Design and development of a novel dry powder inhalation
(DPI) aerosol drug delivery device for the treatment of
acute asthmatic episodes
Annemarie K. Alderson, Annie Saha, Stephanie T. Shaulis, Robert J. Toth
BioE 1160/1161: Senior Design, University of Pittsburgh, Department of Bioengineering
April 29, 2005
Abstract
Asthma is a chronic, constrictive disease of the intrapulmonary airways affecting a significant and
increasing number of individuals. Pharmaceutical therapy of acute asthmatic episodes with drug delivery
directly to the respiratory tract through oral inhalation has been successfully implemented in practically all
asthma sufferers. The use of either metered-dose inhalers (MDIs) and/or dry powder inhalers (DPIs) is
ubiquitous among the vast majority of individuals with the condition. However, current devices of these
types fall short of desired patient preferences, particularly in mobility and robustness. Therefore, was the
goal of this project to design and develop a dry-powder type, single-dose, disposable inhaler. The device is
completely self contained, ruggedly constructed, lightweight, small, ergonomically designed, and actively
mobile. The device is applicable to asthmatic individuals who desire a temporary alternative to traditional
devices during physical activity and/or in extreme environmental settings (running, bicycling, swimming,
skiing, various sports and outdoor activities, at the beach, etc.). The device was designed utilizing
Solidworks solid modeling software, and computationally analyzed with the COSMOSFloWorks
computational fluid dynamics (CFD) functionality within Solidworks. Prototyping of the device was
performed through Quickparts.com – a custom rapid prototyped parts supplier. The functional prototype
was developed within the time table of the project, and aerosol dispersion and flow testing was conducted
in the Aerosol Drug Delivery and Pulmonary Biomechanics Laboratory under the direction of Timothy E.
Corcoran, PhD at the University of Pittsburgh. The time table for detailed design, prototyping, and testing
was four months (January – April, 2005).
Keywords: Asthma; Bronchodilator; Aerosol; Nebulizer; Metered-does inhaler (MDI); Dry powder inhaler
(DPI); Product design specifications (PDS)
Contents
1. Introduction……………………………………………………………………………….......
2. Asthma………………………………………………………………………………………..
2.1. Characterization of the disease………………………………………………………….
2.2. Methods of treatment……………………………………………………………………
2.3. Principles of drug delivery to the respiratory tract……………………………………...
3. Oral inhalation aerosol technology …………………………………………………………..
3.1. Nebulizers……………………………………………………………………………….
3.2. Meter-dose inhalers (MDIs)……………………………………………………………..
3.3. Dry powder inhalers (DPIs)……………………………………………………………..
4. Design considerations………………………………………………………………………...
4.1. MDIs vs. DPIs…………………………………………………………………………...
4.2. Aerosol generation mechanism………………………………………………………….
4.3. Product design specifications……………………………………………………………
5. Design methods…………………………………….…………………………………………
5.1. Initial design decisions………..………………………………………………………….
1
02
03
03
04
06
07
08
09
12
13
13
14
14
17
17
5.2. Solid model design methodology…………………………………………………….…
6. Design results…………………………………………………………………..……………..
7. Design analyses, modification, and validation…….. ………………………………………..
7.1. Model-8 analyses………………………………………………………………………..
7.2. Prototype modifications…………………………………………………………………
7.3. Final analyses and validation……………………………………………………………
8. Conclusions……………………………………………………………………………………
9. Acknowledgements……………………………………………………………………………
References………………………………………………………………………………………..
18
24
26
26
29
32
34
35
35
1. Introduction
One of the most chronic conditions affecting individuals in the United States is
asthma. Simply defined, asthma is an immuno-mediated condition characterized by
increased resistance to airflow in intrapulmonary airways in response to various nonspecific chemical and physical stimuli [1]. The condition manifests itself in common
symptoms including breathlessness, chest tightness, nighttime and/or early morning
coughing episodes, and episodes of “wheezing” – exaggerated and forced breathing [1,2].
Asthma is a rapidly escalating pulmonary disease.
The prevalence of asthma has been increasing since the early 1980s for all age,
sex, and racial groups [3]. The overall age-adjusted prevalence of asthma rose from 30.7
per 1,000 population in 1980 to a 2-year average of 53.8 per 1,000 in 1993-94 [3]. This
represents an increase of 75 percent [3]. The prevalence among children ages 5 to 14
increased 74 percent, from 42.8 per 1,000 in 1980 to an average of 74.4 per 1,000 in
1993-94 [3]. Among children up to four years of age, asthma prevalence increased 160
percent, from 22.2 per 1,000, the lowest prevalence among any age group, to a 2-year
average of 57.8 per 1,000 in 1993-94, the second highest prevalence behind children 5 to
14 [3]. Thus, asthma is an obviously increasingly common condition.
As of 2001, 20.3 million Americans have reported suffering form asthma [4]. In
terms of the effects of the condition on the healthcare industry, in 2000 alone there were
10.4 million asthma-related visits to outpatient hospital clinics and private physician
offices, 1.8 million emergency room visits for asthma related problems, 465,000 inpatient
hospitalizations, and 4,487 deaths attributable to asthma-related complications [4]. This
translated to an estimated $14 billion in related healthcare costs that year [5].
As the above statistics indicate, the market for asthma-related medical devices is
large and is increasing along with the increasing trends in the incidence of the disease. In
addition, the current options in treatment technology do not completely satisfy the needs
and preferences of individuals afflicted with the disease. This is especially true in the
area of prescription drug delivery. In particular, while the pharmaceutical agents and the
devices utilized to deliver those agents to the respiratory tract – the primary site of
localized asthma treatment – have found widespread effectiveness in treating the
condition through symptom mitigation; the delivery devices, inhalers especially, possess
a number of significant limitations that oftentimes prove a hindrance to asthmatic
individuals. These limitations include limited environmental exposure; reduced or
precluded functionality under non-ideal operating conditions; fragile construction;
delicate operating mechanisms; inconvenient and inefficient shape, size, and weight; and
2
significant impairment to active portability. Therefore, it was the goal of this project to
develop, design, prototype, and test a novel aerosol drug delivery device that addresses
the limitation delineated above.
This document presents the preliminary research conducted in preparation for the
design and development of the proposed device as well as the methods and results of the
design work itself. This includes a brief review of asthma, characterization and treatment
of the disease, a review of the current oral inhalation technology, the technical
considerations implemented in the device design, a detailed project plan outlining the
design methods, and the analyses and validation of the developed device.
2. Asthma
2.1. Characterization of the disease
A diagnosis of asthma is generally considered appropriate in patients in whom
episodes of wheezy breathlessness, with intervals of relative or complete freedom from
symptoms, can be shown to be associated with variations in resistance to flow in
intrapulmonary airways [1]. In such patients, abnormal increases in expiratory airflow
resistance can usually be demonstrated in response to various non-specific chemical and
physical stimuli [1]. These stimuli have been shown to include triggers as varied as
secondhand smoke, dust and dust mites, environmental air pollution, cockroach allergen,
pet dander and fur, mold and mildew, high air humidity, freezing temperatures,
thunderstorm-generated ozone, food and/or drug additives, emotional states, and
strenuous physical activity [6]. Moreover, asthma may develop in individuals who suffer
from other chronic bronchopulmonary diseases, such as bronchiectasis or emphysema, in
which the specific pathology of the disease induces asthmatic symptoms in the diseased
respiratory tract [1,2].
Antigen-antibody reactions of several types, due to inhaled triggering stimuli,
have been shown to be concerned in the pathogenesis of asthma [1]. The most frequent
and probably the best understood of these reactions is that observed in a group of patients
who have an inherited genetic susceptibility to develop hypersensitivity to a range of
potentially antigenic substances as a result of the minor exposures to the small amounts
inevitability present from time to time in respired air [1]. The resulting immunologic
reactions are IgE immunoglobin mediated [1,2]. Other types of antigen-antibody reaction
have been shown to be involved in some cases of asthma. For instance, heavy exposure
to inhalation of certain organic and inorganic compounds can cause sensitization, not
dependent upon any genetic susceptibility in the person exposed, and accompanied by the
development of precipitating antibody; and subsequent exposure may then give rise to the
common asthma symptoms [1,6]. These symptoms usually include any number of the
following; breathlessness, chest tightness, coughing episodes, and episodes of wheezy
dyspnoea, otherwise known simply as “wheezing” [1]. All of the above symptoms are
manifestations of the increased resistance to airflow due to constriction of the respiratory
tract, particularly the bronchioles [1,6].
Upon IgE immunoglobin reaction in the respiratory tract, a number of
inflammatory mediators play an interactive role in the constriction of the bronchioles [7].
3
Acute asthmatic episodes are now thought to be connected to IgE reaction through the
actions of mast cells, which are pervasive in bronchial tissue [7]. Mast cells are known to
possess an excess of 10,000 high affinity receptors for activated IgE [7]. Once mast cells
bind with IgE, they have been shown to secrete a number of inflammatory chemicals
including histamine, various leukotrienes, and various prostaglandins [7]. These
substances have been linked to smooth muscle contraction in the bronchioles, thereby
inducing bronchoconstriction and thus asthmatic episodes [7].
The other major triggering mechanism for asthma is exercise and hyperventilation
inducement [8]. There are two major schools of thought on the mediation of exercise and
hyperventilation induced asthma, both of which provide explanations for triggering
mechanisms in addition to the possibility that the increased breathing rate associated with
exercise and hyperventilation simply might bring in more external chemical stimuli [8,9].
It has been shown that changes in the physical environment inside the respiratory tract
due to rapid breathing, particularly changes in air temperature and humidity, can incite
bronchoconstriction through changed osmolarity within the bronchial tissue [8]. Changes
in osmolarity have also been linked to mast cell recruitment and activation in the
respiratory tract [8]. The second school of thought involves direct neurological
mediation of bronchoconstriction [9]. The respiratory tract is innervated throughout its
length, but especially the bronchial regions [9]. The nerves are part of the autonomic
nervous system and retain partial control of bronchial tone through smooth muscle cell
action [9]. In terms of sole nervous effects contributing to asthma, several types of
autonomic defects have been proposed including enhanced cholinergic, -adrenergic, and
non-adrenergic non-cholinergic (NANC) excitatory mechanisms, or reduced -adrenergic
and NANC inhibitory mechanisms [9].
Despite the varied mechanisms implicated in the development and perpetuation of
asthma, they have, for the most part, been reconciled into a contributory theory, where a
number of different triggering stimuli coupled with immuno-inflammatory, neural, and
physical processes all play a role in the condition [7,9]. The complex interplay of
triggering stimuli and physiological response in asthma results in complicated treatment
methods for the disease.
2.2. Methods of treatment
Despite the chronic nature of asthma, it is most commonly treated in an acute
manner. This is primarily due to the lack of development of any successful disease
mitigation therapies [10]. Therefore, all pharmacological treatments for asthma are
disease or symptomatic control measures [10]. The lack of an effective treatment in
eliminating the pathology does not preclude total recovery from or elimination of the
disease in afflicted individuals. Complete recoveries with apparent elimination of all
symptoms have been observed in some asthmatics – predominately in children [3].
However, this occurrence cannot be correlated to pharmacological therapy and is most
likely a case of so-called “growing out of the disease” [3]. The control medication for
asthma is often classified into four categories; immunotherapy or allergy desensitization
shots, anti-IgE monoclonal antibody therapy, long-term control medications, and quickrelief medications [10-12].
4
The first two classifications involve mitigation of the immuno-inflammatory
reactions implicated in asthma [12]. Immunotherapy involves a series of injections of
asthma triggering allergens to induce desensitization [12]. Anti-IgE injections, such as
omalizumab (Xolair), block the action of IgE immunoglobin and thereby eliminate the
initiation of the inflammatory reactions implicated in asthma [12]. These therapies are
less common in the treatment of asthma than are the aerosol-type medications.
Long-term control medications are typically taken on a daily or twice-daily basis
to achieve and maintain control of persistent asthma [10]. They include corticosteroids,
long-acting 2-agonists, leukotiene modifiers, sodium cromoglycate, and theophylline
[10,13]. Table-1 below presents the most common drugs in each category.
Drug name
Table-1: Common drug treatments for asthma [10-13].
Brand name
Drug type
Long-term control medications
fluticasone
Flovent
inhaled corticosteroid
budesonide
Pulmicort
inhaled corticosteroid
triamicinolone
Azmacort
inhaled corticosteroid
flunisolide
Aerobid
inhaled corticosteroid
beclomethasone
Ovar
inhaled corticosteroid
salmeterol
Serevent
long-acting 2-agonists
formoterol
Foradil
long-acting 2-agonists
montelukast
Singulair
leukotiene modifier
zafirlukast
Accolate
leukotiene modifier
sodium cromoglycate
Intal
mechanism not known
nedocromil
Tilade
mechanism not known
theophylline
Uniphyl
mechanism not known
albuterol
Proventil/Ventolin
short-acting 2-agonists
pirbuterol
ipratropium
Maxair
Atrovent
short-acting 2-agonists
anticholinergic
Quick-relief medications
Corticosteroids act in an anti-inflammatory manner by inhibiting local recruitment of
mast cells through suppression of the formation of cytokines – the chemical mediators
responsible for the progression of the asthmatic reaction from IgE-antigen complex to
smooth muscle contraction [10]. Long-acting 2-agonists induce bronchodilation by
binding to -adrenergic receptors in bronchial tissue and inducing smooth muscle cell
relaxation [10]. Leukotriene modifiers function by blocking leukotriene production or
interfering with receptors in the airways. Inhibited leukotriene action reduces
inflammation of the airways and thereby lessens the symptoms of asthma [11]. Sodium
cromoglycate, nedocromil, and theophylline are three drugs whose mechanisms of action
are not well characterized but are known to effectively reduce asthma symptoms [11,13].
Sodium cromoglycate is especially effective. It is thought to prevent antigen-induced
5
release of mediators from mast cells through an unknown mechanism, however, it has
also been shown to possess bronchodilator properties through action on smooth muscles
cells [13]. Effective control of asthma in 60-70% of patients is observed with sodium
cromoglycate [13].
Quick-relief medications, otherwise termed rescue medications, rapidly stop the
symptoms of an abrupt asthma attack [11]. A number of long-term medications also
possess short-term effects, and thus, are utilized in a dual manner [11]. However, there
are medications that act only in a short-term manner; the most common of these are listed
in Table-1 [11]. These drugs act as bronchodilators that act rapidly to relax smooth
muscle cells and dilate the airways [10,11].
While most of these drugs are effective asthma treatments when given
systemically, it has been conclusively shown that their effectiveness is vastly increased
when applied locally to the respiratory tract [11]. Therefore, the majority of asthma
medications are delivered directly to the respiratory tract [11]. However, there a number
of aspects of inhalation drug delivery that complicates this route of administration.
2.3. Principles of drug delivery to the respiratory tract
The respiratory tract is a complex system of branching tubes of progressively
decreasing size [14]. A system of drug administration has to deliver the drugs into and
through these tubes in order for it to reach its site of action which, for drugs to treat
asthma, is regarded as being in the conducting airways [14]. The progressive reduction in
size through an ever-increasing number of airways presents a severe challenge to the
drug, since the drug particles are constantly having to change direction and, in moving
through air of progressively decreasing velocity, have an increasing tendency to deposit
[14]. Premature deposition prior to the bronchial region, where asthma symptoms
actively manifest, significantly decreases the effectiveness of the treatment [14].
Therefore, a drug delivery system to the respiratory tract must account for all major
factors affecting deposition, which for the purposes of design are listed in Table-2.
Table-2: Major factors affecting drug deposition in the lungs (adapted from[14]).
Particle properties
Aerosol properties
Respiratory tract properties
Breathing patterns
Diameter
Density
Shape
Charge
Chemical composition:
Solubility
Hygroscopicity
Concentration
Particle size range
Bolus or continuous cloud
Velocity of spray
Evaporation of propellants
Geometry (variability)
Presence of disease
Humidity
Residence time
(breath-holding)
Volumetric flow rate
(breathing rate, tidal volume)
Mouth or nasal breathing
6
The predominate factors that relate to delivery device design are the aerosol properties.
While the particle properties, respiratory tract properties, and breathing patterns have an
impact on device design, they are oftentimes beyond the control of the specific device
design, and as such, do not have as significant an effect as the aerosol properties. It is the
generation of an effective aerosol that is the primary design goal of the delivery device.
For orally inhaled drugs the rates of availability are usually very rapid and can be
expressed in terms of the extent (i.e. the percentage of the delivered dose) and the
distribution (i.e. where the drug is deposited) [14]. To be an effective delivery device, a
design should maximize the extent and localize the distribution. This is accomplished by
designing for efficient deposition within the proper portion of the respiratory tract.
There are three basic mechanisms by which particles can deposit in the respiratory
tract upon inhalation; inertial impaction, sedimentation, and Brownian motion [14].
Impaction occurs when a particle has sufficient inertia such that it is unable to travel with
the air stream when it changes direction as an airway branches [14]. It then impacts on
the airway surface, often at the bifurcation [14]. For common drugs utilized to treat
asthma, impaction is important for particles above 5m in size [14]. Particles between
0.6 and 5 m are able to move with the air stream but in the lower air velocities existing
in the conducting bronchial airways their mass causes them to sediment, or settle out of
the air stream, and deposit [14]. Deposition of particles less then 0.6 m occurs by
Brownian motion where the individual submicron-sized particles move at velocities and
in directions within the bulk air stream [14]. Particles in this size range to do deposit at
any particular location within the respiratory tract [14]. As these mechanisms indicate,
the 0.6-5 m range is the target for maximally efficient deposition in the bronchioles
[14].
In order to deliver particles of the proper size into the respiratory tract, the
delivery device must fluidize or aerosolize the drug such that agglomeration is eliminated
[15]. There are a number of mechanisms by which this is accomplished depending upon
the phase of the drug, i.e. liquid or solid based [15]. The predominate mechanisms
classify the three major types of oral inhalation aerosol drug delivery devices utilized
currently.
3. Oral inhalation aerosol technology
Oral inhalation drug delivery requires the delivery device to aerosolize the drug
compound. There are three major methods to aerosolize drug compounds; liquid drugs
can be volatilized with compressed air or oxygen mixtures that are subsequently breathed
in, liquid solutions or solid powders can be pneumatically fluidized into a dispersed
aerosol stream that is inhaled, or solid powders can be dispersed into a stream of
passively inhaled air [16,17]. These three mechanisms are the basis for operation of
nebulizers, metered-dose inhalers (MDIs) and dry-powder inhalers (DPIs) respectively
[16,17]. The specific principles and important considerations governing the operation of
these three technologies are discussed in turn below. It is important to note that the
discussions treat the three device types in a general manner and are based upon common
principles. They are not based upon any particular name-brand device.
7
3.1. Nebulizers
Nebulizers are of two primary types; air-driven or jet nebulizers and ultrasonic
nebulizers [17]. Jet nebulizers operate on the principle that by passing air at high speed
over the end of a capillary tube, liquid may be drawn up the tube from a reservoir in
which it is immersed – an example of the Venturi or Bernoulli Effect [17,18]. When the
liquid reaches the end of the capillary tube, it is drawn into the airstream and forms
droplets that disperse to become an aerosol (Figures-1,2) [17,18].
Baffle
…. .. .. .. . .. . . . . . . .. .. .
……… ………... . .
…. .. . .
….Aerosol . . .
Air
Inlet
Air
Outlet
Drug Solution
Air Inlet for
aerosol generation
Capillary Tube
Figure-1: Schematic of jet nebulizer (adapted from [17]).
An ultrasonic nebulizer uses a piezoelectric transducer to induce waves in a reservoir of
drug solution [17]. Interference of these waves at the reservoir surface leads to the
production of droplets in the atmosphere above the reservoir [17]. An airstream is passed
through this atmosphere to transport the droplets as an aerosol [17].
Figure-2: Standard jet-type nebulizer (www.adam.com).
8
Jet nebulizers are by far the most commonly used type [18]. They utilize
compressed gas from an air or oxygen mixture cylinder, hospital air-line, or electric
compressor to provide both the bulk gas flow and the gas to the aerosol generation orifice
[18]. A baffle is a standard component of a jet nebulizer [18]. It functions to provide a
surface for aerosolized droplets to impact and subsequently coalescence, thus draining
back into the reservoir [18]. This is important to ensure proper droplet size for effective
bronchial deposition [18]. Aerosolized droplets are accelerated to a velocity sufficient
for more than 99% of the droplet mass to impact the baffles or on the nebulizer wall [18].
Only 1% of the aerosol mass leaves the nebulizer and is inhaled directly [18]. Due to this
method of operation, a typical treatment takes approximately 5 to 30 minutes to dispense
depending on how much medication is to be administered [18]. Thus, drug delivery with
nebulizers is very time intensive compared to the other drug delivery options. However,
nebulizers are the most efficient inhalation delivery devices due to the manner of aerosol
generation [18]. Specifically, inhaled liquid droplets travel into the respiratory tract in
significantly larger fractions as compared to solid aerosols, which tend to succumb to a
large ingestion effect where significant amounts of the aerosol are swallowed [17].
Nebulizers possess a number of additional disadvantages that tend to favor the use
of alternative delivery devices outside a clinical setting. These include cost, size, and
complications. Nebulizer systems can cost upward of $250 compared to MDIs and DPIs
which cost less than 10% of that cost [18]. Furthermore, nebulizer systems are relatively
large, especially the gas tank and/or compressor, compared to MDIs and DPIs which can
usually fit in the palm of the hand [18]. Finally, the use of nebulizers has been implicated
in the development of respiratory tract and lung infections [18]. Therefore, nebulizer
equipment must be cleaned and sterilized on a regular basis and the air filtered [18].
These disadvantages lead the majority of asthmatic individuals to use MDIs and DPIs as
their respiratory drug delivery devices of choice.
3.2. Metered-dose inhalers (MDIs)
The pressurized metered-dose inhaler (MDI) was conceived of in 1955 by Irvine
Porush and George Maison in response to observed difficulty in inhaling aerosolized drug
treatments from squeeze-bulb glass nebulizers [19]. They were awarded a patent on the
first MDI in 1959, and their invention saw only minor modifications and cosmetic
updates until the late 1980s when advances in materials and metering technology allowed
for miniaturization of the device [19]. MDIs incorporate a propellant, under pressure, to
generate a metered dose of an aerosol through an atomization nozzle [20,21]. MDIs are
the most widely used inhalation drug delivery device, with an estimated 800 million units
produced in 2000 [20,21]. MDIs consist of several components: the active substance
formulated with propellant, surfactants/solvents termed excipients, and the drug; a
container; a metering valve crimped onto the container; an actuator that connects the
metering valve to an atomization nozzle; and a mouth piece [20,21]. Additionally,
holding chambers or spacers may also form part of the delivery system by connection to
the actuator mouthpiece [20,21]. The current incarnation of an MDI is diagrammed in
Figures-3,4 [20,21]. A metered volume (typically between 20 and 100 L) of the
drug/excipient/propellant blend is expelled from the canister via the valve and quickly
passes through the actuator orifice where atomization occurs [20,21].
9
Figure-3: Basic components of an MDI system [20].
Figure-4: Cutaway schematic of and MDI (http://www.3M.com)
Current research and development on MDI systems primarily concerns the
chemical propellants [20]. Chlorofluorocarbons (CFCs) have been utilized as the
chemical propellants in MDIs for almost 50 years [20]. There are three different CFCs
that have found applicability in MDIs: CFC-11, CFC-12, and CFC-114 [20]. However,
due to concern over environmental effects of CFCs the FDA in conjunction with the EPA
has set a series of standards governing the phase-out of CFC-MDIs [20]. Specifically,
once two non-ozone-depleting propellants are marketed with identical or superior
characteristics in MDIs, CFCs will be phased out of use in the devices [20]. Two
promising candidates receiving much attention are from the hydrofluorocarbon family:
HFA 134a and HFA 227ea [20]. These propellants along with the CFCs (Figure-5,
10
Table-3) account for all chemical propellants encountered in all MDI applications (Table4).
Figure-5: Chemical structures of CFC and HFA propellants [20].
Table-3: Physicochemical properties of MDI propellants [20].
Table-4: Common marketed MDIs and their chemical composition [20].
11
The design and development of MDI systems is complicated by the necessary
integration of numerous principles within one device [21]. An effective MDI must
integrate a pressure vessel, nozzle, valve assembly into one device that is intuitive to
operate and which provides a conformable interface with the mouth of a user.
Furthermore, the chemical principles governing the drug/excipient/propellant mixture and
the fluid mechanical principles governing the fluidization of the aerosol must be
accounted for. Consequently, it is understandable why a successful design has been
fundamentally unchanged for over 50 years.
3.3. Dry powder inhalers (DPIs)
A dry powder inhaler (DPI), like a metered dose inhaler, is a handheld device that
delivers a precisely measured dose of asthma medicine into the lungs. Both quick relief
medicine (inhaled bronchodilators) and long-term control medicine (inhaled
corticosteroids) can be delivered to the airways using a DPI [22]. Unlike metered dose
inhalers, where slow inhalation is needed to acquire the full benefit of the medication,
DPIs require the user to breathe in quickly and forcefully to automatically activate the
proper flow of medication [22,23]. Since there are no propellants used in DPIs, the user
must inhale with more force than when using a metered dose inhaler [23]. It is usually
recommended that to receive full benefit from a DPI, the user should hold his or her
breath for approximately ten seconds (or longer, if possible) after inhaling [23]. It is
important that the user does not breathe out through a DPI, because the moisture in the
breath can cause powder agglomeration thereby clogging the mechanism, making it less
efficient for, or precluding, future uses [22,23].
In a DPI, the asthma medication comes in a dry powder form [22,24]. A small
capsule, disk, or compartment inside the inhaler device is used to hold the medication
[22,24]. Manufacturing of DPIs for drug administration requires powders with desirable
characteristics [22]. Specifically, as noted above, upon aerosolization, the powder
particles must be within the 0.6-5 m range [22]. In order to ensure compliance with this
requirement, a number of processing methods are utilized to generate powders of proper
properties [22]. These methods include spray drying, spray freeze drying, controlled
evaporation of droplets, solvent precipitation, recrystallization, fluid energy milling, and
nano-milling [22]. In addition to the drug powder, DPI formulations often include a
number of additional additives – termed excipients as in MDI technology [22]. These
substances include lubricants or anti-adherents which minimize agglomeration upon
aerosolization, desiccants, and for some drugs carrier particles [22]. The most common
dry powder excipient is lactose [22]. Fine lactose (~5 m) can function as a lubricant and
as a carrier particle depending upon the specific formulation [22]. As a lubricant, lactose
functions to ensure particle dispersion upon inhalation by interfering with drug particledrug particle interactions [22]. As a carrier, it functions as a substrate to which the drug
compound is immobilized [22]. The lactose-drug complex is then inhaled and deposits in
the respiratory tract [22]. Because DPIs utilize powdered medications, the need to keep
them dry is crucial. As such, DPIs should not be stored in damp environments.
Moreover, one of the major drawbacks to DPIs is their incompatibility with humid
environments.
12
Many types of DPIs are currently available and each has a different operating
mechanism. Consequently, there is no one general mechanism that could be described as
was the case with MDIs. Furthermore, a description of each mechanism of even the most
common DPIs is beyond the scope of this report. However, the major types bear
mentioning. Some DPIs, including Inhalator®, Spinhaler®, and Rotahaler®, must have
the medication loaded each time they are used [22]. Others, such as Diskhaler®, have
preloaded disks with a certain number of doses [22]. Turbuhaler® and Accuhaler® are
two DPIs that have as many as 200 doses stored in the device [22]. Since all DPIs rely on
the force of the user’s inhalation in order to properly deliver the medication into the
lungs, DPIs are not recommended for children under the age of five, people with severe
asthma or those suffering a severe attack [22].
The primary advantage of using a dry powder inhaler is that it is breath-activated,
so that the user does not need to coordinate activating the inhaler (dispensing the
medication) with inhaling the medication [23]. Instead, the flow of medication is
activated by simply breathing in. Additionally, DPIs do not require propellants so they
are more environmentally-friendly than metered dose inhalers [22]. Several
disadvantages of DPIs are that they are often more expensive than the equivalent metered
dose inhaler and they may be difficult and cumbersome to load [23]. For example, the
Rotahaler® requires the user to carry a supply of medication capsules with them because
it can only hold one capsule at a time [22]. If a single capsule is not sufficient to stop the
asthma attack, the user must load another capsule in order to receive additional
medication.
The design and development of DPI systems is very flexible due to the fact that
there are countless mechanisms whereby dry powder capsules can be aerosolized and
inhaled. A design is only limited by the constraints of the particular specifications for a
device. Accordingly, it is understandable that there are numerous different DPI designs
each with their own unique characteristics.
4. Design considerations
As the above discussions indicate, there are numerous factors that must be
considered when designing an inhalation drug delivery device based on aerosolization.
However, the most pertinent consideration is by far the mechanism that will be utilized to
generate the drug aerosol. Since the proposed drug delivery device is to be by design
mobile and for use in an acute, quick-relief manner, the choice of aerosolization
mechanism is between that of an MDI or a DPI. Accordingly, what follows is a
discussion of the benefits and detriments of the respective technologies and the logic and
motivation behind the ultimate mechanism choice for the proposed device design.
4.1. MDIs vs. DPIs
In terms of effectiveness in the delivery of asthma treatment medications,
numerous clinical studies have reported the therapeutic equivalence of MDIs and DPIs
[14]. While differences in the deposition characteristics under controlled conditions have
been demonstrated, the effective dose delivery is analogous between the two device types
13
with established powder production and processing procedures for DPIs [25]. Moreover,
the goals of this project were not pharmaceutically related, but rather were simply to
develop a device that met the deficiencies of current devices in terms of ruggedness,
mobility, and versatility. Therefore, it was practical and logistic factors that weighed
most heavily on the choice of aerosol generation mechanism.
Ideally, it would have been prudent to design two devices, one MDI-based and
the other DPI-based, each with analogous product specifications so that consumers would
have the choice of which aerosol generation mechanism they prefer. However, given the
scope and resources allocated for this project, such an undertaking was not
accomplishable. Given the limited time frame, monetary resources, personnel, and
production equipment for detailed design, prototyping, and testing; the project had to be
especially limited in scope and of relatively simple execution. In order the meet these
requirements, the overall device design was as simplified as possible. Considering the
principles involved in the operation of MDIs versus DPIs; DPIs were clearly the lesscomplex, more straightforward to execute technology.
A DPI-type novel device would be more rapidly designed, prototyped, and tested
than an MDI-based design. An MDI design would require work with chemical
propellants, metering valves, pressure vessels, filling equipment, and a vast number of
additional considerations that the DPI design would preclude. The design, prototyping,
and testing of a relatively simple mechanical mechanism to activate a dry powder fitted
within a casing that meets the design specifications is well within the constraints of the
project. Therefore, a DPI-type device was chosen to be designed and developed
consistent with the design specifications.
4.2. Aerosol generation mechanism
The aerosol generation mechanisms in current DPIs are as varied as the number of
different devices on the market. Each device generates the aerosol based upon the
functionality of the device, i.e. pre-loaded multi-dose disks, re-loadable devices – both
multi-dose and single-dose, pre-loaded single-dose devices, etc. It was of crucial
importance that the specific mechanism be consistent with the intended specifications and
operation of the device. Accordingly, the utilized mechanism was consistent with
rugged, impact-resistant, and reliable operation. The general principles of the aerosol
generation mechanism were similar to that of current technology, but different enough to
preclude patent infringement.
4.3. Product design specifications
Any respiratory drug delivery device must possess certain general characteristics
and features to be a successful design. These characteristics and features are summarized
in Tables-5 and 6 respectively [19]. These characteristics and features are important to
an effective design regardless of the specific type of mechanism utilized and the unique
features of any one device. However, it is the mechanism itself and the unique features
of a design that provide a market for the device. Therefore, it is important to design an
aerosol drug delivery device with distinctive characteristics that improve upon existing
devices or are novel.
14
Table-5: Desirable characteristics of respiratory drug delivery devices of major interest [19].
15
Table-6: Features of an ideal respiratory delivery device [19].
It is the improvements and novelty of a design that are the important aspects of the
product design specifications (PDS) for the purposes of detailed design work. While a
complete discussion of the product design specifications for the proposed device is
beyond the scope of this report, it is important to note those specifications that
significantly impact the detailed design of the device.
16
As previously noted, the device was of the DPI-type. Thus, the aerosolization
mechanism was consistent with this specification. Moreover, the device was to be
lightweight, single-use-only, disposable, and extremely rugged. These specifications had
a profound effect upon the choice of materials of which the device will be constructed.
In addition, the ergonomic and size properties specified for the device affect the external
shape, which had implications for the internal mechanism. Thus, it was important to take
time to work out the design details before subsequent prototyping and testing in order to
eliminate any waste of the limited time or material/monetary resources.
5. Design methods
5.1. Initial design decisions
The actual development of any computational or physical prototype of the device
was initially precluded by a number of design options that required finalization before
subsequent model development. These specifications included decisions regarding
materials of construction, anthropometry, ergonomics, human factors considerations and
pharmaceutical options and associated dosing constraints.
In accordance with the design specifications of lightweight yet rugged
construction and weather and water resistance, it was initially decided that a polymerbased plastic was the best option. This was consistent with the construction of predicate
devices, which all are constructed from a specified plastic material. Seven plastics were
evaluated in terms of material density (a lightweight material was desired), material
strength (a material that could absorb significant impact was desired), and material
hardness (a soft plastic material was desired). The important physical properties of the
respective materials are presented in Table 7.
Table-7: Plastic materials of construction options and associated properties [27].
Plastics Comparison
Plastic Type
Density (g/cc)
Impact Strength
(J/cm)
Rockwell Hardness
[R]
ABS
1.08
6.40
115
LD-polyethylene
0.91
6.94
60
PET
1.30
1.40
110
Polypropylene
1.07
11.50
91
Polystyrene
1.00
2.94
75
PVC
1.37
13.90
80
ABS/PVC Blend
0.98
12.50
102
17
In consideration of the above materials, it was most desirable to minimize the density and
hardness values while maximizing the impact strength. Unfortunately, no one specific
material possessed all of the optimal properties. Consequently, a compromise had to be
made with the most acceptable choice being low density polyethylene (LDPE). The
choice of LDPE minimized both density and hardness. As a result, the chosen material of
construction would provide for the lightest weight and softest device possible.
Accordingly, the design specifications in terms of product weight and minimal user
impact during physical collisions were effectively met. However, LDPE did not possess
the impact strength of some of the other plastics options. Therefore, the compromise was
made regarding the material’s impact strength. Nonetheless, the impact strength of
LDPE is sufficient for its application in the device since it does not necessarily need to be
maximized as long as it is high enough to withstand the magnitude of any foreseeable
impacts during portability of the device, which it adequately achieves. In addition,
LDPE possesses a unique property that makes it the optimal material of construction
despite its lower impact strength. Specifically, LDPE is heat-sealable [27]. Two separate
components constructed of LDPE can be readily bonded simply through the application
of site-directed heat through filaments, wires, etc. The strength of the heat-sealed bond
can also be modulated through the temperature and time exposure during the sealing
process [27]. This property has exiting implications on the future design of the overall
sealing mechanism of the device, an aspect of the design not considered as a part of this
project.
In order to design a prototype for use for individuals varying in size,
anthropometric considerations were evaluated to determine the average size mouth for
both males and females. Anthropometric considerations were evaluated to determine an
appropriate size for the mouthpiece to maximize comfort and ease of use for a host
of asthmatics. It was also decided to use the current inhaler models on the market as
guidance on the assumption that the companies had done a plethora of research in
designing an optimal mouthpiece. Using various anthropometric data, the lip widths for
people of various races and ages was approximately from commissure to commissure
ranged from 29-38 mm, and the lip heights from base of columella to tubercle ranged
from 19-29 mm [28]. After researching anthropometry, the final mouthpiece size was
chosen to be cylindrical with a 20 mm diameter and 20 mm in length.
Another main component to the design considerations was the mass of the drug to
be loaded into the device. The prototype needed to have adequate space to hold and seal
the amount of medication needed to be inhaled by the user. Delving into the current
research for inhalers on the market, it was determined that about 25 mg 1 mg was the
total load dosage of both the excipient lactose and bronchodilator albuterol sulfate in a
33% w/w composition. [29].
5.2. Solid model design methodology
The design of the DPI device began with conceptual hand-illustrations of how it
would look. These included descriptions of both the internal and external geometry of
the device. These early conceptual drawings were refined in terms of incorporating the
anthropometric considerations regarding hand and mouth size and shape in relation to the
external geometry of the device and the mouthpiece design. In addition, the choice of
18
albuterol sulfate as the bronchodilator drug compound and micronized lactose as the
excipient for a total dosage of 25 mg placed constraints on the internal dimensions of the
device that were reflected in these drawing refinements. The refined drawings were then
converted to solid models in Solidworks. Solidworks is a feature-based dynamic solid
modeling software package. It allows for the creation of single solid part representations
and organized groupings of such parts in dynamic assemblies. Once transferred to solid
model format, the conceptual device models could then be readily modified.
In order to incorporate all of the design considerations into a prototype device
consistent with the project goals, an iterative design process was implemented. A
computational fluid dynamics (CFD) simulation was performed on the initial conceptual
device solid model. The CFD was performed utilizing the COSMOSFloWorks
functionality within Solidworks. The simulations were turbulent and time-independent
with air as the process fluid. Physiologic boundary conditions determined from
pulmonary mechanics were implemented, and the bulk average volumetric flowrate
through the device was chosen as the primary convergence parameter due to its
importance in obtaining aerosol dynamics within the device consistent with the effective
recruitment, dispersion, and deposition of 5 m drug particles. The results of the CFD
simulations were then used as a guide to modify and refine the solid models of the device
in order to improve the CFD predicted fluid behavior within the device. Once these
modifications were implemented, the resulting version of the solid model was used to reperform a comparable CFD simulation. This design cycle was then repeated until a solid
model was developed that adequately met the design considerations and performed as
desired through CFD simulation.
Summarizing the implementation of this process, a total of 10 solid models of
different devices were developed. Each subsequent solid model was developed based
upon the previous model with the aid of the CFD simulation results in performing the
modifications and refinements. The initial CFD results on the early solid models
indicated the presence of significant rotational flow patterns within the body of the
device.
Figure-6: Early solid model representation of device.
19
Figure-7: Initial CFD results indicating rotational flow patterns.
Initially it was thought that these rotational patterns were a consequence of the internal
geometry of the device and the orientation of the two pressure inlets. These patterns were
unacceptable because such air flow would be inefficient at recruiting drug particles from
within the device and dispersing them from the device through the mouth and depositing
them within the airways. In order the effectively achieve such dispersion and deposition,
a bulk flow of air directly through the device would be ideal. Accordingly, the first
modifications made to the device solid model were changes of the internal geometry of
the device and of the orientation and number of the pressure inlets in order eliminate the
rotational patterns and obtain bulk flow directly through the internal body of the device.
These modifications are reflected in Figures 8 and 9.
20
Figure-8: Model-3 CFD results.
Figure-9: Model-6 CFD results.
Various changes were made in the internal geometry and pressure inlet orientation from
Models 1 through 6, with the consistent result being the presence of rotational flow
patterns in the CFD simulation results. It was at Model-6 that the source of these
rotational flow patterns was determined. From Model-1 through Model-6, a mixed set of
boundary conditions consisting of an ambient (atmospheric) static pressure boundary
condition at the inlets and a specified 60 L/min volumetric flowrate as the outlet
boundary condition. This flow was chosen because it has been found to be the median
inspiratory flowrate for individuals during inhalation drug therapy [30]. It was these
mixed boundary conditions that were the source of the rotations. Specifically, by
21
specifying a pressure at the inlets and a flowrate at the outlet, the computational solver
was generating the rotational patterns in order to establish the necessary pressure drop
across the device consistent with the specified flow. In a sense, the mixed boundary
conditions were over constraining the system.
After review of CFD theory and the operational capabilities of
COSMOSFloWorks, it was determined that the boundary conditions should be changed
from the mixed pressure/flow set to a consistent set of pressure boundary conditions. The
inlet boundary condition was kept at ambient pressure and the outlet boundary condition
was set at a total pressure of 1 mmHg below ambient. This pressure was determined
from pulmonary mechanics. Specifically, intralveolar pressure is 2 mmHg below
ambient [31]. It has been shown that the resistive pressure loss from the alveoli through
the airways to the mouth is approximately 1 mmHg [31]. Therefore, during inspiration,
an average of -1 mmHg pressure is drawn at the mouth, and therefore at the mouthpiece
where the boundary condition is implemented.
The improved boundary conditions were implemented for a simple simulation to
test their effectiveness as part of Model-7.
Figure-10: Model-7 CFD results.
Once implemented, the consistent pressure boundary conditions eliminated the rotational
patterns and resulted in bulk flow through the device. Despite the source of the rotational
flow patterns being the over-constraining boundary conditions, the internal geometry of
the device still created a flow pattern that was excessively turbulent and would not
efficiently recruit the drug particles from within the body of the device, disperse them
from the device through the mouth and upper airways, and effectively deposit them
within the lower airways. Therefore, Model-8 was redesigned to simplify the internal
geometry and thereby create aerosol dynamics within the device consistent with the
efficient recruitment, dispersion, and deposition of 5 m sized drug particles. The
resulting model and associated CFD simulation results demonstrated the optimal
characteristics for the device.
22
Figure-11: Model-8 CFD results.
23
These patterns were indicative of a bulk flow through the device consistent with the
desired dynamics. The CFD simulations calculated average fluid flow velocities in the
two regions (pressure inlet, body of the device) corresponding to Reynolds numbers of
10307 and 3436 respectively. Two Reynolds numbers were calculated due to the large
change in diameter between the two regions. These numbers were indicative of a
transitional or turbulent flow regime which was consistent with the choice of a turbulent
simulation. However, the magnitudes of the numbers and the corresponding turbulence
were sufficiently low to still effectively result in drug particle recruitment and dispersion.
Furthermore, the simulation gave a bulk flow through the device consistent with those
observed in inhalation drug therapies, specifically 32.5 L/min. As a consequence of
Model-8 achieving the desired performance and complying with the specified design
considerations, it was subsequently physically prototyped.
6. Design results
Model-8 was the preliminary result of the implementation of the iterative design
methodology. It was physically prototyped through Quickparts.com, an online distributor
of custom rapid prototyped parts. Quickparts.com utilizes uploaded solid model
computational files (.STL format) and creates the part from a proprietary ABS-like
plastic formulation using the stereolithography rapid prototyping technique.
Stereolithography (SLA) is often considered the pioneer of the rapid prototyping
industry with the first commercial system introduced in 1988 by 3D Systems. The system
consists of an Ultra-Violet Laser, a vat of photo-curable liquid resin, and a controlling
system. A platform is lowered into the resin (via an elevator system), such that the
surface of the platform is a layer-thickness below the surface of the resin. The laser beam
then traces the boundaries and fills in a two-dimensional cross section of the model,
solidifying the resin wherever it touches. Once a layer is complete, the platform descends
a layer thickness, resin flows over the first layer, and the next layer is built. This process
continues until the model is complete. Once the model is complete, the platform rises out
of the vat and the excess resin is drained. The model is then removed from the platform,
washed of excess resin, and then placed in a UV oven for a final curing. The model is
then finished by smoothing the "stair-steps" [32].
Unfortunately, the method that would be required to fabricate the device out of
the end-type material – LDPE – would be injection molding. The fabrication of an
injection molded LDPE prototype would have required the production of die-cast custom
molds, which, costing upwards of $10,000 dollars under the most conservative estimates
was beyond the resources available for the project. However, to perform device
validation testing, an analogous device fabricated from any plastic-type material would
suffice and therefore, the rapid-prototyped device was applicable. This was because the
aerosol generation was primarily a function of the boundary conditions and the internal
geometry of the device itself, with the material of construction possessing a negligible
role.
The resulting physical prototype was fabricated in two separate components. The
separate component construction was established in order to simplify the repeated
24
loading and unloading as well as intermediate washes necessary during the testing and
validation of the device.
Figure-12: Model-8 and resulting physical prototype.
25
Figure-13: Model-8 and resulting physical prototype.
7. Design analyses, modifications, and validation
Pursuant with the design considerations and project goals set forth at the onset of
the project, the physical prototype was validated in terms of the effectiveness of the
recruitment and dispersion of drug particles from the device under inhalation therapy
conditions. A major shortcoming of the project as completed was the inability to
fabricate the device out of the LDPE material, and thus, the inability to physically
prototype and validate the sealing mechanism and environmental performance of the
device. However, the testing and validation of the aerosol dynamics within the device
provided a firm foundation upon which to base the success of the project.
7.1. Model-8 analyses
The physical prototype of Model-8 was initially tested in terms of the ability to
effectively recruit the drug particles from the device with a physiological flowrate and
26
disperse those particles from the device in a manner that corresponded to effective airway
deposition. Specifically, it was desired to demonstrate a significant fraction of the
particle distribution from the device representing sizes of 5 m or less. This is pursuant
with respirable particle sizes that would deposit in the bronchial region of the airways as
indicated in Figure-14.
Figure-14: Particle size distributions and the corresponding airway deposition [33].
This was performed using a Malvern Mastersizer device. The Mastersizer utilizes laser
diffraction by the particles to determine the particle size distribution in a stream of
fluidized solids. The device functions by aerosolizing the solid drug compound from the
device by pulling a negative pressure through vacuum line in a manner analogous to
inspiration during drug therapy situations. The fluidized drug particles are then drawn
through a glass chamber where the incident laser beam contacts the solid particles in the
air stream and is diffracted based upon the size of the particles. In relative terms, small
particles bend the beam by large angles whereas large particles bend the beam by small
angles. The device detector is calibrated in order to determine particle size distribution
based upon aspirated samples.
27
Figure-15: The Mastersizer laser sizing device.
The developed device was tested with the Mastersizer using 25 mg of a 33% w/w
blend of micronized atropine sulfate and lactose excipient. In practice the inhaler device
would utilize albuterol sulfate as the active drug compound, however, it was not available
in micronized form at the time of these testing activities, and therefore, the atropine was
used. It should be noted that these tests were on the particle dynamics related to the
device only, and consequently, any micronized solid powder would have functioned
adequately in place of active drug. On average, the vacuum line generates a maximum of
45 L/min of air flow through the inlet port. In order to establish a baseline set of data, the
atropine/lactose blend was sized from open air off of the end of a laboratory spatula.
This provided no additional resistance to flow and the maximal 45 L/min was available
for particle recruitment. Figure-16 presents the results of this test in terms of particle
distribution.
Volume (%) of Blend from Vial
100
90
80
70
60
50
40
30
20
10
0
34%
0.1
1
10
100
Particle Diameter (um)
Figure-16: Open air aerosolization results with atropine/lactose test blend.
28
As indicated in the figure, under optimal conditions, approximately 34% of the total drug
mixture volume was respirable. This corresponded to a large fraction of the 33% by mass
component composition of atropine. Figure-17 presents analogous results for the
physical prototype of Model-8.
Volume (%) of Blend from Inhaler
100
90
80
70
60
50
40
30
20
10
0
11
%
0.1
1
10
100
1000
Particle Diameter (um)
Figure-17: Aerosolization from Model-8 prototype.
From Figure-17 it was evident that the majority of the particles dispersed from the device
were above respirable size. In additional, through qualitative observation it was
discovered that only a small fraction of the total 25 mg dose loaded into the device was
exiting. These results were attributed to insufficient airflow through the device.
Specifically, when attached to the inlet port of the Mastersizer, only 8 L/min of airflow
was being pulled through the device, compared to 45 L/min in open air. As a result, the
prototype and the design as a whole had to be modified.
7.2. Prototype modifications
Due to the inadequate flow through the device, modifications had to be made that
would decrease the fluid resistance within the device and thereby increase flow. Since
the internal geometry of the device body was a function of anthropometry and dosage
considerations, the only option was modification of the pressure inlets. It was initially
decided to increase the diameter of the pressure inlet from 2 mm to 3mm. This
corresponded to a decrease by a factor of 9/4 in the apparent resistance through the
pressure inlet. The resulting Model-9 was evaluated qualitatively through CFD in order
to determine the general effect of the diameter increase.
29
Figure-18: Model-9 CFD results.
The CFD results indicated only a slight increase in the average flow through the device.
It was also determined that the qualitative flow pattern illustrated in Figures 11 and 18,
specifically the flow separation at the top of the internal device chamber, may be
contributing to the inadequate dynamics. This led to the decision to create a second
pressure inlet, symmetric with the first through the mid-plane of the device. These design
changes were reflected in Model-10.
30
Figure-19: Model-10 CFD results.
The CFD results on Model-10 represented a significant improvement in airflow through
the device. Specifically, the predicted bulk air flowrate through the device went up to
42.4 L/min. These promising CFD results led to the implementation of both the diameter
increase and the addition of a second pressure inlet in the prototyped device. The
modifications where made to the device and the resulting Model-10 was tested with the
Mastersizer.
31
7.3. Final analyses and validation
After the modifications were made to the prototype, it was re-tested in the
Mastersizer in an identical manner as the initial tests. Figures 20 and 21 present the
results of these tests.
Volume (%) of Blend from Redesigned Inhaler- First
Run
100
90
80
70
60
50
40
30
20
10
0
51%
0.1
1
10
100
Particle Diameter (um)
Figure-20: Particle distribution from modified prototype.
Volume (%) of Blend from Redesigned InhalerSecond Run
100
90
80
70
60
50
40
30
20
10
0
63%
0.1
1
10
Particle Diameter (um)
Figure-21: Particle distribution from modified prototype.
32
100
Two representative runs of the device through the Mastersizer indicated 51% and 63% of
the total dosage volume was below the 5 m respirable particle size. Visual observation
indicated that the total 25 mg dose was completely evacuated from the device. These
results surpass the baseline open air results in terms of delivering an inhaled dose of
predominately respirable drug. These results were interpreted as demonstrating that the
device successfully delivered a total dosage of micronized bronchodilating drug in an
acceptably respirable form. The successful results led to the fabrication of the final
physical prototype (Model-10) from quickparts.com with the validated modifications
incorporated into the fabrication itself.
Figure-22: Final physical prototype.
33
Measurement of the flowrate through the final device gave an average of 38 L/min.
Comparing this flowrate to the CFD predicted flow of 42.4 L/min, the use of CFD as a
design tool to guide the intuitive design of the device was validated.
Ultimately, it was concluded that the iterative CFD-guided design methodology
was successfully implemented with the resulting physical prototype performing as
designed. Specifically, drug particle dispersion less than 5 m was observed with
effective separation of drug and excipient indicated by the particle size distribution
obtained from the device. Flow through the device was measured at levels consistent
with physiologic inspiration levels. These metrics combined to provide a general
confidence in the device in terms of effective drug deposition within the proper regions of
the airways.
8. Conclusions
Asthma is a chronic disease that shows not signs of being eliminated from the
patient population. Due to the pervasive nature of the disease, and the unique manner in
which it is treated pharmaceutically, the market for respiratory drug delivery devices to
treat asthma is exceptionally large. However, despite the large market, some asthmatics,
especially those who retain an active lifestyle, are not completely satisfied with the
aerosol drug delivery options available in terms of robustness and versatility of the
devices. Thus, there exists a market for rugged and weather-resistant aerosol drug
delivery devices that can be used in outdoor settings and can withstand the rigors of
physical activities. Accordingly, the developed device aimed to meets these demands.
Through the use of a DPI-type mechanism, the developed device provided a
complimentary option for DPI users during sports and other outdoor activities or in any
situation where conventional DPIs are not suited to the physical environment.
The detailed design of the proposed device made use of Solidworks solidmodeling software and the various associated functionalities in order to design and
computationally model the device. The device was prototyped through Quickparts.com,
and was tested with the assistance of Dr. Timothy E. Corcoran in the Aerosol Drug
Delivery and Pulmonary Biomechanics Laboratory at the University of Pittsburgh.
The device has been shown to be capable of full evacuation and effective
dispersion of a 25 mg combined dosage of albuterol sulfate bronchodilator and lactose
excipient. The device therefore shows the potential to effectively function as an aerosol
drug delivery vehicle for the treatment of minor and acute asthmatic episodes.
At this point in the development of the device, the external and internal geometry
of the device has been designed and validated. To complete the design to a functional
level, the internal sealing mechanism for the drug compound and the coupling of that
mechanism to the external sealing mechanism must be finalized. At that point, a fully
functional prototype could be completed, contingent upon the ability to fabricate the
device from the LDPE material.
34
9. Acknowledgements
The authors would like to acknowledge Dr. Timothy E. Corcoran, the director of the
Aerosol Drug Delivery and Pulmonary Biomechanics Laboratory in the University of
Pittsburgh School of Medicine for his mentorship in the initial planning and execution of
the design. His expert knowledge of aerosol drug delivery was proven invaluable in
determining the course of the design for the device and in the validation. Finally, special
thanks to the generous gift of Dr. Hal Wrigley and Dr. Linda Baker, and the Department
of Bioengineering for providing the $500 working budget for the project.
References
[1]
[2]
[3]
[4]
[5]
[6]
[7]
[8]
[9]
[10]
[11]
[12]
[13]
[14]
[15]
[16]
[17]
T.J.H. Clark, Definition and clinical categories of asthma. In T.J.H. Clark, S. Godfrey, T.H. Lee
(Eds.), Asthma (3rd Ed., pp. 1-13); New York: Chapman & Hall.
L.A. Laitinen, A. Laitinen, Pathology and cytology of asthma. In T.J.H. Clark, S. Godfrey, T.H.
Lee (Eds.), Asthma (3rd Ed., pp. 232-253); New York: Chapman & Hall.
P.G.J. Burnery, Epidemiology, In T.J.H. Clark, S. Godfrey, T.H. Lee (Eds.), Asthma (3rd Ed., pp.
254-307); New York: Chapman & Hall.
“Asthma Prevalence, Health Care use, and Mortality, 2000-2001,” National Center for Health
Statistics, Centers for Disease Control and Prevention.
“Morbidity & Mortality: 2002 Chart Book on Cardiovascular, Lung, and Blood Diseases,”
National Heart, Lung, and Blood Institute, National Institutes of Health, May 2002.
T.H. Lee, Mediators of bronchial responsiveness. In T.J.H. Clark, S. Godfrey, T.H. Lee (Eds.),
Asthma (3rd Ed., pp. 162-181); New York: Chapman & Hall.
M.P. Carroll, C. Gratziou, S.T. Holgate, Inflammation and inflammatory mediators in asthma. In
T.J.H. Clark, S. Godfrey, T.H. Lee (Eds.), Asthma (3rd Ed., pp. 182-231); New York: Chapman &
Hall.
S. Godfrey, Exercise- and hyperventilation-induced asthma. In T.J.H. Clark, S. Godfrey, T.H. Lee
(Eds.), Asthma (3rd Ed., pp. 73-107); New York: Chapman & Hall.
P.J. Barnes, Neural mechanism in asthma. In T.J.H. Clark, S. Godfrey, T.H. Lee (Eds.), Asthma
(3rd Ed., pp. 133-161); New York: Chapman & Hall.
R.G. Taylor, Adrenoceptor agonists and steroids in respiratory disease. In D. Ganderton, T. Jones
(Eds.), Drug Delivery to the Respiratory Tract (pp. 27-36); New York: VCH Publishers.
D.C. Thompson, Pharmacology of Theraputic Aerosols. In A.J. Hickey (Ed.), Pharmaceutical
Inhalation Aerosol Technology (2nd Ed., pp. 31-64); New York: Marcel Dekker, Inc.
“Immunotherapy for asthma,” Mayo Clinic: Respiratory Diseases – Treatment and Prevention,
2004, http://www.mayoclinic.com.
R.M. Auty, Sodium cromoglycate. In D. Ganderton, T. Jones (Eds.), Drug Delivery to the
Respiratory Tract (pp. 37-46); New York: VCH Publishers.
J.M. Padfield, Principles of drug administration to the respiratory tract. In D. Ganderton, T. Jones
(Eds.), Drug Delivery to the Respiratory Tract (pp. 75-86); New York: VCH Publishers.
P.A. Crooks, A.M. Al-Ghananeem, Drug Targeting to the Lung: Chemical and Biological
Considerations. In A.J. Hickey (Ed.), Pharmaceutical Inhalation Aerosol Technology (2nd Ed., pp.
90-154); New York: Marcel Dekker, Inc.
P.J. Atkins, T.M. Crowder, The design and development of inhalation drug delivery system. In
A.J. Hickey (Ed.), Pharmaceutical Inhalation Aerosol Technology (2nd Ed., pp. 279-310); New
York: Marcel Dekker, Inc.
A.J. Hickey, Summary of common approaches to pharmaceutical aerosol administration. In A.J.
Hickey (Ed.), Pharmaceutical Inhalation Aerosol Technology (2nd Ed., pp. 385-421); New York:
Marcel Dekker, Inc.
35
[18]
[19]
[20]
[21]
[22]
[23]
[24]
[25]
[26]
[27]
[28]
[29]
[30]
[31]
[32]
[33]
S.P. Newman, P.G.D. Pellow, Evaluation of jet nebulizers. In D. Ganderton, T. Jones (Eds.), Drug
Delivery to the Respiratory Tract (pp. 124-132); New York: VCH Publishers.
R. Dalby, J. Suman, Inhalation therapy: technological milestones in asthma treatment. Advanced
Drug Delivery Reviews, 2003; 55: 779-791.
H.D.C. Smyth, The influence of formulation variables on the performance of alternative
propellant-driven metered-dose inhalers. Advanced Drug Delivery Reviews, 2003; 55: 807-828.
G.W. Hallworth, The formulation and evaluation of pressurized metered-dose inhalers. In D.
Ganderton, T. Jones (Eds.), Drug Delivery to the Respiratory Tract (pp. 87-118); New York: VCH
Publishers.
H.K. Chan, N.Y.K. Chew, Novel alternative methods for the delivery of drugs for the treatment of
asthma. Advanced Drug Delivery Reviews, 2003; 55: 793-805.
C. Fenton, G.M. Keating, G.L. Plosker, Novolizer A Multidose Dry Powder Inhaler. Drugs,
2003; 63(22): 2437-2445.
A.R. Clark, Pulmonary delivery technology: recent advances and potential for the new
millennium. In A.J. Hickey (Ed.), Pharmaceutical Inhalation Aerosol Technology (2nd Ed., pp.
571-592); New York: Marcel Dekker, Inc.
M.R. Feddah, K.F. Brown, E.M. Gipps, N.M. Davies, In-vitro characteristics of metered dose
inhaler versus dry powder inhaler glucocorticoid products: influence of inspiratory flow rates. J
Pharm Pharmaceut Sci, 2000; 3(3): 317-324.
M.B. Dolovich, Practical aspects of imaging techniques employed to study aerosol deposition and
clearance. In A.J. Hickey (Ed.), Pharmaceutical Inhalation Aerosol Technology (2nd Ed., pp. 171214); New York: Marcel Dekker, Inc.
J.A. Brydson, Plastics Materials(7th Ed.), Oxford, UK: Butterworth-Heinemann.
N.N. Prasad, D. V. Reddy. Lip–Nose Complex Anthropometry. International Journal of
Cosmetic Surgery and Aesthetic Dermatology, 2002. 4(3), 155-159.
KV Blake, M Hoppe, E Harman and L Hendeles. Relative amount of albuterol delivered to lung
receptors from a metered- dose inhaler and nebulizer solution. Bioassay by histamine
bronchoprovocation. Chest, 1992. 101, 309-315.
Corcoran, et al. Unpublished data.
W.J. Germann, C.L. Stanfield, Principles of Human Physiology, New York: Benjamin Cummings.
http://www.quickparts.com
Heyder J, et al. Deposition of particles in human respiratory tract in the size range of 0.005-15 µm.
J Aerosol Sci; 17: 811–25.
36