separation forces, i.e. gravitational forces, are larger than the attraction forces). Drug-to-drug interaction is basically an interaction between two spheres of approximately the same size and the same chemical structure (cohesion). In contrast, drug-to-carrier interaction, due to the large difference in size between drug and carrier particles, is rather that between a sphere and a plate (infinite flat surface), each of them having different chemical structures (adhesion). A theoretical derivation for the free energy of attraction per unit area (∆G) for the situation of two hard, flat, non-polar infinite surfaces, separated by a distance x in vacuum, was made by Hamaker:

∆G = -AH/(12.π.x2) [1.4]

where AH is the Hamaker constant. The value of the Hamaker constant is a function of the number of atoms (or molecules) per unit volume (n) and a quantum mechanical constant (A’) which depends on the structure of the atom or the molecule: AH = A’.π.n2 (Myers, 1991). In normal practice, surfaces interact through a medium other than vacuum, for which an effective Hamaker constant can be approximated. Adhesion forces resulting from the free energy of attraction can be considered as the forces necessary to separate the surfaces from each other (Hiestand, 1985). For the model of a sphere attached to a plate (which is more relevant to the drug-to-carrier attachment in adhesive mixtures), the Van der Waals adhesion force (FVDW) is proportional to the diameter (d) of the sphere and inversely proportional to the square of the distance (x) between the plate and the sphere (Leuenberger et al., 1989): FVDW = (AH.d)/(12.x2) [1.5] For perfectly smooth spherical particles the separation distance is about 4 Å, but practically for particles in the size range between 1 and 10 µm, the distance is rather 20 Å (Schönert et al., 1996). According to Hiestand (1985), the force between two spheres with radius r is half as high as that between a sphere and a plate in equation 1.5: FVDW = (AH.r)/(12.x2) [1.6]

For two spheres with different radii (r1 and r2), the complex (or harmonic) radius (rC) can be used: rC = (2.r1.r2)/(r1 + r2) [1.7] The major problem in calculating Van der Waals interaction forces between particles is that of evaluating the Hamaker constant (Shaw, 1983). Hamaker evaluated the individual atomic polarisabilities and densities of the materials involved and assumed that total interaction is the sum of the interactions between all interparticulate atom pairs, which is essentially incorrect because the atomic attraction forces interact with each other. Although the error involved is not extreme, Lifshitz proposed an alternative (macroscopic) approach in

32

which the interacting particles and the intervening medium are treated as continuous phases (Moser and Sommer, 1977). Lifshitz calculations are complex and require solution of Maxwell’s equations. The order of magnitude of the Hamaker constant for single materials is between 1 to 10 x 10-20 J (Shaw, 1983; Li et al., 1996); the Lifshitz constant is considered to be in the range of 1 to 15 x 10-19 J (Otsuka et al., 1988). Different equations for the Van der Waals attraction force (FVDW) on the basis of Lifshitz’s constant are presented (e.g. Stewart, 1986; Otsuka et al., 1988; Schubert et al., 1990; Podczeck et al., 1997), which all show the same difference between the sphere-sphere and sphere-plate model (by a factor 2). According to Schubert et al. (1990):

FVDW = (h.ϖ.d)/(16.π.x2) for sphere-to-plate attraction [1.8] and FVDW = (h.ϖ.d)/(32.π.x2) for sphere-to-sphere attraction [1.9]

where h.ϖ is the Lifshitz constant and the same equation 1.7 (as for the radii) can be applied for the diameter (d) if the attraction is between two spheres of different size. The equations 1.5 to 1.9 suggest that Van der Waals attraction in adhesive mixtures for inhalation is well defined and understood. The contrary is true however. As will be discussed more in detail in Chapter 5, adhesion is not always that between the drug and pure lactose. Minor amounts of impurities may be present on the carrier surface, whereas surface rugosity (on a micronscale) may either result in multiple contact points or reduce the contact area of a single contact point, depending on the size of the surface irregularities relative to the diameter of the adhering drug particle. Besides, micronised drug particles often have a (more or less) irregular shape and their contact area with the surface of the carrier crystal may be larger than that for a perfect sphere. Otsuka et al. (1988) showed the effect of particle shape and asperity on the adhesive force between fine particles and a glass substrate. The equations 1.5 to 1.9 are valid for ideal (hard) particles, but Van der Waals (and electrostatic) forces may gradually deform the surface of the carrier particles. This reduces the separation distance and increases the contact area (Hinds, 1982). The occurrence of plastic deformation is most likely when the carrier has ductile layers of peptide and protein residues on its surface.

Electrostatic (e.g. Coulombic) forces may be only temporary, but they can increase the strength of interparticulate forces permanently by reducing the distance between the drug and carrier particle and/or increasing the contact area (Hinds, 1982). For electrostatic attraction forces many different expressions exist, depending on the charge mechanism, charge distribution and surface conditions, as well as environmental conditions, like the relative air humidity (e.g. Hinds, 1982; Hickey et al., 1994). Because electrostatic forces should preferably be avoided, they are not further discussed in this paragraph. Lam and Newton (1992), without giving relationships, mentioned that Coulombic forces are proportional to d-1.

Capillary forces are the result of a high (local) moisture content in the powder, which

for stability reasons is to be rejected. The capillary force depends on the radius (r) of the particle and the surface tension (γ) of the liquid and is for two smooth spherical particles of the same size (radius r) according Hickey et al. (1994):

FCAP = 2.π.γ.r [1.10]

33

Hinds (1982) presented nearly the same relationship for a sphere attached to a flat surface, except for the particle radius which is replaced by the particle diameter, meaning that this force is twice as high. Many others (e.g. Leuenberger et al., 1979; Stewart, 1986) included the contact angle (θ), yielding the equation: FCAP = 4π.γ.r.cosθ for a sphere on a plate [1.10a] There exists significant disagreement in literature about the numerical value of adhesive forces. An expression for the total adhesive force (FA) at room temperature on experimental basis (for a sphere attached to a surface) is given by Hinds (1982). FA = 150d[0.5 + 0.0045(%RH)] [1.11] where d is the diameter of the sphere in cm, %RH is the percent relative humidity, and the adhesive force is in dyne (10-5 N). According to this equation, FA at 50% RH decreases from 1.1 x 10-5 N for a 100 µm particle to 1.1 x 10-7 N for a 1 µm sphere. Increasing the relative humidity from 20 to 80%, increases FA for a 1 µm particle from 0.9 x 10-7 to 1.29 x 10-7 N. According to Stewart (1986), the total adhesion force (FA) is the sum of the force components that include Van der Waals, Coulombic and capillary forces. Stewart referred to one single example of a 115 µm quartz particle being charged to 0.3 x 10-12 C, which resulted in an attraction force of 4.6 x 10-5 N. This is about the same order of magnitude as calculated for a 100 µm particle with equation 1.11. Yet Stewart considered this force ‘several orders of magnitude greater than ordinary levels of adhesive forces’. It is difficult to compare these data with those presented by Lam and Newton (1992) who used a centrifugal technique to measure the size of the adhesive force for 19 to 34 µm particles (starch and spray dried lactose) attached to a surface of steel, because their particles where subjected to press-on forces of 10-5 N. They presented median adhesive forces between 8.5 x 10-7 (for 19 micron particles) and 17.5 x 10-7 N (d = 34 µm) for starch. For spray dried lactose, median FA appeared to be 2.5x 10-7 for d = 19 µm and 5.5 x 10-7 N for d = 34 µm particles. The data seem to indicate an omission in Equation 1.11, which has no term expressing the material properties and which yields an adhesive force of 0.33 x 10-7 N at 50% RH for 30 micron particles, though this is without applying press-on forces.

Schönert et al. (1996) derived an expression for the ratio of electrostatic (Coulombic) adhesion to Van der Waals attraction on the basis of the Lifshitz equation and calculated that his ratio has a value < 1 for microfine particles in the range between 1 and 10 µm (assuming specific values for the charge density, separation distance and the Lifshitz constant). They concluded that agglomerated microfine particles can hardly be broken down, which is in disagreement with Hinds (1982), Myers (1991) and Li et al., (1996), and also with daily practice with de-agglomeration of pellet formulations in dpi’s. Studies with centrifuge techniques in the 1980s (e.g. Staniforth, 1982) frequently yielded orders of magnitude for the adhesive force of 10-3 N, whereas from studies in the 1990s (e.g. Podczeck, 1996 and 1998) values in the range between 10-9 and 10-12 N were calculated. Podczeck (1996) also used small press-on forces (approx. 5 to 15 x 10-12 N) and measured (at lowest press-on force) adhesive forces between salmeterol-salmeterol, salmeterol-lactose and lactose-lactose combinations of 4.8 x 10-12 to 6.8 x 10-12 N (sphere-to-plate model for micronised spheres with unspecified diameters). Recently, adhesion forces between salbutamol and lactose carrier, respectively between budesonide and lactose, measured with atomic force microscopy (AFM), were found to vary between 3.5 - 6 x 10-7 (for salbutamol) and 3 - 5 x 10-7 N (for budesonide), depending on the relative humidity (Price et al., 2002), after application of a compressive loading (press-on force) of 2 x 10-8 N. These forces also refer to a sphere-to-plate

34

attraction; micronised drug particles were attached to the probe of the AFM and the lactose carrier was modified to exhibit ‘sub-nanometre scale roughness’.

Only adhesive forces with the order of magnitude of 10-12 N seem in agreement with the centrifugal separation forces that are generated in the classifiers presented and evaluated in the Chapters 4 to 11 of this thesis, which are capable of removing a substantial volume fraction of the drug particles from carrier crystals in adhesive mixtures for inhalation. 2.3.2.2. Adhesive mixtures: historical review

Adhesive mixtures consist of relatively large, free flowing carrier crystals with the micronised drug particles attached to their surface by adhesion forces (Fig. 1.6). Different names have been given to this type of mixture, including ordered (e.g. Malcolmson and Embleton, 1998) and interactive mixture (e.g. Louey et al., 2003), which frequently leads to hot disputes about which is the correct nomenclature. The history of adhesive mixtures goes back to the early 1950s when scanning electron microscopy became available. Travers and White (1971) where among the first to review some early studies on this subject. They referred to observations that indentations and irregularities on the surface of large diluent crystals act as adsorption sites for micronised components in the mixture. From exchange experiments between (micronised) bicarbonate rich sucrose carrier fractions and size fractions of pure sucrose, they concluded that strong and weak bonding sites exist on each carrier crystal. Hersey and co-workers were the first to discover that ordered mixtures may have a higher degree of order than expected for random mixtures on the basis of homogeneity indices as derived by for instance Lacey (1943) and Buslik (1973). This observation led to the introduction of a new ‘ordered mixing concept’ by Hersey in 1975, ‘requiring particle interaction, i.e. adsorption, chemisorption, surface tension, frictional, electrostatic or any other form of adhesion’ between the coarse (carrier) and fine drug constituents. In the years between 1975 and 1980, different explanations to the qualification ‘ordered’ were given in terms of homogeneity and type of mixture. Egermann (1980) observed the inconsistencies in nomenclature and proposed to use the degree of homogeneity as the basic criterion for naming powder mixtures. Reactions to this proposal (e.g. by Thiel et al., 1981) resulted in a reconsideration: the term ‘interactive’ would be best to describe mixtures with interacting constituents (in contrast with random), whereas ‘ordered’ should be used to indicate the degree of homogeneity of these mixtures (Egermann and Orr, 1983). However, Staniforth (1987), only a few years later, rejected to speak of interactive mixtures and proposed to use the term ‘adhesive mixtures’ instead, for the simple reason that all matter interacts. For powders, it depends on other (e.g. gravitational or electrical) forces whether the interaction results in an adhesion between particles or not. This reasoning has been adopted for this thesis.

At the time when Staniforth (1987) expressed his views, adhesive mixtures were almost exclusively prepared for tabletting and capsulation, and homogeneity and physical stability were the main concerns in adhesive powder mixing. In many different studies, factors influencing the mixture homogeneity and stability were investigated, e.g. electrostatic charge generation (Staniforth and Rees, 1982a), drug concentration (Chowhan et al., 1981), mixing conditions (Kulvanich and Stewart, 1987) and the shape of the fine particles (Wong and Pilpel, 1988). Additionally, results from various adhesive force measurements and segregation experiments were presented (e.g. Staniforth et al., 1981 and 1982; Staniforth and Rees, 1982b; Lai and Hersey, 1981; Malmqvist and Nyström, 1982). The results induced the development of a new ‘total mixing’ approach (Staniforth, 1981), in which powder mixing is considered as a dynamic process in which the exact balance between adhesive and non-adhesive mixing at a given instant is uncertain. In the one extreme, particles are non-adhesive (non-cohesive), whereas in the other they are adhesive.

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bulk properties on drug particle detachment from the carrier crystals during inhalation are discussed in the Chapters 4 and 5. 2.3.2.3. Adhesive mixtures: the role of carrier properties Mostly, carrier excipients in adhesive mixtures are special size fractions of alpha lactose monohydrate. Occasionally, other lactose modifications (beta anhydrous or spray-dried) have been proposed (e.g. Vanderbist et al., 1998; Harjunen et al., 2002), whereas also other sugars (e.g. sorbitol and mannitol) have been explored as carrier excipient (e.g. Steckel, 2003). Until quite recently, nearly all investigations on adhesive mixtures for inhalation were focussed on characterising and controlling the carrier surface properties, and measuring the adhesive forces between drug and carrier, using centrifugal techniques (e.g. Podczeck, 1998; Clarke et al., 2002) and atomic force microscopy (Price et al., 2002). Mixing theories (Paragraph 2.3.2.2) were developed on the basis of the previously mentioned assumptions that there is a competition during the mixing between cohesion (drug-to-drug interaction) and adhesion (drug-to-carrier interaction), and that the equilibrium between these two types of attraction at any moment during the mixing process is uncertain (Aulton and Clarke, 1996). It has also been postulated that the equilibrium can be driven in a certain direction by modifying the carrier surface properties (Staniforth, 1987). Very much attention has been given to so-called ‘active sites’ on the carrier surface, onto which drug particles are attached with high adhesive forces. In Chapter 5 of this thesis, a review is given of the variety of terms in which these active sites have been expressed by different research groups. They include surface irregularities (pores, clefts, cavities, lattice discontinuities), surface rugosity (coalesced or granular structures), adhering fines, amorphous spots, water of adsorption and impurities (water soluble protein residues, salts, decomposition products, riboflavin, urea, etc.). In Chapter 5, it is also summarised what investigations were undertaken to modify or to control the surface rugosity of carrier crystals (e.g. by corrasion processes, submersion and re-granulation) and what methods were proposed for characterisation of the rugosity (e.g. by permeametry and nitrogen adsorption). Rugosity has been classified into microscale (basically small irregularities on smooth crystal surfaces) and macroscale (large surface discontinuities and coalescent or granular structures) and in some studies it was concluded that a certain micro-rugosity (on a nanometre scale) may be desired, because it reduces the contact area with the drug particle (e.g. Kawashima et al., 1998). In most studies, macro-rugosity resulted in reduced fine particle fractions and therefore, it has been claimed that rugosity (e.g. as measured with permeametry) should not exceed certain values (Ganderton and Kassem, 1991). In neither of these studies, the effects of carrier payload, mixing conditions and type and size of the removal forces during inhalation were regarded relevant to the investigated drug-to-carrier interaction in relation to the carrier rugosity however.

Dickhoff et al. (2003) showed that carrier bulk properties, under certain circumstances, may be more relevant than carrier surface properties to the drug distribution over, and more particularly particle adhesion to the carrier particles. They concluded that the effects of carrier bulk properties and mixing conditions have been underestimated so far. Indeed, only a few references are known in this respect. Shotton and Orr (1971) referred to a much faster mixing with free flowing heavy magnesium carbonate as major component in adhesive mixtures than with low density carrier particles, due to a rapid dispersion of drug agglomerates by free flowing particles. Podczeck (1996) investigated the effect of press-on forces on the adhesion force between different substances and found an increasing adhesion force with increasing press-on force for all materials (and their combinations) investigated. Basically, the effect of press-on forces on the adhesion between drug and carrier particles is quite straightforward, as depicted in Figure 1.10. When two carrier particles make contact with each other during the mixing process, a press-on force is applied to drug particles in between the contacting carrier

37

surfaces. Contact between carrier particles could be the result of a collision (impact force), or occur when carrier particles slide along each other (resulting in a friction force). Both types of press-on force may increase the adhesive force between the drug particle and the carrier surface, and the magnitude of these press-on forces increases with increasing carrier size (as well as with decreasing size distribution). Figure 1.10. Effect of press-on forces (from colliding carrier particles) on the Van der Waals forces between drug and carrier particles. White arrows indicate the size and direction of the Van der Waals forces; black arrows the direction of carrier particle movement. An increase in the adhesive force could have different reasons, like a decreasing distance between the drug particle and the carrier surface or an increased contact area when ductile (e.g. peptide or protein) layers are present on the carrier surface. A drug particle may also migrate from a carrier site with lower, to a site with higher bonding capacity, which could include the migration from one carrier surface to another (Fig. 1.10). Because repeated migration will be necessary to obtain maximal occupation of the carrier sites with the highest bonding capacities, an increase in the adhesive forces in the mixture by the action of press-on forces depends on the mixing time. Even temporary weakening of individual adhesive forces may be possible. The effectiveness of the press-on forces, and the relevance of the carrier bulk properties relative to the carrier surface properties, depend furthermore on many parameters, like the carrier payload, the type and size of the removal forces during inhalation, the size distribution of the drug particles, the presence of lactose fines, the mixing conditions and even the type and scale of carrier rugosity, as will be further explained in Chapter 8. As an example, Dickhoff et al. (2003, 2005) showed that a high carrier rugosity may provide drug particles shelter from the press-on forces. 2.3.2.4. Spherical pellets

The addition of a micronised excipient to the drug, or co-micronisation of drug and excipient, mostly has the intention to prepare so-called spherical pellets (Trofast et al., 2002). Such pellets, in an approximate size range between 200 to 2000 µm, are the result of controlled agglomeration (without adding binder excipient) and subsequent spheronisation of the agglomerates. The porosity of spherical pellets for inhalation is generally quite high (between 60 and 80%) and their mechanical stability is low, as a result of which they are distorted upon impact (Boerefijn et al., 1998), which is intentional during inhalation but occurs also when for instance the inhaler is dropped. This may influence the dose measuring accuracy and the dispersibility of the powder. Micronised glucose (e.g. used by Boehringer

or

Van der Waals force

or

Van der Waals force

38

Ingelheim) and alpha lactose monohydrate (e.g. used by AstraZeneca) both find application as diluent in pellet formulations. A major advantage of spherical pellets is the relatively high drug content in the formulation compared with adhesive mixtures (up to 100% when no excipient is used). Also the relatively low energy required for dispersion is in favour of spherical pellets. The fragile structure of the highly porous pellets easily disintegrates into smaller fragments by the application of shear, friction or collision forces. Besides, as explained in Paragraph 2.3.2.1, the Van der Waals adhesion force between two spheres is theoretically half as high as that between a sphere and a large carrier surface, whereas drug particles in pellets can not find shelter from removal forces in carrier surface irregularities (as in adhesive mixtures), which has been explained in Paragraph 2.3.2.3. These advantages have frequently contributed to the misconception that cohesion forces are by definition lower than adhesion forces. The bulk porosity of pellets may vary considerably from batch to batch, due to variations in the pellet size distribution and intrapellet porosity. This is a major problem regarding consistency of measured dose. A variation in total powder porosity between for instance 60 and 80% includes a variation in metered mass between 133.3 and 66.7% of the nominal dose (corresponding with 70% porosity) in a metering chamber with constant volume, which is a difference by a factor 2. Moreover, the low mechanical pellet stability is a reason to limit the use of spherical pellets to inhalers with pre-metered dose compartments. 2.3.3. Powder de-agglomeration During inhalation, the adhesive forces that exist between the drug and carrier particles in adhesive mixtures, or the cohesive forces between drug particles in spherical pellets, have to be overcome in order to aerosolise primary drug particles. The magnitude of the separation forces during inhalation relative to the adhesive forces in the mixture determines the obtained fine particle fraction. Consequently, optimising a dry powder inhalation system with respect to delivered fine particle dose requires careful balancing between both types of forces. It is quite disappointing to discover that nearly all investigations in literature focus on control of the adhesive forces in the mixture, and that very few studies exist in which improvement of the de-agglomeration efficiency is the objective. It has been described that different types of de-agglomeration forces can be used (Amass, 1996; Voss and Finlay, 2002), but neither the effectiveness of these forces nor the principles to generate these forces have been studied thoroughly.

Basically, three different types of separation forces can be distinguished: friction forces (e.g. from particle sliding along inhaler walls), drag and lift forces (as in turbulent air streams) and inertial forces. Friction forces may result in high internal shear forces for spherical pellets (and therefore be effective for this type of formulation), but they can not get hold of particles in carrier surface discontinuities in adhesive mixtures (as will be discussed more in detail in Chapter 5). The same is true for drag and lift forces in turbulent air flows. Even drug particles attached to smooth crystal planes may not effectively be detached from these planes by so-called turbulent shear inhalers, because micronised particles are in the stationary boundary layer, unless the flow is directed more or less perpendicular to the carrier surface. But even then, air current forces are generally much lower than adhesive forces, except when the air velocity (relative to particle velocity) is extremely high (Hinds, 1982; Li et al., 1996). Most effective are the inertial (e.g. vibratory, centrifugal or impaction) forces, because they are proportional with the third powder of the drug particle diameter (drag and lift forces only on the first or second power), whereas their effectiveness is not necessarily influenced negatively by a high carrier particle rugosity (Chapter 4). Moreover, different technical means can be applied to sustain the action of such forces, like whirl, circulation or cyclone chambers. These considerations have resulted in design and development of the air classifier technology which is explained more in detail in the Chapters 4, 9 and 11, and which

39

is applied in a number of different test inhalers (e.g. Chapter 3), a marketed inhaler (Viatris Novolizer: Chapters 9 and 10) and a very promising disposable inhaler concept for high drug doses (Chapter 11). De-agglomeration of inhalation powders in literature is also referred to as disintegration, dispersion and redispension (in the air stream) of drug particles. The forces used to separate particles from each other during inhalation are described as break-up, separation, detachment, dislodgement or removal forces; the latter three mostly in combination with adhesive mixtures. 2.3.3.1. De-agglomeration principles

Powder de-agglomeration systems that are applied in dpi’s vary considerably in their principle of operation. Classification is often into breath operated systems, utilising the kinetic energy of the inspiratory air flow, and principles using auxiliary energy, such as electromechanical means and pressurised air. It is inherent in breath operated de-agglomeration principles that the magnitude of the detachment forces increases with increasing inspiratory flow rate through the inhaler. In contrast, battery and pressurised air operated dpi’s perform more or less independent of the inhalation manoeuvre in terms of fpf, but they are much more complex in design and therefore, expensive and vulnerable to failure (e.g. in case of flat batteries). Besides, it is arguable whether indeed a constant fine particle fraction at all flow rates gives a more reproducible therapeutic effect than an increasing fpf with increasing flow rate, as has frequently been claimed (to be explained in Chapter 10). Within the class of breath operated systems, subdivision into different categories can be made, of which several representatives have been described previously (de Boer et al., 2003b). For many dpi-designs, using for instance hard gelatin capsules or blisters as dose system, powder de-agglomeration is connected with emptying of the dose system (e.g. Schofeld and Hartley, 1968; Cavazza, 1980; Sumby et al., 1993; Brindley et al., 1995). The whole, or part of the inspiratory flow rate is directed through (or past) the dose compartment, in order to entrain the powder, while dispersing the particles by turbulent shear or by collision forces. Neither of these systems produces a high fine particle dose, particularly because the time during which the powder is subjected to the disruptive forces is quite short. In other capsule inhalers, the capsules are set to a particular motion to discharge and disperse the powder (e.g. Bell et al., 1971; Kladders, 1989; Salvatore, 1976), or a special flow pattern inside the capsule is created (Hovione, 1995): see Fig. 1.7A. Categories of de-agglomeration principles that do not operate in conjunction with the dose system, with their most well known or characteristic representative(s), have been summarised in Table 1.1. The category of miscellaneous principles also includes a powder capsule with propellants (Wetterlin, 1979), a battery driven piston tapping drug from a tape (Smith et al., 1990), a woven cloth from which drug is removed during inhalation with pressurised air ( Hochrainer et al., 1992), a circulation chamber with grinding balls (Aventis Eclipse, no reference available) and many other systems. Only (some) devices from the categories ‘miscellaneous’ and ‘circulation, whirl and cyclone chambers’ have the ability to sustain the action of the separation forces over a certain time period. They all make use of inertial forces, which makes them the most effective types of de-agglomeration principles.

In the Chapters 4 to 11, (multiple) air classifier technology is described as a highly efficient dispersion principle for (high dose) inhalation powders. Classifier technology offers certain advantages for adhesive mixture de-agglomeration in comparison with ordinary cyclone and whirl (or circulation) chambers, like complete carrier retention (with adjustable cut-off diameter) or carrier release with a controlled rate (adjustment of the residence time; Chapters 4 and 9). Carrier retention also appears to be a very useful tool in adhesive mixture optimisation (e.g. Chapter 8). Dependent on the specific classifier design, also spherical

40

pellets can be disintegrated with high efficiency (up to 90% of the dose), without having severe drug accumulation in the classifier chamber (by using bypass channels, creating an internal air barrier). And by designing the classifier as an integrated part of the mouthpiece, also the inhaler resistance to air flow can be controlled and fine particle deposition in the mouth can be reduced (by creating a sheath of clean air around the aerosol cloud), whereas the carrier deposition can be displaced from the throat to the mouth, so as to reduce adverse side effects by mouth rinsing after inhalation (Chapter 9). Table 1.1. Different categories of breath operated dispersion principles in dry powder inhalation. Dispersion principle Example(s)

Aerosol passage through narrow passages (e.g. Venturi tubes)

Orion Easyhaler (Raimo and Pekka, 1986)

Aerosol conducted against impact bodies (baffles, plates, internal inhaler surfaces)

Innovata Biomed ClickHaler (Parry-Billings et al., 2000) Skyehaler, SkyePharma (no reference)*

Aerosol conducted through specially shaped (discharge) channels or channels with (helical) inserts

AstraZeneca Turbuhaler (Wetterlin, 1988) Schering-Plough Twisthaler (Fan et al., 2000) DirektHaler, Direkt-Haler AS (no reference)*

Circulation, whirl or cyclone chambers (with or without control of residence time)

Chiesi Pulvinal (Meakin et al., 1998) Ivax Airmax (Keating and Faulds, 2002) Viatris Novolizer (Narodylo et al., 1998) Focus Taifun (Seppälä et al., 2001)

Pressurised air or vacuum chambers Nektar (Inhale) Pulmonary Inhaler (Burr et al., 1999) Vectura Aspirair (no reference)*

Battery powered (impellor) systems Dura (Elan) Spiros ( Han et al., 2002) Miscellaneous Ratiopharm Jethaler (Rymsa, 1998)

Aventis Ultrahaler (Pitcairn et al., 1997) Aventis Eclipse (no reference)*

* no reference has been found in which a proper description of the working principle is given 2.3.4. Balancing steps in dry powder inhalation

In this paragraph, the factors that influence the fine particle fraction and its site of deposition in the respiratory tract are surveyed. Many of these factors have to be balanced carefully. Balancing has for instance the objective to obtain maximal mixture homogeneity and stability on the one hand (requiring strong adhesive forces) and highest possible fine particle fraction during inhalation on the other (requiring weak forces). This includes (as a first step) balancing between the carrier bulk and surface properties regarding the choice of carrier to be made for adhesive mixture preparation. Good dose measuring requires good flow properties which can be obtained by selecting somewhat coarser carrier fractions. However, with increasing carrier diameter also the carrier surface roughness increases (Chapter 5), as well as the magnitude of the press-on forces that increase the adhesive forces in the mixture (Dickhoff et al., 2003). Also balancing between the obtained fpf and its site of deposition in the respiratory tract is necessary. Considering the increasing fpf with increasing flow rate (for some inhalers: Chapter 10), a high flow rate seems preferable, but this also shifts the particle deposition to higher airways.

To achieve all the objectives, a good understanding of, and control over the drug-to-carrier interaction forces in the adhesive mixture is desired. This includes knowledge of the relevant mechanisms for break-up of drug agglomerates during mixing and particle distribution over (and attachment to) the carrier crystals during the mixing process, as discussed in Paragraph 2.3.2.2. This, in dependence of the drug and carrier properties

41

(Paragraph 2.3.2.3), conditioning of the starting materials and the mixing process, as depicted schematically in Fig. 1.11. Unfortunately, not all properties of the carrier and drug can be fully controlled. Small variations in size and shape distribution are inevitable. Carrier particles are nearly always alpha lactose monohydrate, which is a natural product containing several impurities in the ppm-range, of which most are residues from the mother liquor and therefore, located on the carrier surface. Therefore, the attachment of drug particles is not exclusively to pure lactose. The amounts and composition of the impurities may slightly vary from batch to batch, whereas the decomposition products (e.g. from Maillard reactions) also vary with storage time. Moreover, some of the impurities (e.g. water soluble protein residues) may absorb more than 30 % water, and as a result, capillary forces could play a role in the drug-to-carrier attachment too. Finally, ambient conditions during storage and mixing of the starting materials can not always be controlled to the extreme, meaning that the moisture content (capillary forces) and tribocharge (Coulombic forces) can vary from batch to batch.

Figure 1.11. Simplified scheme of variables affecting the processes of fine drug particle break-up, distribution over, respectively attachment to the carrier particles during mixing, and fine particle detachment from the carrier crystals during inhalation.

Considering all these carrier surface variables, it is not surprising to find that most current academic research in the field of adhesive mixtures is focussed on studying and controlling the carrier surface properties. But it is disappointing to conclude that the majority of these studies involves only a single variable (e.g. carrier rugosity, polymorphism, carrier payload, moisture content, carrier purity, etc.) without taking account of the many (and complex) interactions between the variables (as discussed in Chapter 8). Also the fact that inhalers of different designs (with different types of removal forces, used at different flow rates) have been used, makes comparison of the results from different studies difficult and arguable. As a result, we are still unable to predict the effect of certain variables under other conditions than investigated, in spite of the current knowledge of possible types of interaction forces (Van der Waals, Coulombic, capillary forces) and the recognition of the relevance of carrier (surface and bulk) properties and the mixing conditions to the adhesive forces in the powder mixture.

Drug:- Type of drug- Size distribution- Conditioning- Payload on carrier

Carrier:- Surface properties- Bulk properties- Conditioning- Stability (aging)

Mixing process:- Type of mixer- Mixing time- Batch size

Mixture: - Type of mixture- Homogeneity- Conditioning

Inhalation test:- Type of inhaler- Inhalation manoeuvre- Test system

Fine particle fraction

Drug:- Type of drug- Size distribution- Conditioning- Payload on carrier

Carrier:- Surface properties- Bulk properties- Conditioning- Stability (aging)

Mixing process:- Type of mixer- Mixing time- Batch size

Mixture: - Type of mixture- Homogeneity- Conditioning

Inhalation test:- Type of inhaler- Inhalation manoeuvre- Test system

Fine particle fraction

42

Major challenges remain in: estimating the magnitude (numerical value) of the adhesive forces in the mixture,

depending on the mixing conditions (e.g. the press-on forces), carrier payload and the properties of the starting materials,

estimating the contribution of Coulombic forces to the drug-to-carrier interaction, developing good qualitative or quantitative expressions for carrier rugosity and impurity

(both referred to as active sites), controlling the carrier surface conditions (with respect to rugosity and purity) in a way

that is commercially interesting, measuring the presence of small fractions of amorphous lactose and estimating its

relevance to drug-to-carrier interaction, measuring the changes in the carrier surface properties during storage (stability testing)

and estimating the relevance of the changes in relation to drug-to-carrier interaction, estimating (and controlling) the balance between cohesion and adhesion at any moment

during the mixing process, estimating the role of adhering lactose fines in respect of occupying active sites,

promoting fine particle agglomeration during mixing, saturating large carrier surface discontinuities, providing shelter from press-on forces to drug particles, etc. etc.

The uncertainties of, and lack of control over previous aspects are the main reasons why

many different drug and carrier particle engineering techniques have been proposed to reduce or eliminate these effects (Paragraph 2.3.8).

Table 1.2 summarises some of the factors that influence the adhesive and removal

forces. The table shows that certain factors may influence both types of force. The flow (bulk) properties, batch size and mixing time influence the break-up of drug agglomerates, the drug distribution over the carrier crystals and particularly the press-on forces that are relevant to the attachment.

Table 1.2. Some factors influencing the adhesive forces (FA) during mixing and the removal forces (FR) during inhalation in an air classifier. Factor FA FR Primary particle size distribution of the drug (A) Break-up of drug agglomerates and (re-)agglomeration of fines (B) Increased contact area; reduced distance between drug and carrier: - from high press-on forces (B) - from ductile (e.g. peptide) layers on the carrier surface - from water of adsorption (capillary forces) - from tribocharge (Coulombic forces) (C) - from multiple contact points (high carrier rugosity) Increased inspiratory flow rate Increased inhalation time (e.g. in an air classifier) Flow properties in relation to classifier circulation

X X

X X X X X

X X

X

X X X

(A) Explained by Dickhoff et al. (2002) (B) Increasing the adhesive forces in the mixture, dependent on bulk properties (carrier size

distribution), batch size, mixing time, carrier payload and surface rugosity (C) Electrostatic forces may reduce particle velocity in a classifier

But they may also influence the circulation (average particle velocity) in a classifier

(and by that, the magnitude of the removal forces). The same is true for the occurrence of tribocharge. Not being able to control all these relevant parameters makes balancing between the adhesive and separation forces difficult, and for that reason a force distribution concept

43

(FDC) has been developed (Chapter 4). Balancing between the adhesive forces during mixing and the removal forces during inhalation with air classifier technology is practically also balancing (optimising) between the achievable inhalation manoeuvre and the fine particle fraction. This is a consequence of the possibilities of air classifier technology with which the residence time (equals circulation time) for the powder in the classifier can be controlled. This will be explained in Chapter 9 of this thesis. Considering the particle deposition mechanisms in the respiratory tract, as well as the demand of modern guidelines that administration of a dose has to be completed within 2 l of inhaled air, the aim is to achieve (sub-)maximal fine particle fraction at an average inspiratory flow rate of 60 l/min (1 l/s) within the first 2 s of inhalation.

Summarising all the aspects mentioned in the Paragraphs 2.3.2 and 2.3.3 that are relevant to the fine particle fraction, it can be concluded that optimising the fine particle deposition in the respiratory tract involves a cascade of balancing steps, of which each step is influenced by many variables, and superimposes on the next. Therefore, Fig. 1.11 (which shows these variables) can also be prepared in terms of balances, as shown in Fig. 1.12.

Figure 1.12. Scheme of balancing steps (white arrows) and major determinants (black arrows) in adhesive mixture preparation and inhalation for obtaining optimal dpi performance with respect to fine drug particle generation and deposition.

In Fig. 1.12, the balance between carrier bulk and surface properties (A) is primarily by carrier choice, but because both properties are correlated, the freedom of choice is limited. The balance between the cohesive and adhesive forces in the mixture (B) depends largely on the mixing conditions, but there is also a strong influence of the drug properties (tendency to break-up drug agglomerates), the carrier payload and the carrier surface conditions. The balance between attachment and detachment forces (C) can be controlled by the inhaler design (e.g. type of removal forces) and the flow manoeuvre through the inhaler (magnitude and duration of the removal forces). The final balance (D) requires either that the fpf from a dpi increases in a well controlled way with the inspiratory flow rate (as explained in Chapter

Carrier bulk

properties

Carrier surface

properties

Cohesion drug-drug

Adhesion drug-carrier

Interaction Separation

Fine particle fraction

Deposition

A

B

C

D

Starting materials

Mixing process

Inhaler design

Inhalation manoeuvre

Carrier bulk

properties

Carrier surface

properties

Cohesion drug-drug

Adhesion drug-carrier

Interaction Separation

Fine particle fraction

Deposition

A

B

C

D

Carrier bulk

properties

Carrier surface

properties

Cohesion drug-drug

Adhesion drug-carrier

Interaction Separation

Fine particle fraction

Deposition

Carrier bulk

properties

Carrier surface

properties

Cohesion drug-drug

Adhesion drug-carrier

Interaction Separation

Fine particle fraction

Deposition

A

B

C

D

Starting materials

Mixing process

Inhaler design

Inhalation manoeuvre

35

Figure 1.9. Competition (dynamic balance) between cohesion and adhesion in adhesive mixture preparation. The arrows indicate the direction of expected change in the type of particle-to-particle interaction during the mixing process.

This approach can be further elaborated into a concept specifically for adhesive mixtures, where in the one extreme the fine component consists of agglomerates (cohesive forces) and in the other extreme all fine particles have been attached to the coarser constituent (adhesive forces). Also for this understanding of the total mixing concept, which is in agreement with the competition between cohesive and adhesive forces during mixing as presented by Aulton and Clarke (1996), the equilibrium at any moment is uncertain, but it can be driven in a desired direction by changing the mixing conditions (e.g. the carrier payload, the carrier surface properties and the mixing time). This concept is shown in Fig. 1.9, in which the cohesive state represents the beginning of the mixing process. In this state, de-agglomeration (during inhalation) of the powder (drug agglomerates) requires least energy, but the degree of homogeneity of the powder is lowest. In the other extreme (adhesive state), homogeneity reaches its maximum, but the detachment of particularly primary particles from the carrier surface requires high removal forces, as will be discussed more in detail in the Chapters 4 and 8. The concept (Fig. 1.9) has limited applicability to mixtures with very high carrier payloads exceeding the theoretical monolayer of fines on the carrier particles, as first described by Schmidt and Benke (1985). In what they named ‘supersaturated ordered mixtures’, a mixture of adhesive and cohesive forces is self evident, even in the extreme of maximal drug distribution over the carrier crystals.

In the past fifteen years, investigations on adhesive mixtures have been focussed on their application in dry powder inhalers. The goals for such mixtures are different from those applied to produce tablets and capsules. For inhalation, an optimum is desired between homogeneity, stability and drug particle detachment during inhalation. Meaning that the adhesive forces have to be strong enough for processing (and storage) of the powder, but weak enough to be exceeded by the removal forces that are generated during the inhalation manoeuvre. Therefore, many of the investigations on adhesive mixtures for tabletting from the 1980s have been repeated, such as studies on the effects of electrostatic charge (Carter et al., 1992), changing the moisture content (e.g. Podczeck et al., 1997; Price et al., 2002) and modifying the carrier surface rugosity (e.g. Kawashima et al., 1998). Optimisation of adhesive mixtures for inhalation has been sought in selecting and defining special carrier size fractions (e.g. Arnold et al., 1995), but the possibilities are limited because of the requirements on powder flowability which are the actual reason for adhesive mixture preparation. A brief review of the optimisation approaches mentioned in literature is given in Paragraph 2.3.2.3. The performance of adhesive mixtures in an air classifier and the effects of carrier surface and

Cohesion Intermediate AdhesionLow degree of homogeneity

High degree of homogeneity

Cohesion Intermediate AdhesionLow degree of homogeneity

High degree of homogeneity

44

10) or that the range of attainable flow rates through the dpi is strongly confined. To limit the velocity with which the aerosolised drug particles are inhaled, a moderate to high inhaler resistance is desired. 2.3.5. The dpi as an external resistance to air flow

One of the most relevant characteristics of a dry powder inhaler is a relatively high air flow resistance compared to nebulisers and metered dose inhalers. This is a consequence of the design, which has elements of flow constriction to increase the kinetic energy of the air flow through the inhaler. Local pressure drops or high air velocities are necessary for adequate dose entrainment and powder dispersion. Therefore, a dpi can be considered as an orifice type of flow constrictor. The general relationship between the flow rate and the pressure drop for such type of constrictor is given by Equation 1.12 (Perry and Chilton, 1973):

Φ = Fu(A).(2dP/ρ)0.5 [1.12]

where Φ is the volumetric air flow rate (m3/s)

Fu(A) is a function of the cross section for air flow (m2) dP is the pressure difference across the flow constriction (kPa)

ρ is the density of the fluid (kg/m3)

The function Fu(A) varies with the type of constriction and may be rather complex, including contraction and flow coefficients or expressions with different cross sections (e.g. for a Venturi meter). Because dpi’s differ in their design, different formulas would be necessary to obtain proper relationships between Φ and dP. For that reason, often a simplified empirical formula is used, as for instance described by Clark and Hollingworth (1993):

dP0.5 = R.Φ [1.13] where R is the inhaler resistance, generally in kPa0.5.min.l-1 This is a general formula in terms of driving force (√dP), resistance (R) and resulting

volumetric air flow (Φ), which has similarity with Ohm’s law for electricity or Newton’s law for gravity. It can easily be derived that the overall resistance for serial (combined) or parallel resistances in a dpi follows the same rules as those for partial resistances against electric currents. In equation 1.13, R represents the reciprocal cross section for air flow and includes a correction for the fluid density as well. Relationship 1.13 is satisfactory for all currently marketed dpi’s: generally perfect linear relationships between Φ and √dP are obtained.

Olsson and Asking (1994) proposed a different relationship (dP = C.Φ1.9) in which they termed C the proportionality constant. Their relationship was derived as the best mathematical fit of experimental data, but could have special physical meaning for devices which have a short and narrow powder channel as one of the resistance components for which other Φ-dP relationships apply (particularly when laminar flow through this channel occurs), e.g. Hagen-Poiseuille (for Re < 2000). For the Turbuhaler with a channel diameter of 3.6 x 10-3 m, laminar flow requires an air velocity of less than 8.31 m/s. This velocity corresponds with a partial flow rate < 5.1 l/min. The actual partial flow rate through this channel is already considerably higher at a moderate flow rate of 30 l/min however, which is the lowest flow rate at which the Turbuhaler performs adequately. Therefore, turbulent flow in this channel occurs and the Fanning equation has to be applied. Consequently, there is also a Φ-√dP relationship for this channel. So, there is no physical meaning in the relationship of

45

Olsson and Asking (at least not for inhalers like the Turbuhaler). Practically, for all types of currently used inhalers, equation 1.12 appears to yield a better linearity. Table 1.3. Air flow resistances of some currently marketed dpi’s (kPa0.5.min.lN

-1) and the flow rates through the inhalers corresponding with 4 kPa (lN/min). Inhaler Resistance

Flow rate

corresponding with 4 kPa

GSK Rotahaler Aventis Spinhaler ISF inhaler (e.g. Pharmachemie Cyclohaler)A GSK Diskus Viatris Novolizer GSK Diskhaler (8 dose Becotide) Ratiopharm Jethaler AstraZeneca Turbuhaler Boehringer HandiHaler Inhalator Ingelheim

0.015 0.016 0.019

0.023-0.028 0.028B 0.032 0.036 0.036 0.045

0.051-0.062C

133 125 105

71-87 71 63 56 56 44

32-39 A also: Novartis Aerolizer and Ciba Geigy Foradil inhaler B average value after valve switching C range, depending on capsule position

Many different studies are known in which the air flow resistances of marketed dpi’s are presented (e.g. Sumby et al., 1992; Richards and Saunders, 1993; Clark and Hollingworth, 1993; de Boer et al., 1996). Resistance values for the same type of inhaler may vary between studies, because of inappropriate test arrangements used (or because of changes in the dpi design between studies). Particularly, the uncorrected use of laminar flow types of flow heads (pneumotachographs) is a source of error, as discussed by de Koning et al. (2001). The air flow resistances of some currently marketed dpi’s obtained with the procedures described by de Boer et al. (1996) are listed in Table 1.3. Although fifty percent of the investigated inhalers lies within the range of moderate air flow resistances (0.025 to 0.045 kPa0.5.min.l-1), the total range has an upper and lower value that differ by a factor 3.75, indicating that a 14 times higher pressure drop is necessary for the Inhalator Ingelheim to generate the same flow rate as through the GSK Rotahaler.

A high air flow resistance is favourable from deposition point of view (Paragraphs 1.2.4 and 1.2.6) because it reduces the velocity of the aerosol particles in the respiratory tract, thereby increasing deep lung penetration. It has been postulated that a high air flow resistance requires a high work of breathing (inspiratory effort) to operate the dpi correctly (Sumby et al., 1992), but calculation of the amount of work reveals that this supposition in not true (de Boer et al., 1996). By computing the cumulative area under the curve for a flow rate versus inhalation time curve (yielding the inhaled volume), multiplied for each integration step by the corresponding pressure drop across the inhaler, the total work of breathing (in Nm) is obtained. The mean numerical values for inhalation at maximal inspiratory effort through three different air flow resistances (39 healthy volunteers) appeared to range from 5.7 Nm (when R is 0.084 kPa0.5.min.lN

-1) to 11.2 Nm (when R is 0.015 kPa0.5.min.lN-1), indicating that

nearly twice as much work is done when a dose is inhaled (at maximal effort) from a low resistance dpi. A similar trend, though not so extreme, was observed at comfortable inhalation. The reason is the exponential increase in the flow rate with decreasing air flow resistance (at the same effort: Fig. 1.13), which has much greater effect in the computations than the decrease in pressure drop across the inhaler or the total inhalation time. The higher

46

work of breathing through a low resistance dpi is not necessarily experienced by the patient as a greater effort (or achievement) however. Therefore, this parameter seems to be not very useful. Yet, it is very relevant to inhaler design what patients experience with a certain external air flow resistance and more particularly what inhalation manoeuvre they can achieve through that resistance. Comparison of data from different studies is often difficult because of different units used for the air flow resistance. For the comparison of data, resistance values have been recalculated into kPa0.5.min.l-1. Andersen et al. (1993) reported highest preference for a low resistance (0.015 kPa0.5.min.l-1) by asthmatics and COPD patients and Clark and Hollingworth (1992) noticed that resistances higher than 0.032 kPa0.5.min.l-1 are uncomfortable for healthy volunteers. In contrast, 82% of the healthy volunteers in a study of de Boer et al. (1996) preferred to a moderate (or rather high) resistance between 0.021 and 0.047 kPa0.5.min.l-1.

Most flow manoeuvre experiments with a single inhaler have been performed with the

AstraZeneca Turbuhaler. De Boer et al. (1997b) compared the results of 12 different studies with adult healthy volunteers and asthmatics using this device. Mean PIF values in these studies varied between 55 and 68 l/min (the overall range for the 462 individual values was between 25-118 l/min) and there appeared to be no significant correlation with clinical picture. The difference with non-restricted flow rate (Paragraph 1.2.2) is evident. More recent studies confirm these data, although there may be several factors influencing the PIF through a dpi. Hawksworth et al. (2000) showed that counselling may increase the PIF through the Turbuhaler (the increase in mean value in this study was from 48 to 55 l/min for 24 asthmatics with mean age of 56 years), whereas Persson et al. (1997) concluded that PIF at ‘forceful and deep’ inhalation is on average 20% higher than that at ‘deep’ inhalation; the difference being greatest for patients with lowest PIF value at ‘deep’ inhalation. The results of Dewar et al. (1999) show that also older patients with obstructive airway disease (mean age is 69.1 year; n = 100) can attain a mean PIF of 53 l/min (range 28-78) through the Turbuhaler. Not many reliable data for other dry powder inhalers are known. Nantel and Newhouse (1999) reported PIF values between 40 and 90 l/min through the Clickhaler (Innovata Biomed), which they considered to be a medium resistance device on the basis of a rather arguable measurement. They described a pressure drop of 0.1 cm H2O at a flow rate of 0.5 l/min, which is not only a completely irrelevant flow rate to a dpi, but also yields an extremely high resistance value of 0.20 kPa0.5.min.l-1. Sarinas et al. (1998) used four different resistive loads during pulmonary function testing and found an exponentially increasing PIF (for patients with stable CF and COPD) with decreasing resistance, as may be expected on the basis of equation 1.13. Only two of their resistances were within the range of that for marketed dpi’s (Table 1.3) however. They did not show their exact test arrangement nor did they discuss the need for correction of their flow measurement. De Boer et al. (1996) and de Koning (2001) investigated PIF as function of the external air flow resistance for a more realistic range of values. In spite of completely different test equipment used for their experiments, different groups of healthy volunteers, different instructions and different air flow resistances, they found the same trends, as shown in Fig. 1.13. For both studies, there is an exponential increase in peak flow rate with decreasing air flow resistance and the results confirm that peak flows over 150 l/min are possible for resistances < 0.015 kPa0.5.min.l-1, which equals that of the Glaxo Rotahaler and the Aventis Spinhaler (table 1.3). The difference in numerical values between both studies may have been caused by the indicated differences in procedures. In the study of de Boer et al., patients were seated and were asked to apply maximal inspiratory effort through the resistance, or (in a second experiment) to inhale in (what they considered as) a comfortable manner. In the study of de Koning, patients were seated as well, but wore a nose clip and were asked to ‘inhale forcefully and deeply during a maximal inhalation’.

47

Figure 1.13. Comparison of peak flow rates through external air flow resistances measured by de Boer et al. (1996) en de Koning (2001), for different groups of volunteers. The spread bars indicate the ranges of values obtained by de Koning (for males and females). Closed symbols from the study of de Boer et al. are for male, open are for female volunteers: circles refer to maximal inspiratory effort, triangles to comfortable inhalation.

De Koning (1999) also investigated the effect of clinical picture on the inspiratory performance through a dpi, and concluded that even severe COPD patients (MIP, maximal inspiratory pressure, < 6 kPa; FEV1 < 60% of predicted) may (on average) still have 62% of the PIF that can be generated by healthy volunteers through the same resistance of 0.041 kPa0.5.min.l-1. Furthermore, he investigated the effect of respiratory muscle training on the inspiratory performance, and found an increase in the peak maximal inspiratory pressure (P-MIP) of (on average) 25% after eight weeks of training (by taking five forceful and deep inhalations twice daily). In a control group (without training) no significant increase in P-MIP was observed. 2.3.6. The use of a dpi: variables and interactions The principle of operation of a dry powder inhaler strongly relates to its basic design described in Paragraph 2.3.1. With regard to preparation for inhalation, the patient has to insert a single dose compartment (capsule) or to transport a strip or disk with compartments (e.g. blisters). In most cases, this compartment (blister or capsule) has to be opened (or perforated) next. Alternatively, the patient has to operate a dose metering system. These handlings are conducted according to the instructions given by the manufacturer which may be different for different inhalers although they seem to have comparable designs. For example, the Inhalator Ingelheim and ISF inhaler must be held in a vertical position during perforation (opening) of the capsule, but the Rotahaler may not. Further implications of incorrect handling have been described by de Koning (2001).

The principle of operation also concerns the powder de-agglomeration and dispersion, which in most dpi’s is breath powered. The inspiratory air flow has to discharge the dose system and to entrain the powder to the de-agglomeration principle where its kinetic energy is utilised for further break-up of the agglomerates. For inhalers that are designed to effectively use the available energy within the inhaled air flow through the inhaler, the efficacy of dose entrainment and powder de-agglomeration varies with the inspiratory manoeuvre. Therefore,

0

20

40

60

80

100

120

140

160

180

0 0.02 0.04 0.06 0.08 0.1

air flow resistance (kPa0.5.min.l-1)

flow

rate

(l/m

in)

de Koning

de Boer, max

de Boer, max

de Boer, comf

de Boer, comf

48

it is recommended to speak of breath controlled instead of breath operated (or actuated) devices. Obviously, the efficacy at the same flow manoeuvre varies between inhalers of different designs. With the inhaler design and the inspiratory flow manoeuvre as the major determinants for dpi performance, it is necessary to know which external variables may influence the flow manoeuvre through a dpi (next to the already discussed inhaler resistance), and what interactions exist between all these parameters and the lung deposition fraction.

Fig. 1.14 confirms that the performance in terms of dose entrainment and generated fine particle fraction (fpf) is a function of the flow manoeuvre and the inhaler design, and that the most essential aspects of the inhaler design in this respect are the powder formulation, the dose system and the powder de-agglomeration principle. For most inhalers, the de-agglomeration principle gives the highest contribution to the total inhaler resistance (R). Equation 1.13 (Paragraph 2.3.5) shows that the generated pressure drop (dP) and the resistance (R) together determine the flow rate (Φ) through the inhaler. Which pressure drop is generated during inhalation, depends on the patient’s effort, which is influenced by many patient related factors, including the instructions for use of the dpi as given by the medical or pharmaceutical specialist. Different studies have shown that the PIF may not be the only relevant flow parameter for dpi performance however; some inhalers are more sensitive to the flow increase rate (de Boer et al., 1997b). Particularly for capsule inhalers, the inhalation time may be critical too, as it takes a certain time to fully empty the capsule. For a breath controlled dpi, the inspiratory flow manoeuvre not only governs the fine particle fraction, but it also controls the deposition of that fraction in the respiratory tract (Paragraph 1.2.6).

Figure 1.14. Simplified scheme of variables and interactions in dry powder inhalation therapy.

For different reasons, many of the described variables and interactions are relevant to the air classifier technology in this thesis and its application in the Novolizer multi-dose dry powder inhaler. The classifiers have been designed to maximise utilisation of the kinetic energy of the inspiratory flow rate. Therefore, the flow rate dependence is high. The classifiers also either completely retain the carrier crystals in adhesive mixtures, or control their residence time inside the classifier, meaning that the inhalation time is relevant to the obtained fine particle fraction. The inhalation time correlates inversely with the PIF, as shown

Performance- dose entrainment- fine particle fraction (fpf)- lung deposition

DPI design- powder formulation- dose system- de-agglomeration principle

Flow manoeuvre- peak flow rate (PIF)- flow increase rate (FIR)- inhalation time

Air flow resistance

Inhalation effort

Patient factors- instruction- clinical picture- age, gender, training- smoker/non-smoker

Performance- dose entrainment- fine particle fraction (fpf)- lung deposition

DPI design- powder formulation- dose system- de-agglomeration principle

Flow manoeuvre- peak flow rate (PIF)- flow increase rate (FIR)- inhalation time

Air flow resistance

Inhalation effort

Patient factors- instruction- clinical picture- age, gender, training- smoker/non-smoker

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by the Boer et al. (1996) and de Koning (2001), because the inhaled volume is more or less a patient constant. 2.3.7. DPI performance For the evaluation of dpi performance, different methods are available. Standard in vitro testing includes measurement of the consistency of delivered dose and fine particle dose, using single dose collector systems and inertial (cascade) impactors respectively. In order to assess the de-agglomeration efficiency of an inhaler system, the fine particle dose is expressed as percent of ‘total dose’, yielding the fine particle fraction (fpf) which is often defined as the fraction of particles < 5 µm. Single-dose collectors are frequently filter systems, as for instance described by the European Pharmacopoeia, 3rd edition (1997, supplement 1999). Multi-stage inertial impactors are also described by the Pharmacopoeia, as well as the procedures to operate these devices. Yet, it is quite difficult to compare the results from different in vitro studies with each other. Mostly, because many studies were completed before the pharmacopoeial procedures were well standardised. A great number of different impactor types has been used in the past, having different cut-off diameters and different collection efficiencies (e.g. Hallworth and Andrews, 1976; Olsson et al., 1996). Therefore, fine particle fractions have been defined differently with respect to upper class limit. Moreover, fractions have been expressed as percent of label claim (or nominal dose) (Olsson, 1995; Hill and Slater 1998), of real dose (Broadhead et al., 1996) and of emitted dose (Steckel and Müller, 1997), sometimes even without giving proper notification (e.g. Vidgrén et al., 1987). Most studies in the 1990s were performed at 60 l/min, which flow rate was considered as optimal for most marketed dpi’s (e.g. Vidgrén et al., 1987; Broadhead et al., 1996; Geuns et al., 1997). First in the late 1990s, it was recognised that dpi’s should rather be tested at the same pressure drop across the device which is the driving force for the flow. Therefore, in some studies, high resistance inhalers were tested at lower flow rates than low resistance devices, reflecting approximately the same inspiratory effort (Steckel and Müller, 1997). At present, testing at 4 kPa is dictated by various guidelines and pharmacopoeias (e.g. European Pharmacopoeia 3rd edition, 1997, supplement 2001), but by that, still solely the pressure drop or peak flow rate through the inhaler is controlled, neglecting other relevant flow parameters such as the flow increase rate. Moreover, 4 kPa may not be the most relevant pressure drop for all types of dpi’s, as can be concluded from the studies of de Koning (2001). Due to their differences in air flow resistance, the same effort may result in different pressure drops: the lower its resistance, the lower the maximal attainable pressure drop across a device will be.

In literature, different aspects of dpi-performance are presented and discussed such as delivered dose, inhaler accumulations, flow rate sensitivity, moisture protection by the inhaler, etc. Most frequently investigated devices are the ISF inhaler, the (AstraZeneca) Turbuhaler and the (GSK) Diskus, which are currently the most successful inhalers on the market. Several studies with the Turbuhaler were reviewed before (de Boer et al., 1997b). These studies present fpf’s at 60 l/min ranging from 18 to 52%. The difference in results rather reflects the poor reproducibility of the test procedures than that of the inhaler itself however. Although it has been suggested that the Turbuhaler is sensitive to the inspiratory flow rate between 40 and 80 l/min (e.g. Olsson, 1995), later studies proved that it is rather the flow increase rate that determines the fpf from this device (de Boer et al., 1997b; Everard et al., 1997). And as discussed above, flow increase rate is not defined by the pharmacopoeial procedures, which is one reason to explain the differences in fpf from the Turbuhaler between different studies.

Vidgrén et al. (1987) evaluated the fine particle fractions from the ISF inhaler at 60 l/min for mixtures of Pharmatose 325M carrier lactose and two different batches of disodium

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cromoglycate (dscg), prepared by micronisation and spray drying respectively (dscg : lactose is 1:1 by weight). Using the Sierra Andersen impactor, they found a difference in fpf, reflecting the difference in primary particle size distribution of both dscg batches: 13.0% was delivered within the size range 3.3 - 7.1 µm and 26.4% in the size range 0.3 - 3.3 µm for the spray dried product (having a mean aerodynamic diameter of 2.81 µm). Corresponding fine particle fractions for the micronised product (with a mean diameter of 3.81 µm) were 29.7% (3.3 - 7.1 µm) and 10.1% (0.3 - 3.3 µm) respectively. Broadhead et al. (1996) also tested this ISF inhaler at 60 l/min for a model spray dried protein (β-galactosidase) in different formulations, stored at 0 and 43.2% relative humidity (RH). They obtained substantially lower fpf-values than Vidgrén et al. (1987), varying from only 12.8% (no carrier; storage at 43.2% RH) to 24.7% (Avicel carrier; storage at 0% RH). They concluded that only 10% of this drug from the ISF inhaler has potential for alveolar absorption and that a large portion of the remaining material is retained in the device and capsules. Geuns et al. (1997), using the twin impinger, showed (for beclomethasone dipropionate from the GSK Rotahaler) that storage of the drug formulation in hard gelatin capsules under different relative humidity conditions may result in a dramatic change (generally a decrease) in fpf.

Steckel and Müller (1997) were among the first to test the performance of different

inhalers under the same conditions. They measured the fine particle fractions (< 6.4 µm) collected in the twin impinger and the multi stage liquid impinger (MSLI) for eleven different inhalers at 60 l/min and found comparable values between both testing devices. Highest fpf’s were found for the Turbuhaler (nearly 40% of the total dose) and lowest for the Spinhaler (< 10%). Seven out of eleven devices exhibited fpf’s (from the MSLI) between 30 and 40%. They also tested some low resistance inhalers at 90 l/min and high resistance devices at 30 l/min with the twin impinger and concluded that certain devices (e.g. Spinhaler, ISF inhaler) are sensitive to flow rate, whereas others show less (e.g. Easyhaler, Inhalator M) or no flow rate dependent performance (e.g. Diskhaler, Inhalator Ingelheim). Furthermore, they measured delivered doses at 60 l/min, varying from 60% (for Pulmicort Turbuhaler) to more than 105% of label claim (for ISF inhaler with cromolyn Orion), and mouthpiece accumulations ranging from only 2.5% (Spinhaler) to nearly 30% (Pulmicort Turbuhaler). Most data are in good agreement with those from another comparative study performed in the same year by de Boer et al. (1997a).

Srichana et al. (1998) mentioned the inhaler resistance and powder formulation as the most relevant design parameters for dpi performance. Their flow rate versus pressure drop measurements resulted in a different sequence of device resistances than has been obtained in most other studies, with the Diskhaler exhibiting a surprisingly low value. They also presented a remarkably high fpf for this device (more than 40% of the nominal salbutamol dose at flow rates higher than 60 l/min). In comparison, Olsson (1995), Steckel and Müller (1997) and de Boer et al. (1997a) obtained fpf’s of only approximately 21% (for bdp at moderate inspiratory force), 30% (for salbutamol at 60 l/min) and 23% (for bdp at 60 l/min) respectively from the Diskhaler. In fact, Srichana’s study is one out of many comparative investigations, all from the same period (1997-2001), in which the Diskhaler (or Diskus) was involved. Most of these studies were in comparison with the Turbuhaler (e.g. Prime et al., 1997; Hill and Slater, 1998; Burnell et al., 2001), and it is striking that in all these studies the experimental set-up was designed in such a way that the GSK-inhalers (without proper de-agglomeration principles) were presented as the better performing devices in respect of fpf.

De Koning (2001) was the first to investigate the effect of both PIF and FIR on more than one device and concluded that only the Turbuhaler is highly sensitive to FIR, producing a maximal (budesonide) fpf of 50% of the label claim at 60 l/min and 7.5 l/s2 respectively. His study confirmed also that the maximal fpf from the (fluticason) Diskhaler is only 23% (versus

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33% for the fluticason Diskus). He furthermore observed that the Diskus and Cyclohaler are only slightly FIR-dependent, which supports the idea that in vitro testing with simulated inhalation manoeuvres would yield a more realistic view on the performance of a dpi. Recently, it has been shown that the fpf from breath controlled inhalers (also referred to as passive inhalers) can be doubled when specially engineered drug particles with improved dispersion properties are used (e.g. Duddu et al., 2002; Newhouse et al., 2003; Schiavone et al., 2004). In this thesis, it will be shown (Chapter 11), that multiple air classifier technology can achieve even higher fpf’s from classic (micronised) drug formulations at lower air flow rates.

Selroos et al. (1996) and Pauwels et al. (1997) summarised the results from various

lung deposition (in vivo) studies with currently marketed inhalation devices. They showed that deposition varies not only strongly between devices, but also between studies with the same device and same type of drug. One of the reasons is a difference in the inhalation manoeuvres, which can not be controlled to the same extent as in vitro, but also the inter subject variations with respect to lung morphology may be quite extreme. The total range of presented depositions has a span between 5.5% (for dscg Spinhaler) and 32% (for budesonide Turbuhaler) of the label claim. The arithmetic mean of all studies is 16.0% (in the comparison of Selroos et al.), respectively 16.2% (in the comparison of Pauwels et al.), which indicates that there is still room for considerable improvement. In comparison, mean lung deposition from mdi’s is 15.0% (as reported by Selroos et al.: ranging from 7.2 to 26.2%), respectively 12.3% (Pauwels et al.: ranging from 2.9 to 24.1%). The extreme range of values for the same device tested with the same type of drug at approximately the same flow rate (e.g. Turbuhaler, varying between 16.8 and 26.9% for terbutaline at 55 l/min, or between 15 and 32% for budesonide at, on average, 44 l/min) indicates that lung deposition experiments may be useful in predicting (or explaining) ultimate clinical effects, but they certainly are not useful for device development, also because they are time consuming and expensive.

In the past ten years, several new devices have been presented and evaluated. Many in

vitro investigations of new devices were combined with in vivo deposition (scintigraphic) studies, or clinical effect studies, and some of these data have been compared by Newman et al. (2002). His comparison shows that the lung depositions of the new devices (in the range between 13.6 and 41.5% of label claim) are generally higher than those from the first generation dpi’s. With an exception for the Turbuhaler, which device up to date is still considered as the competitor to beat or the standard for generic formulations. But even the Turbuhaler, and nearly all new devices, have considerable room for improvement regarding not only dose consistency and fpf, but also regarding ease of production (cost aspects, including the formulation), ease of operation, acceptance of high (or extremely low) doses, etc. Finally, some studies are known in which the patients’ acceptance of a particular device, or the patients’ preference between devices was investigated (e.g. Schlaeppi et al., 1996; Sharma et al., 1996). Such studies could be very useful for design of a new inhaler device (Chapters 9 and 11) if performed properly. In practice, they are merely inventories of opinions about one or two existing (competitive) devices, too often on the basis of arguable questioning, and are rather meant to confirm the manufacturer’s views on (and choices for) design than to bring insight into the real desires and needs of the patient (or the physicians). 2.3.8. Recent trends in dry powder inhalation

Many trends in dry powder inhaler (and formulation) design and development can be observed. According to Ashurst et al. (2000), more than 30 dpi’s were under development at the start of the third millennium, and more recently, even higher numbers of inhalers in the

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pipeline are given by different sources on the Internet. Ashurst et al. (2000) classified the driving forces for new inhaler design in pharmaceutical, clinical and commercial factors. Among the pharmaceutical factors, new (pulmonary) delivery systems for new chemical entities were mentioned. Also improved delivery of vaccines and antibiotics may be possible with dry powder inhalers. Vaccines are mono-dose drugs whereas antibiotics bear the hazard of development of bacterial resistance, which could be lethal, particularly in CF. For these, and many other reasons, a growing desire for disposable inhalers exist, which becomes clear from patent literature and activities of pharmaceutical companies in the field of information exchange on the subject. Several single dose inhaler developments are known, e.g. Keldmann and Reipur (1996), Jahnsson (1998), McDerment and Bowman (2001), of which the first is already in the phase of marketing (DirektHaler, Direkt-Hale A/S). This includes RuG’s Twincer development (Fig. 1.15), which is described more in detail in Chapter 11 of this thesis.

Figure 1.15. Basic design (with three plate-like parts and a single blister compartment) of the disposable dry powder inhaler (Twincer) with multiple air classifier technology described in chapter 11 of this thesis.

There is a particular interest in high dose disposable inhalers, because many of the new drugs for dry powder inhalation are either given in relatively high doses (e.g. colistin), or they are produced with one of the so-called particle engineering technologies which frequently include the incorporation of relatively large amounts of excipients in the formulation. The bulk volumes of such formulations are usually rather extreme due to high intraparticulate porosities. Such ‘smart particles’ are for instance produced with (co-)spray drying of drugs and excipients (e.g. Venthoye et al., 2001), emulsification and freeze drying techniques (Edwards et al., 1997) or super critical fluid technology (Beach et al., 1999). Primary objectives are stabilisation of peptide and protein like drugs in the dry state (sugar glass technology), increasing the ratio of geometric to aerodynamic particles in order to avoid phagocytic clearance (large porous particle technology) and increasing the dispersibility (PulmoSphere technology). These particles have reduced interparticulate forces in common, as a result of their irregular surface structure and/or (claimed) reduced surface free energy. They are also expensive to manufacture and (in some cases) highly hygroscopic and instable, and require excellent moisture protection within the inhaler, as well as special measures to prevent exhalation through the device, which makes the inhaler complex and expensive too.

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In some cases, application may even be limited to single use (disposable) inhalers. In addition to the drug particles, carrier crystals may be engineered, or special additives may be used to control the adhesive forces between drug and/or carrier particles. Some trends in this respect, including the use of so-called ‘force control agents’ (FCA’s; ), are discussed more in detail in Chapter 5.

Ashurst et al. (2000) also mentioned power-assisted dispersion and devices that are flow-rate independent as important clinical factors for new dpi-development. In the past decade different devices with battery powered (e.g. Dura Spiros) or pressurised air facilitated de-agglomeration principles (e.g. Nektar dry powder inhaler and Vectura Aspirair) have been introduced. It has been recognised that battery powered systems are vulnerable to failure however, and for that reason the pressurised air systems seem to have better perspectives, particularly when the moment of dispersion is synchronised with the inhalation manoeuvre. Yet, such systems with auxiliary energy are quite complex (and expensive) which makes them inappropriate as disposable devices. Moreover, it may be questioned whether a constant (flow-rate independent) fpf is really desirable if a considerable range of flow rates can be attained through the inhaler (Chapter 10).

3. Aerosol characterisation 3.1. Relevance and principles of aerosol characterisation The most important physicochemical parameter influencing the deposition in the lung is the particle or droplet size (Hickey, 1992). The size distribution in the aerosol, in combination with the inhalation manoeuvre by the patient, is the predictive parameter for lung deposition (Paragraph 1.2.6). But it is also a reference for the performance of the drug delivery system, particularly for systems that are breath operated, like dry powder inhalers. For that reason, particle size analysis is a standard procedure in quality control and device or formulation development. Many different sizing techniques for aerosols are available, as will be reviewed in Chapter 2. Not all techniques are appropriate for the characterisation of medical aerosols however. The analysis has to be performed immediately after the aerosol leaves the inhaler’s mouthpiece to avoid fractionation by sedimentation. Additionally, the aerosol has to be generated at different inspiratory flow manoeuvres, so as to investigate the powder dispersion efficiency (for dpi’s) and the droplet coalescence or evaporation (for nebulisers and mdi’s), etc. as function of this manoeuvre. And finally, the aerodynamic size distribution, rather than a geometric equivalent, is desired for reasons explained in the Paragraphs 1.2.4 and 1.2.6.

Different sizing techniques produce aerodynamic size distributions, but theoretically, only sedimentation analysis yields the aerodynamic diameter by definition (as given in Paragraph 1.2.5). The dynamic force balance during analysis with cascade impactors, wind sifters and equipment for time of flight measurement, may have an effect on the dynamic shape factor of the particles (for instance, particles may start to rotate, as will be explained in Chapter 2). Therefore, these techniques may yield undefined equivalents for the aerodynamic diameter. Some principles should neither be used for other reasons, such as wind sifters and the time of flight (tof) principle for aerosols produced by dry powder inhalers. They break up small agglomerates into primary particles in the violent air streams inside the sifter wheel and the sonic nozzle (e.g. Foss and Egilmez-Reynolds, 2000). Therefore, they measure the primary particle size distribution of the drug rather than the mixture of primary particles and small agglomerates as released by the inhaler. Tof instruments are also susceptible to coincidence measurement problems when measuring high concentration aerosols (Stein et al., 2000). Laser diffraction analysis does not yield aerodynamic size distributions and is therefore

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often rejected as a suitable sizing principle for inhalation aerosols. It is the aim of this thesis to show that the rejection is incorrect and that laser diffraction technique is a powerful tool in inhaler development and quality control. 3.2. Cascade impactor analysis (cia) The principle of cascade impaction is described by Hinds (1982), Hickey (1992) and many others. A major advantage of cascade impactor technique is the chemical analysis of the mass fractions of the drug deposited on the different impactor stages, having different aerodynamic cut-off diameters. This enables the analysis of drug mixtures and/or drug particles in the presence of excipient particles. Cascade impactors have several major drawbacks and limitations too however, as will be discussed in Chapter 2. The facts that they have been adapted by different pharmacopoeias as the technique of first choice, and that the test arrangement, as well as the conditions under which they have to be operated, have been described in detail, do users of the technique believe that the procedures have been completely validated and that the obtained results are reliable and comparable with those from other investigations. A study of Olsson et al. (1996), supervised by a task force of experts in the field, shows that this assumption is incorrect. Studies of the effect of FIR on the performance of certain dry powder inhalers support the criticism (de Boer et al., 1997b; Everard et al., 1997; de Koning et al., 2001), because there are no specifications for FIR in the pharmacopoeial procedures. Consequently, the inspiratory flow curve is not fully defined. Many errors in the analysis procedures (including assembling and sampling) are still possible and the technique is virtually impracticable for device and formulation development, as will be explained in Paragraph 3.3 and Chapter 2. Cascade impactor procedures are furthermore laborious and time consuming for routine measurements, whereas the effect of the inspiratory flow manoeuvre on the performance of so-called breath operated inhalers (dpi’s) can not be investigated without using complex auxiliary equipment, like spacers or flow splitting and compensating devices which influence the size distribution in the aerosol by sedimentation, electrostatic deposition or inertial classification. The introduction of the next generation impactor (NGI) has not brought satisfactory solutions for these problems (Roberts et al., 2001). 3.3. Laser diffraction analysis (lda) Laser diffraction has great potential to become a good alternative. The technique is fast, reliable and accurate, and the theories of light diffraction have well been described (e.g. Müller and Schuhmann, 1996). Major advantages are the large number of classes in the size range that is interesting for lung deposition, and the possibilities for automation. Practical limitations of standard laser diffraction are the absence of means to generate a controlled flow rate through the inhaler and the fact that not the aerodynamic particle size distribution is obtained. These, and other pros and cons, with possible solutions, are discussed more in detail in the Chapters 2 and 3.

Laser diffraction technique has been introduced in studies on nebulisers in the 1990s (e.g. Hurley et al., 1994; McCallion et al., 1996; Coates et al., 1997). The reasons to use this sizing technique for nebulisers are quite obvious. Nebulisers generate aerosols over a long period (5 to 30 min). Many cascade impactors are less appropriate for long term measurements, because of possible dry running of the collection plates (and the occurrence of electrostatic charge). Changes in the aerosol size distribution during nebulisation can not be studied properly with impactors, unless the aerosol from a single dose is divided into a great number of measurements. Moreover, size distributions for aqueous droplets obtained with

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laser diffraction technique do not differ significantly from aerodynamic size distributions, as such droplets are spherical and have (almost) unit density. The only problems to overcome for aqueous aerosol testing are a choice to be made for the diffraction theory (Mie or Fraunhofer) and the control of an inspiratory flow rate through the nebuliser. The latter problem has been solved in a number of studies by directing conditioned air (at higher than atmospheric pressure) through the one-way valve on top of the nebuliser (e.g. Ho et al., 2001). A choice between Mie and Fraunhofer is more problematic (as will be discussed in Chapter 2), but a calibration on (largely monosized) droplets produced with Rayleigh break-up may ease this choice. Droplet size from Rayleigh break-up is theoretically well defined (1.89 times the initial diameter of the jet; Perry and Chilton, 1973) and measurement of Rayleigh break-up droplets of the drug solution could confirm whether the Mie parameters for processing of the diffraction data (‘optical model’) have been selected properly. It could also provide a correction factor for re-calculation of the data, if there is a difference between theoretically expected and experimentally obtained data.

Recently, more applications for laser diffraction technique have been presented (e.g.

Ziegler and Wachtel, 2001; Holmes et al., 2001). However, the application for dry powder inhaler evaluation or development has not been reported yet, except for the studies in this thesis and those presented by Dickhoff et al. (2003, 2005). Most frequently mentioned argument against this application is the fact that the aerosol from a dpi is a mixture of primary particles and small agglomerates, for which the dynamic shape factor and density vary from particle to particle. Consequently, an aerodynamic size distribution can not be calculated from the laser diffraction result. Additionally, it is often mentioned that laser diffraction technique is only applicable for perfectly spherical particles. Both arguments are not relevant however. The aerodynamic size distribution is only important for predicting lung deposition. The objective of dry powder inhaler (or formulation) development is to get the size distribution of the aerosol from the dpi as close as possible to the size distribution of the primary drug particles. This has to be achieved for the highest possible mass fraction of the dose and under inspiratory conditions that are preferably comfortable to the patient and favourable for drug deposition in the target area. Therefore, the size distribution of the drug as obtained from RODOS dispersion is the reference. In most applications, this only requires a pre-separator to remove the larger carrier particles from the air stream and a fine particle collector to measure the mass fraction of the dose for which the size distribution has been measured, yielding a test arrangement as depicted in Fig. 1.16.

With the experimental set-up depicted in Fig. 1.16, with and without pre-separator and

fine particle collector, many applications have already been found (in addition to the studies described in this thesis), even although the current design is still a prototype with many limitations and drawbacks. For instance, exploring and exploiting Rayleigh break-up for aerosol generation (as initiated by MedSpray, Enschede) would have been impossible without laser diffraction technique, as size analysis in the aerosol as function of the pressure profile over the membrane requires the possibility of time sliced measurements. Short-term measurements (in the range of ms) are not possible with cascade impactors. Neither can droplet coalescence (and evaporation) as function of chip geometry (number and ordering of the orifices), mouthpiece design, inspiratory flow rate, properties of the drug solution (etc.) effectively be studied with standard cascade impactor analysis. Development and testing of a dry powder inhaler for antibiotics, as presented by de Boer et al. (2002a) and Le Brun et al. (2002), has required only a few months time from idea generation to proof of principle, With cascade impactor analysis, this development would have cost a few years.

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Figure 1.16. Adapter for laser diffraction analysis of the aerosol cloud from inhaler devices with pre-separator and fine particle collector: prototype.

In another application, measurement of the size distribution of the drug particles as

they are released from carrier crystals in an air classifier at very low inspiratory flow rate (10 l/min) appeared to be possible with laser diffraction technique (de Boer et al., 2002b). With this application it could be shown that the size distribution in the aerosol from the classifier shifts to larger diameters when the mean carrier diameter is increased. The shift is more pronounced at longer mixing times. This observation can be explained with an increasing magnitude of the press-on forces during the mixing process, which forces cause the drug agglomerates to form small agglomerates on the carrier surface. Such experiments are not possible with cascade impactor analysis, if it were only because released drug agglomerates deposit in the inlet tube at such low flow rates. Similar experiments are presented and discussed in Chapter 7.

In spite of all these valuable and successful applications, a critical note is in place.

Laser diffraction results require a different (and critical) way of data interpretation. For dry powder inhalers, it is not the size distribution of the aerosol that matters, but the size distribution in relation to that of the primary drug particles (e.g. from RODOS dispersion), yielding information about the de-agglomeration efficiency of the dpi for a particular type of formulation. Additional reference measurements may be necessary like for instance with the carrier material (treated in the same way as the mixture), in order to make corrections for released fine lactose particles. However, it must be realised that finding the same size distributions for the aerosol (from a dpi) and the primary drug particles (from RODOS dispersion) does not necessarily mean that the complete dose has been fully de-agglomerated. It may be that only a fraction of the dose has been released because of incomplete discharge of the dose system and/or severe inhaler accumulation. The optical concentration of the aerosol or collected fine particle masses may yield valuable additional information in this respect. Over- and underestimation of certain size classes may occur in bimodal distributions and in situations where there is a fluctuating optical concentration in the aerosol, connected (and synchronised) with a varying size distribution. And of course, final conclusions should be checked with cascade impactor experiments.

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4. The aim of this thesis The aim of this thesis is to improve dry powder inhalation therapy by expanding the knowledge of, and challenging existing viewpoints on adhesive mixture performance in respect of fine particle detachment during inhalation. To facilitate the investigations in this respect, laser diffraction technique has been made applicable for size distribution measurement in the aerosol by design and development of a special inhaler adapter (Chapters 2 and 3). Laser diffraction technique increases the speed and accuracy with which developments and comparative evaluations can be made. Air classifier technology (ACT) has been developed as the most appropriate type of de-agglomeration principle for adhesive mixtures (and the other formulation types mentioned in Paragraph 2.3.2) because classifiers generate the most effective type of removal force, and because they provide the possibility to sustain the action of these forces during the entire inhalation manoeuvre. Several applications have already shown that air classifier technology can make the use of complex and expensive particle engineering processes for strongly cohesive materials superfluous (e.g. Zijlstra et al., 2002; Le Brun et al., 2002). As a practical use of this air classifier technology, the principle of carrier residue measurement has been introduced and developed to increase the accuracy of drug detachment experiments (Chapter 4).

In addition to the carrier surface properties, which are generally considered as the key parameter in adhesive mixture preparation for inhalation, the relevance of carrier bulk properties and mixing conditions on the drug-to-carrier interaction is shown in varies studies (Chapters 5 - 8). The possibilities of influencing the adhesive forces in the mixture (e.g. by choice of carrier, carrier payload and mixing conditions) and controlling the removal forces during inhalation (e.g. by type of classifier, inhalation time and flow rate) have resulted in the development of a force distribution concept (FDC), with which the effects of mixing and inhalation can be explained and visualised.

Finally, special classifier concepts have been developed with which the residence time for a certain powder mixture in the classifier can be controlled (Chapters 9 and 10). This provides the possibility to balance between the adhesive forces obtained during mixing and the removal forces during inhalation. Practically, this implies balancing between the time during which a patient is able to inhale with a certain (average) flow rate and the fine particle fraction that is obtained with that inhalation manoeuvre. By using two or more classifiers in a parallel arrangement, the inhaled drug dose can be increased considerably (Chapter 11). The design of a multiple classifier inhaler (Twincer) has been kept very simple, as a result of which it can be used as a disposable device for pulmonary administration of antibiotics (e.g. in CF) and vaccines.

Both the air classifier technology and the laser diffraction adapter have been successfully marketed (in the Novolizer multi-dose inhaler by Viatris, respectively as INHALER 2000 by Sympatec), and for the Twincer already more than five different drug formulations are currently in development. 5. References Allen, T., 1975. Particle size measurement. Chapman and Hall, London, pp 74-112. Amass, J.M., 1996. A study of drug carrier interactions in dry powder inhalers. Thesis, Dept. of Biol. & Chem.

Sci., University of Essex (UK), pp 170-189. Andersen, P.B., Hanssen, N.C.G., 1993. Which magnitude of inhaler resistance against airflow is preferred by

patients using dry powder inhalers. Eur. Resp. J. 6, 148S. Arnold, K., Grass, P., Knecht, A., Roos, R., Sluke, G., Thieme, H., Wenzel, J., 1995. Powders for inhalation.

US-patent 5,478,578. Arppe, J., Vidgren, P., Vainio, P., Tukiainen, H., Vidgren, M., 1994. Effect of hand-lung coordination on the

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