U.S. patent application number 12/471791 was filed with the patent office on 2009-09-24 for methods of treatment with drug loaded polymeric materials.
This patent application is currently assigned to Yale University. Invention is credited to Tarek Fahmy, Peter Fong, William Mark Saltzman.
Application Number | 20090239789 12/471791 |
Document ID | / |
Family ID | 36740923 |
Filed Date | 2009-09-24 |
United States Patent
Application |
20090239789 |
Kind Code |
A1 |
Saltzman; William Mark ; et
al. |
September 24, 2009 |
METHODS OF TREATMENT WITH DRUG LOADED POLYMERIC MATERIALS
Abstract
Polymeric microparticles have been developed which encapsulate
therapeutic compounds such as drugs, cellular materials or
components, and antigens, and can have targeting ligands directly
bound to the microparticle surface. Preferred applications include
use in tissue engineering matrices, wound dressings, bone repair or
regeneration materials, and other applications where the
microparticles are retained at the site of application or
implantation. Another preferred application is in the use of
microparticles to deliver anti-proliferative agents to the lining
of blood vessels following angioplasty, transplantation or bypass
surgery to prevent or decrease restenosis, and in cancer therapy.
In still another application, the microparticles are used to treat
or prevent macular degeneration when administered to the eye, where
agents such as complement inhibitors are administered.
Inventors: |
Saltzman; William Mark; (New
Haven, CT) ; Fahmy; Tarek; (New Haven, CT) ;
Fong; Peter; (New Haven, CT) |
Correspondence
Address: |
Pabst Patent Group LLP
1545 PEACHTREE STREET NE, SUITE 320
ATLANTA
GA
30309
US
|
Assignee: |
Yale University
New Haven
CT
|
Family ID: |
36740923 |
Appl. No.: |
12/471791 |
Filed: |
May 26, 2009 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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11830212 |
Jul 30, 2007 |
7550154 |
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12471791 |
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11170803 |
Jun 30, 2005 |
7534448 |
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11830212 |
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60677991 |
May 5, 2005 |
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60628778 |
Nov 17, 2004 |
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60616821 |
Oct 7, 2004 |
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60585047 |
Jul 1, 2004 |
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Current U.S.
Class: |
514/1.1 |
Current CPC
Class: |
A61P 17/02 20180101;
A61P 9/00 20180101; A61K 9/0048 20130101; A61K 9/1647 20130101;
A61K 9/5192 20130101; A61K 9/167 20130101; A61K 2039/542 20130101;
A61K 39/39 20130101; A61K 2039/6087 20130101; A61K 2039/55555
20130101; A61P 43/00 20180101; A61K 47/6937 20170801; Y10S 977/773
20130101; A61P 27/02 20180101; A61K 9/5153 20130101; A61K 9/0019
20130101; A61K 9/1641 20130101; A61P 35/00 20180101 |
Class at
Publication: |
514/8 |
International
Class: |
A61K 38/17 20060101
A61K038/17 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH
[0002] The U.S. government has certain right in this invention by
virtue of grants from the National Institutes of Health (EB00487
and CA52857) to William Mark Saltzman.
Claims
1. A method of treatment or prevention of undesirable proliferation
of cells comprising administering at a site at or adjacent to a
region of undesired proliferation microparticles delivering a high
density of an anti-proliferative factor in an amount effective to
prevent or decrease cellular proliferation.
2. The method of claim 1 wherein the undesired proliferation is
restenosis arising from endothelial dysfunction.
3. The method of claim 1 comprising administering the
microparticles at the time of or immediately following angioplasty,
vessel grafting, tissue or organ transplantation, synthetic vessel
implants, synthetic joint implants or other medical implants.
4. The method of claim 1 wherein the cells are endothelial cells,
wherein the method is for the treatment or prevention of macular
degeneration and the microparticles contain a high density of an
anti-angiogenic, anti-proliferative or complement inhibitor in an
amount effective to prevent or decrease vascularization of the
retina when the microparticles are administered intraocularly.
5. The method of claim 1 wherein the agent is a cytotoxic,
cytostatic, antiproliferative or anti-angiogenic agent and the
microparticles are administered locally or regionally for the
treatment of cancer.
6. The method of claim 1 wherein the microparticles further
comprise ligands having bound thereto targeting or attachment
molecules.
7. The method of claim 1 wherein the agent is a cytotoxic,
cytostatic, antiproliferative or anti-angiogenic agent, the
targeting molecules are specific for tumor cells and the
microparticles are administered to a individual having the tumor
cells.
8. A method for inhibiting calcification of surgical implants,
stents, prosthesis comprising implanting as part of or adjacent to
the surgical implant, stent, or prosthesis microparticles
comprising molecules inhibiting calcification.
9. The method of claim 8 wherein the molecules are osteopontin.
10. A tissue engineering matrix, wound dressing, or medical implant
comprising high density microparticles for delivery of a
therapeutic, nutritional, diagnostic or prophylactic agent
incorporated in a high density on or within the microparticle.
11. The matrix of claim 10 wherein the microparticles comprise
ligands having a first end incorporated into the surface of the
microparticle and a second end facing outwardly from the surface of
the microparticle, the ligands being present in a high density on
the surface of the microparticle and being bound to an agent to be
delivered selected from the group consisting of therapeutic,
nutritional, diagnostic and prophylactic agents, targeting and
attachment molecules.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application claims priority under 35 U.S.C. 119 to U.S.
Ser. No. 60/677,991 filed May 5, 2005, U.S. Ser. No. 60/628,778
filed Nov. 17, 2004, U.S. Ser. No. 60/616,821 filed Oct. 7, 2004,
and U.S. Ser. No. 60/585,047 filed Jul. 1, 2004.
FIELD OF THE INVENTION
[0003] The present invention relates to polymer microparticles for
treating disorders such as restenosis, macular degeneration, cancer
and transplantation.
BACKGROUND OF THE INVENTION
[0004] Biodegradable polymers have been used to deliver various
therapeutic agents. The therapeutic agents typically are
encapsulated within the biodegradable polymers which are formed
into particles having sizes of 100 .mu.m or less, films, sheets,
disks, pellets, or implants. The biodegradable polymers are
administered to a person, and the encapsulated therapeutic agent is
released within the body of the patient as the polymer degrades
and/or as water diffuses into the polymer to leach out the
encapsulated therapeutic. Biodegradable polymers, both synthetic
and natural, can release encapsulated agents over a period of days
or weeks, which can have benefits in administration of drugs or
other agents.
[0005] These devices have been modified to incorporate drug through
such techniques as solvent encapsulation, melt encapsulation, phase
separation, and other standard methods for processing of polymers.
The surfaces of the polymeric devices have been modified to
incorporate ligands, usually through either derivatization of the
polymer before formation of the device, or after formation of the
device using covalent binding to the polymer or ionic binding to
charged sites on the polymer. Many of these techniques have
disadvantages. Derivatization of the polymer prior to formation of
the device can result in many of the ligands being encapsulated
within the device, lowering the useful number of ligands available
for binding or targeting. Covalent binding after formation can
damage the polymers, lead to cross-reactions that decrease
specificity, and is typically not highly efficient. Ionic binding
is very gentle, but subject to dissociation, frequently not
possible in high density, and of low specificity.
[0006] Biodegradable polymers fabricated from
poly(lactic-co-glycolic acid) (PLGA) have emerged as powerful
potential carriers for small and large molecules of therapeutic
importance as well as scaffolds for tissue engineering
applications. This importance derives from: 1) Physiologic
compatibility of PLGA and its hompolymers PGA and PLA, all of which
have been established as safe in humans after 30 years in various
biomedical applications including drug delivery systems 2)
Commercial availability of a variety of PLGA formulations for
control over the rate and duration of molecules released for
optimal physiological response (Visscher et al. J Biomed Mater Res
1985; 19(3):349-65; Langer R, Folkman J. Nature 1976;
263(5580):797-800; Yamaguchi. J. Controlled Rel. 1993;
24(1-3):81-93.). 3) Biodegradability of PLGA materials, which
provides for sustained release of the encapsulated molecules under
physiologic conditions while degrading to nontoxic,
low-molecular-weight products that are readily eliminated (Shive et
al. Adv Drug Deliv Rev 1997; 28(1):5-24; Johansen et al. Eur J
Pharm Biopharm 2000; 50(1):129-46). 4) Control over its
manufacturing into nanoscale particles (<500 nm) for potential
evasion of the immune phagocytic system or fabrication into
microparticles on the length scale of cells for targeted delivery
of drugs or as antigen-presenting systems (Eniola et al. J Control
Release 2003; 87(1-3); 15-22; Jain R A. Biomaterials 2000;
21(23):2475-90). This unique combination of properties coupled with
flexibility over fabrication has led to interest in modifying the
PLGA surface for specific attachment to cells or organs in the body
(Eniola, et al. 2003; Keegan et al., Biomaterials 2003;
24(24):4435-4443; Lamprecht et al. J Pharmacol Exp Ther 2001;
299(2):775-81; Lathia et al. Ultrasonics 2004; 42(1-9):763-8 Park
et al. J Biomed Mater Res 2003; 67A(3):751-60; Panyam Adv Drug
Deliv Rev 2003; 55(3):329-47) for drug delivery and tissue
engineering applications. With a functional PLGA surface, cells may
be attached specifically to scaffolds enabling control over
interactions that lead to formation of optimal neotissue, or
encapsulated drug or antigen delivered specifically to the site of
interest potentially reducing deleterious drug side effects and
enhancing antigen delivery for vaccine applications.
[0007] A major difficulty associated with coupling ligands to PLGA
particles has been the lack of functional chemical groups on the
aliphatic polyester backbone for linking to target ligands. This
severely hinders the application of traditional conjugation methods
to the PLGA surface. Thus to introduce functionality into PLGA
surfaces several approaches have been studied. These include,
synthesis of PLGA copolymers with amine (Lavik et al J Biomed Mater
Res 2001; 58(3):291-4; Caponetti et al. J Pharm Sci 1999;
88(1):136-41) or acid (Caponetti et al J Pharm Sci 1999;
88(1):136-41) end groups followed by fabrication into particles.
Another approach involves the blending or adsorption of functional
polymers such as polylysine (Faraasen et al. Pharm Res 2003;
20(2):237-46; Zheng et al. Biotechnology Progress 1999;
15(4):763-767) or poly(ethylene-alt-maleic acid) (PEMA) (Keegan et
al. Macromolecules 2004) or PEG (Muller J Biomed Mater Res 2003;
66A(1):55-61) into PLGA and forming particles and matrices from
these blends (Zheng, et al. 1999; Keegan, 2004; Park et al. J
Biomater Sci Polym Ed 1998; 9(2):89-110; Croll Biomacromolecules
2004; 5(2):463-73; Cao et al. Methods Mol Biol 2004; 238:87-112).
Plasma treatment of the PLGA matrix has also been proposed for the
purpose of modifying its surface properties and introducing
hydrophilic functional groups into the polymer (Yang et al. J
Biomed Mater Res 2003; 67A(4):1139-47; Wan et al., Biomaterials
2004; 25(19):4777-83).
[0008] Targeting ligands include any molecule that recognizes and
binds to target antigen or receptors over-expressed or selectively
expressed by particular cells or tissue components. These may
include antibodies or their fragments, peptides, glycoproteins,
carbohydrates or synthetic polymers. The most widely used coupling
group is poly(ethylene glycol) (PEG), because this group creates a
hydrophilic surface that facilitates long circulation of the
nanoparticles. This strategy has been used successfully in making
`Stealth` liposomes with affinity towards target cells.
Incorporating ligands in liposomes is easily achieved by
conjugation to the phospholipid head group, in most cases
phosphotidylethanolamine (PE), and the strategy relies either on a
preinsertion of the functionalized lipid or post insertion into a
formed liposome. Functionality could also be introduced by
incorporating PEG with functional endgroups for coupling to target
ligands.
[0009] While these approaches have had good success in their
specific applications, their general use is hindered by drawbacks
such as difficulty associated with preparing the needed copolymers,
limited density of functional groups and targeting effects that
decrease with time due to desorption or degradation of adsorbed
group as the particle or scaffold erodes. It would be most
desirable to retain ligand function with control over its density
on the surface for prolonged periods of time for improved drug
delivery. There are also still a number of difficulties associated
with preparation of co-polymers, limited density of functional
groups and targeting groups with time due to degradation.
[0010] It is therefore an object of the present invention to
provide a polymer delivery system which can preferentially deliver
therapeutic compositions to selected cells or tissue and/or deliver
high amounts of therapeutic molecules.
[0011] It is another object of the invention to provide high
density, direct attachment to polymer, without harsh cross-linking
or coating requirements.
SUMMARY OF THE INVENTION
[0012] Microparticles are used to deliver therapeutics,
nutritional, diagnostic, or prophylactic agents in tissue
engineering applications, in treatment or prevention of restenosis,
in treatment or prevention of macular degeneration, and in cancer
therapy. In one embodiment, the microparticles are administered
with tissue engineering matrices, wound dressings, bone repair or
regeneration materials, and other applications where the
microparticles are retained at the site of application or
implantation. Another preferred application is in the use of
microparticles to deliver anti-proliferative agents to the lining
of blood vessels following angioplasty, transplantation or bypass
surgery to prevent or decrease restenosis, and in cancer therapy.
In still another application, the microparticles are used to treat
or prevent macular degeneration when administered to the eye, where
agents such as complement inhibitors are administered.
[0013] Polymeric delivery devices have been developed which combine
high loading/high density of molecules to be delivered with the
option of targeting. As used herein, "high density" refers to
microparticles having a high density of ligands or coupling agents,
which is preferably in the range of 1,000 to 10,000,000, more
preferably 10,000-1,000,000 ligands per square micron of
microparticle surface area. Targeting molecules can also be
attached to the surface of the polymers. Specificity is determined
through the selection of the targeting molecules. The effect can
also be modulated through the density and means of attachment,
whether covalent or ionic, direct or via the means of linkers. Drug
to be delivered can be encapsulated within the polymer and/or
attached to the surface of the polymer. The same or different
molecules to be delivered can be encapsulated or attached. This can
provide a two phase delivery or pulsed delivery.
[0014] A general method for incorporating molecules into the
surface of biocompatible polymers using materials with an HLB of
less than 10, more preferably less than 5, such as fatty acids, has
been developed. As demonstrated by the examples, avidin-fatty acid
conjugates were prepared and efficiently incorporated into
polylactic acid-glycolic acid ("PLGA"). In a preferred embodiment,
avidin is used as an adaptor protein to facilitate the attachment
of a variety of biotinylated ligands, although other attachment
molecules can be used. Fatty acids preferentially associate with
hydrophobic polymers, such as a PLGA matrix, rather than the
external aqueous environment, facilitating a prolonged presentation
of avidin over several weeks. Examples demonstrate this approach in
both microparticles encapsulating a model protein, bovine serum
albumin (BSA), and PLGA scaffolds fabricated by a salt leaching
method. Because of its ease, generality and flexibility, this
method has widespread utility in modifying the surface of polymeric
materials for applications in drug delivery and tissue engineering,
as well as other fields. The technology offers advantages over the
prior art: high density, direct attachment to the polymer material
without chemical modification of the PLGA, no harsh crosslinking
reagents required, no need for a coating to provide attachment
surfaces.
[0015] Targeted polymeric microparticles have also been developed
which encapsulate therapeutic compounds such as drugs, cellular
materials or components, and antigens, and have targeting ligands
directly bound to the microparticle surface. These microparticles
can be used to induce cellular immunologic responses or as
therapeutics. Targeting greatly increases specificity, while not
decreasing therapeutic load, such as DNA vaccines, drugs, peptides
proteins or antigens. Another advantage is that more than one
material can be encapsulated and/or coupled to the surface of the
microparticle. This may be a therapeutic and/or targeting material.
In some cases it may be advantageous to provide for an initial
delivery of molecules coupled to the surface of the microparticles,
with a second encapsulated therapeutic load being delivered
following phagocytosis or degradation of the microparticle.
BRIEF DESCRIPTION OF THE DRAWINGS
[0016] FIG. 1A is a scheme to modify a protein with palmitic acid.
NHS-palmitic acid is added to avidin at 10.times. molar excess and
reacted in the presence of 2% deoxycholate detergent. The NHS ester
reacts with avidin amine groups producing a stable amide linkage
and rendering the protein hydrophobic. Both reaction and
purification steps were in the presence of detergent to prevent
palmitate vesicle formation. FIG. 1B is a schematic of a
microparticle showing targeting molecules (antibody) and coupling
agent (avidin) and linkers (polyethylene glycol, PEG) on the
surface.
[0017] FIG. 2 is a graph of the degree of molecular crowding on the
surface of treated particles, determined by titrating
biotin-phycoerythrin ("PE") onto microparticles prepared with
various concentrations of avidin-palmitic acid (micrograms).
Surfaces modified with increasing amounts of the conjugate bound
more of the biotinylated fluorophore, as reflected by the higher
mean channel fluorescence (MCF).
[0018] FIG. 3 is a graph of the fraction of protein release over
time (hours) from avidin-palmitate microparticles versus unmodified
microparticles and surface modified microparticles.
[0019] FIGS. 4A and 4B are graphs of the stimulation of splenocytes
from mice vaccinated by subcutaneous administration of LPS targeted
microparticles encapsulating ovalbumin (closed circles) or with
control microparticles: no ovalbumin (closed diamonds), no LPS
targeting (open circles). FIG. 4A is stimulation of splenocytes
from vaccinated mice; FIG. 4B is stimulation of vaccinated mice in
the absence of ovalbumin antigen.
[0020] FIGS. 5A and 5B are graphs of the stimulation of splenocytes
from mice vaccinated by oral administration of LPS targeted
microparticles encapsulating ovalbumin (closed circles) or with
controls: phosphate buffered saline (closed squares), no LPS
targeting (open circles). FIG. 5A is stimulation of splenocytes
from vaccinated mice; FIG. 5B is stimulation of vaccinated mice in
the absence of ovalbumin antigen.
DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS
I. Polymeric Microparticles
[0021] As used herein, microparticles generally refers to both
microparticles in the range of between 0.5 and 1000 microns and
nanoparticles in the range of between 50 nm to less than 0.5,
preferably having a diameter that is between 1 and 20 microns or
having a diameter that is between 50 and 500 nanometers,
respectively. Microparticles and nanoparticles are also referred to
more specifically.
[0022] The external surface of the microparticles may be modified
by conjugating to the surface of the microparticle a coupling agent
or ligand. As described below, in the preferred embodiment, the
coupling agent is present in high density on the surface of the
microparticle.
[0023] As used herein, "high density" refers to microparticles
having a high density of ligands or coupling agents, which is
preferably in the range of 1,000 to 10,000,000, more preferably
10,000-1,000.000 ligands per square micron of microparticle surface
area. This can be measured by fluorescence staining of dissolved
particles and calibrating this fluorescence to a known amount of
free fluorescent molecules in solution.
[0024] The microparticle may be further modified by attachment of
one or more different molecules to the ligands or coupling agents,
such as targeting molecules, attachment molecules, and/or
therapeutic, nutritional, diagnostic or prophylactic agents.
[0025] A targeting molecule is a substance which will direct the
microparticle to a receptor site on a selected cell or tissue type,
can serve as an attachment molecule, or serve to couple or attach
another molecule. As used herein, "direct" refers to causing a
molecule to preferentially attach to a selected cell or tissue
type. This can be used to direct cellular materials, molecules, or
drugs, as discussed below.
[0026] Improved functionality is the ability to present target for
prolonged periods over the course of controlled release from the
particle (weeks), Functionality is improved because target molecule
remains associated with particle facilitating continuous function
over the duration of controlled release.
[0027] Surface modified matrices as referred to herein present
target that facilitate attachment of cells, molecules or target
specific macromolecules or particles.
[0028] Control over regional modification refers to the ability to
selectively modify sections of a biodegradable scaffold without
modifying the whole.
[0029] By varying the polymer composition of the particle and
morphology, one can effectively tune in a variety of controlled
release characteristics allowing for moderate constant doses over
prolonged periods of time. There have been a variety of materials
used to engineer solid nanoparticles with and without surface
functionality (as reviewed by Brigger et al Adv Drug Deliv Rev 54,
631-651 (2002)). Perhaps the most widely used are the aliphatic
polyesters) specifically the hydrophobic poly (lactic acid) (PLA),
more hydrophilic poly (glycolic acid) PGA and their copolymers,
poly (lactide-co-glycolide) (PLGA). The degradation rate of these
polymers, and often the corresponding drug release rate, can vary
from days (PGA) to months (PLA) and is easily manipulated by
varying the ratio of PLA to PGA. Second, the physiologic
compatibility of PLGA and its hompolymers PGA and PLA have been
established for safe use in humans; these materials have a history
of over 30 years in various human clinical applications including
drug delivery systems. Finally, PLGA nanoparticles can be
formulated in a variety of ways that improve drug pharmacokinetics
and biodistribution to target tissue by either passive or active
targeting.
[0030] A. Polymers
[0031] Non-biodegradable or biodegradable polymers may be used to
form the microparticles. In the preferred embodiment, the
microparticles are formed of a biodegradable polymer.
Non-biodegradable polymers may be used for oral administration. In
general, synthetic polymers are preferred, although natural
polymers may be used and have equivalent or even better properties,
especially some of the natural biopolymers which degrade by
hydrolysis, such as some of the polyhydroxyalkanoates.
Representative synthetic polymers are: poly(hydroxy acids) such as
poly(lactic acid), poly(glycolic acid), and poly(lactic
acid-co-glycolic acid), poly(lactide), poly(glycolide),
poly(lactide-co-glycolide), polyanhydrides, polyorthoesters,
polyamides, polycarbonates, polyalkylenes such as polyethylene and
polypropylene, polyalkylene glycols such as poly(ethylene glycol),
polyalkylene oxides such as poly(ethylene oxide), polyalkylene
terepthalates such as poly(ethylene terephthalate), polyvinyl
alcohols, polyvinyl ethers, polyvinyl esters, polyvinyl halides
such as poly(vinyl chloride), polyvinylpyrrolidone, polysiloxanes,
poly(vinyl alcohols), poly(vinyl acetate), polystyrene,
polyurethanes and co-polymers thereof, derivativized celluloses
such as alkyl cellulose, hydroxyalkyl celluloses, cellulose ethers,
cellulose esters, nitro celluloses, methyl cellulose, ethyl
cellulose, hydroxypropyl cellulose, hydroxy-propyl methyl
cellulose, hydroxybutyl methyl cellulose, cellulose acetate,
cellulose propionate, cellulose acetate butyrate, cellulose acetate
phthalate, carboxylethyl cellulose, cellulose triacetate, and
cellulose sulfate sodium salt (jointly referred to herein as
"synthetic celluloses"), polymers of acrylic acid, methacrylic acid
or copolymers or derivatives thereof including esters, poly(methyl
methacrylate), poly(ethyl methacrylate), poly(butylmethacrylate),
poly(isobutyl methacrylate), poly(hexylmethacrylate), poly(isodecyl
methacrylate), poly(lauryl methacrylate), poly(phenyl
methacrylate), poly(methyl acrylate), poly(isopropyl acrylate),
poly(isobutyl acrylate), and poly(octadecyl acrylate) (jointly
referred to herein as "polyacrylic acids"), poly(butyric acid),
poly(valeric acid), and poly(lactide-co-caprolactone), copolymers
and blends thereof. As used herein, "derivatives" include polymers
having substitutions, additions of chemical groups and other
modifications routinely made by those skilled in the art.
[0032] Examples of preferred biodegradable polymers include
polymers of hydroxy acids such as lactic acid and glycolic acid,
and copolymers with PEG, polyanhydrides, poly(ortho)esters,
polyurethanes, poly(butyric acid), poly(valeric acid),
poly(lactide-co-caprolactone), blends and copolymers thereof.
[0033] Examples of preferred natural polymers include proteins such
as albumin, collagen, gelatin and prolamines, for example, zein,
and polysaccharides such as alginate, cellulose derivatives and
polyhydroxyalkanoates, for example, polyhydroxybutyrate. The in
vivo stability of the microparticles can be adjusted during the
production by using polymers such as poly(lactide-co-glycolide)
copolymerized with polyethylene glycol (PEG). If PEG is exposed on
the external surface, it may increase the time these materials
circulate due to the hydrophilicity of PEG.
[0034] Examples of preferred non-biodegradable polymers include
ethylene vinyl acetate, poly(meth)acrylic acid, polyamides,
copolymers and mixtures thereof.
[0035] In a preferred embodiment, PLGA is used as the biodegradable
polymer.
[0036] The microparticles are designed to release molecules to be
encapsulated or attached over a period of days to weeks. Factors
that affect the duration of release include pH of the surrounding
medium (higher rate of release at pH 5 and below due to acid
catalyzed hydrolysis of PLGA) and polymer composition. Aliphatic
polyesters differ in hydrophobicity and that in turn affects the
degradation rate. Specifically the hydrophobic poly (lactic acid)
(PLA), more hydrophilic poly (glycolic acid) PGA and their
copolymers, poly (lactide-co-glycolide) (PLGA) have various release
rates. The degradation rate of these polymers, and often the
corresponding drug release rate, can vary from days (PGA) to months
(PLA) and is easily manipulated by varying the ratio of PLA to
PGA.
Formation of Microparticles.
[0037] In addition to the preferred method described in the
examples for making a high density microparticle, there may be
applications where microparticles can be fabricated from different
polymers using different methods.
[0038] a. Solvent Evaporation. In this method the polymer is
dissolved in a volatile organic solvent, such as methylene
chloride. The drug (either soluble or dispersed as fine particles)
is added to the solution, and the mixture is suspended in an
aqueous solution that contains a surface active agent such as
poly(vinyl alcohol). The resulting emulsion is stirred until most
of the organic solvent evaporated, leaving solid microparticles.
The resulting microparticles are washed with water and dried
overnight in a lyophilizer. Microparticles with different sizes
(0.5-1000 microns) and morphologies can be obtained by this method.
This method is useful for relatively stable polymers like
polyesters and polystyrene.
[0039] However, labile polymers, such as polyanhydrides, may
degrade during the fabrication process due to the presence of
water. For these polymers, the following two methods, which are
performed in completely anhydrous organic solvents, are more
useful.
[0040] b. Hot Melt Microencapsulation. In this method, the polymer
is first melted and then mixed with the solid particles. The
mixture is suspended in a non-miscible solvent (like silicon oil),
and, with continuous stirring, heated to 5.degree. C. above the
melting point of the polymer. Once the emulsion is stabilized, it
is cooled until the polymer particles solidify. The resulting
microparticles are washed by decantation with petroleum ether to
give a free-flowing powder. Microparticles with sizes between 0.5
to 1000 microns are obtained with this method. The external
surfaces of spheres prepared with this technique are usually smooth
and dense. This procedure is used to prepare microparticles made of
polyesters and polyanhydrides. However, this method is limited to
polymers with molecular weights between 1,000-50,000.
[0041] c. Solvent Removal. This technique is primarily designed for
polyanhydrides. In this method, the drug is dispersed or dissolved
in a solution of the selected polymer in a volatile organic solvent
like methylene chloride. This mixture is suspended by stirring in
an organic oil (such as silicon oil) to form an emulsion. Unlike
solvent evaporation, this method can be used to make microparticles
from polymers with high melting points and different molecular
weights. Microparticles that range between 1-300 microns can be
obtained by this procedure. The external morphology of spheres
produced with this technique is highly dependent on the type of
polymer used.
[0042] d. Spray-Drying In this method, the polymer is dissolved in
organic solvent. A known amount of the active drug is suspended
(insoluble drugs) or co-dissolved (soluble drugs) in the polymer
solution. The solution or the dispersion is then spray-dried.
Typical process parameters for a mini-spray drier (Buchi) are as
follows: polymer concentration=0.04 g/mL, inlet
temperature=-24.degree. C., outlet temperature=13-15.degree. C.,
aspirator setting=15, pump setting=10 mL/minute, spray flow=600
Nl/hr, and nozzle diameter=0.5 mm. Microparticles ranging between
1-10 microns are obtained with a morphology which depends on the
type of polymer used.
[0043] e. Hydrogel Microparticles. Microparticles made of gel-type
polymers, such as alginate, are produced through traditional ionic
gelation techniques. The polymers are first dissolved in an aqueous
solution, mixed with barium sulfate or some bioactive agent, and
then extruded through a microdroplet forming device, which in some
instances employs a flow of nitrogen gas to break off the droplet.
A slowly stirred (approximately 100-170 RPM) ionic hardening bath
is positioned below the extruding device to catch the forming
microdroplets. The microparticles are left to incubate in the bath
for twenty to thirty minutes in order to allow sufficient time for
gelation to occur. Microparticle particle size is controlled by
using various size extruders or varying either the nitrogen gas or
polymer solution flow rates. Chitosan microparticles can be
prepared by dissolving the polymer in acidic solution and
crosslinking it with tripolyphosphate. Carboxymethyl cellulose
(CMC) microparticles can be prepared by dissolving the polymer in
acid solution and precipitating the microparticle with lead ions.
In the case of negatively charged polymers (e.g., alginate, CMC),
positively charged ligands (e.g., polylysine, polyethyleneimine) of
different molecular weights can be ionically attached.
[0044] B. Molecules to be Encapsulated or Attached to the surface
of the Particles
[0045] There are two principle groups of molecules to be
encapsulated or attached to the polymer, either directly or via a
coupling molecule: targeting molecules, attachment molecules and
therapeutic, nutritional, diagnostic or prophylactic agents. These
can be coupled using standard techniques. The targeting molecule or
therapeutic molecule to be delivered can be coupled directly to the
polymer or to a material such as a fatty acid which is incorporated
into the polymer.
[0046] Functionality refers to conjugation of a ligand to the
surface of the particle via a functional chemical group (carboxylic
acids, aldehydes, amines, sulfhydryls and hydroxyls) present on the
surface of the particle and present on the ligand to be attached.
Functionality may be introduced into the particles in two ways. The
first is during the preparation of the microparticles, for example
during the emulsion preparation of microparticles by incorporation
of stabilizers with functional chemical groups. Example 1
demonstrates this type of process whereby functional amphiphilic
molecules are inserted into the particles during emulsion
preparation.
[0047] A second is post-particle preparation, by direct
crosslinking particles and ligands with homo- or heterobifunctional
crosslinkers. This second procedure may use a suitable chemistry
and a class of crosslinkers (CDI, EDAC, glutaraldehydes, etc. as
discussed in more detail below) or any other crosslinker that
couples ligands to the particle surface via chemical modification
of the particle surface after preparation. This second class also
includes a process whereby amphiphilic molecules such as fatty
acids, lipids or functional stabilizers may be passively adsorbed
and adhered to the particle surface, thereby introducing functional
end groups for tethering to ligands.
[0048] In the preferred embodiment, the surface is modified to
insert amphiphilic polymers or surfactants that match the polymer
phase HLB or hydrophile-lipophile balance, as demonstrated in the
following example. HLBs range from 1 to 15. Surfactants with a low
HLB are more lipid loving and thus tend to make a water in oil
emulsion while those with a high HLB are more hydrophilic and tend
to make an oil in water emulsion. Fatty acids and lipids have a low
HLB below 10. After conjugation with target group (such as
hydrophilic avidin), HLB increases above 10. This conjugate is used
in emulsion preparation. Any amphiphilic polymer with an HLB in the
range 1-10, more preferably between 1 and 6, most preferably
between 1 and up to 5, can be used. This includes all lipids, fatty
acids and detergents.
[0049] One useful protocol involves the "activation" of hydroxyl
groups on polymer chains with the agent, carbonyldiimidazole (CDI)
in aprotic solvents such as DMSO, acetone, or THF. CDI forms an
imidazolyl carbamate complex with the hydroxyl group which may be
displaced by binding the free amino group of a ligand such as a
protein. The reaction is an N-nucleophilic substitution and results
in a stable N-alkylcarbamate linkage of the ligand to the polymer.
The "coupling" of the ligand to the "activated" polymer matrix is
maximal in the pH range of 9-10 and normally requires at least 24
hrs. The resulting ligand-polymer complex is stable and resists
hydrolysis for extended periods of time.
[0050] Another coupling method involves the use of
1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDAC) or
"water-soluble CDI" in conjunction with N-hydroxylsulfosuccinimide
(sulfo NHS) to couple the exposed carboxylic groups of polymers to
the free amino groups of ligands in a totally aqueous environment
at the physiological pH of 7.0. Briefly, EDAC and sulfo-NHS form an
activated ester with the carboxylic acid groups of the polymer
which react with the amine end of a ligand to form a peptide bond.
The resulting peptide bond is resistant to hydrolysis. The use of
sulfo-NHS in the reaction increases the efficiency of the EDAC
coupling by a factor of ten-fold and provides for exceptionally
gentle conditions that ensure the viability of the ligand-polymer
complex.
[0051] By using either of these protocols it is possible to
"activate" almost all polymers containing either hydroxyl or
carboxyl groups in a suitable solvent system that will not dissolve
the polymer matrix.
[0052] A useful coupling procedure for attaching ligands with free
hydroxyl and carboxyl groups to polymers involves the use of the
cross-linking agent, divinylsulfone. This method would be useful
for attaching sugars or other hydroxylic compounds with bioadhesive
properties to hydroxylic matrices. Briefly, the activation involves
the reaction of divinylsulfone to the hydroxyl groups of the
polymer, forming the vinylsulfonyl ethyl ether of the polymer. The
vinyl groups will couple to alcohols, phenols and even amines.
Activation and coupling take place at pH 11. The linkage is stable
in the pH range from 1-8 and is suitable for transit through the
intestine.
[0053] Any suitable coupling method known to those skilled in the
art for the coupling of ligands and polymers with double bonds,
including the use of UV crosslinking, may be used for attachment of
molecules to the polymer.
[0054] Coupling is preferably by covalent binding but it may also
be indirect, for example, through a linker bound to the polymer or
through an interaction between two molecules such as strepavidin
and biotin. It may also be by electrostatic attraction by
dip-coating.
[0055] The molecules to be delivered can also be encapsulated into
the polymer using double emulsion solvent evaporation techniques,
such as that described by Luo et al., Controlled DNA delivery
system, Phar. Res., 16: 1300-1308 (1999).
[0056] i. Molecules to be Delivered
[0057] Agents to be delivered include therapeutic, nutritional,
diagnostic, and prophylactic compounds. Proteins, peptides,
carbohydrates, polysaccharides, nucleic acid molecules, and organic
molecules, as well as diagnostic agents, can be delivered. The
preferred materials to be incorporated are drugs and imaging
agents. Therapeutic agents include antibiotics, antivirals
(especially protease inhibitors alone or in combination with
nucleosides for treatment of HIV or Hepatitis B or C),
anti-parasites (helminths, protozoans), anti-cancer (referred to
herein as "chemotherapeutics", including cytotoxic drugs such as
doxorubicin, cyclosporine, mitomycin C, cisplatin and carboplatin,
BCNU, 5FU, methotrexate, adriamycin, camptothecin, and taxol),
antibodies and bioactive fragments thereof (including humanized,
single chain, and chimeric antibodies), antigen and vaccine
formulations, peptide drugs, anti-inflammatories, nutraceuticals
such as vitamins, and oligonucleotide drugs (including DNA, RNAs,
antisense, aptamers, ribozymes, external guide sequences for
ribonuclease P, and triplex forming agents).
[0058] Particularly preferred drugs to be delivered include
anti-angiogenic agents, antiproliferative and chemotherapeutic
agents such as rampamycin. Incorporated into microparticles, these
agents may be used to treat cancer or eye diseases, or prevent
restenosis following administration into the blood vessels,
Exemplary diagnostic materials include paramagnetic molecules,
fluorescent compounds, magnetic molecules, and radionuclides.
[0059] Alternatively, the biodegradable polymers may encapsulate
cellular materials, such as for example, cellular materials to be
delivered to antigen presenting cells as described below to induce
immunological responses.
[0060] Peptide, protein, and DNA based vaccines may be used to
induce immunity to various diseases or conditions. For example,
sexually transmitted diseases and unwanted pregnancy are world-wide
problems affecting the health and welfare of women. Effective
vaccines to induce specific immunity within the female genital
tract could greatly reduce the risk of STDs, while vaccines that
provoke anti-sperm antibodies would function as
immunocontraceptives. Extensive studies have demonstrated that
vaccination at a distal site--orally, nasally, or rectally, for
example--can induce mucosal immunity within the female genital
tract. Of these options, oral administration has gained the most
interest because of its potential for patient compliance, easy
administration and suitability for widespread use. Oral vaccination
with proteins is possible, but is usually inefficient or requires
very high doses. Oral vaccination with DNA, while potentially
effective at lower doses, has been ineffective in most cases
because `naked DNA` is susceptible to both the stomach acidity and
digestive enzymes in the gastrointestinal tract
[0061] Cell-mediated immunity is needed to detect and destroy
virus-infected cells. Most traditional vaccines (e.g. protein-based
vaccines) can only induce humoral immunity. DNA-based vaccine
represents a unique means to vaccinate against a virus or parasite
because a DNA based vaccine can induce both humoral and
cell-mediated immunity. In addition, DNAbased vaccines are
potentially safer than traditional vaccines. DNA vaccines are
relatively more stable and more cost-effective for manufacturing
and storage. DNA vaccines consist of two major components--DNA
carriers (or delivery vehicles) and DNAs encoding antigens. DNA
carriers protect DNA from degradation, and can facilitate DNA entry
to specific tissues or cells and expression at an efficient
level.
[0062] Biodegradable polymer particles offer several advantages for
use as DNA delivery vehicles for DNA based vaccines. The polymer
particles can be biodegradable and biocompatible, and they have
been used successfully in past therapeutic applications to induce
mucosal or humoral immune responses. Polymer biodegradation
products are typically formed at a relatively slow rate, are
biologically compatible, and result in metabolizable moieties.
Biodegradable polymer particles can be manufactured at sizes
ranging from diameters of several microns (microparticles) to
particles having diameters of less than one micron
(nanoparticles).
[0063] Dendritic cells (DCs) are recognized to be powerful antigen
presenting cells for inducing cellular immunologic responses in
humans. DCs prime both CD8+ cytotoxic T-cell (CTL) and CD4+
T-helper (Th1) responses. DCs are capable of capturing and
processing antigens, and migrating to the regional lymph nodes to
present the captured antigens and induce T-cell responses. Immature
DCs can internalize and process cellular materials, such as DNA
encoding antigens, and induce cellular immunologic responses to
disease effectors.
[0064] As used herein, the term "disease effector agents" refers to
agents that are central to the causation of a disease state in a
subject. In certain circumstances, these disease effector agents
are disease-causing cells which may be circulating in the
bloodstream, thereby making them readily accessible to
extracorporeal manipulations and treatments. Examples of such
disease-causing cells include malignant T-cells, malignant B cells,
T-cells and B cells which mediate an autoimmune response, and
virally or bacterially infected white blood cells which express on
their surface viral or bacterial peptides or proteins. Exemplary
disease categories giving rise to disease-causing cells include
leukemia, lymphoma, autoimmune disease, graft versus host disease,
and tissue rejection. Disease associated antigens which mediate
these disease states and which are derived from disease-causing
cells include peptides that bind to a MHC Class I site, a MHC Class
II site, or to a heat shock protein which is involved in
transporting peptides to and from MHC sites (i.e., a chaperone).
Disease associated antigens also include viral or bacterial
peptides which are expressed on the surface of infected white blood
cells, usually in association with an MHC Class I or Class II
molecule.
[0065] Other disease-causing cells include those isolated from
surgically excised specimens from solid tumors, such as lung,
colon, brain, kidney or skin cancers. These cells can be
manipulated extracorporeally in analogous fashion to blood
leukocytes, after they are brought into suspension or propagated in
tissue culture. Alternatively, in some instances, it has been shown
that the circulating blood of patients with solid tumors can
contain malignant cells that have broken off from the tumors and
entered the circulation. These circulating tumor cells can provide
an easily accessible source of cancer cells which may be rendered
apoptotic and presented to the antigen presenting cells.
[0066] In addition to disease-causing cells, disease effector
agents include microbes such as bacteria, fungi, yeast, viruses
which express or encode disease-associated antigens, and
prions.
[0067] The disease effector agents are presented to the antigen
presenting cells using biodegradable polymer microparticles as
delivery vehicles. The loaded microparticles are exposed to
immature antigen presenting cells, which internalize the
microparticles and process the material within the microparticles.
The microparticles may be administered to the patient and the
interaction between the microparticles and the antigen presenting
cells may occur in vivo. In a preferred embodiment, the
microparticles are placed in an incubation bag with the immature
antigen presenting cells, and the microparticles are phagocytosed
by the antigen presenting cells during the incubation period. The
resulting antigen presenting cells are then administered to the
patient to induce an immune response to the disease causing
agent.
[0068] ii. Targeting Molecules
[0069] Targeting molecules can be proteins, peptides, nucleic acid
molecules, saccharides or polysaccharides that bind to a receptor
or other molecule on the surface of a targeted cell. The degree of
specificity can be modulated through the selection of the targeting
molecule. For example, antibodies are very specific. These can be
polyclonal, monoclonal, fragments, recombinant, or single chain,
many of which are commercially available or readily obtained using
standard techniques. Table 1 is a list of ligand-targeted
nanoparticulate systems providing examples of useful ligands and
their targets. Examples of molecules targeting extracellular matrix
("ECM") include glycosaminoglycan ("GAG") and collagen. In one
embodiment, the external surface of polymer microparticles may be
modified to enhance the ability of the microparticles to interact
with selected cells or tissue. The method of example 1 wherein a
fatty acid conjugate is inserted into the microparticle is
preferred. However, in another embodiment, the outer surface of a
polymer microparticle having a carboxy terminus may be linked to
PAMPs that have a free amine terminus. The PAMP targets Toll-like
Receptors (TLRs) on the surface of the cells or tissue, or signals
the cells or tissue internally, thereby potentially increasing
uptake. PAMPs conjugated to the particle surface or co-encapsulated
may include: unmethylated CpG DNA (bacterial), double-stranded RNA
(viral), lipopolysachamide (bacterial), peptidoglycan (bacterial),
lipoarabinomannin (bacterial), zymosan (yeast), mycoplasmal
lipoproteins such as MALP-2 (bacterial), flagellin (bacterial)
poly(inosinic-cytidylic) acid (bacterial), lipoteichoic acid
(bacterial) or imidazoquinolines (synthetic).
TABLE-US-00001 TABLE 1 Selected list of ligand-targeted
nanoparticulate systems evaluated for in vitro or in vivo
therapeutics delivery Ligand Drug System Target Cells Evaluation
Nucleic acids Aptamers.sup.a PLA Prostate In vitro Epithelial cells
ECM Proteins Integrin.sup.b Raf genes Liposomes Melanoma cells In
vivo RGD peptides.sup.c siRNA polyethylene tumor vasculature In
vivo Imine) Fibrinogen.sup.d radioisotopes Albumin tumor
vasculature In vivo Lipids MP Lipid A.sup.e PLGA Dendritic cells In
vitro Carbohydrates Galactose.sup.f retinoic acid PLA Hepatocytes
In vitro Hyaluronic acid.sup.g Doxorubicin Liposomes CD44+ melanoma
cells In vitro Peptidomimetics.sup.h Various mPEG/PLGA Brain cells
Various Antibodies to: HER2 receptor.sup.i gelatin/HAS HER2 cells
In vitro HER2 receptor.sup.j Doxorubicin Liposomes HER2 cells In
vivo CD19.sup.k Doxorubicin Liposomes B cell lymphoma In vivo
Vitamins Folate.sup.l Doxorubicin Liposomes Leukemia cells In vivo
.sup.aPark, J. W. et al. Clin Cancer Res 8, 1172-1181 (2002).
.sup.bHood, J. D. et al. Science 296, 2404-2407 (2002).
.sup.cSchiffelers, R. M. et al. Nucleic Acids Res 32, e149 (2004).
.sup.dHallahan, D. et al. Cancer Cell 3, 63-74 (2003).
.sup.eElamanchili, et al. Vaccine 22, 2406-2412 (2004). .sup.fCho,
C. S. et al. J Control Release 77, 7-15 (2001). .sup.gEliaz, R. E.
& Szoka, F. C., Jr. Cancer Res 61, 2592-2601 (2001).
.sup.hOlivier, J. C. Neurorx 2, 108-119 (2005). .sup.iWartlick, H.
et al. J Drug Target 12, 461-471 (2004). .sup.jPark, J. W. et al.
Clin Cancer Res 8, 1172-1181 (2002) .sup.kLopes de Menezes, et al.
Cancer Res 58, 3320-3330 (1998). .sup.lPan, X. Q. et al. Blood 100,
594-602 (2002).
[0070] In another embodiment, the outer surface of the
microparticle may be treated using a mannose amine, thereby
mannosylating the outer surface of the microparticle. This
treatment may cause the microparticle to bind to the target cell or
tissue at a mannose receptor on the antigen presenting cell
surface. Alternatively, surface conjugation with an immunoglobulin
molecule containing an Fc portion (targeting Fc receptor), heat
shock protein moiety (HSP receptor), phosphatidylserine (scavenger
receptors), and lipopolysaccharide (LPS) are additional receptor
targets on cells or tissue.
[0071] Lectins that can be covalently attached to microparticles to
render them target specific to the mucin and mucosal cell layer
include lectins isolated from Abrus precatroius, Agaricus bisporus,
Anguilla anguilla, Arachis hypogaea, Pandeiraea simplicifolia,
Bauhinia purpurea, Caragan arobrescens, Cicer arietinum, Codiun
fragile, Datura stramonium, Dolichos biflorus, Erythrina
corallodendron, Erythrina cristagalli, Euonymus europaeus, Glycine
max, Helix aspersa, Helix pomatia, Lathyrus odoratus, Lens
culinaris, Limulus polyphemus, Lysopersicon esculentum, Maclura
pomifera, Momordica charantia, Mycoplasma gallisepticum, Naja
mocambique, as well as the letins Concanavalin A,
Succinyl-Concanavalin A, Triticum vulgaris, Ulex europaeus I, II
and III, Sambucus nigra, Maackia amurensis, Limaxfluvus, Homarus
americanus, Cancer antennarius, and Lotus tetragonolobus.
[0072] The attachment of any positively charged ligand, such as
polyethyleneimine or polylysine, to any microparticle may improve
bioadhesion due to the electrostatic attraction of the cationic
groups coating the beads to the net negative charge of the mucus.
The mucopolysaccharides and mucoproteins of the mucin layer,
especially the sialic acid residues, are responsible for the
negative charge coating. Any ligand with a high binding affinity
for mucin could also be covalently linked to most microparticles
with the appropriate chemistry, such as the fatty acid, conjugates
of example 1 or CDI, and be expected to influence the binding of
microparticles to the gut. For example, polyclonal antibodies
raised against components of mucin or else intact mucin, when
covalently coupled to microparticles, would provide for increased
bioadhesion. Similarly, antibodies directed against specific cell
surface receptors exposed on the lumenal surface of the intestinal
tract would increase the residence time of beads, when coupled to
microparticles using the appropriate chemistry. The ligand affinity
need not be based only on electrostatic charge, but other useful
physical parameters such as solubility in mucin or else specific
affinity to carbohydrate groups.
[0073] The covalent attachment of any of the natural components of
mucin in either pure or partially purified form to the
microparticles would decrease the surface tension of the bead-gut
interface and increase the solubility of the bead in the mucin
layer. The list of useful ligands would include but not be limited
to the following: sialic acid, neuraminic acid, n-acetyl-neuraminic
acid, n-glycolylneuraminic acid, 4-acetyl-n-acetylneuraminic acid,
diacetyl-n-acetylneuraminic acid, glucuronic acid, iduronic acid,
galactose, glucose, mannose, fucose, any of the partially purified
fractions prepared by chemical treatment of naturally occurring
mucin, e.g., mucoproteins, mucopolysaccharides and
mucopolysaccharide-protein complexes, and antibodies immunoreactive
against proteins or sugar structure on the mucosal surface.
[0074] The attachment of polyamino acids containing extra pendant
carboxylic acid side groups, e.g., polyaspartic acid and
polyglutamic acid, should also provide a useful means of increasing
bioadhesiveness. Using polyamino acids in the 15,000 to 50,000 kDa
molecular weight range would yield chains of 120 to 425 amino acid
residues attached to the surface of the microparticles. The
polyamino chains would increase bioadhesion by means of chain
entanglement in mucin strands as well as by increased carboxylic
charge.
[0075] Surface Modification with Liposomes
[0076] Microparticles can be further modified by encapsulation
within liposomes.
II. Applications
[0077] A. Drug Delivery
[0078] The submicron size of nanoparticulates offers distinct
advantages over larger systems: First, the small size enables them
to extravasate through blood vessels and tissue. This is especially
important for tumor vessels, which are often dilated and
fenestrated with an average pore size less than a micron, compared
to normal tissue. Second, solid nanoparticles made from
biodegradable polymers and encapsulating drug are ideal for
sustained intracellular drug delivery, especially for drugs whose
targets are cytoplasmic. An example of this application with
dexamethasone-loaded nanoparticles locally delivered to vascular
smooth muscle cells showed greater and sustained anti-proliferative
activity compared to free drug, indicating more efficient
interaction of the drug with cytoplasmic glucorticoid receptors.
The dosage loading varies depending on the nature of encapsulant.
Up to 80% of initial total amount of agent to be incorporated can
be encapsulated in the microparticles.
[0079] The microparticles are useful in drug delivery (as used
herein "drug" includes therapeutic, nutritional, diagnostic and
prophylactic agents), whether injected intravenously,
subcutaneously, or intramuscularly, administered to the nasal or
pulmonary system, administered to a mucosal surface (vaginal,
rectal, buccal, sublingual), or encapsulated for oral delivery. As
noted above, the term "microparticle" includes "nanoparticles"
unless otherwise stated. The dosage is determined using standard
techniques based on the drug to be delivered and the method and
form of administration. The microparticles may be administered as a
dry powder, as an aqueous suspension (in water, saline, buffered
saline, etc), in a hydrogel, organogel, or liposome, in capsules,
tablets, troches, or other standard pharmaceutical excipient.
[0080] In a preferred embodiment for delivery to a mucosal surface,
the microparticles are modified to include ligands for mucosal
proteins or extracellular matrix as described above.
[0081] i. Restenosis and Transplantation
[0082] Percutaneous transluminal coronary angioplasty (PTCA) is a
procedure in which a small balloon-tipped catheter is passed down a
narrowed coronary artery and then expanded to re-open the artery.
It is currently performed in approximately 250,000-300,000 patients
each year. The major advantage of this therapy is that patients in
which the procedure is successful need not undergo the more
invasive surgical procedure of coronary artery bypass graft. A
major difficulty with PTCA is the problem of post-angioplasty
closure of the vessel, both immediately after PTCA (acute
reocclusion) and in the long term (restenosis).
[0083] The mechanism of acute reocclusion appears to involve
several factors and may result from vascular recoil with resultant
closure of the artery and/or deposition of blood platelets along
the damaged length of the newly opened blood vessel followed by
formation of a fibrin/red blood cell thrombus. Restenosis (chronic
reclosure) after angioplasty is a more gradual process than acute
reocclusion: 30% of patients with subtotal lesions and 50% of
patients with chronic total lesions will go on to restenosis after
angioplasty. Although the exact hormonal and cellular processes
promoting restenosis are still being determined, it is currently
understood that the process of PTCA, besides opening the
artierosclerotically obstructed artery, also injures resident
coronary arterial smooth muscle cells (SMC). In response to this
injury, adhering platelets, infiltrating macrophages, leukocytes,
or the smooth muscle cells (SMC) themselves release cell derived
growth factors with subsequent proliferation and migration of
medial SMC through the internal elastic lamina to the area of the
vessel intima. Further proliferation and hyperplasia of intimal SMC
and, most significantly, production of large amounts of
extracellular matrix over a period of 3-6 months, results in the
filling in and narrowing of the vascular space sufficient to
significantly obstruct coronary blood flow.
[0084] The treatment of restenosis requires additional, generally
more invasive, procedures, including coronary artery bypass graft
(CABG) in severe cases. Consequently, methods for preventing
restenosis, or treating incipient forms, are being aggressively
pursued. One possible method for preventing restenosis is the
administration of anti-inflammatory compounds that block local
invasion/activation of monocytes thus preventing the secretion of
growth factors that may trigger SMC proliferation and migration.
Other potentially anti-restenotic compounds include
antiproliferative agents that can inhibit SMC proliferation, such
as rapamycin and paclitaxel. Rapamycin is generally considered an
immunosuppressant best known as an organ transplant rejection
inhibitor. However, rapamycin is also used to treat severe yeast
infections and certain forms of cancer. Paclitaxel, known by its
trade name Taxol.RTM., is used to treat a variety of cancers, most
notably breast cancer.
[0085] However, anti-inflammatory and antiproliferative compounds
can be toxic when administered systemically in
anti-restenotic-effective amounts. Furthermore, the exact cellular
functions that must be inhibited and the duration of inhibition
needed to achieve prolonged vascular patency (greater than six
months) are not presently known. Moreover, it is believed that each
drug may require its own treatment duration and delivery rate.
Therefore, in situ, or site-specific drug delivery using
anti-restenotic coated stents has become the focus of intense
clinical investigation. Recent human clinical studies on
stent-based delivery of rapamycin and paclitaxel have demonstrated
excellent short-term anti-restenotic effectiveness. Stents,
however, have drawbacks due to the very high mechanical stresses,
the need for an elaborate procedure for stent placement, and
manufacturing concerns associated with expansion and
contraction.
[0086] One of the most promising applications for targeted drug
delivery using nanoparticles is in local application using
interventional procedures such as catheters. Potential applications
have focused on intra-arterial drug delivery to localize
therapeutic agents in the arterial wall to inhibit restenosis
(Labhasetwar, et al. J Pharm Sci 87, 1229-1234 (1998); Song, et al.
J Control Release 54, 201-211 (1998)). Restenosis is the
re-obstruction of an artery following interventional procedures
such as balloon angioplasty or stenting as described above. Drug
loaded nanoparticles are delivered to the arterial lumen via
catheters and retained by virtue of their size, or they may be
actively targeted to the arterial wall by non-specific interactions
such as charged particles or particles that target the
extracellular matrix. Surface-modified nanoparticles, engineered to
display an overall positive charge facilitated adhesion to the
negatively charged arterial wall and showed a 7 to 10-fold greater
arterial localized drug levels compared to the unmodified
nano-particles in different models. This was demonstrated to have
efficacy in preventing coronary artery restenosis in dogs and pigs
(Labhasetwar, et al. J Pharm Sci 87, 1229-1234 (1998)).
Nanoparticles loaded with dexamethasone and passively retained in
arteries showed reduction in neointimal formation after vascular
injury (Guzman, et al. Circulation 94, 1441-1448 (1996)).
[0087] The microparticles (and/or nanoparticles) can be used in
these procedures to prevent or reduce restenosis. Microparticles
can be delivered at the time of bypass surgery, transplant surgery
or angioplasty to prevent or minimize restenosis. The
microparticles can be administered directly to the endothelial
surface as a powder or suspension, during or after the angioplasty,
or coated onto or as a component of a stent which is applied at the
time of treatment. The microparticles can also be administered in
conjunction with coronary artery bypass surgery. In this
application, particles are prepared with appropriate agents such as
anti-inflammatories or anti-proliferatives. These particles are
made to adhere to the outside of the vessel graft by addition of
adhesive ligands as described above. A similar approach can be used
to add anti-inflammatory or immunosuppressant loaded particles to
any transplanted organs or tissues.
[0088] In this embodiment, the drug to be delivered is preferably
an anti-proliferative such as taxol, rapamycin, sirulimus, or other
antibiotic inhibiting proliferation of smooth muscle cells, alone
or in combination with an anti-inflammatory, such as the steroidal
anti-inflammatory dexamethasone. The drug is encapsulated within
and optionally also bound to the microparticles. The preferred size
of the microparticles is less than one micron, more preferably
approximately 100 nm in diameter. The polymer is preferably a
polymer such as poly(lactic acid-co-glycolic acid) or
polyhydroxyalkanoate which degrades over a period of weeks to
months. Preferably the microparticles have a high density of an
adhesive molecule on the surface such as one that adds charge for
electrostatic adhesion, or one that binds to extracellular matrix
or cellular material, or otherwise inert molecules such as an
antibody to extracellular matrix component. Biotinylated particles
have a higher level of adhesion to the tissue.
[0089] ii. Treatment of Tumors
[0090] Passive delivery may also be targeted to tumors. Aggressive
tumors inherently develop leaky vasculature with 1.00 to 800 mm
pores due to rapid formation of vessels that must serve the
fast-growing tumor. This defect in vasculature coupled with poor
lymphatic drainage serves to enhance the permeation and retention
of nanoparticles within the tumor region. This is often called the
EPR effect. This phenomenon is a form of `passive targeting`. The
basis for increased tumor specificity is the differential
accumulation of drug-loaded nanoparticles in tumor tissue versus
normal cells, which results from particle size rather than binding.
Normal tissues contain capillaries with tight junctions that are
less permeable to nanosized particles. Passive targeting can
therefore result in increases in drug concentrations in solid
tumors of several-fold relative to those obtained with free
drugs.
[0091] Passive delivery may also be directed to lymphoid organs of
the mammalian immune system, such as lymphatic vessels and spleen.
These organs are finely structured and specialized in eliminating
invaders that have gained entry to tissue fluids. Nanoparticles may
easily penetrate into lymphatic vessels taking advantage of the
thin walls and fenestrated architecture of lymphatic microvessels.
Passive targeting to the spleen is via a process of filtration.
Indeed the spleen filters the blood of foreign particles larger
than 200 nm. This function facilitates splenic targeting with
nanoparticles encapsulating drug for effective treatments against
several hematological diseases.
[0092] Both liposomal and solid nanoparticles formulations have
received clinical approval for delivery of anticancer drugs.
Liposomal formulations include those of doxorubicin (Doxil1/Caelyx1
and Myocet1) and daunorubicin (Daunosome1). The mechanism of drug
release from liposomes is not clear, but is thought to depend on
diffusion of the drug from the carrier into the tumor interstitium.
This is followed by subsequent uptake of the released drug by tumor
cells. The mechanism of release is still poorly understood, which
hinders advanced applications involving the addition of active
ligands for cellular targeting in vivo. Recently, the FDA approved
Abraxane, an albumin-bound paclitaxel nanoparticles formulation as
an injectable suspension for the treatment of metastatic breast
cancer. In addition, other solid nanoparticle-based cancer
therapies have been approved for clinical trials, for example a
Phase 1 clinical trial has been approved that will evaluate the
safety of hepatic arterial infusion of REXIN-GTM (a targeted
nanoparticle vector system with a proprietary mutant cell-cycle
control gene, i.e. anti-cancer gene) as an intervention for
colorectal cancer.
[0093] The particles described herein should be efficacious in the
treatment of tumors, especially those where targeting is beneficial
and delivery of high doses of chemotherapeutic desirable. An
important feature of targeted particle delivery is the ability to
simultaneously carry a high density of drug while displaying
ligands on the surface of the particle. It is well known that other
drug carrier systems, such as immunotoxins or drug-immunoconjugate,
which are made by tethering drug molecules to antibodies or
synthetic polymers, usually deliver less than 10 drug molecules per
carrier to target cells. Targeted high density nanoparticles on the
other hand can deliver thousands of drug molecules on the surface,
and millions of molecules in their interior.
[0094] One important target is E-selectin, which is involved in the
arrest of circulating immune system cells and is differentially
upregulated with inflammatory and immune processes and should be
useful to enhance delivery of therapeutic agents to the vasculature
including tumor blood vessels through selective targeting. A second
important class of targets is receptors involved in the uptake of
vitamin B12, folic acid, biotin and thiamine. These are
differentially overexpressed on the surface of cancer cells
creating a possible target for several types of cancer, including
ovarian, breast, lung, renal and colorectal cancers. One of the
most promising strategies for enhancing active immunotherapy and
inducing potent vaccination is targeting of antigen-loaded
nanoparticles to antigen-presenting cells such as dendritic cells
(DCs). Nanoparticles incorporating toll-like receptors (TLRs) in
biodegradable PLGA have shown efficient delivery of antigen to DC
and potent activation of the T cell immune response.
[0095] The overall strength of nanoparticles binding to a target is
a function of both affinity of the ligand-target interaction and
the number of targeting ligands presented on the particle surface.
Nanoparticles produced by the present techniques have many
thousands of ligands on their surface. This is a particularly
useful feature for ligands that in their monomer form have a weak
affinity to their target receptors, such as single chain variable
fragments (scFv), which in most cases must be reengineered into
multimers to increase their avidity of interaction to target cells
or peptide/Major histocompatability complex (peptide/MHC), which
have weak affinity to target T cell receptors. For example,
multivalency increases the avidity of interaction of peptide/MHC to
the T cell up to 100 fold facilitating enhanced interactions and
effective drug delivery to target antigen-specific T cells.
[0096] iii. Macular Degeneration
[0097] Macular degeneration (MD) is a chronic eye disease that
occurs when tissue in the macula, the part of the retina that is
responsible for central vision, deteriorates. Degeneration of the
macula causes blurred central vision or a blind spot in the center
of your visual field. Macular degeneration occurs most often in
people over 60 years old, in which case it is called Age-Related
Macular Degeneration (ARMD) or (AMD). AMD is the leading cause of
blindness in the United States and many European countries. About
85-90% of AMD cases are the dry, atrophic, or nonexudative form, in
which yellowish spots of fatty deposits called drusen appear on the
macula. The remaining AMD cases are the wet form, so called because
of leakage into the retina from newly forming blood vessels in the
choroid, a part of the eye behind the retina. Normally, blood
vessels in the choroid bring nutrients to and carry waste products
away from the retina. Sometimes the fine blood vessels in the
choroid underlying the macula begin to proliferate, a process
called choroidal neovascularization (CNV). When those blood vessels
proliferate, they leak, causing damage to cells in the macula often
leading to the death of such cells. The neovascular "wet" form of
AMD is responsible for most (90%) of the severe loss of vision.
There is no cure available for "wet" or "dry" AMD.
[0098] The exact causes of AMD are not known, however, contributing
factors have been identified. Factors that contribute to AMD
include reactive oxidants which cause oxidative damage to the cell
s of the retina and the macula, high serum low density cholesterol
lipoprotein (LDL) concentration, and neovascularization of the
choroid tissue underlying the photoreceptor cells in the
macula.
[0099] Treatments for wet AMD include photocoagulation therapy,
photodynamic therapy, and transpupillary thermotherapy. AMD
treatment with transpupillary thermotherapy (TTT) photocoagulation
is a method of delivering heat to the back of the patient s eye
using an 810 nm infrared laser, which results in closure of
choroidal vessels. AMD treatment with photocoagulation therapy
involves a laser aimed at leakage points of neovascularizations
behind the retina to prevent leakage of the blood vessel.
Photodynamic therapy (PDT) employs the photoreactivity of a
molecule of the porphyrin type, called verteporphin or Visudyne,
which can be performed on leaky subfoveal or juxtafoveal
neovascularizations. Macugen is an FDA approved drug that inhibits
abnormal blood vessel growth by attacking a protein that causes
abnormal blood vessel growth.
[0100] Other potential treatments for "wet" AMD that are under
investigation include angiogenesis inhibitors, such as anti-VEGF
antibody, and anti-VEGF aptamer (NX-1838), integrin antagonists to
inhibit angiogenesis has also been proposed, and PKC412, an
inhibitor of protein kinase C. Cytochalasin E (Cyto E), a natural
product of a fungal species that inhibits the growth of new blood
vessels is also being investigated to determine if it will block
growth of abnormal blood vessels in humans. The role of hormone
replacement therapy is being investigated for treatment of AMD in
women.
[0101] There are no treatments available to reverse "dry" AMD.
Treatments shown to inhibit progression of AMD include supplements
containing antioxidants. The use of a gentle "sub-threshold" diode
laser treatment that minimizes damage to the retina is being
investigated for treatment of "dry" AMD. Another potential
treatment for AMD includes rheopheresis, which is a form of
therapeutic blood filtration that removes "vascular risk factor"
including LDL cholesterol, fibrinogen, and lipoprotein A.
Rheopheresis has not yet been FDA-approved, but is available in
Canada and Europe. Other treatments for AMD under investigation
include culturing and transplantation of cells of the Retinal
Pigment Epithelium (RPE), metalloproteinase modulators, inhibitors
of A2E, a vitamin A derivative, which accumulates in the human eye
with age, and carotenoids, zeaxanthin and lutein.
[0102] There have been a number of recent studies indicating that
macular degeneration is caused by, or associated with, a defect in
complement factor H (Haines, et al. Science. 2005 Apr. 15;
308(5720):419-21; Edwards, et al. Science. 2005 Apr. 15;
308(5720):421-4; Klein, et al. Science. 2005 Apr. 15;
308(5720):385-9). This leads to a method of treatment or prevention
of the macular degeneration through administration of one of the
known complement inhibitors, such as antibodies (antibody
fragments, recombinant antibodies, single chain antibodies,
humanized and chimeric antibodies) to C3b or a component thereof.
An example is Pexelizumab.TM. (Alexion Pharmaceuticals, Inc.,
Cheshire, Conn., USA), a humanized, monoclonal, single-chain
antibody fragment that inhibits C5, thereby blocking its cleavage
into active forms. A potential inhibitor is relatively small,
broad-acting C inhibitory protein (termed OmCI), described by Nunn,
et al. J. Immunol. 2005 Feb. 15; 174(4):2084-91.
[0103] Ocular delivery of drug-loaded, sustained-release and
optionally targeted nanoparticles by intravitreal administration is
a promising route for eye disease because it eliminates the need
for multiple injections of drug into the eye. Coupled with the
problem of retention of adequate concentrations of therapeutic
agent in the pre-corneal area (Mainardes, et al. Curr Drug Targets
6, 363-371 (2005)), biodegradable nanoparticles delivered
intravitreally have demonstrated localization in the retinal
pigment epithelium (Bourges, et al. Invest Opthalmol Vis Sci 44,
3562-3569 (2003)) and greater therapeutic efficacy in ocular
disease such as autoimmune uveoretinitis (de Kozak, et al. Eur J
Immunol 34, 3702-3712 (2004)).
[0104] In this embodiment, the drug is encapsulated with, and
optionally also bound to the microparticles. The preferred size of
the microparticles is approximately 100 nm in diameter. The polymer
is preferably a polymer such as poly(lactic acid-co-glycolic acid)
or polyhydroxyalkanoate which degrades over a period of weeks to
months.
[0105] In the preferred embodiment, degradable particles less than
one micron in diameter, preferably about 100 nm in diameter, are
distributed within the eye by subretinal injection or
intravitreally injection, where they degrade over a period of from
several weeks to several months. In the most preferred case, the
microparticles have a high density of adhesive molecules to retinal
epithelial cells.
[0106] B. Tissue Engineering Matrices and Wound Healing
Dressings
[0107] The microparticles can be dispersed on or within a tissue
engineering matrix for delivery of growth factors or modulatory
compounds, as demonstrated in the examples. Many types of materials
are known for use in tissue engineering, including materials formed
of synthetic polymer, decellularized matrix, collagen, and
decellularized tissue. These can be in the form of fibrous matrices
or materials such as those used in bone repair or replacement,
which consist primarily of materials such as hydroxyapatite. In
another embodiment, nanoparticles delivering molecules which are
used to enhance wound healing such as antibiotics, growth,
angiogenesis stimulating molecules, and other types of drugs, can
be applied to wound healing matrices, implants, dressings, bone
cements, and other devices which are applied to the site of injury.
Preferred antibiotics include vancomycin, ciprofloxacin and
anti-infective peptides such as the defensin molecules. In
addition, re-vascularization of these grafts can be a problem,
hence VEGF, FGF and PDGF could be included in the particles.
[0108] The advantage of these particles is that they adhere to the
implanted/applied material, where they are retained at the site of
injury to provide sustained treatment. Mixtures releasing different
amounts or different drugs at different times are particularly
advantageous for treatment of wounds such as diabetic wound ulcers.
Ligands can be selected to enhance the particles being retained at
the site, by binding to extracellular matrix or through
non-specific electrostatic binding. In addition, other ligands can
be selected to enhance the interaction of particles or matrix with
cells that are either added to the material prior to implantation
or migrate into the material after implantation.
[0109] The following examples describe testing performed using
microparticles of the present invention. It should be understood
that these examples are not intended to limit the scope, and are
provided only to present exemplary embodiments.
Example 1
Surface Modification of Biodegradable Polyesters with Fatty Acid
Conjugates for Improved Drug Targeting and Modification of Tissue
Engineering Materials
[0110] Materials
[0111] PLGA with an inherent viscosity of 0.59 dL/g, lot D02022 was
supplied from Birmingham Polymers, Inc. Polyvinyl alcohol
(M.sub.waverage 30-70 Kd), Palmitic acid-N-hydroxysuccinimide ester
(NHS-Palmitate), avidin (affinity purified) from egg white and
biotin-B-phycoerythrin, biotin immobilized on agarose were all
obtained from Sigma Chemical Co. Methylene Chloride and
trifluoroethanol were of chromatography grade and supplied by
Fischer Chemicals. All other reagents were of reagent grade and
used as received.
[0112] Preparation of Avidin-Palmitic Acid Conjugates
[0113] Avidin at 10 mg/ml was reacted with 10-fold excess of
NHS-Palmitic acid in PBS containing 2% deoxycholate buffer. The
mixture was sonicated briefly and gently mixed at 37.degree. C. for
12 hours. To remove excess fatty acid and hydrolyzed ester,
reactants were dialyzed against PBS containing 0.15% deoxycholate.
The resultant avidin-palmitate conjugate was verified by
reverse-phase HPLC on a Prevail.RTM. C18 column with a linear
methanol gradient in PBS as the mobile phase and UV detection at
280 nm.
[0114] Surface Modification and Characterization:
[0115] A modified water-in-oil-in-water (W/O/W) emulsion method was
used for preparation of fatty acid PLGA particles. In the first
emulsion, fluorescent bovine serum albumin (BSA-FITC) in 100 .mu.L
of PBS was added drop wise to a vortexing PLGA solution (5 ml)
dissolved in methylene chloride and trifluoroethanol (4:1) % V/V.
This first emulsion (W/O) was rapidly added to 200 ml of 5% PVA
containing the various concentrations of avidin-palmitic acid
investigated. This external phase underwent vigorous stirring for 4
hours at constant room temperature to evaporate methylene chloride
and trifluoroethanol. The resultant emulsion was then purified by
centrifugation at 12,000 g for 15 minutes then washed 3.times. with
DI water. No subsequent filtration or classification of particles
took place in this study. The particles were freeze-dried then
stored at -20.degree. C. Samples were characterized by Scanning
Electron Microscopy (SEM). Samples were sputter-coated with gold
under vacuum in an argon atmosphere using a sputter current of 40
mA (Dynavac Mini Coater, Dynavac USA). SEM analysis was carried out
with a Philips XL30 SEM using a LaB electron gun with an
accelerating voltage of 5 to 10 kV.
[0116] Surface Density and Functional Specificity
[0117] A colorimetric assay with 2-Hydroxyazobenzen-4'-Carboxylic
Acid (HABA) was used to quantitate the density of surface avidin
groups on PLGA particles. HABA binds to avidin to produce a
yellow-orange colored complex which absorbs at 500 nm. First, a
linear relationship between avidin in solution and HABA absorbance
was obtained by measuring the absorbance at 500 nm. This
standardized relationship was then used to quantitate the density
of surface avidin groups. In this assay 3 mg aliquots of dried
particles were suspended in 1 ml of 10 mM HABA (24.2 mg HABA in 10
mM NaOH). Biotin-phycoerythrin (Biotin-PE), a biotin conjugate of
the red fluorescent protein (PE) (240 kD), was used to monitor
surface functionality. On a rotary shaker the indicated amounts of
biotin-PE in PBS were added to 10 mg of plain and surface modified
particles. These solutions were incubated for 15 min then
centrifuged (10 min/11,000 g) and washed 3.times. in DI water.
Particle fluorescence was measured by flow cytometry.
[0118] Affinity to Target Under Dynamic Conditions:
[0119] Biotinylated agarose beads (2 ml of 4% crosslinked agarose)
were put into a fritted glass column and allowed to settle prior to
addition of plain or modified particles. The bed was briefly
sonicated to eliminate trapped air bubbles. Particles suspended in
PBS were gently added to the top of the packing and allowed to
settle into the packed bed prior to elution with PBS. The volume of
particles added to the bed did not exceed a tenth of the volume of
the packed bed. The column was then carefully filled with buffer
and a constant flow of buffer at 0.2 ml/min was maintained by a
Jasco pump. Fractions were collected every 0.5 ml into polystyrene
UV cuevettes and sample turbidity was analyzed by UV
spectrophotometry at 600 nm. Turbidity of the mixture was an
indicator of particle elution of the column. For modified
particles, when turbidity subsided, a 6M guanidine hydrochloride
was added to the column and fractions were collected as
described.
[0120] Surface Stability and Kinetics of BSA Release:
[0121] Release of encapsulated BSA-FITC and surface-bound biotin-PE
were carried out in phosphate buffer saline at 37.degree. C. At the
indicated time points samples were centrifuged for 10 min at 11,000
g and 1 ml supernatant from the samples was removed and replaced
with fresh buffer preincubated at 37.degree. C. The FITC and PE
content was measured by fluorescence ((.lamda..sub.excitation=480,
.lamda..sub.emission=520) for BSA-FITC and
(.lamda..sub.excitation=529, .lamda..sub.emission=576) for
biotin-PE. The fraction of protein released was calculated by
dividing the amount of BSA-FITC or biotin-PE at the indicated time
points by the total content of both proteins in 10 mg of the same
stock of particles. Total BSA-FITC content was measured by
dissolving 10 mg of particles in 1N NaOH overnight. A standard was
prepared by titrating BSA-FITC in 1N NaOH. Since Biotin-PE was
localized to the surface of the particles, red fluorescence of an
aliquot of (5 mg) particles was measured directly without need for
dissolution.
[0122] Surface Modification of PLGA Scaffolds:
[0123] PLGA 50/50 scaffolds were prepared by a salt-leaching method
(25). PLGA was dissolved in methylene chloride (10 mg in 500 ul).
Sodium chloride particles (100 mg with an averaged diameter,
100<d<250) were sprinkled into a round PVDF containers (Cole
Parmer #H-08936-00) followed by addition of PLGA solution. After
solvent evaporation (24 hthes at room temperature), scaffolds were
washed thoroughly in DI water for three days. Scaffolds were freeze
dried and stored at -20.degree. C. for later use. Avidin-palmitic
acid incorporation was a simple deposition procedure. A 100 ul drop
was regionally placed on top of dried scaffolds and allowed to soak
in for 15 min at RT, followed by washing 5.times. in 1.times.PBS+1%
BSA. For surface staining, the entire scaffold was incubated in a
biotin-PE solution for 10 min at room temperature followed by a
second wash in DI water.
[0124] Results and Discussion:
[0125] Palmitoylation of Avidin
[0126] The overall scheme to modify a protein with palmitic acid is
shown in FIG. 1A. NHS-palmitic acid is added to avidin at 10.times.
molar excess and reacted in the presence of 2% deoxycholate
detergent. The NHS ester reacts with avidin amine groups producing
a stable amide linkage and rendering the protein hydrophobic. Both
reaction and purification steps were in the presence of detergent
to prevent palmitate vesicle formation (Huang J Biol Chem 1980;
255(17):8015-8). Compared to free avidin, which eluted as a single
uniform peak with buffer alone, avidin-palmitic acid exhibited some
aggregation and eluted with methanol in the mobile phase. This
reflects the enhanced hydrophobicity of the conjugate. At higher
methanol concentrations in the mobile phase we observed several
elution peaks indicating different degrees of conjugate association
with the column. A possible explanation is that NHS-palmitic acid
targets individual lysine residues as well as the amino terminus of
the protein for conjugation; a process that can yield heterogeneous
populations of palmitoylated avidin that associate differently with
the hydrophobic stationary phase.
[0127] Effect of Surface Modification on Particle Morphology
[0128] Both plain and palmitoylated avidin particles displayed
heterogeneous size distributions. The average diameter of plain and
surface modified particles ranged from 4-7 um. Therefore, the
presence of avidin-palmitate in the emulsion and at the
concentrations used in this study did not impact significantly on
the size distribution of the particles. Strikingly, microparticles
prepared with conjugate in the emulsion showed a characteristic
texture and surface roughness by SEM. This characteristic varied
with the concentration of avidin-palmitic acid in the emulsion.
These images indicate that palmitic acid in the form of vesicles or
lamellae spread onto the surface of the PLGA during formation of
the particles. Surface spreading is facilitated by mechanical
dispersion or the presence of solvent (methylene chloride and
trifluoroethanol during the solvent evaporation step) or the
presence of low concentrations of detergent (0.15% deoxycholate) in
the final emulsion and during formation of the particles.
[0129] The observed characteristic changes in the surface
morphology of PLGA upon the addition of lipid or other amphiphilic
co-stabilizers have been observed previously in similar systems.
For example, when 1,2-dipalmitoylphosphatidycholine (DPPC) was used
to stabilize PLGA emulsions, significant changes in the surface
chemistry were observed by X-ray photoelectron spectroscopy (Evora
et al. J Control Release 1998; 51(2-3):143-52). The study is
consistent with this observation and supports the fact that the low
surface energy of lipid (DPPC) or palmitic acid, in contrast with
the high surface energy of PVA, dominates the surface chemistry of
PLGA contributing to the observed morphological changes. The study,
however, highlights that these changes may also facilitate the
presentation of surface functional groups for coupling to
proteins.
[0130] Surface Density and Functionality of Avidin-Palmitic Acid on
PLGA Particles
[0131] An increase in the absorbance of HABA at 500 nm correlates
with the presence of avidin in solution. This relationship was used
to verify and quantitate the density of surface avidin groups on
PLGA particles (Table 1). An apparent maximum in surface density
was observed with 0.25 mg of the conjugate per mg of PLGA in
emulsion. The efficiency of avidin-palmitate incorporation into
particles ranged between 14 to 24% with higher efficiencies of
incorporation observed at lower concentrations of the
avidin-palmitate in the emulsion. The presence of an apparent
maximum may therefore reflect the natural tendency of the fatty
acid to aggregate at higher concentrations; limiting its
partitioning into the forming PLGA phase.
[0132] To ascertain the functionality and specificity of
incorporated avidin to target biotinylated ligand, the fluorescence
of plain and modified particles treated with biotin-PE was compared
by flow cytometry. The mean channel fluorescence of surface
modified particles was approximately three orders of magnitude
greater than control microparticles. This functional specificity
was also qualitatively confirmed by fluorescence microscopy.
Fluorescence images showed regions of brighter fluorescence
indicating local high density binding regions on the particles
where conjugate might have localized.
[0133] To determine the degree of molecular crowding on the surface
of treated particles, biotin-PE was titrated onto microparticles
prepared with various concentrations of avidin-palmitic acid (FIG.
2). Surfaces modified with increasing amounts (0, 0.025 wt/v, 0.05
wt/v, 0.15 wt/v, 0.25 wt/v) of the conjugate bound more of the
biotinylated fluorophore, as reflected by the higher mean channel
fluorescence (MCF). A self-quenching of PE was observed with higher
concentrations of biotin-PE added to the particles. Self-quenching
which results in a slight decrease in MCF with increasing
concentration of fluorophore, occurs with the `crowding` of
fluorophores in localized regions in the proximity of 50-100 .ANG.
(Lakowicz J R. Principles of Fluorescence Spectroscopy. New York:
Plenum Press; 1986); an indication of the molecular crowding and
high density of biotin-PE at the surface of the particles.
[0134] Functional Avidity of Surface Modified Microparticles Under
Dynamic Conditions
[0135] In physiological settings injected particles rarely remain
static but undergo a shearing due to flow and encounters with cells
and tissue. Critical to the function of surface active particles in
these settings is their ability to bind their target (Hammer et al.
Annu. Rev. Mater. Res. 2001; 31:387-40). To assess functional
avidity under dynamic conditions, plain and surface modified
microparticles were injected into a column packed with biotinylated
agarose beads followed by elution with saline buffer. Plain
microparticles eluted quickly from the column with PBS; modified
microparticles, however, visibly adhered to the packing and did not
elute even with high buffer flow rates that physically disrupted
the packing. Elution of the modified particles required the
addition of 6M guandium hydrochloride (GuHCl); a strong protein
denaturant known to disrupt the biotin-avidin linkage. A mass
balance showed that while 1-3% wt plain microparticles adhered
nonspecifically to the column packing after buffer elution, 80-90%
of surface modified particles remained associated with the column
prior to GuHCl elution.
[0136] The Effect of Surface Modification on the Encapsulation
Efficiency of BSA
[0137] Because the strategy involved the simultaneous encapsulation
and surface modification of particles at the emulsion stage, the
addition of avidin-palmitic acid might affect the encapsulation
efficiency of BSA. Therefore the amount of encapsulated BSA in PLGA
particles modified with various concentrations of avidin-palmitate
in the emulsion was measured (Table 2).
TABLE-US-00002 TABLE 2 Avidin - % Encapsulation Maximal Palmitate
PVA (mg BSA/mg Polymer).sub.final Avidin density Biotin-PE Binding
(wt/vol) (wt/vol) Particle Yield % (mg BSA/mg Polymer).sub.initial
(ug/mg polymer) (ug/mg polymer) 0 2.5 40 .+-. 5 18.3 .+-. 2 N/A N/A
0.025 2.5 57 .+-. 5 30.7 .+-. 2 6 .+-. 1 1 0.05 2.5 56 .+-. 7 38.1
.+-. 4 9.5 .+-. 2 1.25 0.15 2.5 92 .+-. 6 46.0 .+-. 3 30 .+-. 2 2.0
0.25 2.5 98 .+-. 10 77.8 .+-. 5 35 .+-. 3 2.5
[0138] The results indicated that palmitoylation of microparticles
enhanced BSA encapsulation in a concentration dependent manner. The
encapsulation efficiency of particles modified with 0.25 (wt/vol)
avidin-palmitate was the fold greater than unmodified particles.
There has been an increase in the yield of particles with higher
concentrations of avidin-palmitate in the emulsion (Table 2).
Others have found similar effects on the encapsulation efficiency
and particle yields with the addition of pegylated Vitamin-E or the
lipid DPPC to a PLGA emulsion (Mu et al. J Control Release 2003;
86(1):33-48; Mu et al. J Control Release 2002; 80(1-3):129-44). A
possible mechanism for this general effect might involve the
increased hydrophobic stabilization due to the presence of
co-stabilizing amphipathic molecules such as fatty acids or lipids,
facilitating enhancements in PLGA particle formation and
encapsulation efficiency (Thomas in et al. J Pharm Sci 1998;
87(3):259-68).
[0139] Kinetics of BSA Release and Stability of the
Avidin-Palmitate Layer
[0140] FIG. 3 shows the release profiles of plain and surface
modified microparticles over the duration of a controlled release
experiment at 37.degree. C. for 25 days. Both plain and modified
particles had very similar BSA release kinetics with an initial
release burst during the first 24 hours followed by a gradual
release and a bulk erosion step (12 days) taking place nearly at
the same time for surface modified and unmodified particles. PE
fluorescence was almost negligible in the supernatant. Visually,
centrifuged particles appeared bright red during the entire time
course of the experiment. A cumulative loss of less than 10% PE
fluorescence was detected over this period of time indicating
stable surface functionality over the time of the experiment.
[0141] Using SEM, the morphology of the both plain and modified
particles was examined after 21 days. Surprisingly, while plain
microparticles showed substantial morphological changes at the
endpoint, modified particles were relatively spherical in shape. In
addition to showing less drastic morphologic changes by SEM, a
distinct capping layer was observed in most microparticles
examined. Because of the distinct surface topology associated with
surface modification, coupled with persistent binding avidity over
the time course these of the experiment, it was hypothesize that
the additional surface layer observed in eroded modified
microparticles might be due to surface rearrangement of the
avidin-palmitic acid groups and reorganization during sphere
degradation.
[0142] The fact that surface activity (>90%) was persistent for
several weeks, coupled with greatly reduced changes in morphology
and a possible reorganization of targeting groups during controlled
release suggests a significant robustness and resiliency of the
palmitoylated avidin surface. This is in light of the observation
that the surface likely experiences an acidic microclimate because
of polymer hydrolysis (Mader et al. Pharm Res 1998; 15(5):787-93;
Brunner et al Pharm Res 1999; 16(6):847-53; Shenderova et al. Pharm
Res 1999; 16(2):241-8).
[0143] Surface Modification of PLGA Scaffolds:
[0144] The approach to surface modification of PLGA particles was
translated to an effective strategy for modifying synthetic
matrices for tissue engineering applications. Scaffolds regionally
treated with avidin-palmitic acid displayed bright red
fluorescence, when incubated with biotin-PE, indicative of surface
functionality only in those treated regions. Moreover, these
scaffolds still maintained their red color after 3 weeks in PBS and
37.degree. C. This approach is simple and facilitates three
important aspects for successful tissue growth: 1) The ability of
the matrix to be reliably and easily functionalized for selective
cell attachment, 2) flexibility in terms of attaching a variety of
ligands, and 3) sustained presentation of ligands for long-term
proliferation and differentiation of attached cells on the
matrix.
[0145] A strategy for surface modification of PLGA by introducing a
functionally active amphipathic fatty acid, palmitic acid coupled
to the ligand of interest (avidin) during the emulsion preparation
of PLGA particles. This strategy was also translated to regional
modification of PLGA scaffolds for tissue engineering applications.
Because of the generality of this system and its flexibility,
different ligands may be attached to palmitic acid facilitating
surface modification with a variety of ligands and improving upon
in vivo particle targeting or clearance. For example combinations
of palmitoylated PEG and palmitoylated-avidin incorporated on the
same particle may serve as ideal vehicles that combine high
circulation lifetime with prolonged targeted drug delivery for in
vivo applications. In addition, the combination of regional
modification on PLGA scaffolds and ease of adjusting the density
and type of ligand make for a powerful strategy to adjusting ratios
of different cell types for various applications such as co-culture
and growth of functional tissue composed of several cell types
(Quirk et al. Biotech. Bioeng. 2003; 81(5):625-628)).
Example 2
Non-Specific Targeting with LPS for Delivery of a Protein
[0146] Lipopolysaccharide, LPS, represents the main outer membrane
component of Gram-negative bacteria and plays a key role during
severe Gram-negative infection. LPS is recognized by the TOLL-like
receptor 4 and is one of a class of ligands called PAMPS (Pathogen
Associated Molecular Patterns) which target TOLL receptors
associated with innate immunity (Non-specific immunity). These are
very effective components of adjuvants that help prime the innate
immune response against antigens for vaccination. As a result they
are critical components of adjuvants such as complete Freunds
adjuvant that stimulate a vigorous immune response. LPS is a
polysaccharide backbone with pendant fatty acids.
A. Vaccination by Subcutaneous Administration
[0147] In this particular application ovalbumin antigen is
encapsulated and mice are vaccinated by subcutaneous administration
with particles that have been modified with LPS and the results
compared with mice vaccinated with unmodified particles
encapsulating the same antigen.
[0148] Modified LPS particles induce a powerful response to the
ovalbumin antigen, whereas the unmodified particles showed very
little response. Blank particles also induced no response.
[0149] Methods and Materials.
[0150] LPS is added during formation of the microparticles,
preferably during emulsion formation, in a ratio of between 1 to 10
mg LPS per 200 mg of polymer. Ovalbumin encapsulation is between
100 .mu.g to 10 mg per 200 mg of polymer during emulsion
formation.
[0151] Mice were vaccinated subcutaneously with LPS/OVA particles,
OVA particles with no LPS and blank particles. Three days later
mice were sacrificed and splenocytes isolated. Splenocytes were
stimulated with OVA antigen in vitro to check for immune response.
If successful vaccination took place splenocytes would respond to
OVA antigen in a dose dependent manner. If no vaccination took
place splenocytes would not respond.
[0152] Results
[0153] FIGS. 4A and 4B are graphs of the stimulation of splenocytes
from mice vaccinated by subcutaneous administration of LPS targeted
microparticles encapsulating ovalbumin (closed circles) or with
control microparticles: no ovalbumin (closed diamonds), no LPS
targeting (open circles). FIG. 4A is stimulation of splenocytes
from vaccinated mice; FIG. 4B is stimulation of vaccinated mice in
the absence of ovalbumin antigen.
B. Oral Vaccination
[0154] Similar results were obtained when particles were
administered orally by oral gavage in fasted mice. A good
immunization response was observed after two weeks with one single
dose of particles fed to fasted mice. No boosters were given.
Results are shown in FIGS. 5A and 5B. FIGS. 5A and 5B are graphs of
the stimulation of splenocytes from mice vaccinated by oral
administration of LPS targeted microparticles encapsulating
ovalbumin (closed circles) or with controls: phosphate buffered
saline (closed squares), no LPS targeting (open circles). FIG. 5A
is stimulation of splenocytes from vaccinated mice; FIG. 5B is
stimulation of vaccinated mice in the absence of ovalbumin
antigen.
Example 3
Enhanced Targeting of Microparticles through the use of Star or
Branched PEG Linkers
[0155] An efficient method which facilitates simple attachment of T
cell antigens to a macromolecular carrier which encapsulates a high
density of immunomodulatory drug was developed. Antigen-presenting
drug carriers were constructed from a non-toxic, multi-branched
polyethylene glycol/polyamidoamine (PEG/PAMAM) dendritic vehicle. T
cell antigens were tethered to the branches of this vehicle while
drug was efficiently encapsulated in the core PAMAM which acts as a
`nanoreservoir` of drug. The potency of these vehicles in
modulating the T cell response with antibodies and major
histocompatability ligands to specific T cell populations was
demonstrated. Antigen-presenting carriers encapsulating the
antimitogenic drug, doxorubicin bound their target T cells with
avidities 10-100 fold greater than free antigen and consistently
downregulated the T cell response, while drug-free constructs
elicited strong stimulation of the target populations. Owing to the
flexibility over the nature and density of antigen presented as
well as drug incorporation, these high avidity artificial antigen
presenting vehicles have wide clinical use in a dual role as potent
immunostimulatory or immunosuppressive tools.
[0156] A defining characteristic of the T cell immune response is
its exquisite specific recognition of antigen. This specific
recognition in T cells is governed by the interaction of clonally
distributed T cell receptor (TCRs) with ligands on antigen
presenting cells composed of short peptides derived from
internalized protein antigen and bound to major histocompatability
(MHC) Class I or Class II molecules. Lack of recognition of cells
that have been infected by virus, transformed or otherwise altered
or faulty recognition of self-antigen can mediate the pathogenesis
of malignancies and autoimmune diseases. The T cell receptor
complex is therefore an important target for modulation of these
disease states.
[0157] While the ability to track the intensity and breadth of the
antigen-specific T cell response is clearly useful for disease
diagnosis, the added ability to target and modulate this response
can be used to fix immune system defects and restoring immune
competence. One approach for modulating the antigen-specific
response involves the induction of antigen-specific T cell
unresponsiveness or anergy by exposure to controlled doses of
antibodies to antigen-specific T cells or peptide/major
histocompatability ligands (peptide/MHC). A second approach
involves the conjugation of these reagents to immunosuppressive
drugs for direct delivery to target T cells. Conjugation of drug to
carrier antigens, however requires indirect and often difficult
chemistries to achieve unhindered antigen-presentation coupled with
effective drug delivery. Furthermore, the low-affinity of the
peptide/MHC-TCR, (1-100 .mu.m) coupled with the fact that most
antigen-specific T cell subsets are usually circulating at low
numbers has precluded the use of soluble peptide/MHC monomers for
sustained interactions to antigen-specific T cells. Thus
multimerization of the peptide/MHC is often necessary for enhanced
affinities to target T cells. It was hypothesized that T cell
targeting could be improved by the use of constructs with multiple
T cell antigens, permitting binding to the T cell with enhanced
avidity and significantly lower dissociation rates. If such
constructs could be produced with the added ability to load drug
molecules, they would be attractive reagents for sustaining the
interactions necessary for drug delivery to antigen-specific T
cells.
[0158] Soluble multivalent molecules were combined with a
technology that delivers a high density of drug to the cellular
target, thereby yielding a versatile, physiologically compatible,
multifunctional system that combines high avidity interactions with
targeted drug delivery to T cell subsets. A robust, non-toxic,
antigen-presenting carrier was engineered by linking poly(ethylene
glycol) chains (PEG) to a `nanoreservoir` poly(amidoamine)
spherical core (PAMAM) which functions as a high capacity drug
carrier. Doxorubicin was efficiently encapsulated in the PAMAM core
(32-mol doxorubicin per mol construct). Biotinylated antibodies or
biotinylated MHC were non-covalently attached to the PEG chains via
streptavidin linkers that were covalently linked to PEG.
Approximately 13 streptavidin molecules were attached per
construct. The constructs are specific and bind T cells with an
enhanced avidity, 10-100 times greater than free antibodies or
peptide/MHC chimeras. The complexes are small, with hydrodynamic
diameters in the range of 20-50 nm, allowing efficient
internalization and simultaneous fluorescent detection. In vitro
experiments with T cell specific antibody, anti-CD3s, coupled
constructs loaded with doxorubicin revealed a potent inhibition of
proliferation despite the presence of stimulation. Experiments with
peptide-specific MHC similarly revealed a significant modulation of
the T cell IL-2 response and end-point proliferation.
[0159] Methods and Materials
[0160] Mice: Balb/C mice (6-8 weeks) were obtained from Jackson
Laboratories (Bar Harbor, Me.). 2C TCR transgenic mice breeding
pair were a kind gift from Dr. Fadi Lakkis (Yale University School
of Medicine). 2C mice were maintained as heterozygous by breading
on a C57BL6 background in the animal facility. Phenotypes were
tested with the clonotypic 1B2 antibody, which was provided by Dr.
Jonathan Schneck (Johns Hopkins School of Medicine).
[0161] Cells: All cells used were obtained from homogenized naive
mouse spleens after depletion of RBC by hypotonic lysis. CD8+ cells
were isolated by negative selection from 2C splenocytes using CD8+
T cell subset enrichment columns (R&D systems). Purity>95%
was routinely obtained.
[0162] PEG/PAMAM: PAMAM Generation 6(Aldrich) 10 wt % in methanol
was evaporated under a gentle stream of nitrogen and placed under
high vacuum overnight before further manipulation. To prepare
fluorescently labeled constructs a 24 fold molar excess of
Boc-NH-PEG3400-NHS and a 6 fold molar excess of
fluorescein-PEG5000-NHS (Nektar Pharmaceuticals, Huntsville Ala.)
were added to PAMAM in a 0.2 M borate buffer pH 8.0. For unlabeled
constructs a 30 fold molar excess of PEG3400 was used. The mixture
was vortexed gently and placed on a rotary shaker for 24 hours.
Unreacted PEG was removed by dialysis in a 10,000 MWCO
Slide-a-Lyser (Pierce Chemical, Rockford Ill.) with borate as the
dialysis buffer. To remove the tBoc protecting group, the complex
was lyophilized for 48 hours and redissolved in trifluoroacetic
acid for 30 minutes at room temperature with constant stirring.
Trifluoroacetic acid was removed under vacuum for 1 hour. The
remaining product was dissolved in borate buffer followed by
dialysis in water. The final PEG/PAMAM complex was lyophilized once
more and stored at -20.degree. C. The characterization of these
complexes is discussed in detail in a previous report.sup.12.
[0163] Streptavidin-PEG/PAMAM: Streptavidin (Sigma) was activated
for amine coupling by dissolving at 1 mg/ml in 0.1 M MES, 0.5 M
NaCl buffer pH 5.1. To form active ester functional groups for
coupling NHS and EDC (Pierce Chemical Co.) was added at a
concentration of 5 mM and 2 mM respectively and allowed to react
for 15 min at room temperature. The unreacted EDC was quenched with
2-mercaptoethaol at a final concentration of 20 mM. For amine
coupling to the PEG/PAMAM, a 100 fold molar excess of activated
streptavidin was added to the PEG/PAMAM and reacted for 2 hours at
room temperature. Excess reactant and unconjugated streptavidin was
removed by extensive dialysis in a 200K MWCO CE ester membrane
(Spectrum Laboratories, Rancho Domingeuz Calif.). Homogeneity of
the complexes was assessed by reverse phase HPLC with 30%
acetonitrile as the mobile phase.
[0164] Dynamic light scattering: Sizes were measured by dynamic
light scattering (DLS). The instrument consisted of a diode pumped
laser (Verdi V-2/V-5, Coherent) operating at 532 .mu.m, an ALV-SP
S/N 30 goniometer (ALV-GmbH, Langen, Germany) with index matching
vat filled with doubly filtered (0.1 mm) toluene, and an ALV-500
correlator. Low concentrations of constructs (<5 ug/mL) were
pipetted into a cleaned borosilicate culture tube before measuring
the intensity of the auto-correlation function at a 90.degree.
scattering angle. The hydrodynamic radius, RH, was determined by
non-linear least squares fitting (ALV software) of the resulting
second order cumulants.
[0165] Antibody and MHC coupling: Biotinylated antibodies
(biotin-conjugated hamster anti-mouse CD3.epsilon. and
biotin-conjugated rat anti mouse CD45R/B220) (BD Biosciences
Pharmingen) were used without further purification. Soluble MHC-Ig
dimers L.sup.d-Ig were provided by Dr. Jonathan Schneck (Johns
Hopkins School of Medicine). MHC monomers were prepared from the
same dimer stock used in binding experiments by papain treatment of
the MHC-Ig and purified as described (Pierce Immunopure Fab
preparation kit). Preparation of MHC-Ig Fab fragments by papain
treatment yielded functionally active protein that specifically
bound TCR immobilized to the surface of a biosensor (Biacore) (data
not shown). MHC L.sup.d monomers and dimer were fluorescently
labeled with fluorescein isothiocyanate (FITC) (Molecular probes)
at pH 7.4 and purified by size exclusion chromatography. Protein
concentrations were determined spectrophotometrically by measuring
the absorbance at 280 nm. Both L.sup.d monomers and dimers were
loaded with peptide by stripping under mild acidic conditions (pH
6.5) and refolded in the presence of 40-fold molar excess peptide
and 2-fold molar excess b2-microglobulin. Using a conformationally
sensitive ELISA, it was estimated that >85% of the L.sup.d
monomers were folded properly, Biotinylated antibodies or L.sup.d
monomer were added at a 50 fold molar excess to
streptavidin-coupled PEG/PAMAM and incubated overnight at 4.degree.
C. followed by dialysis in a 300K MWCO CE membrane (Spectrum
Laboratories).
[0166] Doxorubicin loading of PEG/PAMAM constructs: Doxorubicin was
dissolved in water at a final concentration of 2.5 mg/ml and added
to a final concentration of 100 nM to PEG/PAMAM constructs in PBS
pH 7.4. The solution was mixed gently for 2 hours at 37.degree. C.
then 24 hours at 4.degree. C., followed by dialysis in 7000 MWCO
membranes (Pierce Chemical). Encapsulation efficiency was assessed
by fluorescence emission at 570 nm with 488 nm excitation. The
amount of doxorubicin loaded was deduced from a doxorubicin
calibration standard. To assess the magnitude of doxorubicin
fluorescence enhancement in the presence of PEG/PAMAM constructs,
doxorubicin at 2.5 mg/ml in water was titrated in 0.1 uL volumes in
a fluorometer cuevette in the presence or absence of PEG/PAMAM
constructs. Difference spectra were collected in the range 500-600
n with excitation at 488 nm.
[0167] In Vitro proliferation assays: Cells were adjusted to a
concentration of 1.times.10.sup.7 cells/ml in complete media.
Plates were coated with various concentrations of anti-CD3.epsilon.
antibodies according to established protocols. 2.times.10.sup.5
cells were plated per well Cells were treated with 20 nM complexes
either loaded or unloaded with doxorubicin and incubated at
37.degree. C., 5% CO.sub.2. To analyze the kinetics of IL-2
production, supernatants at the indicated time points were
harvested and analyzed by ELISA for IL-2 according to
manufacturer's instructions AD Biosciences, San Diego, Calif.), On
Day 3 T cell proliferation was analyzed with a calorimetric assay
for quantification of cell proliferation and viability, WST-1,
according to manufacturers protocol (Roche Diagnostics GmbH,
Pennsburg, Germany).
[0168] T cell Binding Assay: 1.times.10.sup.5 cells were incubated
with varying concentrations of the reagents discussed constructs
until equilibrium binding was reached (2 hrs, 4.degree. C.), Cells
were washed 3.times. with PBS with 1% Fetal bovine Serum and 0.1%
Sodium azide and analyzed by flow cytometry. The mean channel
fluorescence (MCF) was a measure of the amount of reagent bound.
Specific binding was normalized to the maximum mean channel
fluorescence.
[0169] FRET measurements: PEG/PAMAM constructs at 5 mg/ml were
labeled with a final concentration of 2.5 uM Alex Fluor.RTM. dye
546 (Donor) or Alex Fluor.RTM. 568 (Acceptor) (Molecular Probes,
Eugene, Oreg.) or equimolar mixtures of both fluorophores in a
carbonate buffer pH 8.3. After removal of excess dye by dialysis
the complexes were excited at 540 nm and emission spectra were
collected in the range (550-650 nm). Energy transfer efficiency, E,
was calculated from the relative fluorescence yield in the presence
(F.sub.da) and absence of acceptor (F.sub.d).sup.43,44 and was used
to calculate the energy transfer distance R from:
1 - ( F da F d ) = R 0 6 R 0 6 + R 6 ##EQU00001## where
##EQU00001.2## R 0 = 7.0 nm ##EQU00001.3##
[0170] Results
[0171] A branched, biocompatible, (24-30 arm) artificial
antigen-presenting polymer was constructed from polyethylene glycol
and generation 6 (G6) polyamidoamine dendrimer (PEG-PAMAM) by
methods reported by Luo, Macromolecules 35, 3456-3462 (2002). PAMAM
Starburst dendrimers are unique synthetic macromolecules with a
branched tree-like structure (Tomalia, et al. Angewandte
Chemie-International Edition in English 29, 138-175 (1990); Naylor,
et al. Journal of the American Chemical Society 111, 2339-2341
(1989)). G6 PAMAM tendrils radiate out from a central hydrophobic
core to create a well-defined globular architecture with 128
functional amine groups at the surface. Heterobifunctional PEG
M.sub.w3400 with a protected amine end (HOOC-PEG3400-NH-tBoc) was
covalently attached to the PAMAM tendrils and the amine end
deprotected after attachment. The working construct was a polymer
with radiating amine terminated PEG chains (4.2 nm) linked to a
hydrophobic core (6.7 nm). To facilitate detection of the
constructs, fluorescein terminated PEG chains were covalently
coupled to the dendrimer core at the molar ratio of 1:5 with
respect to amine-terminated PEG chains. The PAMAM cores of the
constructs can function as drug reservoirs, ideally suited as
vehicles for small drugs (Liu, et al. Abstracts of Papers of the
American Chemical Society 216, U875-U875 (1998); Kono, et al.
Abstracts of Papers of the American Chemical Society 221, U377-U377
(2001); Jansen, et al. Journal of the American Chemical Society
117, 4417-4418 (1995); Jansen, et al. Science 266, 1226-1229
(1994)), paramagnetic molecules for contrast enhancement in
magnetic resonance imaging (Kobayashi, et al. Mol Imaging 2, 1-10
(2003)), oligonucleotides (Yoo, et al. Pharm Res 16, 1799-804
(1999)), transgenes (Kobayashi, H. et al. Bioconjug Chem 10, 103-11
(1999)) and radionuclides (Kobayashi, Bioconjug Chem 10, 103-11
(1999)). Because the magnitude of spatial flexibility of the PEG
chains on the construct determines the degree of steric constraint
of proteins attached to the amine ends of PEG, the spatial
flexibility of branched PEGs was assessed by resonance energy
transfer. The amine reactive donor dye, Alexa fluor 546.RTM.
(Molecular Probes) and an acceptor dye, Alexa Fluor 568.RTM., were
conjugated to the amine ends of the unlabeled constructs followed
by purification of the construct by dialysis. The distance at which
fluorescence energy transfer from the donor dye to acceptor dye is
50% (R.sub.o is 7.0 nm) (Molecular Probes). Saturating
concentrations of a 1:1 molar ratio of both dyes conjugated to the
construct resulted in a pronounced decrease in donor fluorescence
and a sensitization of acceptor fluorescence. The transfer
efficiency calculated from the relative fluorescence yields of the
donor in the presence and absence of acceptor was between 50 and
57%. This efficiency was used to estimate a proximity distance
between the dyes of 6.+-.1 nm. This is sufficient distance for
coupling of proteins in the size range of streptavidin (3-4 nm).
Streptavidin coupling facilitates the attachment of a wide variety
of biotinylated ligands. In addition, because the T cell ligands
used in this study were biotinylated with a 2.2 nm biotin spacer
arm (NHS-LC-biotin.RTM.) Pierce Chemicals, it was estimated there
were sufficient flexible spatial interactions between streptavidin
coupled T cell ligands and their target receptors on T cells.
Analysis of the constructs is consistent with this estimate: the
coupling efficiency was approximately 13 streptavidin molecules per
construct with 5-10 fluorescein-terminated pendant chains.
[0172] The homogeneity of construct was verified by reverse phase
HPLC, which revealed a narrow distribution of the PEG/PAMAM and a
slightly wider distribution for streptavidin-PEG/PAMAM
(SA-PEG/PAMAM) constructs. The SA-PEG/PAMAM eluted earlier on a C18
column, probably due to the decrease in hydrophobicity and increase
in molecular size of construct that occurred with streptavidin
conjugation. Sizes of the constructs were also measured by dynamic
light scattering and estimated at 17.1 m and 26.4 nm for PEG/PAMAM
and SA-PEG/PAMAM respectively.
[0173] Antigen-presenting constructs bind their targets with
specificity and high avidity: To evaluate the specificity of
SA-PEG/PAMAM as a multivalent scaffold for T cell ligands,
SA-PEG/PAMAM was coupled to biotinylated antibodies that recognize
the T cell CD3 complex and anti-B220 that recognize the CD45R
antigen on B cells (negative control). Purified multivalent
complexes were incubated at saturating doses with a T cell enriched
(B cell depleted) population of splenocytes from Balb/C mice at
4.degree. C. for 2 his. The cells were then washed and the bound
complexes were analyzed by flow cytometry. Virtually no binding of
the control anti-B220 complexes was seen at the saturating dose
used in this study, but the specific anti-CD3 complex bound
strongly at the same dose. When the anti-CD3 complexes were
incubated at various concentrations with T cells, there was a
striking enhancement in the binding avidity of the constructs in
comparison with native fluorescently labeled anti-CD3 antibody.
Because avidity increases with increased valency of binding, and
because the PEG/PAMAM constructs have a higher valence (>13)
than antibodies, more of the anti-CD3 complexes bound compared to
the native antibody at a fixed ligand concentration. These
multivalent constructs therefore afford a higher sensitivity of T
cell detection at lower concentrations of the reagent.
[0174] Because the affinity of peptide/MHC-T cell interactions is
lower than antigen-antibody interactions, the efficacy of
SA-PEG/PAMAM complexes in increasing the sensitivity of detection
of clonotypic antigen-specific T cells was evaluated in a similar
binding assay. Biotinylated MHC Class I was coupled the constructs
and their binding compared with dimeric MHC constructs to purified
murine CD8+ T cell populations. The model system used was a murine
alloreactive Class I restricted CD8+2C T cell system that
recognizes the self-derived mitochondrial peptide, QLSPFPFDL (QL9)
presented in the context of the alloantigen Class I MHC H-2L.sup.d,
(.sup.Ql9L.sup.d) (Sykulev, Y. et al. Proc Natl Acad Sci USA 91,
11487-91 (1994)), and has little or no affinity to the same MHC
loaded with the negative control peptide YPHFMPNTL (MCMV),
(.sup.MCMVL.sup.d). Monomeric H-2L.sup.d was biotinylated at the
amine terminus and exogenously loaded with peptides QL9 and MCMV
using methods discussed in Fahmy, Immunity 14, 135-43 (2001)).
Modifications to the MHC similar to those discussed here have been
shown to have little or no affect on the MHC-T cell receptor
interaction by in vitro biosensor assays (Fahmy, et al. Immunity
14, 135-43 (2001)). Similar to binding profiles observed with
anti-CD3 constructs, .sup.QL9L.sup.d constructs bound 2C T cells
with enhanced avidity. The enhanced avidity was two orders of
magnitude greater, at half-maximal dose, in comparison with dimeric
forms of the MHC (.sup.QL9L.sup.d-Ig) (Schneck, Immunol Invest 29,
163-9 (2000)).
[0175] It was hypothesized that the enhanced avidity of these
complexes when coupled with the potential capacity of PAMAM for
carrying drug would be a powerful means of drug delivery to
specific T cell populations. To test this hypothesis, the ability
of the constructs to encapsulate the antimitogenic drug doxorubicin
was first assessed.
[0176] High-density encapsulation of doxorubicin by the PAMAM
dendritic core of antigen-presenting constructs. Previous work has
shown that doxorubicin (Dox), an anthracycline which intercalates
into DNA, can exhibit anti-proliferative effects and induce growth
arrest and apoptosis in proliferating T cells. Dox is intrinsically
fluorescent, thus detection of the drug is facilitated by
fluorescent detection with excitation at 488 nm and peak emission
at 570 nm in aqueous solutions. Dox is a weakly basic drug
(pKa=7.6) with limited solubility in aqueous environments.
Motivated by the potential utility of the hydrophobic dendrimer
core as a drug carrier, and the preferential association of Dox
with hydrophobic microenvironments (Dox octanol/water partition
coefficient is 2), the capacity of the constructs for passive
loading of doxorubicin was examined. Constructs were incubated with
a 10 fold molar excess of Dox at 4.degree. C. for 24 hours followed
by extensive dialysis in 7000 MWCO followed by fluorescence
measurements of the complexes. Using a doxorubicin fluorescence
calibration standard, it was estimated that approximately 55.+-.10
moles of Dox associated with each mole of construct. To verify that
the associated Dox is encapsulated in the dendrimer core it was
noted that Dox in an organic-aqueous solution simulating the
microenvironment of the PEG/PAMAM constructs showed an enhancement
in fluorescence. This enhancement in fluorescence was used to
assess the magnitude of Dox association with SA-PEG/PAMAM. A
similar enhancement was observed when comparing Dox fluorescence in
phosphate buffered saline in the presence of the construct. Since
PAMAM constitutes the largest hydrophobic fraction of the complex,
the data indicated an association of Dox with SA-PEG/PAMAM similar
to associations in organic-aqueous media. The magnitude of this
association based on fluorescence enhancement assays was used to
deduce the number of moles of associated drug per mole of
construct. The data peaked at a maximum lower than the amount
deduced from earlier equilibrium measurements. This might have been
due to formation of doxorubicin aggregates in the dialysis chamber
contributing to an overestimate of the amount associated with the
construct.
[0177] The data indicate that Dox is efficiently encapsulated in
the dendritic core of the antigen-presenting constructs.
Doxorubicin is efficiently released from the dendritic core at low
pH. Because drug loaded constructs are small (<100 nm); they are
efficiently internalized by their targets. To examine the level of
association of Dox with constructs in the acidic microenvironment
of endocytic vesicles, drug-construct interactions at pH 5 were
monitored. Dox loaded avidin-coupled constructs were immobilized on
a biotinylated agarose column, and washed with phosphate buffer
saline pH 7.4 before exposure to a low buffer environment
simulating lysosomal pH. Upon lowering the pH of the column, a
striking increase in Dox concentration in the eluent as monitored
by the red fluorescence of the drug was observed. A mass balance
revealed that greater than 90% of the Dox was efficiently released
from the constructs on lowering the pH of the mobile phase. The
data is consistent with a phenomenon known as the `ion trapping
hypothesis`, wherein weak bases with a hydrophobic character such
as doxorubicin become increasingly charged with lower pH and
preferentially partition to acidic compartments. All experiments in
the subsequent studies were performed with constructs saturated
with doxorubicin at the estimated amount of 32 mol Dox/mol
construct.
[0178] To test the efficacy of Dox-loaded anti-CD3 constructs in
downregulating the proliferative response of T cells in culture,
murine Balb/C splenocytes were stimulated with varying doses of
plate-bound anti-CD3 in the presence and absence of Dox-loaded
anti-CD3 and Dox-loaded anti-B220 constructs (negative control) and
measured T cell proliferation after 3 days. In contrast to
anti-B220-dox constructs, which showed little or no effect on
proliferating T cells, anti-CD3 Dox constructs were potent
inhibitors of proliferation. In these experiments, proliferation
was affected by two competing mechanisms: An enhancement in
proliferation due to the additional stimulus provided by the
presentation of anti-CD3-constructs and an inhibition in
proliferation due to specific drug delivery to target T cells.
[0179] To examine the utility of drug loaded antigen presenting
constructs in modulating the response and proliferation of
alloreactive antigen-specific T cell subsets,
.sup.QL9L.sup.d-constructs loaded with Dox (.sup.Ql9L.sup.dDox) and
.sup.MCMVL.sup.d Dox (negative control) were incubated with a
purified naive population of cytotoxic T cells, CD8+ T cells, from
2C mouse splenocytes. T cells were stimulated for 3 days in culture
in anti-CD3 coated plates in the presence or absence of constructs.
To monitor the response of the antigen-specific T cell culture, the
amount of IL-2 produced during the first three days of culture and
the total T cell proliferation after day 3 was measured. IL-2 is an
autocrine cytokine required for growth stimulation and
proliferation of T cells and is thus an important indicator of the
progression of T cell stimulation. The relative difference in IL-2
production between .sup.MCMVL.sup.dDox or .sup.Ql9L.sup.dDox after
day 1 was small and comparable to the amount of IL-2 produced by
untreated cells. This is an expected finding since naive T cells
require at least 20 hours of sustained signaling to be committed to
a vigorous proliferative response. We noticed a discernable change
between specific and non-specific inhibition of IL-2 after day 2.
At day 3 we observed a marked inhibition in IL-2 release from cells
treated with .sup.Ql9L.sup.dDox relative to untreated cells or
cells treated with .sup.MCMVL.sup.dDox. The finding that
.sup.MCMVLd.sup.DOX showed an inhibition effect relative to
untreated cells is consistent with the fact that the MCMV peptide
in the context of H-21.sup.d is not entirely non-specific to
purified 2C T cells in in vitro assays of T cell function.
[0180] At low concentrations of plate-bound anti-CD3 and in the
absence of Dox-loaded constructs, T cells exhibited a pronounced
release of IL-2 and concomitant proliferation which decreased
rapidly with higher levels of stimulation. While
.sup.MCMVL.sup.dDox IL-2 release and proliferation profiles were
lower than untreated cells, probably due to non-specific
interactions with T cells, it was found that by comparison
.sup.Ql9L.sup.dDox profoundly inhibited the production of IL-2 and
the proliferative capacity of antigen-specific T cells by greater
than 60%. Furthermore, .sup.Ql9L.sup.dDox inhibition of IL-2
release was effective over the entire dose range examined. Together
these results demonstrate an ability to selectively inhibit the
proliferation of polyclonal as well as antigen-specific populations
of T cells.
[0181] Discussion
[0182] The goal was to design a multifunctional system, which can
facilitate tracking via high avidity interactions as well as
delivering drugs to specific population of T cells. Because of the
functionality and demonstrated utility of PAMAM dendrimers as
non-toxic, nanoscopic polymers in drug delivery, these polymers
were chosen as a starting point and a core for the design of
multifunctional antigen presenting constructs. Polyethylene glycol
(PEG) was tethered to the dendrimer core for two reasons: First,
PEG is a linear polymer which imparts a flexibility to proteins
attached to the construct and allows for attached proteins to scan
a few nanometers of surface area for attachment to cell surface
receptors. Studies with MHC immobilized on planar membranes
demonstrated that T cells bound and responded most efficiently when
individual MHC molecules were less than 20 nm apart. Second,
proteins attached to PEG take on unusual properties such as
enhanced solubility, biocompatibility, lower immunogenicity and
desirable pharmacokinetics while the main biological functions such
as receptor recognition can often be maintained. These are critical
properties for long-term use of this technology and eventual
utility in clinical settings.
[0183] To accommodate the attachment of a wide variety of expensive
and difficult to prepare ligands, streptavidin was attached to the
PEG chains as an intermediate coupling protein. Streptavidin
facilitates the coupling of smaller amounts of biotinylated reagent
and expands the application of the scaffold to a wide range of
targets. This range of usage with biotinylated reagents that target
whole T cell populations or antigen-specific T cell populations was
demonstrated. Although the antigen-specific T cell studies in this
report have been performed with a class I MHC protein in an
alloreactive setting, the system described could be used in
conjunction with any biotinylated MHC applicable to other model
systems.
[0184] Unlike protein-based delivery systems which must be prepared
de novo and which have a limited capacity for carrying drug, the
PEG/PAMAM complexes described here have the capacity to carry up to
32 mol of doxorubicin per mol of construct. Thus this system offers
a therapeutic potential at lower concentrations comparable to
dose-dense free drug therapy. Control over the construct size,
number of sites available for conjugation and reactivity of the
various sites allows for control over the presentation of mixtures
of peptide/MHC and auxiliary ligands. The technology discussed is
unique because of this versatility. This feature is important for
addressing specific issues that depend on the nature and density of
ligand presented such as T cell tolerance, which is affected by the
density of antigen presented and co-stimulation.
Example 4
Attachment of poly(lactide-co-glycolide) (PLGA) Microparticles to
Decellularized Scaffolds for Drug Delivery in Cardiovascular Tissue
Engineering
[0185] The use of decellularized scaffolds in cardiovascular tissue
engineering is common due to their similar biomechanical properties
to native tissue. Unfortunately, these matrices undergo accelerated
calcification. The phosphoprotein, osteopontin, inhibits
calcification and could be used to decrease mineralization through
microparticle delivery. Furthermore, because cardiovascular tissue
calcifies in a known geometry, it would be of significant utility
if osteopontin could be delivered to specific locations of a
matrix.
[0186] Methods:
[0187] Osteopontin microparticles (125 .mu.g OPN/g PLGA) were
produced by spontaneous emulsification, washed by centrifugation,
and lyophilized for 24 hours. Sections of a porcine heart valve
were harvested, chemically decellularized, and subcutaneously
implanted in mice (n=3). One section was co-implanted with
osteopontin microparticles, while another was implanted alone as a
control. After 7 days the tissue was resected and evaluated for
calcification by atomic absorption spectroscopy. In a separate
experiment, to demonstrate microparticle attachment, decellularized
bovine metatarsal artery was biotinylated and then incubated with
avidin coated PLGA microparticles.
[0188] Results:
[0189] The tissue treated with osteopontin microparticles showed a
45.1% decrease in calcification as compared to untreated tissue.
PLGA microparticles were successfully attached to the fibers of a
decellularized bovine scaffold.
[0190] Conclusions:
[0191] These results demonstrate that osteopontin microparticles
can help inhibit calcification of cardiovascular structures
during/after surgical replacement procedures and can be locally
attached for matrix delivery. These particles can work on other
types of biological vascular grafts as well (i.e. xenografts for
heart valve replacement).
Example 5
Nanoparticles for Delivery of Rapamycin to Prevent Restenosis
[0192] Rapamycin is currently used to prevent restenosis by
application in a polymeric reservoir or coating as part of a stent.
The limitations of these devices are avoided through the separate
application of the nanoparticles at the time of or immediately
after a procedure such as angioplasty, vessel grafting, synthetic
vessel implants, synthetic joint implants or other medical implants
or at the time of bypass surgery. It has been demonstrated that the
short-term application of rapamycin, at the time of implantation,
can have significant long-term effects on restenosis. The advantage
of the nanoparticles is that there is no systemic delivery, and
release of an effective anti-proliferative amount can be achieved
over a period of weeks, during the time period most critical for
treatment.
[0193] A common form of bypass surgery involves resecting the
saphenous vein from the leg for autotransplantation to the coronary
artery. In 50% of the cases these grafts fail within 5
years--largely due to restenosis. Nanoparticles can be used for the
local and sustained delivery of rapamycin, or other
anti-proliferative agent to the autologous graft. After resection
of the saphenous vein the tissue can be, and often is for an hour
or more, suspended in saline while the patient's chest is opened
for graft implantation. The nanoparticles can be administered at
this time. One hour of particle attachment time in saline would be
more than sufficient.
[0194] Preparation Avidin Coated Rapamycin Nanospheres
[0195] Avidin at 10 mg/ml was reacted with 10-fold excess of
NHS-Palmitic acid in PBS containing 2% deoxycholate buffer. The
mixture was sonicated briefly and gently mixed at 37.degree. C. for
12 hours. To remove excess fatty acid and hydrolyzed ester,
reactants were dialyzed against PBS containing 0.15%
deoxycholate.
[0196] A modified double emulsion method was used for preparation
of fatty acid PLGA particles. In this procedure, 1 mg of rhodamine
B in 100 .mu.L of PBS, was added drop wise to a vortexing PLGA
solution (100 mg PLGA in 2 ml MeCl.sub.2). This mixture was then
sonicated on ice three times in 10-second intervals. At this point,
4 ml's of and avidin-palmitate/PVA mixture (2 ml avidin-palmitate
in 2 ml of 5% PVA) were slowly added to the PLGA solution. This was
then sonicated on ice three times in 10-second intervals. After
sonication the material was added drop-wise to a stirring 100 ml's
of 0.3% PVA. This underwent vigorous stirring for 4 hours at
constant room temperature to evaporate methylene chloride. The
resultant emulsion was then purified by centrifugation at 12,000 g
for 15 minutes then washed 3.times. with DI water. The particles
were freeze-dried then stored at -20.degree. C. Samples were
characterized by Scanning Electron Microscopy (SEM). Samples were
sputter-coated with gold under vacuum in an argon atmosphere using
a sputter current of 40 mA (Dynavac Mini Coater, Dynavac USA). SEM
analysis was carried out with a Philips XL30 SEM using a LaB
electron gun with an accelerating voltage of 5 to 10 kV.
[0197] Attachment of Nanoparticles to Ovine Carotid Artery.
[0198] Three 1.times.1 cm pieces of carotid arteries from sheep
were incubated in PLGA avidin labeled nanospheres loaded with
rhodamine (as a marker which is predictive of rapamycin
encapsulation and release) prepared as described above. The
incubation was done in a hybridization oven at 25.degree. C.,
facilitating attachment of the nanospheres through agitation by
placing them in a vial and suspending the vial to a vertically
rotating carousel.
[0199] A fluorescent micrograph at 10.times. magnification of
untreated sheep carotid artery not incubated in avidin
microparticles was compared with a fluorescent micrograph at
10.times. magnification of treated sheep carotid artery incubated
in avidin microparticles. As clearly visible in the micrograph
there is a high degree of fluorescene in the treated tissue as
compared to the untreated tissue-indicative of rhodamine nanosphere
attachment.
[0200] Stability of Attachment in a Sheer Stress Environment.
[0201] A tubular portion of ovine artery was nanosphere coated.
After nanosphere attachment the tube was connected to a bioreactor
where it supported phosphate buffered saline ("PBS") flow for one
hour. After this time, the tissue was removed from the bioreactor,
placed in an Eppendorf tube and incubated in fresh PBS to measure
the amount of rhodamine released from the conduit. After 1 hour the
conduit was placed in a new tube with fresh PBS and the old PBS was
measured for fluorescence. Four fractions were measured in this
manner. This demonstrated that the nanosphere coated conduit was
capable of delivering drug in a controlled fashion without total
washout of the particles after sheer stress.
[0202] Choice of Particle Size.
[0203] Nanoparticles (50-500 nm) were used in the coupling system.
Maximizing the surface area to unit mass of particle should improve
the binding of the particles to the vascular tissue. Nanoparticles
are also better in that washout of the particles will cause
downstream occlusion of smaller vessels (capillaries can be as
small as 5 microns).
[0204] Rapmycin Encapsulation.
[0205] Rapamycin was encapsulated in PLGA nanoparticles and
bioactivity verified using a PBMC assay. Briefly, PBMC cells were
stimulated with IL12 and IL18. In the presence of rapamycin,
interferon secretion is inhibited, resulting in an inverse
correlation between rapamycin concentration and interferon levels.
In this particular experiment, 10 mgs of rapamycin particles were
suspended in 10 mls of PB S. At various time points, 100 .mu.l of
PBS were taken from the 10 mls for subsequent treatment of the
PBMCs. This data indicates that the rapamycin released from the
nanoparticles are bioactive.
[0206] Rapamycin Dosing.
[0207] The desired dosing of rapamycin to autografts based on stent
data has been calculated as a target coating amount of rapamycin of
between one and 500 .mu.g/mm.sup.2, more preferably between 200
.mu.g/mm.sup.2 graft and 2 mg/mm.sup.2 graft, with approximately
75% of rapamycin eluted at 28 days. Release can occur over a range
in dosage from the time of implantation to between three days and
six months after implantation.
Example 6
Microparticles for Delivery of Antibiotics in Tissue Engineered
Matrices, INTEGRA.TM.
[0208] Materials and Methods
[0209] Integra.TM., a tissue engineering product used to treat
burns as a synthetic skin, was treated with nanoparticles that were
designed to adhere to the tissue-like matrix. Three 1.times.1 cm
pieces of INTEGRA.TM. from were incubated in PLGA avidin labeled
nanospheres loaded with rhodamine (as a marker which is predictive
of rapamycin encapsulation and release), prepared as described
above in Example 5. The incubation was done in a hybridization oven
at 25.degree. C., facilitating attachment of the nanospheres
through agitation by placing them in a vial and suspending the vial
to a vertically rotating carousel.
[0210] Results
[0211] A fluorescent micrograph at 10.times. magnification of
untreated INTEGRA.TM. not incubated in avidin microparticles was
compared with a fluorescent micrograph at 10.times. magnification
of treated INTEGRA.TM. incubated in avidin microparticles. As
clearly visible in the micrograph there is a high degree of
fluorescence in the treated tissue as compared to the untreated
tissue-indicative of rhodamine nanosphere attachment.
[0212] INTEGRA.TM. is used as a skin graft for burn victims.
Typically, a patient with second or third degree burns is treated
with INTEGRA.TM. for a couple of weeks before an autologous skin
graft is applied. Unfortunately, infection is a major problem with
this type of treatment. This study demonstrates that the particles
can be used to `dip-coat` INTEGRA.TM. in nanoparticles such that
those nanoparticles attach and deliver agent to the INTEGRA.TM. for
a couple of weeks following application to the wound.
* * * * *