U.S. patent application number 11/177719 was filed with the patent office on 2005-11-03 for aerodynamically light particles for pulmonary drug delivery.
Invention is credited to Ben-Jebria, Abdell Aziz, Caponetti, Giovannia, Edwards, David A., Hanes, Justin, Hrkach, Jeffrey S., Langer, Robert S., Lotan, Noah.
Application Number | 20050244341 11/177719 |
Document ID | / |
Family ID | 27803961 |
Filed Date | 2005-11-03 |
United States Patent
Application |
20050244341 |
Kind Code |
A1 |
Edwards, David A. ; et
al. |
November 3, 2005 |
Aerodynamically light particles for pulmonary drug delivery
Abstract
Improved aerodynamically light particles for delivery to the
pulmonary system, and methods for their preparation and
administration are provided. In a preferred embodiment, the
aerodynamically light particles are made of a biodegradable
material and have a tap density less than 0.4 g/cm.sup.3 and a mass
mean diameter between 5 .mu.m and 30 .mu.m. The particles may be
formed of biodegradable materials such as biodegradable polymers.
For example, the particles may be formed of a functionalized
polyester graft copolymer consisting of a linear a-hydroxy-acid
polyester backbone having at least one amino acid group
incorporated herein and at least on poly(amino acid) side chain
extending from an amino acid group in the polyester backbone. In
one embodiment, aerodynamically light particles having a large mean
diameter, for example greater than 5 .mu.m, can be used for
enhanced delivery of a therapeutic or diagnostic agent to the
alveolar region of the lung. The aerodynamically light particles
optionally can incorporate a therapeutic or diagnostic agent, and
may be effectively aerosolized for administration to the
respiratory tract to permit systemic or local delivery of a wide
variety of incorporated agents.
Inventors: |
Edwards, David A.; (Boston,
MA) ; Caponetti, Giovannia; (Piacenza, IT) ;
Hrkach, Jeffrey S.; (Somerville, MA) ; Lotan,
Noah; (Haifa, IL) ; Hanes, Justin; (Baltimore,
MD) ; Ben-Jebria, Abdell Aziz; (State College,
PA) ; Langer, Robert S.; (Newton, MA) |
Correspondence
Address: |
ELMORE CRAIG & VANSTONE, P.C.
209 MAIN STREET
N. CHELMSFORD
MA
01863
US
|
Family ID: |
27803961 |
Appl. No.: |
11/177719 |
Filed: |
July 8, 2005 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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11177719 |
Jul 8, 2005 |
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10090068 |
Mar 1, 2002 |
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11177719 |
Jul 8, 2005 |
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10090418 |
Mar 1, 2002 |
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10090418 |
Mar 1, 2002 |
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09194068 |
Apr 8, 1999 |
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6503480 |
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09194068 |
Apr 8, 1999 |
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PCT/US97/08895 |
May 23, 1997 |
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Current U.S.
Class: |
424/46 ;
514/10.2; 514/177; 514/5.9 |
Current CPC
Class: |
A61K 38/28 20130101;
A61K 9/0075 20130101; A61K 47/34 20130101; A61K 31/57 20130101;
A61K 9/1647 20130101; A61K 9/1682 20130101; G06F 16/284
20190101 |
Class at
Publication: |
424/046 ;
514/003; 514/177 |
International
Class: |
A61K 038/28; A61L
009/04; A61K 009/14; A61K 031/57 |
Claims
What is claimed:
1. A method of increasing systemic bioavailability of a hormone
administered by inhalation comprising: administering to the
respiratory system of a patient or animal in need of said hormone
aerodynamically light particles that have a mass mean diameter
greater than 5 .mu.m, an aerodynamic diameter less than 4.7 .mu.m
and that include said hormone, wherein the particles are delivered
and deposited to the patient's or animal's lungs and the hormone is
released in the patient's or animal's blood stream for at least 4
hours.
2. The method of claim 1 wherein the hormone is insulin.
3. The method of claim 1 wherein the hormone is testosterone.
4. The method of claim 1 wherein the particles further include a
biodegradable material.
5. The method of claim 1 wherein the mass mean diameter is greater
than 10 .mu.m.
6. The method of claim 1 wherein the mass mean diameter is greater
than 20 .mu.m.
7. The method of claim 1 wherein the hormone is released in the
patient's or animal's blood stream for at least 10 hours.
8. The method of claim 1 wherein the hormone is released in the
patient's or animal's blood stream for at least 24 hours.
9. The method of claim 1 wherein the hormone is released in the
patient's or animal's blood stream for at least 48 hours.
10. A method of delivering a hormone to the pulmonary system to a
patient or animal, comprising: administering, via inhalation,
particles that include a hormone and a biodegradable material, have
an aerodynamic diameter less than about 4.7 .mu.m and a mass mean
diameter greater than about 5 .mu.m, wherein the hormone is
delivered and deposited in the patient's or animal's lungs and is
released in the patient's or animal's blood stream for at least 4
hours.
11. A method of increasing the bioavailabilty of a hormone,
comprising: administering to a patient or animal, via inhalation,
particles that include a hormone and a biodegradable material, have
an aerodynamic diameter less than about 4.7 .mu.m and a mass mean
diameter greater than about 5 .mu.m, wherein the hormone is
delivered and deposited in the patient's or animal's lungs and is
released in the patient's or animal's blood stream for at least 4
hours.
12. A method for making aerodynamically light particles for
administration of a hormone to the respiratory tract by inhalation,
the method comprising forming particles that include the hormone
and a biodegradable material and that have an aerodynamic diameter
that is less than 4.7 .mu.m and a mass mean diameter that is
greater than 5 .mu.m.
13. The method of claim 12 wherein the hormone is insulin.
14. The method of claim 12 wherein the hormone is testosterone.
15. The method of claim 12 wherein the mass mean diameter is
greater than 10 .mu.m.
16. The method of claim 12 wherein the mass mean diameter is
greater than 20 .mu.m.
17. The method of claim 12 wherein the aerodynamically light
particles are made by spray-drying.
18. The method of claim 12 wherein the aerodynamically light
particles are made by solvent evaporation.
Description
BACKGROUND OF THE INVENTION
[0001] The present application relates generally to biodegradable
particles of low density and large size for delivery to the
pulmonary system.
[0002] Biodegradable particles have been developed for the
controlled-release and delivery of protein and peptide drugs.
Langer, R., Science, 249: 1527-1533 (1990). Examples include the
use of biodegradable particles for gene therapy (Mulligan, R. C.
Science, 260: 926-932 (1993)) and for `single-shot` immunization by
vaccine delivery (Eldridge et al., Mol. Immunol., 28: 287-294
(1991)).
[0003] Aerosols for the delivery of therapeutic agents to the
respiratory tract have been developed. Adjei, A. and Garren, J.
Pharm. Res. 7, 565-569 (1990); and Zanen, P. and Lamm, J.-W. J.
Int. J. Pharm. 114, 111-115 (1995). The respiratory tract
encompasses the upper airways, including the oropharynx and larynx,
followed by the lower airways, which include the trachea followed
by bifurcations into the bronchi and bronchioli. The upper and
lower airways are called the conducting airways. The terminal
bronchioli then divide into respiratory bronchioli which then lead
to the ultimate respiratory zone, the alveoli, or deep lung. Gonda,
I. "Aerosols for delivery of therapeutic and diagnostic assents to
the respiratory tract," in Critical Reviews in Therapeutic Drug
Carrier Systems 6:273-313, 1990. The deep lung, or alveoli, are the
primary target of inhaled therapeutic aerosols for systemic drug
delivery.
[0004] Inhaled aerosols have been used for the treatment of local
lung disorders including asthma and cystic fibrosis (Anderson et
al., Am. Rev. Respir. Dis., 140: 1317-1324 (1989)) and have
potential for the systemic delivery of peptides and proteins as
well (Patton and Platz, Advanced Drug Delivery, Reviews, 8:179-196
(1992)). However, pulmonary drug delivery strategies present many
difficulties for the delivery of macromolecules; these include
protein denaturation during aerosolization, excessive loss of
inhaled drug in the oropharyngeal cavity (often exceeding 80%),
poor control over the site of deposition, irreproducibility of
therapeutic results owing to variations in breathing patterns, the
often too-rapid absorption of drug potentially resulting in local
toxic effects, and phagocytosis by lung macrophages.
[0005] Considerable attention has been devoted to the design of
therapeutic aerosol inhalers to improve the efficiency of
inhalation therapies. Timsina et. al., Int. J. Pharm. 101, 1-13
(1995); and Tansey, I. P., Spray Technol. Market 4, 26-29 (1994).
Attention has also been given to the design of dry powder aerosol
surface texture, regarding particularly the need to avoid particle
aggregation, a phenomenon which considerably diminishes the
efficiency of inhalation therapies. French, D. L., Edwards, D. A.
and Niven, R. W., J. Aerosol Sci. 27, 769-783 (1996). Attention has
not been given to the possibility of using large particle size
(greater than 5 .mu.m) as a means to improve aerosolization
efficiency. despite the fact that intraparticle adhesion diminishes
with increasing particle size. French, D. L., Edwards, D. A. and
Niven, R. W. J. Aerosol Sci. 27, 769-783 (1996). This is because
particles of standard mass density (mass density near 1 g/cm.sup.3)
and mean diameters greater than 5 .mu.m are known to deposit
excessively in the upper airways or in the inhaler device. Heyder,
J. et al., J. Aerosol Sci., 17: 811-825 (1986). For this reason,
dry powder aerosols for inhalation therapy are generally produced
with mean diameters primarily in the range of less than 5 .mu.m.
Ganderton. D., J. Biopharmaceutical Sciences 3:101-105 (1992): and
Gonda, I. "Physico-Chemical Principles in Aerosol Delivery," in
Topics in Pharmaceutical Sciences 1991, Crommelin, D. J. and K. K.
Midha, Eds., Medpharm Scientific Publishers, Stuttgart, pp. 95-115,
1992. Large "carrier" particles (containing no drug) have been
co-delivered with therapeutic aerosols to aid in achieving
efficient aerosolization among other possible benefits. French, D.
L., Edwards, D. A. and Niven, R. W. J. Aerosol Sci. 27, 769-783
(1996).
[0006] Local and systemic inhalation therapies can often benefit
from a relatively slow controlled release of the therapeutic agent.
Gonda, I., "Physico-chemical principles in aerosol delivery," in:
Topics in Pharmaceutical Sciences 1991, D. J. A. Crommelin and K.
K. Midha, Eds., Stuttgart: Medpharm Scientific Publishers, pp.
95-117, (1992). Slow release from a therapeutic aerosol can prolong
the residence of an administered drug in the airways or acini, and
diminish the rate of drug appearance in the bloodstream. Also,
patient compliance is increased by reducing the frequency of
dosing. Langer, R., Science, 249:1527-1533 (1990); and Gonda, I.
"Aerosols for delivery of therapeutic and diagnostic agents to the
respiratory tract," in Critical Reviews in Therapeutic Drug Carrier
Systems 6:273-313, (1990).
[0007] The human lungs can remove or rapidly degrade hydrolytically
cleavable deposited aerosols over periods ranging from minutes to
hours. In the upper airways, ciliated epithelia contribute to the
"mucociliary escalator" by which particles are swept from the
airways toward the mouth. Pavia, D. "Lung Mucociliary Clearance,"
in Aerosols and the Lung: Clinical and Experimental Aspects,
Clarke, S. W. and Pavia, D., Eds., Butterworths, London, 1984.
Anderson et al., Am. Rev. Respir. Dis., 140: 1317-1324 (1989). In
the deep lungs, alveolar macrophages are capable of phagocytosing
particles soon after their deposition. Warheit, M. B. and Hartsky,
M. A., Microscopy Res. Tech. 26: 412-422 (1993); Brain, J. D.,
"Physiology and Pathophysiology of Pulmonary Macrophages," in The
Reticuloendothelial System, S. M. Reichard and J. Filkins, Eds.,
Plenum, New York, pp. 315-327, 1985; Dorries. A. M. and Valberg, P.
A., Am. Rev. Resp. Disease 146, 831-837 (1991); and Gehr, P. et al.
Microscopy Res. and Tech., 26, 423-436 (1993). As the diameter of
particles exceeds 3 .mu.m, there is increasingly less phagocytosis
by macrophages. Kawaguchi, H. et al., Biomaterials 7: 61-66 (1986);
Krenis, L. J. and Strauss, B., Proc. Soc. Exp. Med., 107:748-750
(1961); and Rudt, S. and Muller, R. H., J. Contr. Rel., 22: 263-272
(1992). However, increasing the particle size also minimizes the
probability of particles (possessing standard mass density)
entering the airways and acini due to excessive deposition in the
oropharyngeal or nasal regions. Heyder, J. et al., J. Aerosol Sci.,
17: 811-825 (1986). An effective dry-powder inhalation therapy for
both short and long term release of therapeutics, either for local
or systemic delivery, requires a powder that displays minimum
aggregation and is capable of avoiding or suspending the lung's
natural clearance mechanisms until drugs have been effectively
delivered.
[0008] There is a need for improved inhaled aerosols for pulmonary
delivery of therapeutic agents which are capable of delivering the
drug in an effective amount into the airways or the alveolar zone
of the lung. There further is a need for the development of drug
carriers for use as inhaled aerosols which are biodegradable and
are capable of controlled release of drug within the airways or in
the alveolar zone of the lung.
[0009] It is therefore an object of the present invention to
provide improved carriers for the pulmonary delivery of therapeutic
and diagnostic agents. It is a further object of the invention to
provide inhaled aerosols which are effective carriers for delivery
of therapeutic or diagnostic agents to the deep lung. It is another
object of the invention to provide carriers for pulmonary delivery
which avoid phagocytosis in the deep lung. It is a further object
of the invention to provide carriers for pulmonary delivery which
are capable of biodegrading and optionally releasing incorporated
agents at a controlled rate.
SUMMARY OF THE INVENTION
[0010] Improved aerodynamically light particles for delivery to the
pulmonary system, and methods for their preparation and
administration are provided. In a preferred embodiment, the
particles are made of a biodegradable material, have a tap density
less than 0.4 g/cm.sup.3 and a mean diameter between 5 .mu.m and 30
.mu.m. In one embodiment, for example, at least 90% of the
particles have a mean diameter between 5 .mu.m and 30 .mu.m. The
particles may be formed of biodegradable materials such as
biodegradable synthetic polymers, proteins, or other water-soluble
materials such as certain polysaccharides. For example, the
particles may be formed of a functionalized polyester graft
copolymer with a linear .alpha.-hydroxy-acid polyester backbone
with at least one amino acid residue incorporated per molecule
therein and at least one poly(amino acid) side chain extending from
an amino acid group in the polyester backbone. Other examples
include particles formed of water-soluble excipients, such as
trehalose or lactose, or proteins, such as lysozyme or insulin. The
particles can be used for delivery of a therapeutic or diagnostic
agent to the airways or the alveolar region of the lung. The
particles may be effectively aerosolized for administration to the
respiratory tract and can be used to systemically or locally
deliver a wide variety of incorporated agents. The particles
incorporating an agent can optionally be co-delivered with larger
carrier particles, not carrying an incorporated agent, which have,
for example, a mean diameter ranging between about 50 .mu.m and 100
.mu.m.
BRIEF DESCRIPTION OF THE DRAWINGS
[0011] FIG. 1 is a graph comparing total particle mass of
aerodynamically light and non-light, control particles deposited on
the nonrespirable and respirable stages of a cascade impactor
following aerosolization.
[0012] FIG. 2 is a graph comparing total particle mass deposited in
the trachea and after the carina (lungs) in rat lungs and upper
airways following intratracheal aerosolization during forced
ventilation of aerodynamically light poly(lactic
acid-co-lysine-graft-lysine) (PLAL-Lys) particles and control,
non-light poly(L-lactic acid) (PLA) particles.
[0013] FIG. 3 is a graph comparing total particle recovery of
aerodynamically light PLAL-Lys particles and control PLA particles
in rat lungs following bronchoalveolar lavage.
[0014] FIG. 4 is a graph representing serum insulin levels (ng/ml)
over time (hours) following administration via inhalation or
subcutaneous injection of porous PLGA particles.
[0015] FIG. 5 is a graph representing serum insulin levels (ng/ml)
over time (hours) following administration via inhalation or
subcutaneous injection of non-porous PLGA particles. Darkened
circles represent administration via inhalation. Darkened triangles
represent administration via subcutaneous injection. Empty diamonds
represent untreated controls.
[0016] FIG. 6 is a graph-representing serum glucose concentrations
(mg/dl) following administration of porous PLGA particles via
inhalation. Darkened circles represent administration via
inhalation. Darkened triangles represent untreated controls.
[0017] FIG. 7 is a graph representing serum testosterone levels
(ng/ml) over time (hours) following administration via inhalation
or subcutaneous injection of porous PLGA particles with a diameter
of 20.4 .mu.m. Darkened circles represent administration via
inhalation. Darkened triangles represent administration via
subcutaneous injection.
[0018] FIG. 8 is a graph representing serum testosterone levels
(ng/ml) over time (hours) following administration via inhalation
or subcutaneous injection of porous PLGA particles with a diameter
of 10.1 .mu.m. Darkened circles represent administration via
inhalation. Darkened triangles represent administration via
subcutaneous injection.
DETAILED DESCRIPTION OF THE INVENTION
[0019] Aerodynamically light, biodegradable particles for improved
delivery to the respiratory tract are provided. The particles can
incorporate a therapeutic or diagnostic agent, and can be used for
controlled systemic or local delivery of the agent to the
respiratory tract via aerosolization. In a preferred embodiment,
the particles have a tap density less than about 0.4 g/cm.sup.3.
Features of the particle which can contribute to low tap density
include irregular surface texture and porous structure.
Administration of the low density particles to the lung by
aerosolization permits deep lung delivery of relatively large
diameter therapeutic aerosols, for example, greater than 5 .mu.m in
mean diameter. A rough surface texture also can reduce particle
agglomeration and provide a highly flowable powder, which is ideal
for aerosolization via dry powder inhaler devices, leading to lower
deposition in the mouth, throat and inhaler device.
[0020] Density and Size of Aerodynamically Light Particles
[0021] Particle Size
[0022] The mass mean diameter of the particles can be measured
using a Coulter Counter. The aerodynamically light particles are
preferably at least about 5 microns in diameter. The diameter of
particles in a sample will range depending upon depending on
factors such as particle composition and methods of synthesis. The
distribution of size of particles in a sample can be selected to
permit optimal deposition within targeted sites within the
respiratory tract.
[0023] The particles may be fabricated or separated, for example by
filtration, to provide a particle sample with a preselected size
distribution. For example, greater than 30%, 50%, 70%, or 80% of
the particles in a sample can have a diameter within a selected
range of at least 5 .mu.m. The selected range within which a
certain percentage of the particles must fall may be, for example,
between about 5 and 30 .mu.m, or optionally between 5 and 15 .mu.m.
In one preferred embodiment, at least a portion of the particles
have a diameter between about 9 and 11 .mu.m. Optionally, the
particle sample also can be fabricated wherein at least 90%, or
optionally 95% or 99%, have a diameter within the selected range.
The presence of the higher proportion of the aerodynamically light,
larger diameter (at least about 5 .mu.m) particles in the particle
sample enhances the delivery of therapeutic or diagnostic agents
incorporated therein to the deep lung.
[0024] In one embodiment, in the particle sample, the interquartile
range may be 2 .mu.m, with a mean diameter for example of 7.5, 8.0,
8.5, 9.0, 9.5, 10.0, 10.5, 11.0, 11.5, 12.0, 12.5, 13.0 or 13.5
.mu.m. Thus, for example, at least 30%, 40%, 50% or 60% of the
particles may have diameters within the selected ranges of 5.5-7.5
.mu.m, 6.0-8.0 .mu.m, 6.5-8.5 .mu.m, 7.0-9.0 .mu.m, 7.5-9.5 .mu.m,
8.0-10.0 .mu.m, 8.5-10.5 .mu.m, 9.0-11.0 .mu.m, 9.5-11.5 .mu.m,
10.0-12.0 .mu.m, 10.5-12.5 .mu.m, 11.0-13.0 .mu.m, 11.5-13.5 .mu.m,
12.0-14.0 .mu.m, 12.5-14.5 .mu.m or 13.0-15.0 .mu.m. Preferably the
above-listed percentages of particles have diameters within a 1
.mu.m range, for example, 6.0-7.0 .mu.m, 10.0-11.0 .mu.m or
13.0-14.0 .mu.m.
[0025] Particles having a tap density less than about 0.4
g/cm.sup.3 and a mean diameter of at least about 5 .mu.m are more
capable of escaping inertial and gravitational deposition in the
oropharyngeal region than smaller or more dense particles, and are
targeted to the airways of the deep lung. The use of larger
particles (mean diameter greater than 5 .mu.m) is advantageous
since they are able to aerosolize more efficiently than smaller,
denser particles such as those currently used for inhalation
therapies.
[0026] In comparison to smaller, denser particles, the larger
(greater than 5 .mu.m) aerodynamically light particles also can
potentially more successfully avoid phagocytic engulfment by
alveolar macrophages and clearance from the lungs, due to size
exclusion of the particles from the phagocytes' cytosolic space.
For particles of statistically isotropic shape (on average,
particles of the powder possess no distinguishable orientation),
such as spheres with rough surfaces, the particle envelope volume
is approximately equivalent to the volume of cytosolic space
required within a macrophage for complete particle
phagocytosis.
[0027] Aerodynamically light particles thus are capable of a longer
term release of an incorporated diagnostic or therapeutic agent
than smaller, denser particles. Following inhalation,
aerodynamically light biodegradable particles can deposit in the
lungs (due to their relatively low tap density), and subsequently
undergo slow degradation and drug release, without the majority of
the particles being phagocytosed by alveolar macrophages. The agent
can be delivered relatively slowly into the alveolar fluid, and at
a controlled rate into the blood stream, minimizing possible toxic
responses of exposed cells to an excessively high concentration of
the agent. The aerodynamically light particles thus are highly
suitable for inhalation therapies, particularly in controlled
release applications.
[0028] The particles may be fabricated with the appropriate
material, surface roughness, diameter and tap density for localized
delivery to selected regions of the respiratory tract such as the
deep lung or upper airways. For example, higher density or larger
particles may be used for upper airway delivery, or a mixture of
different sized particles in a sample, provided with the same or
different incorporated agent may be administered to target
different regions of the lung in one administration.
[0029] Particle Density and Deposition
[0030] The particles have a diameter of at least about 5 .mu.m and
optionally incorporate a therapeutic or diagnostic agent. The
particles are preferably aerodynamically light. As used herein, the
phrase "aerodynamically light particles" refers to particles having
a tap density less than about 0.4 g/cm.sup.3. The tap density of
particles of a dry powder may be obtained using a GeoPyc.TM.
(Micrometrics Instrument Corp., Norcross, Ga. 30093). Tap density
is a standard measure of the envelope mass density. The envelope
mass density of an isotropic particle is defined as the mass of the
particle divided by the minimum sphere envelope volume within which
it can be enclosed.
[0031] Inertial impaction and gravitational settling of aerosols
are predominant deposition mechanisms in the airways and acini of
the lungs during normal breathing conditions. Edwards, D. A., J.
Aerosol Sci. 26:293-317 (1995). The importance of both deposition
mechanisms increases in proportion to the mass of aerosols and not
to particle (or envelope) volume. Since the site of aerosol
deposition in the lungs is determined by the mass of the aerosol
(at least for particles of mean aerodynamic diameter greater than
approximately 1 .mu.m), diminishing the tap density by increasing
particle surface irregularities and particle porosity permits the
delivery of larger particle envelope volumes into the lungs, all
other physical parameters being equal.
[0032] The low tap density particles have a small aerodynamic
diameter in comparison to the actual envelope sphere diameter. The
aerodynamic diameter, d.sub.aer, is related to the envelope sphere
diameter, d (Gonda, I., "Physico-chemical principles in aerosol
delivery," in Topics in Pharmaceutical Sciences 1991 (eds. D. J. A.
Crommelin and K. K. Midha), pp. 95-117, Stuttgart: Medpharm
Scientific Publishers. 1992) by the formula:
d.sub.aer=d{square root}.rho.
[0033] where the envelope mass .rho. is in units of g/cm.sup.3.
Maximal deposition of monodisperse aerosol particles in the
alveolar region of the human lung (approximately 60%) occurs for an
aerodynamic diameter of approximately d.sub.aer=3 .mu.m. Heyder, J.
et al., J. Aerosol Sci., 17: 811-825 (1986). Due to their small
envelope mass density, the actual diameter d of aerodynamically
light particles comprising a monodisperse inhaled powder that will
exhibit maximum deep-lung deposition is:
d=3/{square root}.rho. .mu.m (where .rho.<1 g/cm.sup.3);
[0034] where d is always greater than 3 .mu.m. For example,
aerodynamically light particles that display an envelope mass
density, .rho.=0.1 g/cm.sup.3, will exhibit a maximum deposition
for particles having envelope diameters as large as 9.5 .mu.m. The
increased particle size diminishes interparticle adhesion forces.
Visser, J., Powder Technology, 58:1-10. Thus large particle size
increases efficiency of aerosolization to the deep lung for
particles of low envelope mass density, in addition to contributing
to lower phagocytic losses.
[0035] Particle Materials
[0036] The aerodynamically light particles preferably are
biodegradable and biocompatible, and optionally are capable of
biodegrading at a controlled rate for release of an incorporated
therapeutic or diagnostic agent. The particles can be made of any
material which is capable of forming a particle having a tap
density less than about 0.4 g/cm.sup.3. Both inorganic and organic
materials can be used. Other non-polymeric materials (e.g. fatty
acids) may be used which are capable of forming aerodynamically
light particles as defined herein. Different properties of the
particle can contribute to the aerodynamic lightness including the
composition forming the particle, and the presence of irregular
surface structure or pores or cavities within the particle.
[0037] Polymeric Particles
[0038] The particles may be formed from any biocompatible, and
preferably biodegradable polymer, copolymer, or blend, which is
capable of forming particles having a tap density less than about
0.4 g/cm.sup.3.
[0039] Surface eroding polymers such as polyanhydrides may be used
to form the aerodynamically light particles. For example,
polyanhydrides such as poly[(p-carboxyphenoxy)-hexane anhydride]
(PCPH) may be used. Biodegradable polyanhydrides are described, for
example, in U.S. Pat. No. 4,857,311.
[0040] In another embodiment, bulk eroding polymers such as those
based on polyesters, including poly(hydroxy acids), can he used.
Preferred poly(hydroxy acids) are polyglycolic acid (PGA),
polylactic acid (PLA) and copolymers and coblends thereof. In one
embodiment, the polyester has incorporated therein a charged or
functionalizable group such as an amino acid.
[0041] Other polymers include polyamides, polycarbonates,
polyalkylenes such as polyethylene, polypropylene, poly(ethylene
glycol), poly(ethylene oxide), poly(ethylene terephthalate), poly
vinyl compounds such as polyvinyl alcohols, polyvinyl ethers, and
polyvinyl esters, polymers of acrylic and methacrylic acids,
celluloses and other polysaccharides, and peptides or proteins, or
copolymers or blends thereof. Polymers may be selected with or
modified to have the appropriate stability and degradation rates in
vivo for different controlled drug delivery applications.
[0042] Polyester Graft Copolymers
[0043] In one preferred embodiment, the aerodynamically light
particles are formed from functionalized polyester graft
copolymers, as described in Hrkach et al., Macromolecules,
28:4736-4739 (1995); and Hrkach et al., "Poly(L-Lactic
acid-co-amino acid) Graft Copolymers: A Class of Functional
Biodegradable Biomaterials" in Hydrogels and Biodegradable Polymers
for Bioapplications, ACS Symposium Series No. 627, Raphael M.
Ottenbrite et al., Eds., American Chemical Society, Chapter 8, pp.
93-101, 1996. The functionalized graft copolymers are copolymers of
polyesters, such as poly(glycolic acid) or poly(lactic acid), and
another polymer including functionalizable or ionizable groups,
such as a poly(amino acid). In a preferred embodiment, comb-like
graft copolymers are used which include a linear polyester backbone
having amino acids incorporated therein, and poly(amino acid) side
chains which extend from the amino acid residues in the polyester
backbone. The polyesters may be polymers of .alpha.-hydroxy acids
such as lactic acid, glycolic acid, hydroxybutyric acid and
hydroxyvaleric acid, or derivatives or combinations thereof. The
polymers can include ionizable side chains, such as polylysine and
polyaniline. Other ionizable groups, such as amino or carboxyl
groups, may be incorporated into the polymer, covalently or
noncovalently, to enhance surface roughness and porosity.
[0044] An exemplary polyester graft copolymer is poly(lactic
acid-co-lysine-graft-lysine) (PLAL-Lys), which has a polyester
backbone consisting of poly(L-lactic acid-co-L-lysine) (PLAL), and
grafted poly-lysine chains. PLAL-Lys is a comb-like craft copolymer
having a backbone composition, for example, of 98 mol % lactic acid
and 2 mol % lysine and poly(lysine) side chains extending from the
lysine sites of the backbone.
[0045] PLAL-Lys may be synthesized as follows. First, the PLAL
copolymer consisting of L-lactic acid units and approximately 1-2%
N .epsilon. carbobenzoxy-L-lysine (Z-L-lysine) units is synthesized
as described in Barrera et al., J. Am. Chem. Soc., 115:11010
(1993). Removal of the Z protecting groups of the randomly
incorporated lysine groups in the polymer chain of PLAL yields the
free e-amine which can undergo further chemical modification. The
use of the poly(lactic acid) copolymer is advantageous since it
biodegrades into lactic acid and lysine, which can be processed by
the body. The existing backbone lysine groups are used as
initiating sites for the growth of poly(amino acid) side
chains.
[0046] The lysine .epsilon.-amino groups of linear poly(L-lactic
acid-co-L-lysine) copolymers initiate the ring opening
polymerization of an amino acid N-.epsilon. carboxyanhydride (NCA)
to produce poly(L-lactic acid-co-amino acid) comb-like graft
copolymers. In a preferred embodiment, NCAs are synthesized by
reacting the appropriate amino acid with triphosgene. Daly et al.,
Tetrahedron Lett., 29:5859 (1988). The advantage of using
triphosgene over phosgene gas is that it is a solid material, and
therefore, safer and easier to handle. It also is soluble in THF
and hexane so any excess is efficiently separated from the
NCAs.
[0047] The ring opening polymerization of amino acid
N-carboxyanhydrides (NCAs) is initiated by nucleophilic initiators
such as amines, alcohols, and water. The primary amine initiated
ring opening polymerization of NCAs allows efficient control over
the degree of polymerization when the monomer to initiator ratio
(M/I) is less than 150. Kricheldorf, H. R. in Models of Biopolymers
by Ring-Opening Polymerization, Penczek, S., Ed., CRC Press, Boca
Raton, 1990. Chapter 1; Kricheldorf, H. R.
.alpha.-Aminoacid-N-Carboxy-Anhydrides and Related Heterocycles,
Springer-Verlag, Berlin, 1987; and Imanishi, Y. in Ring-Opening
Polymerization, Ivin, K. J. and Saegusa, T., Eds., Elsevier,
London, 1984, Volume 2, Chapter 8. Methods for using lysine
.epsilon.-amino groups as polymeric initiators for NCA
polymerizations are described in the art. Sela, M. et al., J. Am.
Chem. Soc., 78: 746 (1956).
[0048] In the reaction of an amino acid NCA with PLAL, the
nucleophilic primary .epsilon.-amino group of the lysine side chain
attacks C-5 of the NCA. This leads to ring opening to form an amide
linkage, accompanied by evolution of a molecule of CO.sub.2. The
amino group formed by the evolution of CO.sub.2 propogates the
polymerization by attacking subsequent NCA molecules. The degree of
polymerization of the poly(amino acid) side chains, the amino acid
content in the resulting graft copolymers and the physical and
chemical characteristics of the resulting copolymers can be
controlled by adjusting the ratio of NCA to lysine .epsilon.-amino
groups in the PLAL polymer, for example, by adjusting the length of
the poly(amino acid) side chains and the total amino acid
content.
[0049] The poly(amino acid) side chains grafted onto or
incorporated into the polyester backbone can include any amino
acid, such as aspartic acid, alanine or lysine, or mixtures
thereof. The functional groups present in the amino acid side
chains, which can be chemically modified include amino, carboxylic
acid, thiol, guanido, imidazole and hydroxyl groups. As used
herein, the term "amino acid" includes natural and synthetic amino
acids and derivatives thereof. The polymers can be prepared with a
range of side chain lengths. The side chains preferably include
between 10 and 100 amino acids, and have an overall amino acid
content between 7 and 72%. However, the side chains can include
more than 100 amino acids and can have an overall amino acid
content greater than 72%, depending on the reaction conditions.
Poly(amino acids) can be grafted to the PLAL backbone in any
suitable solvent. Suitable solvents include polar organic solvents
such as dioxane, DMF, CH.sub.2Cl.sub.2, and mixtures thereof. In a
preferred embodiment, the reaction is conducted in dioxane at room
temperature for a period of time between about 2 and 4 days.
[0050] Alternatively, the particles may be formed from polymers or
blends of polymers with different polyester/amino acid backbones
and grafted amino acid side chains. For example, poly(lactic
acid-co-lysine-graft-ala- nine-lysine) (PLAL-Ala-Lys), or a blend
of PLAL-Lys with poly(lactic acid-co-glycolic acid-block-ethylene
oxide) (PLGA-PEG) (PLAL-Lys-PLGA-PEG) may be used.
[0051] In the synthesis, the graft copolymers may be tailored to
optimize different characteristics of the aerodynamically light
particle including: i) interactions between the agent to be
delivered and the copolymer to provide stabilization of the agent
and retention of activity upon delivery; ii) rate of polymer
degradation and, thereby, rate of drug release profiles; iii)
surface characteristics and targeting capabilities via chemical
modification; and iv) particle porosity.
[0052] Therapeutic Agents
[0053] Any of a variety of therapeutic agents can be incorporated
within the particles, which can locally or systemically deliver the
incorporated agents following administration to the lungs of an
animal. Examples include synthetic inorganic and organic compounds
or molecules, proteins and peptides, polysaccharides and other
sugars, lipids, and nucleic acid molecules having therapeutic,
prophylactic or diagnostic activities. Nucleic acid molecules
include genes, antisense molecules which bind to complementary DNA
to inhibit transcription, ribozymes and ribozyme guide sequences.
The agents to be incorporated can have a variety of biological
activities, such as vasoactive agents, neuroactive agents,
hormones, anticoagulants, immunomodulating agents, cytotoxic
agents, prophylactic agents, antibiotics, antivirals, antisense,
antigens, and antibodies. In some instances, the proteins may be
antibodies or antigens which otherwise would have to be
administered by injection to elicit an appropriate response.
Compounds with a wide range of molecular weight, for example,
between 100 and 500,000 grams per mole, can be encapsulated.
[0054] Proteins are defined as consisting of 100 amino acid
residues or more; peptides are less than 100 amino acid residues.
Unless otherwise stated, the term protein refers to both proteins
and peptides. Examples include insulin and other hormones.
Polysaccharides, such as heparin, can also be administered.
[0055] Aerosols including the aerodynamically light particles are
useful for a variety of inhalation therapies. The particles can
incorporate small and large drugs, release the incorporated drugs
over time periods ranging from hours to months, and withstand
extreme conditions during aerosolization or following deposition in
the lungs that might otherwise harm the encapsulated agents.
[0056] The agents can be locally delivered within the lung or can
be systemically administered. For example, genes for the treatment
of diseases such as cystic fibrosis can be administered, as can
beta agonists for asthma. Other specific therapeutic agents include
insulin, calcitonin, leuprolide (or LHRH), G-CSF, parathyroid
hormone-related peptide, somatostatin, testosterone, progesterone,
estradiol, nicotine, fentanyl, norethisterone, clonidine,
scopolomine, salicylate, cromolyn sodium, salmeterol, formeterol,
albeterol, and valium.
[0057] Diagnostic Agents
[0058] Any of a variety of diagnostic agents can be incorporated
within the particles, which can locally or systemically deliver the
incorporated agents following administration to the lungs of an
animal, including gases and other imaging agents.
[0059] Gases
[0060] Any biocompatible or pharmacologically acceptable Was can be
incorporated into the particles or trapped in the pores of the
particles. The term gas refers to any compound which is a gas or
capable of forming a gas at the temperature at which imaging is
being performed. The gas may be composed of a single compound such
as oxygen, nitrogen, xenon, argon, nitrogen or a mixture of
compounds such as air. Examples of fluorinated gases include
CF.sub.4, C.sub.2F.sub.6, C.sub.3F.sub.8, C.sub.4F.sub.8, SF.sub.6,
C.sub.2F.sub.4, and C.sub.3F.sub.6.
[0061] Other Imaging Agents
[0062] Other imaging agents which may be utilized include
commercially available agents used in positron emission tomography
(PET), computer assisted tomography (CAT), single photon emission
computerized tomography, x-ray, fluoroscopy, and magnetic resonance
imaging (MRI).
[0063] Examples of suitable materials for use as contrast agents in
MRI include the gatalinium chelates currently available, such as
diethylene triamine pentacetic acid (DTPA) and gatopentotate
dimeglumine, as well as iron, magnesium, manganese, copper and
chromium.
[0064] Examples of materials useful for CAT and x-rays include
iodine based materials for intravenous administration, such as
ionic monomers typified by diatrizoate and iothalamate, non-ionic
monomers such as iopamidol, isohexol, and ioversol, non-ionic
dimers, such as iotrol and iodixanol, and ionic dimers, for
example, ioxagalte.
[0065] Particles incorporating these agents can be detected using
standard techniques available in the art and commercially available
equipment.
[0066] Formation of Aerodynamically Light Polymeric Particles
[0067] Aerodynamically light polymeric particles may be prepared
using single and double emulsion solvent evaporation, spray drying,
solvent extraction or other methods well known to those of ordinary
skill in the art. The particles may be made, for example, using
methods for making microspheres or microcapsules known in the
art.
[0068] Methods for making microspheres are described in the
literature, for example, in Mathiowitz and Langer, J. Controlled
Release 5, 13-22 (1987); Mathiowitz et al., Reactive Polymers 6,
275-283 (1987); and Mathiowitz et al., J. Appl. Polymer Sci. 35,
755-774 (1988). The selection of the method depends on the polymer
selection, the size, external morphology, and crystallinity that is
desired, as described, for example, by Mathiowitz et al., Scanning
Microscopy 4,329-340 (1990); Mathiowitz et al., J. Appl. Polymer
Sci. 45, 125-134 (1992); and Benita et al., J. Pharm. Sci. 73,
1721-1724 (1984).
[0069] In solvent evaporation, described for example, in
Mathiowitz, et al., (1990), Benita, and U.S. Pat. No. 4,272,398 to
Jaffe, a polymer is dissolved in a volatile organic solvent, such
as methylene chloride. Several different polymer concentrations can
be used, for example, between 0.05 and 0.20 g/ml. An agent to be
incorporated, either in soluble form or dispersed as fine
particles, is optionally added to the polymer solution, and the
mixture is suspended in an aqueous phase that contains a surface
active agent such as poly(vinyl alcohol). The resulting emulsion is
stirred until most of the organic solvent evaporates, leaving solid
microspheres, which may be washed with water and dried overnight in
a lyophilizer.
[0070] Microspheres with different sizes (typically between 1 and
1000 microns) and morphologies can be obtained. This method is
especially useful for relatively stable polymers such as polyesters
and polystyrene. However, labile polymers such as polyanhydrides
may degrade due to exposure to water. Solvent removal may be a
preferred method for preparing microspheres from these
polymers.
[0071] Solvent removal was primarily designed for use with
polyanhydrides. In this method, a therapeutic or diagnostic agent
can be dispersed or dissolved in a solution of a selected polymer
in a volatile organic solvent like methylene chloride. The mixture
can then be suspended in oil, such as silicon oil, by stirring, to
form an emulsion. As the solvent diffuses into the oil phase, the
emulsion droplets harden into solid polymer microspheres. Unlike
solvent evaporation, this method can be used to make microspheres
from polymers with high melting points and a wide range of
molecular weights. Microspheres having a diameter between one and
300 microns can be obtained using this procedure.
[0072] Targeting of Particles
[0073] Targeting molecules can be attached to the particles via
reactive functional groups on the particles. For example, targeting
molecules can be attached to the amino acid groups of
functionalized polyester graft copolymer particles, such as
PLAL-Lys particles. Targeting molecules permit binding interactions
of the particle with specific receptor sites, such as those within
the lungs. The particles can be targeted by attaching ligands which
specifically or non-specifically bind to particular targets.
Exemplary targeting molecules include antibodies and fragments
thereof including the variable regions, lectins, and hormones or
other organic molecules capable of specific binding to receptors on
the surfaces of the target cells.
[0074] Administration
[0075] The particles can be administered to the respiratory system
alone or in any appropriate pharmaceutically acceptable carrier,
such as a liquid, for example saline, or a powder. In one
embodiment, particles incorporating a prophylactic, therapeutic or
diagnostic agent are co-delivered with larger carrier particles
that do not include an incorporated agent. Preferably, the larger
particles have a mass mean diameter between about 50 and 100
.mu.m.
[0076] Aerosol dosage, formulations and delivery systems may be
selected for a particular therapeutic application, as described,
for example, in Gonda, I. "Aerosols for delivery of therapeutic and
diagnostic agents to the respiratory tract," in Critical Reviews in
Therapeutic Drub Carrier Systems, 6:273-313, 1990; and in Moren,
"Aerosol dosage forms and formulations," in: Aerosols in Medicine.
Principles, Diagnosis and Therapy, Moren, et al., Eds, Elsevier,
Amsterdam, 1985.
[0077] The greater efficiency of aerosolization by aerodynamically
light particles of relatively large size permits more of an
incorporated agent to be delivered than is possible with the same
mass of relatively dense aerosols. The relatively large particle
size also minimizes potential drug losses caused by particle
phagocytosis. When the particles are formed from biocompatible
polymers, the system can provide controlled release in the lungs
and long-time local action or systemic bioavailability of the
incorporated agent. Denaturation of macromolecular drugs can be
minimized during aerosolization since macromolecules are contained
and protected within a polymeric shell. The enzymatic degradation
of proteins or peptides can be minimized by co-incorporating
peptidase-inhibitors.
[0078] Diagnostic Applications
[0079] The particles can be combined with a pharmaceutically
acceptable carrier, then an effective amount for detection
administered to a patient via inhalation. Particles containing an
incorporated imaging agent may be used for a variety of diagnostic
applications, including detecting and characterizing tumor masses
and tissues.
[0080] The present invention will be further understood by
reference to the following non-limiting examples.
EXAMPLE 1
Synthesis of Aerodynamically Light Poly[(p-carboxyphenoxy)-hexane
anhydride] ("PCPH") Particles
[0081] Aerodynamically light poly[(p-carboxyphenoxy)-hexane
anhydride] ("PCPH") particles were synthesized as follows. 100 mg
PCPH (MW approximately 25,000) was dissolved in 3.0 mL methylene
chloride. To this clear solution was added 5.0 mL 1% w/v aqueous
polyvinyl alcohol (PVA, MW approximately 25,000, 88 mole %
hydrolyzed) saturated with methylene chloride, and the mixture was
vortexed (Vortex Genie 2, Fisher Scientific) at maximum speed for
one minute. The resulting milky-white emulsion was poured into a
beaker containing 95 mL 1% PVA and homogenized (Silverson
Homogenizers) at 6000 RPM for one minute using a 0.75 inch tip.
After homogenization, the mixture was stirred with a magnetic
stirring bar and the methylene chloride quickly extracted from the
polymer particles by adding 2 mL isopropyl alcohol. The mixture was
stirred for 35 minutes to allow complete hardening of the
microparticles. The hardened particles were collected by
centrifugation and washed several times with double distilled
water. The particles were freeze dried to obtain a free-flowing
powder void of clumps. Yield, 85-90%.
[0082] The mean diameter of this batch was 6.0 .mu.m, however,
particles with mean diameters ranging from a few hundred nanometers
to several millimeters may be made with only slight modifications.
Scanning electron micrograph photos of a typical batch of PCPH
particles showed the particles to be highly porous with irregular
surface shape. The particles had a tap density less than 0.4
g/cm.sup.3.
EXAMPLE 2
Synthesis of PLAL-Lys and PLAL-Lys-Ala Polymeric and Copolymeric
Particles
[0083] Aerodynamically Light PLAL-Lys Particles
[0084] PLAL-Lys particles were prepared by dissolving 50 mg of the
graft copolymer in 0.5 ml dimethylsulfoxide, then adding 1.5 ml
dichloromethane dropwise. The polymer solution is emulsified in 100
ml of 5% w/v polyvinyl alcohol solution (average molecular weight
25 KDa, 88% hydrolyzed) using a homogenizer (Silverson) at a speed
of approximately 7500 rpm. The resulting dispersion was stirred
using a magnetic stirrer for 1 hour. Following this period, the pH
was brought to between 7.0 and 7.2 by addition of a 0.1 N NaOH
solution. Stirring was continued for an additional 2 hours until
the dichloromethane was completely evaporated and the particles
hardened. The particles were then isolated by centrifugation at
4000 rpm (1600 g) for 10 minutes (Sorvall RC-5B). The supernatant
was discarded and the precipitate washed three times with distilled
water followed by centrifugation for 10 minutes at 4000 rpm each
time. Finally, the particles were resuspended in 5 ml of distilled
water, the dispersion frozen in liquid nitrogen, and lyophilized
(Labconco freeze dryer 8) for at least 48 hours. Particle sizing
was performed using a Coulter counter. Average particle mean
diameters ranged from between 100 nm and 14 .mu.m, depending upon
processing parameters such as homogenization speed and time. All
particles exhibited tap densities less than 0.4 g/cm.sup.3.
Scanning electron micrograph photos of the particles showed them to
be highly porous with irregular surfaces.
[0085] Aerodynamically Light PLAL-Ala-Lys Particles
[0086] 100 mg of PLAL-Ala-Lys was completely dissolved in 0.4 ml
trifluoroethanol, then 1.0 ml methylene chloride was added
dropwise. The polymer solution was emulsified in 100 ml of 1% w/v
polyvinyl alcohol solution (average molecular weight 25 KDa, 80%
hydrolyzed) using a sonicator (Sonic&Material VC-250) for 15
seconds at an output of 40 W. 2 ml of 1% PVA solution was added to
the mixture and it was vortexed at the highest speed for 30
seconds. The mixture was quickly poured into a beaker containing
100 ml 0.3% PVA solution, and stirred for three hours allowing
evaporation of the methylene chloride. Scanning electron micrograph
photos of the particles showed them to possess highly irregular
surfaces.
[0087] Aerodynamically Light Copolymer Particles
[0088] Polymeric aerodynamically light particles consisting of a
blend of PLAL-Lys and PLGA-PEG were made. 50 mg of the PLGA-PEG
polymer (molecular weight of PEG: 20 KDa, 1:2 weight ratio of
PEG:PLGA., 75:25 lactide:glycolide) was completely dissolved in 1
ml dichloromethane. 3 mg of poly(lactide-co-lysine)-polylysine
graft copolymer was dissolved in 0.1 ml dimethylsulfoxide and mixed
with the first polymer solution. 0.2 ml of TE buffer, pH 7.6, was
emulsified in the polymer solution by probe sonication
(Sonic&Material VC-250) for 10 seconds at an output of 40 W. To
this first emulsion, 2 ml of distilled water was added and mixed
using a vortex mixer at 4000 rpm for 60 seconds. The resulting
dispersion was agitated by using a magnetic stirrer for 3 hours
until methylene chloride was completely evaporated and microspheres
formed. The spheres were then isolated by centrifugation at 5000
rpm for 30 min. The supernatant was discarded, the precipitate
washed three times with distilled water and resuspended in 5 ml of
water. The dispersion was frozen in liquid nitrogen and lyophilized
for 48 hours.
[0089] By scanning electron microscopy (SEM), the PLAL-Lys-PLGA-PEG
particles were highly surface rough and porous. The particles had a
mean particle diameter of b 7 .mu.m. The blend of PLAL-Lys with
poly(lactic acid) (PLA) and/or PLGA-PEG copolymers can be adjusted
to adjust particle porosity and size.
[0090] Variables which may be manipulated to alter the size
distribution of the particles include: polymer concentration,
polymer molecular weight, surfactant type (e.g., PVA, PEG, etc.),
surfactant concentration, and mixing intensity. Variables which may
be manipulated to alter the surface shape and porosity of the
particles include: polymer concentration, polymer molecular weight,
rate of methylene chloride extraction by isopropyl alcohol (or
another miscible solvent), volume of isopropyl alcohol added,
inclusion of an inner water phase, volume of inner water phase,
inclusion of salts or other highly water-soluble molecules in the
inner water phase which leak out of the hardening sphere by osmotic
pressure, causing the formation of channels, or pores, in
proportion to their concentration, and surfactant type and
concentration.
[0091] Additionally, processing parameters such as homogenization
speed and time can be adjusted. Neither PLAL, PLA nor PLGA-PEG
alone yields an aerodynamically light structure when prepared by
these techniques.
EXAMPLE 3
Synthesis of Spray-Dried Particles
[0092] Aerodynamically Light Particles Containing Polymer and Drug
Soluble in Common Solvent
[0093] Aerodynamicatiy light 50:50 PLGA particles were prepared by
spray drying with testosterone encapsulated within the particles
according to the following procedures. 2.0 g poly
(D,L-lactic-co-glycolic acid) with a molar ratio of 50:50 (PLGA
50:50, Resomer RG503, B. I. Chemicals, Montvale, N.J.) and 0.50 g
testosterone (Sigma Chemical Co., St. Louis, Mo.) were completely
dissolved in 100 mL dichloromethane at room temperature. The
mixture was subsequently spraydried through a 0.5 mm nozzle at a
flow rate of 5 mL/min using a Buchi laboratory spray-drier (model
190, Buchi, Germany). The flow rate of compressed air was 700 nl/h.
The inlet temperature was set to 30.degree. C. and the outlet
temperature to 25.degree. C. The aspirator was set to achieve a
vacuum of -20 to -25 bar. The yield was 51% and the mean particle
size was approximately 5 .mu.m. The particles were aerodynamically
light, as determined by a tap density less than or equal to 0.4
g/cm.sup.3.
[0094] Larger particle size can be achieved by lowering the inlet
compressed air flow rate, as well as by changing other variables.
Porosity and surface roughness can be increased by varying the
inlet and outlet temperatures, among other factors.
[0095] Aerodynamically Light Particles Containing Polymer and Drug
in Different Solvents
[0096] Aerodynamically light PLA particles with a model hydrophilic
drug (dextran) were prepared by spray drying using the following
procedure. 2.0 mL of an aqueous 10% w/v FITC-dextran (MW 70,000,
Sigma Chemical Co.) solution was emulsified into 100 mL of a 2% w/v
solution of poly (D,L-lactic acid) (PLA, Resomer R206, B. I.
Chemicals) in dichloromethane by probe sonication (Vibracell
Sonicator, Branson). The emulsion was subsequently spray-dried at a
flow rate of 5 mL/min with an air flow rate of 700 nl/h (inlet
temperature=30.degree. C., outlet temperature=21.degree. C., -20
mbar vacuum). The yield is 56%. The particles were aerodynamically
light (tap density less than 0.4 g/cm.sup.3).
[0097] Aerodynamically Light Protein Particles
[0098] Aerodynamically light lysozyme particles were prepared by
spray drying using the following procedure. 4.75 g lysozyme (Sigma)
was dissolved in 95 mL double distilled water (5% w/v solution) and
spray-dried using a 0.5 mm nozzle and a Buchi laboratory
spray-drier. The flow rate of compressed air was 725 nl/h. The flow
rate of the lysozyme solution was set such that, at a set inlet
temperature of 97-100.degree. C., the outlet temperature is between
55 and 57.degree. C. The aspirator was set to achieve a vacuum of
-30 mbar. The enzymatic activity of lysozyme was found to be
unaffected by this process and the yield of the aerodynamically
light particles (tap density less than 0.4 g/cm.sup.3) was 66%.
[0099] Aerodynamically Light High-Molecular Weight Water-Soluble
Particles
[0100] Aerodynamically light dextran particles were prepared by
spray drying using the following procedure. 6.04 g DEAE dextran
(Sigma) was dissolved in 242 mL double distilled water (2.5 % w/v
solution) and spray-dried using a 0.5 mm nozzle and a Buchi
laboratory spray-drier. The flow rate of compressed air was 750
nl/h. The flow rate of the DEAE-dextran solution was set such that,
at a set inlet temperature of 155.degree. C., the outlet
temperature was 80.degree. C. The aspirator was set to achieve a
vacuum of -20 mbar. The yield of the aerodynamically light
particles (tap density less than 0.4 g/cm.sup.3) was 66% and the
size range ranged between 1 and 15 .mu.m.
[0101] Aerodynamically Light Low-Molecular Weight Water-Soluble
Particles
[0102] Aerodynamically light trehalose particles were prepared by
spray drying using the following procedure. 4.9 g trehalose (Sigma)
was dissolved in 192 mL double distilled water (2.5% w/v solution)
and spray-dried using a 0.5 mm nozzle and a Buchi laboratory
spray-drier. The flow rate of compressed air 650 nl/h. The flow
rate of the trehalose solution was set such that, at a set inlet
temperature of 100.degree. C., the outlet temperature was
60.degree. C. The aspirator was set to achieve a vacuum of -30
mbar. The yield of the aerodynamically light particles (tap density
less than 0.4 g/cm.sup.3) was 36% and the size range ranged between
1 and 15 .mu.m.
[0103] Aerodynamically Light Low-Molecular Weight Water-Soluble
Particles
[0104] Polyethylene glycol (PEG) is a water-soluble macromolecule,
however, it cannot be spray dried from an aqueous solution since it
melts at room temperatures below that needed to evaporate water.
PEG was spray-dried at low temperatures from a solution in
dichloromethane, a low boiling organic solvent. Aerodynamically
light PEG particles were prepared by spray drying using the
following procedure. 5.0 g PEG (MW 15,000-20,000, Sigma) was
dissolved in 100 mL double distilled water (5.0% w/v solution) and
spray-dried using a 0.5 mm nozzle and a Buchi laboratory
spray-drier. The flow rate of compressed air was 750 nl/h. The flow
rate of the PEG solution was set such that, at a set inlet
temperature of 45.degree. C., the outlet temperature was
34-35.degree. C. The aspirator was set to achieve a vacuum of -22
mbar. The yield of the aerodynamically light particles (tap density
less than 0.4 g/cm.sup.3) was 67% and the size range ranged between
1 and 15 .mu.m.
EXAMPLE 4
Rhodamine Isothiocyanate Labeling of PLAL and PLAL-Lys
Particles
[0105] Aerodynamically light particles were compared with control
particles, referred to herein as "non-light" particles. Lysine
amine groups on the surface of aerodynamically light (PLAL-Lys) and
control, non-light (PLAL) particles, with similar mean diameters
(between 6 and 7 .mu.m) and size distributions (standard deviations
between 3 and 4 .mu.m) were labeled with Rhodamine isothiocyanate.
The tap density of the porous PLAL-Lys particles was 0.1 g/cm.sup.3
and that of the denser PLAL particles was 0.8 g/cm.sup.3.
[0106] The rhodamine-labeled particles were characterized by
confocal microscopy. A limited number of lysine functionalities on
the surface of the solid particle were able to react with rhodamine
isothiocyanate, as evidenced by the fluorescent image. In the
aerodynamically light particle, the higher lysine content in the
graft copolymer and the porous particle structure result in a
higher level of rhodamine attachment, with rhodamine attachment
dispersed throughout the interstices of the porous structure. This
also demonstrates that targeting molecules can be attached to the
aerodynamically light particles for interaction with specific
receptor sites within the lungs via chemical attachment of
appropriate targeting agents to the particle surface.
EXAMPLE 5
Aerosolization of PLAL and PLAL-Lys Particles
[0107] To determine whether large aerodynamically light particles
can escape (mouth, throat and inhaler) deposition and more
efficiently enter the airways and acini than nonporous particles of
similar size (referred to herein as non-light or control
particles), the aerosolization and deposition of aerodynamically
light PLAL-Lys (mean diameter 6.3 .mu.m) and control, non-light
PLAL (mean diameter 6.9 .mu.m) particles was compared in vitro
using a cascade impactor system.
[0108] 20 mg of the aerodynamically light or non-light
microparticles were placed in gelatine capsules (Eli Lilly), the
capsules loaded into a Spinhaler dry powder inhaler (DPI) (Fisons),
and the DPI activated. Particles were aerosolized into a Mark I
Andersen Impactor (Andersen Samplers, Ga.) from the DPI for 30
seconds at 28.3 l/min flow rate. Each plate of the Andersen
Impactor was previously coated with Tween 80 by immersing the
plates in an acetone solution (5% w/vol) and subsequently
evaporating the acetone in a oven at 60.degree. C. for 5 min. After
aerosolization and deposition, particles were collected from each
stage of the impactor system in separate volumetric flasks by
rinsing each stage with a NaOH solution (0.2 N) in order to
completely degrade the polymers. After incubation at 37.degree. C.
for 12 h, the fluorescence of each solution was measured
(wavelengths of 554 nm excitation, 574 nm emission).
[0109] Particles were determined as nonrespirable (mean aerodynamic
diameter exceeding 4.7 .mu.m: impactor estimate) if they deposited
on the first three stages of the impactor, and respirable (mean
aerodynamic diameter 4.7 .mu.m or less) if they deposited on
subsequent stages. FIG. 1 shows that less than 10% of the non-light
(PLAL) particles that exit the DPI are respirable. This is
consistent with the large size of the microparticles and their
standard mass density. On the other hand, greater than 55% of the
aerodynamically light (PLAL-Lys) particles are respirable, even
though the geometrical dimensions of the two particle types are
almost identical. The lower tap density of the aerodynamically
light (PLAL-Lys) microparticles is responsible for this improvement
in particle penetration, as discussed further below.
[0110] The non-light (PLAL) particles also inefficiently aerosolize
from the DPI; typically, less than 40% of the non-light particles
exited the Spinhaler DPI for the protocol used. The aerodynamically
light (PLAL-Lys) particles exhibited much more efficient
aerosolization (approximately 80% of the aerodynamically light
microparticles typically exited the DPI during aerosolization).
[0111] The combined effects of efficient aerosolization and high
respirable fraction of aerosolized particle mass means that a far
greater fraction of an aerodynamically light particle powder is
likely to deposit in the lungs than of a non-light particle
powder.
EXAMPLE 6
In Vivo Aerosolization of PLAL and PLAL-Lys Particles
[0112] The penetration of aerodynamically light and non-light
polymeric PLAL-Lys and PLAL microparticles into the lungs was
evaluated in an in vivo experiment involving the aerosolization of
the microparticles into the airways of live rats.
[0113] Male Spraque Dawley rats (150-200 g) were anesthetized using
ketamine (90 mg/kg)/xylazine (10 mg/kg). The anesthetized rat was
placed ventral side up on a surgical table provided with a
temperature controlled pad to maintain physiological temperature.
The animal was cannulated above the carina with an endotracheal
tube connected to a Harvard ventilator. The animal was force
ventilated for 20 minutes at 300 ml/min. 50 mg of aerodynamically
light (PLAL-Lys) or non-light (PLA) microparticles were introduced
into the endotracheal tube.
[0114] Following the period of forced ventilation the animal was
euthanized and the lungs and trachea were separately washed using
bronchoalveolar lavage. A tracheal cannula was inserted, tied into
place, and the airways were washed with 10 ml aliquots of HBSS. The
lavage procedure was repeated until a total volume of 30 ml was
collected. The lavage fluid was centrifuged (400 g) and the pellets
collected and resuspended in 2 ml of phenol red-free Hanks balanced
salt solution (Gibco, Grand Island, N.Y.) without Ca.sup.2 and
Mg.sup.2+ (HBSS). 100 ml were removed for particle counting using a
hemacytometer. The remaining solution was mixed with 10 ml of 0.4 N
NaOH. After incubation at 37.degree. C. for 12 h, the fluorescence
of each solution was measured (wavelengths of 554 nm excitation,
574 nm emission). FIG. 2 is a bar graph showing total particle mass
deposited in the trachea and after the carina (lungs) in rat lungs
and upper airways following intratracheal aerosolization during
forced ventilation. The PLAL-Lys aerodynamically light particles
had a mean diameter of 6.9 .mu.m. The non-light PLAL particles had
a mean diameter of 6.7 .mu.m. Percent tracheal aerodynamically
light particle deposition was 54.5, and non-light deposition was
77.0. Percent aerodynamically light particle deposition in the
lungs was 46.8 and non-light deposition was 23.0.
[0115] The non-light (PLAL) particles deposited primarily in the
trachea (approximately 79% of all particle mass that entered the
trachea). This result is similar to the in vitro performance of the
non-light microparticles and is consistent with the relatively
large size of the nonlight particles. Approximately 54% of the
aerodynamically light (PLAL-Lyis) particle mass deposited in the
trachea. Therefore, about half of the aerodynamically light
particle mass that enters the trachea traverses through the trachea
and into the airways and acini of the rat lungs, demonstrating the
effective penetration of the aerodynamically light particles into
the lungs.
[0116] Following bronchoalveolar lavage, particles remaining in the
rat lungs were obtained by careful dissection of the individual
lobes of the lungs. The lobes were placed in separate petri dishes
containing 5 ml of HBSS. Each lobe was teased through 60 mesh
screen to dissociate the tissue and was then filtered through
cotton gauze to remove tissue debris and connective tissue. The
petri dish and gauze were washed with an additional 15 ml of HBSS
to maximize microparticle collection. Each tissue preparation was
centrifuged and resuspended in 2 ml of HBSS and the number of
particles counted in a hemacytometer. The particle numbers
remaining in the lungs following the bronchoalveolar lavage are
shown in FIG. 3. Lobe numbers correspond to: 1) left lung, 2)
anterior, 3) median, 4) posterior, 5) postcaval. A considerably
greater number of aerodynamically light PLAL-Lys particles enters
every lobe of the lungs than the nonlight PLAL particles, even
though the geometrical dimensions of the two types of particles are
essentially the same. These results reflect both the efficiency of
aerodynamically light particle aerosolization and the propensity of
the aerodynamically light particles to escape deposition prior to
the carina or first bifurcation.
EXAMPLE 7
In Vivo Aerosolization of PLGA Porous and Non-Porous Particles
Including Insulin
[0117] Insulin was encapsulated into porous and nonporous polymeric
particles to test whether large particle size can increase systemic
bioavailability. The mass densities and mean diameters of the two
particles were designed such that they each possessed an
aerodynamic diameter (approximately 2 .mu.m) suitable for deep lung
deposition, with the mean diameter of the porous particles 5 .mu.m
and that of the nonporous particles less than 5 .mu.m (see FIGS.
4-6). Identical masses of the porous or nonporous particles were
administered to rats as an inhalation aerosol or injected
subcutaneously (controls).
[0118] Rats were anesthetized and cannulated as previously
described. The animal was force ventilated for between 30 and 20
minutes at 300 ml/min. Two types of aerosols were delivered to the
animal via the endotracheal tube. Following the period of forced
ventilation, the neck of the animal was sutured and the animal
revived within one to two hours. Blood samples (300 .mu.l) were
periodically withdrawn from the tail vein over a period of two to
six days. These samples were mixed with assay buffer, centrifuged,
and the supernatant examined for the presence of (endogenous and
exogenous) insulin or testosterone using radioimmunoassays (ICN
Pharmaceuticals, Costa Mesa, Calif.). Glucose was measured using a
calorimeter assay (Sigma). Control studies involved subcutaneous
injection of the same amount of powder as was inhaled. The
particles were injected into the scruff of the neck.
[0119] Serum insulin concentrations were monitored as a function of
time following inhalation or injection. For both porous (FIG. 4)
and nonporous (FIG. 5) particles, blood levels of insulin reach
high values within the first hour following inhalation. Only in the
case of the large porous particles do blood levels of insulin
remain elevated (p<0.05) beyond 4 h, with a relatively constant
insulin release continuing to at least 96 h (0.04<p<0.2).
[0120] These results are confirmed by serum glucose values which
show falling glucose levels for the first 10 h after inhalation of
the porous insulin particles, followed by relatively constant low
glucose levels for the remainder of the 96 h period, as shown in
FIG. 6. In the case of small nonporous insulin particles, initially
suppressed glucose values rose after 24 h.
[0121] Similar biphasic release profiles of macromolecules from
PLGA polymers have been reported in the literature (S. Cohen et al.
Pharm. Res. 8, 713 (1991)). For the large porous particles, insulin
bioavailability relative to subcutaneous injection is 87.5%,
whereas the small nonporous particles yield a relative
bioavailability of 12% following inhalation. By comparison,
bioavailability (relative to subcutaneous injection) of insulin
administered to rats as an inhalation liquid aerosol using a
similar endotracheal method has been reported as 37.3% (P.
Colthorpe et al. Pharm. Res. 9, 764 (1992)). Absolute
bioavailability of insulin inhaled into rat lungs in the form of a
lactose/insulin powder via a dry powder inhaler connected to an
endotracheal tube has been reported as 6.5% (F. Komada et al. J.
Pharm. Sci. 83, 863 (1994)).
[0122] Given the short systemic half life of insulin (11 minutes),
and the 12-24 h time scale of particle clearance from the central
and upper airways, the appearance of exogenous insulin in the
bloodstream several days following inhalation appears to indicate
that large porous particles achieve long, non-phagocytosed
life-times when administered to the deep lung. To test this
hypothesis, the lungs of rats were lavaged both immediately
following inhalation of the porous and nonporous insulin particles,
and 48 h after inhalation.
[0123] In the case of nonporous particles, 30%.+-.3% of phagocytic
cells contained particles immediately following inhalation, and
39%.+-.5% contained particles 48 h after inhalation. By contrast
only 8%.+-.2% of phagocytic cells contained large porous particles
right after inhalation, and 12.5%.+-.3.5% contained particles 48 h
after inhalation. In the small nonporous particle case,
17.5%.+-.1.5% of the phagocytic cell population contained 3 or more
particles 48 h after inhalation, compared to 4%.+-.1% in the case
of the large nonporous particles. Inflammatory response was also
elevated in the small nonporous particle case; neutrophils
represented 34%.+-.12% of the phagocytic cell population 48 h
following inhalation of the small nonporous particles, compared to
8.5%.+-.3.5% in tlle large porous particle case (alveolar
macrophages represented 100% of phagocytic cells immediately
following inhalation). These results support in vitro experimental
data appearing elsewhere that show phagocytosis of particles
diminishes precipitously as particle diameter increases beyond 3
.mu.m (H. Kawaguchi, et al. Biomaterials 7, 61 (1986). L. J.
Krrenis, and B. Strauss, Proc. Soc. Exp. Med. 107, 748 (1961). S.
Rudt, and R. H. Muller, J. Contr. Rel. 22, 263 (1992)).
EXAMPLE 8
In Vivo Aerosolization of PLGA Porous Particles Including
Testosterone
[0124] A second model drug, testosterone, was encapsulated in
porous particles of two different mean geometric diameters (10.1
.mu.m and 20.4 .mu.m) to further determine whether increased
bioavailability correlates with increasing size of porous
particles. An identical mass of powder was administered to rats as
an inhalation aerosol or as a subcutaneous injection (controls).
Serum testosterone concentrations were monitored as a function of
time following inhalation or injection (FIGS. 7 and 8). Blood
levels of testosterone remain well above background levels
(p<0.05) for between 12 and 24 h, even though the systemic
half-life of testosterone is between 10 and 20 minutes.
Testosterone bioavailability relative to subcutaneous injection is
177% for the 20.4 .mu.m diameter particles (FIG. 7) and 53% for the
10.1 .mu.m diameter porous particles (FIG. 8).
[0125] The increase in testosterone bioavailability with increasing
size of porous particles is especially notable given that the mean
diameter of the 20.4 .mu.m particles is approximately ten times
larger than that of nonporous conventional therapeutic particles
(D. Ganderton, J. Biopharmaceutical Sciences 3, 101 (1992). The
relatively short time scale of testosterone release observed both
for the inhalation and subcutaneous controls is near the several
hour in vitro time scale of release reported elsewhere for 50:50
PLGA microparticles of similar size encapsulating a therapeutic
substance (bupivacaine) of similar molecular weight and
lipophilicity (P. Le Corre et al. Int. J. Pharm. 107. 41
(1994)).
[0126] By making particles with high porosity, relatively large
particles (i.e., those possessing the same aerodynamic diameter as
smaller, nonporous particles) can enter the lungs, since it is
particle mass that dictates location of aerosol deposition in the
lungs. The increased aerosolization efficiency of large, light
particles lowers the probability of deposition losses prior to
particle entry into the airways, thereby increasing the systemic
bioavailability of an inhaled drug.
* * * * *