U.S. patent application number 17/422106 was filed with the patent office on 2022-04-07 for composite viscoelastic hydrogel, and uses thereof for sealing a channel in tissue.
The applicant listed for this patent is The Provost, Fellows, Scholars And Other Mambers Of Board Of Trinity College Dublin. Invention is credited to Colm McGarvey, Garrett Ryan.
Application Number | 20220105232 17/422106 |
Document ID | / |
Family ID | 1000006094182 |
Filed Date | 2022-04-07 |
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United States Patent
Application |
20220105232 |
Kind Code |
A1 |
Ryan; Garrett ; et
al. |
April 7, 2022 |
COMPOSITE VISCOELASTIC HYDROGEL, AND USES THEREOF FOR SEALING A
CHANNEL IN TISSUE
Abstract
A composite viscoelastic hydrogel comprises a continuous phase
of non-crosslinked hyaluronic acid gel and a dispersed phase of
dehydrothermally-crosslinked micron-sized gelatin hydrogel
particles. The hydrogel exhibits a storage modulus (G') of greater
than 400 Pa and a tan .delta. (G''/G') from 0.1 to 0.8 in dynamic
viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate
at 25.degree. C. The dehydrothermally crosslinked micron-sized
gelatin hydrogel particles have an average dimension of less than
100 microns prior to hydration.
Inventors: |
Ryan; Garrett; (Dublin,
IE) ; McGarvey; Colm; (Ringsend, IE) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
The Provost, Fellows, Scholars And Other Mambers Of Board Of
Trinity College Dublin |
Dublin |
|
IE |
|
|
Family ID: |
1000006094182 |
Appl. No.: |
17/422106 |
Filed: |
January 10, 2020 |
PCT Filed: |
January 10, 2020 |
PCT NO: |
PCT/EP2020/050611 |
371 Date: |
July 9, 2021 |
Current U.S.
Class: |
1/1 |
Current CPC
Class: |
C08L 5/08 20130101; A61L
24/0094 20130101; A61L 24/0031 20130101; A61L 24/0042 20130101;
A61L 2430/36 20130101; C08L 89/06 20130101; A61L 2400/06
20130101 |
International
Class: |
A61L 24/00 20060101
A61L024/00; C08L 5/08 20060101 C08L005/08; C08L 89/06 20060101
C08L089/06 |
Foreign Application Data
Date |
Code |
Application Number |
Jan 10, 2019 |
EP |
PCT/EP2019/050597 |
Claims
1. A composite viscoelastic hydrogel comprising: a continuous phase
of non-crosslinked biodegradable hyaluronic acid gel; and a
dispersed phase of dehydrothermally-crosslinked micron-sized
gelatin hydrogel particles.
2. A composite viscoelastic hydrogel according to claim 1 that
exhibits a storage modulus (G') of greater than 400 Pa and a tan
.delta. (G''/G') from 0.1 to 0.8 in dynamic viscoelasticity
measured by a rheometer at 1 Hz and 1% strain rate at 25.degree.
C.
3. A composite viscoelastic hydrogel according to claim 1 that
exhibits an axial compressive stiffness of greater than 800 Pa, as
measured using an axial compression testing machine.
4. A composite viscoelastic hydrogel according to claim 1, in which
the dehydrothermally crosslinked micron-sized gelatin hydrogel
particles have an average dimension of less than 100 microns prior
to hydration.
5. A composite viscoelastic hydrogel according to claim 1, that
possesses an in vivo degradation period in the lung tissue of at
least 2 weeks.
6. A composite viscoelastic hydrogel according to claim 1, in which
the dehydrothermally-crosslinked micron-sized gelatin hydrogel
particles have an in-vivo degradation period in the lung tissue of
less than 2 months.
7. A composite viscoelastic hydrogel according to claim 1 which is
shear-thinning.
8. A composite viscoelastic hydrogel according to claim 1, that is
shear-thinning and demonstrates a storage modulus G' of less than
200 Pa at 100% strain as measured by a rheometer under strain
control and at a test frequency of 1 Hz.
9. A composite viscoelastic hydrogel according to claim 1,
comprising 8-25% gelatin (w/v).
10. A composite viscoelastic hydrogel according to claim 1,
comprising about 0.4-6% of hyaluronic acid (w/v).
11. A composite viscoelastic hydrogel according to claim 1, whereby
the hydrogel is terminally sterilized via steam sterilization.
12. A composite viscoelastic shear-thinning hydrogel according to
claim 1, comprising: a continuous phase of non-crosslinked
biodegradable hyaluronic gel; and a dispersed phase of
dehydrothermally-crosslinked micron-sized gelatin hydrogel
particles, in which the hydrogel exhibits a storage modulus (G') of
greater than 400 Pa and a tan .delta. (G''/G') from 0.01 to 0.8 in
dynamic viscoelasticity measured by a rheometer at 1 Hz and 1%
strain rate at 25.degree. C.
13. A composite viscoelastic shear-thinning hydrogel according to
claim 1, comprising: a continuous phase of non-crosslinked
biodegradable hyaluronic gel; and a dispersed phase of
dehydrothermally-crosslinked micron-sized gelatin hydrogel
particles, in which the hydrogel comprises 8-25% gelatin hydrogel
particles (w/v) having an average dimension of less than 100
microns in a dehydrated state, and about 0.5-2.0% of hyaluronic
acid (w/v).
14. A composite viscoelastic shear-thinning hydrogel according to
claim 1, comprising: a continuous phase of non-crosslinked
biodegradable hyaluronic gel; and a dispersed phase of
dehydrothermally-crosslinked micron-sized gelatin or collagen
hydrogel particles, in which the hydrogel exhibits an in vivo
degradation period in the lung tissue of at least 2 weeks, and in
which dehydrothermally-crosslinked micron-sized gelatin hydrogel
particles exhibit an in-vivo degradation period in the lung tissue
of less than 2 months.
15. (canceled)
16. (canceled)
17. (canceled)
18. (canceled)
Description
FIELD OF THE INVENTION
[0001] The present invention relates to a composite viscoelastic
hydrogel, and uses thereof for sealing a channel in tissue. The
invention also relates to methods, kits and systems employing a
composite viscoelastic hydrogel, including methods, kits and
systems for sealing a channel in tissue.
BACKGROUND TO THE INVENTION
[0002] A number of surgical procedures require puncturing an
instrument through the body to gain access to a target treatment
region, such as puncturing the thoracic wall to gain access to the
thoracic cavity. The most common example is transthoracic needle
lung biopsy where a special needle is used to obtain a sample of
tissue from a suspected cancerous tissue mass. This procedure,
which is presented schematically in FIGS. 1A-1D (Prior art), is
typically carried out by an interventional radiologist using CT
(computed tomography) guidance. When the biopsy needle punctures
the outer surface of the lung air can escape between the lung and
the thoracic wall into a space known as the pleural cavity. The air
gradually pushes the lung away from the thoracic wall causing the
lung to collapse, a complication known as pneumothorax. If the
pneumothorax is large, it can lead to severe pain and distress for
the patient. An unresolved pneumothorax can lead to the patient
being admitted to hospital for treatment and monitoring and often
requires the surgical insertion of a chest drain to withdraw the
air in the pleural cavity. Pneumothorax can result in considerable
pain and morbidity to the patient, increased anxiety and stress to
the attending clinician, and unnecessary and substantial costs to
the hospital. Approximately 33% of patients undergoing a
transthoracic lung biopsy procedure will develop a pneumothorax and
approximately 1 in 3 of these patients will require a chest
drain.
[0003] Methods to prevent pneumothorax are of great interest
because of the concomitant morbidity and hospital expenditures.
Numerous attempts have been described in scientific literature and
have focused on plugging the biopsy needle tract with an adhesive
or plug as the biopsy needle is being withdrawn. A number of
different substances have been injected with this purpose including
gelatine sponge slurry, fibrin adhesive, autologous blood,
supernatant serum and autologous blood mixture, and collagen foam.
These efforts have proven ineffective and have not been widely
adopted. Their lack of efficacy may be as a result of the physical
properties of the substances injected and the lack of control over
their injected location. Additional references which may be
suitable for lung sealing are outlined in U.S. Pat. No. 6,592,608B2
and U.S. Pat. No. 6,790,185B1. This technology is commercially
available as the Biosentry.TM. device from Surgical Specialties
Corp (MA, USA www.biosentrysystem.com). Other publications relevant
to lung and tissue sealing include US2016120528A, US2006025815A,
US2013338636A, US2006009801A, U.S. Pat. No. 6,770,070B,
US2017232138A, US2002032463A, and US2009136589A.
SUMMARY OF THE INVENTION
[0004] The Applicant has discovered that compositions for
generating a sealing plug in tissue should be shear-thinning (which
allows the viscosity of the composition to drop under shear during
injection, and then increase after delivery to form the sealing
plug), should be sufficiently robust to prevent the composition
separating during injection (i.e. prevent the water being squeezed
out of the composition), and requires rheological properties that
allow the composition to self-heal after needle delivery, and to
exhibit tissue apposing properties (i.e. push tissue away). In
addition to these properties, the composition should be
biodegradable but also have sufficient in-vivo persistence time to
perform a sealing function for a sufficient period of time (for
example in the case on a biopsy needle tract, during the period of
healing of the needle tract). The Applicant has a discovered that a
composite viscoelastic hydrogel having a continuous
(non-crosslinked) hyaluronic acid hydrogel phase, and a dispersed
phase comprising crosslinked gelatin or collagen hydrogel
particles, is ideal for providing a sealing plug in tissue,
especially in lung parenchymal tissue. This hydrogel is
shear-thinning, tissue-apposing, and biodegradable over a period of
a few weeks which provides adequate time for tissue (i.e. a needle
tract to heal). In a particular embodiment, the use of a weak
crosslinking method for the collagen or gelatin particles (i.e.
dehydrothermal crosslinking) has been found to be ideal for
providing a hydrogel gel that persists in tissue during a period of
therapy and ultimately biodegrades. In addition, the amount of
collagen or gelatin particles in the hydrogel may be varied to vary
the tissue persistence time An essential benefit of providing a
viscous continuous phase (provided by the hyaluronic acid) is to
ensure that there is no phase separation of the dispersed phase
during injection. As an illustration, if gelatin hydrogel particles
were to be suspended in saline solution as a carrier in a syringe,
during injection through a narrow gauge needle the saline solution
would be expelled from the needle more easily than the suspended
gelatin particles. The gelatin particles would condense behind the
needle opening and the saline would pass through, concentration the
remaining hydrogel in the syringe. This would alter the uniformity
of hydrogel injection, which is a key attribute for several medical
applications such as the one described in this patent. The
composite nature of the gel has another additional benefit in vivo:
the colloidal gelatin provides an excellent framework for cellular
infiltration and vascularisation.
[0005] According to a first aspect of the present invention, there
is provided a composite viscoelastic hydrogel comprising a
continuous phase and a dispersed polymer phase.
[0006] In one embodiment, the dispersed phase is colloidal polymer,
typically polymer hydrogel particles. In one embodiment, the
dispersed phase is crosslinked, ideally dehydrothermally
crosslinked.
[0007] In any embodiment, the dispersed phase polymer is
crosslinked by dehydrothermal (DHT) crosslinking. This method
avoids adding crosslinking chemicals to the gel. In addition, this
method of crosslinking has been found to be ideal for achieving a
relatively weak crosslinking of the dispersed polymer particles,
required to make the particles biodegradable yet persist in the
tissue during a tissue healing period (i.e. a period of weeks).
Typically, the DHT crosslinking is performed under vacuum at a
temperature greater than 100.degree. C., typically 130.degree. C.
to 160.degree. C., for an extended period, for example 24 hours. It
will be clear to a person skilled in the art that an equivalent
level of weak crosslinking may be achieved by using a lower
temperature and longer time, or vice versa. With respect to
gelatin, during the DHT treatment, water is extracted from gelatin
by condensation reactions, such as esterification or amide link
formation, resulting in the formation of intermolecular crosslinks
that improve the materials resistance to dissolution. When placed
in aqueous suspensions, the DHT treated gelatin powders swell by
absorbing the surrounding water to form a colloidal hydrogel. The
amount of swelling is dependent on the level of crosslinking. This
feature greatly influences the final rheological properties of the
gel. An additional benefit of employing the DHT crosslinking
process in relation to gelatin is that it removes residual
endotoxins which may be present in the polymer from processing as
the polymer is derived from an animal source (collagen).
[0008] In one embodiment, the continuous phase is a hydrogel. In
one embodiment, the continuous phase polymer is a glycosaminoglycan
(or a salt thereof), for example hyaluronic acid.
[0009] In one embodiment, the composite viscoelastic hydrogel
comprises 2-25%, 10-20%, 5-15%, 12-18%, 8-12% colloidal
polymer.
[0010] In one embodiment, the colloidal polymer comprises or
consists of gelatin or collagen (or a mixture thereof).
[0011] In one embodiment, the composite viscoelastic hydrogel
comprises about 1-6%, 2-6%, 0.5-2.0%, or 0.6-1.2% continuous phase
polymer (i.e. HA).
[0012] In one embodiment, the continuous phase polymer comprises or
consists of HA (or another glycosaminoglycan).
[0013] In one embodiment, the continuous phase polymer (i.e. HA) is
not cross-linked, or is lightly cross-linked.
[0014] In one embodiment, the invention provides a composite
viscoelastic hydrogel comprising a continuous polymer phase
comprising 2-6% polymer, and a dispersed polymer phase comprising
2-25% colloidal polymer in the form of crosslinked polymer
microbeads In one embodiment, the crosslinked polymer particles
have an average dimension of less than 100 microns prior to
hydration.
[0015] In one embodiment, the continuous phase polymer comprises or
consists of HA and the colloidal polymer comprises gelatin or
collagen.
[0016] In one embodiment, the composite viscoelastic bi-phasic
hydrogel exhibits a storage modulus (G') of greater than 400 Pa and
a tan .delta. (G''/G') from 0.1 to 0.8 in dynamic viscoelasticity
measured by a rheometer at 1 Hz and 1% strain rate at 25.degree. C.
In one embodiment, the composite viscoelastic bi-phasic hydrogel
exhibits an axial compressive stiffness of greater than 800 Pa, as
measured using an axial compression testing machine.
[0017] In one embodiment, the dispersed phase comprises
dehydrothermally crosslinked micron-sized gelatin hydrogel
particles.
[0018] In one embodiment, the composite viscoelastic hydrogel is
formulated to exhibit an in vivo degradation period in the lung
tissue of at least 2 weeks.
[0019] In one embodiment, the dehydrothermally-crosslinked
micron-sized gelatin hydrogel particles are formulated to exhibit
an in-vivo degradation period in the lung tissue of less than 2
months.
[0020] In one embodiment, the dehydrothermally crosslinked
micron-sized gelatin or collagen hydrogel particles have an average
dimension of less than 100 microns prior to hydration (i.e. in a
dehydrated state).
[0021] In one embodiment, the hyaluronic acid has a molecular
weight of at least 1 MDa, for example 1-2 MDa.
[0022] In one embodiment, the DHT gelatin powder is derived from a
high Bloom strength gelatin, for example 200-400, preferably
250-350-, and ideally about 300 gBloom.
[0023] In one embodiment, the composite viscoelastic bi-phasic
hydrogel is shear-thinning. In one embodiment the composite
viscoelastic bi-phasic hydrogel demonstrates a storage modulus G'
of less than 200 Pa at 100% strain as measured by a rheometer under
strain control and at a test frequency of 1 Hz over a strain range
of 0.01-100%.
[0024] In one embodiment, the composite viscoelastic shear-thinning
bi-phasic hydrogel comprises: [0025] a continuous phase of
non-crosslinked biodegradable hyaluronic gel; and [0026] a
dispersed phase of dehydrothermally-crosslinked micron-sized
gelatin or collagen hydrogel particles, [0027] in which the
hydrogel exhibits a storage modulus (G') of greater than 400 Pa and
a tan .delta. (G''/G') from 0.1 to 0.8 in dynamic viscoelasticity
measured by a rheometer at 1 Hz and 1% strain rate at 25.degree.
C.
[0028] In one embodiment, the composite viscoelastic shear-thinning
bi-phasic hydrogel comprises: [0029] a continuous phase of
non-crosslinked biodegradable hyaluronic gel; and [0030] a
dispersed phase of dehydrothermally-crosslinked micron-sized
gelatin or collagen hydrogel particles, [0031] in which the
hydrogel comprises 8-25% gelatin or collagen hydrogel particles
having an average dimension of less than 300 microns in a
dehydrated state.
[0032] In one embodiment, the composite viscoelastic shear-thinning
bi-phasic hydrogel comprises: [0033] a continuous phase of
non-crosslinked biodegradable hyaluronic gel; and [0034] a
dispersed phase of dehydrothermally-crosslinked micron-sized
gelatin or collagen hydrogel particles, [0035] in which the
hydrogel comprises 8-25% gelatin or collagen hydrogel particles
having an average dimension of less than 100 microns prior to
hydration, and about 0.4-2.0% of hyaluronic acid.
[0036] In another aspect, the invention provides a system for
sealing a channel in tissue (for example a channel created during a
minimally invasive percutaneous procedure) comprising: [0037] a
medical device comprising a hydrogel delivery needle (4) with a tip
(5) (generally a piercing tip) and a hydrogel outlet (6), and
[0038] a composite viscoelastic hydrogel of the invention.
[0039] In one embodiment, the composite viscoelastic hydrogel
exhibits a storage modulus (G') of at least 400 Pa in dynamic
viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate
at 25.degree. C.
[0040] In one embodiment, the composite viscoelastic hydrogel
exhibits a tan .delta. (G''/G') from 0.05 to 0.8 in dynamic
viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate
at 25.degree. C.
[0041] In one embodiment, the composite viscoelastic hydrogel is
configured to exhibit an in-vivo residence time of at least 1, 2 or
3 weeks. This enables the gel to persist in tissue, while the
tissue needle tract in the tissue heals. Generally, one week is
sufficient, but at least two weeks in-vivo residence time is
preferred. Hydrogels formed from, or comprising, crosslinked
polymers help with in-vivo residence time. For example, by creating
a composite hydrogel containing 4-5% non-crosslinked hyaluronic
acid and crosslinked gelatin particles (crosslinked by
dehyrothermal treatment) an in-vivo residence time of at least two
weeks in a lung needle biopsy tract was achieved.
[0042] In one embodiment, the composite viscoelastic hydrogel is
configured to exhibit an in vivo residence time of less than 3
months, more preferably less than 2 months, and more preferably
less than 1 month. A fast degradation period for the is an
advantage in several applications, where a long lasting hydrogel
can be confused as a cancerous tumour on CT or MRI scans at
follow-up periods following procedures. In addition, a fast
degrading hydrogel avoids issues with chronic inflammation of the
wound site and allows the tissue to return to normal in line with
the natural wound healing process.
[0043] The composite viscoelastic hydrogel (hereafter "composite
viscoelastic hydrogel" or "viscoelastic hydrogel" or "biphasic
hydrogel" or "hydrogel" or "gel") is generally a tissue apposing
hydrogel of sufficient properties that limits its infiltration of
tissue so that it pushes the tissue away. In this way the hydrogel
can create its own discrete space inside a tissue or organ. To
achieve this the properties must be present on entering the target
injection site. Typically, the viscoelastic hydrogel exhibits a
storage modulus (G') of at least 400 Pa (e.g. 800-6000 Pa), and a
tan .delta. (G''/G') from 0.1 to 0.8 in dynamic viscoelasticity
measured by a rheometer at 1 Hz and 1% strain rate at 25.degree. C.
The hydrogel is configured for injection via a syringe.
[0044] For improved tissue opposing properties and to form a
uniform plug surrounding the needle, it is also preferable that the
viscoelastic hydrogel portrays an axial compressive stiffness of
equal to or greater than lung parenchymal tissue, as measured using
an axial compression testing machine, for example by using a Zwick
universal testing machine with a 5N load cell at a strain rate of 3
mm/min. The viscoelastic hydrogel should preferably have a
compressive modulus of greater than 200 Pa, preferably greater than
400 Pa, and more preferably greater than 800 Pa.
[0045] Optionally in any embodiment, the injectable composite
viscoelastic hydrogel is a shear thinning gel. For example, the
viscoelastic hydrogel may be configured to have a low viscosity
under higher shear stress or shear rates (i.e. during injection
through a needle), and a higher viscosity (under lower shear
stresses or shear rates) after removal of shear stress (i.e. once
delivered to a target location in the body. This enables these
materials to create a singular hydrogel plug at the site of
delivery. Materials which possess these properties are outlined in
the review articles `Shear-thinning hydrogels for biomedical
applications`, Soft Matter, (2012) 8, 260, `Injectable matrices and
scaffolds for drug delivery in tissue engineering` Adv Drug Deliv
Rev (2007) 59, 263-272, and `Recent development and biomedical
applications of self-healing hydrogels` Expert Opin Drug Deliv
(2017) 23: 1-15. Typically, the shear thinning viscoelastic
hydrogel exhibits a storage modulus (G') of less than 200 Pa,
preferably less than 100 Pa in dynamic viscoelasticity at a
frequency of 1 Hz and 100% strain.
[0046] Optionally, in any embodiment, the composite viscoelastic
hydrogel is self-healing. This refers to the hydrogel's ability to
spontaneously form new bonds between molecules when old bonds are
broken within the material.
[0047] Optionally in any embodiment, the colloidal hydrogel is
formed by hydrating biocompatible polymer particles which are
preferably insoluble in biological fluid. Optionally in any
embodiment, the degradation period of the polymer particles is
preferably less than 1 year, more preferably less than 6 months,
and more preferably less than 2 months. Optionally in any
embodiment, the colloidal hydrogel is comprised of a polymer of
biological origin, for example gelatin, collagen, fibrin or
hyaluronic acid. Optionally in any embodiment, the polymer is
crosslinked. Optionally in any embodiment, the colloidal hydrogel
comprises about 0.2-30%, 15-28%, or 20-27% hydrogel forming polymer
(w/v). Optionally in any embodiment, the colloidal hydrogel
exhibits a storage modulus (G') of greater than 400 Pa, more
preferably greater than 800 Pa, more preferably greater than 1000
Pa in dynamic viscoelasticity measured by a rheometer at 1 Hz and
1% strain rate at 25.degree. C.
[0048] Optionally in any embodiment, the continuous phase hydrogel
may be formed by a hyaluronan hydrogel, and may be present at a
concentration of 0.4-6%. Optionally in any embodiment the
hyaluronan hydrogel may be non-crosslinked or lightly
crosslinked.
[0049] Optionally in any embodiment, the colloidal hydrogel may be
present at concentrations of 0.2 to 30%, 8 to 25%, 8 to 15%, 8 to
12%, or about 10% hydrogel forming polymer (w/v).
[0050] Optionally in any embodiment, the colloidal hydrogel is
formed from hydrated polymer particles of <100 .mu.m in average
particle size (for example 5-99, 20-80, or 30-80 microns.
[0051] Optionally in any embodiment the colloidal hydrogel is
insoluble in aqueous solution.
[0052] Optionally in any embodiment the colloidal hydrogel is
formed from crosslinked polymer particles. Optionally in any
embodiment, the colloidal hydrogel is a gelatin hydrogel comprising
dehydrothermally (DHT) crosslinked gelatin powders having an
average particle size (D.sub.50) of about 10-100, 20-50 or 30-40
microns. Optionally in any embodiment, the biphasic hydrogel
exhibits a storage modulus (G') of greater than 400 Pa, more
preferably greater than 800 Pa, more preferably greater than 1000
Pa, and a tan .delta. (G''/G') from 0.1 to 0.6 in dynamic
viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate
at 25.degree. C.
[0053] Optionally in any embodiment, the biphasic hydrogel portrays
an axial compressive stiffness of equal to or greater than lung
parenchymal tissue, as measured using an axial compression testing
machine Optionally in any embodiment, the composite viscoelastic
hydrogel is de-aerated which means it has been removed of air
and/or gas or in other words de-gassed.
[0054] Optionally in any embodiment, the composite viscoelastic
hydrogel comprises a therapeutic agent.
[0055] Optionally in any embodiment, the composite viscoelastic
hydrogel is biodegradable.
[0056] Optionally in any embodiment, the composite viscoelastic
hydrogel is comprised of 0.5-6%, preferably 3-5% high molecular
weight hyaluronan (w/v). Optionally in any embodiment, the
hyaluronan hydrogel may be combined with 0.2 to 30% colloidal
hydrogel to form a biphasic hydrogel. Optionally in any embodiment,
the colloidal hydrogel may be comprised of hydrogel forming polymer
particles. Optionally in any embodiment, the hydrogel forming
polymer particles are gelatin particles, collagen particles or
hyaluronan particles.
[0057] Optionally in any embodiment, the composite viscoelastic
hydrogel described herein may be provided in separate components,
for example in multiple syringes and the means can be provided to
allow mixing of the components prior to injection through the
syringe.
[0058] Optionally in any embodiment, the system and methods
described herein include an initial step of providing the
viscoelastic hydrogel as a dehydrated or semi-dehydrated powder,
and reconstitution of the powder in a suitable fluid to form the
viscoelastic hydrogel.
[0059] Optionally in any embodiment, the composite viscoelastic
hydrogel is a microporous hydrogel which can be described as
hydrogels with interconnected pores that can mechanically collapse
and recover reversibly. When the hydrogel is delivered via
injection with a needle and syringe, water is squeezed out from the
pores, which causes the hydrogel to collapse, allowing it to pass
through the needle. Once the hydrogel has left the needle and the
mechanical constraint imposed by the needle walls is removed, the
hydrogel can recover its original shape almost immediately in the
body. These hydrogels generally behave like a foam and can be
reversibly compressed at up to 90% strain without any permanent
damage to the network.
[0060] Optionally in any embodiment, the composite viscoelastic
hydrogel is provided in a syringe configured for fluidic connection
to a proximal end of the hydrogel delivery needle.
[0061] Optionally in any embodiment, the syringe comprises 200
.mu.L to 5000 .mu.L of viscoelastic hydrogel, 200 .mu.L to 2000
.mu.L of viscoelastic hydrogel, or 200 .mu.L to 1000 .mu.L of
viscoelastic hydrogel.
[0062] Optionally in any embodiment, the hydrogel delivery needle
diameter can range from 10-24 gauge, preferably from 16-20 gauge.
This is the typical needle size range for lung diagnostic
procedures. Larger delivery needles (10-16 gauge) may be employed
for other procedures including therapeutic procedures such as lung,
live and kidney ablation. Smaller needles greater than 20 gauge or
larger than 10 gauge may be used for other medical procedures.
[0063] Optionally in any embodiment, the hydrogel outlet is spaced
proximal to the piercing tip of the needle. The position of the
hydrogel outlet on a side of the needle enables formation of a
closed annular sealing plug around the needle, and the viscoelastic
properties of the hydrogel allow the annular sealing plug to
re-shape upon removal of the device whereby the hole in the middle
of the sealing plug is filled in. Optionally in any embodiment, the
hydrogel outlet is spaced from preferably 1 to 15 mm or more
preferably 3-8 mm, from a piercing tip of the needle.
[0064] Optionally in any embodiment, the hydrogel delivery needle
comprises a plurality of hydrogel outlets disposed on a side of the
needle. The hydrogel outlets may be disposed in a radial fashion
around the circumference of the needle. The hydrogel outlets may be
circular in profile, in which case their size can range from
0.3-1.5 mm in diameter depending on the diameter of the hydrogel
delivery needle. The hydrogel outlets may also take non-circular
and elongated profiles.
[0065] Optionally in any embodiment, the hydrogel outlet consists
of a radiolucent region on the delivery needle where sufficient
material has been removed through cutting or erosion process to
provide a contrast in radiopacity between the delivery needle and
the hydrogel outlet.
[0066] Optionally in any embodiment, the coaxial cannula consists
of an aperture proximal to its distal tip. This aperture may form a
radiolucent region on the coaxial cannula by removing sufficient
material about the circumference of the cannula.
[0067] Optionally in any embodiment, radiolucent regions of both
the delivery needle and coaxial cannula are aligned when the
delivery needle and cannula are engaged. This will provide a
marking function about this radiolucent region during radiographic
guidance and allows the viscoelastic hydrogel to be injected at
this location.
[0068] Optionally in any embodiment, the hydrogel outlet and
coaxial cannula aperture may be created using a laser cut profile
or pattern which removes a portion of material from the delivery
needle wall to create a pathway through which the hydrogel material
can flow to the intended target. Removal of a significant amount of
material will provide radiolucency to this portion of the device
and will provide visual feedback on the position of the hydrogel
outlet under CT guidance or other imaging modality. The
radiolucency (less radiopaque) is achieved by removal of a
significant amount of material from the needle walls using the
laser cut pattern without affecting the structural integrity of the
needle. Laser cut profiles comprising circumferential triangles and
similar structures to those employed in coronary stents can be
employed to maintain structural stability. Alternative material
eroding technology may also be employed to create the cut
pattern.
[0069] Optionally in any embodiment, the medical device comprises
an adjustable positioning mechanism configured to limit the
advancement depth of the hydrogel delivery needle through the
coaxial cannula as indicated by a measurement scale forming part of
the medical device, and typically forming part of the positioning
mechanism.
[0070] Optionally in any embodiment, the positioning mechanism
comprises a fixed housing attached to the hydrogel delivery needle,
a movable hub mounted to the needle for axial movement along the
hydrogel delivery needle relative to the fixed housing and having a
distal-most face configured to abut a proximal face of the coaxial
cannula luer lock.
[0071] Optionally in any embodiment, a visible mark is provided on
the delivery needle proximally to the piercing tip.
[0072] Optionally in any embodiment, the system further comprises a
core needle with penetrating distal tip configured for insertion
through the inner lumen of the coaxial cannula and attachment to
the coaxial cannula luer lock.
[0073] Optionally in any embodiment, the system further comprises a
syringe configured for fluidic connection to the hydrogel delivery
needle, and in which the viscoelastic hydrogel is provided in the
syringe.
[0074] In another aspect, the invention provides a method of
performing a lung needle biopsy, comprising the steps of:
[0075] delivering a composite viscoelastic hydrogel of the
invention to a target location in the lung of a patient adjacent
the visceral pleura of the lung;
[0076] advancing the coaxial cannula distally over the hydrogel
injection needle and through the sealing plug;
[0077] removal of the hydrogel delivery needle through the
cannula;
[0078] advancing a biopsy needle through the cannula to a biopsy
site within the lung; actuating the biopsy needle to take a sample
of lung tissue at the biopsy site;
[0079] withdrawing the biopsy needle through the cannula; and
[0080] withdrawing the cannula whereby the sealing plug seals the
visceral pleura.
[0081] Optionally in any embodiment, after the removal of the
hydrogel delivery needle and prior to advancement of the biopsy
needle, the method includes the steps of insertion of a core needle
into the coaxial cannula, advancement of the core needle and
coaxial cannula to the biopsy site within the lung, and removal of
the core needle.
[0082] Optionally in any embodiment, prior to removal of the
hydrogel delivery needle, the method includes the steps of
advancing the hydrogel delivery needle to the biopsy site within
the lung, and then advancing the coaxial cannula over the hydrogel
delivery needle to the biopsy site within the lung.
[0083] Optionally in any embodiment, the step of advancing the
coaxial cannula distally over the hydrogel injection needle to the
biopsy site in the lung is guided by the cannula depth guide.
[0084] Optionally in any aspect, the invention provides a method of
performing a lung needle biopsy procedure comprising the steps
of:
[0085] injecting a viscoelastic hydrogel (for example, a composite
viscoelastic hydrogel of the invention) through a hydrogel delivery
needle into the lung adjacent the visceral pleura of the lung to
form a sealing plug that embraces the needle and abuts the visceral
pleura; advancing a coaxial cannula along the hydrogel delivery
needle and through the closed annular sealing plug;
[0086] removal of the hydrogel delivery needle through the
cannula;
[0087] advancing a biopsy needle through the cannula to a target
location within the lung;
[0088] actuating the biopsy needle to take a sample of lung tissue
at the target location;
[0089] withdrawing the biopsy needle through the cannula; and
[0090] withdrawing the cannula whereby the sealing plug seals the
visceral pleura preventing pneumothorax.
[0091] In another aspect, the invention provides a method of
performing a lung nodule localisation procedure comprising the
steps of:
[0092] injecting a composite viscoelastic hydrogel of the invention
through a hydrogel delivery needle into the lung adjacent the
visceral pleura of the lung to form a sealing plug that embraces
the needle and abuts the visceral pleura;
[0093] advancing a coaxial cannula along the hydrogel delivery
needle and through the closed annular sealing plug;
[0094] removal of the hydrogel delivery needle through the
cannula;
[0095] advancing a tissue stain delivery needle through the cannula
to a target location within the lung;
[0096] actuating the tissue stain needle to take a sample of lung
tissue at the target location; withdrawing the tissue stain needle
through the cannula; and
[0097] withdrawing the cannula whereby the sealing plug seals the
visceral pleura preventing pneumothorax.
[0098] In another aspect, the invention provides a method
comprising delivery of a composite viscoelastic hydrogel of the
invention into tissue of a patient, especially a lung of a patient
adjacent the visceral pleura of the lung to form a sealing plug
wholly within the lung that abuts the visceral pleura.
[0099] Optionally in any embodiment, the methods of the invention
involve delivering a volume of 100 to 3000 .mu.l of hydrogel.
Optionally in any embodiment, the methods involve delivering a
volume of 100 to 1000 .mu.l of hydrogel. Optionally in any
embodiment, the methods involve delivering a volume of 200 to 900
.mu.l of hydrogel. Optionally in any embodiment, the methods
involve delivering a volume of 200 to 500 .mu.l of hydrogel.
[0100] Optionally in any embodiment, the composite viscoelastic
hydrogel is delivered into the lung through a needle having a
piercing tip and a hydrogel outlet disposed on a side of the needle
spaced apart from piercing tip.
[0101] In another aspect, the invention provides a composite
viscoelastic hydrogel of the invention for use in forming a sealing
plug in a lung of a patient to prevent pneumothorax during a lung
needle biopsy procedure, in which the sealing plug is typically
delivered to the lung adjacent and abutting a visceral pleura.
[0102] Optionally in any embodiment, the biopsy needle is passed
through the sealing plug during the needle biopsy procedure.
[0103] Optionally in any embodiment, a coaxial cannula is passed
through the sealing plug, and the biopsy needle is passed through
the sealing plug via the coaxial needle.
[0104] Optionally in any embodiment, the target location in the
lung is located 0.2 to 6.0 mm distal of the visceral pleura.
[0105] Optionally in any embodiment the target location for
delivery of the hydrogel material is into the pleural cavity. In
this instance the hydrogel outlet will reside inside or across the
pleural cavity.
[0106] Optionally in any embodiment, the hydrogel delivery needle
may have a hydrogel outlet at the tip of the needle as opposed to
the side. It is also possible to have both a hydrogel outlet at the
tip of the needle and/or on the side of the needle. The delivery
device and system described herein may also provide an effective
solution to prevent bleeding during procedures requiring minimally
invasive percutaneous access to other organs such as the liver and
kidney. These procedures may include diagnosis or treatment of part
or all of these organs.
[0107] Optionally in any embodiment, the system and/or composite
viscoelastic hydrogel described herein can be used to separate
tissue during a surgical procedure. This may be required to create
a pathway through tissue for an instrument or to protect tissue
from unwanted stimuli which as tumour ablation or radiotherapy. For
this purpose a greater volume of viscoelastic hydrogel may be
delivered, for example 1-25 ml.
[0108] Optionally in any embodiment, the system and/or the
composite viscoelastic hydrogel described herein can be used as to
fill voids in tissue or organs.
[0109] Optionally in any embodiment, the system and/or the
composite viscoelastic hydrogel described herein can be employed in
the prevention of adhesion between adjacent tissues and organs.
[0110] Optionally in any embodiment, the system and/or composite
viscoelastic hydrogel described herein can be employed as a drug
delivery vehicle. The composite viscoelastic hydrogel may be loaded
with a drug or any other substance having physiological activity
which will slowly diffuse from the hydrogel after its implantation
into the body and the diffusion rate can be conveniently controlled
by changing the compositional parameters of the hydrogel.
[0111] Optionally in any embodiment, the system and/or composite
viscoelastic hydrogel described herein can be used as an embolic
agent for occlusion of an artery or vein. The composite
viscoelastic hydrogel can be deployed into an artery or vein to
occlude the flow of blood, either on a temporary or permanent
basis. In this manner, the hydrogel can be used to treat venous
diseases, for example aneurysm, varicose veins, insufficient veins,
dilated veins and ectasias. Thus, in one embodiment, the invention
provides a method of occluding a lumen (for example a section of
vasculature) in a subject comprising a step of delivering a
composite viscoelastic hydrogel according to the invention into the
lumen to occlude the lumen.
[0112] Optionally in any embodiment, the system and/or composite
viscoelastic hydrogel described herein can be employed as a tissue
bulking agent or tissue augmenting agent.
[0113] Optionally in any embodiment, the devices and components
described herein may be created using biocompatible materials
including polymers, metals and ceramics. Polymers can include
Polyether ether ketone, Polyethylene terephthalate, Nylon,
polyimides, polyurethanes, polyesters, Pebax.RTM. and copolymers
thereof. Metals may include stainless steel, nitinol, titanium and
cobalt chrome. The needles and cannula may also comprise fully or
partially flexible laser cut sections and braided sections to
provide flexibility. The needles and cannula may also be both
elongated and flexible such as in catheter type assemblies.
[0114] In a preferred embodiment, the compositions of the system,
or the system as a whole can be provided sterile for clinical use.
The hydrogel filled syringe can be prepared through an aseptic
formulation, mixing, filling and packaging process. The hydrogel
filling syringe may also be terminally sterilized through a heat or
steam sterilization process for e.g., autoclaving. Sterilization of
the system can also be performed via sterilization processes known
in the field including sterilization by ethylene oxide, hydrogen
peroxide, gamma ray and electronic beam.
[0115] Optionally in any embodiment, the components of the system
can be provided in packaging suitable for sterilization including,
but not limited to, a pouch, a blister pack, a bag, a procedure
set, a tub, a clamshell, a skin pack, a tray (including lid), a
carton, a needle sheath. The components of the system can all be
assembled as a single packaged device. Alternatively, multiple
packages containing the different components of the system can be
prepared and sterilized separately. The components of the system
can include but are not limited to the coaxial cannula with core
needle, the hydrogel delivery needle, the cannula depth lock,
locking arm, one or more syringes filled with viscoelastic
hydrogel, empty syringes, hypodermic needles, scalpels, skin
markers, radiopaque guides, scissors, biopsy needles, surgical
drapes, antiseptic solution, swabs, swab holders, sponges, saline
solution and histology tissue containers.
[0116] Optionally in any embodiment, the methods described herein
include an initial step of flushing the syringe with gel (or saline
or water) prior to insertion of the needle into the body. The
syringe may also be flushed with the hydrogel prior to insertion
into the body.
[0117] Optionally in any embodiment, the piercing tip of the
delivery needle is designed to prevent bleeding on insertion into
the lung, for example it may have a non-cutting atraumatic needle
tip profile, for example a pencil tip style needle or similar will
help prevent bleeding.
[0118] Optionally in any embodiment, the piercing tip is designed
with a sharpened bevel profile to minimise disruption of the
parietal and visceral pleural layers as the needle is being advance
through to the lung.
[0119] Optionally in any embodiment, the tip of the delivery needle
may be blunt. Optionally in any embodiment the hydrogel outlet may
be positioned distal to the blunt tip. Optionally in any embodiment
the tip of the delivery needle may be configured with a veress
needle tip that combines a spring activated blunt core and a sharp
piercing tip.
[0120] Optionally in any embodiment the delivery needle is a single
lumen. Optionally in any embodiment the delivery needle is
comprised of a multi-lumen tube. The multi-lumen tube may be a
single tube, or may be comprised of multiple individual tubes
within another lumen (for example a stainless steel needle). The
tubes may be connected to different delivery outlets. For example,
one tube may be connected to a delivery outlet that is distal to
the needle tip, whereas the other lumen may be connected directly
to the needle tip. Individual delivery lumens may be used to
deliver the hydrogel, deliver instruments, take measurements
(pressure, temperature, impedance), extract tissue (for example FNA
or core biopsies). The tubes may also be used to delivery
crosslinking agents, chemotherapy agents and cellular solution (for
example stem-cells).
[0121] Optionally in any embodiment the delivery needle may be
comprised of a single tube.
[0122] Optionally the single tube may comprise a tissue penetrating
tip. Optionally the delivery needle may be comprised of two or more
tubes bonded together, whereby the distal tube may form a tissue
penetrating tip. The various tubes used to comprise the delivery
needle can be made from radiodensity contrasting materials, for
example stainless steel or polymer.
[0123] Optionally in any embodiment, the delivery needle can be
provided with a central lumen to allow it to pass over a guidewire.
The guidewire can be provided for access to body cavities or
lumens.
[0124] Optionally in any embodiment the delivery needle and coaxial
cannula can be given atraumatic and friction prevention properties
by use of surface coatings and surface modifications such as
polytetrafluorinated ethylene and silicone-based coatings.
Optionally in any embodiment, the coaxial cannula can be provided
with a bevel cut profile, fillet cut or chamfer cut on its
distal-most tip to ease the force of insertion through the bodies
tissues.
[0125] Optionally in any embodiment, the hydrogel delivery needle
and coaxial cannula can be provided with external graduation marks
on their exterior surfaces to monitor the depth of insertion into
tissue and also to determine the position of the coaxial cannula in
relation to the delivery needle. These depth graduations can be
created using laser marking or ink pad printing or similar. Spacing
of 5-10 mm between graduation marks are typical.
[0126] Optionally in any embodiment, the methods described herein
include an aspiration step to ensure no major blood vessel is
punctured. This aspiration step may be conducted when the delivery
needle is inserted into the target location and before the hydrogel
plug is injected. This may be desirable so as to limit or prevent
any hydrogel from entering into the vasculature which may result in
a pulmonary embolism. Aspiration of dark blood would be an
indication that a major blood vessel has been punctured.
[0127] Optionally in any embodiment, the hydrogel filled syringe
employed can be configured to require aspiration before injection
of the hydrogel material. To achieve this, a mechanism can be built
into the syringe to restrict the forward actuation of the syringe
plunger until a retracting aspiration actuation has been
performed.
[0128] Optionally in any embodiment the system describe herein may
include an additional empty syringe for the purpose of performing
the aspiration step.
[0129] Optionally in any embodiment the device may contain a 2- or
3-way medical stopcock fluidically attached to the delivery device.
Any or both of the hydrogel filled syringe and the aspiration
syringe may be attached to the delivery device via the medical
stopcock which can be actuated to change and restrict the fluid
delivery path between aspiration syringe and hydrogel filled
syringe. This may provide the advantage of allowing a faster
aspiration and injection step and reduce the time spend in the lung
prior to injection of the hydrogel plug.
[0130] Optionally in any embodiment, the syringe is an ergonomic
syringe for improved deliverability. Examples are described in
US20090093787 A1 `Ergonomic Syringe` and U.S. Pat. No. 6,616,634 B2
`Ergonomic Syringe`. The system may also include an ergonomic
syringe adapter which can be mounted onto the syringe. An example
is described in USD675317 S1 `Ergonomic syringe adapter`. The
syringe may include a mechanism to inject the viscoelastic hydrogel
under high pressure. This may be in the form of a syringe assist
device Optionally in any embodiment, the coaxial needle may have an
internal sealing/valve feature that prevents any gel from entering
the coaxial needle.
[0131] Optionally in any embodiment, the hydrogel delivery needle
can be employed as a core needle within the coaxial needle.
[0132] Optionally in any embodiment, the positioning mechanism also
comprises a firing mechanism, for example a spring-loaded firing
mechanism, to quickly advance the delivery needle through the
coaxial cannula to a predetermined depth. The required distance can
either be a set distance for penetration depth, or can be
adjustable to take into account the coaxial cannula position in
relation to the target injection site. The device can be positioned
using measurements taken through imaging.
[0133] The system, device and methods of the invention may employ a
coaxial needle with a core that has a radiolucent marker for more
accurate determination of position.
[0134] Optionally in any embodiment the delivery device can be
provided in an elongated and flexible configuration so that it can
be passed through an endoscope to perform injections at
predetermined injection depths via an endoscope. The elongated
members can include both the coaxial cannula and delivery needle
elements of the delivery device.
[0135] Optionally in any embodiment the delivery device can be
provided with one or multiple energy delivery elements that can
deliver sufficient energy into a target location so as to bring
about a therapeutic effect. The elements can be positioned at the
distal-most tip of the needle, or proximal to the distal-most tip.
The delivered energy can be in the form of electrical,
radiofrequency, thermal (including heating and cooling effect),
microwave, short wave or acoustic energy. The energy delivering
device can be connected at its proximal end to a power source which
can include control and feedback capabilities. Irrigation channels
can be incorporated in the delivery device to provide coolant to
the treatment site during treatment. A typical application of this
treatment would include cancer ablation.
[0136] Optionally in any embodiment the delivery device can be
provided with sensors to provide feedback as to the local and/or
surrounding tissue parameters including electrical, chemical,
optical, acoustic, mechanical and thermal. Sensors can be disposed
proximate, distal to and proximal to the hydrogel outlet.
[0137] In another aspect, the invention provides a method of
performing a lung procedure (for example a lung biopsy or a lung
ablation procedure), comprising the steps of:
[0138] advancing a coaxial cannula into the lung, wherein a distal
portion of the coaxial cannula has one or more apertures in a side
wall thereof;
[0139] advancing a lung procedure needle through the cannula to a
procedure site within the lung; actuating the lung procedure needle
to perform a lung procedure at the procedure site;
[0140] withdrawing the lung procedure needle through the
cannula;
[0141] advancing a hydrogel delivery needle through the coaxial
cannula, wherein a distal portion of the hydrogel delivery needle
has one or more apertures in a side wall thereof corresponding to
the one or more apertures in the side wall of the coaxial
cannula;
[0142] aligning the one or more apertures of the coaxial cannula
and hydrogel delivery needle;
[0143] injecting a composite viscoelastic hydrogel of the invention
through the one or more outlets in the hydrogel delivery needle and
one or more outlets of the coaxial cannula into the lung to form a
sealing plug that embraces the coaxial cannula and typically abuts
the visceral pleura; and
[0144] withdrawing the coaxial cannula and hydrogel delivery needle
through the sealing plug.
[0145] In one embodiment, the composite viscoelastic hydrogel is
delivered adjacent the visceral pleura of the lung. In one
embodiment, the lung procedure needle is a biopsy needle. In one
embodiment, the lung procedure needle is a tissue ablation
probe.
BRIEF DESCRIPTION OF THE FIGURES
[0146] FIGS. 1A-1D. Series of lateral views illustrating a
transthoracic needle biopsy procedure and demonstrating how a
pneumothorax occurs (prior art).
[0147] FIGS. 2A-2D. Series of lateral views illustrating
embodiments of the delivery device and a method of delivering a
hydrogel plug to a target location in the lung.
[0148] FIGS. 3A-3C. Series of graphs demonstrating the effect of
crosslinked gelatin powder particle size on the dynamic
viscoelastic properties of biphasic hydrogels.
[0149] FIGS. 4A-4C. Series of graphs demonstrating the influence of
varying concentrations of crosslinked gelatin powder and hyaluronic
acid on the dynamic viscoelastic properties of biphasic
hydrogels.
[0150] FIG. 5. Graph showing the effect of varying concentrations
of crosslinked gelatin powder and hyaluronic acid on the
compressive modulus of biphasic hydrogels.
[0151] FIG. 6. Graph demonstrating the shear thinning behaviour of
biphasic hydrogels.
[0152] FIGS. 7A-7D. Series of charts showing the effect of terminal
steam sterilization on the dynamic viscoelastic properties of the
biphasic hydrogel.
[0153] FIG. 8. Graph showing the effect of dehydrothermal treatment
temperature on the viscosity of the biphasic hydrogel.
[0154] FIGS. 9A-9C. A section of a CT scan of the lung
demonstrating the degradation of the biphasic hydrogel in vivo.
[0155] FIG. 10. Graph showing the injection force required to
inject various biphasic hydrogels through an 18G delivery
needle.
DETAILED DESCRIPTION OF THE INVENTION
[0156] All publications, patents, patent applications and other
references mentioned herein are hereby incorporated by reference in
their entirety for all purposes as if each individual publication,
patent or patent application were specifically and individually
indicated to be incorporated by reference and the content thereof
recited in full.
[0157] The high efficacy demonstrated by exemplary embodiments
disclosed herein is due to the unique viscoelastic properties of
the hydrogel delivered. A hydrogel has both flow and elastic
properties. Elasticity is reversible deformation; i.e. the deformed
body recovers its original shape. The mechanical properties of an
elastic solid may be studied by applying a stress and measuring the
deformation of strain. Flow properties are defined by resistance to
flow (i.e. viscosity) and can be measured by determining the
resistance to flow when a fluid is sheared between two surfaces.
The physical properties of a gel by viscoelasticity can be
expressed by dynamic viscoelastic characteristics such as storage
modulus (G'), loss modulus (G''), tangent delta (tan .delta.) and
the like. Storage modulus characterizes the firmness of a
composition and describes the storage of energy from the motion of
the composition. Viscous modulus is also known as the loss modulus
because it describes the energy that is lost as viscous
dissipation. Tan .delta. is the ratio of the viscous modulus and
the elastic modulus, tan .delta.=G''/G'. A high storage modulus and
a low loss modulus indicate high elasticity, meaning a hard gel.
Reversely, a high loss modulus and a low storage modulus mean a gel
with high viscosity.
[0158] When the hydrogel described herein is used as a biomedical
material, e.g., a biodegradable hydrogel plug for use in the
periphery of the lung to prevent pneumothorax, it is considered
that the increased stiffness and storage modulus of the gel can
bring about improvement in sealing and barrier effect between
tissues. It would also contribute to a prolonged duration
(increased retention) at the target site, especially if the
elasticity is greater than the elasticity of the surrounding
tissues. The flowable nature of the hydrogel is due to its high Tan
.delta. and at rest this allows for improvement in apposition with
the surrounding tissue. This flow property also provides the
hydrogel with its self-healing ability.
[0159] Therefore, it is preferably desirable that the gel for such
use have well-balanced elasticity and viscosity. If the hydrogel
zero shear viscosity is too high and if the gel does not portray
sufficient shear thinning properties, it may become too difficult
to inject through the delivery device into the target site. The gel
may not readily appose surrounding tissue to form a barrier against
fluid leak. Also, the gel may not readily flow back into the needle
tract once the needle has been removed. On the other hand, if tan
.delta. exceeds 0.8, the gel behaves like a solution, and it may
infiltrate the surrounding tissue or be ejected from the needle
tract. That is, the hydrogel described herein is regarded to have
the most suitable physicochemical and rheological properties as a
viscous plug for lung biopsy.
[0160] The term "viscoelastic hydrogel" therefore refers to a
hydrogel that exhibits viscoelastic properties. It generally has a
storage modulus (G') of preferably greater than 400 Pa, more
preferably greater than 800 Pa and even more preferably greater
than 1000 Pa. The viscoelastic hydrogel may exhibit a tangent delta
(tan .delta.; G''/G') of from 0.05 to 0.8, preferably from 0.1 to
0.5 and more preferably from 0.2-0.5 in dynamic viscoelasticity at
a frequency of 1 Hz. Preferably, the viscoelastic hydrogel exhibits
a loss modulus (G'') of from 200 to 6000 Pa, more preferably from
400 to 2000 Pa, in dynamic viscoelasticity at a frequency of 1 Hz
at 25.degree. C. The viscoelastic hydrogel may be free of
crosslinking, lightly crosslinked, or strongly crosslinked to
provide appropriate characteristics, for example to increase its
storage modulus (G') or to increase its in vivo residence time.
[0161] As used herein, the term "shear thinning" as applied to a
hydrogel means that when shear stress is applied to the hydrogel,
the storage modulus (G') reduces, the tan .delta. increases and the
overall viscosity reduces. This property provides injectable
properties to the hydrogel. And allows it to be injected through a
narrow-gauge needle, such as used in minimally invasive procedures
such as lung biopsy (17-20 gauge) or lung ablation (10-14 gauge).
The shear thinning hydrogel described herein typically exhibits a
range of a storage modulus (G') of 1-100 Pa, preferably from 1-50
Pa in dynamic viscoelasticity at a frequency of 1 Hz and 100%
strain. Furthermore, the hydrogel described herein has self-healing
properties and retain their high storage modulus (G') and loss
modulus (G'') when the shear strain is removed.
[0162] The hydrogel described herein possess shear thinning
capabilities. That is, when shear stress is applied, the storage
modulus (G') reduces, the tan .delta. increases and the overall
viscosity reduces. This property allows the gels to be injected
through a narrow-gauge needle, such as used in minimally invasive
procedures such as lung biopsy. The gel described herein portrays
the physical properties with ranges of a storage modulus (G') of
less than 100 Pa, preferably less than 50 Pa in dynamic
viscoelasticity at a frequency of 1 Hz and 100% strain.
Furthermore, the gels described herein portrays rapid thixotropic
recovery properties and retain their high storage modulus (G') and
loss modulus (G'') immediately on removal of the high shear
rate.
[0163] As used herein, the term "self-healing" as applied to a
viscoelastic hydrogel of the invention refers to the ability of the
hydrogel to reform together. "Self-healing" may also be described
as the ability of the hydrogel to spontaneously form new bonds when
old bonds are broken within the material. As an example, when an
annular sealing plug of viscoelastic hydrogel is delivered around a
delivery needle, a self-healing viscoelastic hydrogel will flow
back together once the needle is removed to form a non-annular
sealing plug, typically consisting of a single-bodied cohesive
matrix.
[0164] Optionally in any embodiment the sealing hydrogel plug
should be able to self-heal a channel through its centre
independent of its in vivo environment. By this we refer to the
ability of the hydrogel to fill the channel through a time
dependent viscoelastic flow mechanism.
[0165] Optionally in any embodiment the sealing hydrogel plug
should be able to self-heal a channel through its centre dependent
on its in vivo environment. Stresses from the in vivo environment
imposed on the hydrogel plug may improve its ability to self-heal
in a shorter duration compared to an uninterrupted plug.
[0166] Optionally in any embodiment, the composite viscoelastic
hydrogel should be able to self-heal under its own weight without
any influence from the surrounding environment. This may be
demonstrated by creating a singular mass of the hydrogel, for
example a sphere of the hydrogel created using approximately 0.5 ml
of hydrogel. A cylindrical channel can be created through the
centre of the sphere by passing a 17 gauge needle through its
centre and retracting the needle. The sphere with the cylindrical
channel through its centre can be placed at rest on a bench with
the axis of the cylindrical channel perpendicular to the bend. The
size of the channel can be monitored over time. Referring to the
viscoelastic hydrogels described in this invention, specifically
hydrogels comprising 2-6% hyaluronic acid, the following are the
observations: initially the channel in the ball will be visible,
but over time (1-15 mins, depending on the hydrogel formulation)
this channel will close over as the hydrogel self-heals. This is as
a result of the time dependent flow of the hydrogel.
[0167] Optionally in any embodiment, part or all of the composite
viscoelastic hydrogel is comprised of a hyaluronan hydrogel. The
hyaluronan polymer forms a continuous phase throughout the
three-dimensional matrix. Optionally in any embodiment, the
viscoelastic hydrogel is a high molecular weight hyaluronan
hydrogel. Optionally in any embodiment, the composite viscoelastic
hydrogel is a shear thinning hydrogel (viscosity decreases under
shear strain). Examples of polymer materials that may be employed
to make a viscoelastic hydrogel include hyaluronan, especially high
molecular weight hyaluronan. Other hydrogel materials suitable for
use in the present invention are outlined in the review articles
`Shear-thinning hydrogels for biomedical applications`, Soft
Matter, (2012) 8, 260, `Injectable matrices and scaffolds for drug
delivery in tissue engineering` Adv Drug Deliv Rev (2007) 59,
263-272, and `Recent development and biomedical applications of
self-healing hydrogels` Expert Opin Drug Deliv (2017) 23: 1-15.
[0168] As used herein, the term "hyaluronan" or "hyaluronic acid"
or "HA" refers to the anionic non-sulphated glycosaminoglycan that
forms part of the extracellular matrix in humans and consists of a
repeating
disaccharide.fwdarw.4)-.beta.-d-GlcpA-(1.fwdarw.3)-.beta.-d-GlcpNAc-(1.fw-
darw., or any salt thereof. Hyaluronan is the conjugate base of
hyaluronic acid, however the two terms are used interchangeably.
When a salt of hyaluronic acid is employed, the salt is generally a
sodium salt, although the salt may be employed such a calcium or
potassium salts. The hyaluronic acid or hyaluronan may be obtained
from any source, including bacterial sources. Hyaluronic acid
sodium salt from Streptococcus equi is sold by Sigma-Aldrich under
the product reference 53747-1G and 53747-10G. Microbial production
of hyaluronic acid is described in Liu et al (Microb Cell Fact.
2011; 10:99). The term also includes derivatives of hyaluronic
acid, for example hyaluronic acid derivatised with cationic groups
as disclosed in US2009/0281056 and US2010/0197904, and other types
of functionalised derivatives, such as the derivatives disclosed in
Menaa et al (J. Biotechnol Biomaterial S3:001 (2011)), Schante et
al (Carbohydrate Polymers 85 (2011)), EP0138572, EP0216453,
EP1095064, EP0702699, EP0341745, EP1313772 and EP1339753.
[0169] Hyaluronic acid can be categorised according to its
molecular weight. High molecular weight (preferably>1000 kDa (1
Mda)), medium molecular weight (preferably 250-1000 kDa), low
molecular weight (preferably 10-250 kDa), and oligo hyaluronic acid
(preferably<10 kDa). The effect of molecular weight on
hyaluronic acid hydrogel viscosity has previously been reported.
The stiffness and viscosity of the final gel is dependent on both
molecular weight and solution concentration. In studying the
rheological properties of hyaluronic acid with different molecular
weights, Rheological and cohesive properties of hyaluronic acid J
Biomed Mat Res, 76A, 4, Pg 721-728, Falcone et al found that high
molecular weight hyaluronic acid is considerably more cohesive than
low molecular weight hyaluronic acid. It has been shown that the
presence of high molecular weight hyaluronic acid hydrogels at a
wound site leads to reduction in scarring. High molecular weight
hyaluronic acid has been shown to be anti-inflammatory, enhanced
angiogenesis and enhanced immunosuppression. Jiang et al found that
high molecular weight hyaluronic acid has been shown to protect
from epithelial apoptosis in lung injury "Regulation of lung injury
and repair by Toll-like receptors and hyaluronan" Nature Medicine
(2005) 11, 11 1173-1179. Furthermore, inhalation of high molecular
weight hyaluronic acid has been used to treat lung conditions such
as bacterial rhinopharyngitis, chronic bronchitis, cystic fibrosis
and asthma. In some embodiments, the hyaluronic acid compositions
of the hydrogel are free from crosslinking and are free from other
therapeutic agents. Hyaluronic acid based hydrogels with
characteristics potentially suitable for this application are
described in U.S. Pat. No. 9,492,474B2. `Compositions of`
hyaluronan with high elasticity and uses thereof. This document
describes a material, Elastovisc.TM., comprised of high
concentration and molecular weight hyaluronic acid. Its intended
use is for injection into joints to relieve pain and treat
osteoarthritis.
[0170] As used herein, the term "hyaluronan hydrogel" preferably
includes a three-dimensional network of hyaluronan polymers in a
water dispersion medium. The hyaluronan polymer forms a continuous
phase throughout the three-dimensional matrix. Optionally in any
embodiment, the hyaluronan polymers are non-crosslinked. Optionally
in any embodiment, the hydrogel is free of a crosslinking agent.
Optionally in any embodiment, the matrix is formed with a
homopolymer, typically a hyaluronic acid homopolymer. Optionally in
any embodiment, the hydrogel is pH balanced or buffered to match
the pH of the physiological environment. Optionally in any
embodiment, the matrix is lightly crosslinked. Any crosslinking
agent known to crosslink hyaluronic acid may be used for this
purpose. Crosslinking agents may include epichlorohydrin, divinyl
sulfone, I, 4-bis (2,3-epoxypropoxy) butane (or I, 4-bis
(glycidyloxy) butane or 1,4 butanediol diglycidyl ether=BDDE), the
I, 2-bis (2,3-epoxypropoxy) ethylene, I-(2,3-epoxypropyl)-2,
3-epoxy cyclohexane.
[0171] Optionally in any embodiment, the continuous phase of the
composite viscoelastic hydrogel may be comprised of
`multi-component` hydrogel which refers to at least two hydrogels
that are evenly blended and dispersed together to form a homogenous
hydrogel continuous phase. This construct may also be referred to
as a semi-interpenetrating polymer (hydrogel) network or
interpenetrating polymer (hydrogel) network comprised of two or
more hydrogels. As an example, a hyaluronan hydrogel (concentration
may range from 1-5%) may be blended with a methylcellulose hydrogel
(concentration may range from 3-15%). In the same manner, more than
two hydrogels may be combined to form a single cohesive network
whereby each hydrogel provides improved properties to the overall
network. The properties of each hydrogels may be provided to
increase stiffness and viscosity, to provide improved injectability
(shear thinning), to provide improved self-healing, to prolong the
residence (biodegradation) time of the hydrogel in vivo, to provide
haemostatic properties, to provide antibacterial properties, to
provide anti-inflammatory properties, to provide anti-coagulant
properties, to provide pro-coagulant properties, to provide colour
and marking capability (under visible and radiographic detection),
to provide some diagnostic or therapeutic effect (for example
chemotherapy), to provide resistance to extremes of heat (hot and
cold), to provide improved biocompatibility, and to improve
manufacturability and preparation of the overall hydrogel. One or
more of these hydrogels may be crosslinked to provide improved
properties, for example to increase the residence time of the
hydrogel in vivo
[0172] The viscoelastic hydrogel is generally a "biphasic"
hydrogel, which refers to a hydrogel formed by combining (through
mixing or blending) a colloidal hydrogel with a continuous phase
hydrogel. The colloidal hydrogel will form an evenly dispersed
phase in the continuous hydrogel phase. A variety of natural and
synthetic biodegradable polymers can be used to form the continuous
hydrogel phase. Glycosaminoglycans, for example hyaluronan and its
derivatives form one example. The hyaluronan may be preferably
non-crosslinked or possibly lightly crosslinked so as to retain its
viscoelastic properties, especially its shear thinning and
self-healing ability. Optionally in any embodiment, the hyaluronan
may be provided at concentrations of 0.4-6%. A variety of
biodegradable polymers are also suited to form the colloidal
hydrogel phase as outlined previously (collagen and gelatin are two
examples). The colloidal hydrogel phase can be added in sufficient
quantities to provide the advantage of increased residence time of
the hydrogel in vivo. This can allow the necessary time to provide
for healing of the tissue. An additional benefit is that an
increased residence time can provide a long-term marking function
of the biopsy side for use under video-assisted thoracoscopic
(VATS) surgery. A suitable polymer is one that is insoluble in an
aqueous environment and can be achieved by crosslinking of the
polymer through conventional means. An example would be
dehydrothermally crosslinked gelatin. It should be noted that by
introducing a too large amount of the colloidal hydrogel phase, it
may jeopardize the injectability and self-healing ability of such
compositions. Optionally in any embodiment, the "biphasic" hydrogel
can comprise a colloidal hydrogel at concentrations of 0.2-30%,
15-28%, or 20-25% of hydrogel forming polymer (w/v). Optionally, in
any embodiment, the ratio of continuous phase polymer to dispersed
phase polymer in the viscoelastic hydrogel is about 1:10 to 1:40,
more preferably about 1:10 to 1:20.
[0173] Optionally in any embodiment, the composite viscoelastic
hydrogel exhibits a storage modulus (G') of greater than 400 Pa,
more preferably greater than 600 Pa, more preferably greater than
800 Pa, more preferably greater than 1000 Pa. Optionally in any
embodiment, the viscoelastic hydrogel exhibits tan .delta. (G''/G')
from 0.01 to 0.8, more preferably 0.1 to 0.6 in dynamic
viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate
at 25.degree. C.
[0174] Optionally in any embodiment, the composite viscoelastic
hydrogel may be provided as a powder that is reconstituted in a
physiologically acceptable fluid, for example water, saline,
autologous blood, or autologous plasma prior to the surgical
procedure. Synthetic fluids such as low molecular weight PEG and
glycerol may also be employed. The powder may be comprised of any
suitable biocompatible polymer or combinations of polymers. In one
embodiment, the powder may be provided in the hydrogel delivery
needle. In one embodiment, the powder may be provided in a syringe
with a suitable reconstitution fluid provided in a second syringe.
In one embodiment, the powder has an average particle size of
1-500, 10-100 or 30-40 microns. The powder may be both regular or
irregular in both shape, morphology and size distribution and may
be formed through milling or other means known in the art. In
certain instances, powder hydration can be controlled by varying
the level of de-hydration of the powder particles such as in the
case of collagenous based materials, for example collagen or
gelatin.
[0175] Optionally in any embodiment, the hydrogel described herein
may be provided in separate components, for example in multiple
syringes and the means can be provided to allow mixing of the
components prior to injection through the syringe. Crosslinking
agents can be provided in one or more of these components to
provide the material characteristics necessary to achieve a shear
thinning and self-healing hydrogel. Mixing can be achieved by
reciprocating the contents between the syringes and a static mixer
can be employed to speed up this process.
[0176] In any embodiment the composite viscoelastic hydrogel
composition can be provided in a physiological buffer, e.g., a
phosphate buffer or a bicarbonate buffer. In some embodiments, the
pH of the composition is between pH 5 and pH 9 or between pH 7.5
and pH 8.5. In some embodiments, the pH of the composition is 8.0.
In some embodiments, the pH of the composition is 7.5. In some
embodiments, the pH of the composition is 8.5. If needed, acid
(such as HCL) or base (such as NaOH) can be added to the
composition to attain the desired pH. In a specific embodiment, the
hyaluronic acid hydrogel described herein consists essentially of
hyaluronic acid present at a concentration of 50 mg/ml (or about 5%
W/V, and having an average molecular weight of between 1-2 Mda.
Ranges intermediate to the recited values are also intended to be
part of this invention. For example, hyaluronan content in the
compositions described herein may be between about 0.5% and about
6% (weight/volume). It should further be appreciated that the
amount of hyaluronan in a particular volume may also be expressed
by alternative means (e.g., mg/ml, gram/litre or mol/litre). A
person of ordinary skill in the art would know how to convert the
various means of expressing the amount of hyaluronan in a
particular volume
[0177] As used herein, the term "sealing plug", "hydrogel plug" or
"gel plug" refers to a single body of viscoelastic hydrogel, for
example biphasic composite hydrogel, that is suitable for delivery
through a needle to a locus in the lung and which has sufficient
viscoelasticity to push away the tissue surrounding the needle and
coalesce to form a single closed annular sealing plug around the
needle. The viscoelastic properties and stiffness of the gel
prevents infiltration of the tissue, allowing the gel to precisely
oppose the tissue and form an effective seal around the needle and
subsequently cannula thereby preventing air from lungs leaking past
the plug. The viscoelastic behaviour of the hydrogel allows the
annular plug to coalesce upon removal of the cannula filling the
hole in the annular plug and bearing against the visceral pleura to
seal it after withdrawal of the coaxial cannula.
[0178] Optionally in any embodiment, the hydrogel plug should
exhibit "limited-swelling" behaviour which means that its bulk size
should not increase by any profound extent when placed in vivo, for
example below the surface of the lung to prevent pneumothorax. A
hydrogel plug that swells by a significant degree may cause
unwanted physiological or biological effect. Some swelling of
hydrogels in vivo is to be expected but in order to preserve the
native tissue, swelling of the hydrogel plug should be limited.
Swelling can be characterised by forming a predetermined size of
hydrogel sphere, for example rolling 500 .mu.l of hydrogel into a
sphere, and by placing this ball of hydrogel into an aqueous
solution. This volume 500 .mu.l will initially equate to a sphere
with a diameter of approx. 10 mm. The aqueous solution may be a
saline or simulated body fluid solution and it may also contain the
correct enzyme activity that is found in vivo. The size and shape
and dissolution of the ball of hydrogel can then be monitored over
a prolonged period of time. The swelling ratio can be determine
from:
Swelling (%)=(Ws-Wd)/Wd.times.100
[Wd=Weight of polymer; Ws=weight of swollen polymer]
[0179] Preferably the selling ratio should not exceed 250%, more
preferably it should not exceed 150%, and more preferably it should
not exceed 130%. Sample degradation can be determined by comparing
the dry weight of the polymer over time. Dry weight can be
determined by lyophilising the samples. The degradation rate can be
inferred from the remaining weight of the hydrogel:
Remaining Hydrogel (%)=(W2-W1)/W1.times.100
[W1=Original dry weight of polymer; W2=time dependent dry weight of
polymer]
[0180] Different polymeric materials with thermo-responsive,
shear-thinning, shape memory and biological properties can be
combined to yield composite hydrogels with improved properties for
this application. Improvements can include enhanced
biocompatibility, injectability, viscosity, altered biodegradation,
drug attachment, tissue adhesion, cohesiveness, sealing ability
stability, hydrophilicity. Gelatin and hyaluronic acid are two
examples. Substances which can be combined with these polymer
include methylcellulose, oxidized cellulose, carboxylmethyl
cellulose, and carboxylic acid.
[0181] Optionally in any embodiment, the composite viscoelastic
hydrogel can include contrast medium which refers to an additive
that can be included in the gel in an appropriate amount that
allows the hydrogel to be contrasted against the surrounding
tissue. In this way, the hydrogel plug and injected location can be
visually identified and/or targeted for example during the surgical
procedure or during a follow up surgical procedure.
[0182] Identification can be visual or through guidance systems
such as CT scans, ultrasound or fluoroscopy. Additives which can be
added to the hydrogel in varying concentrations to achieve
effective visual contrast include ionic and non-ionic contrast
medium, methylene blue, indigo carmine, toluidine blue,
lymphazurine, hemotoxylin, eosin, indocyanine green (ICG), India
ink, carbon based powders such as carbon black, carbon nanotubes
and graphene, and ceramic powders such as aluminium oxide, titanium
dioxide, and calcium phosphates. The hydrogel may also comprise
additional detectable marking agents. The detectable marking agent
suitable for use in the hydrogel described herein may include any
composition detectable by spectroscopic, photochemical,
biochemical, immunochemical, electrical, optical or chemical means.
A wide variety of appropriate detectable markers are known in the
art, which include luminescent labels, radioactive isotope labels,
and enzymatic labels. These marking agents can be mixed with the
hydrogel or chemically conjugated to the hydrogel molecules.
[0183] Optionally in any embodiment, the composite viscoelastic
hydrogel can comprise of a therapeutic agent or biologically active
agent. Therapeutic agents which may be linked to, or embedded in,
the hydrogel include, but are not limited to, analgesics,
anaesthetics, antifungals, antibiotics, anti-inflammatories,
anthelmintics, antidotes, antiemetics, antihistamines,
antihypertensives, antimalarials, antimicrobials, antioxidants,
antipsychotics, antipyretics, antiseptics, antiarthritics,
antituberculotics, antitussives, antivirals, cardioactive drugs,
cathartics, chemotherapeutic agents, a colored or fluorescent
imaging agent, corticoids (such as steroids), antidepressants,
depressants, diagnostic aids, diuretics, enzymes, expectorants,
hormones, hypnotics, minerals, nutritional supplements,
parasympathomimetics, potassium supplements, radiation sensitizers,
a radioisotope, sedatives, stimulants, sympathomimetics,
tranquilizers, urinary anti-infectives, vasoconstrictors,
vasodilators, vitamins, xanthine derivatives, and the like.
Optionally in any embodiment, the hydrogel described herein
comprises one or more anesthetics. Exemplary anesthetics include,
but are not limited to, proparacaine, cocaine, procaine,
tetracaine, hexylcaine, bupivacaine, lidocaine, benoxinate,
mepivacaine, prilocalne, mexiletene, vadocaine and etidocaine.
Optionally in any embodiment, the viscoelastic hydrogel can further
comprise foaming agents, foam stabilizers, surfactants, thickeners,
diluents, lubricants, wetting agents, plasticizers.
[0184] Optionally in any embodiment, part or all of the composite
viscoelastic hydrogel can be "biodegradable" and configured to
degrade over time in-vivo. Different phases or components of the
viscoelastic hydrogel can be configured to degrade at different
rates. Biodegradable substances are preferably eliminated by the
body without causing an inflammatory or immune response. For the
viscoelastic hydrogel described herein, the period of time for full
biodegradation can be less than 1 year, preferably less than one
month, more preferably less than 1 week, and more preferably less
than 72 hours. The added benefit of a quick degradation period is
that it allows the lung tissue to return to normal and prevents
excess scar tissue formation at the delivery site. Also, limiting
residence time and scar tissue formation ensures that the delivery
of the hydrogel plug does not interfere with follow up radiological
analysis of the suspected lung lesion. Non-crosslinked systems may
result in a faster in vivo residence period compared to crosslinked
systems. The high molecular weight (>1000 kDa) and high
concentration (40-60 mg/ml) hyaluronic acid hydrogels described
herein have a degradation period of less than 1 week and also less
than 72 hours. Longer degradation periods are possible by modifying
the native hyaluronic acid molecular structure via crosslinking or
by other means. Longer degradation periods are also possible by
combining the hyaluronic acid hydrogel with one or more hydrogels
or colloidal hydrogels to form a composite hydrogel. One of the
hydrogels will remain at the target site for a longer period while
the other is removed. For example, the hyaluronic acid hydrogel may
be combined with a crosslinked polymer (for example hyaluronan,
hylan, collagen or gelatin) to form a composite hydrogel. The
cross-linked polymer can be configured to have a residence time of
greater than 1 week, and often greater than 2 weeks by the use of
various crosslinking modalities known in the art. Cross-linkers
employed as part of the implantable material precursors can include
aldehydes, polyaldehydes, esters, and other chemical functionality
suitable for cross-linking protein(s). Physical crosslinking
methods can also be employed, for example subjecting the polymers
to heat, cold or radiation. Crosslinking agents can be added to
improve cohesion, rigidity, mechanical strength and barrier
properties.
[0185] As used herein, the term "in-vivo residence time" as applied
to a sealing plug of composite viscoelastic hydrogel refers to the
period of time that sealing plug of 0.1-1 ml, preferably 0.2-0.8 ml
and more preferably 0.3-0.5 ml that persists in lung tissue in-vivo
without any significant loss of structure integrity. The in-vivo
residence time should be sufficient to allow healing of the hole in
the visceral pleura to occur, and ideally to allow for healing in
the surrounding lung tissue to occur. Methods of approximating the
in-vivo residence time of hydrogels are described below. To achieve
an appropriate in-vivo residence time to allow healing to occur,
the hydrogel can be comprised of certain unmodified materials
(including proteins) that have a longer residence time. Examples
include collagen, oxidised cellulose, starch, extracellular matrix
(ECM). Crosslinked hydrogels as described herein have been found to
have an in-vivo residence time of more than two weeks. Optionally,
the shear-thinning viscoelastic hydrogel may have an in-vivo
residence time of at least 1 week, preferably at least 2 weeks, and
ideally at least 3 weeks.
EXEMPLIFICATION
[0186] The invention will now be described with reference to
specific examples. These are merely exemplary and for illustrative
purposes only: they are not intended to be limiting in any way to
the scope of the monopoly claimed or to the invention described.
These examples constitute the best mode currently contemplated for
practicing the invention.
[0187] The mechanism of pneumothorax resulting from a transthoracic
needle biopsy is illustrated in FIGS. 1A-1D (Prior art). FIG. 1A
illustrates a cross section of the thoracic cavity A, which
comprises the thoracic (chest) wall muscle B, ribs C, lung tissue
D, and the pleural cavity E defined by the serous membrane of the
thoracic wall (parietal pleura F) and the serous membrane of the
lung (visceral pleura G). During a lung biopsy procedure (FIG. 1B),
a core needle H and coaxial cannula I are advanced percutaneously
through the skin O and through the pleural cavity E towards a
suspected lung nodule J. In FIG. 1C the core needle H has been
withdrawn and replaced with a biopsy needle K which is advanced
through the cannula 2 and obtains a tissue sample from the
suspected lung nodule J. As illustrated in FIG. 1D, removal of the
biopsy needle K and cannula I leaves a void L in the lung tissue D
and also leaves a hole L1 in the visceral pleura G. The dense
muscular tissue of the thoracic wall B contracts around the void
caused by removal of the needles. However, the holes L, L1 created
by the biopsy needles in the lung tissue D and visceral pleura G do
not completely seal over. Due to the pressure gradient between the
lung tissue D and pleural cavity E, air escapes through the hole L1
created in the visceral pleura G and enters the pleural cavity E,
creating a collection of air in the pleural cavity E known as a
pneumothorax M. If a blood vessel of significant size is punctured
during the biopsy procedure the pleural cavity may also fill with
blood, a condition known as a haemothorax. The prevalence of
haemothorax is not as high as pneumothorax. The haemothorax or
pneumothorax M can grow in sufficient size to cause the lung to
partially or fully collapse and bring about respiratory distress
and the need for treatment.
[0188] Referring to FIGS. 2A-2E, a method for overcoming the
shortcomings of the prior art is presented. In FIGS. 2A-2E, a
method of delivering a viscoelastic hydrogel plug to a target
location in the lung is described. This embodiment, employs a
medical device system comprising a coaxial cannula 2 having a
distal-most end 2A and a proximal connector such as a luer lock 2B,
a core needle 3, and a hydrogel delivery needle 4 having a distal
tissue piercing tip 5 and a hydrogel outlet 6 disposed on a side of
the needle proximal of the piercing tip 5. Also contained in the
system is a syringe 15 with reservoir 15b filled with viscoelastic
hydrogel material including any of those described herein. The
syringe may be replaced by any pump, plunger, fluid advancement
mechanism or element suitable for delivering a viscous
hydrogel.
[0189] As shown in FIG. 2A, the core needle 3 and cannula 2
assembly are inserted into the chest wall of the patient to a depth
at which the assembly is located in the chest wall B and does not
penetrate the lung D. A coaxial cannula 2 refers to a needle device
having an inner lumen configured to receive a penetrating device,
for example a core needle 3 where the assembled core needle and
cannula 2 may be used to enter through the skin surface on the
chest. Generally, the coaxial cannula has a gauge size of 10 to 19.
In additional embodiments, the coaxial cannula may also be referred
to as a sheath, an introducer, an obturator/stylet assembly, a
guiding catheter, trocar, port device or other medical introductory
device known in the art.
[0190] As shown in FIG. 2B, the core needle 3 has been withdrawn
from the cannula 2 and a hydrogel delivery needle 4 is advanced
through the cannula 2. The hydrogel delivery needle 4 typically has
a piercing tip, and a hydrogel outlet 6 which is typically disposed
on a side of the needle proximal of the piercing tip 5, for example
0.5-15 mm from the piercing tip 5. The delivery needle 4 has a
distal-most end configured for insertion into the body, and a
proximal end which during use is positioned outside of the body.
The needle is generally formed from a metal, although the
positioning (adjustment) mechanism may be formed from plastic or
polymer or a metal. The needle may comprise polymer tubing at its
proximal end and may include a luer lock to facilitate fluidically
connecting the needle (or polymer tubing part) to a pump or syringe
15. Generally, the hydrogel delivery needle 4 has a gauge of 13 to
20. The hydrogel delivery needle 4 is inserted to a depth at which
the hydrogel outlet 6 is positioned in the lung tissue distal of
the pleural cavity E and visceral pleura G. Positioning of the
hydrogel outlet 5 at this target location may be achieved under CT
guidance by employing a radiopaque or radiolucent marker 32 on the
delivery needle which can be positioned a known distance X from the
hydrogel outlet 6. By overlaying the radiolucent marker 32 with the
pleural cavity E, the hydrogel outlet can be positioned a
predetermined distance X inside the lung D from the pleural cavity
E. The pleural cavity E is a very thin space approximately 25 .mu.m
in width and is often referred to as a virtual cavity. As can be
seen later in FIG. 7, the pleural cavity E can be distinguished
under CT guidance as the transition between the lung (dark area)
and chest wall (bright area). Positioning of the radiolucent marker
32 over the pleural cavity E may be achieved by stepwise scanning
and fine adjustment of the needle 4, or with fine adjustment under
continuous fluoroscopic guidance.
[0191] As shown in FIG. 2C, a syringe 15 with hydrogel filled
reservoir 15B is attached to the delivery needle 4 via a luer lock
12. A predefined quantity of viscoelastic hydrogel is then injected
into the lung through the hydrogel outlet 6 to form a closed
annular viscoelastic sealing plug 7 around the delivery needle 4.
Subsequent to this step, the coaxial cannula 2 is advanced over the
delivery needle 4 through the sealing plug 7 and towards the
suspected lung nodule J. The hydrogel delivery needle 4 is
withdrawn leaving the cannula 2 with surrounding hydrogel sealing
plug 7 in place for receipt of a lung biopsy needle K. As shown in
FIG. 2D a lung biopsy needle K can be then advanced through the
cannula 2 and a lung biopsy carried out, The biopsy needle K and
cannula 2 are both withdrawn after the biopsy has been taken. As
shown in FIG. 2E the sealing plug 7 remains in position in the lung
tissue after the needles have been withdrawn. Due to the physical
properties of the viscoelastic hydrogel material, the sealing plug
7 reflows into the space left behind by the needles, as well as
sealing the hole L1 left in the visceral pleura G by the coaxial
cannula 2. These steps describe a method of performing a lung
biopsy with diminished chance of causing a pneumothorax. The
efficacy of the sealing plug 7 is dependent on its ability to block
any air in the aerated lung tissue D from exiting the hole L1 in
the visceral pleura G.
[0192] For a number of reasons it may be difficult to position the
delivery device as outlined above. Firstly, fluoroscopic guidance
may not be available to the clinician so that the delivery needle 4
with marker band 32 cannot be accurately positioned. Secondly, it
may be harmful to expose the patient to too many CT scans and
resulting high radiation dose to achieve accurate placement of the
needle marker band 32. Furthermore, delayed placement of the
hydrogel plug may lead to potential pneumothorax while the needle
is in the lung tissue unprotected. In order to quickly, easily and
accurately target the required depth of injection in the lung for
the viscoelastic hydrogel to achieve an effective seal, a
positioning mechanism is provided with the hydrogel delivery needle
4 as will be described hereafter.
EXAMPLES
[0193] Example 1: A composite viscoelastic hydrogel comprising
hyaluronic acid and crosslinked gelatin was created using the
following method. Type A porcine derived gelatin (300Bloom) was
dissolved fully in water at 7% w/v at 40.degree. C. and allowed to
set at 4.degree. C. overnight. The resulting gel were subsequently
freeze dried by freezing at -40.degree. C. and drying at 25.degree.
C. under a constant vacuum of 0.1 mbar. The dried constructs were
then heated under vacuum conditions (0.001 mbar) for 24 hours at
140.degree. C. to induce crosslinking. The sponge was then roughly
diced before being milled to form a fine powder using a cryo-mill
(Model: 75 Spex SamplePrep, LLC.). The powder was sieved using a
125 .mu.m sieve and the resultant powders had a powder particle
size distribution of D.times.10=7.4 .mu.m, D.times.50=32.8 .mu.m,
D.times.90=95 .mu.m as measured using a Mastersizer 3000 laser
diffraction particle size analyser (Malvern Panalyticlal ltd). The
dehydrothermally crosslinked gelatin powder was mixed with sodium
hyaluronate powder (molecular weight: 1.8-2 MDa) and the powder
mixture was hydrated with phosphate buffered saline solution at the
following concentration: Gelatin: 130 mg/ml, Sodium hyaluronate: 35
mg/ml. The resulting hydrogel was loaded into a syringe. The
hydrogel was employed to prevent pneumothorax during a CT-guided
transthoracic needle biopsy procedure as outlined in FIG. 8A-8F.
This procedure was performed in a porcine model. The hydrogel
formed an annular sealing plug around the needle during the biopsy
procedure and after the needles were withdrawn, the hydrogel
self-healed to prevent pneumothorax. The hydrogel persisted at the
site for at least 1 week as was evidence from CT-scan
follow-up.
[0194] Example 2: A composite viscoelastic hydrogel comprising
hyaluronic acid and crosslinked gelatin was created using the
following method. A type A porcine derived gelatin powder
(300bloom) was ground to a fine powder using a cryo-mill (Model: 75
Spex SamplePrep, LLC.). The powder was sieved using a 125 .mu.m
sieve and the resultant powders had a powder particle size
distribution of D.times.10=5.4 .mu.m, D.times.50=35.5 .mu.m,
D.times.90=90 .mu.m as measured using a Mastersizer 3000 laser
diffraction particle size analyser (Malvern Panalyticlal ltd). The
resultant fine powder was heat treated under vacuum conditions
(0.001 mbar) for 24 hours at 160.degree. C. to induce crosslinking.
The DHT crosslinked gelatin powder was mixed with sodium
hyaluronate powder (molecular weight: 1.8-2 MDa) and the powder
mixture was hydrated with phosphate buffered saline solution at the
following concentration: Gelatin: 100 mg/ml, Sodium hyaluronate: 45
mg/ml. The resulting hydrogel was loaded into a syringe. The
hydrogel was employed to prevent pneumothorax during a CT-guided
transthoracic needle biopsy procedure similar to that outlined in
FIG. 8A-8F. This procedure was performed in a porcine model. The
hydrogel formed an annular sealing plug around the needle during
the biopsy procedure and after the needles were withdrawn, the
hydrogel self-healed to prevent pneumothorax.
[0195] Using the above method, various concentrations of the
biphasic gel were evaluated rheologically and experimentally. The
measurement of the dynamic viscoelasticity and dynamic viscosity of
the hydrogels was made using a rheometer Model AR2000 manufactured
by TA Instruments under the following conditions.
[0196] Method of measurement: oscillation test method, strain
control
[0197] Measuring temperature: 25.degree. C.
[0198] Geometry: 4.degree. cone plate angle
[0199] Measuring geometry: 4 cm
[0200] Truncation gap: 112 .mu.m
[0201] Frequency: 1 Hz
TABLE-US-00001 Storage Crosslinked Sodium Modulus @ Tan.delta. @
Gelatin Hyaluronate 1 Hz & 1 Hz & Zero shear Viscosity @
Concentration Concentration 1% Strain 1% Strain viscosity 100
s.sup.-1 100 mg/ml 45 mg/ml 5,813 Pa 0.4 18,367 Pa s 6.8 Pa s 150
mg/ml 45 mg/ml 11,667 Pa 0.27 43,317 Pa s 10.0 Pa s 100 mg/ml 35
mg/ml 2,722 Pa 0.45 6,700 Pa s 4.2 Pa s 150 mg/ml 35 mg/ml 6,406 Pa
0.37 14,150 Pa s 5.9 Pa s
[0202] Example 3: FIGS. 3A-3C. represents viscoelastic properties
of composite viscoelastic hydrogels of the invention as measured
using a dynamic oscillatory test method. Various biphasic
viscoelastic hydrogels comprising hyaluronic acid and crosslinked
gelatin powder were fabricated using the following method. A type A
porcine derived gelatin powder (300bloom) was ground to a fine
powder using a cryo-mill (Model: 75 Spex SamplePrep, LLC.). The
powder was sieved using sieves of varying mesh size to yield powder
particles of the following size ranges: <53 .mu.m, 50-100 .mu.m,
100-200 .mu.m, 200-300 .mu.m. The resultant gelatin powders were
heat treated under vacuum conditions (0.001 mbar) for 24 hours at
140.degree. C. to induce dehydrothermal (DHT) crosslinking. The DHT
crosslinked gelatin powders were subsequently mixed with 7.5 mg/ml
sodium hyaluronate powder (molecular weight: 1.8-2 MDa) at the
following concentrations: 120 mg/ml, 140 mg/ml, 160 mg/ml. The
powder mixtures were hydrated with phosphate buffered saline to
form the biphasic hydrogels. The test rheometer used to measure the
rheological properties of the hydrogels was a model MCR102 by Anton
Paar GmbH. Dynamic oscillatory tests were conducted under stress
control, with a 25 mm flat plate geometry, a gap of 1 mm, an
analysis temperature of 25.degree. C. and over the frequency range
0.1-10 Hz. The graphs present the Storage Modulus (G') (FIG. 3A),
Loss Modulus (FIG. 3B), and Tan Delta (G'/G'') (FIG. 3C) for the
biphasic hydrogels taken at a frequency of 1 Hz. The data
demonstrates how the Storage (G') and Loss (G'') moduli both
increase for increasing powder particle size, however the Tan
.delta. (G'/G'') decreases for increasing powder particle size.
[0203] Example 4: FIGS. 4A-4C. represents viscoelastic properties
of composite viscoelastic hydrogels of the invention as measured
using a dynamic oscillatory test method as outlined above. Various
biphasic viscoelastic hydrogels comprising hyaluronic acid and
crosslinked gelatin powder were fabricated using the method
outlined above. A uniform gelatin powder particle size (Pm) of 125
.mu.m<Pm<300 .mu.m was used to fabricate the hydrogels. The
following hyaluronic acid concentrations were employed: 5 mg/ml,
7.5 mg/ml, 10 mg/ml, 15 mg/ml, 20 mg/ml. Each concentration was
combined with one of the following gelatin concentrations: 120
mg/ml, 140 mg/ml, 160 mg/ml, 180 mg/ml to create a matrix of 20
biphasic hydrogels for rheological analysis. The graphs present the
Storage Modulus (G') (FIG. 4A), Loss Modulus (FIG. 4B), and Tan
Delta (G'/G'') (FIG. 4C) for the biphasic hydrogels taken at a
frequency of 1 Hz. The general trends are an increase in Storage
(G') and Loss (G'') moduli for increasing concentrations of both
gelatin and hyaluronic acid. We also see a trend towards higher Tan
.delta. (G'/G'') for increasing hyaluronic acid concentration.
[0204] Example 5: FIG. 5 presents results of compression testing to
determine the compression stiffness of the composite viscoelastic
hydrogels of the invention. Biphasic hydrogels, formed in a
phosphate buffered saline carrier of varying hyaluronic acid and
gelatin concentrations were prepared as previously described.
Hyaluronic acid with a molecular weight of 1.8-2 MDa was
employed.
[0205] Hydrogels were formed into 4 mm thick sheet and using a core
biopsy punch, 6 mm diameter cylinders with a height of 5 mm were
created. Compression tests of the cylindrical samples of hydrogel
and lung tissue were performed using a Zwick universal testing
machine with a 5N load cell at a strain rate of 3 mm/min. The
results show an increase in compressive stiffness with increasing
concentrations of gelatin and hyaluronic acid. A large spike in
compressive stiffness is seen at DHT gelatin concentrations in
excess of 140 mg/ml. Similar tests were performed with samples of
parenchymal lung tissue that yielded a compressive stiffness of
approx. 800 Pa.
[0206] Example 6: FIG. 6 shows the strain sweep data for composite
viscoelastic hydrogels of the invention as measured using a dynamic
oscillatory test. Biphasic hydrogels, formed in a phosphate
buffered saline carrier of varying hyaluronic acid and gelatin
concentrations were prepared as previously described. The test
rheometer used to measure the rheological properties of the
hydrogels was a model MCR102 by Anton Paar GmbH. Tests were
conducted under strain control, with a 25 mm flat plate geometry, a
gap of 1 mm, an analysis temperature of 25.degree. C., a frequency
of 1 Hz and over a strain range of 0.01-100%. The graph
demonstrates how all gels exhibit shear thinning behaviour and all
gels demonstrate a storage modulus G' of less than 200 Pa at 100%
strain.
[0207] Example 7: FIG. 7 demonstrates the effect of terminal steam
sterilization on the dynamic viscoelastic properties of the
composite viscoelastic hydrogel of the invention. Biphasic
hydrogels were formulated consisting of varying concentrations of
dehydrothermally crosslinked gelatin: 160 mg/ml, 170 mg/ml, 200
mg/ml. The gelatin was combined with hyaluronic acid in phosphate
buffered saline as previously described. The gels were loaded into
1 ml glass syringes and steam autoclaved at 128.degree. C. for 15
mins. The rheological characteristics of the hydrogels were
determined both pre and post-sterilization The test rheometer used
to measure the rheological properties of the hydrogels was a model
MCR102 by Anton Paar GmbH. Dynamic oscillatory tests were conducted
under stress control, with a 25 mm flat plate geometry, a gap of 1
mm, an analysis temperature of 25.degree. C. and over the frequency
range 0.1-10 Hz. The charts present the Storage Modulus (G') (FIG.
7A), Loss Modulus (FIG. 7B), and Tan Delta (G'/G'') (FIG. 7C) for
the biphasic hydrogels taken at a frequency of 1 Hz before and
after sterilization. FIG. 7D represents the yield stress data.
Yield stress was measured by exposing a sample to a steady stress
ramp until the sample began to undergo plastic deformation; the
stress at which plastic deformation begins was taken to be the
yield stress. The shear rate was ramped logarithmically and had an
initial value 0.0011/s and final value of 2001/s.
[0208] Example 8: FIG. 8 presents viscosity data for composite
viscoelastic hydrogels of the invention produced using different
dehydrothermal treatment conditions. Milled gelatin powder was
heated under vacuum conditions (0.001 mbar) for 24 hours at three
different temperatures (130.degree. C., 140.degree. C. and
150.degree. C.) to induce dehydrothermal (DHT) crosslinking. This
resulted in different levels of crosslinking for the powders, with
increased temperature leading to a higher degree of crosslinking.
Equal concentrations of gelatin powder from each batch were
combined with hyaluronic acid (8 mg/ml) in phosphate buffered
saline. Hydrogels were then evaluated rheologically. The test
rheometer used to measure the rheological properties of the
hydrogels was a model MCR102 by Anton Paar GmbH. Dynamic
oscillatory tests were conducted under shear rate control, with a
25 mm flat plate geometry, a gap of 1 mm, an analysis temperature
of 25.degree. C. The viscosity was recorded at a shear rate of
0.0011/s. The results show a reduction in viscosity with increasing
dehydrothermal treatment. This is likely owning to the fact that
the higher level of crosslinking reduces the ability of the
hydrogel particles to swell, therefore more residual fluid is left
in the sample.
[0209] Example 9: FIG. 9A to FIG. 9C presents partial CT-scans of
the cross-sections of a composite viscoelastic hydrogel plug used
to prevent pneumothorax during transthoracic lung biopsy procedure
in a porcine model. The bi-phasic hydrogel is formed by combining
hyaluronic acid and dehydrothermally crosslinked gelatin as
described previously. The procedure used to deliver the biphasic
hydrogel plug is similar to the one described in FIG. 2A to FIG.
2E. At the start of the biopsy procedure a hydrogel plug of volume
0.3 ml is injected to just below the surface (visceral pleura) of
the lung. The hydrogel material is tissue opposing and does not
infiltrate the lung tissue. It therefore creates a plug surrounding
the delivery needle and coaxial cannula. The hydrogel plug thereby
creates an air-tight seal between the cannula and the surface of
the lung. After the biopsy procedure the needles are removed, the
biphasic hydrogel self-heals by flowing back to occupy the space
taken up by the needle. This further prevents any air from escaping
the lung. FIG. 9A shows a partial CT scan of the lung and in
particular a cross-section of the hydrogel plug immediately after
the lung biopsy procedure. The black arrow indicates the hydrogel
plug. FIG. 9B shows a partial CT scan of the lung and the same
hydrogel plug after 13 days. FIG. 9C shows a partial CT scan of the
lung at the same location of the hydrogel plug after 30 days. It is
event in FIG. 9C that the hydrogel plug has fully degraded leaving
healthy lung tissue in it's place.
[0210] Example 10: FIG. 10 represents the injection force employed
to delivery various composite hydrogels of the invention through an
18G delivery needle of 150 mm in length. The hydrogels were
prepared as described previously with varying concentrations of DHT
treated gelatin powder and 8 mg/ml hyaluronic acid in phosphate
buffered saline. To perform the test, hydrogels were loaded into 1
ml glass syringes and the syringes were mounted onto a Zwick
universal testing machine with a 200N load cell. An 18G delivery
needle of length 150 mm was attached to the 1 ml glass syringe via
a luer lock connector. A compression test was performed at a rate
of 120 mm/min so that the load cell depressed the syringe plunger
and the hydrogel was injected through the delivery needle. Results
show that the maximum injection force achieved for all gels lay
between 10-25N. This is believed to be below the clinically
acceptable level of 45N.
[0211] Example 10: In order to determine the DHT gelatin powder
particle size a laser diffraction technique was employed (Malvern
Instruments Master Sizer 3000) using the following methods. Data
was analysed using the respective software (Mastersizer 3000,
v.0.1). An even dispersion of dry particles was achieved using the
Aero S dispersion unit. The refractive index of gelatin was set to
1.543 while the absorption index was set to 0.010. Data was
recorded when obscuration levels were between 0.10% and 10.00%. An
even dispersion of wet particles was achieved using the Hydro LV
dispersion unit. The refractive index of PBS and gelatin was set to
1.33 and 1.543, respectively. The absorption index of the dispersed
gelatin particles was set to 0.01. For measurement of wet particle
size distribution, gelatin was hydrated at room temperature for 24
hours prior to testing. Particles were added to approximately 125
mL of PBS in the dispersion unit until an obscuration of
approximately 10% was obtained.
[0212] DHT particles were prepared by firstly grinding 300bloom
type A porcine gelatin powder. This powder was sieved with a mesh
size of 1125 .mu.m and subjected to a DHT treatment of 140.degree.
C. for 24 hours. Laser diffraction revealed the following dry
particle size distribution: 10.sup.th (Dv(10)), 50.sup.th (Dv(50)),
and 90th (Dv(90)) percentiles of 19, 58, and 123 microns
respectively. Gelatin particles were then hydrated for 24 hours
prior and wet particle size distribution revealed the following
results: Dv(10), Dv(50) and Dv(90) was 27, 107 and 229 microns
respectively. This indicates a volume swelling factor of
approximately 6.2 for the particles post hydration.
[0213] In a preferred embodiment, the composite viscoelastic
hydrogel is capable of preventing pneumothorax during procedures
requiring transthoracic needle access by being injected just below
the visceral pleura of the lung and by having the following
properties: [0214] 1. The hydrogel has low enough viscosity under
shear stress exerted by the syringe to enable the hydrogel to be
injected to the target site through a needle, catheter or other
luminal device. [0215] 2. Once exiting the needle the hydrogel
undergoes a rapid thixotropic recovery to a stiffness sufficient to
prevent infiltration of lung tissue. [0216] 3. Once the needle has
been removed, an element of viscous flow enables the gel to flow
back to form a single entity. The gel flows back to fill the void
left by the needle in the lung tissue and in the visceral pleura.
It may achieve this by having a sufficient flowable nature which is
preferably dependent on having a high tan .delta.. [0217] 4. The
gel has sufficient rigidity and storage modulus (G') that it is not
prematurely ejected from the lung and remains at the delivery site
until healing has occurred.
EQUIVALENTS
[0218] The foregoing description details presently preferred
embodiments of the present invention. Numerous modifications and
variations in practice thereof are expected to occur to those
skilled in the art upon consideration of these descriptions. Those
modifications and variations are intended to be encompassed within
the claims appended hereto.
* * * * *
References