Note : Les descriptions sont présentées dans la langue officielle dans laquelle elles ont été soumises.
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ENHANCED STABILITY OF INVERSE THERMAL GELLING COMPOSITE
HYDROGELS
FIELD
The present invention relates to a composite hydrogel comprising a blend of an
aqueous solution of
an anionic polysaccharide or a derivative thereof, such as hyaluronan (also
commonly referred to as
hyaluronic acid) or a derivative thereof and an aqueous solution of
methylcellulose or another water
soluble cellulose derivative thereof, having dispersed polymeric particles,
such as polymeric
hydrophobic particles therein selected from micro particles and nanoparticles,
and wherein the
stability of the hydrogel is enhanced relative to the stability of the
hydrogel alone. The polymeric
particles may contain at least one therapeutic agent, in which case each
therapeutic agent exhibits a
linear sustained release rate that can be tuned or altered by selecting the
appropriate polymer
formulation of the micro particles and/or nanoparticles. The composite may be
injectable, and in the
absence of a therapeutic agent may be used as a bulking agent for
reconstructive and cosmetic
surgery or may act as a platform for subsequent delivery of therapeutic
agents.
BACKGROUND
About 11,000 new cases of traumatic spinal cord injury (SCI) are reported in
the United States
annually, primarily affecting young adults [Wosnick, J., Baumann, M.D.,
Shoichet, M.S., Tissue
Therapy in the Central Nervous System, in Principles of Regenerative Medicine,
A. Atala, Lanza,
R., Thomson, LA., Nerem, R.M., eds., Editor. 2007, Elsevier: New York]. A
majority of these cases
are compression injuries wherein the cord is bruised under displacement of the
spinal column,
resulting in formation of a cystic cavity in the days after injury. As tissue
degenerates, the degree of
paralysis increases, causing further permanent loss of motor control and
sensory perception. For
this reason compression injuries are normally described as occurring in two
stages, the immediate
primary injury and subsequent secondary injury. Various treatment strategies
are being developed
with a view of limiting degeneration after the primary injury and! or
promoting regeneration after
secondary injury. Currently, however, there is no standard clinical treatment,
other than application
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of methylprednisolone, the efficacy of which is still debated [Miller, S.M.,
Methylprednisolone in
acute spinal cord injury: a tarnished standard. J Neurosurg Anesthesiol, 2008.
20(2): P. 140-2;
Rozet, I., Methylprednisolone in acute spinal cord injury: is there any other
ethical choice? J
Neurosurg Anesthesiol, 2008. 20(2): p. 137-9]. Similarly, there is no standard
of care for traumatic
brain injury or stroke. For example, there is no cure for stroke, and the only
FDA approved
treatment is tissue plasminogen activator (tPA), a thrombolytic agent with
limited therapeutic
benefit [Stroke and cerebrovascular accidents. World Health Organization,
Circulation, 2009].
Therapies designed to enhance cell survival during the trauma of secondary
injury are focused on
the hours to days following the primary injury and seek to limit vascular
damage, excitotoxicity, and
the inflammatory response around the injury site [Ramer, L.M., M.S. Ramer, and
J.D. Steeves,
Setting the stage for functional repair of spinal cord injuries: a cast of
thousands. Spinal Cord,
2005. 43(3): p. 134-61.]. Neuroprotective strategies target one or more of
these mechanisms with
the goal of minimizing the death of motor and sensory neurons. For example,
methylprednisolone
targets acute inflammation and inhibits lipid peroxidation [Bracken, M.B., et
al.,
Methylprednisolone or naloxone treatment after acute spinal cord injury: 1-
year follow-up data.
Results of the second National Acute Spinal Cord Injury Study. J Neurosurg,
1992. 76(1): p. 23-31],
while the sodium channel antagonist NBQX minimizes excitotoxicity [Li, Y., et
al., Effects of the
AMPA receptor antagonist NBQX on the development and expression of behavioral
sensitization to
cocaine and amphetamine. Psychopharmacology (Berl), 1997. 134(3): p. 266-76]
and nimodipine
limits vasospasm [Scriabine, A., T. Schuurman, and J. Traber, Pharmacological
basis for the use of
nimodipine in central nervous system disorders. Faseb J, 1989. 3(7): p. 1799-
806].
Neuroregenerative therapies enhance axonal outgrowth by either direct action
or suppression of the
inhibitory environment after injury. For example, numerous neurotrophins
stimulate proliferation
and regeneration, including: nerve growth factor [Romero, M.I., et al.,
Functional regeneration of
chronically injured sensory afferents into adult spinal cord after
neurotrophin gene therapy. J
Neurosci, 2001. 21(21): P. 8408-16], brain derived neurotrophic factor
[Tobias, C.A., et al., Delayed
grafting of BDNF and NT-3 producing fibroblasts into the injured spinal cord
stimulates sprouting,
partially rescues axotomized red nucleus neurons from loss and atrophy, and
provides limited
regeneration. Exp Neurol, 2003. 184(1): p. 97-113], epidermal growth factor
(EGF) [Kitchens, D.L.,
E.Y. Snyder, and D.I. Gottlieb, FGF and EGF are mitogens for immortalized
neural progenitors. J
Neurobiol, 1994. 25(7): p. 797-807] and basic fibroblast growth factor (FGF-2)
[Bikfalvi, A., et al.,
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Biological roles of fibroblast growth factor-2. Endocr Rev, 1997. 18(1): P. 26-
45]. FGF-2 has also
been reported to prevent neuronal cell death [Lee, T.T., et al.,
Neuroprotective effects of basic
fibroblast growth factor following spinal cord contusion injury in the rat. J
Neurotrauma, 1999.
16(5): p. 347-56; 13; Teng, Y.D., et al., Basic fibroblast growth factor
increases long-term survival
of spinal motor neurons and improves respiratory function after experimental
spinal cord injury. J
Neurosci, 1999. 19(16): p. 7037-47] and promote angiogenesis [Loy, D.N., et
al., Temporal
progression of angiogenesis and basal lamina deposition after contusive spinal
cord injury in the
adult rat. J Comp Neurol, 2002. 445(4): p. 308-24]. The family of antibodies
targeting NogoA
[Schwab, M.E., Nogo and axon regeneration. Curr Opin Neurobiol, 2004. 14(1):
p. 118-24], rho
kinase inhibitors [McKerracher, L. and H. Higuchi, Targeting Rho to stimulate
repair after spinal
cord injury. J Neurotrauma, 2006. 23(3-4): p. 309-17] and cyclic AMP [Hannila,
S.S. and M.T.
Filbin, The role of cyclic AMP signaling in promoting axonal regeneration
after spinal cord injury.
Exp Neurol, 2008. 209(2): p. 321-32] are well known anti-inhibitory molecules
that act by blocking
or overriding the inhibitory environment present post-injury. Similarly,
chondroitinase abc requires
local delivery as it cannot cross the blood-spinal cord barrier (or the blood-
brain barrier) and
requires sustained delivery, which is not easily obtained by other delivery
methods. Chondroitinase
abc acts to degrade the chondroitin sulfate proteoglycan present in the
injured central nervous
system and thereby facilitates axonal regeneration. These molecules are often
delivered for extended
periods, ranging from 7-28 days.
Whether neuroprotective or neuroregenerative, delivery is limited to local
strategies as most
molecules are unable to cross the blood-spinal cord barrier and blood-brain
barrier, confounding
systemic delivery. Current local delivery strategies are inadequate: bolus
delivery often results in
rapid clearance due to cerebrospinal fluid flow in the intrathecal space
[Terada, H., et al., Reduction
of ischemic spinal cord injury by dextrorphan: comparison of several methods
of administration. J
Thorac Cardiovasc Surg, 2001. 122(5): p. 979-85; 19; Yaksh, T.L., et al.,
Intrathecal ketorolac in
dogs and rats. Toxicol Sci, 2004. 80(2): p. 322-34], whereas the indwelling
catheter/external pump
is associated with scarring and infection [Jones, L.L. and M.H. Tuszynski,
Chronic intrathecal
infusions after spinal cord injury cause scarring and compression. Microscopy
Research and
Technique, 2001. 54(5): p. 317-324]. With a view toward developing a minimally-
invasive drug
delivery system that would provide sustained, local release of factors, a
delivery paradigm is
presented in which a drug loaded thermo-sensitive hydrogel is injected
intrathecally and remains
localized at the site of injection, delivering the drug load to the cerebral
spinal fluid (CSF) with
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concomitant access to the brain and spinal cord [Jimenez Hamann, M.C., et al.,
Novel intrathecal
delivery system for treatment of spinal cord injury. Exp Neurol, 2003. 182(2):
p. 300-9] and then
biodegrading, has been described. In this manner the hydrogel provides a
platform for localized
release over the life of the material. Evidence shows that intrathecal
injection bypasses the dura and
arachnoid mater and limits convective drug redistribution from CSF flow, all
barriers that negatively
impact epidural delivery [Chvatal, S.A., et al., Spatial distribution and
acute anti-inflammatory
effects of Methylprednisolone after sustained local delivery to the contused
spinal cord.
Biomaterials, 2008. 29(12): p. 1967-75]. Subsequently, a biocompatible and
biodegradable blend of
2 wt% hyaluronan and 7 wt% methylcellulose (2:7 HAMC) has been developed for
this application
[Gupta, D., C.H. Tator, and M.S. Shoichet, Fast-gelling injectable blend of
hyaluronan and
met hylcellulose for intrathecal, localized delivery to the injured spinal
cord. Biomaterials, 2006.
27(11): p. 2370-9]. The role of MC is to form a physical hydrogel through
hydrophobic junctions
[Schupper, N., Y. Rabin, and M. Rosenbluh, Multiple stages in the aging of a
physical polymer gel.
Macromolecules, 2008. 41(11): p. 3983-3994] and HA to increase solution
viscosity and to enhance
MC gel strength at lower temperatures through the salting out effect. 2:7 HAMC
was found to
degrade within 4-7 days in vivo, making it well suited for neuroprotective
delivery strategies but
unsuitable for drug delivery over the 2-4 weeks necessary for regenerative
strategies [Kang, C.E., et
al., A New Paradigm for Local and Sustained Release of Therapeutic Molecules
to the Injured
Spinal Cord for Neuroprotection and Tissue Repair. Tissue Eng Part A, 2008].
Accordingly, these
injectable hydrogels were used to deliver erythropoietin [Kang, C.E., et al.,
A New Paradigm for
Local and Sustained Release of Therapeutic Molecules to the Injured Spinal
Cord for
Neuroprotection and Tissue Repair. Tissue Eng Part A, 2008], as well as EGF
and FGF-2 via simple
diffusion [Jimenez Hamann, M.C., C.H. Tator, and M.S. Shoichet, Injectable
intrathecal delivery
system for localized administration of EGF and FGF-2 to the injured rat spinal
cord. Exp Neurol,
2005. 194(1): p. 106-19]. For soluble molecules, the release profile is
determined principally by
diffusivity and occurs within 24 hours due to the short diffusive path length
in vivo [Jimenez
Hamann, M.C., C.H. Tator, and M.S. Shoichet, Injectable intrathecal delivery
system for localized
administration of EGF and FGF-2 to the injured rat spinal cord. Exp Neurol,
2005. 194(1): p. 106-
19; Kong, C.E.; Tator, C.H.; Shoichet, M.S. 2010. Poly (ethylene glycol)
Modification Enhances
Penetration of Fibroblast Growth Factor 2 to Spinal Cord Tissue from an
Intrathecal Delivery
System J.Control Release; doi: 10.1016/j.jconre1.2010.01.029].
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As mentioned above, US parent Patent Application No.11/410831 describes a
polymer matrix
comprising an inverse thermal gelling polymer and an anionic polymer, for
example HAMC that
exists as a solid gel. This polymer matrix has a faster gelling rate than the
inverse gelling polymer,
and may be used alone or as a drug delivery vehicle for many applications. In
particular, the
polymer matrix can be used for localized, targeted delivery of pharmaceutical
agents upon injection
providing sustained release. A particular use of this invention is in delivery
of a therapeutic agent to
a fluid-filled space, such as the intrathecal space, in a highly localized,
targeted manner, wherein the
polymer matrix-contained therapeutic agent is able to circumvent the blood-
spinal cord barrier or
blood-brain barrier and enter the target tissue directly.
U.S. Patent No. 6,335,035 ('035) to Drizen, etal. is a divisional of U.S.
Patent No. 6,063,405 to
Drizen et al. which teaches sustained release compositions comprising a drug
dispersed within a
polymer matrix, methods of producing the same and treatments with the complex.
The '035 patent
discloses a sustained drug delivery system, which comprises a drug dispersed
within a polymer
matrix solubilized or suspended in a polymer matrix. The polymer matrix is
composed of a highly
negatively charged polymer material selected from the group consisting of
polysulfated
glucosoglycans, glycoaminoglycans, mucopolysaccharides and mixtures thereof,
and a nonionic
polymer selected from the group consisting of carboxymethylcellulose sodium,
hydroxypropylcellulose and mixtures thereof. Nonionic polymers are generally
used in amounts of
0.1% to 1.0% and preferably from 0.5% to 1.0%. Nonionic polymers in amounts
above 1.0% are
not used as they result in the formation of a solid gel product when employed
in combination with
an anionic polymer.
U.S. Patent No. 6,692,766 to Rubinstein et al. concerns a controlled release
drug delivery system
comprising a drug which is susceptible to enzymatic degradation by enzymes
present in the
intestinal tract; and a polymeric matrix which undergoes erosion in the
gastrointestinal tract
comprising a hydrogel-forming polymer selected from the group consisting of
(a) polymers which
are themselves capable of enhancing absorption of said drug across the
intestinal mucosal tissues
and of inhibiting degradation of said drug by intestinal enzymes; and (b)
polymers which are not
themselves capable of enhancing absorption of said drug across the intestinal
mucosal tissues and of
inhibiting degradation of said drug by intestinal enzymes.
U.S. Patent No. 6,716,251 to Asius et al. discloses an injection implant for
filling up wrinkles, thin
lines, skin cracks and scars for reparative or plastic surgery, aesthetic
dermatology and for filling up
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=
gums in dental treatment. The invention concerns the use of biologically
absorbable polymer
microspheres or micro particles suspended in a gel.
U.S. Patent No. 6,586,493 to Massia et al. discloses hyaluronate-containing
hydrogels having
angiogenic and vascularizing activity and pre-gel blends for preparing the
hydrogels. The hydrogels
contain a cross-linked matrix of a non-angiogenic hyaluronate and a
derivatized polysaccharide
material, in which cross-linking is effected by free-radical polymerization.
JP2003-342197 discloses a heat gelling pharmaceutical preparation containing
methylcellulose and
hyaluronic acid that is liquid at room temperature and gels upon
administration to the eye.
The literature also teaches the properties of polymer matrices and their use
as drug delivery vehicles
(Xu et al. Langmuir, (2004) 20(3): 646-652, Liang et al. Biomacromolecules,
2004. 5(5):1917-25,
Ohya et al. Biomacromolecules (2001) 2:856-863, Cho et al. International
Journal of Pharmaceutics
(2003) 260:83-91, Kim etal. Journal of Controlled Release (2002) 80:69-77,
Tate etal.
Biomaterials (2001) 22:1113-1123, and Silver et al., Journal of Applied
Biomaterials (1994) 5:89-
98).
SUMMARY
The present disclosure provides a composite hydrogel comprising a blend of an
aqueous solution of
an anionic polysaccharide or a derivative thereof, in particular hyaluronan or
a derivative thereof
and methylcellulose or other water soluble cellulose derivative which is
inverse thermal gelling and
that gels through hydrophobic interactions, together with dispersed
hydrophobic polymeric particles
selected from microparticles and nanoparticles.
In the hydrogel composite, some or all of the dispersed hydrophobic polymer
particles may be
encapsulated microparticles or nanoparticles that comprise at least one
therapeutic agent and each of
the at least one therapeutic agents has its own linear sustained release
profile. Each therapeutic
agent may be released independently, and when the one or more therapeutic
agents are encapsulated,
a tunable release rate is provided that is a sustained linear release profile
from a hydrogel composite
with enhanced stability.
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In another aspect, the disclosure provides a method for manufacturing a
hydrogel composite which
comprises the steps of I) providing an aqueous solution of methylcellulose or
other cellulose
derivative; 2) mixing an anionic polysaccharide or a derivatives thereof,
which may be hyaluronan
or a derivative thereof into the aqueous solution; 3) dispersing hydrophobic
polymeric particles
selected from micro particles and nanoparticles into the aqueous solution to
form a stable hydrogel
composite that has enhanced stability relative to a hydrogel without the
dispersed particles. In
another form of this method, the dispersed hydrophobic polymer particles may
be encapsulated
micro particles or nanoparticles that comprise at least one therapeutic agent
and each of the at least
one therapeutic agents has its own linear sustained release profile.
In another aspect of the present disclosure, there is provided a method for
the treatment of spinal
cord injury comprising delivering into the intrathecal space, a composite
hydrogel an anionic
polysaccharides or a derivatives thereof, such as hyaluronan or a derivative
thereof and
methylcellulose or other water soluble cellulose derivative, and at least one
therapeutic agent
selected from FGF2, FGF1, vascular endothelial growth factor (VEGF) and
platelet derived growth
factor (PDGF) dispersed in the composite hydrogel as microparticles and
nanoparticles and/or as
encapsulated microparticles and nanoparticles to provide a linear sustained
release profile of the at
least one therapeutic agent to the injured spinal cord to promote endothelial
cell proliferation and
blood vessel formation.
In another aspect, there is provided a hydrogel composite comprising: a
hydrogel comprising a
blend of a solution containing dissolved methylcellulose and a hyaluronan
wherein the
hyaluronan and the methylcellulose are present in a weight ratio of hyaluronan
to
methylcellulose in an amount of about 2:3; and dispersed hydrophobic polymeric
particles
selected from microparticles, being a particle size of 1 micron to 30 microns
and nanoparticles
being a particle size of from 10 nm to 1000 nm, wherein some or all of which
hydrophobic
polymeric microparticles and nanoparticles encapsulate at least one
therapeutic agent and
wherein the dispersed hydrophobic polymeric particles interact with the
hydrogel through
hydrophobic interactions between the hydrogel and the dispersed hydrophobic
polymeric
particles to alter the release of the at least one therapeutic agent, and
wherein the stability of the
hydrogel composite with the dispersed hydrophobic polymeric particles is
enhanced relative to
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CA 02703807 2016-10-17
the stability of the hydrogel alone, and each of the at least one therapeutic
agent has a sustained
release profile that extends for about 28 days or more.
In another aspect, there is provided a method of manufacturing the hydrogel
composite according
to claim 1, which comprises the steps of 1) providing an aqueous solution of
dissolved
methylcellulose; 2) mixing hyaluronan into the aqueous solution, wherein the
hyaluronan and the
methylcellulose are present in a weight ratio of hyaluronan to methylcellulose
in an amount of
about 2:3; and 3) dispersing hydrophobic polymeric particles being a particle
size of 1 micron to
30 microns and nanoparticles being in a particle size of from 10 nm to 1000 nm
into the aqueous
solution to form a stable hydrogel composite that has enhanced stability
relative to a hydrogel
without the dispersed hydrophobic polymer particles.
In another aspect, there is provided use of the hydrogel composite described
above for the
promotion of angiogenesis in a subject with spinal cord injury by intrathecal
or sub-dural
administration, to promote endothelial cell proliferation and blood vessel
formation; wherein the
at least one therapeutic agent promotes angiogenesis.
Surprisingly, the stability of the hydrogel with the polymeric particles
dispersed therein is enhanced
relative to the stability of the hydrogel alone. In this form, the hydrogel
composite may be used as a
bulking agent for reconstructive or cosmetic surgery or as a lubricating
agent, or matrix for in situ
tissue growth. Because methylcellulose is currently used in food, the hydrogel
composite could be
used, for example in molecular gastronomy.
The stability of the hydrogel composite is highly advantageous and unexpected
as it offers a stable
composite without a therapeutic agent or a stable composite for simultaneous
delivery of therapeutic
agents that has been shown to be stable for periods of up to 50 days.
In the aforementioned pending, parent US Patent Application No. 11/410831, it
was thought release
would be tuned by altering the properties of the encapsulating polymer.
Examples of these
properties are: polymer degradation rate, molecular weight; particle porosity,
size; and drug load.
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Drug release from the resulting composite was not expected to differ
substantially from drug release
from the particles alone. This is because HAMC (the gel) presents a minimal
diffusive barrier to
drug release. This is also the case with HMW(high molecular weight) HAMC in
the present
disclosure when a drug is dissolved in the hydrogel: as release is fast and
occurs by Fickian
diffusion. Unexpectedly, when hydrophobic PLGA particles containing a
therapeutic agent were
dispersed in HMW HAMC to form composite HMW HAMC it was found that the
particles and gel
interacted synergistically to alter drug release. As seen in the accompanying
drawings, release of
particle encapsulated drugs from composite HMW HAMC is significantly slower
and more linear
than release from the same particles in aqueous suspension. This previously
unidentified interaction
provides an additional mechanism to alter the release of particle encapsulated
drugs that is not
dependent on particle or gel formulation. This is advantageous because
existing methods in the field
to sustain and linearize drug release negatively impact the utility of the
drug delivery platform. For
example, if drug release is sustained through alteration to the encapsulating
polymer, the deliverable
drug load is reduced or particle size is increased. This reduces therapeutic
efficacy or injectability,
respectively. Alternatively, if drug release is slowed by reducing diffusivity
through hydrogel, for
example by using high concentrations of gel forming polymer, injectability is
reduced and gel
stiffness may be increased. Increased gel stiffness may be detrimental when
the device is brought in
contact with soft tissue. Necrosis and scarring of biomaterial contacting
tissue has been recorded
when the biomaterial is stiffer than the adjacent tissue. Thus, the
interaction of HAMC with drug
loaded hydrophobic polymer particles maximizes the deliverable drug load and
injectability without
undue increase to gel stiffness.
DETAILED DESCRIPTION
The other water soluble cellulose derivatives may be selected from the group
comprising
hydroxypropyl methylcellulose, ethylcellulose, 3-0-ethylcellulose,
hydroxypropyl methylcellulose
phthalate, hydrophobically modified hydroxyethyl cellulose selected from
ethyl(hydroxyethyl)cellulose, 6-0-alkylated cellulose, cellulose octanoate
sulfate, cellulose lauroate
sulfate, cellulose stearoate sulfate, and cationic derivatives thereof, 6-0-
benzylcellulose, 2,3-di-O-
methy1-6-0-benzylcellulose, 2,3-di-0- benzylcellulose, 2,3-di-O-benzy1-6-0-
methylcellulose, 2,3,6-
tri-O-benzylcellulose, hydroxypropyl methylcellulose acetate succinate, 0-
24242-
methoxyethoxy)ethoxy]acetyl cellulose, when formulated such that the
cellulosic solution is inverse
thermal gelling.
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The anionic polysaccharide or derivative thereof functions to salt out the
water soluble cellulose
derivative, lowering the temperature of gelation and to increase the viscosity
of the composite
immediately after injection during the period of gelation. The anionic
polysaccharide or derivative
thereof may therefore comprise polycarboxylates, a specific example of which
is hyaluronan and its
derivatives; xanthan gum, dextran sulphate, and other glycosaminoglycans.
Derivatives of hylaruonan include esters of HA, typically formed by treating
quaternary ammonium
salt of HA with an esterifying agent. Esterification may be carried out using
a number of different
classes of alcohols such as aliphatic, cycloaliphatic and heterocyclic. Thus,
a number of different
derivatives can be synthesized and these derivatives have a wide range of
physicochemical
properties. Examples of HA esters are described by Campoccia D et al. Fidia
Advanced
Biopolymers, Albano Terme (PD), Italy. Biomaterials. 1998 Dec;19(23):2101-27.
Glycosaminoglycans are well known, naturally occurring, polysaccharides
containing disaccharide
repeating units of hexosamine and hexose or hexuronic acid, and may contain
sulfate groups.
Representative glycosaminoglycans are: heparin; heparan; chondroitin; keratan;
dermatan; and
sulfates of such materials. Glycosaminoglycans are rendered anionic when the
amine group is other
than a quaternary form (e.g. all other than R3NH+) or when the number of
deprotonanted sulphate,
carbxoylate, or other anionic moieties are greater than the number of
protonated amines.
The anionic polysaccharide or derivative thereof and/or methylcellulose or
other water soluble
hydrophobic cellulose derivatives may additionally be chemically modified
using known methods to
bear contain increased functionality. Non-limiting examples of functionality
include; carboxylic
acid, primary amine, aldehyde, hydrazide, maleimide, thiol, furan, alkyne,
azide, alkene, urethane,
or primary alcohol. These chemical modifications permit subsequent biological
utility, for example
through covalent linkage of the anionic polysaccharide or derivative thereof
and/or methylcellulose
or other water soluble hydrophobic cellulose derivatives to a drug molecule or
ligand for cellular
interaction. For example, the tethering of opiates to HA [Gianolio, DA., et al
"Hyaluronan tethered
opioid depots: synthetic strategies and release kinetics in vitro and in vivo.
Bioconjugate Chem.
(2008). Sep;19(9):1767-74.] and the tethering of cellular adhesion peptides to
HA [Cui, F.Z., et al.
"Hyaluronic acid hydrogel immobilized with RGD peptides for brain tissue
engineering. J. Mater.
Sci. Mater.Med. (2006). Dec;17(12):1393-401.1.
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The hydrogel composite when formulated for delivery to a body cavity must meet
the following
criteria: injectable through a fine needle (via injection delivery) which
allows for minimally invasive
surgery; fast gelling to ensure localized drug delivery at a site of injury;
in situ degradable to avoid
additional surgeries for device-removal; and scar formation biocompatible to
limit foreign-body
reaction.
The gelation temperature of thermal gelling materials as well as the kinetics
of gelation is
concentration dependent. Immediately upon injection of a temperature sensitive
polymer into a
fluid filled cavity, the polymer disperses prior to gelling. The dispersion
causes the gelation rate to
decrease. This phenomenon also occurs with chemically cross-linked gels where
the kinetics are
concentration dependent. To overcome this obstacle, it is necessary to have a
highly viscous
material so that once injected, will not disperse and thereby suffer from a
decreased gelation rate.
At the same time, however, the viscous material must still be injectable and
this can be achieved
with the use of a shear-thinning material. Blending a highly negatively
charged anionic
polysaccharide or derivative thereof with an inverse thermal gelling polymer
like methylcellulose or
other cellulose water soluble derivatives at certain molar ratios can achieve
this effect.
The anionic polysaccharides or derivatives introduced above may have a
molecular weight range of
between about 100,000 and about 7,000,000 kg/mol. Hyaluronan and derivatives
of hyaluronan
may be employed. Exemplified herein is hyaluronic acid (HA). HA is a linear
polysaccharide
composed of repeating disaccharide units of N-acetyl-glucosamine and D-
glucuronic acid. HA
forms highly viscoelastic and shear-thinning solutions and has been used for
drug delivery, tissue
engineering applications as well as for soft tissue augmentation. HA is known
to have wound-
healing effects such as anti-inflammation, as well as to minimize tissue
adhesion and scar formation.
It is degraded enzymatically by hyaluronidase, which can be produced by all
cells. Its polymeric
chains, of lengths 10-15 thousand disaccharides, form random coils with large
spheroidal hydrated
volumes of up to 400-500 nm in diameter. Because of the high solubility of HA
in water, it must be
chemically modified or blended with a gelling polymer to form a gel. Chemical
modification can
occur at the carboxyl group or the hydroxyl group of HA and also at the amino
group when the N-
acetyl group is removed. Blends of unmodified HA with a gelling polymer are
injectable upon an
application of force to a syringe because the shear-thinning properties of HA
cause the polymer
chains to straighten and align themselves permitting flow through the needle.
HA then returns to its
DOCSTOR: 1938001\1
CA 02703807 2010-05-12
high viscosity, zero shear structure upon exiting the needle as the polymeric
chains once again
become entangled amongst themselves.
The methylcellulose or other water soluble cellulose are inverse thermal
gelling polymers capable of
gelling upon an increase in temperature. The methyl cellulose or other water
soluble cellulose may
have a molecular weight in the range of between about 2000 and about 1,000,000
g/mol. Exemplary
of other suitable polymers include a chitosan and quadrature-glycerophosphate
solution, collagen, a
tri-block copolymer of poly(ethylene glycol)-poly(lactic-co-glycolic acid)-
poly(ethylene glycol), a
tri-block copolymer of poly(propylene glycol)-poly(ethylene glycol)-poly
(propylene glycol),
poly(N-isopropyl acrylamide), copolymers of poly-N-isopropylacrylamide,
polysaccharides and
mixtures thereof.
Exemplified herein is methylcellulose (MC), a carbohydrate and derivative of
cellulose. MC is an
example of a temperature sensitive gel, or a thermally reversible gel, that
gels upon increase in
temperature. When the degree of substitution of hydroxyl groups with methyl
groups is between
1.4-1.9, methylcellulose has inverse thermal gelling properties whereby it
gels upon an increase of
temperature. As the temperature increases, hydrogen bonds with the surrounding
solvent break and
hydrophobic junctions form to produce a gel. Methylcellulose generally forms
weak gels at 37 C
when in water, but the gelation temperature can be decreased by an increase in
salt concentration.
This occurs because the water molecules surround the salts, effectively
reducing the number of
polymer-solvent interactions. Methylcellulose has previously been considered
as a scaffold for
experimental traumatic brain injury where in vivo tests in rats indicated
biocompatibility over a span
of two weeks. MC has also been used as a scaffold in the peripheral nervous
system for nerve
regeneration with promising results, without any adverse pathological
reactions over 8 weeks.
Although it is not found to degrade enzymatically, the weak gel structure does
dissolve at 37 C and
swells minimally.
Through the manipulation of polymer structure, concentration, and molecular
weight, the hydrogel
composites may not be in gel form at the time of administration or formation,
however, they do form
gels with an increase in temperature such as to body temperatures.
To take advantage of the thermal gelling properties of MC and the shear-
thinning properties of HA,
MC and HA are blended. The combination of an aqueous solution of MC and
lyophilized HA
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results in dispersal of HA within the solution. The resulting hydrogel
composite is comprised of
dissolved MC and dissolved HA. It is a fast-gelling polymer and is referred to
as HAMC. Methods
of blending polymer matrices for drug delivery are well known. In general,
methods to prepare
HAMC involve preparation of a sterile solution of MC in a buffered salt
solution, which is cooled to
4 C prior to the addition of sterile, lyophilized HA which dissolves over
time. Because of the high
viscosity of this material prior to gelation, HAMC does not flow significantly
at room temperature.
This allows the polymer blend to maintain some structure as it gels. It is
expected that since HA
strongly interacts with the solvent, the presence of HA in a MC solution
likely dehydrates the MC,
similar to the effect of salt on MC gelation, effectively decreasing the
gelation temperature. Hence,
HA also functions to lower the gelation temperature of MC.
HAMC is unique amongst reversible physical hydrogels in its ability to return
to the gel state more
rapidly after injection. Typically, physical gelling polymers undergo a phase
transition from a
solution to a gel after injection whereas HAMC can be formulated such that it
is a gel both prior to
and following injection. The shear thinning properties of HA enable the HAMC
gel to be injectable
while the thermal gelling properties of MC return HAMC to a gel following
injection. The
properties of the gel are highly sensitive to the amount of HA, and altering
the concentration of HA
would be expected to affect the injectability of the hydrogel composite and
the gelation rate. For
example, higher molecular weights of HA are likely to have enhanced shear
thinning properties.
Varying the concentrations of the individual polymers as well as the use of
polymers of different
molecular weights enhances the properties of the hydrogel composite for
injectable delivery.
The hydrogel composite of this invention can be used to target delivery of a
pharmaceutical agent,
particularly by means of injection. It is well known in the art that
pharmaceutical agents can be
loaded into polymer matrices with high loading efficiency while retaining the
agent's bioactivity.
Common methods include imbibing the pharmaceutical agents into pre-formed
matrices or
incorporating the pharmaceutical agent in the preparation of the polymer
matrix itself [Liang et al.
Biomacromolecules 5:1917-1925 (2004), Cho et al. Int. J. Pharmaceutics 260:83-
91 (2003), Kim et
al. J. Controlled Release 80:69-77 (2002)]. For HAMC, both methods will work.
Preferably, the
therapeutic agent(s), protein(s) or peptide(s) will have some solubility in
the MC solution prior to
the addition of HA. The solution is maintained overnight to allow the HA to
completely dissolve in
the solution. The injectable hydrogel composite of this invention provides the
following
advantages: localized drug release, improved drug distribution, and controlled
release rates.
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Localised drug release at the site of injury enhances therapeutic efficacy,
thereby minimizing the
risks of systemic toxicity and side effects. Since less drug is lost
systemically, localized release also
allows for lower doses of drug to be released for therapeutic efficacy. Drug
distribution is improved
through the localized delivery and by sustained release rates. The advantage
offered by the current
hydrogel composite is enhanced stability and a release profile that is linear
and can be determined
based on the particles incorporated therein.
In a specific form of the hydrogel composite, the hyaluronan or a derivative
thereof may comprise
from about 100 to about 7,000 kg/mol and the methylcellulose or a derivative
thereof may comprise
from about 1,500 to about 3,000 kg/mol. More particularly the hyaluronan or a
derivative thereof
may comprise from about 1500 to about 3000 kg/mol and the methylcellulose or a
derivative thereof
may comprise from about 10 to about 400 kg/mol. When other combinations are
used to form the
hydrogel composite these amounts can be readily adjusted. This applies to all
the ratios, quantities
provided hereafter.
In this specific hydrogel composite, the ratio of hyaluronan or a derivative
thereof to the
methylcellulose or a derivative thereof may comprise from about 1:20 to about
1:1 w/w, more
particularly the ratio of hyaluronan or a derivative thereof to the
methylcellulose or a derivative
thereof may comprise from about 1:5 to about 2:3 w/w.
The ratio of hyaluronan or a derivative thereof in this hydrogel composite may
comprise from about
0.5% to about 5.0% by weight and the methylcellulose or a derivative thereof
may comprise from
about 1.0% to about 10% by weight of the composite. More particularly, the
amount of hyaluronan
or a derivative thereof may comprise from about 1.0% to about 2.0% by weight
and the
methylcellulose or a derivative thereof may comprise from about 3.0% to about
7.0% by weight,
based on the composite.
The polymeric particles may be hydrophobic polymer particles selected from
degradable polymers
selected from the group consisting of aliphatic polyesters, aliphatic-aromatic
polyesters, aliphatic
polyamides, amide ester copolymers, urethane ester copolymers, urethane amide
copolymers and
urea ester copolymers; and from non-degradable polymers selected from the
group consisting of
cellulose, starch, polystyrene, polyethylene, polypropylene, and alkylated
poly(acrylates). The
encapsulated therapeutic agents may be prepared in a known manner in the art.
Exemplified here
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are therapeutics that have been encapsulated in poly(lactic-co-glycolic acid)
(PLGA) microparticles
and nanoparticles. Hydrogels containing polymer particles are denoted
composite hydrogels.
As used herein, "microparticles" refers to particles having a diameter of less
than 1.0 mm, and more
specifically between 1.0 and 100.0 microns. Microparticles include micro
spheres, which are
typically solid spherical microparticles. Microparticles also include
microcapsules, which are
spherical microparticles typically having a core of a different polymer, drug,
or composition.
As used herein, "nanoparticles" refers to particles or structures in the
nanometer range, typically
from about 1 nm to about 1000 nm in diameter, which are encapsulated within
the polymer.
Lipid/polymer liposomes and polymeric microspheres are known in the art. A
method of producing
such lipid/polymer liposomes is described, for example, in U.S. Patent No.
6,277,413.
Suitable biodegradable polymers for producing the microparticles are
polyesters such as polylactide,
polyglycolide, copolymers of lactide and glycolide, polyhydroxybutyrate,
polycaprolactone,
copolymers of lactic acid and lactone, copolymers of lactic acid and PEG,
copolymers of a-hydroxy
acids and a -amino acids (polydepsipeptides), polyanhydrides, polyorthoesters,
polyphosphazenes,
copolymers of hydroxybutyrate and hydroxyvalerate, poly ethylene carbonate),
copoly(ethylene
carbonate), polyethyleneterephthalate or mixtures of these polymers. Examples
of
resorbable/biodegradable polymers are lactide homopolymers poly(L-lactide),
poly(D,L-lactide),
and copolymers of lactide and glycolide such as 50:50 poly(DL lactide co-
glycolide)(PLG). While
polyethylene glycol (PEG) is the preferred water soluble polymer for mixing
with the biodegradable
polymer, other suitable water soluble polymers include poly(oxyethylene
oxide)(PEO),
poly(oxyethylene)-poly(oxypropylene) [PEO-PPO] block copolymers such as tri-
block PEO-PPO-
PEO copolymers (Poloxamers, Pluronics) and tetra-functional block copolymers
derived from the
sequential addition of propylene oxide and ethylene oxide to ethylene diamine
(Poloxamines,
Tetronics), copolymers of PEG with poly(lactic acid), oligomers of poly(lactic
acid), lactides,
copolymers of PEG and amino acids, conjugates of PEG with polysaccharides for
example a
conjugate produced from 40000 MW dextran and polyoxyethylene-glycol monomethyl
ether and
others as described by Duval et al. in Carbohydrate Polymers, 15, (1991), 233-
242, conjugates of
PEG with proteins such as those described by Nucci et al., in Advances in Drug
Delivery Review, 6,
(1981),113-151, or with collagen as described by Rhee et al in Po[y(ethylene
glycol) chemistry.
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Biotechnical and Biomedical Applications. Ed. J. Milton Harris, Plenum Press
(1992), or conjugates
of PEG with colony Stimulating Factor (CSF-1) as described by Katre N.V. in
The conjugation of
proteins with polyethylene glycol and other polymers. Adv. Drug Delivery
Reviews, 10, 91-114
(1993).
Suitable biocompatible, non-biodegradable polymers include, but are not
limited to, polyacrylates;
ethylene-vinyl acetates; acyl substituted cellulose acetates; non-degradable
polyurethanes;
polystyrenes; polyvinyl chlorides; polyvinyl fluorides; poly(vinyl
imidazoles); chlorosulphonate
polyolefins; polyethylene oxides; or blends or copolymers thereof.
To form microspheres, in particular, a variety of techniques known in the art
can be used. Methods
of producing microspheres are described, for example, in U.S. Pat. Nos.
5,552,133; 5,310,540;
4,718,433; and 4,572,203; European Patent Publication No. EP 458,745; and PCT
Publication No.
WO 92/05806. Methods include, for example, single or double emulsion steps
followed by solvent
removal. Solvent removal may be accomplished by extraction, evaporation or
spray drying among
other methods.
In the solvent extraction method, the polymer is dissolved in an organic
solvent that is at least
partially soluble in the extraction solvent such as water. The bioactive
molecule, either in soluble
form or dispersed as fine particles, is then added to the polymer solution,
and the mixture is
dispersed into an aqueous phase that contains a surface-active agent such as
poly (vinyl alcohol).
The resulting emulsion is added to a larger volume of water where the organic
solvent is removed
from the polymer/bioactive agent to form hardened microparticles.
In the solvent evaporation method, the polymer is dissolved in a volatile
organic solvent. The
bioactive molecule, either in soluble form or dispersed as fine particles, is
then added to the polymer
solution, and the mixture is suspended in an aqueous phase that contains a
surface-active agent such
as poly (vinyl alcohol). The resulting emulsion is stirred until most of the
organic solvent
evaporates, leaving solid microspheres.
In the spray drying method, the polymer is dissolved in a suitable solvent,
such as methylene
chloride (e. g., 0.04 g/m1). A known amount of bioactive molecule (drug) is
then suspended (if
insoluble) or co-dissolved (if soluble) in the polymer solution. The solution
or the dispersion is then
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spray-dried. Microspheres ranging in diameter between one and ten microns can
be obtained with a
morphology, which depends on the selection of polymer.
Other known methods, such as phase separation and coacervation, and variations
of the above, are
known in the art and also may be employed in the present invention.
Formulation of nanoparticles can be achieved essentially as described above
for microparticles
except that high speed mixing or homogenization is used to reduce the size of
the polymer/bioactive
agent emulsions to below about 2.0 microns, preferably below about 1.0 micron.
For example,
suitable techniques for making nanoparticles are described in PCT Publication.
No. WO 97/04747.
In the examples presented herein, in the case of the dispersed hydrophobic
polymeric particles, the
particle load may comprise from about 1 to about 20 wt%, more particularly
from about 2.5 to about
10 wt%, based on the composite. In the case where some or all of the dispersed
hydrophobic
polymer particles are encapsulated micro particles or nanoparticles that
comprise at least one
therapeutic agent, the dispersed particle load may comprise from about 1 to
about 20 wt%, more
particularly from about 10 to about 20 wt%, based on the composite.
In the examples presented herein the hydrophobic polymeric particles may have
particle sizes
selected from particle sizes of about 150 nm to about 40 pm, and more
particularly, from about 220
nm to about 830 nm. When the particles include the at least one therapeutic
agent, the particle sizes
are selected to provide the desired release profile. A suitable sustained
release profile has been
found to be provided by dispersing polymeric particles selected from particle
sizes of from about
220 nm to about 37 pm.
The delivered therapeutic agent load from the encapsulated particles is
generally in the range of
about 0.1 to about 30 wt% (drug mass as a percentage of the drug loaded
particle mass), and more
particularly in the range of from about 1.0 to about 10 wt% (drug mass as a
percentage of the drug
loaded particle mass).
The aqueous solution of hyaluronan or a derivative thereof and methylcellulose
or other cellulose
derivative may be selected from the group comprising water, saline, artificial
cerebrospinal fluid,
and buffered solutions.
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The drug delivery hydrogel composite of this invention has multiple
applications and may be
delivered via injection, transdermal, oral, sub-cutaneous, intranasal,
vaginal, buccal, intrathecal,
subdural, epidural, ocular space, dental, intratumoral, intramuscular,
intraarticular, and
intraveneously. The drug delivery synergistic hydrogel composite is designed
for delivery into a
fluid-filled (or partially-filled) cavity. These include all cavities
throughout the body, including but
not limited to the intrathecal space, the intra-articular cavity, among
others.
The hydrogel composite components can be modified to alter the degradation
rate of the hydrogel
composite and, hence, affect the rate of release of the pharmaceutical agent
from the hydrogel
composite. One such modification involves addition of salts to alter the pH.
Since the charge of the
anionic gelling polymer causes its viscosity to be pH sensitive, it is
possible that the hydrogel
composite blend is also pH sensitive. The pH can be varied to control
properties such as
formulation for delivery or processing. A pH-sensitive hydrogel composed of
methylcellulose and
alginate was previously demonstrated by Liang et al. (Biomacromolecules 5:1917-
1925, 2004) to be
capable of increased load release at a higher pH (pH 7.4) compared to a lower
pH (pH 1.2).
Blending of the hydrogel composite with a salt could be performed to achieve a
pH-dependent
delivery vehicle.
Another alternative to creating a more stable gel for slower degradation is to
functionalize the
polymers with thiol groups and acrylate groups. The hydrogel composite is
injected and gels
quickly at the site of injection because, at physiological conditions, a
Michael-type addition reaction
occurs between the polymer end terminated with thiol and the polymer
terminated with acrylate
chains. This technique results in a product that is fast gelling with a high
degree of gel strength,
achieved as a result of linking multiple crosslinked polymers. For example,
using a methacrylated
polymer, such as methacrylated dextran, and a thiol conjugated polymer, such
as PEG-dithiol or a
peptide-dithiol, a crosslinked dextran gel can be achieved. Using a specific
amino acid sequence
that is enzymatically cleaved, a degradable, injectable crosslinked
polysaccharide gel can be
synthesized.
Another method of controlling degradation rates is to increase the
hydrophobicity of HA, which
helps to maintain the integrity of gel through the formation of more
hydrophobic junctions resulting
in less water penetration. To render HA more hydrophobic, the reactive
functional groups, hydroxyl
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or carboxyl, can be modified with hydrophobic molecules. For example, it is
possible to modify the
carboxyl group of HA with acetic hydrazide using standard coupling agents,
such as carbodiimides
like EDC. It should be noted that the carboxyl group is important for the
highly viscous nature of
the hydrogel composite.
Another means to enhance sustained release of the pharmaceutical agent is to
take advantage of
ionic interactions between the agent and the polymer. The highly negatively
charged anionic gelling
polymer engages in ionic interactions with positively charged molecules. In
cases where there is no
significant drug-polymer interaction or the charges are the same such that
there are no attractive
forces, the charge can be modified with the use of charged stabilizers.
Cationic particles or a
mixture of cationic and anionic particles are used within the anionic gelling
polymer to prevent the
particles from dispersing away from the gel, as well as to promote increased
gel strength through
ionic crosslinks. Methods for incorporating cationic or cationic/anionic
charge stabilizers into
pharmaceutical compositions may be employed and are known to those of skill in
the art. Examples
are known in the art [Schexnailder, P. and G. Schmidt, Nanocomposite polymer
hydrogels. Colloid
And Polymer Science, 2009. 287(1): p. 1-11. Hooper, J.B. and K.S. Schweizer,
Theory of phase
separation in polymer nanocomposites. Macromolecules, 2006. 39(15): p. 5133-
5142. Hooper, J.B.
and K.S. Schweizer, Contact Aggregation, Bridging, and Steric Stabilization in
Dense Polymer-
Particle Mixtures. Macromolecules, 2005. 38: p. 8858-8869.]
Another alternative to further controlling drug release is by tethering or
covalently bonding the
pharmaceutical agent to the polymer. The agent releases from the hydrogel
composite upon
breakage of the covalent bond or upon dissolution of the chain from the
hydrogel composite
network. Methods of covalently bonding pharmaceutical agents to polymers may
be employed and
are known to those of skill in the art. Examples are described in Hoffman et
al. (Clinical Chemistry
46(9):1478-1486).
Chitosan, an amino-polysaccharide, is another example of an inverse thermal
gelling polymer that
can be used in the hydrogel composite. It is obtained by the alkaline
deacetylation of chitin.
Chitosan is both biocompatible and biodegradable and has inherent wound
healing properties, in
addition to a wide range of applications in drug delivery and tissue
engineering. Chitin and chitosan
are generally found as copolymers, and it is the chitin segments that are
enzymatically degradable by
lysozyme. It is a cationic polymer which is soluble in acidic conditions.
Recently, Chenite et al.
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(Biomaterials 21:2155-2161, 2000) developed a thermogelling polymer by mixing
beta-
glycerophosphate (quadrature-GP) into a chitosan solution. Chitosan/beta-GP
gels upon an increase
in temperature where the gelation temperature is affected by both pH and beta-
GP concentration.
The negatively charged beta-GP molecules are attracted to the positively
charged amine groups of
chitosan, preventing chitosan from aggregating and precipitating at
physiological pH. Upon an
increase in temperature, a gel is formed because of the formation of physical
junction zones which
occur when hydrophobic and hydrogen bonding forces outweigh the interchain
electrostatic
repulsion forces.
It is possible to alter the rate of degradation of the composite by increasing
the hydrophobicity of
hyaluronan or the derivative thereof. More specifically, an altered rate of
degradation rate may be
provided by the addition of at least one functional group to the hyaluronan or
the derivative thereof
or the methylcellulose or other cellulose derivative selected from the group
consisting of carboxylic
acid, primary amine, aldehyde, hydrazide, maleimide, thiol, fiiran, alkyne,
azide, alkene, urethane,
and primary alcohol.
The therapeutic agent may be encapsulated in a microsphere, nanoparticle or
liposome.
A charge stabilizer may be added to promote an interaction between the blend
and the therapeutic
agent. The therapeutic agent may be covalently bonded to the hyaluronan or the
derivative thereof.
Non-limiting examples of the therapeutic agent include, but are not limited to
the group of
therapeutic agents comprising anaesthetics for use in caudal, epidural,
inhalation, injectable,
retrobulbar, and spinal applications; analgesics, selected from the group
comprising acetaminophen,
baclofen, ibuprofen, fluriprofen, ketoprofen, voltaren, phenacetin and
salicylamide; anti-
inflammatories selected from the group comprising naproxen and indomethacin;
antihistamines,
selected from the group comprising chlorpheniramine maleate, phenindamine
tartrate, pyrilamine
maleate, doxylamine succinate, henyltoloxamine citrate, diphenhydramine
hydrochloride,
promethazine, brompheniramine maleate, dexbrompheniramine maleate, clemastine
fumarate and
triprolidine; antitussives selected from the group comprising dextromethorphan
hydrobromide and
guaifenesin; expectorants; decongestants, selected from the group comprising
phenylephrine
hydrochloride, phenylpropanolamine hydrochloride, pseudoephedrine
hydrochloride, and ephedrine;
antibiotics selected from the group comprising amebicides, broad and medium
spectrum, fungal
medications, monobactams and viral agents; bronchodilators selected from the
group comprising
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theophylline, albuterol and terbutaline; cardiovascular preparations selected
from the group
comprising diltiazem, propranolol, nifedepine, clonidine, alpha adrenoceptor
agonists, alpha
receptor blocking agents, alpha and beta receptor blocking agents, antiotensin
converting enzyme
inhibitors, beta blocking agents, calcium channel blockers, and cardiac
glycosides; central nervous
system drugs selected from the group comprising thioridazine, diazepam,
meclizine, ergoloid
mesylates, chlorpromazine, carbidopa and levodopa; metal salts selected from
the group comprising
potassium chloride and lithium carbonate; minerals selected from the group
consisting of iron,
chromium, molybdenum and potassium; immunomodulators; immunosuppressives
selected from the
group comprising minocycline, cyclosporine A; thyroid preparations selected
from the group
comprising synthetic thyroid hormone, and thyroxine sodium; peptide and
glycoprotein hormones
and analogues selected from the group comprising human chorionic gonadotrophin
(HCG),
corticotrophin, human growth hormone (HGH--Somatotrophin) and erythropoietin
(EPO); steroids
and hormones selected from the group comprising ACTH, anabolics, androgen and
estrogen
combinations, androgens, corticoids and analgesics, estrogens, glucocorticoid,
gonadotropin,
gonadotropin releasing, hypocalcemic, menotropins, parathyroid, progesterone,
progestogen,
progestogen and estrogen combinations, somatostatin-like compounds,
urofollitropin, vasopressin,
methyl prednisolone, GM1 ganglioside, cAMP, and others; vitamins selected from
the group
comprising water-soluble vitamins and veterinary formulations; growth factors
selected from the
group comprising EGF, FGF2 and neurotrophin; peptides, peptide mimetics and
other protein
preparations; DNA; and, small interfering RNAs; with or without a
pharmaceutically acceptable
carrier or preservative.
DETAILED EMBODIMENTS
The present disclosure may be understood more readily by reference to the
following detailed
description taken in connection with the accompanying figures and examples,
which form a part of
this disclosure. It is to be understood that this disclosure is not limited to
the specific compositions,
methods, conditions or parameters described and/or shown herein, and that the
terminology used
herein is for the purpose of describing particular embodiments by way of
example only and is not
intended to be limiting of the claims.
Exemplified herein are biocompatible and biodegradable blends of 2 wt%
hyaluronan and 7 wt%
methylcellulose (2:7 HAMC) using specific molecular weights of HA and MC, and
biocompatible
and biodegradable blends of between 1-3 wt% HA with 3% MC using a second pair
of polymer
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molecular weights. The later pair of molecular weights are denoted "high
molecular weight
(HMW)" relative to the former. The role of MC is to form a physical hydrogel
through hydrophobic
junctions and HA to increase solution viscosity and to enhance MC gel strength
at lower
temperatures through the salting out effect. Additionally, based on the anti-
inflammatory action of
HA, this component is likely responsible for the beneficial, anti-inflammatory
effect of 2:7 HAMC
in a compression model of SCI. 2:7 HAMC was found to degrade within 4-7 days
in vivo, making it
well suited for neuroprotective delivery strategies but unsuitable for drug
delivery over longer time
periods (ie. 1-4 weeks) weeks necessary for regenerative strategies.
Accordingly, these injectable
hydrogels were used to deliver erythropoietin, as well as EGF and FGF-2 via
simple diffusion. For
soluble molecules, the release profile is determined principally by
diffusivity and occurs within 24
hours due to the short diffusive path length in vivo.
BRIEF DESCRIPTION OF THE DRAWINGS
Figure 1 illustrates that the rheological behaviour of certain HMW HAMC
formulations are similar
to that of 2:7 HAMC. Similar G' and G" values indicate comparable
injectability. The frequency
sweeps were conducted at 37 C. Storage (G', open symbols) and loss modulus
(G", filled symbols)
of injectable gels are represented. 2:3 HMW HAMC (0) and 2:7 HAMC (0) are gels
are shown in
the left graph. 1:3 HMW HAMC (o) and 3:3 HMW HAMC (A) are shown in the right
graph;
Figure 2 illustrates swelling of hydrogel composites over time. The traces
are; 0 mg/mL (s), 25
mg/mL (o), and 75 mg/mL (a) of 220 nm polystyrene particles dispersed in the
hydrogel.
Specifically, equilibrium swelling of 1:3 and 2:3 HMW HAMC and their
composites and composite
2:7 HAMC is similar to or less than 2:7 HAMC alone, suggesting that these
formulations can be
used in confined volumes such as the intrathecal space, in the same manner as
2:7 HAMC:
Figure 3 illustrates utilizes normalized swelling as a proxy for mass loss and
a measure of stability.
In other experiments (data not shown) mass loss from the composite does not
occur until normalized
swollen volume falls below 80%;
Figure 4 illustrates residual particle load over time for a number of hydrogel
composites. The traces
are: 1:3 HMW HAMC (a), 2:3 HMW HAMC (A), 3:3 HMW HAMC (*), each loaded with 25
mg/mL of 220 nm PS particles. Particles embedded in composite HAMC remain
within composite
HAMC until the composite begins to degrade: that is, they are not prematurely
lost from the
injectable composite until the normalized swollen volume falls below about
80%;
Figure 5 illustrates the cumulative therapeutic agent release from a number of
hydrogel composites.
The traces are; 1:3 HMW HAMC (0), 2:3 HMW HAMC (0), and 3:3 HMW HAMC (A). a-
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Chymotrypsin release is represented by filled symbols and Iga release by open
symbols. When a protein is
dispersed in soluble form within the hydrogel or composite it is released in a
primarily Fickian fashion, being
linear with ea. The use of chymotrypsin and IgG illustrates this property is
conserved across a wide range of
molecular weights;
Figure 6 illustrates the in vitro release of dissolved NBQX (a) and FGF-2 (A)
from 2:3 HMW RAMC. The
release of the therapeutic molecules NBQX and FGF-2 from the gel is also
Nokia%
Figure 7. illustrates cumulative release normalized to the amount of
therapeutic agent encapsulated in
particles. The open symbols indicate release from free particles for (a)
dbeAMP, (b) EGF, (c) a-
chymotrypsin, (d) Iga, and filled symbols from composite 2:3 BIM RAMC. Release
of polymer micro or
nanopartiole encapsulated drugs is unexpectedly slower and more linear when
embedded in BMW RAMC
that in aqueous suspension. The fast Fickian diffusion of dissolved drugs
illustrated in Figures 5 and 6
indicates the sustained release is not the result of slow release through the
hydrogel.
Figure 8 illustrates graphically the results found in the Method of Medical
Treatment Section hereafter. It
includes plots of Dorsal Horn Blood Vessel Counts v. Distance from Epicenter
(pun) for artificial cerebral
spinal fluid; 1:3 HMW HAMC; 1:3 BMW RAMC + Blank V; and 1:3 MEW RAMC +FGF2NP.
Examples
The present release formulation provides longer term release suitable for
combination neuroregenerative and
neuroprotective strategies. High molecular weight blends of hyaluronic acid
(HA) or a derivative thereof
and methylcellulose (MC) (RMW RAMC) which remain injectable and are stable for
more than 28 days in
vitro are described To achieve longer-term release profiles, formulations of
drug loaded poly(lactic-co-
glycolic acid) (PLGA) nano- and micro particles were dispersed in the IIMW
RAMC gel. Use of the drug
delivery platform was demonstrated using six therapeutic molecules or models
thereof, shown in Table 1, for
periods of between 1 and 28 days. This contribution demonstrates the composite
HMW RAMC hydrogel is a
flexible, localized drug delivery platform for the evaluation of therapies
targeting protection and repair of the
injured spinal cord.
Table 1. Molectges released from HAM HAPAC
Mo Desiredlecular Mauro-
Miura-
Molecule (model) Treatment
Weight (ighnoll protectiveregenerative
Term
MCP( 0.336 days
discAMP 0.469 = weeks
EDF 6.2 = weeks
F6F-2 17 days
NeurotropNo-3 (ce-Chymotrypsin) 29(25) = weeks
end-NogoA (Ie) 150(350) = weeks
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Recombinant human basic fibroblast growth factor (FGF-2, >95% wt) was
purchased from
Biovision (Santa Clara, USA). 2,3-Dioxo-6-nitro-1,2,3,4-
tetrahydrobenzo[f]quinoxaline-7-
sulfonamide disodium salt (sodium NBQX, >98% wt) was purchased from A.G.
Scientific (San
Diego, USA). a-Chymotrypsin (type II from bovine pancreas), human IgG, and
N6,2'-0-
dibutyryladenosine 3',5'-cyclic monophosphate sodium salt (dbcAMP) were
purchased from Sigma-
Aldrich (Oakville, CA). Recombinant human epidermal growth factor (EGF) was
purchased from
Peprotech TM (Rocky Hill, USA).
Sodium hyaluronate of 1700 kg/mol was purchased from FMC Biopolymer (Sandvika,
Norway) and
of 2600 kg/mol from Lifecore (Chaska, USA). Methylcellulose of 300 kg/mol was
purchased from
Shin-Etsu (Tokyo, Japan). Poly(DL-lactic-co-glycolic acid) 50:50 of inherent
viscosity 0.15-0.25
dL/g and methylcellulose (13 kg/mol) were purchased from Sigma-Aldrich.
Poly(DL-lactic-co-
glycolic acid) 50:50 of inherent viscosity 0.20 dL/g and 0.37 dL/g were
purchased from Durect
(Cupertino, USA). Poly(vinyl alcohol), 6 kg/mol and 80% mol hydrolyzed, was
purchased from
Polysciences Inc. (Warrington, USA). Polystyrene (220 nm, 510 nm, 830 nm, 3.09
Am, 15.5 inn)
and poly(acrylic acid) (60nm) particles were received as suspensions in water
from Bangs
Laboratories (Fishers, USA), scrubbed of surfactant by exposure to cation
exchange resin
(Amberlyst-15, Sigma-Aldrich), neutralized with 1M NaOH and lyophilized
(Labconco, Kansas
City, USA) prior to use.
HPLC grade dichloromethane (DCM) and dimethyl sulfoxide (DMSO) were supplied
by Caledon
Labs (Georgetown, CA). All buffers were made with distilled and deionized
water prepared using a
Millipore Milli-RO 10 Plus and Milli-Q UF Plus at 18M0 resistance (Millipore,
Bedford, USA).
Phosphate buffered saline powder was purchased from MP Biomedicals (pH 7.4,
9.55 g/L, Solon,
USA). Artificial cerebrospinal fluid (aCSF) at a pH of 7.4 was prepared as
previously described
[Gupta, D., C.H. Tator, and M.S. Shoichet, Fast-gelling injectable blend of
hyaluronan and
methylcellulose for intrathecal, localized delivery to the injured spinal
cord. Biomaterials, 2006.
27( 1 1 ) : p. 2370-9]. All other solvents and reagents were supplied by Sigma-
Aldrich and used as
received.
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Preparation of HAMC hydrogels and composite hydrogels
Physical hydrogel blends of hyaluronan (HA) and methylcellulose (MC) were
prepared in the
following compositions in aCSF; 2 wt% 1700 kg/mol HA, 7 wt% 13 kg/mol MC (2:7
HAMC); 1
wt% 2600 kg/mol HA, 3 wt% 300 kg/mol MC (1:3 HMW HAMC); 2% 2600 kg/mol HA, 3
wt%
300 kg/mol MC (2:3 HMW HAMC); and 3 wt% 2600 kg/mol HA, 3 wt% 300 kg/mol MC
(3:3
HMW HAMC). In each case MC was mechanically dispersed using a planetary mixer
(Flacktek
Inc., Landrum, USA) and left to dissolve overnight at 4 C. HA was then added
to the MC solution,
dispersed, and dissolved in the same way. Cold solutions were centrifuged to
remove entrained
bubbles, resulting in transparent hydrogels.
Drug loaded hydrogels were prepared by dispersing dry drug formulations or
concentrated, buffered
solutions in HAMC using a planetary mixer and left to dissolve overnight at 4
C before use. Drug
loaded PLGA particles and model polystyrene particles were dispersed in HAMC
immediately prior
to use.
Rheological characterization of HAMC hydrogels
The storage and loss moduli of 2:7 HAMC and the HMW HAMC hydrogels were
determined as a
function of oscillation frequency on an AR-1000 rheometer fitted with a 40 mm,
2 cone and plate
geometry (TA Instruments, New Castle, USA). An amplitude sweep was performed
to confirm that
the frequency and strain were within the linear viscoelastic region.
Temperature was controlled at
37 C using the integrated Peltier plate and sample evaporation was minimized
using a solvent trap.
After 5 min equilibration the frequency sweep was conducted from 0.1-100 rad/s
at 12% strain for
all materials.
In vitro stability and swelling of composite HAMC hydrogels
Approximately 150 mg of HMW HAMC loaded with 0, 25, or 75 mg/mL of 220 nm or
830 nm
polystyrene nanoparticles was deposited into pre-weighed polypropylene sample
tubes, weighed,
warmed to 37 C, and combined with 800 AL of warm aCSF. Samples were incubated
at 37 C on a
rotary shaker at 2 Hz throughout the experiment and the aCSF buffer sampled
with total replacement
after: 1 hour, 6 hours, and 1, 3, 7, 14, 21, and 28 days. Recovered buffer was
sonicated to disrupt
residual hydrogel (Sonics, Newtown, USA) and create a uniform nanoparticle
suspension. The
concentration of 220 nm particles was determined by a turbidity assay at 500
nm on an Agilent
8453TM spectrophotometer (Agilent Technologies, Santa Clara, USA).
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The swelling ratio, Q, of HAMC and composites was determined by accurately
weighing each gel
sample in the stability study after the aCSF had been removed, correcting for
residual buffer, and
dividing by the original hydrated sample mass. To compare the degradation of
composites with
different swelling characteristics, normalized Q was defined as:
Q(t)
(1)
Qmax
where ana, is the maximum recorded swelling ratio of the sample.
Determination of drug diffusivity in HMW HAMC
The diffusivity of a given drug in HMW HAMC, D, was normalized to its
diffusivity in water, Do,
to determine the impact of HMW HAMC on molecular diffusion [N.A. Peppas, C.T.
Reinhart,
Solute diffusion in swollen membranes.] . A new theory, J. Membr. Sci. 15 (3)
(1983) 275-287.
Approximations of D for NBQX, a-chymotrypsin, and IgG in HMW HAMC were
estimated
according to the one dimensional, unidirectional, thin film approximation for
non-swelling samples
at short times [J. Crank, The Mathematics of Diffusion, Clarendon Press,
Oxford, 1956.].
Normalized diffusion coefficients were then determined with respect to
previously reported values
of Do according to:
( \ 2
t
71"
M00 2
¨ =
(2)
D, Dot
Where Mr/M. is the cumulative mass of drug detected at time, t, divided by the
total mass released
and 1 is the sample thickness.
Approximately 100 mg of HAMC with a drug loading of 100-1000 p.g/mL was
deposited into a
cylindrical sample tube to yield a gel with thickness of 0.3 cm and one
exposed surface. Samples
were warmed to 37 C and combined with 900 AL of warm aCSF. The buffer was
sampled with total
replacement at 0.5, 1, 2, and 5 hours and analyzed and described below.
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Drug release from HAMC hydrogels
Release profiles of each particle encapsulated drug were obtained by
depositing approximately 150
mg of drug loaded 2:3 HMW HAMC into polypropylene sample tubes, warming to 37
C, and
adding 600 AL warm aCSF. Samples were incubated at 37 C on a rotary shaker at
2 Hz throughout
the study and the aCSF buffer sampled with total replacement after 0.5, 1, 3,
7, 14, 21, and 28 days.
Particle loads were typically 30-40 mg PLGA per gram of composite. Release
profiles of dissolved
drugs were obtained in the same manner with more frequent sampling, typically
at 0.5, 1, 3, 6, 24,
and 72 hours. All sample aliquots were immediately frozen and stored at -20 C
until analysis. IgG,
a-chymotrypsin, and EGF concentrations were determined by the bicinchoninic
acid assay (BCA)
(Thermo Fisher Scientific, Rockford, USA); NBQX and dbcAMP by UV absorbance at
425 and 273
nm, respectively; and FGF-2 by ELISA (R&D Systems, Minneapolis, USA). Each
sample was
thawed immediately prior to assay and clear supernatant analyzed for dissolved
drug.
Preparation and characterization of drug loaded PLGA particles
PLGA microparticle synthesis was optimized for each factor encapsulated and
thus synthesis varied
slightly for each factor, described below for EGF, dbcAMP, IgG and a-
chymotrypsin. Microparticle
size was determined by laser diffraction (Mastersizer 2000TM, Malvern
Instruments, Malvern, UK):
nanoparticle size was determined by dynamic light scattering (Zetasizer Nano
ZSTM, Malvern
Instruments). Encapsulation efficiency was defined as the fraction of drug
detected per unit mass of
particle compared to the theoretical maximum. Particle yield is the mass of
recovered particulate
PLGA adjusted for drug content, divided by the initial PLGA mass. Drug loading
is the mass
fraction of drug in the particles expressed as microgram of drug per milligram
of particles.
Preparation of EGF loaded PLGA microparticles
EGF loaded microparticles were prepared by a water-oil-water double emulsion
(w/o/w) method
with an inner aqueous phase of 100 AL, 20 mg/mL EGF in PBS, an organic phase
of 1.5 mL, 100
mg/mL PLGA (0.15-0.25 dL/g) in DCM and an outer aqueous phase of 50 mL, 10
mg/mL PVA and
100 mg/mL NaCl. The primary emulsion was created by 10 s of vortexing
(Scientific Industries,
Bohemia, USA) followed by 15 s of sonication. The secondary emulsion was
formed by addition of
the outer aqueous phase and homogenization by a Kinematica PT3000Tm
(Brinkmann, Mississauga,
CA). The double emulsion was then added to 150 mL of a 100 mg/mL NaC1 and 1
mg/mL PVA
solution and stirred for 4 h at room temperature. EGF loaded PLGA
microparticles were isolated
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CA 02703807 2010-05-12
and washed 4 times by centrifugation, lyophilized, irradiated with 2.5 kGy
gamma rays, and stored
at -20 C.
EGF content was determined by degrading the PLGA in 1 M NaOH for 24 h at 37
C, centrifuging
the resulting suspension and assaying the supernatant for total protein using
the BCA assay.
Preparation of dbcAMP loaded PLGA microparticles
dbcAMP loaded microparticles were prepared from a w/o/w double emulsion, with
an inner aqueous
phase of 75 pt, 267 mg/mL dbcAMP in ddH20, an organic phase of 600 AL, 217
mg/mL PLGA
(0.20 dL/g) in a 75:25 v/v solution of DCM and acetone, and an outer aqueous
phase of 25 mL, 25
mg/mL PVA and 100 mg/mL NaCl. The primary and secondary emulsions were created
by 45 s of
sonication and homogenization, respectively. The double emulsion was then
added to 200 mL of a
100 mg/mL NaC1 and 2.5 mg/mL PVA solution and stirred for 3 h at room
temperature. dbcAMP
loaded PLGA microparticles were isolated and washed with ddH20 over a 200 nm
nylon filter,
lyophilized, irradiated with 2.5 kGy gamma rays, and stored at -20 C.
dbcAMP content was determined by DCM/water solvent extraction. PLGA was
dissolved in DCM
and extracted 3x. Aqueous dbcAMP concentration was determined by UV absorbance
at 273nm on
a Nanodrop NID-1000TM (Thermo Fisher Scientific).
Preparation of IgG and ot-Chymotrypsin loaded PLGA nanoparticles
IgG loaded nanoparticles were prepared from a w/o/w double emulsion, with an
inner aqueous phase
of 100 AL, 10 mg/mL IgG in aCSF, an organic phase of 0.9 mL, 50 mg/mL PLGA
(0.15-0.25 dL/g)
and 0.5 mg/mL Pluronic F127TM in DCM, and an outer aqueous phase of 3 mL, 25
mg/mL PVA.
The primary emulsion was created by 10 min of sonication over ice. The
secondary emulsion was
formed by addition of the outer aqueous phase and sonication for a further 10
min over ice. The
double emulsion was then added to 40 mL of a 25 mg/mL PVA solution and stirred
for 20 h at room
temperature. a-Chymotrypsin loaded nanoparticles were produced in an identical
manner with the
addition of 1 mg/mL of DCM to the final aqueous volume prior to combination
with the double
emulsion. Protein loaded PLGA nanoparticles were isolated and washed 4 times
by
ultracentrifugation, lyophilized, and stored at -20 C.
27
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Protein content was determined based on a method by Wong [Wong, H.M., Wang,
J.J., and Wang,
C, In Vitro Sustained Release of Human Immunoglobulin G from Biodegradable
Microspheres. Ind.
Eng. Chem. Res., 2001. 40: p. 933-948]. Briefly, nanoparticles were dissolved
in DMSO at 37 C
and then diluted with 50 mM NaOH. The resulting suspension was allowed to
settle and the
supernatant assayed for total protein using the BCA assay.
Statistical analysis
Data are expressed as means standard deviation unless otherwise noted.
Comparisons of groups of
means were determined by ANOVA and pairs of mean by Student's t-test where
appropriate.
Significance was assigned at p<0.05.
Rheology of Hyaluronan and Methylcellulose Hydrogels
While previous 2:7 HAMC formulations degraded in vivo within 4-7 days [Kang,
C.E., et al., A New
Paradigm for Local and Sustained Release of Therapeutic Molecules to the
Injured Spinal Cord for
Neuroprotection and Tissue Repair. Tissue Eng Part A, 2008], the goal here was
to develop a more
stable HAMC for longer term delivery while maintaining the properties of
injectability and fast
gelation. Higher molecular weight HA and MC provided enhanced stability.
Blends of 1-3% HA
(2600 kg/mol) and 3% MC (300 kg/mol) met the qualitative criteria of fast
gelation and injectability
through a 30G / 200 Am inner diameter needle and were compared to 2:7 HAMC by
rheology. The
frequency sweeps shown in Figure 1 were conducted at 37 C and revealed the new
compositions
were of a similar stiffness to 2:7 HAMC and that 2:3 HMW HAMC was most like
the original
hydrogel. This observation was of practical importance because the stiffness
of 2:7 HAMC
approached the upper limit of what can be injected in the intrathecal space
using the method
described by Jimenez-Hamann, et al [Jimenez Hamann, M.C., et al., Novel
intrathecal delivery
system for treatment of spinal cord injury. Exp Neurol, 2003. 182(2): p. 300-
9]. Referring to Figure
1, it should be noted that Storage (G', open symbols) and loss modulus (G",
filled symbols) of
injectable gels. 2:3 HMW HAMC (0) and 2:7 HAMC (0) are gels and behave
similarly at low
frequencies shown in the left graph. 1:3 HMW HAMC (o) and 3:3 HMW HAMC (A) are
shown in
the right graph.
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CA 02703807 2010-05-12
Hydrogel Swelling & Degradation
As a cross-linked hydrogel containing the polyelectrolyte HA, HAMC swells when
placed in a
reservoir of aCSF in vitro or CSF in vivo. It was thought possible that an
intrathecally injected
hydrogel, resting between the pia and arachnoid mater, may put pressure on the
spinal cord as the
material swells in CSF. The swelling ratio, Q, at early times and maximal
swelling ratio, an,a, were
both of interest because the spinal cord can tolerate larger gel volumes if
the swelling force is
progressively applied over longer periods [Uchida, K., et al., Progressive
changes in neurofilament
proteins and growth-associated protein-43 immunoreactivities at the site of
cervical spinal cord
compression in spinal hyperostotic mice. Spine, 2002. 27(5): p. 480-6]. It has
previously been
demonstrated that injections of 20 fit of collagen gel or 10 iLL of 2:7 HAMC
were safe in a rat
model of SCI [Gupta, D., C.H. Tator, and M.S. Shoichet, Fast-gelling
injectable blend of
hyaluronan and methylcellulose for intrathecal, localized delivery to the
injured spinal cord.
Biomaterials, 2006. 27(11): p. 2370-9, Kang, C.E., et al., A New Paradigm for
Local and Sustained
Release of Therapeutic Molecules to the Injured Spinal Cord for
Neuroprotection and Tissue
Repair. Tissue Eng Part A, 2008]. Although the maximum safe gel volume in vivo
has not been
well characterized, the 2:7 HAMC formulation reached a maximum swelling ratio,
Qm, of 2.2 at 3
days in vitro and was shown to be safe in vivo, suggesting that this value was
acceptable. By
comparison, at three days the ana, for 1:3 HMW HAMC was 1.4 and 2:3 HMW HAMC
was 1.8
whereas 3:3 HMW HAMC was 2.4, nominally higher than 2:7 HAMC. The increase in
ana, as a
function of HA concentration reflects an increase in osmotic pressure common
to polyelectrolytes
[Gerdin, B. and R. Hallgen, Dynamic role of hyaluronan (HYA) in connective
tissue activation and
inflammation. Journal of Internal Medicine, 1997. 242(1): p. 49-55] . Swelling
at early times,
shown at 6 hours in Figure 3, was similarly comparable between 2:7 HAMC and
the HMW HAMC
blends. The three HMW HAMC formulations met the swelling criteria, permitting
an in vivo
injection volume of ¨10 AL and maximum gel volume similar to 2:7 HAMC over
time.
As evidence of composite drug delivery the swelling behaviour of HAMC
composites was
determined at particle loads of 25 and 75 mg/mL, and particle diameters of 220
and 830 nm.
Polystyrene (PS) particles were used as model hydrophobic polymer beads to
simulate PLGA
microspheres because of their narrow size distribution and range of available
diameters. No
difference in swelling was observed as a function of particle diameter for any
HAMC formulation at
any loading (data not shown). As shown in Figure 2 for 220 nm particles
dispersed in various
HAMC hydrogels, these formulations with nanoparticles reached maximum swelling
on day 3, at
which time an effect of particle loading on swelling became significant.
29
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CA 02703807 2010-05-12
The swelling behaviour of the composites can be partially explained if the
particles are considered
as non-interacting spheres which increase the volume fraction of polymer,
v2,5, in the gel by
displacing aCSF. According to the Peppas-Merrill equation of hydrogel
swelling, Q decreases as v2,s
is increased [Peppas, N.A. and E.W. Merrill, Polyvinyl-Alcohol) Hydrogels -
Reinforcement of
Radiation-Crosslinked Networks by Crystallization. Journal of Polymer Science
Part a-Polymer
Chemistry, 1976. 14(2): p. 441-457]. For example, a 75 mg/mL particle load
increases the aqueous
MC content in HMW HAMC from 3.0% to 3.2% by displacing aCSF. This mechanism
accounts for
the reduced swelling in HMW HAMC when particle loading was increased from 25
to 75 mg/mL
for all blends, and why Q is independent of particle diameter, since only the
total mass of buffer
displaced is considered. It does not, however, account for the increase in Q
from zero to 25 mg/mL
particles in each HMW HAMC. This unexpected increase in swelling may be a
kinetic effect not
well described by equilibrium swelling theory. Importantly, both 1:3 and 2:3
HMW HAMC at 25
and 75 mg/mL swelled similarly to, or less than, the pre-existing drug
delivery system, supporting
the safety of these materials in vivo.
The graphs in Figure 2 show that each of the HMW HAMC formulations swell
similarly or less than
2:7 HAMC after six hours. All materials reached a maximum swelling ratio at 3
days. The traces
are; 0 mg/mL (m), 25 mg/mL (o), and 75 mg/mL (a) of 220 nm polystyrene
particles dispersed in
the hydrogel.
The in vitro degradation of the three HMW HAMC blends with and without PS
particles was
followed by measuring the swelling ratio over time relative to Qm. For the HMW
HAMC swelling
traces in Figure 3, 1:3 HMW HAMC was most stable, followed by 2:3 and 3:3 HMW
HAMC.
Since HAMC gels through physical cross-links between methylcellulose, the
higher concentration of
MC in the minimally swollen gels results in slower degradation. This is
supported by Peppas and
Merrill who showed that lower swelling ratios are the result of more physical
cross-links and are
therefore predictive of slower degradation / dissolution [Peppas, N.A. and
E.W. Merrill, Polyvinyl-
Alcohol) Hydrogels - Reinforcement of Radiation-Crosslinked Networks by
Crystallization. Journal
of Polymer Science Part a-Polymer Chemistry, 1976. 14(2): p. 441-457].
Dispersing nanoparticles in all HMW HAMC resulted in significantly slower
degradation regardless
of particle loading (25 or 75 mg/mL) or diameter (220 or 830 nm). For 1:3 HMW
HAMC, inclusion
DOCSTOR: 1 938001 \1
CA 02703807 2010-05-12
of nanoparticles resulted in gels that retained 80% of anc, at 28 days
relative to gels alone that
retained only 60% of Qmax. For 2:3 and 3:3 HMW HAMC, dispersing hydrophobic
nanoparticles
stabilized the gel and left a majority of the composite intact when the blank
hydrogels had
completely degraded. For 2:3 HMW HAMC, which had degraded completely by 28
days, inclusion
of nanoparticle resulted in 70-80% retention of Qmax. For 3:3 HMW HAMC, which
had completely
degraded at 21 days, the inclusion of particles resulted in 60-80% retention
of ana, then, and 30-
60% at 28 days. As with initial swelling, particle diameter did not measurably
affect degradation.
Based on previously observed differences between HAMC degradation in vitro and
in vivo, where
2:7 HAMC was observed to degrade faster after intrathecal injection [C.E.
Kang, et al., A new
paradigm for local and sustained release of therapeutic molecules to the
injured spinal cord for
neuroprotection and tissue repair, TissueEng Part A 15 (3) (2009) 595-6041, it
was desirable that
the new drug delivery platform remain substantially intact at 28 days in
vitro. 1:3 HMW HAMC
met this criterion under all conditions and 2:3 HMW HAMC was satisfactory in
the presence of
nanoparticles.
The residual particle load was quantified for the 220 nm particle loaded
composites as a direct
measure of the composite's utility as a drug delivery platform. The residual
particle load, defined as
the mass of particles in the gel at time, t, divided by the initial particle
mass, was followed by
injecting 150 mg of nanoparticle loaded hydrogel in 800 AL of aCSF.
Substantial numbers of the
220 nm and larger particles were not predicted to diffuse from the composite
in the absence of
HMW HAMC degradation because the particles were significantly larger than the
mesh sizes
typically reported for hydrogels [Lin, C.C. and A.T. Metters, Hydrogels in
controlled release
formulations: network design and mathematical modeling. Adv Drug Deliv Rev,
2006. 58(12-13): p.
1379-408]. In Figure 4, a small release of less than 2% of nanoparticles was
seen from each of the
composites in the hours after injection, followed by a delay in particle loss
as the composites shrunk
with the reorganization of the MC hydrophobic network and formation of
optimized cross-links
[Schupper, N., Y. Rabin, and M. Rosenbluh, Multiple stages in the aging of a
physical polymer gel.
Macromolecules, 2008. 41(11): p. 3983-3994]. At 28 days the residual particle
load was 40% for
3:3 HMW HAMC, 84% for 2:3 HMW HAMC and 98% for 1:3 HMW HAMC, following the
degradation pattern of the hydrogels. Residual particle loads were not
dependent on initial mass
loading (25 or 75 mg/mL) after 28 days. These results were consistent with the
characterization of
1:3 HMW HAMC as having the lowest swelling ratio, greatest stability, highest
weight fraction of
MC, and lowest predicted molecular weight between cross-links.
31
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The mechanism of nanoparticle-mediated stabilization remains unclear, although
one possibility is
based on network optimization [Schupper, N., Y. Rabin, and M. Rosenbluh,
Multiple stages in the
aging of a physical polymer gel. Macromolecules, 2008. 41(11): p. 3983-3994],
supported by the
reduction in normalized swelling without loss of embedded particles in 1:3 HMW
HAMC over 28
days and 2:3 HMW HAMC over 14 days. Considering dynamic association of MC
chains, those
near the surface of the gel have fewer neighbouring hydrophobic regions and
are likely lost from the
gel by diffusion. Dispersed hydrophobic particles may slow this process by
associating with MC
[Saunders, F.L., Adsorption of Methylcellulose on Polystyrene Latexes. Journal
of Colloid and
Interface Science, 1968. 28(3-4): p. 475], slowing diffusion and resulting in
a higher fraction of MC
chains re-forming hydrophobic junctions and stabilizing the composite.
Figure 3 illustrates the degradation of HMW HAMC and composite HMW HAMC
quantified by the
swelling ratio and represented a change in composite volume relative to Q..
Degradation of
HMW HAMC (o) was faster than composite HMW HAMC of the same formulation in all
cases.
No significant effect of particle diameter on degradation was found. Open
circles are 25 mg/mL of
220 nm particles (0), and filled circles are 75 mg/mL.
Figure 5 illustrates the loss of particles from composite HMW HAMC as a strong
function of gel
composition. The traces are: 1:3 HMW HAMC (N), 2:3 HMW HAMC (A), 3:3 HMW HAMC
(*),
each loaded with 25 mg/mL of 220 nm PS particles.
To study the possible effect of embedded particles on the mechanical
properties of HMW HAMC
the injectability of composite HMW HAMC through a 30G needle from a Hamilton
250 AL glass
syringe was tested. Surfactant free suspensions of low polydispersity PS
particles ranging from 60
nm to 15.5 itm in diameter were lyophilized, added dry to each of the
hydrogels up to 150 mg/mL,
and mechanically dispersed. Given the inherent stiffness of HMW HAMC and the
reported
difficulty in evenly suspending hydrophobic particles in hydrogels [Liu, X.,
K. Nakamura, and A.M.
Lowman, Composite hydrogels for sustained release of therapeutic agents. Soft
Materials, 2003.
1(3): p. 393-408; Mackay, M.E., et al., General strategies for nanoparticle
dispersion. Science,
2006. 311(5768): p. 1740-3], large diameter and higher weight percent
formulations were expected
not to be injectable due to incomplete dispersion and occlusion of the needle
by particle aggregates.
Surprisingly, each formulation was injectable at room temperature. The
injection of higher
32
DOCSTOR: 1938001\1
CA 02703807 2010-05-12
concentrations and larger diameter particles enhances the utility of the drug
delivery platform by
both increasing the deliverable drug load and capturing the distinct release
profiles reported for
PLGA particles of different diameters [Sinha, V.R. and A. Trehan,
Biodegradable microspheres for
protein delivery. Journal of Controlled Release, 2003. 90(3): p. 261-280;
Soppimath, K.S., et al.,
Biodegradable polymeric nanoparticles as drug delivery devices. J Control
Release, 2001. 70(1-2):
p. 1-20.].
Drug Diffusivity in Hydrogels
Next examined were the diffusion of two high molecular weight proteins, IgG
(150 kg/mol) and a-
chymotrypsin (25 kg/mol), from HMW HAMC since they were most likely to be
restricted by
sieving effects [Hennink, W.E., et al., Controlled release of proteins from
dextran hydrogels.
Journal of Controlled Release, 1996. 39(1): p. 47-55]. It was desirable that
the normalized diffusion
coefficient, D/Do, be sufficiently large that the rate-limiting step in
release of particle encapsulated
drugs was not diffusion through the gel. In this manner long-term drug release
is controlled by the
particle formulation and the release profile is decoupled from molecular
diffusivity. Plotting the
fractional release of IgG and a-chymotrypsin against til2 yielded a linear
relationship, as predicted
for Fickian diffusion [Ritger, P.L. and N.A. Peppas, A Simple Equation for
Description of Solute
Release 1. Fickian and Non-Fickian Release from Non-Swellable Devices in the
Form of Slabs,
Spheres, Cylinders or Discs. Journal of Controlled Release, 1987. 5: p. 23-
36]. Applying a Do of
6.4 x 10-7 cm2/s for IgG [Cruise, G.M., D.S. Scharp, and J.A. Hubbell,
Characterization of
permeability and network structure of interfacially photopolymerized poly
(ethylene glycol)
diacrylate hydrogels. Biomaterials, 1998. 19(14): p. 1287-94] and an estimated
Do of 1.5 x 10-6
CM2/s for a-chymotrypsin [Han, J.H., et al., Lactitol-based poly(ether polyol)
hydrogels for
controlled release chemical and drug delivery systems. Journal of Agricultural
and Food Chemistry,
2000. 48(11): p. 5278-5282], the normalized diffusion coefficient of IgG
ranged from 0.04 in 3:3
HMW HAMC to 0.25 in 1:3 HMW HAMC and a-chymotrypsin ranged from ¨0.3 to ¨0.8
in the
same materials. It is clear from the non-zero intercept in Figure 5 that
swelling impacted drug
release at these early times as penetration of aCSF into the gel retards drug
release. This result was
expected given that for all gels within 3 hours Q for all gels exceeded 1.25,
the ratio above which
swelling is a significant factor in drug release [Ritger, P.L. and N.A.
Peppas, A Simple Equation for
Description of Solute Release 2. Fickian and Anomalous Release from Swellable
Devices. Journal
of Controlled Release, 1987. 5: p. 37-42] and reduced the calculated value of
D. The in vitro
estimation of D is therefore smaller than what can be expected in vivo where
the restrictive
33
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CA 02703807 2010-05-12
environment may prevent the gel from swelling more than 1.25x normal to the
spinal cord. Both
IgG and a-chymotrypsin diffuse from HMW HAMC relatively quickly and release is
predicted to be
complete within 24 hours based on the planar geometry of the gels after
injection in vivo. This rate
is advantageous, being slow enough to allow prolonged release of dissolved
molecules yet fast
enough that release of PLGA encapsulated molecules is not diffusion limited.
Figure 4 illustrates that the slope of IgG and a-chymotrypsin release from HMW
HAMC decreases
as HA concentration increases, indicating HA slowed diffusion. DID, was lowest
for both
molecules in 3:3 HMW HAMC, on the order of 0.3 and 0.04 for a-chymotrypsin and
IgG,
respectively. The non-zero intercept indicates that swelling has slowed drug
release, reducing D.
The traces are; 1:3 HMW HAMC (0), 2:3 HMW HAMC (0), and 3:3 HMW HAMC (A). a-
Chymotrypsin release is represented by filled symbols and IgG release by open
symbols.
Drug Release from Composite Hydrogels
After establishing that both 1:3 and 2:3 HMW HAMC composites met the in vitro
design criteria,
the release of six therapeutic molecules from 2:3 HMW HAMC was assessed. These
were chosen
because the HA content matched 2:7 HAMC and is thought to be an important
component in
HAMC's anti-inflammatory action. NBQX and FGF-2 have been shown to play a role
in
neuroprotection and thus fast release from the hydrogel is desirable. Both
molecules were released
in a diffusion-limited manner from 2:3 HMW HAMC, shown in the inset of Figure
5, with a
normalized diffusion coefficient on the order of 0.1. The release profiles,
also plotted versus linear
time in Figure 5 for ease of comparison to particle mediated release in Figure
6, show that NBQX is
released from the hydrogel faster that FGF-2. The diffusive release in vitro
is likely slower than that
in vivo, where the release rate is bounded by fast unidirectional diffusion
from a thin film [Crank, J.,
The Mathematics ofDiffusion. 1956, Oxford: Clarendon Press] and the slow
diffusion of many
molecules through tissue [Kang, C.E., et al., A New Paradigm for Local and
Sustained Release of
Therapeutic Molecules to the Injured Spinal Cord for Neuroprotection and
Tissue Repair. Tissue
Eng Part A, 2008, Jimenez Hamann, M.C., C.H. Tator, and M.S. Shoichet,
Injectable intrathecal
delivery system for localized administration of EGF and FGF-2 to the injured
rat spinal cord. Exp
Neurol, 2005. 194(1): p. 106-19, Krewson, C.E., M.L. Klarman, and W.M.
Saltzman, Distribution of
nerve growth factor following direct delivery to brain interstitium. Brain
Res, 1995. 680(1-2): p.
196-206].
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Four neuroregenerative molecules (or models thereof) were encapsulated in
formulations of PLGA
particles and individually dispersed in 2:3 HMW HAMC for long term release.
Encapsulation in
PLGA particles is widely used to control temporal drug release. In these
systems the release profile
results from drug diffusion through pores in the hydrogel composite formed by
dissolution of
entrapped protein and degradation of PLGA. The data for the encapsulation of
dbcAMP, EGF, a-
chymotrypsin, and IgG are summarized in Table 2 and the drug release profiles
from free particles
are reported in Figure 7.
Figure 6 illustrates the in vitro release of dissolved NBQX (m) and FGF-2 (=)
from 2:3 HMW
HAMC is diffusion limited (inset) and complete within 1 and 4 days,
respectively (main graph).
Fractional release normalized to the total initial drug mass for NBQX and to
total detectable protein
for FGF-2.
Table 2. Synthesis of drug loaded PLGA particles
Molecular Encapsulation Drug Loading
Molecule
Particle Yield (%) Particle Size (pm)
Weight (kg/mol) Efficiency (%) (pg/mg)
dbcAMP 0.469 51 68 50 37 14
EGF 6.2 36 4.8 63 10 2
a-Chymotrypsin 25 32 7 53
0.285, polydisperse
IgG 150 56 14 61 0.272
0.103
Figure 7 illustrates the cumulative release normalized to the amount
encapsulated in the particles.
The open symbols indicate release from free particles for (a) dbcAMP, (b) EGF,
(c) a-chymotrypsin,
(d) IgG, and filled symbols from composite 2:3 HMW HAMC. The release of
individual drugs
from particles dispersed in 2:3 HMW HAMC is longer than from the corresponding
particles alone.
As shown in Figure 7, extended release over 28 days was achieved for these
formulations. In a
majority of these trials, the initial burst release characteristic of PLGA
particles was reduced and
subsequent release was typically more linear and longer from PLGA particles
dispersed in HMW
HAMC composite gels than from the particles alone. This observation was in
agreement with the
work of Ying et al [Ying, L., et al., In vitro evaluation of lysozyme-loaded
microspheres in
thermosensitive methylcellulose-based hydrogel. Chinese Journal of Chemical
Engineering, 2007.
15(4): p. 566-572], although in the current case only a small portion of the
effect was due to
molecular diffusion through the gel. In HMW HAMC diffusion can only prolong
release for 1-4
DOCSTOR: 1938001\1
CA 02703807 2010-05-12
days after release from the particle, as reported in Figure 5 for IgG and a-
chymotrypsin and in
Figure 6 for NBQX and FGF-2. Visual observation of particle loaded composites,
which remained
opaque during these experiments, suggested decreased degradation of PLGA
particles dispersed in
HMW HAMC relative to particles dispersed in aqueous buffer. In each case the
particles, which
scatter light and cause the composite to appear opaque and white, remained
intact. If the PLGA
were hydrolyzed at the rate reported for free particles [Dunne, M., I.
Corrigan, and Z. Ramtoola,
Influence of particle size and dissolution conditions on the degradation
properties of polylactide-co-
glycolide particles. Biomaterials, 2000. 21(16): p. 1659-68], opacity would
have decreased over
time as particle size and number were reduced. If instead the rate of PLGA
degradation was reduced
in the composite, drug release would be slower than from particles alone and
the composites would
remain opaque as observed. This may be the result of MC adsorption onto the
particle surface,
resulting in slower diffusion of drug and degraded PLGA through pores in the
hydrogel composite,
mechanisms supported by the reduction in burst release observed for dbcAMP, a-
chymotrypsin, and
IgG. A MC / particle interaction is also supported by the increase in HMW HAMC
stability on
particle addition, discussed under the heading Hydrogel Swelling &
Degradation. The atypical
behaviour of the EGF loaded particles may indicate more of the drug was near
the particle surface
and less subject to variation in PLGA degradation.
As an injectable drug delivery platform, the particle-loaded hydrogels allow
different drug
formulations to be dispersed within the hydrogel to create a combination
therapy while maintaining
control of the resulting release profiles. Existing drug and particle
formulations can be directly
dispersed in the hydrogel without modification, and the combination of fast
diffusion limited release
from the dissolved phase and slow release of particle-borne drugs can be
exploited such that release
can be substantially decoupled from molecular weight. It has been demonstrated
that the in vitro
release of the small molecules NBQX and dbcAMP are over 1 and 28 days,
respectively, and the
release of proteins spanning 6-150 kg/mol over 4-28 days. Particle loads up to
15 wt% were
injectable, resulting in a deliverable drug load of 1.1-10.1 mg per gram of
composite as a function of
PLGA particle properties. The 2:3 HMW HAMC used to evaluate drug release
retained greater than
80% of the initial particle load after 28 days, suggesting that a high
percentage of the drug loaded
will be locally delivered at the site of injection.
This specific disclosure is directed toward a clinically acceptable drug
delivery platform for the
treatment of spinal cord injury. There are described a series of physical
hydrogels composed of
36
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CA 02703807 2010-05-12
hyaluronan and methylcellulose which meet the design criteria of
injectability, safe swelling,
satisfactory diffusivity of molecules up to 150 kg/mol, high residual particle
load, and significantly
slower in vitro degradation relative to earlier reports. The slow degradation
rate of HMW HAMC
with particles dispersed therein suggests this as a platform for 28 day
combination drug therapy.
Composites with particle loads up to 15 wt% and 0.06-15.5 jim diameter
remained injectable for all
blends and that greater than 95% of the initial particle load was retained
after 28 days in vitro in 1:3
HMW HAMC. Utilizing a combination of diffusion limited and particle mediated
drug delivery,
release of six neuroprotective and neuroregenerative drugs from 1 to 28 days
was shown.
Method of medical treatment
In spinal cord injury, small blood vessels are ruptured and lead to hemorrhage
and edema in the
tissue parenchyma [ I. Koyanagi, C.H. Tator, and P.J. Lea, Three-dimensional
analysis of the
vascular system in the rat spinal cord with scanning electron microscopy of
vascular corrosion
casts. Part 1: Normal spinal cord, Neurosurgery. 33 (1993) 277-83; discussion
283-4; I. Koyanagi,
C.H. Tator, and P.J. Lea, Three-dimensional analysis of the vascular system in
the rat spinal cord
with scanning electron microscopy of vascular corrosion casts. Part 2: Acute
spinal cord injury,
Neurosurgery. 33 (1993) 285-91; discussion 292]. Reduced blood flow leads to
widespread
ischemia, ultimately contributing to tissue degeneration in a large area
surrounding the initial injury.
To induce blood vessel growth in tissue and thus reduce the ischemic impact,
the angiogenic protein
fibroblast growth factor 2 (FGF2) [R. Montesano, J.D. Vassalli, A. Baird, R.
Guillemin, and L. Orci,
Basic fibroblast growth factor induces angiogenesis in vitro, Proc Natl Acad
Sci U S A. 83 (1986)
7297-301; Y. Shing, J. Folkman, C. Haudenschild, D. Lund, R. Crum, and M.
Klagsbrun,
Angiogenesis is stimulated by a tumor-derived endothelial cell growth factor,
Journal of Cellular
Biochemistry. 29 (1985) 275-87; M. Relf, S. LeJeune, P.A. Scott, S. Fox, K.
Smith, R. Leek, A.
Moghaddam, R. Whitehouse, R. Bicknell, and A.L. Harris, Expression of the
angiogenic factors
vascular endothelial cell growth factor, acidic and basic fibroblast growth
factor, tumor growth
factor beta-1, platelet-derived endothelial cell growth factor, placenta
growth factor, and
pleiotrophin in human primary breast cancer and its relation to angiogenesis,
Cancer Res. 57
(1997) 963-9] was delivered with composite HAMC into the intrathecal space.
Physical hydrogel blends of hyaluronan (HA) and methyl cellulose (MC) were
prepared in aCSF at
1 wt% 2600 kg/mol HA and 3 wt% 300 kg/mol MC as previously reported (1:3 HMW
HAMC [
M.D. Baumann, C.E. Kang, J.C. Stanwick, Y. Wang, H. Kim, Y. Lapitsky, and M.S.
Shoichet, An
37
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injectable drug delivery platform for sustained combination therapy, J Control
Release. (2009)].
Briefly, MC and HA were sequentially mechanically dispersed in aCSF and
allowed to dissolve at
4 C. To form composite 1:3 HMW HAMC, slurries of particles were added together
for a final
concentration of 100 mg/mL. PLGA nanop articles were prepared from a w/o/w
double emulsion as
previously described [Li, Y., et al., Effects of the AMPA receptor antagonist
NBQX on the
development and expression of behavioral sensitization to cocaine and
amphetamine.
Psychopharmacology (Berl), 1997. 134(3): p. 266-76], either as blank
nanoparticles or encapsulating
FGF2.
All animal procedures were performed in accordance with the Guide to the Care
and Use of
Experimental Animals (Canadian Council on Animal Care) and protocols were
approved by the
Animal Care Committee of the Research Institute of the University Health
Network. Sprague
Dawley rats were anaesthetized by inhalation of halothane and a laminectomy
was performed at the
Ti -T2 vertebral level. All animals sustained a moderate compressive spinal
cord injury at T2 using
a modified aneurysm clip calibrated to a closing force of 26 g for 60 s, as
previously described [
A.S. Rivlin and C.H. Tator, Effect of duration of acute spinal cord
compression in a new acute cord
injury model in the rat, Surg Neurol. 10 (1978) 38-43]. After a laminectomy,
the dura was punctured
with a bevelled 30G needle at T2, and a 30G blunt-tipped needle was inserted
into the intrathecal
space. Each animal received 10 IA of either: 1) artificial cerebrospinal fluid
(aCSF), 2) HMW
HAMC, 3) composite HAMC, or 4) composite HAMC loaded with FGF2 pre-heated to
37 C.
Following injection, the overlying muscles and fascia were sutured closed and
the rats were
ventilated with pure oxygen and placed under a heat lamp for recovery.
Buprenorphine was
administered every 12 h for 3 days post-surgery for pain management. The
animals were sacrificed
28 days after surgery and perfused intracardially with 4% paraformaldehyde
under terminal
anaesthesia with sodium pentobarbital. A 20 mm segment of spinal cord
encompassing the injury
site was harvested from each animal, dehydrated in 30% sucrose, and stored at -
80 C until
cryoprocessing. Cords were then cut cross-sectionally into 20 um sections and
every third section
rostrocaudally from the epicenter stained with SMI-71, a marker for mature
blood vessels. Blood
vessels were then counted in the dorsal horns of the spinal cord. Statistical
significance was
determined by ANOVA followed by Tukey's post-hoc test.
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CA 02703807 2016-10-17
Results for localized and sustained delivery FGF2 with Composite HAMC
Following a moderate spinal cord injury, FGF2 was delivered from composite
HAMC on the dorsal
surface of the injured spinal cord. FGF2 promotes endothelial cell
proliferation, which leads to
blood vessel formation [ C. Basilico and D. Moscatelli, The FGF family of
growth factors and
oncogenes, Adv Cancer Res. 59 (1992) 115-65; A. Bikfalvi, S. Klein, G.
Pintucci, and D.B. Rifkin,
Biological roles offibroblast growth factor-2, Endocr Rev. 18 (1997) 26-45].
Mature blood vessels
counted in the dorsal horns of the injured spinal cord showed that FGF2
delivery with composite
HAMC results in angiogenesis. Angiogenesis can reduce the ischemic injury in
tissue.
Thus it has been demonstrated that sustained delivery of FGF2 with the
composite HAMC provides
enhanced blood vessels in the dorsal horns after spinal cord injury. FGF2 is
an angiogenic
molecule, and can be substituted with FGF1, vascular endothelial growth factor
(VEGF), or platelet
derived growth factor (PDGF) to induce angiogenesis in tissue.
While the invention has been described in connection with specific embodiments
thereof, it will be
understood that it is capable of further modifications and this application is
intended to cover any
variations, uses, or adaptations of the invention following, in general, the
principles of the invention
and including such departures from the present disclosure as come within known
or customary
practice within the art to which the invention pertains and as may be applied
to the essential features
hereinbefore set forth, and as follows in the scope of the appended claims.
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