Note: Descriptions are shown in the official language in which they were submitted.
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LOCAL DELIVERY OF MEDICATIONS TO THE HEART
Cross Reference To Related Applications
The priority of Provisional Application No. 60/098,975, filed September 2,
1998 is claimed and its content is incorporated by reference herein in its
entirety.
Summary of the Invention
This invention relates to the use of drug containing microspheres with
acoustical characteristics that allow the ultrasound release of drugs for
treatment of
the heart. In particular, the invention relates to specific microparticle
characteristics
in combination with specific types of cardiac drugs for ultrasound mediated
therapy.
Description of the Preferred Embodiments
As used herein the term microparticles is intended to include microcapsules,
15 microspheres and microbubbles which are hollow particles enclosing a core
filled
with a gas. It is not necessary for the microparticles to be precisely
spherical
although they generally will be spherical and described as having average
diameters.
If the microparticles are not spherical, then the diameters are referred to or
linked to
the diameter of a corresponding spherical microparticle having the same mass
and
enclosing approximately the same volume of interior space as a non-spherical
microparticle.
The microparticles are preferred to have an extended circulatory half life ~s
compared to bolus injections of free drug. In order to treat the heart for
several
minutes, the microparticles will necessarily need to recirculate back to the
heart. The
microparticles will need to avoid biological clearance and degradation in
order to
recirculate and enable drug treatment for several minutes to hours from a
single bolus
injection. Alternatively, the microparticles may be constantly administered by
an
infusion pump.
The types of agents to be released by the microparticles are typically
30 cardiovascular drugs with short circulatory half lives that affect the
cardiac tissues,
vasculature and endothelium to protect and treat the heart. Drugs which target
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2
platelets and white cells which may plug the microvasculature of the heart
after a
heart attack are also useful for local cardiac delivery. A third type of drug
useful for
local delivery is one for which a local effect is required but where the
systemic effects
of the drug would be detrimental. These are typically drugs with high
toxicity, for
example, locally administered potent vasodilators which would increase blood
flow to
hypoxic tissue, but if delivered systemically would cause a dangerous drop in
blood
pressure. Suitable drugs include fibrinolytic agents such as tissue
plasminogen
activator, streptokinase, urokinase, and their derivatives, vasodilators such
as
verapamil, multifunctional agents such as adenosine, adenosine agonists,
adenosine
monophosphate, adenosine diphosphate, adenosine triphosphate, and their
derivatives,
white cell or platelet acting agents such as GPllb/llla antagonists, energy
conserving
agents such as calcium channel blockers, magnesium and beta blockers,
endothelium
acting agents such as nitric oxide, nitric oxide donors, nitrates, and their
derivatives,
free-radical scavenging agents, agents which affect ventricular remodeling
such as
ACE inhibitors and angiogenic agents, agents that limit ischemic or
reperfusion injury
to the heart, and agents to limit restenosis of coronary arteries after
balloon
angioplasty or stenting.
In addition to therapeutic agents delivered locally to the heart, the use of
vasodilators in the microparticles will have enhanced diagnostic application.
20 Vasodilators are used in cardiology to assess the coronary blood flow
reserve by
comparing blood flow in heart with and without the maximal vasodilation by the
pharmacological agent. Coronary blood flow reserve correlates well with
patient
prognosis since the reserve capacity enables the myocardium to remain viable
during
a heart attack. Adenosine and other vasodilators are used during
interventional
25 cardiology and nuclear imaging to determine coronary reserve. A
microparticle
which contains gas and a vasodilator will be useful in echocardiography to
examine
the myocardium under normal conditions, and then trigger to release
vasodilator by
the ultrasound beam conditions to stimulate local vasodilation. The coronary
blood
flow reserve may be estimated non-invasively using ultrasound imaging by the
extent
30 of hyperemia of the myocardium, Doppler regional flow, or by other well
known
methods of characterizing the ultrasound imaging data.
It is advantageous to use the microparticles since the half life of the drug
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incorporated into the microparticle may be longer than the half life of the
drug in the
circulatory system if administered without the microparticles. This gives a
prolonged
effect of the drug compared to the effect when the drug is administered alone.
The microparticles according to the present invention have a bi-layered shell.
The outer layer of the shell will be a biologically compatible material or
biomaterial
since it defines the surface which will be exposed to the blood and tissues
within the
body. The inner layer of the shell will be a biodegradable polymer, which may
be a
synthetic polymer, which may be tailored to provide the desired mechanical and
acoustic properties to the shell or provide drug delivery properties. The
10 microparticles will contain gas, typically air or nitrogen, and for drug
delivery
purposes, a drug incorporated into the microparticle. To make the
microparticles
rupturable by a low intensity ultrasound energy, they must contain a gas to
allow
acoustic coupling and particle oscillation. Microparticles are constructed
herein such
that the majority of those prepared in a composition will have diameters
within the
15 range of about one to ten microns in order to pass through the capillary
system of the
body.
Since the microparticles have an outer and inner layer, the layers can be
tailored to serve different functions. The outer shell which is exposed to the
blood
and tissues serves as the biological interface between the microparticles and
the body.
20 Thus it will be made of a biocompatible material which is typically
amphiphilic, that
is, has both hydrophobic and hydrophilic characteristics. Blood compatible
materials
are particularly preferred. Such preferred materials are biological materials
including
proteins such as collagen, gelatin or serum albumins or globulins, either
derived from
humans or having a structure similar to the human protein, glycosoaminoglycans
such
25 as hyaluronic acid, heparin and chondroitin sulphate and combinations or
derivatives
thereof. Synthetic biodegradable polymers, such as polyethylene glycol,
polyethylene
oxide, polypropylene glycol and combinations or derivatives may also be used.
The
outer layer is typically amphiphilic, as well as having a chemistry which
allows
charge and chemical modification. The versatility of the surface allows for
such
30 modifications as altering the charge of the outer shell, such as by
selecting a type A
gelatin having an isoelectric point above physiological pH, or by using a type
B
gelatin having an isoelectric point below physiological pH. The outer surfaces
may
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also be chemically modified to enhance biocompatibility, such as by
PEGylation,
succinylation or amidation, as well as being chemically binding to the surface
targeting moiety for binding to selected tissues. T'he targeting moieties may
be
antibodies, cell receptors, lectins, selectins, integrins or chemical
structures or
5 analogues of the receptor targets of such materials. The mechanical
properties of the
outer layer may also be modified, such as by cross linking, to make the
microparticles
suitable for passage to the left ventricle, to provide a particular resonant
frequency for
a selected harmonic of the diagnostic imaging system, or to provide stability
to a
threshold diagnostic imaging level of the ultrasound radiation.
10 The inner shell will be a biodegradable polymer, which may be a synthetic
polymer. An advantage of the inner shell is that it provides additional
mechanical or
drug delivery properties to the microparticle which are not provided or
insufficiently
provided by the outer layer, or enhances mechanical properties not
sufficiently
provided by the outer layer, without being constrained by surface property
1 S requirements. For example, a biocompatible outer layer of a cross-linked
proteinaceous hydrogel can be physically supported using a high modulus
synthetic
polymer as the inner layer. The polymer may be selected for its modulus of
elasticity
and elongation, which define the desired mechanical properties. Typical
biodegradable polymers include polycaprolactone, polylactic acid, polylactic-
20 polyglycolic acid co-polymers, co-polymers of lactides and lactones, such
as epsilon-
caprolactone, delta-valerolactone, polyalkylcyanoacrylates, polyamides,
polyhydroxybutryrates, polydioxanones, poly-beta-aminoketones, polyanhydrides,
poly-(ortho)esters, polyamino acids, such as polyglutamic and polyaspartic
acids or
esters of polyglutamic and polyaspartic acids. References on many
biodegradable
25 polymers are cited in Lan eg r-et. al. (1983) Macromol.Chem.Phys.C23, 61-
125.
The inner layer permits the modification of the mechanical properties of the
shell of the microparticle which are not provided by the outer layer alone.
Moreover,
the inner layer may provide a drug carrier and/or drug delivery capacity which
is not
sufficient or providable by the outer layer alone. For use as an ultrasonic
contrast
30 agent, the inner layer will typically have thickness which is no larger
than is
necessary to meet the minimum mechanical or drug carrying/delivering
properties, in
order to maximize the interior gas volume of the microparticle. The greater
the gas
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volume within the microparticle the better the echogenic properties.
Since the multilayer nature of the microparticle allows the ultrasound
response
to be tunable based on the choice of materials and combinations of materials
in the
different layers, the microparticles may be made so that they are detectable
by
S ultrasound without rupturing. This allows a microparticle sample to be
imaged under
one set of conditions, then, by changing the ultrasound conditions so that a
threshold
is reached where they rupture, the drug can be selectively released when the
microparticles are in the vicinity of the heart.
The combined thickness of the outer and inner layers of the microparticle
shell
10 will depend in part on the mechanical and drug carrying/delivering
properties
required of the microparticle, but typically the total shell thickness will be
in the
range of 25 to 750 nm.
The microparticles may be prepared by an emulsification process to control
the sequential interfacial deposition of shell materials. Due to the
amphiphilicity of
the material forming the outer layer, stable oil/water emulsions may be
prepared
having an inner phase to outer phase ratio approaching 3:1, without phase
inversion,
which can be dispersable in water to form stable organic phase droplets
without the
need for surfactants, viscosity enhancers or high shear rates.
Two solutions are prepared, the first being an aqueous solution of the outer
biomaterial. The second is a solution of the polymer which is used to form the
inner
layer, in a relatively volatile water-immiscible liquid which is a solvent for
the
polymer, and a relatively non-volatile water-immiscible liquid which is a non-
solvent
for the polymer. The relatively volatile water-immiscible solvent is typically
a C5.-C7
ester, such as isopropyl acetate. The relatively non-volatile water-immiscible
non-
25 solvent is typically a C6-C20 hydrocarbon such as decane, undecane,
cyclohexane,
cyclooctane and the like. In the second solution containing the polymer for
the inner
layer, the polymer in water-immiscible solvents are combined so that the
polymer
fully dissolves and the two solvents are miscible with agitation. The polymer
solution
(organic phase) is slowly added to the biomaterial solution (aqueous phase) to
form a
30 liquid foam. Typically about three parts of the organic polymer solution
having a
concentration of about 0.5 to 10 percent of the polymer is added to one part
of the
aqueous biomaterial solution having a concentration of about 1 to 20 percent
of the
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6
biomaterial. The relative concentrations of the solutions and the ratio of
organic
phase to aqueous phase utilized in this step essentially determine the size of
the final
microparticle and wall thickness. After thorough mixing of the liquid foam, it
is
dispersed into water and typically warmed to about 30 - 35 °C with mild
agitation.
While not intending to be bound by a particular theory, it is believed that
the
biomaterial in the foam disperses into the warm water to stabilize an emulsion
of the
polymer in the organic phase encapsulated within a biomaterial envelope. To
render
the biomaterial envelope water insoluble, a cross linking agent, such as
glutaraldehyde, is added to the mixture to react with the biomaterial envelope
and
render it water insoluble, stabilizing the outer shell. Other cross-linking
agents may
be used, including the use of carbodiimide cross-linkers.
Since at this point the inner core contains a solution of a polymer, a solvent
and a non-solvent with different volatilities, as the more volatile solvent
evaporates,
or is diluted, the polymer precipitates in the presence of the less volatile
non-solvent.
This process forms a film of precipitate at the interface with the inner
surface of the
biomaterial shell, thus forming the inner shell of the microparticle after the
more
volatile solvent has been reduced in concentration either by dilution,
evaporation or
the like. The core of the microparticle then contains predominately the
organic non-
solvent. The microparticles may then be isolated by centrifugation, washed,
formulated in a buffer system, if desired, and dried. Typically, drying by
lyophilization removes not only the non-solvent liquid core but also the
residual water
to yield gas-filled hollow microparticles.
It may be desirable to further modify the surface of the microparticle, for ..
example, in order to passivate surfaces against macrophages or the
reticuloendothelial
system (RES) in the liver. This may be accomplished, for example by chemically
modifying the surface of the microparticle to be negatively charged since
negatively
charged particles appear to better evade recognition by macrophages and the
RES
than positively charged particles. Also, the hydrophilicity of the surface may
be
changed by attaching hydrophilic conjugates, such as polyethylene glycol
(PEGylation) or succinic acid (succinylation) to the surface, either alone or
in
conjunction with the charge modification.
The biomaterial surface may also be modified to provide targeting
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characteristics for the microparticle. The surface may be tagged by known
methods
with antibodies or biological receptors. The drug is released at the target
site, for
example, by increasing the ultrasonic energy to rupture the particles at the
appropriate
time and location.
S The microparticles may also be sized or processed after manufacture. This is
an advantage over lipid-like microparticles which may not be subjected to
mechanical
processing after they are formed due to their fragility.
The final formulation of the microparticles after preparation, but prior to
use,
is in the form of a lyophilized cake. The later reconstitution of the
microparticles may
be facilitated by lyophilization with bulking agents which provide a cake
having a
high porosity and surface area. The bulking agents may also increase the
drying rate
during lyophilization by providing channels for the water and solvent vapor to
be
removed. This also provides a higher surface area which would assist in the
later
reconstitution. Typical bulking agents are sugars such as dextrose, mannitol,
sorbitol
and sucrose, and polymers such as PEG's and PVP's.
It is undesirable for the microparticles to aggregate, either during
formulation
or during later reconstitution of the lyophilized material. Aggregation may be
minimized by maintaining a pH of at least one to two pH units above or below
the
isoelectric point(P;) of the biomaterial forming the outer surface. The charge
on the
surface is determined by the pH of the formulation medium. Thus, for example,
if the
surface of the biomaterial has a P; of 7 and the pH of the formulation medium
is
below 7, the microparticle will possess a net positive surface charge.
Alternatively, if
the pH of the formulation medium is greater than 7, the microparticle would
possess a
negative charge. The maximum potential for aggregation exist when the pH of
the
formulation medium approaches the P; of the biomaterial used in the outer
shell.
Therefore by maintaining a pH of the formulation medium at least one to two
units
above or below the P; of the surface, microparticle aggregation will be
minimized. As
an alternative, the microparticles may be formulated at or near the P; with
the use of
surfactants to stabilize against aggregation. In any event, buffer systems of
the final
formulation to be injected into the subject should be physiologically
compatible.
The bulking agents utilized during lyophilization of the microparticles may
also be used to control the osmolality of the final formulation for injection.
An
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osmolality other than physiological osmolality may be desirable during the
lyophilization to minimize aggregation. However, when formulating the
microparticles for use, the volume of liquid used to reconstitute the
microparticles
must take this into account.
5 Other additives may be included in order to prevent aggregation or to
facilitate
dispersion of the microparticles upon formulation. Surfactants may be used in
the
formulation such as poloxomers (polyethylene glycol-polypropylene glycol-
polyethylene glycol block co-polymers). Water soluble polymers also may assist
in
the dispersion of the microparticles, such as medium molecular weight
polyethyleneglycols and low to medium molecular weight polyvinylpyrolidones.
The microparticles may be soaked in a solution of the drug whereby the
solution diffuses into the interior. In particular, the use of bilayered
microparticles
where the inner shell has a porous characteristic allows for rapid diffusion
of a drug
solution into the hollow core. The microparticles may be re-dried such as by
15 lyophilization to produce a gas filled, drug containing microparticle.
Alternatively,
the drug may be dissolved in the organic phase with the biopolymer during the
microparticle forming process. Evaporation of the organic solvents causes the
drug to
coprecipitate with the biopolymer inside the microparticle.
It will be realized that various modifications of the above-described
processes
20 may be provided without departing from the spirit and scope of the
invention. For
example, the wall thickness of both the outer and inner layers may be adjusted
by
varying the concentration of the components in the microparticle-forming
solutions.
The mechanical properties of the microparticles may be controlled, not only by
the
total wall thickness and thicknesses of the respective layers, but also by
selection of
25 materials used in each of the layers by their modulus of elasticity and
elongation, and
degree of cross-linking of the layers. Upon certain conditions involving rapid
deposition of the inner polymer or very low inner polymer content, porosity of
the
inner polymer shell is observed. The pores range from approximately 0.1 to 2
micron
in diameter as observed under electron microscopy. Mechanical properties of
the
30 layers may also be modified with plasticizers or other additives.
Adjustment of the
strength of the shell may be modified, for example, by the internal pressure
within the
microparticles. Precise acoustical characteristics of the microparticle may be
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achieved by control of the shell mechanical properties, thickness, as well as
narrow
size distribution. The microparticles may be ruptured by ultrasonic energy to
release
gases trapped within the microparticles into the blood stream. In particular,
by
appropriately adjusting the mechanical properties, the particles may be made
to
5 remain stable to threshold diagnostic imaging power, while being rupturable
by an
increase in power and/or by being exposed to its resonant frequency. The
resonant
frequency can be made to be within the range of transmitted frequencies of
diagnostic
body imaging systems or can be a harmonic of such frequencies. During the
formulation process the microparticles may be prepared to contain various
gases,
including blood soluble or blood insoluble gases. It is a feature of the
invention that
microparticle compositions may be made having a resonant frequency greater or
equal
to 2 MHz, and typically greater or equal to 5 MHz.
EXAMPLE 1
1 S Controlled Rupture of Microcapsules
with Ultrasound In-Vitro
A 6% aqueous solution was prepared from a 25% solution of USP grade
human serum albumin (Alpha Therapeutic Corp) by dilution with deionized water.
The solution was adjusted to a pH of 3.49 using 1 N HCI. Separately, 8 parts
by
weight polycaprolactone (M.W. 50,000) and 45 parts cyclooctane were dissolved
in
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resulting rnicroparticles were retrieved by centrifugation and washed 2 times
in an
aqueous solution of 0.25% poloxamer.
Microscopic inspection of the suspension revealed spherical particles having a
thin-walled polymer shell with an outer protein layer and an organic liquid
core. The
5 peak diameter as, determined by the Malvern Micro particle size analyzer,
was 4.12
microns.
The suspension was then lyophilized from a suspension in 25mM glycine,
0.5% Pluronic f 127, 0.1 % sucrose, 3.0% mannitol and 5.0% PEG-3400. The
resulting dry cake was reconstituted with deionized water and examined under
the
10 microscope to reveal that the microparticles were spherical, discrete, and
contained a
gaseous core.
A test was performed to determine the ultrasound power threshold required to
disrupt the capsule of this bilaminate microcapsule (POINT Biomedical
biSphereTM)
made of albumin and polycaprolactone. The system was set up as follows. A
water
15 tank 18" long by 12" wide by 4" deep was constructed using acrylic side
walls and a
plate glass bottom. A liquid flow loop was constructed to circulate the
microsphere
agent under test, consisting of a length of vinyl tubing, running through a
peristaltic
pump and connected to a section of thin walled silicone tubing to create a
loop. The
imaging was performed with a Hewlett-Packard Sonos 2500 ultrasound scanner. A
20 harmonic imaging transducer (1.8/3.6 MHz) was connected to the Sonos 2500,
and
the transducer head mounted on a test stand such that only the transducer face
was
immersed in the test tank. The flow Loop was filled with de-gassed de-ionized
water
and the transducer was positioned to image the silicone tube.
With the transducer properly aligned, the Sonos 2500 parameters were
25 adjusted to give a clear image of the flow tube. The peristaltic pump was
set to the
lowest speed setting. After the set-up was complete, the Sonos 2500 was then
switched to harmonic imaging mode ( 1.8/3.6 MHz) to observe the harmonic image
response of the flow system. Without any contrast agent in the system, the
silicone
flow loop was barely visible. A small amount of the microcapsule suspension
was
30 injected into the flow loop. Under harmonic imaging mode, the contrast
agent was
immediately visible flowing through the silicone imaging tube.
The Sonos 2500 was then switched into Pulse Wave (PW) Doppler mode and
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a region of interest for the PW scan was centered on the apex of the silicone
flow
tube. Under PW mode, the ultrasound scanner converts the Doppler signal into
an
aural signal within the range of human hearing, and displays a repeating time
vs
intensity scan for the received signal. With the contrast agent flowing
through the
5 imaging tube, the transmission power was varied and the results evaluated on
the
ultrasound scanner screen and by listening to the aural signal from the
converted data.
At transmission power below 70% of the maximum, the PW received signal is a
steady baseline, and the aural signal is a steady flow sound overlaid by the
repetitive
sound of the peristaltic flow from the pump. A sample of the flowing liquid
was
10 withdrawn from the injection port and visualized under a standard light
microscope.
The microcapsules were still intact.
The transmission power was then increased above 70% and the signal scan
showed increasing high frequency spikes indicating the disruption of the
capsules of
the agent under test. The quantity of capsule disruption increased steadily
with
1 S increasing power, up to the maximum output. The aural signal was a series
of "click"
sounds which also increased with increasing power. Again, a sample of the
flowing
liquid was withdrawn from the loop and checked with a light microscope. After
ultrasound power input greater than 70%, the majority of the microsphere
capsules
appeared to be disrupted in some manner, usually fractured along an spherical
plane,
20 but in some cases reduced to wall fragments.
The results of the experiment were repeatable, even with the same sample
flowing in the loop, i.e. decreasing the power below the 70% threshold stopped
the
capsule disruption and increasing back above the 70% threshold began
disruption; a
process that could be repeated many times until all of the agent capsules
under test
25 had been disrupted.
EXAMPLE 2
Controlled Rupture of Microcapsules
with Ultrasound In-Vivo
30 A test similar to Example 1 was performed within the heart of a canine
subject. Suspensions of a bilaminate albumin/polycaprolactone microsphere
(POINT
Biomedical biSphereTM) were injected into an anesthetized dog and the wash-in
of the
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12
contrast effect observed in the chambers of the heart with a Hewlett Packard
Sonos
2500 medical ultrasound imaging system, using a 1.8/3.6 MHz harmonic
transducer
placed over the heart. After several minutes, an apparent steady state of
contrast
brightness was achieved in the heart. The ultrasound system was set to Pulse
Wave
5 (PW) Doppler mode as performed in Example 1, and the heart imaged. With the
region of interest set within the left ventricle, an increase in ultrasound
output power
created a series of audible clicks and corresponding sharp peaks on the
graphic
Doppler display of the ultrasound system, indicating the rupturing of the
microcapsules as demonstrated in-vitro. Moving the region of interest to the
10 myocardium still produced audible clicks and graphic peaks, but a reduced
level,
indicating microcapsule rupture within the reduced blood volume of the
myocardium.
EXAMPLE 3
Dye Loadin~~ Of Albumin Polylactide Microparticles
15 A 6% aqueous solution was prepared from a 25% solution of USP grade
human albumin by dilution with deionized water. Ion exchange resin ( AG 501-
X8,
BioRad Laboratories) was then added to the solution at a ratio of 1.5 gm resin
to 1.0
gm dry weight of albumin. After 3 hours the resin was removed by filtration
and the
pH of the solution was adjusted from 4.65 to 5.5 Separately, 0.41 gm d-1
lactide (0.69
20 dL/gm in CHC13: at 30 C) and 5.63 gm cyclooctane were dissolved in 37.5 gm
isopropyl acetate. The organic solution was then slowly incorporated into 25.0
gm of
the prepared albumin solution with mild stirring while the mixture was
maintained at
30 C. The resulting coarse o-w emulsion was then circulated through a
stainless steel
sintered metal filter element having a nominal pore size of 7 microns.
Recirculation of
25 the emulsion was continued for 8 minutes. The emulsion was then added with
stirring
to 350 ml deionized water maintained at 30 C and containing 1.0 ml of 25%
gluteraldehyde. During the addition, the pH of the bath was monitored to
insure that it
remained between 7 and 8. Final pH was 7Ø Low shear mixing was continued for
approximately 2'/2 hours until the isopropyl acetate had completely
volatilized.
30 Polyoxamer 188 in the amount of 0.75 gm was then dissolved into the bath.
The
resulting microspheres were retrieved by centrifugation and washed 2 times in
an
aqueous solution of 0.25% polyoxamer.
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13
Microscopic inspection revealed hollow spherical polymer microparticles
having an outer protein layer and an inner organic liquid core. The suspension
was
formulated with a glycine/PEG 3350 excipient solution, then lyophilized. The
resulting dry cake was reconstituted with deionized water and examined under
the
microscope to reveal that the microparticles were spherical, discrete, and
contained a
gaseous core.
A lyophilized cake in a 10 ml serum vial, composed of excipient and the
lactide-containing microparticles was placed into a 50 ml centrifuge tube.
Enough
isopropyl alcohol was added to cover the cake and it was allowed to soak for
30
10 seconds. Aqueous Pluronic F68 solution (0.25% w/w) was added to fill the
tube.
After centrifuging, the supernatant was removed and another rinse performed. A
saturated, filtered solution of rhodamine B was added to the microparticles
and
allowed to soak overnight. Under the microscope, the microparticles appeared
filled
with dye solution. A dye saturated F68 solution was made to use as a
lyophilization
15 excipient. Four ml of the excipient was combined with the approximately 2
ml of
microcapsule containing solution and the resulting mixture was split between
two 10
ml serum vials. The vials were frozen at -80°C and lyophilized in a FTS
tray dryer.
Both vials were purged with perfluorobutane gas by five pump-down purge cycles
with a vacuum pump. Observation showed some microparticles that were half full
of
20 red solution and half full of gas. There was no obvious leakage of the dye
from these
microparticles during observation. The microparticles were rinsed with four,
20 ml
portions of F68 solution on a vacuum filter. The microparticles were placed in
a
cuvette, centrifuged, and an initial spectra was taken. The cuvette was
sonicated in an
ultrasonic bath, centrifuged, and another spectra taken.
Abs. Initial (553-800) Abs. Sonicated (553-800)
1.164 1.86
The higher absorption after sonication indicates that encapsulated dye was
released upon insonation of the microparticles.
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14
EXAMPLE 4
Dve loading of Human Serum Albumin
Pol ~~caprolactone Microparticles
A lyophilized cake in a 10 ml serum vial, composed of excipient and paraffm-
containing microparticles prepared in accordance with example 1 was placed
into a 50
ml centrifuge tube. The only modification was that 0.2 gm paraffin was added
with
the polycaprolactone in cyclooctane. The cake was covered with methanol and
allowed to soak for 30 seconds. The tube was then filled with an aqueous
solution of
0.25% (w/w) Pluronic F68, gently mixed, and centrifuged in order to
precipitate the
now fluid-filled microcapsules. The supernatant was removed and the tubes were
again filled with pluronic solution. The microparticles were resuspended by
vortexing and again centrifuged. After removing the supernatant solution, two
ml of a
saturated, filtered solution of brilliant blue G dye in 0.25% (w/w) aqueous
F68 was
added. The microparticles were allowed to soak for approximately 72 hours.
15 Microscopic examination revealed 90-95% of the microparticles to be filled
with dye
solution. A lyophilization excipient was prepared. Four ml of the excipient
was
added to the microparticle solution and mixed by vortexing. Two 10 ml serum
vials
were filled with 3 ml each of solution and frozen at -80°C. The vials
were lyophilized
on a FTS flask lyophilizer. Both vials and a portion of deionized water were
purged
with perfluorobutane for i 0 minutes. Both vials were reconstituted with
deionized
water and rinsed with two 40 ml portions of 0.25% (w/w) F68 solution on a
vacuum
filter. The resulting microparticle solution was split into two 3 ml portions.
One
portion was sonicated in an ultrasonic bath to rupture the bubbles. Both
portions were
diluted 1/10 with F68 solution and placed into UV-visible cuvettes. The
cuvettes
25 were centrifuged and a visible spectra was taken.
Absorption (at 605nm-800nm)
Sonicated 0.193
Non-sonicated 0. I 36
The higher absorption after sonication indicates that encapsulated dye was
released upon insonation of the microcapsules.
