U.S. patent application number 13/866764 was filed with the patent office on 2013-12-05 for swellable particles for drug delivery.
This patent application is currently assigned to STC.UNM. The applicant listed for this patent is STC.UNM. Invention is credited to Martin J. Donovan, Hugh D. Smyth.
Application Number | 20130323310 13/866764 |
Document ID | / |
Family ID | 38625503 |
Filed Date | 2013-12-05 |
United States Patent
Application |
20130323310 |
Kind Code |
A1 |
Smyth; Hugh D. ; et
al. |
December 5, 2013 |
SWELLABLE PARTICLES FOR DRUG DELIVERY
Abstract
Swellable particles for delivering a working agent to the
pulmonary system comprise a plurality of biodegradable particles
each formed from a polymer network, each of the plurality of
biodegradable particles having a mass mean aerodynamic diameter not
exceeding 5 .mu.m, the particles being swellable by hydration to a
size that is greater than 6 .mu.m volume mean diameter, and a
working agent entrapped in the polymer network of each of the
plurality of biodegradable particles.
Inventors: |
Smyth; Hugh D.; (Austin,
TX) ; Donovan; Martin J.; (Austin, TX) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
STC.UNM |
Albuquerque |
NM |
US |
|
|
Assignee: |
STC.UNM
Albuquerque
NM
|
Family ID: |
38625503 |
Appl. No.: |
13/866764 |
Filed: |
April 19, 2013 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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13601532 |
Aug 31, 2012 |
8440231 |
|
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13866764 |
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|
11732489 |
Apr 3, 2007 |
8257685 |
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13601532 |
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60788942 |
Apr 4, 2006 |
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Current U.S.
Class: |
424/491 ;
424/489; 424/490; 514/40; 514/454 |
Current CPC
Class: |
A61K 9/0073 20130101;
A61K 9/0075 20130101; A61K 47/60 20170801; A61P 35/00 20180101;
A61K 47/6949 20170801; A61K 9/0078 20130101; B82Y 5/00 20130101;
A61P 11/00 20180101; Y10T 428/2982 20150115; A61K 48/00 20130101;
A61K 9/1641 20130101; A61P 11/06 20180101 |
Class at
Publication: |
424/491 ;
424/489; 424/490; 514/40; 514/454 |
International
Class: |
A61K 9/00 20060101
A61K009/00 |
Claims
1. Swellable particles for delivering a working agent to the
pulmonary system, the particles comprising: a plurality of
biodegradable particles each formed from a polymer network, each of
the plurality of biodegradable particles having a mass median
aerodynamic diameter not exceeding 5 .mu.m, the particles being
swellable by hydration to a size that is greater than 6 .mu.m
volume mean diameter; and a working agent entrapped in the polymer
network of each of the plurality of biodegradable particles.
2-3. (canceled)
4. The particles of claim 1, wherein the working agent is entrapped
in nanoparticles, wherein the nanoparticles are incorporated in the
biodegradable particles.
5. The particles of claim 1, wherein the working agent is
chemically bonded with the polymer network of the biodegradable
particle.
6. The particles of claim 1, wherein the working agent comprises
one or more of a therapeutic treating agent, a diagnostic agent, a
prophylactic agent, or an imaging agent.
7. The particles of claim 6, wherein the working agent comprises a
mucolytic agent.
8. The particles of claim 1, wherein the working agent comprises at
least one of an antibiotic agent, a cytotoxic agent, an RNA
interfering agent, and a gene.
9-11. (canceled)
12. The particles of claim 1, wherein the working agent comprises
multiple cytotoxic agents.
13. The particles of claim 1, wherein the working agent comprises a
cytotoxic agent and an RNA interfering agent.
14. The particles of claim 1, wherein at least 90% of the particles
have an aerodynamic diameter of 5 .mu.m or less and swell to a size
greater than 6 .mu.m volume mean diameter.
15. The particles of claim 1, wherein the polymer network of the
biodegradable particles comprises a material selected from the
group consisting of a biodegradable natural polymer, a synthetic
polymer, a protein, and a carbohydrate, or combinations
thereof.
16. The particles of claim 1, wherein the plurality of
biodegradable particles comprise hydrogel particles.
17. The particles of claim 1, including a coating on the plurality
of biodegradable particles that controls a rate of particle
swelling.
18. The particles of claim 17, wherein the coating comprises an
excipient selected from the group consisting of a carbohydrate, a
lipid, a protein, or a biocompatible salt of sodium, potassium,
calcium, magnesium or lithium.
19. The particles of claim 17, wherein the coating comprises a
targeting agent for binding to receptors or to a target within a
diseased site.
20. An aerosol comprising the swellable particles of claim 1.
21. The particles of claim 1, wherein the polymer network of each
of the plurality of biodegradable particles comprises a
cross-linked polymer network.
22. The particles of claim 21, wherein the cross-linked polymer
network physically entraps the working agent in the polymer
network.
23. The particles of claim 21, further comprising a crosslinker to
form crosslinking of polymers within the polymer network.
24. The particles of claim 23, wherein the crosslinker comprises
dithiothreitol,
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation of U.S. application Ser.
No. 13/601,532, filed Aug. 31, 2012, which is a divisional of U.S.
application Ser. No. 11/732,489, filed Apr. 3, 2007, which is a
non-provisional of and claims the benefit of priority under 35
U.S.C. .sctn.119(e) to U.S. Provisional Application No. 60/788,942,
filed Apr. 4, 2006, all of which are incorporated herein by
reference in their entirety.
BACKGROUND OF THE INVENTION
[0002] The present invention relates generally to swellable
biodegradable particles for use in delivering a therapeutic or
other agent to the pulmonary system and, more particularly, to
swellable biodegradable particles that have dehydrated sizes for
delivery to the pulmonary system and that swell on hydration in the
pulmonary tract to larger sizes to achieve controlled release of a
drug or other agent from the particle structure.
BACKGROUND OF THE INVENTION
[0003] Delivery of therapeutic agents to the pulmonary system has
been used for the treatment of local lung diseases such as asthma,
cystic fibrosis, and chronic obstructive pulmonary disease (A. J.
Hickey, editor, Inhalation Aerosols: Physical and Biological Basis
for Therapy, New York: Marcel Dekker, Inc. 1996). Relative to
systemic oral or injection drug delivery, local delivery of
respiratory drugs to the lungs provides advantages because it: (1)
requires smaller doses of the drug; and (2) minimizes systemic
toxicity by allowing delivery directly to the site of the disease.
Delivery of systemically acting agents has also been investigated,
such as for the administration of proteins and peptides (e.g.
insulin) as described by Patton et al., in "Inhaled insulin", Adv.
Drug. Deliv. Rev. 35, pp. 235-247 (1999). However, pulmonary
delivery of drugs is limited by several major issues including poor
efficiency of deposition in the respiratory tract and excessive
removal of the drug by the oropharyngeal cavity, poor control over
the site of deposition of the drug within the respiratory tract,
poor reproducibility of dosing due to the dependence on breathing
patterns of the patient, and too rapid clearance and/or absorption
of the drug from the pulmonary system potentially resulting in
inappropriate drug concentrations at the target site and even toxic
effects.
[0004] A controlled release delivery system for drugs delivered
locally to the lung would provide a very desirable method to
effectively treat respiratory and systemic diseases. Moreover,
controlled release of respiratory drugs may offer significant
clinical benefit to millions of patients with respiratory disease
by allowing them to take treatments for such diseases as asthma
less frequently and to receive more prolonged and controlled
relief. Controlled delivery of drugs to the lung also offers the
potential for improved safety by moderating the drug peaks and
troughs of immediate release drugs, which can cause added toxicity
or reduced efficacy. Also, controlling the release of two or more
therapeutic agents from a single particle system delivered to the
pulmonary system would have significant benefits for
co-localization of the agents within the respiratory tract. The
likelihood of synergism or additive effects between agents would be
significantly increased.
[0005] Currently available pulmonary delivery systems are not
ideal, delivering inaccurate doses, requiring frequent dosing and
losing significant amounts of drug in the delivery process. Most
asthma drugs delivered via inhalation are immediate-release
formulations that must be inhaled multiple times per day (Cochrane
et al. Inhaled Corticosteroids for Asthma Therapy: Patient
Compliance, Devices, and Inhalation Technique Chest. 117, pp.
542-550, 2000). This frequent inhalation tends to discourage
patient compliance. When patients forget to take their medicine
they may experience complications which may result in increased
emergency room visits and hospitalizations. In a recent analysis of
published studies of patient compliance with asthma medications,
patients took the recommended doses of medication on only 20 to 73%
of days (Cochrane et al. Inhaled Corticosteroids for Asthma
Therapy: Patient Compliance, Devices, and Inhalation Technique,
Chest. 117, pp. 542-550, 2000). The percentage of under-use days
ranged from 24 to 69%. In addition, immediate release formulations
often deliver drug levels that peak and trough, causing undesirable
toxicity or inadequate efficacy.
[0006] Although promising, inhaled formulations face difficult
challenges in maintaining effective drug concentration in the lungs
for extended periods. Factors contributing to the short duration of
drug action following pulmonary delivery include: (1) the rapid
mucociliary clearance rates (approximately 1.7-4.9 mm/min)
resulting in a very short half-life for inhaled particles
(approximately 0.5-2 hr) (Lansley, A. B., 1993. Mucociliary
clearance and drug delivery via the respiratory tract. Adv Drug Del
Rev. 11, 299-327); (2) phagocytosis of particles by the alveolar
macrophages; and; (3) rapid absorption of drug molecules
(Mw<1000 Da) to the systemic circulation with a mean half-time
for absorption of <2 hr.
[0007] The pulmonary region has several particle clearance
mechanisms. The relative importance of each clearance mechanism
varies depending on the physicochemical properties of the particle.
Particle retention in the pulmonary region is longer than that of
the ciliated airways. After deposition, uptake of particles by
alveolar macrophages is very rapid. An initial fast phase of
clearance is related to phagocytosis by alveolar macrophages.
[0008] There are limited technologies available to circumvent the
natural clearance mechanisms of the airways that largely prevent
sustained release particles from being effective. A number of prior
art references, including, but not limited to U.S. Pat. No.
6,136,295 to Edward, et al.; and U.S. Pat. No. 6,730,322 to
Bernstein, et al., describe particles that have been designed to
have low densities (large porous particles). Although geometrically
large, those particles are aerodynamically much smaller.
[0009] Generally to achieve sustained release, particles must be
delivered to the airways and avoid mucociliary clearance, uptake by
alveolar macrophages, and prevention of rapid absorption from the
lung. Avoiding mucociliary clearance can be achieved by avoiding
particle deposition in the tracheobronchial region where ciliated
epithelia are present. Generally an aerodynamic particle size must
be less than around 5 .mu.m to accomplish this. Once particles are
deposited in the peripheral airways where the mucociliary clearance
mechanism is not present, particles must avoid alveolar macrophage
uptake that can rapidly clear therapeutic compounds. Avoidance of
macrophages can be accomplished by (1) creating particles that are
not recognizable as foreign particulates (stealth particles); (2)
providing particles that are physically too large to be engulfed by
macrophages or which delay engulfment; or (3) providing particles
that are too small to be recognized by macrophages
(nanoparticles).
[0010] Current sustained release pulmonary systems as described by
the cited prior art generally comprise large porous particle
technologies. The main problems with these systems is the low drug
loading possible in the particle matrix, the special
physicochemical properties of the drug required for inclusion in
these particle systems, and the limits on how long drug may be
sustained. The present invention overcomes these problems by using
swelling particles to improve sustained release. The swelling
particles of the present invention include the drug or other
working agent being delivered on and/or in a biocompatible and
biodegradable swellable matrix that preferably enables deep lung
delivery and avoids clearance by the alveolar macrophages. In
addition, the matrix materials can be modified to modulate the drug
release characteristics or to improve compatibility of the drug
with the matrix system.
SUMMARY OF THE INVENTION
[0011] The present invention provides improved swellable particles
for delivery to the pulmonary system, and to a method for their
incorporation and administration of a working agent, such as
including but not limited to a therapeutic agent, diagnostic agent,
prophylactic agent or imaging agent. The swellable particles
include dehydrated (dry) aerodynamic particle diameters to enable
delivery to the respiratory tract, such as for example to the
tracheo-bronchial airways of the upper respiratory tract and/or to
the alveolar regions of the deep lung, and hydrated particle
diameters that are greater than 6 .mu.m volume mean diameter to
retard or prevent their phagocytosis by the macrophages present in
airways of the respiratory tract.
[0012] In an illustrative embodiment of the invention, the
dehydrated (dry) particles are made of a biodegradable material,
have a mass median aerodynamic particle diameter between 0.5 .mu.m
and 5 .mu.m, and are capable of swelling to a hydrated geometric
particle diameter greater than 6 .mu.m volume mean diameter. In a
preferred embodiment of the invention, at least 90%, more
preferably 95% to 99%, of the particles have an aerodynamic
particle diameter not exceeding 5 .mu.m and swell to a size of
greater than 6 .mu.m volume mean diameter. The particles may be
formed of biodegradable materials such as including, but not
limited to, a biodegradable natural or synthetic polymer, a
protein, a carbohydrate, or combinations thereof. For example, the
particles may be formed of a multi-branched polyethylene glycol
(PEG) hydrogel polymer. Other examples include particles formed of
biodegradable polymers such as dextrans,
hydroxyethylmethylacrylate, or other biocompatible and
biodegradable swellable polymeric systems. The swellable particles
can be used for enhanced delivery of one or more working agents to
the airways of the respiratory tract, including to the alveolar
region of the lung. The particles incorporating one or more working
agents may be effectively aerosolized for administration to the
respiratory tract to permit systemic or local delivery of a wide
variety of therapeutic and other agents. They optionally may be
co-delivered with larger carrier particles (not carrying a
therapeutic or other agent) which have for example a mean diameter
ranging between about 50 .mu.m and 150 .mu.m.
[0013] The present invention is advantageous in that the dehydrated
(dry) particles possess an aerodynamic diameter such that a) they
are able to reach one or more target regions of the respiratory
tract, including the tracheo-bronchial airways of the upper
respiratory tract and/or to the alveolar regions of the deep lung,
b) they can deliver a payload of one or more working agents without
premature release, c) they are swellable by hydration in the
airways to a size that retards or prevents their uptake and removal
by macrophages, and d) they can provide controlled release of the
working agent(s) at predictable rates following hydration.
[0014] Other features and advantages of the present invention will
become apparent from the following detailed description.
DETAILED DESCRIPTION OF THE INVENTION
[0015] The present invention provides swellable, biodegradable
particles for improved delivery of therapeutic and other working
agents to the respiratory tract. The working agents which can be
delivered via the particles include, but are not limited to, a
therapeutic agent, diagnostic agent, prophylactic agent, imaging
agent, or combinations thereof.
[0016] In an illustrative embodiment of the invention, the
swellable particles initially comprise dehydrated (dry) powder
particles having mass median aerodynamic particle diameter of 5
.mu.m or less to enable delivery to the respiratory tract, such as
for example to the tracheo-bronchial airways of the upper
respiratory tract and/or to the alveolic regions of the deep lung,
and having hydrated particle diameter that is greater than 6 .mu.m
volume mean diameter to retard or prevent their phagocytosis by the
macrophages present in airways of the respiratory tract. The
dehydrated (dry) particles typically have a mass median aerodynamic
diameter between 0.5 .mu.m and 5 .mu.m, and typically are capable
of swelling to a hydrated geometric diameter greater than 6 .mu.m
to 50 .mu.m volume mean diameter. At least 90%, more preferably 90%
to 95%, of the particles have an aerodynamic diameter of 5 .mu.m or
less and swell to a size of greater than 6 .mu.m volume mean
diameter.
[0017] The mass median aerodynamic diameter (MMAD) is typically
obtained from conventional aerosol sizing instruments, such as
cascade impactors and/or time of flight instruments such as the TSI
Aerodynamic Particle Sizer (TSI Incorporated, Shoreview, Minn.).
This size determination occurs where 50% of the mass of the
particles, when classified by their aerodynamic size, are below
this diameter (i.e. the MMAD). That is, 50% of the mass of
particles have a diameter lower than the MMAD and 50% of the mass
of particles have a diameter higher than the MMAD. This measure of
particle size converts the particle in question (which can have
different densities, shapes, surface and aerodynamic drag) into a
sphere having a density equal to 1 and provides an equivalent
sphere diameter, even though the particle may have a flake,
acicular or other non-spherical shape. For example, the aerodynamic
diameter is defined as the diameter of the spherical particle with
a density of 1 g/cm.sup.3 (the density of a water droplet) that has
the same settling velocity as the particle and is given by the
following equation (see Hinds, W. I., "Uniform Particle Motion," in
Aerosol Technology 1999, pp. 42-74, John Wiley and Sons Inc.):
d.sub.a=d.sub.p(.rho..sub.p).sup.1/2 where d.sub.p is the diameter
of the particle and .rho..sub.p is its density in g/cm.sup.3. The
aerodynamic diameter may be thought of as the diameter of a
spherical droplet of water possessing the same aerodynamic
properties as the particle. For example, if a particle has an
aerodynamic diameter of 1 .mu.m, it acts aerodynamically identical
to a 1 .mu.m water droplet regardless of the particle's actual
size, shape, or density. By adjusting the mean diameter and density
of an aerosol population, the particles can be tailored to possess
the exact aerodynamic diameter necessary for delivery to a
specified lung region. The diameters of the swellable particles in
a sample can range depending upon 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.
[0018] The equivalent sphere diameter allows one to directly
compare particles with different particle geometries and compare
them only on the basis of aerodynamics, which is the functional
size characteristic of importance for delivery of the particles to
the respiratory tract. The MMAD is selected to describe the
dehydrated (dry) particles because it is useful in predicting the
deposition of dry powder particles within the airways.
[0019] The volume mean diameter is used to describe the hydrated
particles because the particles will have swollen in size after
hydration in the airways. The volume mean diameter will increase
from the dry state to the hydrated state regardless of the particle
shape. The volume mean diameter is also an equivalent sphere
diameter whereby the particle in question is converted to a sphere
of equivalent volume, even though the particle may have a flake,
acicular or other non-spherical shape. In terms of functionality,
the volume of the particle is important because it is related to
how well the macrophage cells in the airways can clear the
particles; i.e. their volume is important for retarding or avoiding
the clearance mechanisms of the respiratory tract. The volume mean
diameter can be determined by testing as follows: using hydrogel
particles dispersed in buffer solution, a laser light scattering
instrument can be used to measure volume diameter changes in the
particle geometry upon hydration. A low energy liquid dispersion
attachment may also be used to minimize particle aggregation.
Alternatively, particle swelling can be quantified via confocal
microscopy.
[0020] The swellable particles may be formed from any
biodegradeable, and preferably biocompatible polymer, copolymer, or
blend, which is capable of forming particles with a mass median
aerodynamic diameter between 0.5 and 5 .mu.m, but can also swell to
a geometric diameter of greater than 6 .mu.m volume mean diameter.
For purposes of illustration and not limitation, the particles can
be formed of a swellable biodegradable natural polymer, synthetic
polymer, protein, carbohydrate, or combinations thereof. For
example, the particles may be formed of a multi-branched
polyethylene glycol (PEG) hydrogel polymer. Other examples include
particles formed of biodegradable polymers such as dextrans,
hydroxyethylmethylacrylate, or other biocompatible and
biodegradable swellable polymeric systems.
[0021] For purposes of further illustration and not limitation, the
swellable particles also can be made from bulk eroding hydrogel
polymers, such as those based on polyesters including poly(hydroxy
acids) can be used in practice of the invention. Moreover, surface
eroding polymers such as polyanhydrides may be used to form the
swelling particles. For example, polyanhydrides such as
poly[(p-carboxyphenoxy)-hexane anhydride] (PCPH) may be used. For
example, polyglycolic acid (PGA) or polylactic acid (PLA) or
copolymers thereof may be used to form the swellable particles,
wherein the polyester has incorporated therein a charged or
functionalizable group such as an amino acid as described below.
Other polymers include polyamides, polycarbonates, polyalkylenes
such as polyethylene, polypropylene, 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 which are capable of forming swellable particles described
above. Polymers may be selected with or be modified to have the
appropriate stability and degradation rates in vivo for different
controlled drug delivery applications.
[0022] Features of the swellable particle which can contribute to
swelling include degree of polymer cross-linking, monomer size, and
porosity. A rough particle 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. Moreover,
administration of the swellable particles to the lung by
aerosolization permits deep lung delivery of therapeutic aerosols
where the particles can swell after hydration on the airway
surfaces). In order to serve as efficient drug carriers in drug
delivery systems, the swellable particles preferably are
biodegradable and biocompatible, and optionally are capable of
biodegrading at a controlled rate for delivery of a drug.
[0023] In an illustrative embodiment of the invention,
biodegradeable and/or biocompatible hydrogel particles are formed
from acrylated 8-arm PEG (20 kDa) crosslinked with dithiothreitol,
as described by Hubbel et al. Journal of Controlled Release
76:11-25 (2001), the teachings of which are incorporated herein by
reference. Alternate methods of crosslinking the acrylated polymers
include photopolymerization, other covalent crosslinking methods
(polycysteine), ionic crosslinking, and physical crosslinking
(entanglements between highly branched and high molecular weight
polymers).
[0024] In the synthesis, both the molecular weight, degree of
branching (e.g. 8-arm to 4-arm), and the concentrations of the
reactants may be altered to change the pore size of the hydrogels,
and thereby adjust the release rate of the therapeutic within the
polymer matrix.
[0025] Alternatively, the biocompatible hydrogel particles may be
formed of multiple polymer molecules, copolymeric hydrogels, chosen
to grant specific and advantageous characteristics to the system.
In one preferred embodiment, hydrogels are constructed from a block
copolymer configuration of repeating units of poly-lactic acid
(PLA) and polyethylene glycol. The PLA confers rapidly hydrolyzable
ester bonds, whereas the PEG backbone prevents both the rapid
degradation of the polymer and adsorption of proteins to the
hydrogel surface and subsequent removal by the immune system.
[0026] The swellable particles can be made using a variety of
particle-forming processes and in a variety of particle shapes. For
example, swelling polymeric particles can be prepared using spray
drying, solvent evaporation, polymer micronization, and other
methods well known to those of ordinary skill in the art. Swellable
particles comprising hydrogels can be synthesized during
emulsification with an aqueous phase with a non-aqueous phase to
form microspheres of hydrogel that can be modulated in size by
changing parameters of the emulsion (e.g. non-aqueous phase
composition, concentration of reactants, mixing speed and shear
introduced into the emulsion, presence of surfactants and
surfactant types, etc). Swellable particles comprising hydrogels
also can be synthesized during spraying so that the hydrogel
particles form and are cross linked while dispersed as a droplet.
Alternately, disks, spheres, cubes, irregular shapes, thin films of
hydrogels can be synthesized and then broken into (comminuted)
small respirable, swellable particles using milling and
micronization methods. The comminuted particles can have a flake,
acicular or other non-spherical shape.
[0027] For example, the dried hydrogels made can be broken down
into macroparticles using a rotary blade mill (M 20 Universal mill,
IKA.RTM. Werke GmbH, Germany). Size reduction to narrowly dispersed
micron size particles suitable for inhalation can then be performed
using a fluid energy mill that uses impinging air jets to finely
micronize the material. Particle size of the resultant material is
controlled by the parameters of fluid energy milling (air
pressures) and can also be separated using an air classifier.
Particle size can be determined using a Sympatec Helos laser
diffraction instrument and also conventional cascade impaction
techniques. Milled particles exhibit much faster rates of swelling
than unmilled particles as a result of the increased surface area
available for water uptake.
[0028] The swellable particles may be fabricated or separated, for
example by sieving or air separation methods, to provide a particle
sample with a preselected size distribution to provide the desired
MMAD in the dry powder state. As mentioned above, the selected
range within which a certain percentage of the particles must fall
preferably is controlled such that at least 90%, or even more
preferably 95% or 99%, have an aerodynamic particle diameter
between 0.5 .mu.m and 5 .mu.m.
[0029] A particular process for making swellable particle starts
with hydrated films or hydrogel masses, which can then be processed
for particle size reduction and micronization. For example, the
hydrated films or hydrogel masses can be extruded through a fine
orifice or mesh to reduce particle size. The particles can then be
dehydrated. Alternatively, particles can be produced after
dehydration of the hydrogel films. Dehydration is achieved using
methods such as a vacuum oven drying, lyophilization, solvent
displacement by volatiles, among others. The dried hydrogels can be
broken down into particles using a rotary blade mill (M 20
Universal mill, IKA.RTM. Werke GmbH, Germany). Size reduction to
narrowly dispersed micron size particles suitable for inhalation
can then be performed using a fluid energy mill that uses impinging
air jets to finely micronize the material.
[0030] The swellable particles can be used for enhanced delivery of
one or more working agents to the airways of the respiratory tract,
including the alveolar region of the lung. The particles
incorporating one or multiple working agents may be effectively
aerosolized for administration to the respiratory tract to permit
systemic or local delivery of a wide variety of therapeutic agents.
For example, particle size of an inhaled aerosol is the primary
determinant of the deposition pattern within the airways. They
optionally may be co-delivered with larger carrier particles (i.e.
not carrying a working agent) which have for example a mean
diameter ranging between about 50 .mu.m and 150 .mu.m.
[0031] Incorporation of a therapeutic or other working agent within
the particle can be accomplished using a variety of methods. For
example, inclusion/incorporation of a working agent inside the
particles comprising hydrogels described herein can be achieved by
entrapping the working agent in the hydrogel polymer network so as
to control the release of the drug (or other working agent) from
the particle at specifically desired rates. Entrapment can be
achieved by performing the cross linking of the polymers in the
presence of the working agent (e.g. drug) such that the working
agent (e.g. drug) is entrapped within the polymer network that
forms the hydrogel network. Entrapment also can be achieved by
performing the cross linking of the polymers prior to placing the
hydrogels in the presence of the working agent (e.g. drug) such
that the working agent is entrapped within the polymer network by
diffusing into the hydrogel. Moreover, multiple working agents
(e.g. drugs) can be loaded into the same swellable particles using
the same or different methods since beneficial release rates may be
achieved by loading differently or using the same methods.
[0032] The therapeutic agent (or other working agent) to be loaded
into the swellable hydrogel particles can take various forms
including a drug in solution such that the drug is a molecular
dispersion throughout the hydrogel particle, a drug in suspension,
a colloidal dispersion, a nanoparticle system dispersed throughout
the hydrogel particle, or a drug in liposomes, lipid dispersions,
nanocapsules, polymeric nanoparticles, etc dispersed throughout the
hydrogel particle.
[0033] For purposes of further illustration and not limitation,
incorporation of a therapeutic agent (or other working agent)
within the particle can be accomplished by the following
methods:
[0034] a) Encapsulation of the therapeutic agent within a
nanoparticle and placement of the nanoparticle within the particle.
For example, Ibuprofen can be encapsulated within Lecithin
(phophotidylcholine) nanoparticles 200-400 nm in diameter. These
nanoparticles were prepared using the method of Chen et al, 2002
(Chen, X., Young, T. J., Sarkari, M., Williams, R. O., Johnston, K.
P., Preparation of cyclosporine A nanoparticles by evaporative
precipitation into aqueous solution, International Journal of
Pharmaceutics 242, (2002) 3-14). The nanoparticles are incorporated
into the swellable particles by physical entrapment in the hydrogel
network by performing crosslinking reactions in the presence of the
nanoparticles in the reactant solution of the 20 kDa 8-arm
acrylated PEG with dithiothreitol as described below for the
illustrative embodiment. Nanoparticles can be designed to have
differential rates of release from the swellable particles based on
their relative sizes, hydrophilic nature, ionic properties, and
diffusion coefficients.
[0035] b) Direct encapsulation of the molecule within the matrix of
the polymer network (e.g. PEG hydrogel network) of the swellable
particle. For example, Rhodamine therapeutic agent can be trapped
within the polymer network (PEG hydrogel network) during the
crosslinking of the 20 kDa 8-arm acrylated PEG with dithiothreitol
as described below for the illustrative embodiment.
[0036] c) Attachment of the therapeutic agent to the polymer
network itself (PEG hydrogel network) through chemical interactions
(covalent, ionic, and hydrogen bonds). For example,
N-acetylcysteine mucolytic agent can be attached to the polymer
network of 8-arm acrylated PEG crosslinked with dithiothreitol
through covalent bonds between the thiol group on the
N-acetylcysteine and one thiol group on the dithiothreitol with the
other thiol group of dithiothreitol bonded to an acrylate group of
the polymer network. The N-acetylcysteine mucolytic agent can be
reversibly covalently bound to the hydrogel network of the particle
such that the N-acetylcysteine functions to decrease the viscosity
of the mucus by disrupting the disulfide bonds formed between
adjacent cysteine residues. This disruption of mucus disulfide
bonds is readily achieved since the disulfide bonds are transferred
between cysteine residues. This facilitates prolonged and localized
mucolytic release around the hydrogel particle structure,
increasing transport rates through the CF mucus environment.
[0037] Using these encapsulation methods, numerous therapeutic and
other working agents, ranging from small and hydrophilic to large
and hydrophobic, may be incorporated into the swellable particles
for the aerosolized treatment of cystic fibrosis, lung cancer,
asthma, chronic obstructive pulmonary disease (COPD), acute
bronchitis, emphysema, tuberculosis, or systemic diseases. For
example, loading of Rhodamine therapeutic agent in PEG hydrogel
particles described herein has been achieved during polymerization
or after the particles were made. High loading of this drug was
obtained using both methods. For example, approximately 35% w/w
Rhodamine therapeutic agent was present after washing surface drug
from the hydrogel particles.
[0038] Any of a variety of therapeutic treating agents,
prophylactic agents, diagnostic agents, imaging agents such as
radio-isotopes, or other active working agents can be incorporated
within the swellable particles. The swellable particles can be used
to locally or systemically deliver a variety of therapeutic agents
to the respiratory tract. Examples of working agents include
synthetic inorganic and organic compounds, proteins and peptides,
polysaccharides and other sugars, lipids, and nucleic acid
sequences having therapeutic, prophylactic, diagnostic or imaging
activities. Nucleic acid sequences include genes, antisense
molecules which bind to complementary DNA to inhibit transcription,
and ribozymes. The working 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 can be
encapsulated, for example, between 100 and 500,000 .mu.m per
mole.
[0039] Proteins are defined as comprising 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.
[0040] The swellable polymeric aerosols are useful as carriers for
a variety of inhalation therapies. They can be used to encapsulate
small and large drugs, release encapsulated 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 therapeutic.
[0041] For example, the swellable particles may include a
therapeutic agent for local delivery within the lung, such as
agents for the treatment of asthma, emphysema, or cystic fibrosis,
or for systemic treatment. 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, but
are not limited to, insulin, calcitonin, leuprolide (or LHRH),
G-CSF, parathyroid hormone-related peptide, somatostatin,
testosterone, progesterone, estradiol, nicotine, fentanyl,
norethisterone, clonidine, scopolamine, salicylate, cromolyn
sodium, salmeterol, formeterol, albuterol, and vallium.
[0042] The particles including a therapeutic agent may be
administered alone or in any appropriate pharmaceutical carrier,
such as an inert sugar particle system typically used in a powder
inhaler, for administration to the respiratory system. They can be
co-delivered with larger carrier particles (not including a
therapeutic agent) possessing mass mean diameters for example in
the range 50 .mu.m to 150 .mu.m.
[0043] 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 Drug 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, Esevier,
Amsterdam, 1985, the disclosures of which are incorporated herein
by reference.
[0044] The relatively large size of swollen aerosol particles
deposited in the deep lungs minimizes potential drug losses caused
by particle phagocytosis. The swellable polymeric matrix also
facilitates as a therapeutic carrier to provide the benefits of
biodegradable polymers for controlled release in the lungs and
long-time local action or systemic bioavailability. Denaturation of
macromolecular drugs can be minimized during aerosolization since
macromolecules are contained and protected within a polymeric
matrix shell. Coencapsulation of peptides with peptidase-inhibitors
can minimize peptide enzymatic degradation.
[0045] For purposes of still further illustration and not
limitation, the swellable particles can include a working agent
that comprises a mucolytic agent alone or together with another
working agent such as antibiotic agent, a cytotoxic agent, other
mucolytic agents, an RNA interfering agent which includes siRNA and
miRNA, a gene, or combinations thereof. The working agent also can
comprise multiple cytotoxic agents, a cytotoxic agent and an RNA
interfering agent, or other combinations of working agents for
purposes of further illustration.
[0046] In comparison to non-swellable particles, the swellable
particles pursuant to the present invention 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. Phagocytosis of
particles by alveolar macrophages diminishes precipitously as
particle diameter increases beyond 3 .mu.m 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). 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.
[0047] Swellable particles thus are capable of a longer term
release of a therapeutic or other working agent. Following
inhalation, swellable biodegradable particles can deposit in the
lungs (due to their relatively small size), and subsequently
undergo swelling, slow degradation and drug release, without the
majority of the particles being phagocytosed by alveolar
macrophages. A drug 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 drug. The swellable particles
thus are highly suitable for inhalation therapies, particularly in
controlled release applications. The preferred mass median
aerodynamic diameter for swellable particles for inhalation therapy
is between 0.5 to 5 .mu.m (prior to swelling). After swelling, the
particles have geometric sizes of greater than 6 .mu.m volume mean
diameter in the airways.
[0048] The particles may be fabricated with the appropriate
material, surface roughness, diameter, density, and swelling
properties for localized delivery to selected regions of the
respiratory tract such as the deep lung or upper airways. For
example, larger particles or more dense particles may be used for
upper airway delivery, or a mixture of different sized particles in
a sample, provided with the same or different therapeutic agent may
be administered to target different regions of the lung in one
administration.
[0049] The swellable particles can be delivered by inhalation
methods using propellant driven metered dose inhalers wherein
hydrofluoroalkane and/or alkane liquefied gas propellants are used
in these formulations with other excipients included for
stabilization of the preparations. Dry powder inhalers can use
swellable hydrogel particles prepared for aerosolization and
inhalation. Use of a dry powder inhaler may require blending with
so called "carrier" particles (See Smyth and Hickey, Carriers in
Drug Powder Delivery: Implications for Inhalation System Design,
American Journal of Drug Delivery, Volume 3, Number 2, 2005, pp.
117-132). These carrier particles are typically lactose, sucrose,
glucose or other particles that are blended with the swellable
particles for aerosolization. Typically, the carrier particles are
sized between 50-500 .mu.m as determined by sieve analysis and make
up from 90-99% w/w of the powder placed in the inhaler for aerosol
dispersion and inhalation. Dry powder insufflation, liquid spray
systems, and nebulizers also can be used to deliver the swellable
particles.
[0050] Moreover, modulation of the aerodynamic particle size of the
aerosol particles can be used to target different regions of the
airways. Attachment of targeting ligands on the surface of the
swellable particles can result in their localization at specific
sites within the respiratory systems, such as for lung cancer a
targeting ligand may be used to bind to a receptor that is overly
expressed in that lung cancer such as a folate receptor. Targeting
also can be achieved by causing the hydrogel particle to change
chemical bonding or conformation when the particle is in a
microenvironment that is unique to the disease site, such as in
infection in the lung where inflammatory response of the lungs to
the microorganisms causes higher concentrations of chemicals and
mediators that can cause the hydrogel to actively change its
nature. This could be to cause pH sensitive changes in the hydrogel
network so that the drug loaded in the hydrogel particle is rapidly
released when pH decreases so as to concentrate the drug release to
areas where the disease is most pronounced.
Example 1
Synthesis of Peg Hydrogel
[0051] An eight-arm, hydroxyterminated PEGs with total number
average molecular weights (Mw) of approximately 10 and 20 kD are
acrylated to a degree of functionalization exceeding 95% after
azeotropic distillation followed by reaction with acryloyl chloride
in the presence of triethylamine as has previously been described
by Elbert, D. L., et al., Self-selective Reactions in the Design of
Materials for Controlled Delivery of proteins. Journal of
Controlled Release, 2001. 76: p. 11-25. Hydrogels are formed by
mixing the chosen PEG-acrylate with either dithiothreitol or
PEG-dithiol, Mw 3.1 kD at a 1:1 stoichiometric ratio of acrylates
to thiols in 50 mM phosphate buffered solution (PBS, pH 7.8). Each
reactant is dissolved separately in an aliquot of PBS. The amount
of PBS is varied to give the desired total precursor concentration
(wt %) upon mixing. The two solutions are mixed vigorously in 1.5
mL plastic tubes and centrifuged to remove bubbles. The sealed tube
containing the mixture is placed at 37.degree. C., and allowed to
react overnight to ensure complete conversion.
[0052] There are significant toxicological and compatibility
advantages of using such a reaction to form gels under
physiological conditions for drug delivery applications Peppas, N.
A., et al., Physicochemical foundations and structural design of
hydrogels in medicine and biology. Annu Rev. Biomed. Eng., 2000. 2:
p. 9-29. The degradation rate of the polymer is determined by
various factors including the initial water content of the hydrogel
network. This is initial water content is controlled by the
crosslinker and the molecular weight of the PEG acrylate. Drug
release from the hydrogel is determined by the relative sizes of
the drug molecule and the mesh size of the crosslinked network. If
drug size is assumed constant (though drug suspension particles
could conceivably be modulated in some cases), drug release can be
modulated by decreasing the mesh size. This is achieved by
decreasing the length of the polymers (molecular weight).
[0053] Poly(acrylic acid-co-acrylamide) hydrogels were synthesized
using similar methods to those described by Chen, J. et al.
"Synthesis and characterization of superporous hydrogel
composites", Journal of Controlled Release 65: pp. 73-82 (1999),
the teachings of which are incorporated herein by reference.
[0054] Alternatively, the swelling particles for pulmonary drug
delivery 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-colysine-graft-alanine-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.
[0055] In the synthesis, the graft copolymers may be tailored to
optimize different characteristics of the swelling 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.
Example 2
Dual Action Mucolytic-Therapeutic Drug Delivery Vector for Cystic
Fibrosis
[0056] From the moment an aerosolized drug is expelled from the
metered-dose inhaler or nebulizer and enters the mouth, through its
journey past the pharynx, down the trachea and bronchioles into the
deeper recesses of the airways toward its site of action, and
finally to its degradation and removal, aerosolized agents are
under the influence of a multitude of factors, which may be grouped
into two general categories. Those determinants which govern the
deposition of the aerosolized agent onto the airway lumen surface
are termed physical properties, and include a particle's diameter
and density. The properties that determine the fate of the drug
subsequent to its impaction on the luminal surface, including its
absorption, metabolism and excretion, are referred to
pharmacokinetic factors.
[0057] Accordingly, the physical and pharmacokinetic factors of an
aerosolized particle must be precisely tailored to complement one
another as a means of delivering the most effective dosage possible
while simultaneously minimizing drug waste and circumventing
undesired collateral reactions.
[0058] With these considerations in mind, this EXAMPLE pursuant to
another illustrative embodiment provide a novel dual action
mucolytic-therapeutic hydrogel drug delivery vector for the
treatment of cystic fibrosis as a means of significantly improving
the efficacy of current FDA approved cystic fibrosis therapeutics.
Furthermore, although this delivery system was initially designed
specifically for cystic fibrosis, it can also serve as a
therapeutic delivery vector for other pulmonary disorders,
including lung cancer, COPD, and asthma.
[0059] Due to the large amounts of pathogens and debris that we
inhale with each breath, the lung possesses multiple lines of
defense to prevent infection and maintain homeostasis. From the
trachea to the terminal bronchioles, an area collectively referred
to as the central airways, the surface of the luminal epithelium is
coated with a film of fluid that is composed of a sol and gel phase
and referred to respectively as the periciliary and mucus layers.
Each of these two layers play an important role in keeping the lung
clear of pathogens, and their precise composition is essential to
the effective clearance of foreign particles from the airways. The
overlying thick and viscous mucus acts as a barrier to prevent the
passage of inhaled pathogens and other foreign debris to the
underlying epithelial cells below. The periciliary fluid, while not
a direct obstacle in the manner of the superjacent mucus layer, is
no less important to ensuring the lung is kept clean of pathogens.
Through the rhythmic and concerted beating of the epithelial cilia
within the periciliary fluid, mucus is propelled in the cephalic
direction towards the pharynx and removed from the airway via
expectoration or ingestion into the gastrointestinal tract, a
process referred to as the mucociliary escalator. Therefore, any
aerosolized agent remaining trapped within the mucus layer is
carried along and summarily removed from the respiratory tract.
[0060] The periciliary fluid is maintained at an optimal volume
which ensures that only the tips of the cilia contact the overlying
mucus, propelling it onward. The volume and ionic composition of
the periciliary fluid layer is the end result of the delicate
balance between absorption and secretion of H.sub.20 and ions
(specifically Na.sup.+ and C.sup.I-). This balance is tightly
regulated through the concerted action of ion channels located in
the apical surface of pulmonary epithelial cells. Absorption of
fluid is controlled primarily by the activity of the
amiloride-sensitive epithelial Na channel (ENAC). The absorption of
osmotically active Na.sup.+ ions from the periciliary fluid into
the epithelial cells drives water from apical to the basolateral
surface of the cell, decreasing the volume of the periciliary
fluid. This action is countered by both the outward rectifying
chloride channel (ORCC) and the cystic fibrosis transmembrane
conductance regulator (CFTR) channel, which allows the passage of
chloride ions out of the cell and into the lumen, carrying water
along and restoring the volume of the periciliary fluid.
[0061] In cystic fibrosis (CF), the most prevalent autosomal
recessive genetic disorder in Caucasians, the gene encoding the
CFTR channel contains a mutation which results in the translation
of a defective protein. The majority of the mutated CFTR protein is
degraded in the endoplasmic reticulum by the 26S proteosome, and
the small amount that does reach the plasma membrane of the apical
surface does not function properly and no longer allows the passage
of chloride ions into the pulmonary lumen. This decreased transport
of chloride ions results in a reduced amount of water entering the
periciliary fluid, and since the Na.sup.+ channels are still
absorbing water at a normal rate, the volume of the periciliary
fluid is considerably reduced. The upper region of the cilia become
embedded in the mucous layer, significantly hindering their
movement and leading to the cessation of the mucociliary escalator.
The mucus layer becomes increasingly viscous and stagnant, allowing
bacteria colonies to rapidly accumulate throughout the lung. This
increased infestation produces a vigorous response from the immune
cells of the body, which over time significantly deteriorates the
lung, and eventually results in death.
[0062] The ideal solution would be to administer gene therapy that
would replace the defective copy in the chromosome with one that
encodes for a functioning CFTR channel. Unfortunately, while this
is a very active area of research, due to numerous setbacks gene
therapy is still not a viable option, and treatment of the physical
manifestations of the genetic disorder remains the only route to
treat CF. The current therapy consists primarily of the mucolytic
N-acetylcysteine, which severs the bonds between adjacent mucin
glycoproteins and reduces the viscosity of the mucus, and
administration of various antibiotics targeting the numerous
bacterial colonies infesting the lungs of CF patients. These
treatments are administered in aerosolized form and may be
prescribed either separate or in tandem. While this therapy does
help alleviate some of the complications of CF, it is an extremely
wasteful procedure and much can be done to improve its efficacy
while simultaneously reducing its cost in both time and money. For
instance, the antibiotic will only be effective against the
bacteria that it encounters as soon as it is deposited in the
airways. Once the therapeutic contacts the viscous mucus it will be
become entrapped, completely abrogating its bactericidal activity.
Conversely, no matter where the mucolytic lands it will be able to
decrease the viscosity of the stagnant mucus in its vicinity.
However, the mucolytic alone does nothing to combat the numerous
bacterial colonies populating the airways, and which are the direct
source of the complications of CF.
[0063] This EXAMPLE provides swellable particles for delivering
both the mucolytic and therapeutic simultaneously within the same
aerosol dose. In this manner the mucolytic will reduce the
viscosity of the mucus and increase the radius of diffusion of the
therapeutic, bringing it into contact with a greater number of
bacteria and significantly improving its effectiveness. The use of
swellable particles to this end takes into consideration that any
therapy administered this way, where the therapeutic and mucolytic
are delivered in their free (i.e. unencapsulated) form, will be
short-lived. The immediate impact of the therapy will permit the
cilia to beat with increased frequency and improve the function of
the mucociliary escalator, which in turn will remove much of the
mucus and bacteria, accompanied by the therapeutics, from the lung.
However, as this treatment does nothing to address the underlying
genetic defect, the mucus will once again become stagnant and
viscous, allowing the remaining bacteria an opportunity to divide
and multiply, so that soon after administration the condition in
the lung is returned to its pre-therapy state. On average, cystic
fibrosis patients must spend approximately three hours per day
self-medicating. Not only does this strict and inflexible routine
have a tremendous impact on their quality of life, it also gives
rise to patient non-compliance with their prescribed therapy. Due
to the rigorous nature of their treatment, a patient may either
accidentally or intentionally forgo a day or two of therapy,
believing that such a brief absence of medication will not be
severely detrimental to their health. However, when fighting
against bacteria which require only hours to replicate this
negligence may yield severe consequences. How then does one solve
this paradox of alleviating the stringency of the therapy as a
means of improving patient compliance while simultaneously ensuring
the delivery of the prescribed dose of medication? The solution to
this question is found in sustained release therapy, where the
therapeutics are released continuously over a period of days, if
not weeks, allowing the patient to receive their prescribed dosage
while increasing the interval between administration.
[0064] The use of swellable particles to this end takes into
consideration the problem that delivery via aerosolized particles
completely ignores the aerodynamics which governs the deposition of
aerosolized particles in the lung. In aerosols a very important
property is the aerodynamic diameter (d.sub.a) of a particle. The
aerodynamic diameter is a method used to standardize the
aerodynamical properties of particles regardless of shape, density,
size, or actual diameter, and may be thought of as the diameter of
a water droplet having the same aerodynamic properties as described
above. The most significant information derived from the
aerodynamic diameter is that a particle with a given aerodynamic
diameter will be aerodynamically indistinguishable from other
particles of different size, shape, diameter or density having the
same aerodynamic diameter.
[0065] As the preceding paragraph demonstrates, the aerodynamic
properties of an aerosolized particle are heavily dependant upon
both its' diameter and density, and the slightest alteration of
either property will result in the deposition of the particle in a
different region of the lung. Therefore, attempting to deliver a
mixture of mucolytic and therapeutic either in their free form or
encapsulated within discreet delivery vectors will result in
non-uniform deposition, since the two do not have identical
diameters and densities and therefore will not possess identical
aerodynamic properties. Consequently, the EXAMPLE provides for
delivery of both the mucolytic agent and therapeutic agent together
using the swellable particles.
[0066] Moreover, the swellable particles provide a preferred CF
therapeutic vector to provide sustained release of the therapeutic
agent and also capability to carry both mucolytic and therapeutic
agents. Furthermore, when designing a delivery vector for human
consumption, it is essential that it is biocompatible, non-toxic,
and non-immunogenic. The hydrogel particles described above
encapsulating the therapeutic and mucolytic agents within a
crosslinked, hydrophilic polymer network achieves the sustained
release of the agents. For example, when a hydrogel is placed in an
aqueous environment it readily imbibes water and swells, stretching
its polymer chains and producing numerous pores that will permit
the drugs within to diffuse out. Hydrogels are especially relevant
for drug delivery in the lung, as the initial surface that a
delivery vector encounters is the viscous mucus, which is primarily
composed of hydrophilic mucin glycoproteins, which readily attract
water. Therefore, hydrophilic polymer hydrogels are ideally suited
to siphon water away from the mucus, allowing the hydrogel to swell
and release its therapeutic cargo.
[0067] The most important factor when developing a hydrogel network
is the selection of a suitable hydrophilic polymer, and due its
numerous beneficial properties, polyethylene glycol (PEG) is
described above. Among the attributes that make this FDA approved
polymer attractive include that it is non-toxic, biocompatible,
non-immunogenic and strongly hydrophilic, allowing it to draw water
away from the mucus to promote hydrogel swelling. Furthermore, PEG
exhibits almost no protein adsorption, allowing it to elude the
ubiquitous macrophages patrolling the labyrinth of the pulmonary
passages. These characteristics combine to make PEG the most
promising candidate for employment in pulmonary drug delivery.
[0068] The antibiotics used to treat CF include tobramycin and
gentamycin, which are large, bulky, hydrophobic macromolecules.
While their hydrophobic nature prevents the antibiotics from being
incorporated in their free form, it permits them to be readily
encapsulated within the core of a liposome nanoparticle. The
liposome nanoparticle in turn possesses a hydrophilic surface,
allowing it be easily incorporated inside the hydrogel matrix.
Furthermore, by adjusting the size of both the liposome
nanoparticle and the pores of the polymer matrix, the release rate
of the antibiotic from the hydrogel may be accurately
controlled.
[0069] The next step is to encapsulate the mucolytic (e.g.
N-acetylcysteine designated NACS) within the hydrogel. Due to the
small size of NACS and its hydrophilic nature, if it is simply
loaded into the hydrogel in its free state, the moment the hydrogel
begins to swell NACS will be released from the lung in an
undesirable single, rapid burst. To overcome this, the NACS is
covalently bonded to the PEG polymer network, while still retaining
its mucolytic activity. The advantage of this method is that as the
hydrolytically labile ester bonds are broken NACS will be released
into the environment. This ensures that mucolytic will be released
throughout the entire lifetime of the hydrogel, providing the
sought after sustained release to accompany the diffusion of the
antibiotic from the gel.
[0070] As previously mentioned, NACS is extremely soluble in water
(100 mg/mL), thus precluding its incorporation into nanoparticles
similar to those encapsulating the hydrophobic antibiotic.
##STR00001##
[0071] The structure of N-acetylcysteine (pictured above) affords
us two functional groups with which to create a hydrolytically
labile bond to attach the mucolytic agent to the PEG-polymer
network. The release rate of the therapeutic will depend on both
the kinetics of the cleavage/degradation of the drug-network
linkage (which is described by an appropriate rate constant), and
upon the diffusion rate of the free molecule from the matrix of the
polymer network. The crosslinker employed in our hydrogels is
dithiothreitol, which contains two thiol moieties that readily
react with our functionalized PEG octa-acrylates without the need
of organic solvents at body temperature and biological pH.
##STR00002##
[0072] However, since NACS contains only one thiol, it could either
form a disulfide bond with another NACS molecule, or it may react
with the acrylate group of the polymer. Either way, its function as
a mucolytic agent would be significantly disrupted, if not
abrogated entirely.
##STR00003##
[0073] As shown in the reaction scheme (which is only showing the
acrylate end group of the 8-arm PEG-acrylate coupling with the
thiol functionalized end of DTT), the initial step is the Michael
addition of the thiol to the acrylate group. Following this
reaction, the next step is an ester hydrolysis in which an alcohol
and a carboxylic acid are formed (the exact opposite of the Fisher
esterification reaction). But as can be seen, the sulfur atoms
(since what is formed is actually a dicarboxylic acid, since
dithiothreitol contains two S--H groups per molecule) are no longer
in their reduced thiol forms. Instead, as shown below, they are
bonded to two carbon atoms, forming C--S--C linkages, and no longer
able to participate in disulfide bonding with the thiol moieties of
the mucin polymers.
##STR00004##
[0074] Therefore, control of the release of NACS can be achieved by
reacting dithiothreitol with N-acetylcysteine, in perhaps a molar
ratio of at least 4:1 (DTT:NACS), prior to the crosslinking
reaction with the 8-arm PEG polymer. This will allow the thiols of
DTT to react with the thiol of NACS. The advantage of NACS is that
it is already an FDA approved drug, and similar to other thiol
compounds, will react readily without the need for harmful solvents
or reaction conditions. The structure of the cleaved molecule
(following ester hydrolysis) is shown below:
##STR00005##
[0075] As opposed to the previous compound, this molecule contains
a disulfide S--S linkage, which readily competes with mucin
polymers for their disulfide bonds, since the only exchange that is
occurring is one disulfide bond for another (requiring no
significant cost in energy, unlike an exchange from a low-energy
bond to a high-energy bond, which would be thermodynamically
unfavorable and would possess a large transition barrier) this
exchange of disulfide bonds occurs under very mild conditions
(essentially mixing the chemicals in a PBS buffer).
[0076] Although there will certainly be some NACS-NACS disulfide
bonding, this is not an irreversible linkage, and a greater amount
of DTT will ensure that the majority of the NACS will be bound to
the crosslinker (theoretically, every molecule of NACS will yield
one S--S active mucolytic bond). By increasing the amount of DTT
used we can ensure that the hydrogel has the majority of the
crosslinkable moieties occupied, thus maintaining structural
stability, while simultaneously possessing enough DTT-NACS to
exhibit a significant enhancement in hydrogel permeability through
the mucus. And while it is true that any DTT that is bound to NACS
will not be able to form a crosslink with another PEG molecule,
recent experiments show that firm hydrogels can be formed with at
least 30% concentration of crosslinker compared to the polymer
(that is to say that there are enough DTT molecules to
theoretically occupy 30% of the PEG acrylate groups).
Synthesis of Biocompatible Hydrogels from 8-Arm PEG Acrylate
Containing Rhodamine and N-Acetylcysteine
[0077] The initial synthesis of the acrylated 8-arm PEG (Mw=10 kDa
and 20 kDa) was based upon the work of Hubbell et al. Journal of
Controlled Release 76:11-25 (2001) as described above. For example,
ten (10) g of 8-arm PEG (20 kDa, Nektar) was dissolved in 200 mL of
toluene and distilled azeotropically for 2 hrs. The resulting
solution was then allowed to cool to 50 C under argon. Two (2) mL
of triethylamine was added to 50 mL of dichloromethane which was
then added to the reaction solution. An amount (1.3 mL) of acryloyl
chloride was then added dropwise and the reaction proceeded under
argon in the dark for 20 hrs. The resulting opaque pale yellow
solution was then filtered multiple times until clear. Anhydrous
sodium carbonate was added to the solution and stirred for two
hours to remove any water that was present. The solution was
filtered to remove the sodium carbonate and was then evaporated
under reduced pressure. Diethylether was then added to the solution
and the reaction flask was placed in an ice bath to allow the
product (acrylated PEG) to precipitate. The product was collected
by filtration and repeatedly washed with diethylether. The average
yield was around 85%.
[0078] Firm and stable PEG hydrogels can be formed with an
acrylate:thiol stoichiometric ratio >1 and with a thiol amount
as low as 60% (0.60 ratio of thiol/acrylate) of the amount of
acrylate moieties present. Generally, dithiothreitol, having two
thiols per molecule, forms crosslinks between polymer molecules to
form the hydrogel network. To form hydrogels with N-acetylcysteine
mucolytic agent covalently bound to the hydrogel network, we
reacted 5 mg of dithiothreitol with 1.1 mg of N-acetylycysteine,
each dissolved in 20 ul of 1.times.PBS (7.4 pH). N-acetylcysteine
has the capacity to form a bond between it's thiol group and either
the acrylated polymer or to dithiothreitol via S--S disulfide bonds
formed from two thiol groups. Stiochiometrically, at the above
mentioned concentrations of reactants, if all of the thiol groups
on the N-acetylcysteine are each bound to one unique dithiothreitol
molecule, there will still remain enough free thiol groups on
dithiothreitol molecules to form sufficient crosslinking between
the polymermolecules such that stable hydrogels are formed. This
solution was added to 0.160 g of acrylated 8-arm PEG (20 kDa)
dissolved in 200 .mu.l of 1.times.PBS (pH 7.4). To this solution
was added 10 .mu.l of Rhodamine therapeutic agent (40 mg/mL
1.times.PBS (pH 7.4)). Seventy (70) .mu.l aliquots of the solution
was placed between microscope slides coated with SigmaCote (Sigma
Chemical Co.), and separated by 1 mm spacers. The gels were allowed
to cure for 24 hours in a humid environment at 37 C. The cured gels
were milled after drying, in a micro-ball mill (from Dentsply Rinn,
Elgin, Ill.) cooled using liquid nitrogen, to produce swellable
particles having volume mean diameters of between 1.1 and 3 .mu.m
and a span of 2.2 (span=(D90-D10)/D50) where D50 is median diameter
and D10 and D90 are respective 10.sup.th and 90.sup.th percentile
diameters (e.g. for D10, 10% of particles are less than this
diameter).
[0079] As discussed above, in this EXAMPLE, dual delivery of both
mucolytic and therapeutic agents would serve to significantly
enhance the effectiveness of the present CF therapy by increasing
the radius of diffusion of the therapeutic, allowing it to contact
a larger number of bacteria, while simultaneously improving the
function of the mucociliary escalator. However, this dual-action
aerosolized hydrogel is not limited for use to cystic fibrosis
therapy. As previously mentioned, during physiological conditions
the lumen of the pulmonary passages are coated with a layer of
mucus serving as a barrier to inhaled particles. This mucus is
continuously removed via the mucociliary escalator, requiring
approximately 10 hours to eject inhaled debris from the furthest
reaches of the lung, and much less time to clear the upper passages
where the majority of malignant tumors dwell. Incapable of
distinguishing between therapeutic aerosols and pathogens, the
mucociliary escalator expels friend and foe alike with equal vigor.
Accordingly, any inhaled therapeutic has only a brief window in
which to penetrate the mucous barrier, attain the underlying
epithelium, and deliver its medicinal cargo, else it forever loses
any opportunity for efficacy. Furthermore, when one also considers
that the tumor is not uniformly distributed throughout the lung and
that there is only a specific region where the therapeutic will be
effective, the aforementioned brief window of action is further
narrowed. By controlling the aerodynamic properties of a particle
via its diameter and density, an aerosol can be tailored so that
the majority of the particles will arrive at the desired location
in the lung. However, once the particle lands on the surface of the
lumen, the onus is entirely upon the particle to penetrate the
mucus prior to its expulsion from the lung.
[0080] Therefore, it becomes evident that the quicker a particle
can pass through the viscous mucus, the greater its chances to
provide a beneficial effect. By combining a mucolytic with a
cytotoxic agent, corticosteroid or bronchodilator for the treatment
of lung cancer, COPD, and asthma respectively, the efficacy of the
treatments will be increased and the amount of drug wasted via the
action of the mucociliary escalator will be markedly reduced.
[0081] Targeting molecules can be attached to the swelling
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
interaction of the particle with specific receptor sites, such as
those within the lungs. The particles can be targeted by attachment
of 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 for example to receptors on the surfaces of the
target cells.
[0082] Although the invention has been described above with respect
to certain illustrative embodiments, those skilled in the art will
appreciate that changes, modifications and the like can be made
thereto without departing from the spirit and scope of the
invention as defined in the appended claims.
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