U.S. patent application number 13/609588 was filed with the patent office on 2013-01-03 for polymer nanofilm coatings.
This patent application is currently assigned to The Regents of the University of California. Invention is credited to Mark Chen, Genhong Cheng, Edward K. Chow, Dean Ho, Houjin Huang, Erik Pierstorff.
Application Number | 20130004560 13/609588 |
Document ID | / |
Family ID | 40160830 |
Filed Date | 2013-01-03 |
United States Patent
Application |
20130004560 |
Kind Code |
A1 |
Ho; Dean ; et al. |
January 3, 2013 |
Polymer Nanofilm Coatings
Abstract
Disclosed herein are nanofilm coatings for implantable medical
devices comprising a diblock or triblock copolymer (PEO-PMMA or
PMOXA-PDMS-PMOXA, respectively). Such nanofilms, may be used, for
example, as amphiphilic supports for therapeutic agents. These
materials are conducive towards the formation of active substrates
for a suite of biological and medical applications.
Inventors: |
Ho; Dean; (Evanston, IL)
; Chen; Mark; (Chicago, IL) ; Pierstorff;
Erik; (Highland Park, IL) ; Huang; Houjin;
(Evanston, IL) ; Chow; Edward K.; (Rancho Palos
Verdes, CA) ; Cheng; Genhong; (Calabasas,
CA) |
Assignee: |
The Regents of the University of
California
Oakland
CA
Northwestern University
Evanston
IL
|
Family ID: |
40160830 |
Appl. No.: |
13/609588 |
Filed: |
September 11, 2012 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
12135640 |
Jun 9, 2008 |
8263104 |
|
|
13609588 |
|
|
|
|
60942885 |
Jun 8, 2007 |
|
|
|
60981688 |
Oct 22, 2007 |
|
|
|
Current U.S.
Class: |
424/443 ;
428/141; 428/216; 428/220; 977/755; 977/906 |
Current CPC
Class: |
Y10T 428/24975 20150115;
A61L 27/18 20130101; A61L 2400/12 20130101; A61L 27/56 20130101;
A61L 2300/416 20130101; A61L 2300/606 20130101; Y10T 428/24355
20150115; A61L 27/54 20130101; A61L 27/18 20130101; A61L 27/34
20130101; A61L 2300/41 20130101; C08L 71/02 20130101; A61L 27/18
20130101; C08L 83/04 20130101 |
Class at
Publication: |
424/443 ;
428/220; 428/216; 428/141; 977/755; 977/906 |
International
Class: |
A61K 9/70 20060101
A61K009/70; B32B 27/08 20060101 B32B027/08; B32B 7/02 20060101
B32B007/02; B32B 27/28 20060101 B32B027/28 |
Goverment Interests
[0002] This invention was made with government support under U54
AI065359 awarded by the National Institutes of Health. The
government has certain rights in the invention.
Claims
1. A membrane comprising: a) a first layer comprising parylene, and
b) a second layer comprising a copolymer selected from the group
consisting of polyethylene oxidepolymethylmethacrylate,
polymethyloxazoline-polydimethylsiloxiane-polymthoxazoline, and a
mixture thereof, wherein the membrane has a thickness between 0.1
nm and 20 nm.
2. The membrane of claim 1, wherein said membrane comprises a third
layer, wherein the third layer comprises parylene.
3. The membrane of claim 2, wherein said second layer is between
the first and second layers.
4. The membrane of claim 1, further comprising at least one
therapeutic agent.
5. The membrane of claim 4, wherein the at least one therapeutic
agent is within the second layer.
6. The membrane of claim 1, wherein the first layer comprises
nonoporous parylene.
7. The membrane of claim 2, wherein the third layer comprises rough
parylene C.
8. The membrane of claim 1, wherein the copolymer comprises a
single layer having a thickness from about 1 nm to about 10 nm.
Description
[0001] This application is a continuation of U.S. patent
application Ser. No. 12/135,640, filed Jun. 9, 2008, which claims
priority to U.S. Provisional patent application 60/942,885, filed
Jun. 8, 2007, and U.S. Provisional Application 60/981,688, filed
Oct. 22, 2007, each of which are herein incorporated by reference
in their entireties.
FIELD OF THE INVENTION
[0003] The present invention relates generally to the field of
coatings for medical and pharmaceutical applications. In
particular, the present invention relates to medical and
pharmaceutical products having biocompatible nanofilms comprising
therapeutic agents.
BACKGROUND
[0004] Implantable medical devices are becoming increasingly common
and more complex. Advances in medical device technology have led to
smaller and more complex implants that provide a greater standard
of living to an increasingly aging population. However, not all of
the materials used in medical implants are entirely biocompatible,
i.e. the surrounding tissues may become inflamed when in contact
with the surface of the implant. Inflammation results from the
infiltration of immune cells such as neutrophils and macrophages to
the tissue-implant interface as these cells attempt to repair
damage that occurs following implantation. Thus, inflammatory
responses against implants remains a problem with respect to both
tolerance and maintenance of function for a variety of these
implants, ranging from cardiovascular devices (e.g. coronary
stents) and electrical devices (e.g. pacemakers and glucose sensors
to prostheses such as hip joint replacements). Accordingly, there
is a need for biocompatible materials which can coat the implant
and serve as vehicles for the targeted delivery of drugs to the
surrounding tissues.
[0005] Existing delivery technologies include the use of
poly(lactic-co-glycolic acid) (PLGA) microspheres as drug-eluting
particles for suppression of inflammation. The use of fluidic
delivery of anti-inflammatories has also been explored. In
addition, poly(3,4-ethylenedioxythiophene (PEDOT) has been used in
nanotube and planar formats. Unfortunately, each of the delivery
systems provide a relatively thick material associated with an
implant. For example, the PLGA microspheres are on the order of 200
microns in diameter, while fluidic delivery materials are several
microns in thickness. As a result, these materials significantly
impact the design parameters of the implants they serve and
generate adverse effects on surrounding tissues because they
increase the overall dimensions of the implants and preclude
non-invasive behaviors.
[0006] In addition, all invasive biomedical devices inherently
possess the challenge of overcoming three primary obstacles. These
concerns are biocompatibility, bio-longevity, and efficacy.
Biomedical devices desire to leap these hurdles and be completely
biocompatible to reduce the risk of patient complications, but
often, current technologies result in new drawbacks. For example,
many types of drug-eluting stents serve to reduce clotting of the
stent in implant patients, but are associated with increased
clinical complication rates. Because recent advances do not
optimally address implant biocompatibility, newer, more effective
mechanisms are a necessity.
[0007] Pioneering drug delivery as a new field for parylene-based
applications fills the need for a biocompatible, functionalized
membrane capable of slow-releasing drug to a localized region for
targeted delivery. Currently, chemotherapeutic drugs are capable of
killing cancerous cells, but cannot selectively kill only cancerous
cells; they exhibit universal cytotoxicity. A similar issue exists
in administering anti-inflammatory compounds; indiscriminately
introducing these compounds significantly dilute drug efficacy and
weaken the global immune response, which opens a window for
infection. Therefore, it is imperative that drugs are delivered in
a targeted fashion.
[0008] Drug delivery is an important aspect of medicine as an
essential mechanism that bridges drug development and treatment.
During the past decade, there has been much attention focused on
improving control of drug delivery and the advent of newer
technologies including tissue scaffolds and drug-eluting stents is
evidence of the desire to have more control over how, where, and
when pharmaceutical agents are delivered. Nevertheless, it has been
a challenge to create a biocompatible coating capable of eluting
drug due to several developmental barriers. Such a material would
need to be biologically inert and stable, possess anti-inflammatory
mechanisms, and pliable for use in a wide variety of applications.
With these criteria in mind, the biomedical industry has sought to
create a biocompatible coating that does not interfere with device
operation, may abate inflammatory responses, and bolster efficacy
of the coated mechanism.
[0009] An example of such a search led to the creation of the
drug-eluting stent. The drug-eluting stent was created to improve
stent longevity and minimize clotting on the stent itself to reduce
the risk of neointimal hyperplasia, an excessive immune response to
bare metal stent implants that results in narrowing of vessels due
to clots, and thus medical complications for the patient. By
eluting immunosuppressive drugs from the stent, the inflammatory
response was lessened near the implant location of the stent, thus
inhibiting platelet activation and preventing neointimal
hyperplasia. The drug-eluting stent has revolutionized cardiology,
but many stents rely on less than ideal materials, or an unsuitable
combination of materials.
[0010] For example, while the FDA has concluded that drug-eluting
stents are safer and more effective than bare metal stents, certain
complications may result from using a drug-eluting stent including
severe thrombotic (formation of clot inside a blood vessel) events
and restenosis (abnormal narrowing of vessels), despite restenosis
and thrombosis being some of the very issues drug eluting stents
were created to solve. Therefore, while modern stents are a step
forward in reaching complete biocompatibility of implanted devices,
much of the present technology is not progressive enough, and the
search and necessity for a more biocompatible coating
continues.
SUMMARY OF THE INVENTION
[0011] The present invention relates generally to the field of
coatings for medical and pharmaceutical applications. In
particular, the present invention relates to medical and
pharmaceutical products having biocompatible nanofilms comprising
therapeutic agents.
[0012] The present invention relates to nanoscale copolymer thin
films associated with therapeutic agents to provide
copolymer-therapeutic agent complexes. The invention further
relates to the use of the thin films as coatings in medical
applications to provide localized therapeutic agent delivery.
Polyethleneoxide-polymethylmethacrylate (PEO-PMMA) diblock
copolymers and
polymethyloxazoline-polydimethylsiloxane-polymethyloxazoline
(PMOXA-PDMS-PMOXA) triblock copolymers are examples of copolymers
that may be used in the present thin films. The use of amphiphilic
diblock and triblock copolymers in the formation of multifunctional
nanofilms for the coating of implants is advantageous because these
copolymers facilitate the incorporation of a wide variety of
therapeutic agents into the coatings. In certain embodiments, the
copolymer materials of the present invention are combined with an
anti-inflammatory molecule. The present invention contemplates any
type of suitable anti-inflammatory molecule. Such molecules are
known in the art and can be easily located in the literature.
[0013] In addition to providing a mechanism for localized
therapeutic agent delivery, the present coatings, in certain
embodiments, facilitate chronic device functionality and prevent,
delay or minimize bio-fouling of implanted medical devices, such as
implants. In addition, the coatings may also be useful for
single-cell studies to examine polymeric activation of cellular
gene expression pathways for biotic-abiotic interfacing studies
(e.g. mechano-sensation, metabolism, etc.). The
copolymer-therapeutic agent complex coatings optionally may be used
in conjunction with vesicular therapeutic agent delivery structures
to provide replenishment of the therapeutic agents in the coatings.
In such an embodiment, copolymers may also be assembled into
vesicular structures that can serve as reservoirs for the
therapeutic agents. Surface functional groups on these vesicles
interact with the copolymer coatings on a medical implant to enable
intelligent, localized delivery of the therapeutic agents to
specific areas, including thin film coatings that require drug
replenishment. In addition, the vesicular structures provide the
ability to target and facilitate on-demand vesicular endocytosis to
prevent excessive dosing. As such, these materials may serve as a
modality for novel medical capabilities, as well as a platform for
fundamental cellular studies.
[0014] The present copolymer coatings may be produced using
Langmuir-Blodgett (LB) deposition. Details regarding the LB
deposition process are provided in Example 2, below. Using this
method, a self-assembled copolymer-therapeutic agent complex may be
fabricated by adding a therapeutic agent (e.g., a hydrophilic
agent) atop a pre-deposited copolymer layer on a Langmuir trough.
The use of the LB method is advantageous because it allows for the
deposition of very thin copolymer films (e.g., .ltoreq.5 nm) and
provides films with controlled molecular spacing that adsorb in a
very robust manner to a variety of underlying substrates. In
addition, the method can be used to deposit a spectrum of
therapeutic agents with a variety of hydrophilic/hydrophobic
properties that resist the generation of clumps or aggregates that
may adversely affect cell behavior or the biocompatibility of cells
that come into contact with the nanofilm coatings. Furthermore, the
LB method allows for a layer-by-layer deposition of
copolymer-therapeutic agent complex thin films, making it possible
to tune the coating thickness and therapeutic agent concentration
to control how much of the therapeutic agent is associated with an
implant surface. The present LB methods and the resulting LB
nanofilms are easily distinguished from thin films produced by the
Langmuir technique which creates a thin film over an opening or
aperture, rather than on an underlying substrate.
[0015] In certain embodiments, the present invention provides an
implantable medical device having one or more of its surfaces
coated with a nanofilm composition comprising a copolymer, wherein
the copolymer may be, for example, (i) a diblock copolymer
comprised of polyethylene oxide-polymethyl methacrylate or (ii) a
triblock copolymer comprised of
polymethyloxazoline-polydimethylsiloxiane-polymthoxazoline.
[0016] In one aspect, the present invention provides an implantable
medical device having one or more of its surfaces coated with an
nanofilm composition comprising: (a) at least one therapeutic agent
(e.g., wherein the therapeutic agent is not a protein); and (b) a
copolymer, wherein the copolymer may be, for example, (i) a diblock
copolymer comprised of polyethylene oxide-polymethyl methacrylate
or (ii) a triblock copolymer comprised of
polymethyloxazoline-polydimethylsiloxiane-polymthoxazoline.
[0017] The copolymers desirably, but not necessarily, comprise
monomer units having acrylate endgroups, which may facilitate
crosslinking and film stability. The nanofilm coatings of the
present invention may also comprise endgroups attached to a
receptor or a ligand, which may facilitate subsequent target
delivery of additional therapeutic agents, as described in more
detail below. A single layer of the copolymer in a nanofilm may be
designed to have a thickness from about 1 nm to about 10 nm (e.g.,
1, 2, 3, 4, 5, 6, 7, 8, 9, or 10 nm), desirably less than about 4
nm. However, the nanofilm may include multiple layers (e.g., from
about 2 to about 10 layers) of the copolymer-therapeutic agent
complexes, wherein each layer has a thickness from about 1 to about
10 nm (e.g., about 4 nm or less).
[0018] The therapeutic agent may be selected from the group
consisting of, for example: thrombin inhibitors, antithrombogenic
agents, thrombolytic agents, fibrinolytic agents, vasospasm
inhibitors, calcium channel blockers, vasodilators,
antihypertensive agents, antimicrobial agents, antibiotics,
inhibitors of surface glycoprotein receptors, antiplatelet agents,
antimitotics, microtubule inhibitors, anti-secretory agents, actin
inhibitors, remodeling inhibitors, antisense nucleotides, anti
metabolites, antiproliferatives, anticancer chemotherapeutic
agents, anti-inflammatory steroid or non-steroidal
anti-inflammatory agents, immunosuppressive agents, growth hormone
antagonists, growth factors, dopamine agonists, radiotherapeutic
agents, extracellular matrix components, inhibitors, free radical
scavengers, chelators, antioxidants, anti polymerases, antiviral
agents, photodynamic therapy agents, and gene therapy agents. In
one embodiment, the therapeutic agent is an anti-inflammatory
compound, e.g. Dexamethasone or an LXR agonist. In particular
embodiments, the LXR agonist is
3-((4-Methoxyphenyl)amino)-4-phenyl-1-(phenylmethyl)-1H-pyrrole-2,5-dione-
. In particular embodiments, the LXR agonist functions as a
non-steroidal anti-inflammatory to block transcriptional machinery
associated with cell stress among other disorders. In certain
embodiments, the therapeutic agent is one or more of the following:
sirtuin Activators, cytokines, interferons of all kinds (e.g.
alpha, beta, gamma, etc), as well as any other suitable therapeutic
molecule.
[0019] The nanofilm coatings of the present inventions may be used
on a variety of medical substrates, including any implantable
medical device. Such medical devices may be made of a variety of
biocompatible materials including, but not limited to, polymers and
metals. Medical substrates onto which the nanofilms may be coated
include, neural/cardiovascular/retinal implants, leads and stents,
and dental implants (e.g. nanofilms to seed bone growth). In some
embodiments, the nanofilm may be coated onto the electrode of an
implantable medical device. In fact, coating the present nanofilms
onto an electrode may provide an important medical advantage
because the copolymer films prevent or minimize bio-fouling which
often begins at the site of a metal electrode. In addition, unlike
more conventional implant coatings, the present nanofilms may be
made thin enough that they do not interfere with electrode function
(e.g., electrical conductivity or redox reactions at
electrodes).
[0020] In another aspect, the present invention provides a method
of delivering a therapeutic agent to a target site in a subject,
the method comprising: (1) coating an implantable device with a
nanofilm composition comprising: (a) at least one therapeutic agent
(e.g., wherein the therapeutic agent is not a protein); and (b) a
copolymer, wherein the copolymer may be, for example, (i) a diblock
copolymer comprised of polyethylene oxide-polymethyl methacrylate
or (ii) a triblock copolymer comprised of
polymethyloxazoline-polydimethylsiloxiane-polymthoxazoline; and (2)
implanting the device into the subject near the target site,
wherein elution of the therapeutic agent from the nanofilm delivers
the therapeutic agent to the target site.
[0021] This method may further comprise the step of administering
to the subject a vesicle, wherein the vesicle comprises (a) at
least one additional therapeutic agent, wherein the additional
therapeutic agent is the same or a different therapeutic agent as
used in the device; and (b) a copolymer, wherein the copolymer may
be, for example, (i) a diblock copolymer comprised of polyethylene
oxide-polymethyl methacrylate or (ii) a triblock copolymer
comprised of
polymethyloxazoline-polydimethylsiloxiane-polymthoxazoline; and
wherein the copolymer comprises monomer units having endgroups
covalently attached to a second receptor or ligand, wherein the
second receptor or ligand is capable to specifically binding to a
first receptor or ligand attached to endgroups on the nanofilm
coating. This additional step allows for the interaction of the
vesicle with the nanofilm for the purpose of releasing the
additional therapeutic agent into the nanofilm.
[0022] In some embodiments, the present invention provides
membranes comprising: a) a first layer comprising parylene, and b)
a second layer comprising a copolymer selected from the group
consisting of polyethylene oxidepolymethylmethacrylate,
polymethyloxazoline-polydimethylsiloxiane-polymthoxazoline, and a
mixture thereof In particular embodiments, the membrane comprises a
third layer, wherein the third layer comprises parylene. In other
embodiments, the second layer is between the first and second
layers. In other embodiments, the membranes further comprise at
least one therapeutic agent. In further embodiments, the at least
one therapeutic agent is within the second layer. In other
embodiments, the first layer comprises nonoporous parylene. In
particular embodiments, the third layer comprises rough parylene
C.
[0023] In some embodiments, the present invention provides a device
comprising an active parylene-encapsulated co-polymeric (APC)
membrane for slow release drug delivery. Such devices may be
implantable medical devices and comprise a base layer, which may
serve as a backbone of the membrane, an upper layer coated with a
nanofilm composition, that may act as a semi-permeable membrane;
and a copolymeric matrix capable of being conjugated with a
molecule, wherein the matrix comprises a network matrix of a
copolymer selected from the group consisting of di-block copolymers
such as polyethylene oxide-polymethyl methacrylate, tri-block
copolymers such as
polymethyloxazoline-polydimethylsiloxiane-polymthoxazoline, and
mixtures thereof, and a therapeutic agent, wherein the matrix is
conjugated with the therapeutic agent. Thus, in some embodiments, a
membrane is provided comprising a base layer having a first and a
second surface; a copolymeric layer having a first and a second
surface and comprising a copolymer selected from the group
consisting of polyethylene oxide-polymethyl methacrylate,
polymethyloxazoline-polydimethylsiloxiane-polymthoxazoline, and a
mixture thereof; a coating having a first and a second surface and
comprising a nanoporous, parylene film; and a therapeutic agent;
wherein the copolymeric layer is conjugated with the therapeutic
agent; the first surface of the base layer is in contact with the
second surface of the copolymeric layer and the second surface of
the coating is in contact with the first surface of the copolymeric
layer, such that a sandwich structure is achieved.
[0024] In another aspect, parylene, an FDA approved biologically
inert nanomaterial capable of being deposited uniformly on
virtually any surface, may be used to prepare an active
parylene-encapsulated copolymer (APC) membrane that can be applied
to implants, sensors, stents, and a wide array of other invasive
biomedical devices. This novel membrane technology can serve as
both a coating based modality or a stand-alone device for targeted
therapeutic delivery, demonstrating its versatile range of
medically-significant functionalities. APC membranes are capable of
being functionalized with a diverse selection of compounds for
specialized treatment including, but not limited to cancer,
inflammation, and anti-viral therapies Furthermore, APC membranes
have slow-release capabilities that enhance the bio-longevity of a
device by several days, as revealed through RT-PCR gene expression
studies. In addition, the slow release mechanism targets drug
elution to a localized area near the membrane, thereby augmenting
the efficacy of the drug by concentrating the effective dosage to
the targeted region and limiting drug that is haphazardly flushed
through the patient's entire system. This method of drug delivery
offers significant benefits especially for chemotherapy and
anti-inflammation therapies that have serious consequences when
drugs are arbitrarily carried through the patient's system.
[0025] In yet another aspect, an APC membrane is described. Such
APC membranes are parylene-based functionalized drug delivery
membranes capable of targeted and slow-release drug elution
capabilities, and also having the flexibility to be tailored to any
surface as well as to many therapeutic applications. The APC
membrane fulfills the need for localized delivery through its
flexibility and slow-release mechanism by acting as a reservoir for
a spectrum of therapeutic compounds. Because the nanopore layer of
the APC membrane is capable of trapping the drug and eluting it
slowly through the membrane as it comes into contact with fluids,
the effective dose of the drug will be applied in a very controlled
and precise fashion that contains the drug concentration near the
surface of the membrane. This is ideally applicable for the patient
because the drug effect is localized and the effective dose is
utilized in its entirety. The prime level of specificity and
biocompatibility that has been effectively engineered into the APC
membrane and will serve as a flexible, functionalized platform
capable of creating a controlled window where neither infection nor
implant rejection occurs, and localizing drug activity to cancerous
regions that will enhance chemotherapy potency and ultimately,
benefit the patient's life and well being.
[0026] In a related aspect, stable packagings of the tri-block
copolymeric nanofilm, are provided, so that device technologies can
be developed while using the copolymeric membrane as a foundational
drug elution matrix. As parylene is coated on a device through room
temperature vapor deposition of individual molecules, the resulting
coating is extremely uniform, and conforms to practically any
surface shape. This property is important for many biomedical
devices that require isolation from the body to preserve the
function of the device. Furthermore, the middle layer of the APC
membrane is a network matrix of tri-block copolymer that may be
conjugated with a very broad spectrum of drug molecules.
BRIEF DESCRIPTION OF THE DRAWINGS
[0027] FIG. 1. The process for copolymer-mediated Langmuir-Blodgett
deposition is shown here. 1) The copolymer was deposited at the
air-water interface to a starting pressure of 10-15 mN/m.sup.2). 2)
The functional molecule (e.g. SWNT or Dex) was added to the surface
of the pre-formed membrane in a dropwise fashion to minimize
copolymer film perturbation. 3) The film was compressed at a rate
of 1 mm/min to fabricate a uniform, tightly packed composite
material that could be tethered to the substrate via the copolymer.
4) The deposition process onto the substrate was carried out at a
rate of 1 mm/min to ensure complete film formation.
[0028] FIG. 2A. Tapping mode atomic force microscopy was utilized
for LB film imagery. An AFM micrograph is shown here with a scan
size of 2 .mu.m at a frequency of 0.7016 Hz. The copolymer nanofilm
exhibited a very robust adhesion to the glass substrate. FIG. 2B. A
scan of a SWNT-Copolymer composite film is shown here with a scan
size of 0.4922 .mu.m with a scan rate of 1.001 Hz. The presence of
SWNT's integrated with the copolymer molecules was also confirmed
via Langmuir isotherm analysis as well as cyclic voltammetric
measurement of cytochrome c-mediated oxidation-reduction reactions
(triblock).
[0029] FIG. 3A. Langmuir isotherm is shown here comparing
measurements for polymer-only films with SWNT-Copolymer films. It
can be seen that composite films exhibit a higher surface pressure
versus trough area indicating the presence of SWNT's at the
air-water interface. FIG. 3B. Cyclic Voltammetry further confirmed
the functional presence of SWNT's as well as the anti-adsorption
abilities of the copolymer via measurement of cytochrome c-mediated
electron transfer to an underlying gold electrode which served as
the substrate for composite film deposition (triblock).
[0030] FIG. 4A. FITC-Dex complexes were deposited atop PEO-PMMA
copolymers and fluorescence imagery was performed to examine Dex
transfer onto the substrate. Amphiphile-assisted Dex deposition
resulted in the transfer of the drug to the substrate while
deposition of Dex alone did not result in its transfer. FIG. 4B.
Application of LPS to the Raw 264.7 macrophages resulted in the
induction of p65-mediated inflammatory gene programs which were
targeted by the Dex-Copolymer composites to attenuate cell stress
(diblock).
[0031] FIG. 5A. Using TNF.alpha. as an example of a cytokine that
was expressed as a result of LPS induction, it could be seen that
cells cultured atop the composite films showed significant
reductions in inflammation compared to cells cultured atop the
polymer and stimulated with LPS. FIG. 5B. Incubation of the Raw
264.7 macrophage cell line with a 0.1 g/ml of copolymer in solution
did not affect cell growth behavior nor did the incubation affect
cell growth rates. Images were taken at 4 hrs, 24 hrs, and 48 hrs.
This observation was an indication that the copolymer was
biologically inert with no qualitative adverse effects upon
cellular behavior.
[0032] FIG. 6. Water-soluble Dex deposited atop the copolymer was
suspended at the surface of the Langmuir trough and integrated with
the copolymer (circle) while Dex was shown to submerse into the
subphase when deposited without the copolymer (square).
[0033] FIG. 7 RT-PCR examination TNF-.alpha., IP-10, I1-6, and
I1-12 mRNA show major reductions in inflammation based on the
number of DEX- films deposited. These films may be useful towards
eliminating initial onset of cellular inflammation that may trigger
device fouling.
[0034] FIG. 8. Dex-functionalized films show an early complete
suppression of TNF production from RAW cells.
[0035] FIG. 9. For samples where Dex was self-assembled onto the
glass substrate without the polymer, it could be seen that stress
attenuation was precluded. This showed that the copolymer enabled
both interfacial Dex deposition as well as the ability to tether
the Dex to the substrate and sustain the interface between the
macrophages and anti-inflammation capabilities of the composite
material.
[0036] FIG. 10 illustrates drug elution from a nanofilm into
macrophages. Immobilized dexamethasone was shown to exhibit
internalization capabilities following in vitro cell culture.
Dexamethasone is known to enter into a binding relationship with
the nuclear glucocorticoid receptor whereby Dex-GR complex binding
with p65 interferes with p65 interaction with transcriptional
machinery for inflammatory cytokines Evidence also shows that p65
interaction with co-activators is also blocked via Dex-GR
binding.
[0037] FIG. 11. Cross-section view of an APC membrane. The base
layer is a layer of parylene that serves as the backbone of the
membrane. The top layer is a layer of ultra-thin parylene with
nanopores that acts as a semi-permeable membrane. Sandwiched
between the parylene layers is a tri-block copolymer matrix that
can be rapidly functionalized with molecules. The drug used in one
embodiment was dexamethasone (Dex), which was incorporated into the
tri-block copolymer matrix and deposited onto the parylene base
layer with a Langmuir-Blodgett trough. The figure shows drug
eluting through the nanopore layer by diffusing across the
membrane. The size of the nanopores may be engineered to control
the rate of drug elution. Multiple layers of the Dex/tri-block
copolymer matrix may be applied at the cost of only a few
nanometers to extend the bio-longevity of the membrane for many
more hours or even days.
[0038] FIG. 12. AFM images of APC membrane surfaces (top) and
parylene only surfaces (bottom) at two different nanopore layer
thicknesses. Depressions are shown on the surface of the membranes
as darker areas suggesting possible nanopore formation. The
geography of parylene exhibits rough terrain capable of acting as
an inflammatory stimuli.
[0039] FIG. 13A. Cellular expression levels for iNOS were shown to
decrease when RAW264.7 cells were cultured atop parylene copolymer
composite surfaces versus being cultured atop bare nanostructmed
parylene. FIG. 13B. Cellular expression levels for TNFa were shown
to decrease when RAW264.7 cells were cultured atop
parylene-copolymer composite surfaces versus being cultured atop
bare nanostructured parylene. These results indicate that murine
macrophages can possibly sense their substrate stiffness to elicit
innate inflammatory responses as polymer-coated parylene may
present a more biomimetic interface with the cells. AFM images
indicated that a bare parylene surface contained
micro/nano-structured features that could serve as inflammatory
stimuli.
[0040] FIG. 14. Incubation of RAW 264.7 macrophages on glass,
parylene, and APC membrane surfaces display differences in growth
patterns. LPS stimulation occurred for 4 hours beginning after 20
hours of growth. Images were taken 8 hrs, 20 hrs, and 24 hrs.
Qualitatively, cell growth was less confluent on the parylene and
APC surfaces compared to glass. This observation suggests that
growth and activity of macrophages are inhibited on the Parylene
and APC surfaces. Closer inspection reveals that growth was almost
completely halted on the APC membrane, while macrophages continued
to grow on the parylene surface. FIG. 14 shows that parylene and
the APC membrane have comparable cell adhesion properties; RT-PCR
results reveal contrasting levels of inflammatory cytokine
expressed.
[0041] FIG. 15A. RT-PCR relative gene expression data for unsoaked
slides confirm the presence of dexamethasone on the functionalized
slides. Based on past studies, IL-6 expression values were most
consistent and reliable possibly due to IL-6 occurring as a
critical cytokinetic component for the immune response in murine
macrophages. The expression of IL-6 is at least relatively two
times higher for the parylene and glass slides compared to an APC
membrane. This shows that Dex can successfully survive vapor
deposition and soaking, and still be an effective inflammatory
cytokine inhibitor. FIG. 15B. The Same pattern is shown in relative
TNF-a expression levels. FIG. 15C. While this is not apparent in
the relative iNOS expression levels, LPS induced inflammation may
not have had a significant effect on iNOS levels. When taking the
error bars into account, the levels of iNOS are comparable for the
three surfaces.
[0042] FIG. 16 RT-PCR relative gene expression data for soaked
slides confirm the presence of dexamethasone on the functionalized
APC membranes. Slide A3 is a functionalized slide with a parylene
base and layers of Dex incorporated into a tri-block copolymer
matrix, but lacks the final nanopore layer. The gene expression of
IL-6 is higher for slide A3 compared to the APC membrane that has
the nanopore layer engineered for slow elution capabilities. These
data suggest that there is an advantage to slow elution owing to
localization of drug elution to near the surface of the membrane.
The APC membrane inhibits inflammatory cytokines, up to 18 times
more effectively for IL-6 than does plain parylene on glass. The
bare terrain of parylene was rough and littered with
micro/nano-structured features that could serve as inflammatory
stimuli.
[0043] FIG. 17 shows histological analysis of Dex-copolymer
nanofilm implantation as described in Example 10. Untreated tissue
(17A), as well as tissue containing implanted uncoated disks (17B),
as well as Dex-copolymer nanofilm coated disks (17C) were stained
to examine cellular recruitment to the implant surface, which is a
commonly observed mechanism of foreign body formation and eventual
implant fouling.
[0044] FIG. 18 shows that biologically inert polymer in solution
has no effect upon cellular proliferation. Triblock copolymer in
0.01 mg/ml and 0.1 mg/ml concentrations were incubated with
cultured macrophages. Samples were imaged using brightfield
microscopy at 4 hrs, 24 hrs, and 48 hrs. Results show from in vitro
standpoint that cellular proliferation is unaffected, confirming
bio-inert properties of nanofilm.
[0045] FIG. 19 shows that biologically inert polymer in solution
has no effect upon cellular inflammatory gene programs. Triblock
copolymer in 0.01 mg/ml and 0.1 mg/ml concentration were incubated
with cultured macrophages. Samples were lysed and RT-PCR was
performed to examine the release of inflammatory
cytokines/molecules. Results show negligible change compared with
control readings (-) as changes exhibited changes of less than
2-fold.
DETAILED DESCRIPTION
[0046] Disclosed herein are functionalized copolymers that provide
for tailored biology, where specific functionalities within these
materials can be rapidly engineered. Embedding effector molecules
into these copolymers transforms an otherwise inactive component
into a functional matrix for in vivo applications, including, for
example, suppression of cellular inflammation/stress or
immuno-regulatory implant coating. Furthermore, the ability to
utilize these copolymers in both a vesicular or planar
configuration, coupled with their versatility in block composition,
allows for dedicated drug delivery with targeting capabilities, or
modalities to replenish exhausted effector molecule stores within
planar copolymer thin films. As such, these copolymeric membrane
materials impact multiple medically-relevant fields in both
scientific and technological contexts.
[0047] The present inventors have discovered that the materials
described herein have bio-inert properties when applied as a thin
coating to implantable medical devices. This includes preventing
any stress response of macrophages to the chemical/topographical
stimuli presented by these materials. In addition to the
biocompatibility of the polymers themselves, the present inventors
have discovered that various therapeutic agents can be incorporated
into the copolymer in order to deliver a therapeutic agent to a
tissue.
[0048] In one aspect, the present invention provides a nanofilm
coating for an implantable medical device (e.g., a stent, or
under-skin device, or catheter, or surgical instrument, or
implantable device such as a pacemaker, etc.). The term "coating"
as used herein, will refer to one or more vehicles (e.g., a system
of solutions, mixtures, emulsions, dispersions, blends etc.) used
to effectively coat a surface with therapeutic agent and a
copolymer component, either individually or in any suitable
combination. The present invention further provides a method for
using the nanofilm coating to coat a surface with a therapeutic
agent, for instance to coat the surface of an implantable medical
device in a manner that permits the surface to release the
therapeutic agent over time when implanted in vivo.
[0049] The coatings include a copolymer complexed with a
therapeutic agent in the form of a nanofilm that may be constructed
to extremely thin dimensions. In one embodiment, a the thickness of
the nanofilms may be from about 0.1 nm to about 20 nm, preferably
about 4 nm or less. As such, the coatings can significantly
suppress cellular inflammation while possessing dimensions that do
not impact the implant with which the coatings are interfaced.
These dimensions also enable interfacing with electrodes, or
implant leads while still enabling electrode function specifically
because the material is so thin and does not preclude electron
transport. For example, the present nanofilm coatings may be
applied to gold electrodes (e.g., leads) while still supporting
oxidation-reduction reactions to occur at the electrode
surface.
[0050] In some embodiments, multiple layers of the copolymers are
used to generate the coatings. Nanofilms can be sequentially
deposited to form multilayers to `tune` the amount of therapeutic
agent that has been added to an implant surface. Due to the thin
dimensions of each layer in the coating, even coatings made
multiple copolymer layers may have a negligible impact upon the
device dimensions because each layer may be very thin (e.g., only
about 4 nm thick). As such, even a coating that includes 10
copolymer layers, for example, would be orders of magnitude thinner
in dimension than a PLGA microsphere.
[0051] The coatings may be fabricated/deposited using the
Langmuir-Blodgett method which is rapid, low-cost, and accomplished
in a parallel fashion with multiple substrates being deposited
simultaneously. Furthermore, the materials and deposition
modalities employed by this technology enables film deposition on
any type of surface or form factor, making this an exceptionally
versatile technology that is easily adaptable to multiple
applications.
[0052] The copolymers in the coating may be di- to tri-block
copolymers. Block copolymers have been shown to be effective
matrices to support protein function for the mimicry of key natural
biological processes such as energy conversion, as well as
voltage-gated ion transport [1-7]. While conventional lipid-based
systems have enabled single protein characterization and mechanisms
of functionality to be elucidated [8-11], block copolymers
represent a highly versatile approach towards nanoscale/engineered
medicine, whereby specific properties can be engineered into the
material to accommodate specific protein geometries, or desired
block lengths, compositions, and charge properties, to name a few.
Furthermore, the addition of endgroups, such as acrylate, can be
made to enable UV-induced polymeric crosslinking with reported
steric contraction of the film to enhance material stability for
enhanced device robustness over conventional lipid systems [12,
13]. The copolymeric materials are extremely robust, and can
successfully coat a substrate while left in the ambient environment
for months. Furthermore, the copolymers are amenable to the
demonstrated drug/effector molecule/protein elution studies because
they can be cross-linked using both chemical and UV-induced
methods. Several polymers/lipids exhibit rapid breakdown when
exposed to UV light where interfaced molecules also exhibit
breakdown upon degradation of the supporting matrix. In this case,
UV crosslinking actually serves as a beneficial condition to
further enhance material stability.
[0053] In some embodiments, the copolymer is a triblock copolymer
and possesses the structure of
polymethyloxazoline-polydimethylsiloxane-polymethyloxazoline
(PMOXA-PDMS-PMOXA). As noted above, the end groups may be
terminated with acrylate, which enables rapid crosslinking to
enhance material stability. As previously mentioned, these
endgroups can be rapidly functionalized with a spectrum of
molecules including membrane proteins (e.g. Bacteriorhodopsin,
Cytochrome C Oxidase, etc., as well as effector molecules that
suppress cellular processes such as inflammation. Furthermore, the
PMOXA endgroup is biologically-inert, meaning it does not freely
enter into interactions (e.g., chemical) with its surrounding
biological environment. Thus, this material is resistant to
cellular adhesion.
[0054] The copolymers may optionally be functionalized at the
endgroups to enable directed targeting which will enable rapid
replenishment of therapeutic agent-depleted nanofilms. More
specifically, nanofilm vesicles, or hollow spheres that are
carrying a specific drug can have their outside shell
functionalized with a receptor, or ligand, that can directly target
a planar-deposited film atop an implant that has already eluted the
complete stores of the integrated drug. These vesicles can then
unravel to restore additional drug into the planar film while
providing a very minimal impact upon nanofilm thickness because the
material is already inherently extremely thin. As such, this highly
versatile material possesses significant improvements over existing
films given its several improvements in possessing non-invasive
dimensions, interfacing capabilities with a spectrum of molecules,
as well as several other key advantages.
[0055] Therapeutic agents are incorporated into or complexed with
the copolymer such that a therapeutically effective amount of the
agent may be delivered to the target site upon implantation of the
medical device. A "therapeutically effective amount" of a compound
refers to an amount of the compound that alleviates, in whole or in
part, symptoms associated with a disorder or disease, or slows or
halts further progression or worsening of those symptoms, or
prevents or provides prophylaxis for the disease or disorder in a
subject at risk for developing the disease or disorder.
[0056] The copolymers serve as a platform system, meaning they can
be interfaced with a collection of molecules and substances for a
wide-array of applications. Virtually any therapeutically active
agent for which localized delivery of is desired may be associated
with (e.g., complexed with) the copolymer nanofilms in accordance
with the present invention. The term "therapeutically active agent"
is intended to encompass any substance that will produce a
physiological response when administered to a host. Because the
triblock copolymer itself has alternating hydrophilic and
hydrophobic groups within its structure, both hydrophilic and
hydrophobic drugs can be integrated into the copolymer films.
Furthermore, crosslinkable endgroups enhance materials robustness
and may play a major role in regulating elution rate, as there is a
steric hindrance of the film structure which will `tighten` up the
copolymer to reduce drug release.
[0057] In general, the term therapeutically active agent includes
therapeutic or prophylactic agents in all major
therapeutic/prophylactic areas of medicine as well as nutrients,
cofactors, and xenobiotics. Suitable substances include, but are
not restricted to, antifungals such as amphotericin B,
griseofulvin, miconazole, ketoconazole, tioconazole, itraconazole,
and fluconazole; antibacterials such as penicillins,
cephalosporins, tetracyclines, aminoglycosides, erythromicin,
gentamicins, polymyxin B; anti-cancer agents such as
5-fluorouracil, bleomycin, methotrexate, hydroxyurea;
antiinflammatories such as glucocorticoids, including
dexamethasone, hydrocortisone, colchicine; nonsteroidal
antiinflammatory agents including ibuprofen, indomethacin, and
piroxicam; antioxidants, such as tocopherols, carotenoids, metal
chelators, ubiquinones, or phytate; antihypertensive agents such as
prazosin, verapamil, nifedipine, and diltiazem; analgesics such as
acetaminophen and aspirin; antiviral agents such as acyclovir,
ribavarin, and trifluorothyridine; antiandrogens such as
spironolactone; androgens such as testosterone; estrogens such as
estradiol; progestins such as modified progestogens; opiates;
muscle relaxants such as papaverine; vasodilators such as
nitroglycerin; antihistamines such as cyproheptadine; antitussives
such as dextromethorphan; neuroleptics such as clozaril;
antiarrhythmics; antiepileptics; proteins, polypeptides,
neuropeptides such as somatostatin, substance P, vasoactive
intestinal peptide (VIP), calcitonin-gene related peptide (CGRP),
capsaicin, insulin, and gastrin; and protein enzymes, such as
superoxide dismutase or neuroenkephalinase or psychotropics
including penothiazines and tricyclics, carbohydrates,
glycoproteins, glycolipids, other lipids and cytokines. Cytokines
include tumor necrosis factors, the interleukins, growth factors,
colony stimulating factors, and interferons. Other useful drugs, in
approved commercially available formulations, and their recommended
dosages are listed in the annual publication of the Physicians'
Desk Reference, published by Medical Economics Company, a division
of Litton Industries, Inc.
[0058] In an exemplary embodiment, the coatings of the present
invention can be interfaced with multiple types of therapeutic
agents or drugs including Dexamethasone, which is a glucocorticoid
anti-inflammatory molecule that interacts with the glucocorticoid
receptor (GR) in the cellular cytoplasm, which then causes the GR
to translocate into the nucleus and block the production of
cytokines (e.g.. indicators of adverse cellular conditions) such as
tumor necrosis factor-alpha (TNF-.alpha.), interleukin-6 (IL-6),
interleukin-12 (IL-12), 1P-10, as well as inducible Nitric Oxide
Synthase (NOS). Nanofilm materials of the present invention which
incorporate Dexamethasone significantly impair the production of
these inflammatory molecules. This embodiment of the invention is
described in more detail in Example 6, below.
[0059] In a second exemplary embodiment, Liver-X-Receptor (LXR)
agonists are interfaced with the nanofilm materials of the present
invention. LXR agonists are potent inhibitors of LPS-mediated
inflammatory responses. Other embodiments include the combination
of the polymer with cancer-suppressing as well as bone-growth
inducing molecules for their respective applications in cancer
therapeutics, as well as seeding bone growth for dental and
orthopedic implants.
[0060] Recent studies have been done on natural polymers in the
field of tissue engineering with hopes to mimic nature [19B].
However, these natural polymers alone cannot be incorporated as a
coating because they are often only a few nanometers thick. In the
past, the triblock copolymer used in the conventional APC
membranes, though effective as a nano scale drug molecule
reservoir, was very difficult to manually handle and therefore, was
challenging to use in a clinical setting as a stand-along device or
patch and is generally applicable as a drug eluting coating.
Inability to manually handle tri-block copolymer in the
conventional APC membranes were addressed by incorporating
parylene, an extremely conformal FDA approved bio-inert polymer, as
a backbone for the drug/copolymer layer [20B, 21B].
[0061] In another exemplary embodiment, Dex was successfully
conjugated to the copolymer matrix; the presence of Dex was
confirmed through extensive gene expression studies that illustrate
the inflammatory response was significantly abated in macrophages
cultured on the APC membrane versus plain glass and plain parylene
(Example 8, below) therefore, the APC membrane can serve as an
effective drug carrying interface for the purpose of coating
implants, or directed drug delivery. In addition, this embodiment
demonstrates tunable slow release capabilities of the APC membrane
by incorporating an ultra-thin parylene layer that contains
nanopores. The presence of these nanopores has been confirmed
through AFM images and slow-elution properties were apparent from
gene expression studies (Example 8, below). The nanopore parylene
layel is important to the APC membrane because it acts as a
semi-permeable membrane through which pharmaceutical agents may
slowly diffuse. Without the porous layer, drug delivery would not
be localized and have absolutely no long-term value to the patient
due to immediate exposure and dilution of the drug through the
patient's whole system; such a result would have severe medical
complications especially with anti-inflammatory and
chemotherapeutic agents that are universally cytotoxic [22B]. The
importance of localized drug delivery is unparalleled in overcoming
numerous ailments including cancer, infections, and aggressive
immune responses. Several criteria that should be considered when
constructing a novel coating include biocompatibility, preserved
efficacy, and longevity. This strategy enables the APC membrane to
satisfy these criteria through focusing on using biocompatible
materials such as parylene, which has an extended shelf-life
lasting many years due to its enhanced biostability [23B].
[0062] In the related aspect, the membrane has been engineered with
slow-elution capabilities to maximize exposure to the effective
dose for the maximum amount of time. Most importantly, these
properties have combined into one single entity to serve as a
flexible platform for drug delivery capable of diverse application
as a pharmaceutical agent transplant, device coating, drug-eluting
patch, and tissue regeneration scaffold among others. Therefore,
the APC membrane disclosed herein satisfies the criteria necessary
for a biomedical membranes, and offers a significant advantage over
contemporary methods of drug delivery and modern biomedical
coatings due to its ability to slow-release a plethora of
pharmaceutical agents, localize drug delivery, and conform to
virtually any device surface giving it even greater potential to be
incorporated into future biomedical devices.
[0063] Experiments conducted during the development of the present
invention also showed that nanofilms of the present invention were
non-toxic. In particular, in a mouse liver toxicity model injected
with high concentrations of nanofilm showed favorable, non-toxic
results.
[0064] In certain embodiments, the compositions of the present
invention are used as a nutrient delivery wrap for plants. In other
embodiments, the compositions of the present invention are used to
treat hepatitis by allowing sustained release of interferon (e.g.,
in a subdermal/subcutaneous configuration). In some embodiments,
the compositions of the present invention are used on the skin for
tattoo removal. In further embodiments, the compositions of the
present invention are used for space medicine (e.g., transdermal
release of nutrients). In particular embodiments, the compositions
of the present invention are employed for veterinary purposes, such
as treating cancer and infection by helping to reduce toxicity and
enhance efficacy.
[0065] In certain embodiments, commercial applications include the
application of the nanofilm technology as cardiovascular and neural
implant coatings, lead and stent coatings, dental implant coatings,
as well as orthopedic implant coatings. The nanofilm compositions
of the present invention allow interfacing of the
anti-inflammatory/anti-cell adhesion nanofilm with implant surfaces
due to its demonstration as an effective suppressor of macrophage
adhesion/recruitment in vivo. This ability plays a key role towards
the enhancement of implant chronicity in functionality, or how long
the implant is able to function while resisting impairment by
macrophage aggregation. In addition, in certain embodiments, the
nanofilm can be interfaced with leads and stents because it can
still enable normal electrode function due to its very thin
dimensions. As such, this nanofilm serves as a highly efficient
modality for the localized suppression of cellular inflammation
which in an applications context can combat the implant fouling
process. The highly versatile and robust nature of the polymer also
allows for vesicular targeting to replenish polymeric films that
have completed the drug elution process. This very robust material
can also be applied towards the coating of dental implants as
current technology results in the imminent breakdown of the implant
due to fouling of the implant-bone interface. This polymer could be
integrated with bone growth-promoting molecules (e.g. bone
morphogenic proteins) to form highly robust interfaces to promote
bone growth. Because block copolymers are already used as
ingredients in toothpastes, precedent already exists with respects
to the FDA approval process for this material. Furthermore, a
substantial amount of investment has been provided by private
corporations, as well as the federal government across all major
agencies (e.g. NIH, NSF, DoD, etc.) to develop the next generation
of implant technologies for neuromedicine, cardiomedicine, and
wireless health monitoring. At the interface of this emerging class
of implantables, and chronic functionality within the human body
will be the nanofilm material of the present invention which can be
integrated with all of these technologies and will hence, serve as
a gateway material to empower this new class of implants to
function with unparalleled fouling resistance capabilities.
[0066] For the purposes of this disclosure and unless otherwise
specified, "a" or "an" means "one or more"
[0067] One skilled in the art will readily realize that all ranges
discussed can and do necessarily also describe all subranges
therein for all purposes and that all such subranges also form part
and parcel of this invention. Any listed range can be easily
recognized as sufficiently describing and enabling the same range
being broken down into at least equal halves, thirds, quarters,
fifths, tenths, etc. As a non-limiting example, each range
discussed herein can be readily broken down into a lower third,
middle third and upper third, etc
[0068] All publications, patent applications, issued patents, and
other documents referred to in the present disclosure are herein
incorporated by reference as if each individual publication, patent
application, issued patent, or other document were specifically and
individually indicated to be incorporated by reference in its
entirety. Definitions that are contained in text incorporated by
reference are excluded to the extent that they contradict
definitions in this disclosure.
EXAMPLES
[0069] The present invention is further illustrated by the
following examples, which should not be construed as limiting in
any way.
Example 1
Polymer Preparation and Characterization
[0070] The PEO-PMMA and PMOXA-PDMS-PMOXA copolymers were
solubilized to 0.1 mg/ml in chloroform and stirred overnight to
result in a translucent, homogeneous solution for LB deposition.
The thickness of the materials may be characterized, using a
PMOXA-PDMS-PMOXA triblock structure as an example, by performing a
secondary treatment of the chloroform-solubilized copolymer with
toluene and suspending a droplet across a 25 .mu.m-thick
hydrophobic septum flanked by 2 Ag/AgCl electrodes for capacitance
measurements as given by: C=.di-elect cons.o.di-elect cons.1A/d,
where C is the membrane capacitance, .di-elect cons.0 is the
permittivity of free space (=8.9.times.10-19 F/m), .di-elect cons.1
is the relative dielectric constant of the PDMS hydrophobic block,
A is the measured area of the annulus in the septum (=250 .mu.m),
and d is the thickness of the membrane.
[0071] The resultant thickness of the materials were determined to
range between 3-4 nm [14-16] for the triblock structure. As such,
we expected the diblock copolymer to possess a thickness with an
upper limit of .about.3 nm.
Example 2
General Langmuir-Blodgett Film Fabrication Protocol for Polymeric
Substrates
[0072] LB films were fabricated using a KSV 2000 Standard Langmuir
Trough with a Teflon.RTM. base and a subphase of water. To preclude
sample contamination from ambient particles, the entire trough was
covered with a plastic case and a small door was integrated to
allow for manual manipulation/cleaning of the trough. The base was
cleaned with chloroform using a cotton swab and large tweezers and
then rinsed thoroughly with nanopure water. The water was then
swept using cotton swabs into the central reservoir of the trough
and suctioned off using a vacuum pump. This step was performed
three times to ensure trough cleanliness. The trough was then
filled with nanopure water while paying careful attention not to
deposit water droplets along the edge of the trough. The Wilhemy
platinum pressure sensing plate (stored in MeOH was then thoroughly
rinsed using nanopure water, and subsequently sterilized using a
torch. The pressure sensor was then zeroed and the ready for film
deposition (FIG. 1).
Example 3
Single-Walled Carbon Nanotube Preparation/Deposition for
Bioelectricity Measurements
[0073] Single walled carbon nanotubes (SWNT) (Sigma-Aldrich, Inc.
St. Louis, Mo., USA) were solubilized to 1 mg/ml in chloroform with
overnight stirring. Following the addition of PEO-PMMA or
PMOXA-PDMS-PMOXA polymer solutions onto the Langmuir trough to a
starting pressure of 15-20 mN/m, the SWNT's were then added in a
dropwise fashion to minimize Langmuir film perturbation. Changes in
surface pressure as a result of SWNT addition were noted and
following 30 minutes to allow for the film to reach equilibrium,
compressions were performed at a rate of 1 mm/min to maximum
pressures of 30-40 mN/m for LB deposition onto gold-coated glass
slides at a rate of 1 mm/min. (VWR Scientific, Inc.) (Films were
compressed to >50 mN/m until collapse for Langmuir film
characterization of film properties). Gold coated slides were
utilized in this case as the SWNT films were used to enhance
electron harvesting capabilities for the gold working electrode
measurement of cytochrome c-mediated oxidation-reduction. For the
SWNT-copolymer experiments, variations to the film deposition as
well as preparation methods were also employed given that the
SWNT's were also soluble in chloroform, and for the purposes of
cytochrome c activity measurement the PEO-PMMA/PMOXA-PDMS-PMOXA
copolymers also possessed anti-protein adsorption capabilities to
facilitate electron transfer. A down-dip method was also employed
where the substrate was lowered into the trough such that the
floating Langmuir film facing the air was deposited directly onto
the substrate, while the side facing the subphase became the top
layer after deposition onto the substrate. Chloroform-solubilized
SWNT/Copolymer solutions were also deposited at the air-water
interface. For oxidation-reduction measurement, 30 mg of cytochrome
c (Horse heart muscle, Sigma-Aldrich, Inc. St. Louis, Mo., USA) was
dissolved in 2462.5 .mu.l of water (Nanopure). 37.5 .mu.l of a 1
mg/ml solution of ferricyanide (Sigma-Aldrich, Inc.)
[0074] was then added to the cytochrome c solution. This composite
solution was then added to a Desalting column (Amersham
Biosciences) to complete the process to produce CytC3+. The
oxidized CytC was then concentrated down in a swing bucket rotor at
4000 g and 4.degree. C. to a final volume of 250-300 .mu.l.
[0075] Cyclic voltammetry was performed using an electrochemical
workstation (Solartron, Inc.). Platinum, gold, and a saturated
Ag/AgCl (205 mV vs SHE) were used as the counter, working, and
reference electrodes, respectively. A buffer of 20 mM Mops, pH 7.0,
50 mM Na2SO4, 50 mM K2SO4, 2.5 mM MgSO4, 0.2 mM EDTA was prepared
for CV experiments. Experiments were performed at a scan rate of 30
mV/s between -0.18V and 0.5V versus open circuit.
Example 4
Glucocorticoid Preparation/Deposition for Inflammation Attenuation
Studies
[0076] Water soluble dexamethasone (Sigma-Aldrich, Inc.) was
dissolved in nanopure water to a concentration of 1 mg/ml. The drug
was then added to an interfacial pre-formed 10 mN/m copolymer film
and changes in surface pressure were monitored to confirm
dexamethasone presence at the air-water interface. After 30 minutes
of allowing the film to reach equilibrium, compressions were also
performed at a rate of 1 mm/min (FIG. 6) to a maximum pressure of
30 mN/m for LB deposition onto glass slides (25 mm.times.75 mm) at
a rate of 1 mm/min. (VWR Scientific, Inc.) (Films were also
compressed to >50 mN/m until collapse for Langmuir film
characterization of film properties). The slides were then used for
Raw 264.7 murine macrophage culture and quantitative Polymerase
Chain Reaction (qPCR) analysis.
[0077] Raw 264.7 (ATCC) were cultured at 37.degree. C. in DMEM
supplemented with 10% FBS and 5% Penicillin/Streptomycin. Following
the acquisition of cultures of adequate density, cells cultured on
bare glass as well as the composite films were exposed to
lipopolysaccharide (LPS) for 4 hours, and slides were subsequently
transferred to new Petri dishes and 1 ml of TRIzol cell lysis
solution was added to wash the slides and collect the genetic
material. RNA isolation was done according to the manufacturer's
protocol. Subsequent conversion of the RNA to cDNA was performed
using the I-script enzyme (Bio-Rad) (Applequist et al., 2002.
International Immunity 9:1065-1074; Perry et al., 2004. The Journal
of experimental medicine. 199:1651-1658; Doyle et al., 2003.
Journal of immunology 170:3565-3571).
[0078] Following conversion of the isolated mRNA to cDNA, qPCR
analysis (Bio-Rad, Richmond, Calif., USA) was performed to examine
the expression of tumor necrosis factor-alpha (TNF.alpha.),
following LPS induction both with and without dexamethasone
activity.
Example 5
SWNT-Copolymer Deposition for Bioelectric Measurement
[0079] Atomic force microscopy was utilized to image the polymeric
nanofilm substrates as well as analyze SWNT deposition (FIGS.
2A/2B). FIG. 2A represents an AFM micrograph with a scan size of 2
.mu.m at a frequency of 0.7016 Hz. FIG. 2B represents an AFM
micrograph of a SWNT deposited at the air-water interface. In
addition to AFM analysis, Langmuir isotherms were performed for
polymer-only membranes as well as films containing both the polymer
and the SWNT's (FIGS. 3A/B). FIG. 3A shows an increase in surface
pressure versus trough area for the SWNT-containing film indicating
the presence of the nanostructures along with the PMOXA-PDMA-PMOXA
copolymer molecules at the air-water interface. In addition, the
compression isotherm data indicates that carbon nanotubes are
interacting with the polymer blocks at the air-water interface and
not completely falling into the subphase. The polymer solution was
spread at the surface and a compression isotherm was run. The
barriers were relaxed and SWNT's solubilized in chloroform were
spread across the air-water interface, resulting in a slight rise
in the surface pressure, indicating an increase in the mean
molecular area at the air-water interface. This data suggests that
a portion of the SWNT's remained at the surface, possibly due to an
interaction with the hydrophobic block of the copolymer for
substrate deposition.
[0080] The greatest current output occurred when polymer/SWNT films
were deposited by the down dip method (FIG. 3B). In this method,
the substrate was lowered into the trough such that the floating
Langmuir film facing the air is deposited directly onto the
substrate, while the side facing the subphase becomes the top layer
once deposited onto the substrate. As previously mentioned, to
further ensure a high density of SWNT's in the film, polymer and
carbon nanotubes were incubated and dissolved into the same
chloroform solvent prior to spreading. After spreading the
polymer/SWNT solution on top of the subphase, additional SWNT
solution was spread across the surface. The film was compressed to
30 mN/m, and additional SWNT solution was spread on top of the
film. It was believed that the compressed polymer film could
sustain the additional deposited SWNT's, such that during the down
dip method, a layer of SWNTs can be directly deposited onto the
gold electrode. SWNT solution was added until the chloroform no
longer spread rapidly across the surface, indicating that the
surface was saturated. The floating film was then compressed to 40
mN/m for deposition onto a gold electrode. Cyclic voltammetry (CV)
analysis of cytochrome c-mediated electron transfer to the carbon
nanotube films resulted in high current outputs. It was found that
incorporating SWNT's into the polymer membrane could improve the
conductivity of the polymer while maintaining both anodic and
cathodic current peaks. In addition, the presence of the copolymer
was believed to play an additional role to supporting SWNT
interfacial deposition. As the stability of the cytochrome c-gold
interaction was previously shown to be influenced by the
unperturbed interaction between the protein, buffer, and electrode
by preventing the adsorption of protein to the electrode which in
turn precludes high fidelity electron transfer and reaction
reversibility, the hydrophilic copolymer endgroups were believed to
play an important role in precluding cytochrome c adsorption much
like previous reports of using hydroxythiol (HT) self assembled
monolayers [20,21].
Example 6
Glucocorticoid-Copolymer Films for the Study of Cellular
Inflammation
[0081] Dexamethasone (Dex) deposition was conducted atop both the
PEO-PMMA/PMOXA-PDMS-PMOXA copolymers. Dex presence was confirmed
via the application of a FITC-Dex compound atop the pre-formed
copolymer monolayer (FIG. 4A). FITC-Dex added without the copolymer
resulted in the inability to transfer the anti-inflammatory to
solid substrates following film compression (Data not shown).
[0082] Following confirmation of Dex-copolymer composite film
fabrication, Raw 264.7 were cultured atop the active substrates as
well as bare glass slides and LPS was utilized to induce cellular
stress and the production of a suite of inflammatory cytokines and
signaling molecules (FIG. 4B, FIG. 10). LPS binds to the
membrane-bound Toll-like receptor 4 (TLR4) that simulates bacterial
infection and elicits the activation of transcriptional factors
(e.g. Nuclear Factor kappa B (NF.kappa.B) for inflammatory cytokine
production. For the purposes of this study, TNF.alpha. was selected
as the cytokine marker for macrophage stress, and qPCR was utilized
to examine the expression of TNF.alpha. mRNA (FIG. 5A). Macrophages
activated by LPS that were cultured atop bare substrates resulted
in significantly higher levels of inflammation over samples where
LPS was introduced to macrophages that were cultured atop either 3
layers of Dex-copolymer composites. In addition, on samples where
solutions of Dex were incubated directly with the glass culture
slides/self-assembled without the polymer, LPS treatment of
cultured macrophages resulted in the absence of inflammatory
suppression that was previously observed with samples cultured atop
the Dex/copolymer composites (data not shown).
[0083] Furthermore, studies were conducted to evaluate the effects
of copolymer interaction upon cell growth to evaluate its potential
application as a medically-relevant material such as an implant
coating. Studies also examined the potential effects of solubilized
polymer (e.g. polymer in solution) upon cellular growth and
morphology. We incubated the polymer with the cell culture
solutions (0.1 mg/ml-H.sub.2O with sonication) to examine the
impact of the polymer upon cell growth over multiple time points
including 4 hrs, 24 hrs, and 48 hrs. These studies utilized a
triblock copolymer which served as a larger structure to evaluate
potential impact upon cellular growth. FIG. 5B shows that the cell
growth and morphology were unaffected over the measured time period
as the cells were able to undergo normal cell growth behavior. As
such, this study showed that 1) Interfacially deposited Dex atop
the polymeric amphiphiles could be dispensed to macrophages
cultured atop the active composite, and 2) The copolymer is a
biologically-amenable material that interacts favorably with cells
in culture. As demonstrated by the qPCR trials, the copolymer
served as both a buoying element at the air-water interface for the
deposition of Dex, as well as a tethering component to maintain the
integrity between the Dex and the glass substrate. This was an
expected observation as the nature of the copolymer-substrate
interaction, believed to be based upon a high density Van der Waals
interaction, was previously demonstrated. As such, the
intercalation of the Dex molecules within the strongly attached
PEO-PMMA or PMOXA-PDMS-PMOXA copolymers was expected to be
maintained, and potentially enhanced following endgroup
crosslinking via UV exposure, a condition that may not be amenable
towards more conventional, lipid-based amphiphiles that are less
robust than the copolymer molecules. Suppression of LPS-mediated
induction of TNF.alpha. in Raw 264.7 cells demonstrates the ability
of our copolymers in suppressing immune response through Dex
activation of glucocorticoid receptor, GR, which results in GR
binding and inhibition NF-.kappa.B subunit, p65. The activation of
GR by Dex has been previously shown to robustly inhibit NF-.kappa.B
target genes through binding of p65 and blocking of p65 recruited
transcriptional machinery, such as P-TEFb [22].
[0084] This study has demonstrated the concept of applying
copolymer amphiphiles as the foundational element of fabricating
thin films with versatile functionalities. Using both
bioelectro-activity measurement and stress-suppression as exemplary
cases, we have shown that the copolymer element enabled functional
material (SWNT/Dex) deposition, and played a secondary role as an
anti-protein adsorption and tethering component as well. As such,
this methodology represents a broadly applicable technique for life
science studies, drug delivery, as well as the fabrication of
multifunctional electrodes for medical applications.
Example 7
Drug Dosing Based on Copolyer Layers
[0085] For the purposes of this study, TNFa was selected as the
cytokine marker for macrophage stress, and RT-PCR was utilized to
examine the expression of TNF.alpha. mRNA (FIG. 7). Macrophages
activated by LPS that were cultured atop bare substrates resulted
in significantly higher levels of inflammation over samples where
LPS was introduced to macrophages that were cultured atop either 3,
or 7 layers of Dex-copolymer composites. This outcome demonstrated
that the drug-dosing capabilities of the hybrid material could be
tailored based upon the number of layers deposited. Because the
copolymer membrane dimensions are within the range of 2 nm
(diblock)-4 nm (triblock) in thickness, coating substrates (e.g.
implants) with multiple layers would result in minimal impact upon
overall device dimensions. In addition, on samples where solutions
of Dex were incubated directly with the glass culture
slides/self-assembled without the polymer, LPS treatment of
cultured macrophages resulted in the absence of inflammatory
suppression that was previously observed with samples cultured atop
the Dex/copolymer composites (FIG. 8). Self-assembled Dex without
the copolymer substrate was unable to stay tethered to the glass
slide for subsequent transfer to the macrophages (FIG. 9).
Example 8
APC Membrane Functionalized with Anti-Inflammatory Drug
[0086] In the initial studies the APC membranes the were
functionalized with Dexamethasone (Dex), an anti-inflammatory drug.
FIG. 11 provides a cross section view of the_APC membrane. The base
layer was comprised of a tangible layer of parylene that was
deposited through vapor deposition with an SCS Coater. This
resulted in an even layer of_parylene free of pinholes to serve as
the backbone of the APC membrane. In FIG. 11, the middle layer is
the functionalized layer. In the present embodiment, the middle
layer was functionalized with 300 uL of 5 mglmL Dex. The Dex was
incorporated via Langmuir-Blodgett deposition into a tri-block
copolymer matrix, which secured the Dex molecules in a copolymer
network matrix. Finally, another parylene layer was deposited to
complete the membrane. The final parylene layer was an ultra-thin
layer that had nanopores. FIG. 12 shows AFM images that were taken
that suggest the possibility of nanopores on the surface of
parylene. The darker areas indicate depressions on the surface that
could be nanopores. As FIG. 12 shows, the surface of parylene is
very rough terrain full of micro/nano-structures that potentially
act as inflammatory stimuli. Comparing the AFM images of 0.03 g
parylene with drug to 0.03 g parylene without drug, it can be
clearly seen that the layer of Dex incorporated into a copolymer
matrix changes the terrain and is definitely present under the
porous parylene layer. The same conclusion may be drawn from the
comparison between the 0.05 g parylene with and without drug. FIG.
13 suggests that macrophages may be able to sense substrate
stiffness and roughness to prompt an inflammatory response. The
graph in FIG. 13A shows relative levels of iNOS expression between
RAW 2643 murine macrophages cultured on parylene and parylene
covered with a layer of copolymer. The bottom graph in FIG. 13B
shows relative levels of TNF-a expression between RAW 264.7 murine
macrophages grown on the different substrates. The graphs suggest
that macrophages may sense the copolymer/parylene surface as a more
biocompatible medium than parylene alone. However, the copolymer
cannot be handled tangibly and parylene is needed as a backbone for
the copolymer. The unique combination of copolymer and parylene is
an integral and distinctive aspect of the APC membrane that
simultaneously enables the copolymer to be handled, while taking
advantage of its biomimetic properties and ability to be
functionalized with an array of molecules.
[0087] Tissue culture images are shown in FIG. 14, which provides
qualitative input on effects of growth on glass, parylene, and APC
membrane substrates on cell development and morphology. The images
of RAW 264.7 murine macrophage growth on glass, parylene, and APC
membrane were taken over three time points. At 8 hours, the
macrophages had been cultured in DMEM with 10% FBS and 1% Pen/Strep
for 8 hours on their respective surfaces for the purpose of
reaching the desired cell confluence on the various substrates; the
plain glass slide served as the control. At the 8 hour time point,
macrophage growth on the plain glass is about 60-70% confluent with
minimal visible clumping. Contrasting growth on glass with growth
on parylene and the APC membrane, it can be seen that the cell
development is stunted; there is very little growth on the parylene
and APC membrane substrates. This observation suggests that growth
on parylene and the APC membrane is difficult for murine
macrophages due to poor adhesion. This quality is very desirable
and ideal for any implant and invasive biomedical device because
cells are less likely to adhere to the device in problematic
regions that could impair or even disable the function of the
device. Images were taken again at 20 hours and immediately
afterwards, 75 uL of I ug/mL lipopolysaccharide solution (LPS) was
added to each sample to induce inflammation in the macrophages. At
20 hours before LPS induced inflammation, it can be seen that the
cell growth is still virtually halted on the APC membrane and
parylene surfaces; macrophages continue to multiply on plain glass
reaching 90-95% confluence by 20 hours of growth.
[0088] LPS stimulation occurred for 4 hours and images were taken
at the third time point (24 hours). In these images the macrophages
growing on plain glass were well over 100% confluent. Growth on
parylene increased slightly over these 4 hours with LPS
stimulation, and there was no visible change in macrophage
development on the APC membrane. These data warrant an observation
made previously: a parylene or APC membrane layer on biomedical
devices hinders macrophage adhesion due to inadequate traction on
the aforementioned substrates. Diminished adhesion results in a
more biomimetic material for invasive biomedical devices, but a
suitable coating must also address the issue-of inflammation:
the-APC membrane accomplishes this goal.
[0089] FIG. 15 shows the data from RT-PCR gene expression studies
done on the macrophages cultured on plain glass, parylene, and APC
membrane. The macrophages were incubated on the substrates for 24
hours. After 20 hours, 75 uL of 1 ug/mL LPS was added to each
sample to inflame the macrophages; LPS stimulation lasted for 4
hours. After 24 hours the macrophages were harvested. RNA isolation
was achieved through cell lysis with Trizol reagent washes followed
by chloroform extraction, centrifugation and isopropyl alcohol
precipitation. cDNA was synthesized with 5.times. buffer, Oligo DT
primer, reverse transctiptase, and the purified RNA samples. After
a water bath at 37.degree. C., the cDNA was mixed with primers for
RT-PCR analysis. Three cytokines that control inflammation were
investigated: TNF-a, IL-6 and iNOS (3-Actin was used to normalize
the relative quantities of genes expressed). Relative IL-6
and_TNF-a expression show an apparent decrease in secretion of IL-6
and TNF-a. The relative_iNOS expression data did not show an
advantage for any particular substrate. From past RT-PCR
experiments conducted in the laboratory, iNOS was identified as an
unpredictable cytokine; essentially, the results for iNOS were
seldom consistent. FIG. 15C, when the error bars are taken into
account iNOS expression levels are comparable across the different
substrates, Additionally, past experiments have also determined
IL-6 to be the most consistent inflammatory cytokine we have
studied. The IL-6 data presented in FIG. 15A clearly show the APC
membrane decreases inflammation up to 2.5 times compared to plain
glass and parylene surfaces. The advantage that the functionalized
drug-eluting, biocompatible APC membrane has over plain parylene
was not apparent in our cell images, but RT-PCR shows inflammation
is significantly reduced when macrophages are cultured on the APC
membrane functionalized with Dex. Therefore, the APC membrane is a
very biocompatible coating capable of eluting drug resulting in
significantly reduced inflammation in RAW 2647 munine
macrophages.
[0090] FIG. 16 illustrates results of slow-elution studies through
RT-PCR gene expression data. The samples represented in FIG. 16
were prepared differently from the samples in FIG. 15. To test
slow-elution capabilities, the APC membrane was soaked in a Petri
dish with DMEM (containing 10% FBS and 1% Pen/Strep) for 3 days
prior to tissue culture. The plain glass and parylene samples were
also soaked for 3 days to maintain uniformity in experimentation.
After 3 days of soaking in media, macrophages were cultured on the
samples. Soaking allowed the dexamethasone to elute slowly from the
APC membrane. FIG. 16A shows relative IL-6 expression was 18 times
less for the macrophages cultured on APC membrane than cells on
parylene, and 13 times less than macrophages grown on plain
glass.
[0091] These data suggest that a 3 day soak allowed the Dex to
slowly diffuse through the nanopore layer, thus eluting more Dex
into the media for interaction with macrophages. The samples
presented in FIG. 15 were not soaked in media prior to tissue
culture. Therefore, the macrophages harvested for FIG. 15 data were
exposed to less drug because the Dex had less time to elute into
the media. The effects of the difference between soaking and not
soaking the membrane may be seen in relative gene expression
graphs. The relative IL-6 expression of macrophages cultured on
unsoaked membrane was over 2.5 times less than IL-6 expression of
macrophages grown on parylene, but the soaked membrane resulted in
over 18 times less IL-6 expression of macrophages grown on
parylene. Therefore, RT-PCR data demonstrates that there was a
clear advantage in soaking the membrane in vitro studies, which
translates to effective and prolonged bio-longevity of the
biomedical device coated with APC membrane in a patient.
[0092] FIG. 16 also compares the APC membrane to a functionalized
copolymer membrane with no nanopore layer engineered for slow
elution. These data show that slow-elution is advantageous because
it keeps the Dex that is eluted localized near the surface of the
membrane where most of the effective dose can interact with
macrophages to decrease inflammation. Even in a Petri dish, we can
see the inefficiency of haphazard drug delivery; localized drug
delivery through slow-elution is nearly 3 times more effective in
reducing inflammation when compared to the functionalized copolymer
membrane that elutes all the drug at once. This effect will be more
pronounced in an actual patient where the strictly defined
boundaries of a Petri dish are removed because the drug will have
more area to diffuse through, which causes the effective dose to be
diluted, and thus results in diminished treatment and drug
efficacy.
[0093] In vitro studies indicate that the APC membrane is a
versatile material capable of delivering Dex in a controlled
fashion to maximize drug efficacy. Ultimately, the biocompatibility
of the prylene and copolymer materials that was incorporated into
the functionalized membrane, makes the APC membrane a low adhesion,
biomimetic slow drugeluting tangible membrane and therefore, a
prime candidate as an ideal biomedical device coating and platform
for chemotherapeutics, anti-inflammatory treatments, and
nanomedicine in the future.
Example 9
Industrial Applicability of the Embodiments
[0094] The APC membrane is a very versatile coating suitable for a
broad scope of applications. The APC membrane has the potential to
serve as a flexible drug delivery platform capable of diverse
functionalization. Novel drugs are developed for countless
therapeutics, but these drugs all need a method of delivery for
their therapeutic potential to be realized. This method must be
precise, biocompatible, and ideally, slow-eluting to prolong the
time the drug may act in a patient; the APC membrane meets these
requirements. Therefore, the APC membrane may be easily made into a
drug-eluting patch for organs to deliver drugs to a specific part
of the organ. An example of this application is a heart patch that
could aid cardiac tissue regeneration by acting as a growth
scaffold, for instance, after a heart attack by incorporating
growth hormones into the copolymer matrix, the APC membrane may be
used as a foundation for tissue regeneration and augmentation of
cardiac, muscle, and vascular tissue among others.
[0095] In addition to serving as a drug conjugation and delivery
platform, the APC membrane may also effectively coat implantable
and invasive biomedical devices. The surface properties of the
parylene surface may be "tuned" to enable effective switching of
the parylene properties to either promote or resist cell adhesion
depending on the specific application desired or implant being
engineered. Such tuning can be employed to further enhance the
versatility of the device as bio-interfacial properties are vital
to the longevity of the device itself Furthermore, when
functionalized with Dex, the APC membrane significantly decreases
inflammation and minimizes cell adhesion making it a biomimetic
"skin" capable of boosting the biocompatibility and versatility of
many biomedical devices, establishing the membrane as a_relevant
application for all future invasive medical technologies.
Example 10
In Vivo Characterization of Tri-Block Polymer Nanofilms
[0096] This Example describes in vivo studies that were carried out
to examine the material-mediated blockage of cell aggregation
around an implant site when nanofilm material is employed.
Histological analysis was performed using Hematoxylin and Eosin
(H&E) to image cell recruitment activity. Untreated tissue,
tissue containing implanted uncoated disks, as well as
Dex-copolymer nanofilms were stained to examine cellular
recruitment to the implant surface which is a commonly observed
mechanism of foreign body formation and eventual implant fouling
(FIG. 17). C57b1/6 mice (n=6) were subcutaneously implanted
dorsally with two polyethylene disks (Uncoated or PolyDex coated).
Following 7 days, disks were excised, formalin fixed and
Hematoxylin and Eosin stained. Samples were analyzed at 10.times.
magnification. Representative images are presented. Black arrows
indicate site of cell infiltration at the interface of the dermis
and disk. FIG. 17A showed that the untreated tissue exhibits the
presence of equally dispersed cells as shown by the nuclear stains.
FIG. 17B showed that the implantation of disks only without
nanofilm coatings exhibited a rapid recruitment of cells to the
implant surface. Furthermore, FIG. 17C showed that the tissue that
resided at the interface of the implant with the nanofilm showed
the suppression of tissue inflammation. As such, the ability to
translate the observed results from the in vitro study to in vivo
applications was demonstrated using H&E analysis. The
translational applicability of the material was further confirmed
using F4/80 staining of the cells that were observed to aggregate
around the implant site. This macrophage-specific stain showed that
the cells that aggregated around the uncoated implant, and the
cells that were precluded from aggregation around the
polymer-coated implant were indeed macrophages.
[0097] Continued studies examined the potential effects of
solubilized polymer (e.g. polymer in solution) upon cellular
proliferation and normal cellular activity. Varying concentrations
of the polymer were incubated with the cell culture solutions (0.01
mg/ml, 0.1 mg/ml) to examine the impact of the polymer upon cell
growth over the multiple time points including 4 hrs, 24 hrs, and
48 hrs. FIG. 18 shows that the cell growth and morphology was
unaffected which indicated that the polymer was indeed
biologically-inert, further strengthening its application as an
implant coating. The observed lack of adverse effects upon cell
growth was further confirmed by examining the cellular inflammation
programs of the macrophages exposed to the 0.01 mg/ml and 0.1 mg/ml
polymeric concentrations (highest concentrations applied). FIG. 19
shows that the mRNA levels of TNF.alpha., IL-6, IL-12, and iNOS are
virtually unaffected, and actually decrease slightly when compared
with the controls where no polymer was added, further confirming
quantitatively that cellular activity is unaffected by the
biologically-inert material. This Example has produced a
comprehensive assessment of the potential that this
nanopolymer-drug composite possesses as a robust and versatile
implant coating that can provide localized cellular regulation
using a variety of effector molecules situated on a spectrum of
implant surfaces.
[0098] All publications, patent applications, issued patents, and
other documents referred to in this specification are herein
incorporated by reference as if each individual publication, patent
application, issued patent, or other document was specifically and
individually indicated to be incorporated by reference in its
entirety. Definitions that are contained in text incorporated by
reference are excluded to the extent that they contradict
definitions in this disclosure.
[0099] The present invention, thus generally described, it should
be understood that changes and modifications can be made therein in
accordance with ordinary skill in the art without departing from
the invention in its broader aspects as defined in the following
claims.
REFERENCES
[0100] 1. D. Ho, B. Chu, H. Lee, and C. D. Montemagno. 2004.
Protein-driven energy transduction across polymeric biomembranes.
Nanotechnology 15:1084-1094.
[0101] 2. J. Xi, D. Ho, B. Chu, and C. D. Montemagno. 2005. Lessons
Learned From Engineering Biologically-Active Hybrid
Nano/Micro-devices. 2005 Advanced Functional Materials
15:1233-1240.
[0102] 3. D. Ho, S. Chang, and C. D. Montemagno. 2006. Fabrication
of biofunctional nanomaterials via Escherichia coli OmpF protein
air-water interface insertion/integration with copolymeric
amphiphiles. Nanomedicine 2:103-112.
[0103] 4. A. Graff, M. Sauer, P. van Gelder, and W. Meier. 2002.
Virus-Assisted Loading of Polymeric Nanocontainers. Proc. Nat.
Acad. Sci. 99:5064-5068.
[0104] 5. C. Nardin, T. Hirt, J. Leukel and W. Meier. 2000.
Polymerized ABA-triblock copolymer vesicles. Langmuir
16:1035-1041.
[0105] 6. D. Ho, B.Chu, H.Lee, E. K.Brooks, K. Kuo, and C. D.
Montemagno. 2005. Fabrication of biomolecule-copolymer hybrid
nanovesicles as energy conversion systems. Nanotechnology.
16:3120-3132.
[0106] 7. H. Lee, D. Ho, and C. D. Montemagno. 2006. Fluorometric
Measurement of Vectorially-Inserted Purple Membrane Activity Across
Block Copolymer Thin Films. Polymer 47:2935-2941.
[0107] 8. W. Stoeckenius, R. H. Lozier, and R. A. Bogomolni. 1979.
Structure of biological membranes. Biochem. Biophys. Acta
505:215-278.
[0108] 9. J. F. Rathman, and P. Sun. 2005. Biocomposite films
synthesized at a fluid/fluid interface. Faraday
Disc.129:193-203.
[0109] 10. G. Grant, D. Koktysh, B. Yun, R. Matts, and N. Kotov.
2001. Layer-By-Layer Assembly of Collagen Thin Films: Controlled
Thickness and Biocompatibility. Biomed. Microdev.3:301-306.
[0110] 11. M. M. Ghannam, M. M. Mady, and W. Khalil. 1999.
Interaction of type-I collagen with phospholipid monolayer.
Biophys. Chem. 80:31-40.
[0111] 12. W. Meier, C. Nardin, and M. Winterhalter. 2000.
Reconstitution of Channel Proteins in (Polymerized) ABA Triblock
Copolymer Membranes. Angew. Chim. Int. Ed. 39:4599-4602.
[0112] 13. C. Nardin, M. Winterhalter, W. Meier. 2000. Giant
Free-Standing ABA Triblock Copolymer Membranes. Langmuir.
16:7708-7712.
[0113] 14. D. Ho, B. Chu, J. J. Schmidt, E. Brooks, and C. D.
Montemagno. 2004. Hybrid Protein/Polymer Biomimetic Membranes. IEEE
Trans. Nanotechnology. 3:256-263.
[0114] 15. Lee, H., D. Ho, J. J. Schmidt, and C. D. Montemagno.
2003. Biosolar Powered Fabric. IEEE Proceedings on Nanotechnology
2:733 -736.
[0115] 16. D. Ho, B. Chu, K. Kuo, and C. D. Montemagno. 2004.
Functionalizing Biomimetic Membranes with Energy Transducing
Proteins. Proc. of the Mat. Res. Soc. 823:W11.8.1-W11.8.6.
[0116] 17. S. Applequist, R.P.A. Wallin, and H. G. Ljunggren. 2002.
Variable expression of toll-like receptor in murine innate and
adaptive immune cell lines. International Immunity 9:1065-1074.
[0117] 18. A. K. Perry, E. K.Chow, J. B.Goodnough, W. C. Yeh, and
G. Cheng. 2004. Differential requirement for TANK-binding kinase-1
in type I interferon responses to toll-like receptor activation and
viral infection. The Journal of experimental medicine.
199:1651-1658.
[0118] 19. S. E. Doyle, R. O'Connell, S. A. Vaidya, E. K. Chow, K.
Yee, and G. Cheng. 2003. Toll-like receptor 3 mediates a more
potent antiviral response than Toll-like receptor 4. Journal of
immunology 170:3565-3571.
[0119] 20. S. Terrettaz, J. Chen, C. J. Miller, and R. D. Guiles.
1996. Kinetic Parameters for Cytochrome c via Insulated Electrode
Voltammetry. Journal of the American Chemical Society.
118:7857-7858.
[0120] 21. A. Szucs, and M. Novak. 1995. Stable and Reversible
Electrochemistry of Cytochrome-C on Bare Electrodes .2. Effects of
Experimental Conditions. Journal of Electroanalytical Chemistry,
383:75-84.
[0121] 22. H. F. Luecke and K. R. Yamamoto. 2005. The
glucocorticoid receptor blocks P-TEFb recruitment by NFkappaB to
effect promoter-specific transcriptional repression. Genes Dev.
19:1116-1127.
[0122] 1B. E. Grube et al, "Six-month clinical and angiograplric
results of a dedicated drug-eluting stent for the treatment of
coronary bifurcation narrowings," The American Journal of
Cardiology, vol 99, pp 1691-1697,2007.
[0123] 2B. M Sokolsky-Papkov, K. Agaslri, A , Glaye, K. Shakesheff;
A J Domb, "Polymer carriers for drug delivery in tissue
engineering," Advanced Drug Delivery Reviews, voL 59, pp
187-206,2007.
[0124] 3B. H M Butt W. L Hunter, "Thug-eluting stents: A
multidisciplinary success Story," Advanced Drug Delivery Reviews,
vol. 58, pp 350-357,2006.
[0125] 4B H M Butt, W L. Hunter, "Drug-eluting stents: an
innovative multidisciplinary drug delivery platform," Advanced Drug
Delivery Reviews, vol 58, pp. 345-0346, 2006.
[0126] 5B. M. M. Krucoff, A Bomn, D G. Schultz, "Drug-eluting
stents `deliver heartburn`-How do we spell relief going forward,"
Circulation, vol 115, pp. 2990-2994, 2007.
[0127] 6B D. Ho, R Chu, H Lee, and C. D. Montemagno. 2004
Protein-driven energy transduction across polymeric biomembranes
Nanotechnology 15:1084-1094.
[0128] 7B J. Xi, n. Ho, B Chu, and C. D Montemagno 2005. Lessons
Learned From Engineering Biologically-Active Hyblid
NanolMicro-devices. 2005 Advanced Functional Materials
15:1233-1240.
[0129] 8B. D Ho, S. Chang, and C D. Montemagno. 2006. Fabrication
of biofunctional nanomaterials via Eschericlria coli OmpF protein
air water interface insertion/integration with copolymeric
amphiphiles Nanomedicine 2: 103-112.
[0130] 9B. A Graff, M Sauer, P van Gelder, and W Meier. 2002
Virus-Assisted Loading of Polymeric Nanocontainers. Proc Nat Acad.
Sci. 99:5064-5068.
[0131] 10B. C Nardin, I. Hirt, 1 Leukel and W. Meier 2000
Polymelized ABA-tri-block copolymer vesicles. Langmuir
16:1035-1041.
[0132] 11B. D. Ho, B Chu, H Lee, E. K. Brooks, K. Kuo, and C D
Montemagno.2005 Fabrication of biomolecule-copolymer hybrid
nanovesicles as energy conversion systems. Nanotechnology
16:31203132.
[0133] 12B H. Lee, D. Ho, and CD. Montemagno. 2006 Fluorometric
Measurement of Vectorially Inserted Purple Membrane Activity Across
Block Copolymer Thin Films Polymer 47:2935-2941.
[0134] 13B W Stoeckenius, R H Lozier, and R. A Bogomolni. 1979.
Structure of biological membranes, Biochem. Biophys Acta
505:215-278.
[0135] 14B J F Rathman, and P. Sun 2005. Biocomposite films
synthesized at a fluid/fluid interface, Faraday Disc 129:
193-203.
[0136] 15B G Grant, D. Koktysh, B Yun,R Matts,and N Kotov2001
Layer-By-Layer Assembly of Collagen Thin Films: Controlled
Thickness and Biocompatibility Biomed Microde v 3:301.306.
[0137] 16B M. M Ghannam, M. M. Mady, and W Khalil 1999 Interaction
of type-1 collagen with phospholipid monolayer, Biophys. Chem
80:31-40.
[0138] 17B W. Meier, C Nardin, and M. Winterhalter. 2000
Reconstitution of Channel Proteins in (polymerized) ABA Tri-block
Copolymer Membranes. Angew Chim Int Ed 39:4599-4602
[0139] 18B. C Nardin, M Winterhalter, W Meier 2000. Giant
Free-Standing ABA Tri-block Copolymer Membranes Langmuir.
16:7708-7712.
[0140] 19B P. B Malafaya, G A, Silva, R. L Reis, "Natural-origin
polymers as carriers and scaffolds for biomolecules and cell
delivery in tissue engineering application," Advanced Drug Delivery
Reviews, vol 59, pp 207-233,2007.
[0141] 20B. N Stark, "Literature review: Biological safety of
parylene C," Medical Plastics and Biomaterials, p. 30, March
1996
[0142] 21B. L. Wolgemuth, "Assessing the performance and
suitability of patylene coating," Medical Device and Diagnostic
Industry, p 42, August 2000.
[0143] 22B. H. L, Wong et aL, "Chemotherapy with anticancer drugs
encapsulated in solid lipid nanoparticles," Advanced Drug Delivery
Reviews, 2007.
[0144] 23B. D. W ,Grattan and M, Bilz, "The thermal aging of
parylene and the effect of antioxidant," Studies in Conservation,
voL 36, pp 44-52, 1991.
[0145] 24B. D Ho, R Chu, I J Schmidt, E Brooks, and C D Montemagno.
2004 Hybrid Protein Polymer Biomimetic Membranes IEEE Trans.
Nanotechnology 3:256-263
[0146] 25B. Lee, H, D Ho, J J Schmidt, and C D. Montemagno 2003
Biosolar Powered Fabric IEEE, Proceedings on Nanotechnology 2:733
-736.
[0147] 26B. D. Ho, B. Chu, K Kuo, and C D. Montemagno 2004
Functionalizing Biomimetic Membrane with Energy Transducing
Proteins. Proc.. of the Mat Res. Soc 823:W11 8 1-W118.6.
[0148] 27B. S Applequist, R P A Wallin, and H. G. Ljunggren. 2002
Variable expression of toll-like receptor in murine innate and
adaptive immune cell lines. International Immunity 9:1065-1074.
[0149] 28B. A K Perry, E K Chow, I B Goodnough, W. C Yeh, and G.
Cheng 2004. Differential requirement for TANK-binding kinase-I in
type I interferon responses to toll-like receptor activation and
vital infection. The journal of experimental medicine
199:1651-1658.
[0150] 29B S. E. Doyle, R O'Connell, S. A Vaidya, E. K Chow, K Yee,
and G Cheng 2003 Toll-like receptor 3 mediates a more potent
antiviral response than Toll-like receptor 4. Journal of
immunology, 170:3565-3571.
[0151] 30B S. Terrettaz, I Chen, C L Miller, and R. D. Guiles.
1996. Kinetic Parameters for Cytochromec via Insulated Electrode
Voltammetry Journal of the American Chemical Society
118:7857-7858.
[0152] 31B A Szucs, and M Novak 1995 Stable and Reversible
Electrochemistry of Cytochrome-C on Bare Electrodes Effects of
Experimental Conditions. Journal of Electoanalytical Chemistry,
383:7584.
[0153] 32B. H. F. Luecke and K R. Yamamoto 2005. The glucocorticoid
receptor blocks P-TEFb recruitment by NFkappaB to effect
promoter-specific transcriptional repression Genes Dev
19:11161127.
* * * * *