Patent Application
20240274037
The present invention generally relates to a filter apparatus for aerosol sampling, and more specifically to a specialized filter apparatus that mimics tracheobronchial (TB) deposition of pharmaceutical aerosols emitted by commercially available pharmaceutical inhalers, such as dry powder inhalers (DPIs) having filters to replicate deposition in specific regions of the human respiratory tract over a range of physiologically relevant flowrates.
Obtaining in vivo regional deposition measurements of inhaled aerosols directly in identified anatomical structures throughout the lungs remains an elusive goal, in part because it demands spatial and temporal imaging resolution that is beyond current abilities (1,2). Despite these challenges, Usmani et al. (3) demonstrated that targeted regional lung deposition of albuterol in the conducting airways improved therapeutic response relative to a similar dose depositing more deeply in the lung. Conversely, in situations involving occupational exposure to hazardous aerosols, localized deposition may lead to increased severity of adverse outcomes. Indeed, Schlesinger & Lippmann (4) pointed out that areas in the lung predisposed to increased particle deposition are also those areas most likely to see the onset of cancer. These examples highlight the importance of estimating regional deposition in the lungs, be it for optimizing therapeutic outcomes or for toxicology studies related to environmental aerosol exposure.
In the field of aerosol science, in vitro testing involves the use of physical experiments to determine various properties of aerosols. Such methods are commonly used in the early stages of product development for inhaled pharmaceutical aerosols to characterize device performance. Many have argued for improved in vitro methods that can accurately predict device performance in vivo to reduce the costs and complexities related to the drug development process (Byron, et al., 2010). Of specific interest is the prediction of the dose of drug depositing in different regions of the respiratory tract using in vitro methods.
Particle size and inhalation flow rate are the two external variables that most strongly determine the eventual deposition site of the inhaled aerosol particles. During tidal breathing of ambient aerosols at rest, inhalation flow rate is set by metabolic need, so that to assess risks of exposure to atmospheric aerosols the single most important variable is the size distribution of the aerosol. With knowledge of the size distribution of the particles in question, in silico deposition models can be used to predict the average dose that would deposit in the lungs or specific regions therein. Alternatively, a sampling device can be used to provide exposure estimates, assuming particles are captured with the same efficiency that they are expected to deposit in the body. Koehler et al. (5) describe a sampler that uses polyurethane foam as a filter medium to mimic deposition in the respiratory tract (inclusive of the nasal airway) for an average adult under light exercise breathing conditions (6). This sampler was shown to provide a good estimate of the internal dose of metals associated with exposure to welding fumes and is a step forward from common measurement methods that simply provide a measurement of ambient concentrations and particle size (7). This sampler was further developed by Koehler & Volckens (8) to include multiple filtration stages composed of polyurethane foams and nylon screens representative of the three regions of the respiratory tract, i.e., the extrathoracic, tracheobronchial, and alveolar regions.
While these samplers show great promise in their intended field of occupational hygiene, they are not suitable for use with pharmaceutical aerosols primarily because of their use of a constant flow rate that differs from the expected actual inhalation flow maneuver. It is also not expected that these samplers can be scaled and used directly to mimic regional deposition of pharmaceutical aerosols because the sampler is explicitly designed to mimic tidal breathing deposition efficiencies. For tidal breathing it is known that a significant amount of aerosol is exhaled (9), but for most single breath inhalers the exhaled dose is nearly eliminated due to deposition occurring during breath holding. Furthermore, assay of pharmaceutical ingredients may be difficult when using polyurethane and nylon substrates due to solubility of the substrate in solvents used for assay, resulting in interfering signals of dissolved pharmaceutical ingredients and dissolved substrate.
For estimating regional deposition of pharmaceutical aerosols, the inhalation flow rate generated through the device is of critical importance. As for all aerosols, this flow rate plays a major role in determining deposition efficiencies throughout the respiratory tract, but it is also the source of energy for aerosolizing and deagglomerating the powder contained in passive dry powder inhalers (DPIs). This relationship introduces a dependence of inhaler performance on inhalation maneuvers (flow rates vs. time) and further complicates the dependence of regional deposition on inhalation flow rate, necessitating direct testing of these inhalers using the inhalation maneuvers that patients are expected to perform during actual usage.
Current standards related to quality control and regulatory compliance for inhaled pharmaceutical aerosols, such as the standard set up shown in
In vitro quantification and qualification of orally inhaled drug products is extensively used as a measure of quality assurance and quality control (QA/QC): the same methods are also used to inform new product development and demonstrate equivalence of generic alternatives to previously approved products. When using these methods in the latter two capacities, researchers must understand their limitations and implement enhancements to improve the correlation of in vitro results to expected in vivo outcomes.
For QA/QC purposes, delivered dose uniformity (measured with a dose uniformity sampling apparatus) and the aerodynamic particle size distribution (APSD, measured by cascade impaction) can be determined using the methods described in the appropriate pharmacopeia documents [34, 35]. Delivered dose uniformity provides a measure of the entire delivered dose of active pharmaceutical ingredient (API) while the APSD relates to the anticipated in vivo deposition profile. However, the simplifications that make cascade impactor methods standardized, repeatable, and simple preclude meaningful estimation of the deposition site of the emitted API. In particular, the use of a highly simplified induction port results in coarse particles being characterized by the impactor. This problem is partially addressed by the determination of the fine particle fraction (FPF), which is generally defined as the fraction of the delivered dose with an aerodynamic diameter less than 5 μm. While this size is explicitly defined by the European Pharmacopeia [34], the United States Pharmacopeia allows the specific size to be defined by the individual product monograph; this combined with users making the convenient choice of using a specific impactor stage to define the FPF (thereby eliminating the need to interpolate between stages) results in FPF upper size limits ranging in the literature between 4.5 and 6.8 μm [36]. Also referred to as the respirable dose, the FPF is sometimes considered to be commensurate with the lung dose, but the deposition site (extrathoracic vs. lung) of particles around the FPF limit will depend strongly on the kinetics of the inhaler and the inhalation maneuver of the patient. Finally, the FPF alone omits any measure of the width (or geometric standard deviation [GSD]) of the particle size distribution: two aerosol clouds could have an identical FPF yet significantly different widths, resulting in differing deposition patterns for the two aerosol clouds in vivo [36].
Currently, the preferred experimental setup designed to provide quantitative in vitro predictions of inhaler performance involves placing a cascade impactor downstream of an extrathoracic airway model: if the extrathoracic airway model accurately mimics extrathoracic deposition, then the aerosol classified by the impactor represents the total lung dose and can be divided into regional deposition estimates by applying in silico deposition models (10). While in theory, this methodology can generate helpful insights into inhaler performance, it has many disadvantages. First, cascade impactors must operate at a constant flow, which makes testing inhalers with realistic breath profiles more complicated. While this problem can be overcome in large part by using mixing inlets and providing a bias flow to the impactor in order to generate time-varying flow through the inhaler (11-13), upper limits on inhalation flow rates are still imposed. More important is the issue of the resources required to perform cascade impactor experiments. If an experimenter is using the Next Generation Impactor (NGI), the aerosol is classified into seven size stages (eight if the pre-separator is used); each stage must be collected and assayed individually to determine the mass of aerosol in each size bin. If regional deposition estimates are desired, this size distribution data must then be coupled with a regional deposition model calculation before a final result can be reached. This procedure is so involved that a well trained laboratory analyst can only complete 1-2 cascade impactor measurements in a single work day, which precludes high volume testing.
The determination of aerodynamic particle size distributions using cascade impactors is standard practice (Hickey, 2004). Various cascade impactor designs have been described in the literature (U.S. Pat. No. 3,795,135, 1974; U.S. Pat. No. 5,343,767, 1994; U.S. Pat. No. 6,453,758 B1, 2002). Common limitations among existing designs of cascade impactors are inter-stage losses, costs associated with operator time and solvent usage, and the requirement for operation at a constant flowrate (precluding the use of realistic inhalation patterns). Methods have been proposed to utilize mixing inlets such that a realistic inhalation profile can be delivered through an inhaler upstream of a cascade impactor while still satisfying the requirement for a constant flowrate in the impactor itself. The introduction of a mixing inlet, however, may lead to increased losses and significantly complicates experimental testing procedures. Standard practice involves the definition of a fine particle dose (or fine particle fraction), representing the portion of particles in an aerosol with a diameter below a specified limit (delineating fine and coarse particles) based on measured aerodynamic particle size distributions. Fine particles are often thought of having a size allowing for penetration and deposition in the airways. However, as particle deposition depends on flowrate as well as particle size, this treatment masks the effects that changing flowrates over the course of an inhalation can have on respiratory tract deposition, and there is no consensus on the appropriate size range for such fine particles (Clark & Borgstrom, 2002).
Flowrate effects on regional lung deposition can be captured using mathematical models of deposition in the respiratory tract. One-dimensional lung deposition models using particle size and inhalation flowrate as input parameters have been developed to predict regional deposition (ICRP, 1994), and can be used in conjunction with cascade impactor measurements to predict respiratory tract deposition. However, the correlations used in mathematical models are valid only for a specified range of parameters, and most existing models prove inadequate in predicting extrathoracic deposition a priori. Fortunately, accurate measures of extrathoracic deposition can be obtained using in vitro models of the mouth-throat (e.g. the Alberta Idealized Throat, Copley Scientific). The combination of realistic mouth-throat models placed upstream of cascade impactors may provide a method for estimating regional deposition, but is still subject to the limitations associated with cascade impactor measurements and the incorporation of time-variant inhalations.
The first and simplest step in improving the pharmacopeial methods is to use an induction port that mimics extrathoracic deposition in the human population [36, 37] such as the Alberta Idealized Throat (AIT). When such an induction port is used, the dose depositing in the cascade impactor is representative of the expected lung dose and can be characterized by its MMAD and GSD, allowing further analysis using advanced in silico methods to predict regional deposition of this dose within the lung [10]. Others [38, 39] have proposed using impactors with fewer stages, called fast screening impactors (FSI), for abbreviated impactor measurements (AIM) to fractionate this in vitro lung dose directly into FPF and extra-fine particle fractions to reduce the resource burden associated with full resolution impactor measurements. These fractions cannot be directly associated with in vivo deposition in regions of the lungs since deposition site/proportions will depend on the inhalation profile generated by the patient. Further, this approach prohibits accurate estimation of regional lung deposition because the resolution of the distribution is lost.
An additional step in improving the pharmacopeial methods involves generating realistic inhalation patterns through the inhaler and extrathoracic model. This is most important for dry powder inhalers, where the inhalation provides the energy source for aerosolization and deagglomeration of the powder, but is also useful for testing delayed inhalation with pressurized metered dose inhalers (pMDIs) and soft mist inhalers (SMIs) for which poor patient coordination could reduce device performance. Cascade impactors, however, must be operated at specified, fixed flow rates for which they are calibrated. Additional equipment (mixing inlet, makeup flow supply line with flow monitoring, and breath generator) can be incorporated into the cascade impactor apparatus to enable time varying inhalations to be performed through the inhaler while the cascade impactor remains under constant flow [40, 12]. This is a significantly more complicated experiment to conduct and also sets an upper limit on flow rate, imposed by the maximum flow rate for which the cascade impactor has been calibrated.
Filters provide an alternative method for classifying aerosol particles. The theory of operation for aerosol filters is described in the literature (Davies, 1973). Filters are typically designed to clean an airstream by removing as many particles as possible. For such filters, the filtering efficiency is ideally 100% for all particles, though in practice there is a most-penetrating particle size for which the filter efficiency is a minimum.
Polyurethane foam filters have been used to characterize ambient aerosols (Brown, 1980; Kuo, et al., 2005; Koehler & Volckens, 2013; Sleeth, Balthaser, Collingwood, & Larson, 2016; Tomyn, Sleeth, Thiese, & Larson, 2016; Courbon, Wrobel, & Fabries, 1988: Chung, Aitken, & Bradley, 1997; Vincent, Aitken, & Mark, 1993; Chen, Lai, Shih, & Yeh, 1998). Empirical models describing the penetration of particles through such filters are available and known in the art (Clark, Koehler, & Volckens, 2009). Such models have been used in the development of ambient aerosol samplers that attempt to replicate respiratory tract deposition based on the various deposition models. The state of the art involves the replication of ICRP (ICRP, 1994) deposition efficiency curves for extrathoracic, tracheobronchial, and alveolar deposition (Koehler & Volckens, 2013). The current art advocates the use of constant flowrates, precluding the use of time variant inhalation profiles, and relies on mathematical models to model extrathoracic deposition efficiency curves. These limitations prevent the use of such filters in characterizing therapeutic aerosols delivered from inhalation devices, as time-variant inhalation profiles and complicated deposition behavior in the extrathoracic region are observed when using such devices.
Recently, Tavernini et al. (14) presented a filter that could be used to provide regional deposition estimates without the use of cascade impactor measurements. This filter was shown to have a filtration efficiency similar to predicted tracheobronchial (TB) deposition for a wide range of particle sizes and flow rates. The authors suggest the filter can be used downstream of an extrathoracic airway model and upstream of an absolute filter, the latter collecting any aerosol passing through the TB filter. The two filters separately capture the expected tracheobronchial and alveolar dose (assuming that there is no exhaled dose). It was not shown whether this filter functions as intended with pharmaceutical powder aerosols, since Tavernini et al. examined only the filtration efficiency of their TB filter sampling from a chamber containing laboratory-generated oil aerosol.
For DPI aerosols, a complicating issue is the electrostatic charge state of the particles exposed to the TB filter. It is known that particles emitted from DPIs can carry considerable bipolar charge (15). Charged particles were not considered in the development of the Tavernini et al. (14) TB filter. Traditional methods for evaluating inhaler performance also do not take particle charge into consideration, and the effect that aerosol charge has on deposition in the lungs remains an area of study. Majid et al. (16) implemented electrostatic deposition in a stochastic lung deposition model and found TB deposition was enhanced by approximately 10% for one-micron and sub-micron sized particles inhaled at 30 L/min, while enhancement in the alveolar region was approximately 20%. The authors also compared deposition calculated at 18 and 50 L/min and found that at higher flow rates the enhancement was reduced in the TB region but increased in the alveolar region. Flow rates generated with the use of single breath inhalers are often higher than the authors studied, indicating TB deposition enhancement may be insignificant for those applications. Similar results were reported by Koullapis et al. (17), who employed computational fluid and particle dynamics to investigate the effects of particle charge on deposition in a realistic geometry of the human respiratory tract inclusive of the mouth-throat to the seventh generation of the lungs. At a flow rate of 60 L/min (the highest investigated in that study) deposition in the domain was nearly independent of particle charge for 2.5 μm particles and increased from about 20% to 30% for 1 μm particles. The authors show that for a flow rate of 15 L/min the deposition enhancement occurs primarily in the extrathoracic region, but they do not specify if this is the case at higher flow rates. Nevertheless, these two studies show that for pharmaceutically relevant particle sizes and inhalation flow rates, deposition in the tracheobronchial airways is largely independent of particle charge for charge levels and flow rates expected with pharmaceutical aerosols.
By contrast, particle charge could have major effects on the filtration efficiency of the Tavernini et al. (14) TB filter, since it is comprised of two layers of stainless-steel mesh. The square pores of the steel mesh are ˜30 μm wide, so that particles passing through the filter come into close proximity to the filter fibers. These distances are an order of magnitude smaller than even the final generation of the alveolar region, while the tracheobronchial region terminates with bronchioles approximately 600 μm in diameter (18). Since the fibers of the Tavernini et al. (14) TB filter are electrical conductors, image charges will readily form in the presence of a nearby charged particle, causing large electrostatic forces and thus higher filtration efficiency. While particle charge may not have a significant effect in the tracheobronchial region of the lungs in vivo, it may cause significant departure from the designed filtration efficiency of the Tavernini et al. idealized TB filter. This issue could be resolved by neutralizing the aerosol before exposing it to the TB filter.
Accordingly, there remains a need in the art for a filter apparatus that mimics tracheobronchial (TB) deposition of pharmaceutical aerosols emitted by commercially available pharmaceutical inhalers, of which, dry powder inhalers (DPIs) and pMDIs are subtypes.to provide an in vitro method to rapidly evaluate regional lung doses delivered by pharmaceutical inhalers directly using variable inhalation profiles without any additional equipment or changes to the experimental configuration. There additionally remains a need for the ability of such a filter-based apparatus (FBA) to provide regional deposition estimates in agreement with an in silico model that uses the results of traditional cascade impactor measurements as input to generate regional lung deposition estimates for a selection of pMDI devices.
The present invention provides a regional lung deposition system including a throat simulation device, an electrostatic neutralizer, and a first filter positioned within a first filter housing. The electrostatic neutralizer is positioned downstream of an in fluid communication with the throat simulation device. The filter housing is positioned downstream of and in fluid communication with the electrostatic neutralizer. The first filter that is positioned within the filter housing is downstream of the throat simulation device and the electrostatic neutralizer. The regional lung deposition system mimics tracheobronchial (TB) deposition of pharmaceutical aerosols emitted by commercially available pharmaceutical inhalers to provide an in vitro method to rapidly evaluate regional lung doses delivered by pharmaceutical inhalers directly using variable inhalation profiles. The present invention additionally provides a method of providing regional deposition estimates in agreement with an in silico model that uses the results of traditional cascade impactor measurements as input to generate regional lung deposition estimates for a selection of pharmaceutical inhalers.
The present invention is further detailed with respect to the following drawings that are intended to show certain aspects of the present of invention, but should not be construed as limit on the practice of the invention, wherein:
The present invention has utility as a filter apparatus that mimics tracheobronchial (TB) deposition of pharmaceutical aerosols emitted by commercially available dry powder inhalers (DPIs) to provide an in vitro method to rapidly evaluate regional lung doses delivered by pharmaceutical inhalers directly using variable inhalation profiles without any additional equipment or changes to the experimental configuration. The present invention additionally has utility as a method for using such a filter-based apparatus (FBA) to provide regional deposition estimates in agreement with an in silico model that uses the results of traditional cascade impactor measurements as input to generate regional lung deposition estimates for a selection of pharmaceutical inhalers.
It is to be understood that in instances where a range of values are provided that the range is intended to encompass not only the end point values of the range but also intermediate values of the range as explicitly being included within the range and varying by the last significant figure of the range. By way of example, a recited range of from 1 to 4 is intended to include 1-2, 1-3, 2-4, 3-4, and 1-4.
According to embodiments, the inventive regional lung deposition system 50 that mimics tracheobronchial (TB) deposition of pharmaceutical aerosols includes an Alberta Idealized Throat 40, an electrostatic neutralizer 52, and a filter housing 51 having a filter 20 that mimics TB deposition. The inventive system 50 is able to fractionate in vitro lung dose into tracheobronchial and alveolar regional deposition estimates for dry powder inhaler aerosols to provide an in vitro method to rapidly evaluate regional lung doses delivered by pharmaceutical inhalers directly using variable inhalation profiles.
A custom bipolar ion source 52 is designed to allow neutralization of inhaled pharmaceutical aerosol formulations prior to their passage through a filter 20 that mimics TB deposition. Since the intent of the present TB filter device 50 is to enable direct testing of inhalers with physiologically realistic inhalation profiles, the ion source 52 must be integrated into the TB filter apparatus 50 to rapidly neutralize aerosol after it exits the extrathoracic airway model before traversing the TB filter media 20. The internal volume is minimized so that the aerosol emitted by the inhaler is exposed to the TB filter 20 at an appropriate time in the breath inhalation maneuver.
An estimate for the maximum required ion concentration can be made by considering one of the most highly charged dry powder inhalers, the Pulmicort Turbuhaler (15). A total charge is assigned to each bin of the particle size distribution reported by Ruzycki et al. (10) by multiplying mass by charge-to-mass ratios reported by Yu et al. (15). The number of ions needed to neutralize the aerosol can then be determined by summing the total charge and dividing by the elementary charge. While Ruzycki et al. reported mass as a function of particle size for five actuations of a 200 μg per dose budesonide Turbuhaler, the present invention does not divide the number of ions by five in order to extrapolate a worst case ion requirement. An ion concentration can then be determined by identifying the volume of the aerosol bolus. The emission kinetics of DPIs remains an area of study but depends on the device type, powder formulation, and inhalation maneuver (19). For the fastest emptying device studied by Dorosz et al., the bulk of aerosol appears to be emitted in as little as 0.4 s after only 0.1 s of inhalation, giving an aerosol bolus of about 400 mL for the flow profile used by Ruzycki et al. (10). Assuming that the aerosol is evenly distributed throughout the bolus leads to an estimate for an ion requirement of approximately 1.9×108 ions/mL to neutralize the aerosol. Note that this value represents only the ions required to neutralize the aerosol. In actuality, an excess of ions is required since neutralization is not expected to be 100% efficient.
The ion requirement found with this approach happens to be very similar to the requirement reported by Hinds & Kennedy (20), who demonstrate that the radioactive source required to provide such an ion concentration is enormous. However, the authors also demonstrate the ability of a custom-made corona discharge ion generator to almost entirely neutralize a unipolarly charged aerosol in their application. Their ion generator employs direct current applied to preferentially sized electrodes in order to generate the desired polarity of corona. For the present application, 60 Hz alternating current (A C) is used to generate both polarities of ions, albeit at opposite phases of each A C cycle rather than simultaneously.
Bearing in mind the above considerations, embodiments of the present disclosure provide a custom ion generator 52 to incorporate into the TB filter housing 51. A simple A C high-voltage source is built using a 10 kV transformer (2721-628G: Allanson International, Markham, Canada) which uses standard 120 V 60 Hz input to excite the secondary coil. Benchtop experimentation leads to an electrode design in which one high-voltage lead 54 is attached to an internally serrated stainless-steel ring 56, with the second high-voltage lead attached to a cylinder centered within the ring 56. Corona discharge is observed to form on each point of the serrated ring 56, emitting ions and creating an annular region of high ion concentration. Passing charged aerosol through this annulus provides the desired neutralization. To do so, however, necessitates redesigning the TB filter housing of Tavernini et al. (14) from a simple conical expansion (shown in
This design change is also inspired by the potential for the laryngeal jet (21,22) to persist into the TB filter housing 51 and thereby cause non-uniform face velocity and thus off-spec filtration due to the previously mentioned dependence of filtration efficiency on face velocity. By introducing an abrupt expansion in the filter housing 51, the jet is forced off the anterior wall rather than remaining attached due to the Coanda effect. After rapid expansion from the extrathoracic model outlet, the free jet impinges on the top of the annular core to further distribute the energy of the jet into the bulk flow, allowing relatively uniform flow down the annular gap towards the TB filter. Since the jet impinges on the top of the annular core it is assumed that the largest particles remaining in the flow will be deposited here, so a collection dish is incorporated into the design at this location (herein referred to as the flow disruptor plate).
Regional deposition estimates of four commercially available DPIs delivering different active pharmaceutical ingredients (APIs) are obtained. Each device differs in either its metering mechanism, use of carrier particles (or not), or resistance to flow through the device. Inhalers tested included Onbrez Breezhaler® (Lot SWN96; Novartis Pharmaceuticals Canada Inc., Dorval, Canada), Flovent Diskus®) (Lot W94G; GlaxoSmithKline Inc., Mississauga, Canada), Pulmicort Turbuhaler®) (Lot PAVE: AstraZeneca Canada Inc., Mississauga, Canada), and Asmanex Twisthaler®) (Lot T032728: Merck Canada Inc., Kirkland, Canada). Details of each device are listed in Table I. Inhalation profiles are different for each device because of differences in device resistance. Delvadia et al. (23) measured in vivo profiles generated through different resistance inhalation flow cells and provided the method for selecting appropriate inhalation profiles for in vitro inhaler performance testing are used herein. The inhaler resistance defines the peak inspiratory flow rate (PIFR). The 50th percentile PIFRs are used and expected to be generated after professional instruction (Equation 5 in [23]), except in the case of the Breezhaler®, which by this formula would be tested with a PIFR higher than the maximum allowed when using the NGI. The time at which the PIFR is generated (IPIFR) and the total volume inhaled (V) are independent of inhaler resistance and are therefore constant for all devices. Furthermore, tPIFR=0.49 s and V=2.7 L is used, which are the 50th percentile values for mixed-gender data (23). The shape of the breath profile is similar for all inhalers and is made up of two sinusoidal quarter waves. The first quarter wave is the “ramp up” period where inhalation flow rate quickly accelerates to PIFR. The second quarter wave is the majority of the breath from tPIFR to the end of the breath, ttotal, which is set by the total volume inhaled (Equation 10 in [23]). The unique inhalation parameters for each inhaler are also listed in Table I.
Two different experiments are performed (with n=5 measurements each) for each inhaler with these inhalation profiles. One experiment involves direct measurements of the TB filter device performance downstream of an Alberta Idealized Throat 40. The other involves cascade impactor measurements to characterize the inhaler particle size distribution for use in a regional deposition model (see next section).
Regional lung deposition estimates for each inhaler are established based on cascade impactor measurements, the results of which are used as input for an in silico regional deposition model. The method used very closely follows the method used by Ruzycki et al. (2020). For the upstream mouth-throat inlet, the Alberta Idealized Throat (AIT) 40 is used. The inhaler being tested is coupled to the inlet of the AIT 40 with a 3D-printed custom adapter unique to that inhaler type, providing an airtight seal and ensuring consistent and proper orientation of the inhaler relative to the AIT 40. A Mixing Inlet (MSP Corporation, Shoreview, M N, USA) downstream of the AIT enables use of time-varying inhalation profiles through the inhaler/extrathoracic model while providing a constant flow of 100 L/min (volumetric) to an NGI with a pre-separator stage (MSP Corporation). An inhalation profile is generated by a custom-built breathing machine that draws from the supply to the mixing inlet. The resulting deficit in makeup air causes the NGI to pull air through the AIT 40 and inhaler. Flow rates are monitored using mass flow meters (4040 and 4043: TSI Incorporated, Shoreview, M N, USA) and the volumetric flow generated through the inhaler is calculated as described by Ruzycki et al. (10) using a LabVIEW program (LabVIEW Professional Development System 2017; National Instruments, Austin, TX, USA). This experimental configuration is demonstrated in
The cascade impactor data obtained from the above experimental configuration represent the total lung dose, provided that the appropriate fraction of particles is removed in the AIT 40. One mechanism that can cause decreased collection efficiency in the AIT is particle bounce, where a particle which would otherwise be deposited on the extrathoracic airway wall instead rebounds back into the flow. This mechanism artificially inflates the mass of API reaching the lungs but is also difficult to detect. Zhou et al. (24) investigated the difference in particle collection of liquid droplets vs. solid particles with and without coating the walls of extrathoracic models with silicone lubricant fluid. They suggest that for particles<10 μm, bounce should not be a concern when wall coatings are used. However, for particles>10 μm some bounce may occur and is evident by the plateau in deposition observed for polystyrene latex particles around 90% efficiency, which does not occur for oleic acid droplets. While few of these relatively massive particles may bounce through the AIT, they carry significant mass. Conveniently, when using the pre-separator at 100 L/min the cut-point is 10 μm, meaning that particles collected in the pre-separator are 10 μm and larger and thus could be those subject to bouncing through the AIT. Therefore, in the present work any mass detected on the pre-separator is added to that collected in the AIT and are referred to as the Upper Respiratory Tract fraction (URT). Thus, only the sized mass collected from the stages of the NGI are considered to contribute to the lung dose in this work.
Regional lung deposition estimates from NGI data are obtained using the regional deposition model presented by Javaheri et al. (25), which uses deposition efficiency equations for inertial impaction (26), sedimentation (27,28), and diffusion (29) to estimate total deposition in each generation of a symmetric lung model (30). In this lung model the trachea is represented by generation 0, the tracheobronchial (or conducting airway) region by generations 0-14, and the alveolar region by generations 15-23. Extrathoracic deposition is not present in the in silico model, since it is already accounted for in the URT deposition noted above. Particle diameters used for evaluating deposition efficiencies are the geometric centers of the stage-bounding 50% collection efficiencies. Particle sizes are assumed to be constant, and hygroscopic effects were neglected. The model uses a constant flow rate to propagate a constant density aerosol through the lung generations; since the aerosol is expected to interact with the TB filter at approximately the PIFR, the breath time in the in silico model is adjusted so that the resulting average flow rate (inhaled volume divided by inhalation time) is equal to the PIFR used for that inhaler. A 10 second breath hold and 7 second exhalation are used for all inhalers (although the latter has little effect on our results since the breath hold reduces the exhaled aerosol to negligible amounts). Total inhaled volume is the same 2.7 L used in the TB filter runs described above. Tracheobronchial deposition is then estimated by summing the calculated particle deposition in generations 0-14. Alveolar deposition is taken to be the total NGI mass less the TB estimate, as this approach best aligns with the experimental method, in which there is no exhaled fraction of aerosol.
These estimates provide target regional lung deposition values for the TB filter device to achieve. Because of variability in the conception of different regional deposition models and inter-subject variability known to exist in vivo, these targets are considered approximate.
Using equations governing the theory of deposition of particles in filters and the respiratory tract a filter 20 with properties able to mimic deposition in different regions of the respiratory tract is provided. According to embodiments of the present disclosure, the inventive filters 20 mimic deposition in different regions of the human respiratory tract over the wide range of flow rates, such as those seen during inhalation from an inhaler, where the flow rate starts at zero, reaches a maximum and then decreases until inhalation stops at the end of a breath. Filters 20 of the present disclosure approximately match curves of in vivo trachea-bronchial deposition versus particle size for the range of typical flow rates expected in vivo, as seen in
Filters 20 according to the present disclosure, such as that shown in
Important filter properties for fibrous filters include the volume fraction of fibers a, the fiber diameter df, the filter thickness t, and the single fiber deposition efficiency EΣ. EΣ represents a deposition fraction per unit length of fiber and is a function of multiple mechanisms. The efficiency of a fibrous filter, E, can be calculated according to Equation 1, given knowledge of the above parameters.
The focus of single-fiber theory involves characterizing the single-fiber deposition efficiency EΣ. Various expressions are available in the literature for a number of mechanisms, the most commonly used of which are summarized by Hinds (1999).
Deposition due to impaction E1, caused by the inertia of a particle leading to deviation from fluid streamlines near the fiber surface, is given as a function of particle Stokes number Stk, the Kuwabara hydrodynamic factor Ku, and a factor J accounting for filter properties, as calculated by Equation 2.
The particle Stokes number (Stk) can be expressed in terms of the particle relaxation time t, the face velocity U0, and the fiber diameter, as calculated by Equation 3.
The Kuwabara hydrodynamic factor is a function of fiber volume fraction, as calculated by Equation 4.
The factor J is a function of the ratio of particle diameter to fiber diameter R and the fiber volume fraction, as calculated by Equation 5.
This expression is valid for R<0.4. An approximate value of J=2.0 is be used when R exceeds this value.
Interception occurs when a particle following a streamline comes close enough to a fiber such that the particle's radius causes it to deposit. As such, it is a consequence of the finite size of particles. Deposition due to interception ER can be estimated according to Equation 6.
Deposition due to diffusion ED occurs when small particles collide with filter fibers through stochastic random motion. Deposition due to diffusion ED is calculated according to Equation 7.
Pe is the Peclet number, whose value depends on the particle diffusion coefficient D and is calculated according to Equation 8:
An additional deposition mechanism accounts for the interception of diffusing particles and is calculated according to Equation 9.
Deposition due to gravitational settling EG depends on the orientation of the airflow relative to gravity. In situations in which the flow velocity and the settling velocity are in the same direction, EG is expressed according to Equation 10.
Here, G is the ratio of settling velocity to face velocity according to Equation 11.
The overall filtration efficiency EΣ can be approximated by a summation of the above mechanisms according to Equation 12.
Control of fiber diameter, face velocity, filter thickness, and fiber volume fraction allow for the design of filters that replicate deposition in various regions of the respiratory tract.
Embodiments of the present disclosure further provide filters for different regions of the lung in addition to the trachea-bronchial and alveolar regions described above. According to some embodiments, the inventive filters are used for testing inter-subject variability in deposition in the different regions of the lung by being designed to mimic the lower and upper limits of expected in vivo regional lung deposition as opposed to mimicking the average regional deposition. Embodiments of the present disclosure also provide filters that mimic regional deposition in children or infants, with breath simulation supplying realistic pediatric inhalation profiles.
According to further embodiments, the inventive apparatus 50, shown in
On each day of experimentation the ambient conditions are recorded. Pressure is recorded from one of the flowmeters while the apparatus is stagnant. Temperature and relative humidity are measured using a digital hygrometer/thermometer (MI70 Measurement Indicator and HMP75B Humidity and Temperature Probe: Vaisala, Vantaa, Finland). Ambient conditions (temperature, pressure, and relative humidity) ranged from 21-23° C., 90.7-94.5 kPa, and 13-26%.
For all experiments the interior surfaces of the clean, dry AIT 40 are coated with silicone release spray (Molycote 316; Dow Corning, Midland, MI, USA). This spray is also used on all stages of the NGI impactor plates and the pre-separator impactor plate. The propellant/solvent is allowed to evaporate for a minimum of 15 minutes before the equipment is assembled. A specialized 3D-printed adapter for the inhaler being tested is affixed to the AIT 40 inlet and the appropriate inhalation profile loaded into the breathing machine. For NGI runs, the vacuum pump pulling flow through the impactor is turned on and the flow set to 100 L/min. The makeup air supply to the mixing inlet is then adjusted to match, producing zero flow at the AIT 40 inlet. For runs with the TB filter device, no flow monitoring/matching is required. The only additional preparation step with the TB filter setup is to place a small amount of solvent on the flow disruptor plate. Using the assay solvent here eliminates bounce (similar to a liquid impinger) and makes for simple and effective collection of any API landing on this plate.
With the apparatus assembled and ready for the inhaler to be tested, the inhaler is primed and loaded following the directions in the leaflet supplied with the inhaler. When the dose is ready to be inhaled, the inhaler is brought to the AIT 40 and held firmly in place. The breathing machine is actuated and a single inhalation maneuver is generated through the device. For TB filter 20 runs, the neutralizer 52 is manually switched on just prior to generation of the inhalation and switched off when the machine is stopped. The inhaler is then removed from the adapter at the AIT 40 inlet. Valves in the system piping allow the piston of the breathing machine to return to the starting position with the resulting exhalation being directed out of a release port rather than through the system. In the case of the Breezhaler, the capsule is inspected to ensure it is completely empty: if powder remains, the inhalation procedure is repeated as per the product leaflet instructions. Subsequent doses (as per Table I) are delivered following the same procedure, in accordance with all product leaflet instructions.
When the predetermined number of doses is been delivered, the collection surfaces of the components are washed with HPLC-grade methanol ensuring API is collected in full. The AIT 40 is washed twice with 10 mL, NGI collection stages are washed with 5 mL, and the pre-separator and flow disruptor plate are also washed with 5 mL except in the case of the Twisthaler runs, where 10 mL is used due to increased deposition. The stainless-steel TB filters are washed twice with 10 mL of methanol, while the fibrous alveolar filters are washed with a total of 25 mL divided into three parts. The mass of API contained in each sample is quantified via ultraviolet light absorbance relative to a series of prepared standards. Absorption is measured using a U V-Vis spectrophotometer (Cary 8454: Agilent, Santa Clara, C A, USA) at 236, 243, 248, and 260 nm for fluticasone propionate (Diskus), budesonide (Turbuhaler), mometasone furoate (Twisthaler), and indacaterol maleate (Breezhaler), respectively. For each inhaler, the above procedures are repeated five times for each configuration to provide a measure of repeatability. Statistical comparisons of regional deposition estimates are performed using the two-tailed, unpaired student's t-test in Microsoft Excel 2016 (Microsoft, Redmond, W A, USA).
The results of characterizing the inhalers with the NGI are given in Table II. Similarly, the results of testing the same inhalers with the regional deposition filter apparatus are given in Table III. In silico regional deposition estimates arising from NGI data are presented in Table I V alongside the in vitro regional deposition estimates from the TB filter and neutralizer apparatus.
Since the masses collected on the flow disruptor plate and pre-separator are similar, it is safe to assume the aerosol classified by the NGI is representative of the aerosol travelling through the filter. Indeed, the total lung dose is the same for each experimental configuration (p>0.05). Excluding the flow disruptor and pre-separator dose from the tracheobronchial deposition estimates results in a pure comparison of the TB filter's ability to mimic the in silico deposition model. Comparison of the regional lung deposition between the two approaches can be done in terms of mass, fraction of recovered dose, or fraction of lung dose estimated to deposit in each location; all three such comparisons are shown in
The above comparisons of regional deposition estimates show that the TB filter 20 satisfactorily mimics the estimated tracheobronchial deposition given by a regional deposition model applied to the sized mass as measured by cascade impaction. In some cases, the TB filter appears to capture slightly more mass than predicted by the regional deposition model, a discrepancy that could be due to incomplete neutralization or else a tendency of the filter to capture slightly more mass than is predicted by the model used here. Given the variations in predictions by different in silico models (14) and the large in vivo variability, we believe these differences are acceptable.
As expected, Table I V shows there is generally no significant difference in URT deposition since the AIT is used in both experimental configurations and the inhalation pattern is the same. Though URT is statistically different for the Asmanex Twisthaler®), the small difference is aligned with the statistically significant difference in total recovered API: these differences are not detected for any other inhaler.
While the ability of the custom ion generator 52 to neutralize the test aerosol is not directly measured, it is indirectly assessed by comparing the mass of budesonide recovered from the TB filter in runs where the neutralizer was not powered. These experiments are merely exploratory in scope, so a full data set is not presented here. However, for a Pulmicort TurbuhalerR) (a different individual inhaler device than presented in Tables 2-4) the TB filter mass decreased from 101.4±16.7 μg when the neutralizer was not used to 47.4±14.0 μg when the neutralizer was powered at 10 K V (for 3 actuations per run, n=5 runs). This represents a reduction from 23.3±2.1% of the recovered dose to 10.3±2.1%.
Also being monitored during those exploratory experiments is the amount of API collected from a swab of the interior walls upstream of the TB filter in the neutralizing region. With the neutralizer off, 19.6±4.7 μg was recovered from the walls, but with the neutralizer at 10 kV this increased to 53.8±17.3 μg, or an increase from 4.5±0.9% of recovered mass to 11.5±1.7%. Further experimentation indicated that powering the neutralizer with 5 kV reduced wall losses to only 3.3±0.5% of the recovered dose while still adequately neutralizing the aerosol. This finding led to using the neutralizer with a potential difference of 5 kV in all further experiments. Wall losses are not monitored further as they are considered acceptable. Indeed, total recovery is consistent between all runs when comparing NGI runs and those using the TB filter device, indicating that the wall losses are similar in the two experiments.
While the total recovered dose is not of primary interest, in this research it serves as a measure of the integrity of the experiment being performed. For the Breezhaler®, Diskus®), and Twisthaler® devices total recovery as a percent of the label claim was 79.5±4.1%, 79.8±3.3%, and 87.7±3.1%, respectively. However for the Turbhuhaler®), the recovery dropped to 50.4±3.3% of label claim. The low recovery when using the Turbuhaler® is surprising when considering the satisfactory recovery of the other inhalers. It is not, however, without precedent. Hill & Slater (31) found an emitted dose from the Turbuhaler® of 46.0±14.6% of the label claim, while Hindle & Byron (32) found similarly low recovery and showed significant variability between devices.
The API collected from the pre-separator and the flow disruptor plate allows for some interesting speculation. The pre-separator is a carefully designed impactor stage intended to operate at a constant flow rate and to provide a sharp change in deposition efficiency (otherwise known as a cut-point). The flow disruptor plate is substantially different in that it operates under variable flow rate defined by the inhalation pattern used. For this reason, a calibrated cut-point cannot be defined. Despite this fundamental difference, a similar amount of API was collected on the pre-separator and the flow disruptor plate for each inhaler. While the pre-separator was always operated at 100 L/min, the flow disruptor plate saw peak flow rates as low as 60 L/min while still collecting a similar amount of API. Although it cannot be stated with certainty, it is likely that particles are sized similarly by the two and that the particles collected on these surfaces have bounced through the AIT and are relatively massive. This is supported by physical observations made during preliminary experiments with the Twisthaler® before solvent was used on the flow disruptor plate. Originally, the disruptor plate is sprayed with the same silicone release spray used to reduce bounce in the AIT and on the impactor stages. Despite the use of this spray in the AIT and on the flow disruptor plate, exceptionally large particles are observed on the TB filter after experimentation. Since these particles are visible to the naked eye, it was not believed that they could be avoiding contact with the flow disruptor. However, to investigate this possibility, solvent is placed in the flow disruptor plate and the experiment is repeated. The effects of this change are immediately visible as these large particles re no longer observed on the TB filter surface but instead are found in a pile in the solvent on the flow disruptor plate.
The notion that particle bounce is responsible for API being found on the flow disruptor plate and in the pre-separator is also supported by the variability observed between API recovery from the AIT versus the variability in URT deposition. Recall that URT deposition here is the sum of API recovered from the AIT and from the pre-separator or the flow disruptor plate (depending on the experimental configuration). Variability in URT deposition is reduced when compared to AIT deposition (see Table II, Table III, and Table I V) for the Twisthaler® and Breezhaler®). The amount of bounce occurring in the AIT for a given flow rate will depend on the AIT coating quality and the extent of deagglomeration occurring in the inhaler. For example, if slightly less coating is applied in a particular run more bounce may occur; however, if the inhaler powder deagglomerates well, the amount of bounce may decrease. Regardless of this interplay between coating quality and inhaler performance, if the pre-separator or flow disruptor plate is catching these bouncing particles, there should be an inverse relationship between the mass collected on these surfaces and in the AIT. That is, if particles bounce through the AIT (thereby reducing mass collected in the AIT) they should be recovered from the pre-separator or flow disruptor, where mass will be increased. This trend is confirmed for the Twisthaler® and Breezhaler® by plotting the mass recovered in the AIT against the mass recovered from the pre-separator or flow disruptor plate, as shown in
The converse is observed for the Turbuhaler (
These trends suggest that without further knowledge of the aerosol size distribution, the experimenter could identify whether mass landing on the flow disruptor plate should be considered URT or TB deposition by comparing the relationship between AIT and flow disruptor plate deposition
The ability of a filter mimicking tracheobronchial deposition to collect the correct fraction of particles exiting the Alberta Idealized Throat has been demonstrated with commercially available dry powder inhalers. This filter provides a fast method to estimate in vivo regional lung deposition using in vitro inhaler testing. Electrostatic charge on the aerosol particles has the potential to cause the TB filter to remove more particles than intended, but this issue is resolved by using an alternating current corona discharge ion generator integrated into the filter housing to quickly and adequately neutralize the aerosol just upstream of the filter. This apparatus can be used directly downstream of the AIT and has a small enough internal volume to allow direct testing of inhalers with the inhalation maneuvers expected to be performed by patients using the device. The apparatus presented here is useful as a tool for inhaler product research and development.
Regional deposition estimates are obtained for three different commercially available pMDI devices: Asmanex HFA (Lot U019924, Merck & Co., Inc., Whitehouse Station, N J, USA), QVAR (Lot GUL037A, Valeant Canada L P, Laval, Canada), and Ventolin HFA (Lot RX4E, GlaxoSmithKline Inc., Mississauga, Canada). Additional inhaler details are available in Table. Each inhaler is tested under constant flow conditions of 30 and 60 L/min. For all experiments, an AIT is used as the induction port to mimic extrathoracic filtration. The aerosol exiting the AIT is interpreted as the in vitro lung dose. The lung dose is further divided into TB and alveolar doses using two distinct methods. In the first method, the aerosol exiting the AIT is supplied directly to the FBA, with the TB and alveolar doses collected on their respective, representative filters. In the second method, the aerosol exiting the AIT is characterized using a cascade impactor, with the results then being used as input to an in silico regional lung deposition model. For each experiment, regardless of the method employed, the inhaler is actuated into the apparatus five times and each experiment is repeated five times. Regional deposition estimates arising from the two experimental methods are then compared.
While the fundamental component of the filter-based approach when estimating regional deposition is the carefully-designed TB filter, the housing that holds this filter includes various additional functional design elements. Most notable are the use of an annular flow field and an integrated electrostatic neutralizer, as shown in
As the aerosol flows down the annular channel, it passes through a neutralizing section where alternating current corona discharge is used to generate bipolar ions. The outer electrode is an internally serrated ring of 0.003″ (0.08 mm) thick 304 stainless steel located 33 mm downstream from the top of the annulus core. The tips of the serrations act as corona onset locations and are separated from the ‘core electrode’ by a distance of 17 mm; an electrode separation larger than the annular gap was used to keep the electrode tips out of the flow path. The outer electrode is axially centred on the core electrode, which is a smooth stainless-steel ring 15 mm long forming part of the annular core (diameter of 45 mm). A potential difference of 5 kV is applied between the two electrodes which ionizes the air near the serrations of the outer electrode: the ions then travel across the annular gap towards the core electrode. If an ion encounters an oppositely charged aerosol particle along the way, it will be attracted to the particle and reduce its static charge. The exact charge state of aerosol leaving the neutralizing region has not been measured; however, filtration measurements have shown that the use of the neutralizer improves agreement between TB filter measurements and traditional methods of estimating regional deposition [41].
After leaving the neutralizing region, the aerosol encounters the TB filter. This filter has been carefully designed to have a filtration efficiency that mimics the average deposition efficiency of the TB region of the lungs [14], including both aerosol size and inhalation flow rate effects. The face area of this filter is an essential metric that dictates the annular radii. The filter is located 42 mm downstream of the outer electrode ring and 75 mm downstream of the top of the annular core. Two layers of highly specific stainless-steel mesh comprise the filter; the first layer uses 25 μm diameter wire woven with 500 wires per inch in each direction, while the second layer uses 30 μm wire and 400 wires per inch (SV-26/25tw, SV-34/30tw; ASADA MESH Co., Osaka, Japan). The stainless-steel mesh is spot welded between 1 mm thick stainless-steel rings to hold the mesh taught. As a result, the mesh filter elements are separated by 2 mm. After traversing the TB filter, the annular section extends for an additional 30 mm, then with a 45° draft angle the annulus core and outer walls constrict. The annulus cone ends when the diameter reaches 15 mm and the outer walls return to a straight pipe flow when the diameter reaches 18 mm. Finally, this straight pipe section extends for 29 mm to the outlet of the TB filter housing. A commercially available filter with a standard 22 mm ID connection (H8-6301; KEGO Corporation, London, Canada) installs on the outlet of the TB filter housing to collect the remaining aerosol. The total internal volume, or ‘dead space’, of the TB Filter housing is 488 mL, while the dead space from the inlet of the housing to the TB filter elements is 310 mL. A rendering of the TB Filter assembly is shown in
The desired flow rate through the apparatus was generated by a vacuum pump (0523, Gast Manufacturing Inc., Benton Harbor, MI, USA), controlled using a needle valve, and monitored with a mass flow meter (4043, TSI Incorporated, Shoreview, M N USA). A schematic representation of this experimental configuration is shown in
In the more traditional method of estimating regional deposition, the inhaler is actuated into the AIT, which is upstream of a Next Generation Impactor (NGI) including a pre-separator stage (MSP Corporation, Shoreview, M N, USA). A vacuum pump (Model KSV, General Signal, Canton, Massachusetts) generates the desired flow through the apparatus, which is set manually with a ball valve and monitored with a mass flow meter (4040, TSI Incorporated). A schematic representation of this experimental configuration is shown in
Aerodynamic particle size distributions determined using the NGI are used as input to the in silico regional deposition model described by Javaheri et al. [25]. Since the AIT is already used to remove the extrathoracic fraction, extrathoracic deposition is not modeled. The model divides the breath into 1000 bolus fractions and propagates each one sequentially through a ID symmetrical branching lung model with dimensions given by [30]. The aerosol classified by the NGI is evenly distributed between all bolus fractions and thus not all particles will reach the final generation of the model lung. Deposition equations for sedimentation [27, 28], diffusion [29], and impaction are used to estimate deposition in each generation of the lung under inhalation, breath hold, and exhalation. In this study the inhaled volume was 2.7 L, the inhalation time is set to 5.4 s or 2.7 s corresponding to 30 and 60 L/min inhalation flow rate, a breath hold of 10.0 s is used and exhalation time is 7.0 s. The model permits aerosol that evades deposition in the lung to be exhaled, which is important in predicting lung deposition for some nebulized aerosols inhaled under tidal breathing [9]. However, for single breath inhalers it is expected that most aerosol will deposit due to gravitational settling during breath holding [42]. Indeed, the exhaled dose calculated for the conditions used here is less than 10% of the lung dose, so, to draw a closer comparison to the filter-based apparatus, only the TB estimate is used from this model, while the alveolar dose is considered to be the lung dose less the TB estimate.
Ambient conditions are recorded on each day of experimentation: ambient air temperature and relative humidity are recorded from a digital thermometer/hygrometer (TH-1, Amprobe, Everett, W A, USA), and atmospheric pressure is recorded from the mass flow meters while they are open to the atmosphere under zero flow.
Collection surfaces of the AIT, NGI, and pre-separator are coated with silicone release spray (Molycote 316; Dow Corning, Midland, MI, USA) to reduce particle bounce and re-entrainment. To allow the propellant and solvents in the spray to evaporate, 15 minutes is allowed to elapse after spraying before the instruments are assembled. To ensure full collection of any particles impacting at the flow disruptor plate, ˜2 mL of the assay solvent is added to the collection dish: a disc of stainless-steel mesh was fitted into the bottom of the collection dish to promote wetting of the surfaces while using less solvent to reduce the chance of solvent splashing over the collection dish walls (which was observed to occur occasionally during preliminary experiments). No surface preparation is required for the TB Filter elements or the commercial final filter.
A 3D-printed adapter, unique to each inhaler tested, is affixed to the AIT inlet. These adapters closely match the mouthpiece geometry, ensuring a snug, airtight fit and consistent positioning of the inhaler relative to the AIT inlet. The adapters used in this study hold the inhaler in the ‘transverse’ orientation to the AIT, i.e., the axis of the mouthpiece and the aerosol plume is horizontal despite the axis of the oral cavity of the AIT being angled upwards at 29°, as shown in
With the collection surfaces adequately prepared and the appropriate adapter affixed to the AIT inlet, the equipment is assembled into the configuration being used (NGI or FBA). The vacuum pump is turned on and throttled to the appropriate volumetric flow rate; at this point the apparatus is ready for the inhaler to be actuated into it. Inhalers are primed following the leaflet directions before use for the first time: Ventolin HFA and Asmanex HFA are shaken well before every actuation, whereas QVAR is not, as per leaflet instructions. The vacuum pump remains on for the duration of the experiment: for runs with the TB Filter apparatus, the neutralizer is switched on just prior to actuating the inhaler and switched off approximately two seconds afterwards. The inhaler is removed from the AIT inlet and approximately ten seconds is allowed to elapse before the next actuation is performed.
After five actuations of the inhaler, the vacuum pump is switched off and API is recovered from collection surfaces by rinsing with HPLC grade methanol (Asmanex HFA and QVAR) or a 1:1 v/v mixture of HPLC grade methanol and deionized micron filtered water (Ventolin HFA). The AIT and alveolar filters are extracted with 25 mL of solvent divided into three washings but collected into a single 25 mL sample. The NGI and pre-separator impactor plates and the flow disruptor plate are washed with 5 mL of solvent, while the TB Filter is washed with 10 mL. When testing with QVAR, only 6 mL is used to extract API from the TB Filter to increase the sample concentration since TB deposition with QVAR is low. Ultraviolet light absorbance is used to quantify the mass of API in each sample, absorbance is measured using a spectrophotometer (Cary 8454, Agilent, Santa Clara, C A, USA) and concentration was quantified relative to absorbance measured in a series of dilutions of a prepared stock solution. Absorbance is measured at 226, 240, and 248 nm for salbutamol sulfate (Ventolin HFA), beclomethasone dipropionate (QVAR), and mometasone furoate (Asmanex HFA), respectively. Blank samples are made by preparing and washing the collection surfaces following the normal experimental procedure but without actuating the inhaler into the apparatus: absorbance is measured relative to these blank samples to eliminate interference of external compounds potentially absorbing near the API peak. Five replicate measurements are performed for each combination of inhaler, flow rate, and experimental configuration. Statistical comparisons are performed in Microsoft Excel 2019 (Microsoft, Redmond, W A, USA) using two-tailed, unpaired students t-tests.
Experiments are performed at ambient conditions (temperature, pressure, relative humidity) ranging from 21.0 to 24.7° C., 91.6 to 94.2 kPa, and 8.0 to 29.3% R H. Total recovered dose is independent of the experimental configuration (p>0.05) and is 110.5±4.1% of the label claim for Ventolin HFA, 96.7±4.3% of the label claim for Asmanex HFA, and 65.5±5.5% of the label claim for QVAR. For Asmanex HFA and QVAR, no API is detected on either the pre-separator or the flow disruptor plate. For Ventolin HFA, API detected on the pre-separator is independent of flowrate (p>0.05) and is only 3.3±1.9 μg (less than 1% of recovered API). API is detected on the flow disruptor plate for only two runs with Ventolin at 60 L/min (13.3 and 2.8 μg) when testing with the TB Filter apparatus. Since API on the pre-separator and flow disruptor plate is generally zero or negligible, it is added to the mass recovered in the AIT, consistent with previous study [41], referred to herein as aggregate upper respiratory tract (URT) deposition.
The aerodynamic particle size distribution data from the cascade impactor runs is shown in
The FBA is developed with the intention of alleviating the burden associated with compendial testing methods when estimating regional deposition. The resource burden is vastly reduced in the laboratory by reducing the number of collection surfaces to assay. There are only four collections surfaces to assay when using the FBA (AIT, flow disruptor plate, TB Filter, alveolar filter) but 10 collections surfaces when using the NGI (AIT, pre-separator, 8 NGI stages). The resource burden associated with interpreting the results using the FBA is nearly completely eliminated because there is no post processing required to arrive at regional deposition estimates since they are measured directly. The lung dose is well quantified by adopting the AIT as the induction port: if the TB Filter removes the appropriate dose from the flow, then both the TB and alveolar fractions will be accurately estimated.
The comparison of regional deposition estimates given by the two methods in Error! Reference source not found. demonstrates that the filter-based method provides reasonable estimates that account for differences due both to particle size and to inhalation flow rate. The particle size dependence is not explicitly shown in Error! Reference source not found.; however, the inhalers have been grouped from smallest to largest MMAD along the x-axis to demonstrate the effect of particle size. From this manner of grouping, it appears that the TB Filter somewhat over predicts TB deposition relative to the regional deposition model for coarser aerosols. This result is not unexpected as the filtration curves for the TB Filter given by were given only for particle sizes up to 3.3 μm, and the filter curves appear to cross above the deposition model curve around this point. For Asmanex (where the difference between estimates is greatest), there is significant mass reaching the in vitro lungs in particles larger than 3.3 μm, which will be collected in greater proportions than the in silico regional deposition model predicts. This difference is considered to be acceptable for the reasons that follow. First, the in silico regional deposition model is used here only as a guide, since the in silico result will depend on which regional deposition model (of which there are many) is used, and in vivo regional deposition will vary considerably from person to person. Furthermore, statistically significant differences in regional deposition estimates here are largely the result of the high repeatability of each method (with commensurately small standard deviations), and it is unlikely these differences would translate to clinically significant differences.
Exploratory experiments are conducted to investigate the effect of electrostatic charge on the filtration of the TB Filter and the necessity of using the neutralizer. While this has not explicitly measured the charge state of the pMDI aerosols used here before and after neutralization by our custom ion generator, limited measurements have been conducted with and without powering the neutralizer to investigate its effectiveness. These measurements are performed for the QVAR and Ventolin HFA inhalers, both of which are also studied by Leung et al. who measured the bipolar charge distribution of aerosols emitted by the pMDIs using a Bipolar Charge Analyzer (BOLAR).
We have previously discussed electrostatic effects in detail and do not expect the charge levels associated with pharmaceutical aerosols to significantly affect in vivo deposition in the TB region. However, electrostatic charge on the aerosol can significantly affect the filtration efficiency of the stainless-steel TB Filters, where image charges will readily be induced due to their conductive nature. The pore sizes of the TB Filters are 25 and 30 μm, whereas the diameter of the most distal generation of the TB region of the lung is expected to have a diameter around 20x larger [12]. Electrostatic attractive forces are inversely proportional to the squared separation distance, so the electrostatic attractive force in the TB Filter could be 400x higher than in the most distal generation of the TB region, and even higher in more proximal generations.
In the present experiments, when the neutralizer is not powered, deposition on the TB filter increases by 82% and 30% relative to experiments where the neutralizer is active for QVAR and Ventolin HFA, respectively. These results are aligned with the charge distributions measured by Leung et al. [44], who found that QVAR and Ventolin produced aerosols with very different charge characteristics. Ventolin HFA produced aerosol with relatively low charge-to-mass ratios (˜25 pC/μg) and no API was detected in the smallest size bin of the BOLAR (dp<0.95 μm), whereas the smallest bin collected the most mass when testing the QVAR, and the charge-to-mass ratio magnitude associated with these particles was >200 pC/μg. The highly charged, small particles associated with the QVAR are likely the most susceptible to electrostatic effects due to their high physical mobility. Further, more extremely charged particles could drive the deposition enhancement while mildly charged particles are less affected. The increase in deposition on the TB Filter when the neutralizer is not used is in relative agreement with the magnitude of previously measured charge profiles for these inhalers.
The two experimental configurations provide similar results for upper respiratory deposition, as the AIT is used as an induction port in both experiments. For this study, the total recovered dose from each inhaler is used to compare the recovery efficiency of the filter-based method to the established, cascade impactor method. No significant difference in recovered dose is seen between the two methods, indicating the particle losses within the NGI and the TB Filter housing are similar. While this is the primary comparison, it is also expected that total recovered dose lies between 75 and 125% of the label claim as required by [35]. The recovered dose found for QVAR is not within this range, but this is largely explained by the fact that the label claim is the ex-valve dose. The product monograph states that each actuation delivers 80 μg from the actuator. Using 80 μg as the ‘label claim’ changes the recovered dose here to 81.9±6.9%. Furthermore, QVAR is the only inhaler studied where total recovery depends on the flow rate. When testing at 30 L/min, total recovery is 76.7±3.3% of the ex-actuator claim, whereas total recovery was 87.0±5.7% of the ex-actuator claim when testing at 60 L/min (p=0.00018). While the dose remaining in the actuator is not quantified in this study, it is surmised that there is increased deposition in the actuator at lower flow rates, possibly due to better entrainment of the aerosol plume at higher flow rates. In terms of mass of API, 65.5±5.5 μg per actuation (61.4±2.7 μg and 69.6±4.5 μg at 30 and 60 L/min, respectively) are recovered, a result that compares reasonably to those reported by who recovered 74.3±13.9 μg per actuation for QVAR. Finally, the recovered API is measured when actuating the inhaler directly into a filter through which 60 L/min is being generated by a downstream vacuum pump (n=3 replicate measurements). In these confirmatory tests 71.1±4.3 μg of API per actuation (88.9±5.3% of ex-actuator claim) are recovered, indicating that the mass of API recovered in the regional deposition measurements is typical for this particular inhaler.
The differences in regional deposition due to flow rate seen for each inhaler require considering more than just the particle dynamics of an inhaled aerosol. If particle dynamics alone dictated the regional deposition estimate, one might expect the TB fraction to increase with flow rate (due to increased deposition due to impaction) and thus cause the alveolar fraction to decrease. However, for QVAR the fraction of delivered dose reaching the alveolar region increases with flow rate while the TB dose remains relatively unchanged. The TB Filtration curve is relatively flat for the small particle sizes associated with the QVAR aerosol plume, a result that agrees with the small changes to TB Filtration. The increase in alveolar dose is due to reduced extrathoracic deposition, which in part results from improved entrainment of the aerosol plume. Most important is that this trend is captured consistently by both experimental methods. Liu et al. measured the droplet velocities and impaction forces of the aerosol plumes emitted from various pMDIs, including QVAR and Ventolin HFA. They showed that droplet velocities and plume impaction forces are positively correlated and significantly lower for the QVAR device than for the Ventolin HFA. While the droplet mass at impaction could not be measured, these measurements provide a good estimate of plume momentum, as is used by Ruzycki et al. to describe the dependence of extrathoracic deposition on inhaler orientation observed for the Ventolin Evohaler (equivalent to Ventolin HFA) but not for the QVAR device. The moderate difference in inhalation flow rate used in this study does not affect extrathoracic deposition for the relatively high momentum plumes of Ventolin HFA and Asmanex HFA. Combined with the larger particle sizes associated with these inhalers, the regional lung deposition estimates follow the expected trend of increasing TB deposition with flow rate. Overall, the expected change in TB deposition due to flow rate is well captured by the filter-based apparatus as compared to the in silico deposition model used here. It should be noted that, due to the use of a constant flow rate in this study, the effect on deposition due to flow rate identified here is void of any effects relating either to the shape of inhalation profile performed by the patient or to (in) coordination of inhaling and actuating the pMDI. These effects are eliminated in the present work both because doing so enables straightforward comparison of regional deposition estimates obtained at different flow rates and because the FBA has previously been validated for use with time varying inhalation profiles [41].
A novel, filter-based apparatus downstream of an Alberta Idealized Throat extrathoracic airway is used to fractionate the in vitro lung dose into TB and alveolar regional deposition estimates for various commercially available pMDI devices. The same devices are tested in a more conventional manner using an NGI to classify the in vitro lung dose, with the results then used as input to an in silico regional deposition model. The inhalers used here provide a wide variety of aerosol sizes (MMAD ranging from 0.9 to 3.3 μm) and were each tested at 30 and 60 L/min. The regional deposition estimates from each method are in good agreement, indicating the filter-based approach can capture effects due to particle size and inhalation flow rate while requiring far less time and fewer resources to perform. This approach therefore has the potential both to expedite research and development of novel inhalation therapies exploiting targeted delivery to a distinct region of the lungs, and to expedite comparisons of equivalence between innovator products and second-entry generics.
The references cited herein are hereby incorporated by reference in their entireties.
While at least one exemplary embodiment has been presented in the foregoing detailed description, it should be appreciated that a vast number of variations exist. It should also be appreciated that the exemplary embodiment or exemplary embodiments are only examples, and are not intended to limit the scope, applicability, or configuration of the described embodiments in any way. Rather, the foregoing detailed description will provide those skilled in the art with a convenient roadmap for implementing the exemplary embodiment or exemplary embodiments. It should be understood that various changes may be made in the function and arrangement of elements without departing from the scope as set forth in the appended claims and the legal equivalents thereof.
Testing. J Aerosol Med Pulm Drug Deliv. 2016:29(2): 196-206. https://doi.org/10.1089/jamp.2015.1215
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/CA2022/050987 | 6/20/2022 | WO |
| Number | Date | Country | |
|---|---|---|---|
| 63212217 | Jun 2021 | US |
