Note: Descriptions are shown in the official language in which they were submitted.
CA 02637578 2013-09-18
Temporarily Stiffened Mesh Prostheses
This application claims priority under 35 U.S.C. 119(e)(5) of U.S.
Provisional Patent
Application No. 60/772,827, filed February 8, 2006.
Field of the Invention
[0001] The present invention relates to medical prostheses and methods of
manufacturing
those devices. In particular, the prostheses are temporarily stiffened meshes
with particular
coatings to provide initial stiffness and thereby permit easier surgical
handling for treatment
or reconstruction of soft tissue defects. Preferred embodiments include
surgical meshes
coated with one or more biodegradable polymers that can act as a stiffening
agent by coating
the filaments or fibers of the mesh to temporarily immobilize the contact
points of those
filaments or fibers and/or by increasing the stiffness of the mesh by at least
1.1 times its
original stiffness. The devices of the invention can also provide relief from
various post-
operative complications associated with their implantation, insertion or
surgical use. By
including biologically active agents and/or drugs in the coating, the devices
provide
prophylaxis for and can alleviate side effects or complications associated
with the surgery or
use of prostheses in general.
Background of the Invention
[0002] Prosthetic implants such as meshes, combination mesh products or
other porous
prostheses are commonly used to provide a physical barrier between types of
tissue or extra
strength to a physical defect in soft tissue. However, such devices are often
associated with
post-surgical complications including post-implant infection, pain, excessive
scar tissue
formation and shrinkage of the prosthesis or mesh. Excessive scar tissue
formation, limited
patient mobility, and chronic pain are often attributed to the size, shape,
and mass of the
implant and a variety of efforts have been undertaken to reduce the amount of
scar tissue
CA 02637578 2008-08-07
formation. For example, lighter meshes using smaller fibers, larger weaves,
and/or larger
pore sizes as well as meshes woven from both non-resorbable and resorbable
materials are in
use to address these concerns.
[0003] For treating acute pain and infection, patients with implanted
prostheses are
typically treated post-operatively with systemic antibiotics and pain
medications. Patients
will occasionally be given systemic antibiotics prophylactically; however,
literature review of
clinical trials does not indicate that systemic antibiotics are effective at
preventing implant-
related infections.
[0004] Many types of soft tissue defects are known. For example, hernias
occur when
muscles and ligaments tear and allow the protrusion of fat or other tissues
through the
abdominal wall. Hernias usually occur because of a natural weakness in the
abdominal wall
or from excessive strain on the abdominal wall, such as the strain from heavy
lifting,
substantial weight gain, persistent coughing, or difficulty with bowel
movements or urination.
Eighty percent of all hernias are located near the groin but can also occur
below the groin
(femoral), through the navel (umbilical), and along a previous incision
(incisional or ventral).
Almost all hernia repair surgeries are completed with the insertion of a
barrier or prosthesis to
prevent their reoccurrence. Therefore products used in the management of
hernias require
some measure of permanent strength. The most commonly employed woven meshes
are
crafted from polypropylene fibers using various weaves. Tightly woven meshes
with the
highest strength characteristics and stiffness are very easy for the surgeon
to implant;
however, there appears to be a positive correlation between the tightness of
the weave
(correlated to surface area and stiffness), lack of patient mobility, and
chronic pain. Newer
meshes have larger pore structures and while they are more flexible, they are
also more
difficult to implant by surgeons. They are extremely difficult for
laparoscopic repair, as they
have very little recoil associated with them and, when rolled up to insert,
they cannot be
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reflattened and positioned in a quick and efficient manner by the surgeon.
Hence, a need still
exists for surgical meshes, including hernia meshes, that have sufficient
stiffness to facilitate
handling and ease of insertion during surgery, yet are or can become
sufficiently flexible to
be comfortable after implantation.
[0005] Surgical meshes that have been manipulated to improve handling,
insertiona and
positioning post-insertion are known in the art, but do not employ larger-pore
mesh
construction. For example, a laparoscopic surgical mesh with extruded
monofilament PET
coils or rings (e.g., the Bard Composix Kugel hernia patch) increases the
overall stiffness
of the device and gives a shape memory to the device but does not readily
allow for drug
loading of the mesh, can not provide temporary stiffening of the mesh
component, and can
not be further shaped into a fixed three-dimensional structure after
manufacture without
further processing or alteration. Similarly, meshes with reinforced edges have
been produced
(e.g., Bard Visilex ). These meshes have the same disadvantages as those with
coils or
rings. Additionally, the Kugel patch ring has been reported to break under
conditions of use,
causing patient morbidity and mortality.
[0006] Meshes produced from a co-weave of a biodegradable material with a
non-
biodegradable material have been described, e.g., the Johnson & Johnson Vypro
and Vypro
II meshes. In these meshes, polypropylene and polyglactin filaments are
braided together
before being knitted into a mesh. Such meshes do not change stiffness upon
implant as the
polyglactin fibers are very fine and flexible. The biodegradable fibers in the
VyPro meshes
in concert with its particular fiber weave imparts additional flexibility to
the mesh such that it
distends more easily than the surrounding tissue so that it is more flexible
than an equivalent
polypropylene fiber mesh with the same weave. Moreover, because the
biodegradable
polymers of that mesh may be subjected to high temperatures to produce fibers
and filaments
suitable for weaving, it drastically limits the drugs or biologically active
agents that can be
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included in a biodegradable layer since, under such conditions, the vast
majority of
biologically-active agents and drugs are unable to withstand the manufacturing
temperatures
involved in fiber and filament formation. If three-dimensional structures are
desired for such
meshes, they must undergo further processing to attain such shapes. Finally,
these meshes
are often more difficult for surgeons to anchor in place because the
polyglactin fiber cannot
withstand the suturing tension.
[0007] The present invention overcomes these disadvantages by providing
temporarily
stiffened and shapeable meshes.
Summary of the Invention
[0008] The present invention is directed to a medical prosthesis comprising
a mesh and
one or more coating which temporarily stiffens the mesh to at least 1.1 times
its original
stiffness. The coatings on such meshes do not alter the integrity of the mesh
and thus allow
the mesh to remain porous. In general, the coatings do not substantially alter
the porosity of
the mesh. More particularly, the medical prostheses of the invention comprise
a mesh with
one or more coatings with at least one of the coatings comprising a stiffening
agent that coats
the filaments or fibers of the mesh so to temporarily immobilize the contact
points of those
filaments or fibers. Again, the coatings on such meshes do not alter the
integrity strength of
the underlying mesh and thus allow the mesh to remain porous after coating.
The meshes are
capable of substantially reverting to their original stiffness under
conditions of use.
[0009] The stiffening agents of the invention can selectively, partially or
fully coat the
contact points of the filaments or said fibers of the mesh to create a
coating. The contact
points generally include the knots of woven meshes. Such coating are
preferably positioned
on the mesh in a templated pattern or in an array such as might be deposited
with ink-jet type
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technology, including computer controlled deposition techniques. Additionally,
the coatings can
be applied on one or both sides of the mesh.
[0010] In accordance with the invention, the medical prostheses of the
invention include
meshes that have been formed into three-dimensional shapes with the shape
being maintained by
the strength imparted by the coating. Such meshes flat or substantially flat
before coating and the
application of the coating provides the structural support for the three-
dimensional shape. Any
shape can be formed, including curved meshes and conical shapes. The mesh may
be a woven
polypropylene that has a backing of expanded polytetrafluoroethylene (ePTFE).
[00111 The stiffening agents include but are not limited to hydrogels
and/or biodegradable
polymers. One or more biodegradable polymers can be used per individual
coating layer.
Preferred biodegradable polymer comprises one or more tyrosine-derived
diphenol monomer
units as polyarylates, polycarbonates or polyiminocarbonates. The
biodegradable polymer may
be a polyarylate, which may be desaminotyrosyl-tyrosine (DT) and an
desaminotyrosyl-tyrosyl
ester (DT ester).
[0012] In another aspect of the invention, the medical prosthesis of the
invention have at least
one of the coatings that further comprises one or more drugs. Such drugs
include, but are not
limited to, antimicrobial agents, anesthetics, analgesics, anti-inflammatory
agents, anti-scarring
agents, anti-fibrotic agents and leukotriene inhibitors. The antimicrobial
agent may be
vancomycin. One of the coatings may be an anti-inflammatory agent selected
from non-selective
cox-1 and cox-2 inhibitors. One of the coatings may be an anti-inflammatory
agent selected from
selective cox-1 and cox-2 inhibitors.
[0013] Yet another aspect of the invention is directed to a process for
coating a mesh with a
stiffening agent that coats the filaments or fibers of the mesh to temporarily
immobilize the
contact points of the filaments or fibers of said mesh by (a) preparing a
coating solution
CA 02637578 2013-09-18
comprising a solvent and the stiffening agent; (b) spraying a mesh one or more
times to provide
a sufficient amount of solution on the said mesh to produce a coating having a
thickness and
placement sufficient to temporarily immobilize the contact points of the
filaments or fibers of
said mesh that coats filaments or fibers; and(c) drying the mesh to produce
the desired coating.
[00141 Still another aspect of the invention provides a method for
producing a shaped or
three-dimensional medical prosthesis by (a) forming a mesh into a desired
shape or three-
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dimensional shape; (b) applying a stiffening agent to the mesh to coat the
filaments or fibers
of the mesh; (c) allowing the agent to dry, set or cure causing the mesh to
temporarily
immobilize contact points of the filaments or fibers of said mesh and thereby
to retain the
desired shape. Shapes can be formed by affixing the mesh to a mold or
framework to created
the desired shape, applying the agent to the mesh while so affixed, and
removing the mesh
from the mold or framework after the agent has dried, set or cured
sufficiently to allow the
mesh to retain its shape.
[0015] The coated meshes of the invention are capable of releasing one or
more drugs
into surrounding bodily tissue such that the drug reduces or prevents an
implant- or surgery-
related complication. For example, the surgical mesh coatings can include an
anesthetic
agent such that agent seeps into the surrounding bodily tissue, bodily fluid,
or systemic fluid
in a predictable manner and at rate sufficient to attenuate the pain
experienced at the site of
implantation. In another example, the surgical meshes coatings can include an
anti-
inflammatory agent such that the anti-inflammatory agent seeps into the
surrounding bodily
tissue, bodily fluid or systemic fluid in a predictable manner and at a rate
sufficient to reduce
the swelling and inflammation associated implantation of the mesh. Still a
further example,
the surgical mesh coatings can include an antimicrobial agent such that the
antimicrobial
agent is released into the surrounding bodily tissue, bodily fluid, or
systemic fluid in a
predictable manner and at a therapeutically-effective dose to provide a rate
of drug release
sufficient to prevent colonization of the mesh (and/or surgical implantation
site) by bacteria
for a minimum of the period of time following surgery necessary for initial
healing of the
surgical incision.
[0016] In yet another embodiment, the coated surgical meshes of the
invention can be
formed to encapsulate a pacemaker, a defibrillator generator, an implantable
access system, a
neurostimulator, or any other implantable device for the purpose of securing
them in position,
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providing pain relief, inhibiting scarring or fibrosis and/or inhibiting
bacterial growth. Such
coated meshes formed into an appropriate shape either before or after coating
with the
biodegradable polymers.
[00171 The surgical meshes of the invention can deliver multiple drugs from
one or more
independent layers.
100181 The invention thus provides a method of delivering drugs at
controlled rates and
for set durations of time using biodegradable polymers.
Brief Description of the Drawings
[0019] Fig 1. graphically depicts the zone of inhibition (ZOI) for
polyarylate-coated
meshes containing rifampin and minocycline hydrochloride that have been
incubated on
Staphylococcus aureus lawns for the indicated times (Example 1). The symbols
represent
the following meshes: P22-25 20 passes; N, P22-25 40 passes; A, P22-25 80
passes; x,
P22-27.5 20 passes; *, P22-27.5 40 passes;., P22-27.5 80 passes; and I ,
catheter.
[0020] Fig. 2 graphically depicts cumulative bupivacaine release from
multilayer
polyarylate-coated meshes.
[0021] Fig. 3 graphically depicts cumulative bupivacaine release from
multilayer
polyarylate-coated meshes having various loadings of bupivacaine. The symbols
represent
the following meshes: *, P22-27.5 (11 passes, 1 dip); N, P22-27.5 (11 passes,
2 dips); and A,
P22-27.5 (2 passes, 2 dips).
[0022] Fig. 4 graphically depicts the time course of dermal anesthesia from
1 x 2 cm
surgically implanted, polyarylate meshes containing 7.5 mg/cm2 bupivacaine.
Meshes were
implanted in rats by subcostal laparotomy, pin-prick responses were determined
and are
shown as % pain response inhibition (see Examples for details). The "*"
indicates statistically
significant response at p<0.05 compared to the baseline pin-prick response.
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[0023] Fig. 5 graphically depicts mesh stiffness. The bars, from top to
bottom, represent
the stiffness for (1) a PPM3 mesh without a polyarylate coating and without
sterilization, (2)
a ProleneTm (Ethicon) mesh sterilized with ethylene oxide, (3) a polyarylate-
coated PPM3
mesh 12 months after coating and sterilized by gamma irradiation with a
nitrogen flush, and
(4) a polyarylate-coated PPM3 mesh 12 months after coating and sterilized by
gamma
irradiation.
[0024] Fig. 6 graphically depicts the change in mesh stiffness over time
during the course
of polymer degradation for a polymer-coated polypropylene mesh soaking in PBS.
[0025] Fig. 7 depicts micrographs of a tyrosine polyarylate-coated mesh.
The top left
panel shows the woven nature of the mesh and the contact points of the
filaments. The
bottom left panel demonstrates the coating over the contact points of the mesh
filaments. The
right panel is a scanning electron micrograph of a coated filament.
[0026] Fig. 8 provides an optical image of a mesh having a tyrosine
polyarylate coating
containing rifampin and minocycline. On the left, the optical image; on the
right, a schematic
thereof indicating the areas of intense orange color by the circled areas
filled with diagonal
lines.
Detailed Description Of The Invention
[0027] The present invention is directed to medical prostheses that
comprise a mesh and
one or more coating which temporarily stiffen the mesh, preferably by at least
1.1 times its
original stiffness. The coatings on such meshes do not alter the integrity of
the mesh and thus
allow the mesh to remain porous. In general, the coatings do not substantially
alter the
porosity of the mesh. In some embodiments, the medical prosthesis comprises a
mesh with
one or more coatings with at least one of the coatings comprising a stiffening
agent that coats
the filaments or fibers of the mesh so to temporarily immobilize the contact
points of those
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filaments or fibers. Again, the coatings on such meshes do not alter the
integrity strength of
the underlying mesh. and thus allow the mesh to remain porous after coating.
In general, the
coatings do not substantially alter the porosity of the mesh. The prostheses
of the invention
are useful for surgical repair and reconstruction of soft tissue defects.
[0028] In one embodiment, the mesh of the medical prostheses can be shaped
into a
three-dimensional structure, e.g., on a mold or other form, and a coating
applied. The
coatings (and stiffening agents) have sufficient strength to hold the mesh in
that shape once
the coating has dried, cured or set, as appropriate to the particular agent,
and the mesh
removed from the mold or form. Accordingly, the present invention provides
medical
prostheses that are three dimensional structures with coatings that have
temporary stiffness in
accordance with the invention as described herein.
[0029] A mesh in accordance with the invention is any web or fabric with a
construction
of knitted, braided, woven or non-woven filaments or fibers that are
interlocked in such a
way to create a fabric or a fabric-like material. As used in accordance with
the present
invention, "mesh" also includes any porous prosthesis suitable for temporarily
stiffening.
[0030] Surgical meshes are well known in the art and any such mesh can be
coated as
described herein. The meshes used in the present invention are made from
biocompatible
materials, synthetic or natural, including but not limited to, polypropylene,
polyester,
polytetrafluroethylene, polyamides and combinations thereof. One of the
advantages of the
present invention is that the coatings can be used with any commercially
available mesh. A
preferred mesh is made from woven polypropylene. Pore sizes of meshes vary.
For example
the Bard Marlex mesh has pores of 379 +/- 143 micrometers or approx. 0.4 mm,
whereas
the Johnson and Johnson Vypro mesh has pores of 3058 +/- 62 micrometers or
approx. 3
mm.
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[0031] The stiffening agents of the invention include hydrogels,
biodegradable polymers
and any other compound capable of imparting temporary stiffness to the mesh in
accordance
with the invention. Temporary stiffness means that, relative to the
corresponding uncoated
mesh material, there is an increase in stiffness when one or more coatings are
applied in
accordance with the invention. Upon use, those coatings then soften or degrade
over time in
a manner that causes the mesh to revert back to its original stiffness, revert
nearly back to its
original stiffness or sufficient close to its original stiffness to provide
the desired surgical
outcome and the expected patient comfort. To determine if the medical
prosthesis has
temporary stiffness, the prosthesis can be evaluated in vitro or in vivo. For
example, a
coating can be applied to the mesh and then the mesh left in a physiological
solution for a
period of time before measuring its stiffness. The time period of stiffness is
controlled by the
degradation rate (for biodegradable polymers) or absorption ability (for
hydrogels). The time
period can vary from days, to weeks or even a few months and is most
conveniently
determined in vitro. Meshes with that revert to their original stiffness in
vitro within a
reasonable time (from 1 day to 3-4 months) are considered to be temporarily
stiffened.
Additionally, animal models can be used to assess temporary stiffness by
implanting the
mesh and then removing it from the animal and determining if its stiffness had
changed.
Such in vivo results can be correlated with the in vitro results by those of
skill in the art.
Methods to measure stiffness of a mesh or a coated mesh are known in the art.
[00321 A hydrogel is composed of a network of water-soluble polymer chains.
Hydrogels
are applied as coatings and dried on the mesh. Upon use, e.g., implantation in
the body, the
hydrogel absorbs water and become soft (hydrogels can contain over 99% water),
thereby
increasing the flexibility of the mesh and reverting to the original or near
original stiffness of
the mesh. Typically, hydrogels possess a degree of flexibility very similar to
natural tissue,
due to their significant water content. Common ingredients for hydrogels,
include e.g.
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polyvinyl alcohol, sodium polyacrylate, acrylate polymers and copolymers with
an
abundance of hydrophilic groups.
[0033] Meshes can have one or more polymer coatings and can optionally
include drugs
in the coatings. Meshes with a single coating are useful to improve handling
of the mesh
during surgical implantation and use. Meshes with drugs can be coated with
single or
multiple layers, depending on the amount of drug to be delivered, the type of
drug and
desired release rate. For example, a first coating layer can contain drug,
while the second
layer coating layer contains either no drug or a lower concentration of drug.
[0034] The coated implantable surgical meshes of the invention comprise a
surgical mesh
and one or more biodegradable polymer coating layers with each coating layer
optionally,
and independently, further containing a drug. The physical, mechanical,
chemical, and
resorption characteristics of the coating enhance the clinical performance of
the mesh and the
surgeon's ability to implant the device without affecting the overall or
primary performance
characteristics of the mesh, especially when used as a permanent implant in
the patient.
[0035] These characteristics are accomplished by choosing a suitable
coating thickness
for the selected biodegradable polymer.
[0036] The biodegradable coating deposited onto the surface of the mesh
gives the mesh
superior handling characteristics relative to uncoated meshes and facilitates
surgical insertion
because it imparts stiffness to the mesh and thereby improves handling
thereof. Over time,
however, the coating resorbs, or the stiffening agents degrades or softens, to
leave a flexible
mesh that provides greater patient comfort without loss of strength.
[0037] The surgical mesh can be coated with the biodegradable polymer using
standard
techniques such as spray or dip coating to achieve a uniform coating having a
thickness that
provides at least 1.1 to 4.5 and more preferably 1.25 to 2 times the stiffness
of the uncoated
mesh. In addition, the coating is optimized to provide for a uniform,
flexible, non-flaking
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layer that remains adhered to the mesh throughout the implantation and initial
wound healing
process. Typically, the polymer coating must maintain its integrity for at
least 1 week.
Optimal coating solutions are obtained by choosing a biodegradable polymer
with a solubility
between about .01 to about 30% in volatile solvents such as methylene chloride
or other
chlorinated solvents, THF, various alcohols, or combinations thereof.
Additionally, it is
preferred to use biodegradable polymers with a molecular weight between about
10,000 and
about 200,000 Daltons. Such polymers degrade at rates that maintain sufficient
mechanical
and physical integrity over about 1 week at 37 C in an aqueous environment.
[0038] Additionally, a biodegradable polymer-coated implantable mesh is
described in
which the biodegradable polymer layer (i.e., the coating) has a chemical
composition that
provide relatively good polymer-drug miscibility. The polymer layer can
contain between 1-
80% drug at room temperature as well as between 1-95%, 2-80%, 2-50%, 5-40%, 5-
30%, 5-
25% and 10-20% drug or 1, 2, 3, 4, 5, 6, 7, 8, 9, or 10% drug as well as 5%
increments from
10-95%, i.e., 10, 15, 20, 25, etc. In one embodiment, the biodegradable
polymer coating
releases drug for at least 2 ¨ 3 days. Such release is preferred, for example,
when the drug is
an analgesic to aide in localized pain management at the surgical site. Such
loading and
release characteristics can be also be obtains for drug polymer-combinations
that do not have
good miscibility by using multiple layering techniques.
[0039] To achieve an analgesic affect, the anesthetic and/or analgesic
should be delivered
to the injured tissue shortly after surgery or tissue injury. A drug or drugs
for inclusion in the
coatings of surgical meshes include, but are not limited to analgesics, anti-
inflammatory
agents, anesthetics, antimicrobial agents, antifungal agents, NSAIDS, other
biologics
(including proteins and nucleic acids) and the like. Antimicrobial and
antifungal agents can
prevent the mesh and/or the surrounding tissue from being colonized by
bacteria. One or
more drugs can be incorporated into the polymer coatings of the invention.
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[0040] In another embodiment, a mesh of the invention has a coating
comprising an
anesthetic such that the anesthetic elutes from the implanted coated mesh to
the surrounding
tissue of the surgical site for between 1 and 10 days, which typically
coincides with the
period of acute surgical site pain. In another embodiment, delivery of an
antimicrobial drug
via a mesh of the invention can create an inhibition zone against bacterial
growth and
colonization surrounding the implant during the healing process (e.g., usually
about 30 days
or less) and/or prevent undue fibrotic responses.
[0041] Using biodegradable polymer coatings avoids the issue of drug
solubility,
impregnation or adherence in or to the underlying device since a coating
having suitable
chemical properties can be deposited onto the mesh, optionally in concert with
one or more
drugs, to provide for the release of relatively high concentrations of those
drugs over
extended periods of time. For example, by modulating the chemical composition
of the
biodegradable polymer coating and the coating methodology, a clinically-
efficacious amount
of anesthetic drug can be loaded onto a mesh to assure sufficient drug elution
and to provide
surgical site, post-operative pain relief for the patient.
[0042] To provide such post-operative, acute pain relief, the mesh should
elute from
about 30 mg to about 1000 mg of anesthetic over 1-10 days, including, e.g.,
about 30, 50,
100, 200, 400, 500, 750 or 1000 mg over that time period.
[0043] The prosthesis should elute clinically effective amounts of
anesthetic during the
acute post-operative period when pain is most noticeable to the patient. This
period, defined
in several clinical studies, tends to be from 12 hours to 5 days after the
operation, with pain
greatest around 24 hours and subsiding over a period of several days
thereafter. Prior to 12
hours, the patient is usually still under the influence of any local
anesthetic injection given
during the surgery itself. After the 5-day period, most of the pain related to
the surgery itself
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(i.e., incisional pain and manipulation of fascia, muscle, & nerves) has
resolved to a
significant extent.
[0044] Bupivacaine has a known toxicity profile, duration of onset, and
duration of
action. Drug monographs recommend the daily dose not to exceed 400 mg. Those
of skill in
the art can determine the amount of anesthetic to include in a polymer coating
or a hydrogel
coating to achieve the desired amount and duration of pain relief.
[0045] There are numerous reports of reduction or complete elimination of
narcotic use
and pain scores after open hernia repair during days 2-5 with concomitant use
of catheter pain
pump system. In these cases, the pump delivers either a 0.25% or 0.5% solution
of
bupivacaine to the subfascial area (Sanchez, 2004; LeBlanc, 2005; and Lau,
2003). At a 2
mL/hour flow rate, this translates into constant "elution" of approximately
120 mg of
bupivacaine per day. However, this system purportedly suffers from leakage, so
the 120 mg
per day may only serve as an extremely rough guide for the amount of
bupivacaine that
should be delivered to provide adequate post-operative pain relief.
[0046] One of the most well characterized sustained release depot systems
for post-
operative pain relief reported in the literature is a PLGA microsphere-based
sustained release
formulation of bupivacaine. This formulation was developed and tested in
humans for relief
of subcutaneous pain as well as neural blocks. Human trials indicated that
subcutaneous
pain was relieved via injection of between 90 to 180 mg of bupivacaine which
then eluted
into the surrounding tissue over a 7-day period, with higher concentrations in
the initial 24-
hour period followed by a gradual taper of the concentration. Other depot
sustained release
technologies have successfully suppressed post-operative pain associated with
inguinal hernia
repair. For example, external pumps and PLGA microsphere formulations have
each
purportedly release drug for approximately 72 hours.
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[0047] To achieve loading at the lower limit of the elution profile, for
example, one can
choose a relatively hydrophilic biodegradable polymer and combines it with the
anesthetic
hydrochloride salt so that the anesthetic dissolves in the polymer at a
concentration below the
anesthetic's saturation limit. Such a formulation provides non-burst release
of anesthetic. To
achieve loading at the upper limit of the elution profile, one can spray coat
a layer of an
anesthetic-polymer mixture that contains the anesthetic at a concentration
above its saturation
limit. In this formulation, the polymer does not act as a control mechanism
for release of the
anesthetic, but rather acts as a binder to hold the non-dissolved, anesthetic
particles together
and alters the crystallization kinetics of the drug. A second coating layer,
which may or may
not contain further anesthetic, is sprayed on top of the first layer. When
present in the second
coating, the anesthetic concentration is at a higher ratio of polymer to
anesthetic, e.g., a
concentration at which the anesthetic is soluble in the polymer layer.
[0048] The top layer thus can serve to control the release of the drug in
the bottom layer
(aka depot layer) via the drug-polymer solubility ratio. Moreover, it is
possible to alter the
release rate of the drug by changing the thickness of the polymer layer and
changing the
polymer composition according to its water uptake. A polymer that absorbs a
significant
amount of water within 24 hours will release the contents of the depot layer
rapidly.
However, a polymer with limited water uptake or variable water uptake (changes
as a
function of its stage of degradation) will retard release of the water soluble
anesthetic agent.
[0049] Biodegradable polymers suitable for coatings of the invention
include but are not
limited to, polylactic acid, polyglycolic acid and copolymers and mixtures
thereof such as
poly(L-lactide) (PLLA), poly(D,L-lactide) (PLA);
[0050] polyglycolic acid [polyglycolide (PGA)], poly(L-lactide-co-D,L-
lactide)
(PLLA/PLA), poly(L-lactide-co-glycolide) (PLLA/PGA), poly(D, L-lactide-co-
glycolide)
CA 02637578 2008-07-07
(PLA/PGA), poly(glycolide-co-trimethylene carbonate) (PGA/PTMC), poly(D,L-
lactide-co-
caprolactone) (PLA/PCL), poly(glycolide-co-caprolactone) (PGA/PCL),
poly(oxa)ester;
[0051] polyethylene oxide (PEO), polydioxanone (PDS), polypropylene
fumarate,
poly(ethyl glutamate-co-glutamic acid), poly(tert-butyloxy-carbonylmethyl
glutamate),
polycaprolactone (PCL), polycaprolactone co-butylacrylate, polyhydroxybutyrate
(PHBT)
and copolymers of polyhydroxybutyrate, poly(phosphazene), poly(phosphate
ester),
poly(amino acid) and poly(hydroxy butyrate), polydepsipeptides, maleic
anhydride
copolymers, polyphosphazenes, polyiminocarbonates, poly[(97.5% dimethyl-
trimethylene
carbonate)-co-(2.5% trimethylene carbonate)], poly(orthoesters), tyrosine-
derived
polyarylates, tyrosine-derived polycarbonates and tyrosine-derived
polyphosphonates, PEG
derivatives, polyethylene oxide, hydroxypropylmethylcellulose, polysaccharides
such as
hyaluronic acid, chitosan and regenerate cellulose, and proteins such as
gelatin and collagen,
and mixtures and copolymers thereof, among others.
[0052] Methods of making biodegradable polymers are well known in the art.
[0053] The preferred biodegradable polymers of the invention have tyrosine-
derived
diphenol monomer units that are copolymerized with the appropriate chemical
moiety to form
a polyarylate, a polycarbonate, a polyiminocarbonate, a polyphosphonate or any
other.
[0054] The preferred biodegradable polymers are tyrosine-based polyarylates
including
those described in U.S. Patent Nos. 4,980449; 5,099,060; 5,216,115; 5,317,077;
5,587,507;
5,658,995; 5,670,602; 6,048,521; 6,120,491; 6,319,492; 6,475,477; 6,602,497;
6,852,308;
7,056,493; RE37,160E; and RE37,795E; and as well as U.S. Patent Application
Publication
Nos. 2002/0151668; 2003/0138488; 2003/0216307; 2004/0254334; 2005/0165203; and
as
well as PCT Publication Nos. W099/52962; WO 01/49249; WO 01/49311;
W003/091337.
These patents and publications also disclose other polymers containing
tyrosine-derived
diphenol monomer units, including polyarylates, polycarbonates,
polyiminocarbonates,
16
CA 02637578 2013-09-18
polythiocarbonates, polyphosphonates and polyethers. Likewise, the foregoing
patents and
publications describe methods for making these polymers, some methods of which
may be
applicable to synthesizing other biodegradable polymers. Finally, the
foregoing patents and
publications also describe blends and copolymers with polyalkylene oxide,
including
polyethylene glycol (PEG). All such polymers are contemplated for use in the
present
invention.
[0055] The representative structures for the foregoing polymers are provide
in the above-
cited patents and publications.
[0056] As used herein, DTE is the diphenol monomer desaminotyrosyl-tyrosine
ethyl
ester; DTBn is the diphenol monomer desaminotyrosyl-tyrosine benzyl ester; DT
is the
corresponding free acid form, namely desaminotyrosyl-tyrosine. BTE is the
diphenol
monomer 4-hydroxy benzoic acid-tyrosyl ethyl ester; BT is the corresponding
free acid form,
namely 4-hydroxy benzoic acid-tyrosine.
[0057] P22 is a polyarylate copolymer produced by condensation of DTE with
succinate.
P22-10, P22-15, P22-20, P22-xx, etc., represents copolymers produced by
condensation of
(1) a mixture of DTE and DT using the indicated percentage of DT (i.e., 10,
15, 20 and xx%
DT, etc.) with (2) succinate.
[0058] Additional preferred polyarylates are random copolymer of
desaminotyrosyl-
tyrosine (DT) and an desaminotyrosyl-tyrosyl ester (DT ester), wherein the
copolymer
comprises from about 0.001% DT to about 80% DT and the ester moiety can be a
branched
or unbranched alkyl, alkylaryl, or alkylene ether group having up to 18 carbon
atoms, any
group of which can, optionally have a polyalkylene oxide therein. Similarly,
another group
of polyarylates are the same as the foregoing but the desaminotyrosyl moiety
is replaced by a
4-hydroxybenzoyl moiety. Preferred DT or BT contents include those copolymers
with from
about 1% to about 30%, from about 5% to about 30% from about 10 to about 30%
DT or BT.
17
CA 02637578 2013-09-18
Preferred diacids (used informing the polyarylates) include succinate,
glutarate and glycolic
acid.
[0059] Additional biodegradable polymers useful for the present invention
are the
biodegradable, resorbable polyarylates and polycarbonates disclosed in U.S.
provisional
application Serial No. 60/733,988, filed November 3, 2005 and in its
corresponding PCT
Appin, No. PCT/US06/42944, filed November 3, 2006. These polymers, include,
but are not
limited to, BTE glutarate, DTM glutarate, DT propylamide glutarate, DT
glycineamide
glutarate, BTE succinate, BTM succinate, BTE succinate PEG, BTM succinate PEG,
DTM
succinate PEG, DTM succinate, DT N-hydroxysuccinimide succinate, DT
glucosamine
succinate, DT glucosamine glutarate, DT PEG ester succinate, DT PEG amide
succinate, DT
PEG ester glutarate and DT PEG ester succinate.
[0060] The most preferred polyarylates are the DTE-DT succinate family of
polymers,
e.g., the P22-xx family of polymers having from 0-50%, 5-50%, 5-40%, 1-30% or
10-30%
DT, including but not limited to, about 1, 2, 5, 10, 15, 20, 25, 27.5, 30, 35,
40%, 45% and
50% DT.
[0061] Additionally, the polyarylate polymers used in the present invention
can have
from 0.1-99.9 % PEG diacid to promote the degradation process.
[0062] The prostheses of the invention can be used to reconstruct,
reinforce, bridge,
replace, repair, support, stabilize, position or strengthen any soft tissue
defect.
[0063] For example, soft tissue defects that can be treated in accordance
with the instant
invention include hernias, including but not limited to inguinal, femoral,
umbilical,
abdominal, incisional, intramuscular, diphragmatic, abdornino-throacic and
thoracic hernias.
The prosetheses of the invention can also be used for structural reinforcement
for muscle
flaps, to provide vascular integrity, for ligament repair/replacement and for
organ
18
CA 02637578 2008-07-07
support/positioning/repositioning such as done with a bladder sling, a breast
lift, or an organ
bag/wrap. The prosetheses of the invention can be used in recontruction
procedures
involving soft tissue such as an orthopaedic graft support/stabilization, as
supports for
reconstructive surgical grafts and as supports for bone fractures.
[0064] Examples of drugs suitable for use with the present invention
include anesthetics,
antibiotics (antimicrobials), anti-inflammatory agents, fibrosis-inhibiting
agents, anti-scarring
agents, leukotriene inhibitors/antagonists, cell growth inhibitors and the
like. As used herein,
"drugs" is used to include all types of therapeutic agents, whether small
molecules or large
molecules such as proteins, nucleic acids and the like. Those of skill in the
art can readily
determine the amount of a particular drug to include in the coatings on the
meshes of the
invention.
[0065] Any pharmaceutically acceptable form of the drugs of the present
invention can be
employed in the present invention, e.g., the free base or a pharmaceutically
acceptable salt or
ester thereof. Pharmaceutically acceptable salts, for instance, include
sulfate, lactate, acetate,
stearate, hydrochloride, tartrate, maleate, citrate, phosphate and the like.
[0066] Examples of non-steroidal anti-inflammatories include, but are not
limited to,
naproxen, ketoprofen, ibuprofen as well as diclofenac; celecoxib; sulindac;
diflunisal;
piroxicam; indomethacin; etodolac; meloxicarn; r-flurbiprofen; mefenamic;
nabumetone;
tolmetin, and sodium salts of each of the foregoing; ketorolac bromethamine;
ketorolac
bromethamine tromethamine; choline magnesium trisalicylate; rofecoxib;
valdecoxib;
lumiracoxib; etoricoxib; aspirin; salicylic acid and its sodium salt;
salicylate esters of alpha,
beta, gamma-tocopherols and tocotrienols (and all their d, 1, and racemic
isomers); and the
methyl, ethyl, propyl, isopropyl, n-butyl, sec-butyl, t-butyl, esters of
acetylsalicylic acid.
[0067] Examples of anesthetics include, but are not limited to, licodaine,
bupivacaine,
and mepivacaine. Further examples of analgesics, anesthetics and narcotics
include, but are
19
CA 02637578 2008-07-07
not limited to acetaminophen, clonidine, benzodiazepine, the benzodiazepine
antagonist
flumazenil, lidocaine, tramadol, carbamazepine, meperidine, zaleplon,
trimipramine tnaleate,
buprenorphine, nalbuphine, pentazocain, fentanyl, propoxyphene, hydromorphone,
methadone, morphine, levorphanol, and hydrocodone.
[0068] Examples of antimicrobials include, but are not limited to,
triclosan,
chlorhexidine, rifampin, minocycline, vancomycin, gentamycine, cephalosporins
and the like.
In preferred embodiments the coatings contain rifampin and another
antimicrobial agent. In
another preferred embodiment, the coatings contains a cephalosporin and
another
antimicrobial agent. Preferred combinations include rifampin and minocycline,
rifampin and
gentamycin, and rifampin and minocycline.
[0069] Further antimicrobials include aztreonam; cefotetan and its disodium
salt;
loracarbef; cefoxitin and its sodium salt; cefazolin and its sodium salt;
cefaclor; ceftibuten
and its sodium salt; ceftizoxime; ceftizoxime sodium salt; cefoperazone and
its sodium salt;
cefuroxime and its sodium salt; cefuroxime axetil; cefprozil; ceftazidime;
cefotaxime and its
sodium salt; cefadroxil; ceftazidime and its sodium salt; cephalexin;
cefamandole nafate;
cefepime and its hydrochloride, sulfate, and phosphate salt; cefdinir and its
sodium salt;
ceftriaxone and its sodium salt; cefixime and its sodium salt; cefpodoxime
proxetil;
meropenem and its sodium salt; imipenem and its sodium salt; cilastatin and
its sodium salt;
azithromycin; clarithromycin; dirithromycin; erythromycin and hydrochloride,
sulfate, or
phosphate salts ethylsuccinate, and stearate forms thereof; clindamyein;
clindamycin
hydrochloride, sulfate, or phosphate salt; lincomycin and hydrochloride,
sulfate, or phosphate
salt thereof; tobramycin and its hydrochloride, sulfate, or phosphate salt;
streptomycin and its
hydrochloride, sulfate, or phosphate salt; vancomycin and its hydrochloride,
sulfate, or
phosphate salt; neomycin and its hydrochloride, sulfate, or phosphate salt;
acetyl
sulfisoxazole; colistimethate and its sodium salt; quinupristin; dalfopristin;
amoxicillin;
CA 02637578 2008-07-07
ampicillin and its sodium salt; clavulanic acid and its sodium or potassium
salt; penicillin G;
penicillin G benzathine, or procaine salt; penicillin G sodium or potassium
salt; carbenicillin
and its disodium or indanyl disodium salt; piperacillin and its sodium salt;
ticarcillin and its
disodium salt; sulbactam and its sodium salt; moxifloxacin; ciprofloxacin;
ofloxacin;
levofloxacins; norfloxacin; gatifloxacin; trovafloxacin mesylate;
alatrofloxacin mesylate;
trimethoprim; sulfamethoxazole; demeclocycline and its hydrochloride, sulfate,
or phosphate
salt; doxycycline and its hydrochloride, sulfate, or phosphate salt;
minocycline and its
hydrochloride, sulfate, or phosphate salt; tetracycline and its hydrochloride,
sulfate, or
phosphate salt; oxytetracycline and its hydrochloride, sulfate, or phosphate
salt;
chlortetracycline and its hydrochloride, sulfate, or phosphate salt;
metronidazole; dapsone;
atovaquone; rifabutin; linezolide; polymyxin B and its hydrochloride, sulfate,
or phosphate
salt; sulfacetamide and its sodium salt; and clarithromycin.
[0070] Examples of antifungals include amphotericin B; pyrimethamine;
flucytosine;
caspofungin acetate; fluconazole; griseofulvin; terbinafin and its
hydrochloride, sulfate, or
phosphate salt; ketoconazole; micronazole; clotrimazole; econazole;
ciclopirox; naftifine; and
itraconazole.
[0071] Other drugs that can be incorporated into surgical meshes include,
but are not
limited to, keflex, acyclovir, cephradine, malphalen, procaine, ephedrine,
adriamycin,
daunomycin, plumbagin, atropine, quinine, digoxin, quinidine, biologically
active peptides,
cephradine, cephalothin, cis-hydroxy-L-proline, melphalan, penicillin V,
aspirin, nicotinic
acid, chemodeoxycholic acid, chlorambucil, paclitaxel, 5-flurouracil and the
like.
[0072] Examples of useful proteins include cell growth inhibitors such as
epidermal
growth factor.
[0073] Examples of anti-nflammatory compound include, but are not limited
to,
anecortive acetate; tetrahydrocortisol, 4,9(11)-pregnadien-17.alpha.,21-dio1-
3,20-dione and
21
CA 02637578 2008-07-07
its -21-acetate salt; 11-epicortisol; 17.alpha.-hydroxyprogesterone;
tetrahydrocortexolone;
cortisona; cortisone acetate; hydrocortisone; hydrocortisone acetate;
fludrocortisone;
fludrocortisone acetate; fludrocortisone phosphate; prednisone; prednisolone;
prednisolone
sodium phosphate; methylprednisolone; methylprednisolone acetate;
methylprednisolone,
sodium succinate; triamcinolone; triamcinolone-16,21-diacetate; triamcinolone
acetonide and
its -21-acetate, -21-disodium phosphate, and -21-hemisuccinate forms;
triamcinolone
benetonide; triamcinolone hexacetonide; fluocinolone and fluocinolone acetate;
dexamethasone and its 21-acetate, -21-(3,3-dimethylbutyrate), -21-phosphate
disodium salt, -
21-diethylaminoacetate, -21-isonicotinate, -21-dipropionate, and -21-palmitate
forms;
betamethasone and its -21-acetate, -21-adamantoate, -17-benzoate, -17,21-
dipropionate, -17-
valerate, and'-21-phosphate disodium salts; beclomethasone; beclomethasone
dipropionate;
diflorasone; diflorasone diacetate; mometasone furoate; and acetazolamide.
[0074] Those of ordinary skill in the art will appreciate that any of the
foregoing
disclosed drugs can be used in combination or mixture in coatings of the
present invention.
Methods
[0075] Another aspect of the invention is directed to a process for coating
a mesh with a
stiffening agent that coats the filaments or fibers of the mesh to temporarily
immobilize
contact points of the filaments or fibers of said mesh. The method is
comprises (a) preparing
a coating solution comprising a solvent and said stiffening agent; (b)
spraying a mesh one or
more times to provide an amount of said solution on said mesh to produce a
coating having a
thickness and placement sufficient to temporarily immobilize contact points of
the filaments
or fibers of said mesh that coats filaments or fibers; and (c) drying said
mesh to produce said
coating. An example of ratio of coating thickness to polymer coating is shown
in the
22
CA 02637578 2008-07-07
scanning electron micrograph of Fig. 7. When used with a drug (or combination
of drugs),
the drug is included in the coating solution at the desired concentration.
[0076] Spraying can be accomplished by known methods. For example, the
coating can
be applied to the entire mesh or to that portion of the mesh necessary to
stiffen it. One
technique is to dip the mesh in the coating material; another is to push the
mesh through
rollers that transfer the coating on the mesh. Spraying the mesh with a
microdroplets is also
effective. Techniques for selectively coating only those areas necessary to
stiffen the mesh
include deposition the coating through a template that exposes only the
desired areas of
coverage for the coating, including dispensing the coating with micro needles
or similar
means. More preferably the coating can be applied using a photoresist-like
mask that expose
the desired portions, applying the coating over the photomask and the removing
the
photomask.
[0077] Still another aspect of the invention relates to a method for
producing a shaped or
three-dimensional medical prosthesis mesh by forming a mesh into a desired
shape or three-
dimensional shape. This step can be accomplished with the aid of a mold or
other form on
which to affix and shape the mesh or by holding the mesh in a frame in the
desired shape or
structural configuration. Once configured into the desired shape, a stiffening
agent of the
invention is applied to the mesh to coat filaments or fibers and the agent is
allowed to dry, set
or cure as appropriate, causing the the mesh stiffen and hold the desired
shape. The mesh is
then removed or released from the mold, form or device that had been holding
the mesh to
produce the three dimensional medical prosthesis which is capable of retaining
its shape
without any further structural support or aid.
[0078] It will be appreciated by those skilled in the art that various
omissions, additions
and modifications may be made to the invention described above without
departing from the
scope of the invention, and all such modifications and changes are intended to
fall within the
23
CA 02637578 2013-09-18
=
scope of the invention, as defined by the appended claims.
Example 1
Antibiotic Release from DTE-DT succinate Coated Mesh
A. Preparation of mesh by spray-coating
[0079] A 1% solution containing a ratio of 1:1:8
rifampin:minocycline:polymer in 9:1
tetrahydrofuran/methanol was spray-coated onto a surgical mesh by repeatedly
passing the
spray nozzle over each side of the mesh until each side was coated with at
least 10 mg of
antimicrobial-embedded polymer. Samples were dried for at least 72 hours in a
vacuum oven
before use.
[0080] The polymers are the polyarylates P22-xx having xx being the % DT
indicated in
Table I. In Table 1, Rxx or Mxx indicates the percentage by weight of rifampin
(R) or
minocycline (M) in the coating, i.e., RI OMIO means 10% rifampin and 10%
minocycline
hydrochloride with 80% of the indicated polymer. Table 1 provides a list of
these
polyarylates with their % DT content, exact sample sizes, final coating
weights and drug
coating weights.
24
CA 02637578 2008-07-07
Table I
Polyarylate Coated Meshes with Rifampin and Minocycline
Sample Coating Parameters Avg. Coating
Coating Wt. Rifampin Minocycline
No. (No. Spray Passes) Wt. per 116 cm2 per cm2 (lig)
HCI
(mg) (mg) (jig)
1 P22-25 R10M10 100 0.86 86 86
(20)
2 P22-25 RIOM10 150 1.29 129 129
_____ (40)
3 P22-25 R10M10 200 1.72 172 172
(80)
4 P22-27.5 R10M10 20 0.17 17 17
_____ (1)
P22-27.5 RIOMIO 40 0.34 34 34
(2)
6 P22-27.5 R10M10 60 0.52 52 52
(3)
B. Zone of Inhibition (Z01) Studies
[0081] The ZOI for antibiotic coated meshes was determined according to the
Kirby-
Bauer method. Staphylococcus epiderrnidis or Staphylococcus aureus were
inoculated into
Triplicate Soy Broth (TSB) from a stock culture and incubated at 37 C until
the turbidity
reached McFarland # 0.5 standard (1-2 hours). Plates were prepared by
streaking the
bacteria onto on Mueller-Hinton II agar (MHA) three times, each time swabbing
the plate
from left to right to cover the entire plate and rotating the plate between
swabbing to change
direction of the streaks.
[0082] A pre-cut piece (1-2 cm2) of spray-coated mesh was firmly pressed
into the center
of pre-warmed Mueller Hinton II agar plates and incubated at 37 C. Pieces were
transferred
every 24 h to fresh, pre-warmed Mueller Hinton IT agar plates using sterile
forceps. The
distance from the sample to the outer edge of the inhibition zone was measured
every 24 h
and is reported on the bottom row in Table 2 and 3 for each sample. The top
row for each
sample represents difference between the diameter of the ZOI and the diagonal
of the mesh.
CA 02637578 2013-09-18
Table 2 shows the ZOI results for meshes placed on S. epidermidis lawns and
Table 3 show s
the ZOI results for meshes placed on S. aureus lawns. Additionally, three
pieces were
removed every 24 h for analysis of residual minocycline and rifampin.
[0083] Fig. 1 shows the total ZOI on S. aureus for meshes with 10% each of
minocycline
hydrochlorideand rifampin in a DTE-DT succinate polyarylatc coating having 25%
or 27.5%
TM
DT. The catheter is a COOK SPECTRUM venous catheter impregnated with rifampin
and
minocycline hydrochloride.
Table 2
S. epidermidis ZOI
Sample Coating Parameters Day 1 Day 2
Day 3 Day 4 Day 6 Day 7
No. (mm) (mm)
(mm) (mm) (mm) (mm)
1 P22-25 R10M10 18.65 31.70
33.04 29.63 25.43 15.66
31.30 44.36 45.70 42.29 38.08 28.31
2 P22-25 RIOMIO 19.28 30.59
33.67 31.74 0.60 8.56
32.10 43.45 46.53 44.60 13.45 21.42
3 P22-25 RIOMIO 26.59 34.70 30.31 31.75 23.65 17.29
39.48 47.59 43.20 46.16 36.54 30.18
4 P22-27.5 RIOMIO 18.33 31.58
35.25 30.45 2.08 6.72
31.06 44.31 47.98 43.18 14.81
19.45
P22-27.5 R10M10 17.48 32.81 33.68
28.06 7.89 12.86
30.17 45.51 46.38 40.76 20.59 25.56
6 P22-27.5 RIOM10 31.73 29.81
35.03 24.99 12.55 16.22
44.42 42.50 47.72 37.68 25.24 28.91
26
CA 02637578 2008-07-07
Table 3
S. aureus ZOI
Sample Coating Parameters Day 1 Day
2 Day 3 Day 4 Day 5 Day 7
No. (mm) (mm)
(mm) (mm) (mm) (mm)
1 P22-25 RIOMIO 12.75
17.90 18.22 22.44 12.35 11.94
25.84 30.66 30.97 35.20 25.11 24.69
2 P22-25 R10M10 14.23
11.28 20.04 28.24 16.31 10.35
26.90 23.94 32.71 40.91 28.98 23.02
3 P22-25 R10M10 17.87
21.52 23.45 25.36 17.42 14.72
30.57 34.22 36.15 36.02 30.12 27.42
4 P22-27.5 RIOMIO 9.77
19.02 19.06 23.01 13.81 5.61
22.76 32.01 32.05 36.00 26.80 18.6
P22-27.5 RIOM10 9.70 21.77 19.55
24.00 11.84 3.89
22.30 34.36 35.48 36.60 24.44 16.49
6 P22-27.5 R1OM10 20.92
21.29 22.40 24.27 11.06 4.99
33.68 34.05 35.15 37.02 23.82 17.75
10084] Table 4 shows that the duration of in vitro drug release increases
with the
hydrophilicity of the resorbable polymer. Solvent cast films were soaked in
PBS and
antibiotic release was monitored by HPLC.
Table 4
Antibiotic Release as a Function of Polymer Hydrophilicity
Films Days releasing Days releasing
Rifampin MinocyclineHC1
P22-15 RIOM10 32 32
P22-20 R10M10 25 25
P22-25 RIOM10 7 7
P22-27.5 R10M10 10 10
P22-30 R10M10 4 4
27
CA 02637578 2008-07-07
Example 2
Bupivacaine Release from DTE-DT succinate Coated Mesh
A. Preparation of mesh
[0085] For the experiment shown in Fig. 2, a first depot coating containing
540 mg of
bupivacaine fIC1 as a 4% solution with 1% P22-27.5 polyarylate in a mixture of
TI-IF
Methanol was spray coated onto a mesh. A second layer consisting of 425 mg of
the same
polyarylate alone was deposited on top of the first layer.
[0086] For the experiment shown in Fig. 3, a solution of approximately 4%
bupivacaine
in DTE-DT succinate polymer having 27.5% DT was sprayed onto a mesh using the
indicted
number of passes followed by the indicated number of dips into a solution of
the same
polyarylate in THF:Methanol (9:1)
B. Anesthetic Release
[0087] Pre-weighed pieces of mesh were placed in PBS at 37 C and a sample
withdrawn
periodically for determination of bupivacaine by HPLC. Fig. 2 shows the
cumulative release
of bupivacaine into PBS from the multilayer polyarylate coating as a function
of time.
Nearly 80% of the bupivacaine had been released after 25 hours of incubation.
[0088] Figure 3 is an example of the changes in release characteristics
that can be
achieved by altering both the amount of drug in the depot layer and the
thickness of the outer
layer. These coated surgical meshes are much stiffer than their uncoated
counterparts.
Example 3
In Vivo Bupivacaine Release from DTE-DT succinate coated Meshes
A. Overview
[0089] Rats with jugular cannulas for pharmacokinetic studies were
surgically implanted
with a 1 x 2 cm P22-27.5 polyarylate-coated mesh containing 7.5 mg of
bupivacaine/cm2.
Before surgery, baseline pin-prick responses to nociception were measured at
the planned
28
CA 02637578 2008-07-07
surgical incision site, and baseline blood samples were obtained. A hernia was
created by
incision into the peritoneal cavity during via subcostal laparotomy, and a
Lichtenstein non-
tension repair was performed using the bupivacaine-impregnated polyarylate-
coated mesh.
Blood samples were drawn at 3, 6, 24, 48, 72, 96, and 120 hours after
implantation. Prior to
drawing blood, the rats were subjected to a pin prick test to assess dermal
anesthesia from
bupivacaine release. The behavioral results indicate that moderate levels of
dermal
anesthesia appeared from 3 to 120 hours, with the amount at 6 and 48 hours
significantly
above baseline (p<0.05). Pharmacokinetic analysis indicates that the plasma
bupivacaine
levels fit a one-compartment model with first-order absorption from 0 to 24
hours.
B. Preparation of Surgical Mesh
[0090] A polypropylene mesh was spray coated as described in the first
paragraph of
Example 2. Individual meshes were cut to 1 x 2 cm, individually packaged, and
sterilized by
gamma irradiation. The mesh was loaded with 7.5 mg/cm2 of bupivacaine HC1 for
a total of
15 mg of bupivacaine loaded per 1 x 2 cm mesh.
C. Surgical Implantation of Mesh
[0091] Eight male rats, 59-63 days old and weighing from 250-275 g, were
obtained from
Taconic Laboratory (Germantown, NY) with an external jugular cannula (SU007).
Each rat
was anesthetized with isoflurane to a plane of surgical anesthesia, as
determined by the
absence of a response to toe pinch and corneal reflex and maintained at 2%
isoflurane during
surgery. The subcostal site was shaved, washed with 10% providone iodine and
rinsed with
70% ethanol. Sterile drapes were used to maintain an aseptic surgical field,
and sterilized
instruments were re-sterilized between rats using a hot-bead sterilizer. A 2.5
cm skin incision
was made 0.5 cm caudal to and parallel to the last rib. The underlying
subcutaneous space (1
29
CA 02637578 2013-09-18
cm on both sides of the incision) was loosened to accommodate the mesh. A 2 cm
incision
was made through the muscle layers along the same plane as the skin incision,
penetrating the
TM
peritoneal cavity and the peritoneum was closed with 6-0 Prolene sutures in a
continuous
suture pattern. Rather than suturing the inner and outer oblique muscles using
the classic
"tension closure," a Lichtenstein "non-tension" repair was undertaken using
the mesh as the
repair material. The mesh prepared in Section A was positioned over the
incisional hernia,
and sutured into the internal and external oblique muscles using 6-0 Prolene
sutures. The
subcutaneous tissue was then sutured in a continuous pattern with 6 to 8 6-0
Prolene sutures
to prevent the rats from accessing the mesh, followed by 6 to 8 skin sutures.
Total surgical
time was 10 min for anesthetic induction and preparation and 20 min for the
surgery.
[0092] The rats were allowed to recover in their home cages, and monitored
post-
surgically until they awoke. Blood samples were drawn for determination of
plasma
bupivacaine levels at 3, 6, 24, 48, 72, 96, and 120 hours after surgery. The
rats were assessed
for guarding the incision, and the incision was assessed for signs of
inflammation, swelling or
other signs of infection. No rats exhibited toxicity or seizures, or were in a
moribund state
from infection or the release of bupivacaine.
D. Dermal Anesthetic Tests
[0093] The nociceptive pin prick test was used to assess dermal anesthesia
(Morrow and
Casey, 1983; Kramer et al., 1996; Haynes et al., 2000; Khodorova and
Strichartz, 2000).
Holding the rat in one hand, the other hand was used to apply the pin.
Nociception was
indicated by a skin-flinch or by a nocifensive (i.e., startle or attempt to
escape) response from
the rat. While the presence of the mesh interfered with the skin flinch
response, nocifensive
response remained completely intact.
CA 02637578 2008-07-07
100941 Baseline
nocifensive responses to 10 applications of the pin from a Buck
neurological hammer were obtained at the planned incision site prior to mesh
implantation.
After surgery, the pin prick test was applied rostral to the incision. The
nerves caudal to the
incision were transected during the procedure, and therefore did not respond
to pin
application and were not tested. The post-implantation test was repeated using
the same
force as before surgery and with 10 pin applications, and the percent
inhibition of nocifensive
responding was calculated by: [1 ¨ (test responses/10 base responses)] X 100.
The data was
analyzed using repeated measures ANOVA followed by post hoc analysis using the
Tukey's
test. The results are shown in Fig. 4.
Example 4
Mesh Stiffness
[00951 A. Meshes prepared as described in Example 1 were subject to
stiffness testing
according to the method of TyRx Pharma Inc. Mesh Stiffness Test Protocol, ATM
0410,
based on ASTM 4032-94. Meshes were sealed in foil bags before sterilization
using gamma
irradiation. Where indicated by "Gamma N2", the bags were flushed with
nitrogen before
sealing and irradiation. Meshes were tested in triplicate. The results are
shown in Table 5
and indicate that aging does not affect the flexibility of the coated meshes.
Table 5
Stiffness Testing
Sample 1 Sample 2 Sample 3 Average
Mesh
(Newtons) (Newtons) (Newtons) (Newtons) t-test
PPM3, Gamma, 12 month aged
coating 1.84 2.36 1.62 1.94 0.016
PPM3, Gamma N2 flush, 12 month
aged coating 2.2 2.24 2.56 2.3 0.014
Prolene, Ethylene oxide
sterilization 2.78 2.16 1.94 2.29 0.019
PPM3, No Sterilization, No
Coating 1.2 1.3 1 1.17
31
CA 02637578 2008-07-07
[0096] B. Meshes were prepared by spray coating a solution of P22-27.5 onto
a PPM3
mesh as generally described in Example 1. the coated meshes were cut into 3"
by 3" squares
to provide 80 mg polymer coating per square. The squares were incubated in 1 L
of 0.01 M
PBS for the indicated times then removed for stiffness testing as described in
part A of this
Example. All experiments were done in triplicate. As a control, non-coated
PPM3 meshes
were incubated under the same conditions. The stiffness of the control when
dry was 1.42
0.23 N when dry and 1.12 N after both 1 hour and 24 hour in 0.01 M PBS. The
results are
shown in Fig. 6.
Example 5
Micrographs of Coated Meshes
[0097] A tyrosine polyarylate-coated mesh without antibiotics, i.e., only a
polymer
coating, was prepared as described in Example 1 and omitting the antibiotics
in the spray
coating solution. An optical image of the coated mesh is shown in the top left
panel of Fig. 7
at a magnification that readily shows the woven nature of the mesh and the
contact points of
the filaments. A close up of a contact point is shown in the bottom left panel
of Fig. 7 and
demonstrates that the coating immobilizes the contact points of the mesh
filaments. The right
panel of Fig. 7 is a scanning electron micrograph of a coated filament.
[0098] Fig. 8 shows an optical image of a mesh from Example 1, i.e., coated
with
polymer, rifampin and minocycline. In color, this photograph shows the mesh on
a blue
background with the filaments appearing greenish with some orange and the
knots (or
filament contact points) appearing mostly solid orange. The orange color is
due to the
antibiotics and is more visible on the knots due to the greater surface area
of the mesh in that
region. The color differentiation is difficult to visualize in the black and
white version of this
32
CA 02637578 2008-07-07
photograph so on the right panel the areas of orange are indicated by circled
areas filled with
diagonal lines.
References
[00991 Hayes, B.B., Afshari, A., Millecchia, L., Willard, P.A., Povoski,
S.P., Meade,
B.J., 2000. Evaluation of percutaneous penetration of natural rubber latex
proteins. Toxicol.
Sci. 56, 262-270.
[00100] Khodorova, A.B., Strichartz, G.R., 2000. The addition of dilute
epinephrine
produces equieffectiveness of bupivacaine enantiorners for cutaneous analgesia
in the rat.
Anesth. Analg. 91, 410-416.
[00101] Kramer, C., Tawney, M., 1998. A fatal overdose of transdermally
administered
fentanyl. J. Am. Osteopath. Assoc. 98, 385-386.
[001021 Lau H, Patil NG, Lee F. Randomized clinical trial of postoperative
subfascial
infusion with bupivacaine following ambulatory open mesh repair of inguinal
hernia. Dig
Surg. 2003;20(4):285-9.
[001031 LeBlanc KA, Bellanger D, Rhynes VK, Hausmann M Evaluation of a
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Following Open Inguinal Hernia Repair. J Am Coll Surg 2005;200(2):198-202.
[00104] Morrow, T.J., Casey, K.L., 1983. Suppression of bulboreticular unit
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[00105] Sanchez B, Waxman K. Local anesthetic infusion pumps improve
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33
