Note: Descriptions are shown in the official language in which they were submitted.
CA 02705083 2016-12-05
TUNABLE SUSTAINED RELEASE OF A SPARINGLY SOLUBLE HYDROPHOBIC
THERAPEUTIC AGENT FROM A HYDROGEL MATRIX
FIELD:
The incorporation of polymeric excipients into an injectable hydrogel matrix,
for example,
methyl cellulose in the case of a hydrogel matrix comprising hyaluronan and
methylcellulose
(HAMC) has been found to increase the solubility of sparingly soluble
hydrophobic drugs and
tune their rate of release. The hydrogel matrix may also include other
sparingly soluble
hydrophobic food or cosmetic agents.
BACKGROUND:
Traumatic spinal cord injury (SCI) is a devastating condition for which there
is no cure.
Currently, there is no standard of care for traumatic brain injury or stroke.
There is also 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, 20091. There is a need generally to
provide therapies
for all traumatic injuries to the central nervous system. The initial
mechanical trauma, termed
the primary injury, causes damage to blood vessels and localized cell death
[C.H. Tator,
Strategies for recovery and regeneration after brain and spinal cord injury.
Injury Prevention 8
(2002) Iv33-1v361. These in turn lead to excitotoxicity, inflammation,
hemorrhage, vasospasm,
and edema, which result in functional deficits in the patient [J. Krieglstein,
Excitotoxicity and
neuroprotection. Eur J Pharm Sei 5(4) (1997) 181-187; A. Scriabine, T
Schuurman, J. Traber,
Pharmacological Basis for the Use of Nimodipine in Central Nervous-System
Disorders. Faseb
3(7) (1989) 1799-1806]. These pathological events can occur from days to
months after injury
and are known as the secondary injury [A. Arun, B.S.R. Reddy, In vitro drug
release studies from
the polymeric hydrogels based on HEA and HPMA using 4-{(E)-[(3Z)-3-(4-
1
CA 02705083 2010-05-21
(acryloyloxy)benzylidene)-2-hexylidenelmethyl}lphenyl acrylate as a
crosslinker. Biomaterials
26(10) (2005) 1185-1193; MD. Norenberg, I Smith, A. Marcillo, The pathology of
human spinal
cord injury: Defining the problems. J Neurotraum 21(4) (2004) 429-440]. Both
neuroregenerative and neuroprotective therapeutics are being pursued to limit
the devastation that
occurs after injury, yet their delivery remains challenging.
There are three common delivery strategies ¨ systemic, pump/catheter, and
bolus ¨ yet each has
its drawbacks. Systemic delivery is limited because most molecules cannot
cross the blood-
spinal cord barrier and those that do may have profound systemic side effects
[C.H. Tator,
Strategies for recovery and regeneration after brain and spinal cord injury.
Injury Prevention 8
(2002) Iv33-1v36]. The external pump/catheter system pumps drugs from a
reservoir into the
intrathecal space through a catheter. While a constant dose can be
administered, this method is
open to infection and has not been approved for long-term delivery in SCI
patients in the USA.
Bolus injection into the intrathecal space is compromised by cerebral spinal
fluid (CSF) flow,
which disperses the drug, thereby requiring repeated administration.
Much research effort has been devoted to improving the therapeutic efficacy
and delivery of
hydrophobic drugs which is often limited by low solubility [L. Zema, A.
Maronii, A. Foppoli, L.
Palugan, ME. Sangalli, A. Gazzaniga, Different HPMC viscosity grades as
coating agents for an
oral time and/or site-controlled delivery system: An investigation into the
mechanisms governing
drug release. J Pharm Sci 96(6) (2007) 1527-1536; Z.G. He, D.F. Zhong, XY.
Chen, XH. Liu, X
Tang, L.M Zhao, Development of a dissolution medium for nimodipine tablets
based on
bioavailability evaluation. Eur J Pharm Sci 21(4) (2004) 487-491; E. Lu, Z.Q.
Jiang, Q.Z.
Zhang, XG. Jiang, A water-insoluble drug monolithic osmotic tablet system
utilizing gum arable
as an osmotic, suspending and expanding agent. J Control Release
92(3)(2003)375-382 ]. In
solid pharmaceutical formulations, polymeric excipients similar to MC, such as
hydroxypropyl
methylcellulose or poly(vinylpyrrolidone), are incorporated into the drug
particles to increase the
solubility of sparingly soluble drugs [Z G. He, D.F. Zhong, XY. Chen, XH. Liu,
X Tang, L.M
Zhao, Development of a dissolution medium for nimodipine tablets based on
bioavailability
evaluation. Eur J Pharm Sci 21(4) (2004) 487-491; H Wen, KR. Morris, K Park,
Synergic
effects of polymeric additives on dissolution and crystallization of
acetaminophen. Pharmaceut
Res 25(2) (2008) 349-358; B.C. Hancock, M Parks, What is the true solubility
advantage for
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CA 02705083 2010-05-21
amorphous pharmaceuticals? Pharmaceut Res 17(4) (2000) 397-404; ME. Matteucci,
B.K
Brettmann, TL. Rogers, E.J. Elder, R. 0. Williams, KP. Johnston, Design of
potent amorphous
drug nanoparticles for rapid generation of highly supersaturated media.
Molecular
Pharmaceutics 4(5) (2007) 782-793; S.L. Raghavan, A. Trividic, A.F. Davis, J
Hadgraft,
Crystallization of hydrocortisone acetate: influence of polymers. Int J Pharm
212(2) (2001) 213-
221]. This is typically achieved by disrupting the crystalline drug particle
structure [B.C.
Hancock, M Parks, What is the true solubility advantage for amorphous
pharmaceuticals?
Pharmaceut Res 17(4) (2000) 397-404; ME. Matteucci, B.K Brettmann, T.L.
Rogers, E.J. Elder,
R. 0. Williams, K.P. Johnston, Design of potent amorphous drug nanoparticles
for rapid
generation of highly supersaturated media. Molecular Pharmaceutics 4(5) (2007)
782-793],
thereby producing a less-stable amorphous drug particle that can be up to
orders of magnitude
more soluble than the crystalline drug [B.C. Hancock, M Parks, What is the
true solubility
advantage for amorphous pharmaceuticals? Pharmaceut Res 17(4) (2000) 397-404;
ME.
Matteucci, B.K. Brettmann, T.L. Rogers, El Elder, R. 0. Williams, K.P.
Johnston, Design of
potent amorphous drug nanoparticles for rapid generation of highly
supersaturated media.
Molecular Pharmaceutics 4(5) (2007) 782-793; V.M Rao, J.L. Haslam, V.J.
Stella, Controlled
and complete release of a model poorly water-soluble drug, prednisolone, from
hydroxypropyl
methylcellulose matrix tablets using (SBE)(7M)-beta-cyclodextrin as a
solubilizing agent. J
Pharm Sci 90(7) (2001) 807-8161 These polymeric excipients are also used as
stabilizing
additives in supersaturated solutions [ME. Matteucci, B.K. Brettmann, TL.
Rogers, E.J. Elder,
R. 0. Williams, K.P. Johnston, Design of potent amorphous drug nanoparticles
for rapid
generation of highly supersaturated media. Molecular Pharmaceutics 4(5) (2007)
782-793; S.L.
Raghavan, A. Trividic, A.F. Davis, J. Hadgraft, Crystallization of
hydrocortisone acetate:
influence ofpolyrners. Int J Pharm 212(2) (2001) 213-221; S.L. Raghavan, A.
Trividic, A.F.
Davis, J. Hadgraft, Effect of cellulose polymers on supersaturation and in
vitro membrane
transport of hydrocortisone acetate. Int J Pharm 193(2) (2000) 231-237; K
Yamashita, T
Nakate, K Okimoto, A. Ohike, Y. Tokunaga, R. Ibuki, K Higaki, T. Kimura,
Establishment of
new preparation method for solid dispersion formulation of tacrolimus. Int J
Pharm 267(1-2)
(2003) 79-.91; S.L. Raghavan, K. Schuessel, A. Davis, J. Hadgraft, Formation
and stabilisation of
triclosan colloidal suspensions using supersaturated systems. Int J Pharm
261(1-2) (2003) 153-
158; U Kumprakob, J. Kawakami, I. Adachi, Permeation enhancement of ketoprofen
using a
supersaturated system with antinucleant polymers. Biological & Pharmaceutical
Bulletin 28(9)
3
CA 02705083 2010-05-21
(2005) 1684-1688] and gels [Si. Raghavan, A. Trividic, A.F. Davis, J.
Hadgraft, Crystallization
of hydrocortisone acetate: influence of polymers. Int J Pharm 212(2) (2001)
213-221; S.L.
Raghavan, A. Trividic, A.F. Davis, J. Hadgraft, Effect of cellulose polymers
on supersaturation
and in vitro membrane transport of hydrocortisone acetate. Int J Pharm 193(2)
(2000) 231-237]
for oral and transdermal drug delivery, where a layer of adsorbed,
"antinucleating" polymer on
the surface of the nascent crystal is believed to inhibit further
crystallization of the drug [X G.
Ma, J. Taw, C.M Chiang, Control of drug crystallization in transdermal matrix
system. Int J
Pharm 142(1) (1996) 115-119; P.N. Kotiyan, P.R. Vavia, Eudragits: Role as
crystallization
inhibitors in drug-in-adhesive transdermal systems of estradiol. Eur J Pharm
Biopharm 52(2)
(2001) 173-180].
Given the limitations associated with current delivery strategies as described
previously, a
minimally-invasive injectable, thermally-responsive hydrogel comprised of
hyaluronan (HA) and
methylcellulose (MC) was designed for sustained and localized release. [D.
Gupta, C.H. Tator,
MS. Shoichet, Fast-gelling injectable blend of hyaluronan and methylcellulose
for intrathecal,
localized delivery to the injured spinal cord. Biomaterials 27(11) (2006) 2370-
2379]. This
physical blend has been shown to be safe and provide greater neuroprotection
when used to
deliver erythropoietin to the intrathecal space than traditional delivery
strategies such as
intraperitoneal and intrathecal bolus [Kong CE, Poon PC, Tator CH, Shoichet
MS, A New
Paradigm for Local and Sustained Release of Therapeutic Molecules to the
Injured Spinal Cord
for Neuroprotection and Tissue Repair. Tissue Engineering Part A 15(3) (2009)
595-604].
US parent Patent Application No.11/410831 describes a polymer blend comprising
an inverse
thermal gelling polymer and an anionic polymer, for example HAMC that exists
as a gel. This
polymer mixture has a shorter time to gelation than the inverse gelling
polymer alone, and may
be used alone or as a drug delivery vehicle for many applications. In
particular, the polymer
mixture can be used for localized, targeted delivery of pharmaceutical agents
upon injection
providing sustained release. A particular use of this polymer mixture is in
delivery of a
therapeutic agent 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. This can be achieved, for example, by
injection of the matrix (or
mixture) into the intrathecal space, a fluid-filled space wherein cerebral
spinal fluid flows.
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CA 02705083 2010-05-21
U.S. Patent No. 6,335,035 ('035) to Drizen, et al. 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 injectable implant for
filling up wrinkles,
thin lines, skin cracks and scars for reparative or plastic surgery, aesthetic
dermatology and for
filling up 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.
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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 gel-forming polymer mixtures 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 etal. Biomacromolecules (2001) 2:856-863, Cho et al.
International
Journal of Pharmaceutics (2003) 260:83-91, Kim et al. Journal of Controlled
Release (2002)
80:69-77, Tate et al. Biomaterials (2001) 22:1113-1123, and Silver etal.,
Journal of Applied
Biomaterials (1994) 5:89-98).
SUMMARY:
Whether for the delivery of medical or non-medical applications, such as
cosmetic or food
applications, the delivery of hydrophobic molecules is difficult to achieve in
water-based
systems. Methyl cellulose itself solubilises the hydrophobic molecules (drugs
or otherwise),
thereby increasing the amount of hydrophobic molecule released. This is
beneficial for the
delivery of hydrophobic drugs for medical applications and hydrophobic
molecules for cosmetic
or food applications. The MC hydrogel may be used alone for this purpose ¨
that is MC mixed
with hydrophobic molecules ¨ or MC may be used together with hyaluronan, an
anionic polymer
that lowers the gelation temperature of MC and allows the hydrogel blend to be
shear-thinning.
Importantly, other hydrophobic cellulose derivatives will be useful in
solubilising hydrophobic
molecules.
This disclosure relates to a hydrogel matrix comprising an aqueous mixture
methylcellulose or
other water soluble hydrophobic cellulose derivative, in which particles of at
least one selected
size of at least one sparingly soluble hydrophobic agent are dispersed and
solubilised, which is
blended with an anionic polysaccharide or a derivative thereof to form the
hydrogel matrix, the
solubility of the sparingly soluble hydrophobic agent being increased in the
presence of the
methylcellulose or other water soluble hydrophobic cellulose derivative, which
together with the
selected particle sizes determine the sustained release profile of the at
least one sparingly soluble
hydrophobic agent from the hydrogel matrix.
In another aspect, the disclosure provides a method for the manufacture of a
hydrogel matrix as
described above, wherein the matrix has a tunable therapeutic agent release
profile comprising
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the steps of 1) providing particles of at least one selected size of the at
least one sparingly soluble
hydrophobic agent ; 2) dispersing and solubilising the particles in an aqueous
solution of
methylcellulose or other water soluble hydrophobic cellulose derivative; and
3) blending the thus
formed dispersion with an anionic polysaccharide or a derivative thereof to
form the hydrogel
matrix containing the at least one solubilised sparingly soluble hydrophobic
agent; wherein, the
solubility of the sparingly soluble hydrophobic agent is increased in the
presence of the
methylcellulose or other water soluble hydrophobic cellulose derivative which
together with the
selected particle sizes of the sparingly soluble hydrophobic agent determine
the sustained release
profile of the at least one sparingly soluble hydrophobic agent from the
hydrogel matrix.
This disclosure also provides for a hydrogel matrix without the use of the
anionic polysaccharide
or a derivative thereof. In this form the hydrogel matrix comprises an aqueous
mixture of
methylcellulose or other water soluble hydrophobic cellulose derivative, in
which particles of at
least one selected size of at least one sparingly soluble hydrophobic agent
are dispersed and
solubilised, the solubility of the sparingly soluble hydrophobic agent being
increased in the
presence of the methylcellulose or other water soluble hydrophobic cellulose
derivative, which
together with the selected particle sizes determine the sustained release
profile of the at least one
sparingly soluble hydrophobic agent from the hydrogel matrix
There is also disclosed in a related aspect, a method for the manufacture of a
hydrogel matrix as
described above, wherein the matrix has a tunable agent release profile
comprising the steps of 1)
providing particles of at least one selected size of the at least one
sparingly soluble hydrophobic
agent; 2) dispersing and so lubilising the particles in an aqueous mixture of
methylcellulose or
other water soluble hydrophobic cellulose derivative; and 3) increasing the
temperature of the
thus formed solution to form the hydrogel matrix containing the at least one
solubilised sparingly
soluble hydrophobic therapeuti agent; wherein, the solubility of the sparingly
soluble
hydrophobic agent is increased in the presence of the methylcellulose or other
water soluble
hydrophobic cellulose derivative which together with the selected particle
sizes of the sparingly
soluble hydrophobic agent determine the sustained release profile of the at
least one sparingly
soluble hydrophobic agent from the hydrogel matrix.
While the primary purpose of the hydrogel matrix described herein is
pharmaceutical, there are
many sparingly soluble hydrophobic substances that can benefit from this form
of delivery and
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administration, in applications such as cosmetics and food. In these forms,
the hydrogel
composite may also 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. Thus,
the sparingly soluble hydrophobic agent may be selected from therapeutic, food
and cosmetic
agents, and combinations thereof
In one embodiment, there is provided a hydrogel matrix comprising: (1) an
aqueous solution of
methylcellulose, in which particles of at least one hydrophobic therapeutic
agent chosen from
the group consisting of a drug and a hydrophobic vitamin are predissolved; (2)
solid particles
of the at least one hydrophobic therapeutic agent; and (3) hyaluronan or a
derivative thereof
that forms the hydrogel matrix, the solubility of the hydrophobic therapeutic
agent being
increased in the presence of the methylcellulose, which together with the size
of the solid
particles determine the sustained release profile of the at least one
hydrophobic therapeutic
agent from the hydrogel matrix.
In another embodiment, there is provided a hydrogel matrix that provides a
biphasic release
profile comprising: a blend of (1) a pre-made aqueous solution of
methylcellulose or a methyl
cellulose derivative chosen from the group consisting of: hydroxypropyl
methylcellulose,
hydroxypropyl methylcellulose phthalate, 2,3-di-O-methyl-6-0-benzylcellulose,
2,3-di-0-
benzy1-6-0-methylcellulose, and hydroxypropyl methylcellulose succinate and
solubilized
particles of at least one hydrophobic vitamin or hydrophobic drug that are pre-
dissolved in the
methylcellulose or the methylcellulose derivative prior to blending where the
methylcellulose
or the methylcellulose derivative enhances the solubility of the solubilized
particles of the at
least one hydrophobic vitamin or hydrophobic drug; (2) a first solid
particulate form of a
particle size of the at least one hydrophobic vitamin or hydrophobic drug; and
(3) a hyaluronan
or a derivative thereof; wherein the blend forms the hydrogel matrix; and
wherein the hydrogel
matrix provides the solubilized particles of the at least one hydrophobic
vitamin or
hydrophobic drug as a faster rate than the first solid particle form of a
particle size of the at
least one hydrophobic vitamin or hydrophobic drug and wherein the water
solubility of the at
least one hydrophobic vitamin or hydrophobic drug is increased in presence of
the
methylcellulose or the methylcellulose derivative and wherein the solubilized
particles and the
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particle size of the first solid particulate form of the at least one
hydrophobic vitamin or
hydrophobic drug create a biphasic release of the at least one hydrophobic
vitamin or
hydrophobic drug.
DETAILED DESCRIPTION:
The term "solubilization" as used herein is meant to have its ordinary meaning
which is
generally understood to be "to make a substance more soluble or soluble in
water"
It is well-known in the art that MC gels as temperature increases. This
gelation process is
entropically driven by MC coming together to form hydrophobic interactions and
water being
liberated from interactions with MC. It is also well-known that the gelation
temperature of MC
can be reduced by the addition of salt. We have previously shown that the
gelation
temperature of MC can also be reduced by the addition of anionic
polysaccharides, such as
hyaluronan, which acts as a viscosity enhancer as well.
The specific use of hyaluronan (HA) and methylcellulose (MC), HAMC for the
sustained
release of sparingly soluble hydrophobic drugs is shown in this disclosure
with nimodipine, a
hydrophobic, sparingly-soluble vasodilator and calcium channel blocker used
for treating
central nervous system (CNS) disorders [A. Scriabine, T Schuurman, I Traber,
Pharmacological Basis for the Use of Nimodipine in Central Nervous-System
Disorders. Faseb
J 3(7) (1989) 1799-1806; Y.S.R. Krishnaiah, P. Bhaskar, V. Satyanarayana,
Penetration-
enhancing effect of ethanol-water solvent system and ethanolic solution of
carvone on
transdermal permeability of nimodipine from HPMC gel across rat abdominal
skin.
Pharmaceutical Development and Technology 9(1) (2004) 63-74]. The
incorporation of
polymeric excipients into an injectable hydrogel (e.g., MC in the case of
HAMC) has been
found to increase the solubility of sparingly soluble drugs, such as
nimodipine, and tune their
rates of release. The rate of nimodipine dissolution in MC solution is slow,
and depends on the
initial drug particle size. Experimental and model analyses indicate that
these differences in
particle dissolution kinetics are reflected in the nimodipine release profiles
from HAMC, and
can be exploited in tailoring drug release rates.
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Thus based on these results, injectable hydrogel matrices can accelerate the
delivery of sparingly
soluble hydrophobic drugs, and can yield highly tunable release profiles that
are dependent on
how the drugs are introduced into the in situ gelling scaffold and on their
particle sizes.
The data in the Examples presented herein indicate that the solubilization of
the sparingly soluble
hydrophobic pharmaceutical agent is mediated by the adsorption of MC to the
hydrophobic drug
particle. The adsorption is supported by (1) the steric stabilization of the
drug particles observed
in the presence of MC and not HA, and (2) the slow drug particle dissolution,
which is more than
1000 times slower than it would be if it were controlled by solution- or gel-
phase diffusion. The
slow dissolution suggests that a particle-bound polymer layer is acting as a
dissolution barrier.
Once the drug is dissolved, it diffuses at a rapid rate that is characteristic
of a small molecule
rather than a micelle or molecular aggregate. This suggests that the increased
solubility is not due
to solubilization in micelle-like MC aggregates, but is rather caused by some
interfacial
interaction between the MC and the solid drug particle (i.e., adsorption of MC
to the hydrophobic
drug particle). Because MC adsorption is, at least in part likely driven by
hydrophobic
interactions ¨ where hydrophobic segments of MC bind to the hydrophobic
surface of the drug
particle ¨ this indicates that the solubilization effect can be extended to
other hydrophobic drugs.
The literature has relatively few reports of hydrophobic interactions, one
example being Wen et
al. "Hydrogen bonding interactions between adsorbed polymer molecules and
crystal surface of
acetaminophen", J. Colloid Interface Sci (2005) 325-335 ¨ that ascribe HPMC
and PVP
adsorption to drug particles to hydrogen bonding alone, and do not discuss the
hydrophobic
interactions at all.
More specifically, an injectable hydrogel matrix, comprised of hyaluronan and
methylcellulose
(HAMC), can be used for localized, sustained delivery of growth factors for
treatment of spinal
cord injury (SCI) and other injuries to the CNS such as traumatic brain injury
and stroke. To
better understand the ability of HAMC for the delivery of small molecules, the
release of
sparingly soluble neuroprotectant, nimodipine, was investigated experimentally
and via
continuum modeling. This revealed that the MC in HAMC increased the solubility
of this
sparingly soluble drug by over an order of magnitude, and enabled highly
tunable release profiles
to be achieved by varying the method by which the drug was introduced into the
matrix.
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When nimodipine was introduced into HAMC in solubilized form, it was rapidly
released from
the matrix within 8 hours. Conversely, when solid nimodipine particles were
blended into
HAMC in particulate form, the release rates were greatly reduced, giving rise
to complete release
over 2 ¨3 days for small, sub-micron particles, and longer times for large,
100 gm particles. The
nimodipine particle-loaded gels yielded particle size-dependent, biphasic
release profiles, which
reflected rapid release of the solubilized drug followed by the slow,
dissolution-limited release of
solid nimodipine. This demonstrates that injectable hydrogel matrices can act
as polymeric
excipients that accelerate the delivery of poorly soluble hydrophobic drugs
and yield highly
tunable release rates.
The anionic polysaccharide or a derivative thereof may comprise from about 100
to about 7,000
kg/mol and the methylcellulose or other water soluble hydrophobic cellulose
derivative may
comprise from about 2 to about 3,000 kg/mol. The ratio of anionic
polysaccharide or a derivative
thereof to the methylcellulose or other water soluble hydrophobic cellulose
derivative may
comprise from about 1:20 to about 1:1 w/w.
The amount of anionic polysaccharide or a derivative thereof may comprise from
about 0.5% to
about 5.0% by weight and the methylcellulose or other water soluble
hydrophobic cellulose
derivative may comprise from about 1.0% to about 20.0% by weight, more
particularly, from
about 5.0% to about 10% by weight of the matrix.
The dispersed hydrophobic therapeutic agent particles may be micro particles
or nanoparticles.
As used herein, "microparticles" refers to particles having a diameter of less
than 1.0 mm, and
more specifically between 1 and 1000 microns. Microparticles include
microspheres, which are
typically solid spherical microparticles. As used herein, "nanoparticles"
refers to particles or
structures in the nanometer range, typically from about 1 nm to about 1000 nm
in diameter. The
microparticles may be prepared in accordance with known methods, such as
sonication as
exemplified herein. The nanoparticles may be produced in accordance with
suitable known
methods as well. The selection of the size or sizes of the particles will
determine the amount and
rate of solubilisation that occurs, as well as the delivery profile for each
sparingly soluble
hydrophobic therapeutic agent present.
CA 02705083 2010-05-21
Given that the methyl cellulose or other cellulose derivative is hydrophobic
and the
pharmaceutical agent is selected from sparingly soluble hydrophobic drugs, the
same solubilising
effect will be found for any sparingly soluble hydrophobic therapeutic agent
selected, and
particularly for those that are specifically mentioned herein.
The delivered therapeutic agent load from the matrix may be in the range of
from about 0.0001 to
about 30 wt% (drug mass as a percentage of the matrix).
In the hydrogel matrix, the aqueous solution may be selected from the group
comprising water,
saline, artificial cerebrospinal fluid, and buffered solutions.
The hydrogel matrix components can be modified to alter the degradation rate
of the hydrogel
matrix and, hence, affect the rate of release of the pharmaceutical agent from
the hydrogel matrix.
One such modification involves addition of salts to alter the gelation
temperature of the MC.
Another alternative to creating a more stable hydrogel matrix for slower
degradation is to
functionalize the polymers with thiol groups and acrylate groups. The hydrogel
matrix 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 the anionic
polysaccharide, in particular the 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 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 matrix.
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CA 02705083 2010-05-21
Another means to enhance sustained release of the pharmaceutical agent is to
take advantage of
ionic interactions between the therapeutic agent and the polymer. The highly
negatively charged
anionic gelling polymer engages in ionic interactions with positively charged
molecules. 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
matrix upon breakage
of the covalent bond or upon dissolution of the chain from the hydrogel matrix
network.
Methods of covalently bonding pharmaceutical agents to polymers may be
employed that 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 a polymer which can
be inverse
thermal gelling polymer in a properly formulated hydrogel matrix. 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. (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.
The other hydrophobic 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-methyl-6-0-benzylcellulose, 2,3-di-0- benzylcellulose, 2,3-di-O-
benzy1-6-0-
methylcellulose, 2,3,6-tri-O-benzylcellulo se, hydroxypropyl methylcellulose
acetate succinate,
0-242-(2-methoxyethoxy)ethoxy]acetyl cellulose.
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CA 02705083 2010-05-21
The derivatives of hyaluronan may be esters of hyaluronan resulting from
esterification with
different classes of alcohols such as aliphatic, cycloaliphatic and
heterocyclic.
The drug delivery hydrogel matrix 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 matrix 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 drug
delivery system can also be injected into tissue.
While nimopidine has been used to illustrate the present disclosure, it should
be understood that
the at least one sparingly soluble hydrophobic therapeutic agent may be
selected from any
suitable sparingly soluble hydrophobic pharmaceutical agent or other type of
agent for the non-
pharmaceutical applications. The group of pharmaceutical agents is exemplified
by the
following: analgesics and anti-inflammatory agents: aloxiprin, auranofin,
azapropazone,
benorylate, diflunisal, etodolac, fenbufen, fenoprofen calcim, flurbiprofen,
ibuprofen,
indomethacin, ketoprofen, meclofenamic acid, mefenamic acid, nabumetone,
naproxen,
oxyphenbutazone, phenylbutazone, piroxicam, sulindac.; anthelmintics:
albendazo le, bephenium
hydroxynaphthoate, cambendazole, dichlorophen, ivermectin, mebendazole,
oxamniquine,
oxfendazole, oxantel embonate, praziquantel, pyrantel embonate, thiabendazole;
anti-arrhythmic
agents: amiodarone HC1, disopyramide, flecainide acetate, quinidine sulphate:
anti-bacterial
agents: benethamine penicillin, cinoxacin, ciprofloxacin HC1, clarithromycin,
clofazimine,
cloxacillin, demeclocycline, doxycycline, erythromycin, ethionamide, imipenem,
nalidixic acid,
nitrofurantoin, rifampicin, spiramycin, sulphabenzamide, sulphadoxine,
sulphamerazine,
sulphacetamide, sulphadiazine, sulphafurazole, sulphamethoxazole,
sulphapyridine, tetracycline,
trimethoprim; anti-coagulants: dicoumarol, dipyridamole, nicoumalone,
phenindione; anti-
depressants: amoxapine, maprotiline HC1, mianserin HCL, nortriptyline HC1,
trazodone HCL,
trimipramine maleate; anti-diabetics: acetohexamide, chlorpropamide,
glibenclamide, gliclazide,
glipizide, tolazamide, tolbutamide, anti-epileptics: beclamide, carbamazepine,
clonazepam,
ethotoin, methoin, methsuximide, methylphenobarbitone, oxcarbazepine,
paramethadione,
phenacemide, phenobarbitone, phenytoin, phensuximide, primidone, sulthiame,
valproic acid;
anti-fungal agents: amphotericin, butoconazole nitrate, clotrimazole,
econazole nitrate,
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CA 02705083 2010-05-21
fluconazole, flucytosine, griseofulvin, itraconazole, ketoconazole,
miconazole, natamycin,
nystatin, sulconazole nitrate, terbinafine HC1, terconazole, tioconazole,
undecenoic acid; anti-
gout agents: allopurinol, probenecid, sulphin-pyrazone; anti-hypertensive
agents: amlodipine,
benidipine, darodipine, dilitazem HC1, diazoxide, felodipine, guanabenz
acetate, isradipine,
minoxidil, nicardipine HC1, nifedipine, nimodipine, phenoxybenzamine HC1,
prazosin HCL,
reserpine, terazosin HCL; anti-malarials: amodiaquine, chloroquine,
chlorproguanil HC1,
halofantrine HC1, mefloquine HC1, proguanil HC1, pyrimethamine, quinine
sulphate; anti-
migraine agents: dihydroergotamine mesylate, ergotamine tartrate, methysergide
maleate,
pizotifen maleate, sumatriptan succinate; anti-muscarinic agents: atropine,
benzhexol HC1,
biperiden, ethopropazine HC1, hyoscyamine, mepenzolate bromide,
oxyphencylcimine HC1,
tropicamide; anti-neoplastic agents and Immunosuppressants: aminoglutethimide,
amsacrine,
azathioprine, busulphan, chlorambucil, cyclosporin, dacarbazine, estramustine,
etoposide,
lomustine, melphalan, mercaptopurine, methotrexate, mitomycin, mitotane,
mitozantrone,
procarbazine HC1, tamoxifen citrate, testolactone; anti-protazoal agents:
benznidazole, clioquinol,
decoquinate, diiodohydroxyquino line, diloxanide furoate, dinitolmide,
furzolidone,
metronidazo le, nimorazo le, nitrofurazone, ornidazole, tinidazole; anti-
thyroid agents:
carbimazo le, propylthiouracil; anxiolytic, sedatives, hypnotics and
neuroleptics: alprazo lam,
amylobarbitone, barbitone, bentazepam, bromazepam, bromperidol, brotizo lam,
butobarbitone,
carbromal, chlordiazepoxide, chlormethiazole, chlorpromazine, clobazam,
clotiazepam,
clozapine, diazepam, droperidol, ethinamate, flunanisone, flunitrazepam,
fluopromazine,
flupenthixol decanoate, fluphenazine decanoate, flurazepam, haloperidol,
lorazepam,
lormetazepam, medazepam, meprobamate, methaqualone, midazo lam, nitrazepam,
oxazepam,
pentobarbitone, perphenazine pimozide, prochlorperazine, sulpiride, temazepam,
thioridazine,
triazo lam, zopicione; beta ¨blockers, acebutolol, alprenolol, atenolol,
labetalol, metoprolol,
nadolol, oxprenolol, pindolol, propranolol; cardiac inotropic agents:
amrinone, digitoxin, digoxin,
enoximone, lanatoside C, medigoxin; corticosteroids: beclomethasone,
betamethasone,
budesonide, cortisone acetate, desoxymethasone, dexamethasone, fludrocortisone
acetate,
flunisolide, flucortolone, fluticasone propionate, hydrocortisone,
methylpredniso lone,
predniso lone, prednisone, triamcinolone; diuretics: acetazolamide, amiloride,
bendrofluazide,
bumetanide, chlorothiazide, chlorthalidone, ethacrynic acid, frusemide,
metolazone,
spironolactone, triamterene; anti-parkinsonian agents: bromocriptine mesylate,
lysuride maleate;
gastro-intestinal agents: bisacodyl, cimetidine, cisapride, diphenoxylate HC1,
domperidone,
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famotidine, loperamide, mesalazine, nizatidine, omeprazole, ondansetron HCL,
ranitidine HCI,
sulphasalazine; histamine 1-1; receptor antagonists: acrivastine, astemizole,
cinnarizine, cyclizine,
cyproheptadine HC1, dimenhydrinate, flunarizine HCI, loratadine, meclozine
HCI, oxatomide,
terfenadine; lipid regulating agents: bezafibrate, clofibrate, fenofibrate,
gemfibrozil, probucol;
nitrates and other anti-anginal agents: amyl nitrate, glyceryl trinitrate,
isosorbide dinitrate,
isosorbide mononitrate, pentaerythritol tetranitrate; nutritional agents:
betacarotene, vitamin A,
vitamin B2, vitamin D, vitamin E, vitamin K; opioid analgesics: codeine,
dextropropyoxyphene,
diamorphine, dihydrocodeine, meptazinol, methadone, morphine, nalbuphine,
pentazocine; sex
hormones; clomiphene citrate, danazol, ethinyl estradiol, medroxyprogesterone
acetate,
mestranol, methyltestosterone, norethisterone, norgestrel, estradiol,
conjugated oestrogens,
progesterone, stanozolol, stibestrol, testosterone, tibolone; stimulants:
amphetamine,
dexamphetamine, dexfenfluramine, fenfluramine, mazindol; and mixtures of
hydrophobic drugs
may, of course, be used where therapeutically effective.
Examples of sparingly soluble hydrophobic molecules that may be used for food
formulations
include, but are not limited to hydrophobic neutracueticals, examples of which
include
flavonoids, isoflavones, and theobromine; hydrophobic vitamins, examples of
which are Vitamin
A (retinol, retinoids and carotenoids), Vitamin D (ergocalciferol and
cholecalciferol), Vitamin E
(tocopherol and tocotrienol) and Vitamin K (phylloquinone and menaquinone).
Examples of cosmetic agents that may be included in the hydrogel matrices
include a variety of
cosmetic additives that fall in the category of sparingly soluble hydrophobic
agents, examples of
which include but are not limited to fragrances, examples of which are
dihydromyrcenol,
limonene, benzyl acetate, Romascone; antibacterial agents, examples of which
are chlorhexidine,
triclosan; and for skin therapy, examples are some cosmetic applications also
use vitamins and
neutraceuticals listed above.
When more than one pharmaceutical agent or other agent is present, each will
have its own
release profile which will be determined by its solubilisation in the methyl
cellulose or other
water soluble cellulose derivative and its particle size or sizes. A mixture
of particle sizes may
also be selected to provide a more tailored release rate.
CA 02705083 2010-05-21
DETAILED DESCRIPTION OF EXEMPLIFIED EMBODIMENTS:
BRIEF DESCRIPTION OF THE DRAWINGS:
Fig. 1 illustrates a comparison of nimodipine solubility in MC and HA, with
different nimodipine
particle sizes: (*) 100- m nimodipine particles in 0.25 wt% HA; and (0) 100-
gm, (A) 900-nm
, and (0) 380-nm nimodipine particles in 7 wt% MC. The upper shaded
concentration range
indicates nimodipine solubility values achieved in 7 wt% MC, while the lower
shaded range
indicates aqueous nimodipine solubility reported in the literature [A.
Yoshida, M Yamamoto, T
Itoh, T. Irie, F. Hirayama, K Uekama, Utility of 2-Hydroxypropyl-Beta-
Cyclodextrin in an
Intramuscular Injectable Preparation of Nimodipine. Chemical & Pharmaceutical
Bulletin 38(1)
(1990) 176- 179]. Nimodipine Preparation;
Fig. 2 illustrates release of solubilized nimodipine plotted against time
(mean standard
deviation, n=3). The curve represents the model fit (Eq. 3b) for the release
data. The inset shows
that the drug release scales linearly with the square root of time, according
to Eq. 2, for the first
70-80% of released nimodipine; and
Fig. 3 illustrates a comparison of model predictions to experimental data for:
(*) solubilized
nimodipine,(0) 380 nm particulate nimodipine, (A) 900 nm particulate
nimodipine, (0) 100 gm
particulate nimodipine (mean standard deviation, n=3). The solid lines (¨)
depict the model
predictions, the dashed line (---) represents the slowest release predicted by
Eq. 6, and the shaded
region indicates the range of release profiles obtained by varying the
formulation of nimodipine.
This range is bounded by the Fickian model (upper limit) and that described by
Eq. 6 (lower
limit).
The specific use of hyaluronan (HA) and methylcellulose (MC), (HAMC) for the
sustained
release of low molecular weight drugs is exemplified in this disclosure with
nimodipine, a
hydrophobic, sparingly-soluble vasodilator and calcium channel blocker used
for treating central
nervous system (CNS) disorders [A. Scriabine, T Schuurman, I Traber,
Pharmacological Basis
for the Use of Nimodipine in Central Nervous-System Disorders. Faseb J 3(7)
(1989) 1799-1806;
Y.S.R. Krishnaiah, P. Bhaskar, V. Satyanarayana, Penetration-enhancing effect
of ethanol-water
solvent system and ethanolic solution of carvone on transdermal permeability
of nimodipine from
HPMC gel across rat abdominal skin. Pharmaceutical Development and Technology
9(1) (2004)
63-74]. The incorporation of polymeric excipients into an injectable hydrogel
(e.g., MC in the
16
CA 02705083 2010-05-21
case of HAMC) has been found to increase the solubility of sparingly soluble
drugs, such as
nimodipine, and tune their rate of release. The rate of nimodipine dissolution
in MC solution is
slow, and depends on the initial drug particle size. Experimental and model
analyses indicate that
these differences in particle dissolution kinetics are reflected in the
nimodipine release profiles
from HAMC, and can be exploited in tailoring drug release rates.
Thus based on these results, injectable hydrogel matrices can accelerate the
delivery of
hydrophobic, sparingly soluble drugs, or any food or cosmetic sparingly
soluble hydrophobic
agent and can yield highly tunable release profiles that are dependent on how
the drug or agent is
introduced into the in situ gelling scaffold.
An injectable hydrogel, comprised of hyaluronan and methylcellulose (HAMC),
shows promise
for localized, sustained delivery of growth factors for treatment of spinal
cord injury (SCI). To
better understand its potential for the delivery of small molecules, the
release of sparingly soluble
neuroprotectant, nimodipine, was investigated experimentally and via continuum
modeling. This
revealed that the MC in HAMC increased the solubility of this sparingly
soluble drug by over an
order of magnitude, and enabled highly tunable release profiles to be achieved
by varying the
method by which the drug was introduced into the matrix.
When nimodipine was introduced into HAMC in solubilized form, it was rapidly
released from
the scaffold within 8 hours. Conversely, when solid nimodipine particles were
blended into
HAMC in particulate form, the release rates were greatly reduced, giving rise
to complete release
over 2 ¨3 days for small, sub-micron particles, and longer times for large,
100 gm particles. The
nimodipine particle-loaded gels yielded particle size-dependent, biphasic
release profiles, which
reflected rapid release of the solubilized drug followed by the slow,
dissolution-limited release of
solid nimodipine. This indicates that injectable hydrogel matrices can act as
polymeric excipients
that accelerate the delivery of poorly soluble drugs and yield highly tunable
release rates.
EXAMPLES
To prepare nimodipine (Sigma Aldrich, Oakville, ON, Canada) for the release
study, two types of
nimodipine formulations were prepared at room temperature: (1) 0.5 mg/ml of
nimodipine
17
CA 02705083 2010-05-21
particles dissolved in 20 v/v % ethanol in water; and (2) 0.5 mg/ml of
nimodipine particles
dispersed in a 0.1 wt% methylcellulose (MC, Sigma Aldrich) solution in
artificial CSF (aCSF).
To vary the drug particle size, the particulate dispersions were either used
as received (non-
sonicated particles); or sonicated to reduce particle size for 1 or 5 min at
20 kHz, 40% amplitude,
using a Sonics Vibra Cell CV18TM tip sonicator (Sonics & Materials Inc.,
Newtown, CT, USA).
The nimodipine particles dispersed in MC were sized via dynamic light
scattering (DLS, Malvern
Zetasizer Nano ZSTM, Worcestershire, UK) for the sonicated sub-micron
particles, and laser
diffraction (Malvern Mastersizer 2000TM, Worcestershire, UK) for the non-
sonicated particles.
The particle dispersions were left for 3 days at room temperature to ensure
that particles were
stable in dispersion before incorporation into HAMC.
Incorporation of Nimodipine in HAMC
Thermogelling, sterile-filtered HAMC blends were prepared as previously
described [D. Gupta,
C.H. Tator, MS. Shoichet, Fast-gelling injectable blend of hyaluronan and
methylcellulose for
intrathecal, localized delivery to the injured spinal cord. Biomaterials
27(11) (2006) 2370-2379]
by dissolving hyaluronan (HA, NovamatrixTM, Sandvika, Norway) at 2 wt% into MC
at 7 wt%.
For nimodipine release studies, 100 pi of nimodipine formulations was added to
900 [11 of MC
dissolved in aCSF, yielding a 7 wt% MC/nimodipine dispersion, with a
nimodipine concentration
of 50 [ig/m1. HA was then dissolved into the MC/nimodipine dispersion at 2
wt%. The HAMC
solution was then vortexed (Vortex-Genie 2TM, 120V, 60Hz 0.65 amps, Scientific
Industries Inc.,
New York, NY, USA) until a clear, homogeneous, highly-viscous solution was
obtained [D.
Gupta, C.H. Tator, MS. Shoichet, Fast-gelling injectable blend of hyaluronan
and
methylcellulose for intrathecal, localized delivery to the injured spinal
cord. Biomaterials 27(11)
(2006) 2370-2379].
Nimodipine Dissolution Kinetics
To determine the effect of MC and HA on nimodipine dissolution in HAMC, 0.5
mg/ml
nimodipine dispersions (composed of either non-sonicated larger particles or 1
or 5 minute
sonicated sub-micron particles) was dispersed in 10 ml of either 7 wt% MC or
0.25 wt% HA in
aCSF. Here, the HA solution composition was adjusted to match the viscosity of
the 7 wt% MC
(ca. 0.5 Pa-s,), to maintain similar hydrodynamic conditions and mass transfer
coefficients
between the two receiving mediums. The dispersions were allowed to dissolve at
25 C under
18
CA 02705083 2010-05-21
constant stirring. The concentration of so lubilized nimodipine was tracked
over a 9 day period
using the NanoDropTM Spectrophotometer ND-100Tm (Wilmington, DE, USA, X = 275
nm,
extinction coefficient = 4.217 ml mg-1
Drug Release Studies
One hundred microliters of HAMC containing nimodipine was injected onto the
bottom of a 2 ml
eppendorf tube containing 900 I of aCSF at 37 C [D. Gupta, C.H. Tator, MS.
Shoichet, Fast-
gelling injectable blend of hyaluronan and methylcellulose for intrathecal,
localized delivery to
the injured spinal cord. Biomaterials 27(11) (2006) 2370-2379], thereby
mimicking the 10%
volume dilution expected in the intrathecal space of a rat animal model. At
various time points,
the supernatant was removed and replaced with the same volume of fresh aCSF.
To determine the
amount of drug released between each time point, the absorbance of supernatant
containing
released nimodipine was measured using the NanoDropTM Spectrophotometer.
Nimodipine Dissolution Kinetics
The particle size and the properties of the dissolution medium are expected to
influence the
dissolution kinetics and release profiles of nimodipine. To investigate the
effects of particle size
and the presence of MC and HA on the solubilization of nimodipine, the
dissolution of three
polydisperse populations of nimodipine particles with diameters of 100 30
pm, 900 60 nm
(second order polydispersity factor, PI = 0.48), and 380 20 nm (PI = 0.64),
as sized by laser
diffraction and DLS, were quantified over time in well-stirred MC and HA
solutions. To ensure
that the liquid phase mass transfer coefficients would be similar for the two
polymer solutions,
the viscosity of the HA solution was matched to that of 7 wt% MC, resulting in
an I-TA
concentration of 0.25 wt%. Figure 1 shows that the concentration of nimodipine
ultimately
solubilized in 7 wt% reached a plateau at approximately 30 ¨ 40 g/ml, which
may be interpreted
as its solubility limit. This solubility is an order of magnitude higher than
nimodipine's literature
aqueous solubility of 2 ¨4 g/m1 [A. Yoshida, M. Yamamoto, T. Itoh, T. Irie,
F. Hirayama, K.
Uekama, Utility of 2-Hydroxypropyl-Beta-Cyclodextrin in an Intramuscular
Injectable
Preparation of Nimodipine. Chemical & Pharmaceutical Bulletin 38(1) (1990) 176-
179], which
was also observed in 0.25 wt% HA. In contrast to previous work where polymeric
excipients
prevented crystallization of supersaturated drug solutions over time [Y.S.R.
Krishnaiah, P.
Bhaskar, V. Satyanarayana, Penetration-enhancing effect of ethanol-water
solvent system and
19
CA 02705083 2010-05-21
ethanolic solution of carvone on transdermal permeability of nimodipine from
HPMC gel across
rat abdominal skin. Pharmaceutical Development and Technology 9(1) (2004) 63-
74; S.L.
Raghavan, K Schuessel, A. Davis, J. Hadgraft, Formation and stabilisation of
triclosan colloidal
suspensions using supersaturated systems. Int J Pharm 261(1-2) (2003) 153-
158], here the
addition of MC leads to amplified drug solubilization. Conversely, the
presence of HA had no
measurable impact on nimodipine solubility.
The dissolution of each particle type in 7 wt% MC appeared to occur in two
stages. The first
stage corresponded to the solubility of nimodipine in water (2 - 4 gimp and
occurred within
minutes, whereas the second stage corresponded to its solubility in MC (30 -40
g/ml) and
occurred within several days. Interestingly, the larger 100-gm (and to a
lesser extent the 900-nm)
nimodipine particles showed an induction period in their dissolution profiles
in MC, where the
enhanced solubilization mediated by MC was not observed until 1 - 3 days into
the dissolution
process. The induction time increased with particle size. For all nimodipine
particles studied, the
plateau of solubilized drug was attained after approximately 1 week in MC
solution.
Visual observation of the nimodipine particles revealed that MC also affects
their dispersion
properties. All three particles formed large -0(1000 gm) aggregates in HA, and
smaller -0(10
gm) aggregates in MC (data not shown). This indicates that when dispersed in
MC solution the
large, 100-gm particles fragment and dissolve, while the small, sub-micron
particles undergo
some aggregation. The improved colloidal stability that is mediated by MC
suggests that MC
adsorbs to the surface of the nimodipine particles and prevents flocculation
of nimodipine into
larger particles through steric stabilization. This improved colloidal
stability is consistent with
the polymer-mediated stabilization reported for colloidal drug dispersions in
aqueous
hydroxypropyl methylcellulose solutions [S.L. Raghavan, K Schuessel, A. Davis,
J. Hadgraft,
Formation and stabilisation of triclosan colloidal suspensions using
supersaturated systems. Int J
Pharm 261(1-2) (2003) 153-158].
To further probe the mechanism of nimodipine dissolution in the presence of
MC, the mass
transfer coefficient, kin, for the slower second stage of particle dissolution
in MC was estimated
via [S. W. Bird RB, Lightfoot EN, Transport Phenomena, John Wiley and Sons,
2006]:
CA 02705083 2010-05-21
dC k a
A m cSat
(1)
dt V A
where a is the total surface area of the 10- m nimodipine particle flocs, V is
the volume of the
receiving MC solution, CAsat is the saturation concentration of nimodipine in
7 wt% MC, and
dCA/dt is the approximate slope of the dissolution curves estimated to be
¨0(10 ug/ml-day) from
Fig. 1. This analysis yields a mass transfer coefficient of km ¨0(10-6 cm/s),
which is more than
three orders of magnitude lower than the minimum mass transfer coefficient
predicted for
solution mass transfer-controlled dissolution. The km for solution mass
transfer-controlled
dissolution is ¨2 x 10-3 cm/s, estimated for the 10- um nimodipine aggregates
in the absence of
convection, where the Sherwood number (Sh) is equal to two [I. Tosum, Modeling
in Transport
Phenomena, a Conceptual Approach, 2nd ed., Elsevier, 2007]: Sh = kmdp/DA ,
where dp is the
drug particle diameter, and DA is the molecular diffusivity of the drug. This
suggests that the
solubilization of nimodipine is not limited by the solution mass transfer of
nimodipine, but is
rather governed by another slower process. A layer of adsorbed polymer may be
slowing down
the dissolution of nimodipine particles. The hypothesis that adsorption of MC
improves drug
solubility is supported by following: (1) the steric stabilization of the
nimodipine particles
observed in the presence of MC; and (2) the high diffusivity of nimodipine
observed in HAMC
(DA ¨ 0(10-5 cm2/s), see Section 3.2.1.). This high DA value is characteristic
of small molecule
diffusion, suggesting that once the nimodipine is solubilized in HAMC, it
remains in a molecular
state. Importantly, while the presence of MC retards the rate of drug particle
dissolution, the
solubility is enhanced. Thus, the increased amount of solubilized drug at the
beginning of the
release process (e.g., from < 4 to < 40 },tg/m1 nimodipine) should accelerate
the rate of drug
release when MC is present. From these results, the slow, particle size-
dependent process of
MC-mediated drug dissolution was expected to enable tunable acceleration of
hydrophobic drug
release from HAMC (and other similar injectable gels) by modulating the drug
fraction that is
solubilized at the beginning of the release profile. This can be achieved by
either varying the size
of the drug particles that are used in the hydrogel preparation or the time
period between the
preparation and application of the HAMC blend.
21
CA 02705083 2010-05-21
Nimodipine Release from HAMC
Release of Solubilized Nimodipine
HAMC may be classified as a matrix drug delivery system where nimodipine is
distributed
throughout the gel network [B.1V. Nalluri, C. Milligan, J.H. Chen, P.A.
Crooks, A.L. Stinchcomb,
In vitro release studies on matrix type transdermal drug delivery system of
naltrexone and its
acetyl prodrug. Drug Dev Ind Pharm 31(9) (2005) 871-877; C. C. Lin, A. T.
Metters, Hydrogels in
controlled release formulations: Network design and mathematical modeling. Adv
Drug Deliver
Rev 58(12-13) (2006) 1379-1408]. Solubilized nimodipine, which was
predissolved in ethanol to
produce a 50 ug/mlnimodipine and 2% v/v ethanol solution in HAMC, was fully
released within
8 h (Fig. 2). The square root scaling of the release profile (see Fig. 2
inset) suggests that it is
diffusion-controlled. For a planar geometry, such as the release of nimodipine
from the top of a
cylindrical HAMC gel, drug release can be estimated by the analytical
approximation [C.S.
Brazed, N.A. Peppas, Modeling of drug release from swellable polymers. Eur J
Pharm Biopharm
49(1) (2000) 47-58]:
M, 2 ID
M co L\ n-
(2)
where Mt /Mo, is the fraction of drug molecules released from the hydrogel at
time t, DA is the
diffusivity of the drug in the matrix, and L is the scaffold thickness [CS.
Brazel, NA, Peppas,
Modeling of drug release from swellable polymers. Eur J Pharm Biopharm 49(1)
(2000) 47-58].
For an estimated gel thickness of 0.37 cm, the fitted diffusivity value of 1.0
x le cm2is is
characteristic of the diffusion of small molecules and suggests that the drug
remains dissolved
during the release process. The proportionality to the square root of time is
maintained for the
first 70-80% of release [C. C. Lin, A. T. Metters, Hydrogels in controlled
release formulations:
Network design and mathematical modeling. Adv Drug Deliver Rev 58(12-13)
(2006) 1379-1408;
J. Siepmann, N.A. Peppas, Modeling of drug release from delivery systems based
on
hydroxypropyl methylcellulose (HPMC). Adv Drug Deliver Rev 48(2-3) (2001) 139-
157], after
which drug depletion affects the concentration gradient, thus reducing the
driving force for drug
release. This also supports our previous findings that diffusion is the
dominant mechanism of
drug release from HAMC [Y.S.R. Krishnaiah, P. Bhaskar, V. Satyanarayana,
Penetration-
enhancing effect of ethanol-water solvent system and ethanolic solution of
carvone on
22
CA 02705083 2010-05-21
transdermal permeability of nimodipine from HPMC gel across rat abdominal
skin.
Pharmaceutical Development and Technology 9(1) (2004) 63-74].
Release of Particulate Nimodipine
When nimodipine was introduced in particulate form, its release from HAMC was
significantly
slower than the soluble form. The complete release of 380 nm and 900 nm
nimodipine particle
formulations from HAMC was achieved at 48 h and 72 h, respectively (Fig. 3).
For 100 gm
nimodipine particles, only ¨40% of the drug was released from HAMC after 3
days, likely
because only a fraction of the total nimodipine is soluble and able to diffuse
from the gel at a
given time.
The release profiles obtained using gels loaded with nimodipine particles were
biphasic.
Submicron particles yielded a high initial burst release (ca. 80% for the 380
nm particles, and ca.
60% for the 900 nm particles), occurring within the first few hours, similar
to that of the release
of solubilized nimodipine. This initial burst release phase was followed by a
second slower
release phase, which takes place over 2-3 days. These two phases correspond to
the rapid release
of the drug that is solubilized at the beginning of the release process
followed by slower
dissolution-limited release of the drug that remains in particulate form.
Likewise, the 100- gm
particles yielded a 5 ¨ 10% burst release followed by the slow dissolution-
controlled release.
This burst release is consistent with the slower solubilization rates of
larger nimodipine particles
in MC solution, and indicates that the release profiles can be tuned over a
wide range of release
rates by varying the method by which sparingly soluble drug is introduced into
the gel.
Model Analysis of the Release Profiles
To analyze the release of nimodipine from HAMC, a generalized model was
developed based on
diffusion- and particle dissolution-controlled mass transport. It was assumed
that the nimodipine
particles were uniformly distributed within HAMC, and that the solubilized
drug concentration
and the radii of the dissolving drug particles varied with respect to both
time and spatial position
within the gel. Using these assumptions, the temporal variation in drug
particle size and
solubilized drug concentrations can be estimated using two coupled
differential equations:
dR ____ =¨k m MW A (cSai c
(3a)
dt PA A A
23
CA 02705083 2010-05-21
ac, a2c,
_______ = DA 2A + 47rk,õRI2ni, (CAsa` ¨ CA )
(3b)
at az
Here, CA is the drug concentration at specific spatial (z) and temporal (t)
points within the matrix,
CASat is the saturation concentration of the drug in the gel, and np is the
number of particles per
unit volume within the matrix (number of particles/cm3), MWA is the molecular
weight of the
drug, and PA is the density of the drug particle. DA is the diffusivity of
drug molecules in the
hydrogel matrix, and km is the mass transfer coefficient for drug particle
dissolution. R1 is the
drug particle radius, which varies with respect to time, t and position, z. R1
is a function of drug
particle position within the gel because dissolution is driven by the
concentration gradient of
dissolved molecules around the particle. For regions closer to the surface of
the gel, the drug
diffuses out more quickly compared to the interior regions of the gel, leading
to faster particle
dissolution. Equation 3a describes the dissolution of the drug particles over
time [NA. Peppas, A
Model of Dissolution-Controlled Solute Release from Porous Drug Delivery
Polymeric Systems.
Journal of Biomedical Materials Research 17(6) (1983) 1079-1087; MI. Cabrera,
JA. Luna,
R.JA. Grau, Modeling of dissolution-diffusion controlled drug release from
planar polymeric
systems with finite dissolution rate and arbitrary drug loading. Journal of
Membrane Science
280(1-2) (2006) 693-704]. Likewise, Equation 3b provides a microscopic
materials balance on
the solubilized drug in the gel matrix, where the change in the local
solubilized drug
concentration reflects the balance between the dissolution of the drug
particles and the diffusion
of the drug out of the gel. Using the appropriate boundary conditions, where
flux at the inner
boundary and the drug concentration at the outer boundary are both equal to
zero, this system of
equations was solved numerically with MATLABTm via finite difference
approximation. Using
the DA-value fitted to Equation 2 in Section 3.2.1 (1.0 x10-5 cm2/s) and an
approximate CASat_
value of 40 g/ml and initial particle diameter of 10 tirn, the numerical
solutions given in terms
of CA(z,t) and R1(t,z) (not shown) were obtained. These profiles were then
integrated over the
volume of the scaffold to generate the release curves showing the amount of
drug released over
time. The model release curves were fitted to the experimental release
profiles by varying the
fraction of the drug that was dissolved at the beginning of the release
experiment (fdissoived, which
affects the initial solubilized drug concentration and the drug particle
radius at the start of the
experiment) and km (see Table 1).
24
CA 02705083 2010-05-21
Table 1. Model parameters used for fitting Equations 3a and 3b to the
experimental data
CASat
Formulation DA (CM2/S) (.tg/m1) fdissolved km (cm/s)
Solubilized nimodipine 1.0 x10-5 40 1.00 N/A
380 nm nimodipine
particles 1.0 x10-5 40 0.85 2.5x10-5
900 nm nimodipine
particles 1.0 X10-5 40 0.62 2.3x10-5
100 gm nimodipine
particles 1.0 x10-5 40 0.06 7.5x10-6
The model fits were in excellent agreement (see Fig. 3) with all four
experimental release
profiles. Although there was some uncertainty in the initial particle size
and, to a lesser extent,
CAsat, the models support the interpretation of the biphasic release
mechanism. They revealed
consistent km-values on the order of 10 cm/s and fdissmved values that varied
from 6% for the 100
gm particles, to 62% for the 380 nm and 85% for the 900 nm particles, to 100%
for the
solubilized nimodipine. This suggests that a full range of faissolved-values
can be achieved by
varying the way in which the drug is introduced into HAMC.
Given the broad range of release profiles that can be achieved using HAMC, it
is useful to define
"limiting" analytical expressions for the fastest and slowest possible release
profiles. The fastest
possible release occurs when the entire amount of drug is dissolved, as in the
case of the
solubilized nimodipine, where the release profile can be estimated using
Equation 2. Conversely,
the release profile is slowest when all loaded drug starts out in the
particulate state (i.e., f
-dissolved =
0.00). A simple analytical expression for the release profile in this
situation can be obtained
under two sets of circumstances: (1) the release rate is controlled only by
diffusion through the
gel matrix, where drug particle dissolution is faster than the diffusion of
the drug out of the gel,
or (2) the release rate is controlled only by the dissolution of the drug
particle, where the
diffusion of the drug out of the gel is faster than the drug particle
dissolution. The time scales of
these two processes can be compared by defining a dimensionless number (4 )
that represents the
CA 02705083 2010-05-21
ratio between the characteristic times of drug diffusion out of the gel and
drug particle
dissolution:
kmnpRi2L2
= _________ (4)
DA
When 4 1, the release profile is governed exclusively by the diffusion of
the drug through the
aqueous gel matrix, and the release profile can be described by the Higuchi
shrinking core model
[WI Higuchi, Diffusional Models Useful in Biopharmaceutics - Drug Release Rate
Processes. J
Pharm Sci 56(3) (1967) 315-324]:
M,- DA 3Cl MWA (3Cl MWA 2 p0,5
ASa ASa
(5)
M.õ \ 27tRop An 47TR, Anp
Conversely, when 4 <<1, release is governed exclusively by slow dissolution of
the drug
particles. The release of the particulate nimodipine from HAMC 0(10-2-10-
3)) exemplifies
this situation, and enables the determination of a limiting release profile
equation through the
solution of Equation 3a. In this case, since the diffusion of the drug out of
the gel is rapid relative
to the particle dissolution rate, it is reasonable to assume that CA is
negligible relative to CAse.
This decouples Equation 3a from Equation 3b, and enables the analytical
solution for RI(t),
yielding:
S
MI =1 [1 _____________ al A t)3
(6)
M P A R1,0
which is the scaling predicted by the Hixson-Crowell model [A. W. Hixson, IH.
Crowell,
Dependence of reaction velocity upon surface and agitation I - Theoretical
consideration.
Industrial and Engineering Chemistry 23 (1931) 923-931]. Assuming the kin-
value that was fitted
in the case of the 100 gm particles, which is the closest condition to the
limit that was tested,
Equation 6 predicts a limiting release profile (dashed lines in Fig. 3) that
is similar to the
experimental profile obtained for the large nimodipine particles, but has a
starting point at the
origin.
As can be seen from the shaded region of Fig. 3, by varying the method by
which the nimodipine
is introduced into HAMC a broad range of release profiles can be achieved.
These are bounded
26
CA 02705083 2010-05-21
by the Fickian release obtained in the case of the fully-solubilized
nimodipine and the nearly
linear release that is predicted by Equation 6. Significantly, under each
condition described
above, the drug release rate increases with increasing CAsat, as indicated by
Eq. 3a, 3b, 5, and 6.
HAMC as a Delivery Platform for Hydrophobic Drugs
It has been shown that the incorporation of polymeric excipients as structural
elements of an
injectable hydrogel, such as MC in the case of HAMC, can increase the aqueous
solubility of
hydrophobic drugs. Unlike previous work where polymer additives reduced the
rate of drug
crystallization [Y.S.R. Krishnaiah, P. Bhaskar, V. Satyanarayana, Penetration-
enhancing effect of
ethanol-water solvent system and ethanolic solution of carvone on transdermal
permeability of
nimodipine from HPMC gel across rat abdominal skin. Pharmaceutical Development
and
Technology 9(1) (2004) 63-74; L. Zema, A. Maronii, A. Foppoli, L. Palugan, ME.
Sangalli, A.
Gazzaniga, Different HPMC viscosity grades as coating agents for an oral time
and/or site-
controlled delivery system: An investigation into the mechanisms governing
drug release. J
Pharm Sci 96(6) (2007) 1527-1536; S.L. Raghavan, A. Trividic, A.F. Davis, J.
Hadgraft, Effect of
cellulose polymers on supersaturation and in vitro membrane transport of
hydrocortisone
acetate. Int J Pharm 193(2) (2000) 231-237; K Yamashita, T Nakate, K Okimoto,
A. Ohike, 1'.
Tokunaga, R. Ibuki, K Higaki, T Kimura, Establishment of new preparation
method for solid
dispersion formulation of tacrolimus. hit J Pharm 267(1-2) (2003) 79-91; S.L.
Raghavan, K
Schuessel, A. Davis, J. Hadgraft, Formation and stabilisation of triclosan
colloidal suspensions
using supersaturated systems. Int J Pharm 261(1-2) (2003) 153-158; U
Kumprakob,
Kawakami, 1. Adachi, Permeation enhancement of ketoprofen using a
supersaturated system with
antinucleant polymers. Biological & Pharmaceutical Bulletin 28(9) (2005) 1684-
1688; XG. Ma,
J. Taw, C.M Chiang, Control of drug crystallization in transdermal matrix
system. Int J Pharm
142(1) (1996) 115-119; P.N. Kotiyan, P.R. Vavia, Eudragits: Role as
crystallization inhibitors in
drug-in-adhesive transdermal systems of estradiol. Eur J Pharm Biopharm 52(2)
(2001) 173-
1801, the presence of 7 wt% MC in aqueous solution gives rise to a tenfold
amplification in
nimodipine solubility. This increase in drug solubility significantly
accelerates drug release from
the hydrogel, and suggests that injectable hydrogel matrices can act as
polymeric excipients that
accelerate the delivery of hydrophobic, poorly soluble drugs.
27
CA 02705083 2010-05-21
The differences in the size-dependent particle dissolution kinetics are
reflected in the nimodipine
release profiles from HAMC, and can be exploited in tailoring drug release
rates. In the case
where the drug is completely solubilized at the beginning of the release
process, its release is
rapid and governed by Fickian diffusion [I. Tosum, Modeling in Transport
Phenomena, a
Conceptual Approach, 2nd ed., Elsevier, 2007]. In the case of particulate
nimodipine, however,
release occurs over longer time scales due to the slow dissolution of the drug
particles. The
release profiles obtained from these particulate formulations are biphasic and
dependent on the
size of the drug particles introduced into HAMC. Experimental and model
analysis of the drug
dissolution and release reveals that the biphasic release profiles reflect a
rapid release of
solubilized drug, followed by a slow dissolution-controlled release of the
solid nimodipine.
Because the amount of nimodipine that is solubilized at the beginning of the
release process
varies with the initial drug particle size, the release profiles depend
strongly on the size of the
drug particles that were used in its preparation. This suggests that HAMC and
its homologues
can yield highly tunable release profiles that are dependent on how the drug
is introduced into the
in situ gelling scaffold. Similarly, because the dissolution of the drug
particles in MC occurs over
the course of several days, these variations indicate that these release
profiles can also be adjusted
by modulating the time between the preparation and application.
Additionally, the general model developed for predicting the release of
nimodipine from HAMC
is transferrable to other similar systems where sparingly soluble drugs are
released from hydrogel
scaffolds. The model allows both a better understanding of the mechanism that
controls the
release of drugs from these systems, as well as to predict drug release
behaviour in future studies.
Likewise, the use of MC and its homologues in other products, such as foods
and personal care
formulations, suggests that hydrogels such as HAMC can be used to achieve
highly tunable,
accelerated delivery of other types of active ingredients, such as hydrophobic
neutraceuticals
[K.P. Velikov, E. Pelan, Colloidal delivery systems for micronutrients and
nutraceuticals. Soft
Matter 4(10) (2008) 1964-1980], the use of which are limited by their low
solubilities.
This description demonstrates that the incorporation of polymeric excipients
in an injectable
hydrogel can accelerate the release of hydrophobic drugs, and that the
addition of MC to water
increases the aqueous solubility of sparingly soluble nimodipine. The effect
of varying initial
28
CA 02705083 2016-09-07
particle sizes on the particle dissolution rates to obtain a broad range of
release profiles has
been demonstrated. These release profiles depend on the method by which the
nimodipine is
introduced into HAMC, namely the fraction of pre-dissolved drug. Model
analysis of these
release profiles supports the release mechanism described above, and indicates
that an
injectable hydrogel bearing MC and its homologues can provide a versatile
platform for rapid
and controlled release of hydrophobic drugs and other sparingly-soluble
compounds.
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 herein before set forth, and as follows in
the scope of the
appended claims.
29
