U.S. patent application number 17/312423 was filed with the patent office on 2022-02-17 for hydrogel for in-vivo directional release of medication.
The applicant listed for this patent is BO-IP B.V., SENTRYX B.V., UMC Utrecht Holding B.V.. Invention is credited to Bas Jeroen OOSTERMAN, Susanna PILUSO, Jasper Gerard STEVERINK, Joannes Jacobus VERLAAN.
Application Number | 20220047856 17/312423 |
Document ID | / |
Family ID | |
Filed Date | 2022-02-17 |
United States Patent
Application |
20220047856 |
Kind Code |
A1 |
STEVERINK; Jasper Gerard ;
et al. |
February 17, 2022 |
HYDROGEL FOR IN-VIVO DIRECTIONAL RELEASE OF MEDICATION
Abstract
The invention provides a hydrogel for in-vivo release of
medication comprising at least one medication, wherein the surface
of the hydrogel comprises a coating such that the surface has one
or more sub-surfaces with permeability that is at least 2.times.
higher than the average permeability of the entire surface, wherein
the hydrogel has an elastic modulus of between 50 and 1000 kPa.
Inventors: |
STEVERINK; Jasper Gerard;
(Austerlitz, NL) ; PILUSO; Susanna; (Utrecht,
NL) ; VERLAAN; Joannes Jacobus; (Zeist, NL) ;
OOSTERMAN; Bas Jeroen; (Zeist, NL) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
UMC Utrecht Holding B.V.
BO-IP B.V.
SENTRYX B.V. |
Utrecht
Zeist
Austerlitz |
|
NL
NL
NL |
|
|
Appl. No.: |
17/312423 |
Filed: |
June 11, 2019 |
PCT Filed: |
June 11, 2019 |
PCT NO: |
PCT/NL2019/050352 |
371 Date: |
June 10, 2021 |
International
Class: |
A61M 31/00 20060101
A61M031/00; A61L 27/52 20060101 A61L027/52; A61L 27/22 20060101
A61L027/22; A61L 27/20 20060101 A61L027/20 |
Foreign Application Data
Date |
Code |
Application Number |
Dec 12, 2018 |
NL |
PCT/NL2018/050832 |
Claims
1. A hydrogel for in-vivo release of medication comprising at least
one medication in the form of a small molecule, wherein the
hydrogel has a surface and the surface of the hydrogel comprises a
coating that is composed of a material that is less permeable to
the medication than the material of the hydrogel itself such that
the surface has one or more sub-surfaces with permeability that is
at least 2.times. higher than the average permeability of the
entire surface, wherein the hydrogel has an elastic modulus of
between 50 and 1000 kPa.
2. The hydrogel of claim 1, the hydrogel has an elastic modulus of
between 100 and 600 kPa.
3. The hydrogel of claim 1, having a degree of swelling in the
range of 2-20 calculated as (swollen weight-dry weight)/dry
weight.
4. The hydrogel of claim 1, comprising a cross-linked
biopolymer.
5. The hydrogel of claim 4, wherein the cross-linked biopolymer is
a protein-based and/or polysaccharide-based polymer.
6. The hydrogel of claim 1, wherein between 10 and 90% of the
surface is covered by the coating.
7. The hydrogel of claim 1, wherein the coating has a thickness
between 10 nm to 200 .mu.m.
8. The hydrogel of claim 1, wherein the coating is based on a
biodegradable polymer.
9. A method for the preparation of the hydrogel of claim 1, wherein
the coating is comprised of a precursor material and some of the
precursor material to the coating is allowed to partially diffuse
into the hydrogel.
10. The hydrogel according to claim 1 for use in the treatment of
musculoskeletal disorders.
11. The hydrogel of claim 2, having a degree of swelling in the
range of 2-20 calculated as (swollen weight-dry weight)/dry
weight.
12. The hydrogel of claim 2, comprising a cross-linked
biopolymer.
13. The hydrogel of claim 3, comprising a cross-linked
biopolymer.
14. The hydrogel of claim 11, comprising a cross-linked
biopolymer.
15. The hydrogel of claim 12, wherein the cross-linked biopolymer
is a protein-based and/or polysaccharide-based polymer selected
from the group consisting of hyaluronic acid, chitosan, cellulose,
gelatin, and combinations thereof.
16. The hydrogel of claim 13, wherein the cross-linked biopolymer
is a protein-based and/or polysaccharide-based polymer selected
from the group consisting of hyaluronic acid, chitosan, cellulose,
gelatin, and combinations thereof.
17. The hydrogel of claim 4, wherein the cross-linked biopolymer is
gelatin.
18. The hydrogel of claim 1, wherein between 20 and 80% of the
surface is covered by the coating, optionally wherein between 30
and 70% of the surface is covered by the coating.
19. The hydrogel of claim 1, wherein the coating is selected from
the group consisting of PLGA, PCL, gelatin, alginate, and
combinations thereof.
20. The hydrogel according to claim 1 for use in the treatment of
musculoskeletal disorders for treatment of infection, inflammation,
malignant processes, growth disorders, degenerative disorders,
treatment of pain arising from said disorders, or treatment of pain
arising from surgical treatment of said disorders.
Description
TECHNICAL FIELD
[0001] The present invention relates to a hydrogel for in-vivo
directional release of medication. In particular it concerns a
controlled and local release of medication. More in particular, the
present invention relates to a hydrogel for close contact to organs
and skeletal structures.
BACKGROUND ART
[0002] Hydrogels are three-dimensional, physically or chemically
cross-linked networks of water-soluble polymers. Their hydrophilic
nature, water content similar to living tissue and elasticity, make
them excellent candidates for biomedical applications. There is
therefore quite some prior art on biodegradable hydrogels that are
designed to release medication in the (human or animal) body in a
sustained way.
[0003] For instance, in the Journal of Advanced Research, volume 8,
Issue 3, May 2017, pages 217-233, a thorough review by E. A. Kamoun
et al may be found on hydrogels and their medical application. As
indicated in the introduction of this article, a further overview
may be found in European Polymer Journal, volume 65, April 2015,
pages 252-267 by E. Calo et al, "Biomedical applications of
hydrogels: A review of patents and commercial products".
[0004] Q. Feng et al describes "Mechanically resilient, injectable,
and bioadhesive supramolecular gelatin hydrogels crosslinked by
weak host-guest interactions assist cell infiltration and in situ
tissue regeneration" in Biomaterials, Volume 101, September 2016,
Pages 217-228.
[0005] In RSC Adv., 2017, 7, 34053, T. T. H. Thi et al describe
injectable hydrogels as a novel platform for the release of
hydrophobic drugs. An additional Schiff base reaction was
introduced into a phenol-phenol crosslinked gelatin hydrogel to
increase adhesiveness. .beta.-cyclodextrin possessing a hydrophobic
cavity and oxidized to present aldehyde groups (hereinafter
"o.beta.-CD") was grafted to the gelatin backbone via Schiff base
reaction, with the cavity providing encapsulation for hydrophobic
drugs. Simply blending gelatin-tyramine (hereinafter "GTA") and
o.beta.-CD in the presence of horseradish peroxidase and hydrogen
peroxide (hereinafter "HRP/H.sub.2O.sub.2") rapidly and
controllably formed GTA-o.beta.-CD hydrogels in situ. The optimal
composition of GTA-o.beta.-CD hydrogels was found to be 5 wt % GTA
with 1 wt % o.beta.-CD. Their elastic modulus and degradation rate
were 1.8- and 1.5-fold higher than those of GTA hydrogels owing to
additional imine bonds. Hydrophobic drugs (e.g., dexamethasone and
curcumin) could be dissolved homogeneously in GTA-o.beta.-CD
matrices with greater loading efficiencies than in GTA matrices. An
in vitro test of cell viability using human dermal fibroblasts
demonstrated that GTA-o.beta.-CD hydrogels were cytocompatible. In
summary, dual-functional injectable GTA-o.beta.-CD hydrogels can be
used as a promising platform to improve tissue adhesion and
hydrophobic drug delivery.
[0006] Important factors to consider during the design of these
hydrogels include 1) duration of delivery, and 2) location of
delivery with respect to its working mechanism. For example, for
effective local pain relief it is essential that an anaesthetic is
delivered and remains in situ for a period in close proximity to
the origin of pain. The problem of sustained release is
particularly challenging for small molecules, such as Bupivacaine
(hereafter "Bupi").
[0007] Bupi is a very effective and relatively inexpensive local
anaesthetic. However, the duration of its effect is limited to
approx. 8 hours. Increasing the dose or concentration of
conventional bupivacaine solutions to obtain prolonged durations of
effect can lead to both systemic and local toxicity, cf., Gitman M,
Barrington M J "Local Anesthetic Systemic Toxicity: A Review of
Recent Case Reports and Registries" in Regional Anesthesia &
Pain Medicine 2018; 43:124-130. Cardio- and central nervous system
toxicity are well-known systemic toxic effect of bupivacaine. It is
therefore of interest to find a way of releasing Bupi locally and
in a delayed fashion, whereby it may work longer and with a
decreased incidence of local and systemic cytotoxicity compared to
conventional bupivacaine applications such as local bolus
injection.
[0008] In the yet unpublished NL patent application 2020071 by the
present applicant a deformable body and combination of such
deformable body and a surgical screw element is described. The
deformable body may be made of a visco-elastic material, a
degradable felt material, a sponge-like material, a gelatine
material, a gel, in particular a hydrogel, a polymer or any
combination thereof. The deformable body may comprise an
anaesthetic and/or another pharmaceutical compound. It has a
surface through which the anaesthetic may be released, for example,
the bone contact surface. The release in the surrounding area may
be avoided, by use of a further substantially non-pervious wall of
the deformable body.
[0009] As further improvement on the deformable body, the present
inventors set out to design a biocompatible, biodegradable hydrogel
with controlled, sustained and directional release of medication.
Moreover, the inventors set out to design a hydrogel that is
versatile and easy to produce on a large scale, is easy to
cross-link and can be cross-linked in a controlled manner to
produce a hydrogel that is both flexible and strong. In this regard
it should be understood that the hydrogel must be both sufficiently
flexible and strong as to allow it to be implanted and to withstand
local circumstances and forces so as to stay at the location of
implantation for sufficient time to release the medication and not
break or otherwise be damaged. This means that the hydrogel can
adapt to a shape of a surface of a skeletal structure against which
it is pressed, whereby intimate contact with the outer bone surface
of the skeletal structure is achieved. Typically this requires a
hydrogel with an elastic/compressive modulus of between 50 and 1000
kPa, more preferably between 100 and 600 kPa.
SUMMARY OF THE INVENTION
[0010] The present invention provides a hydrogel for in-vivo
release of medication comprising at least one medication, wherein
the surface of the hydrogel comprises a coating such that the
surface has one or more sub-surfaces with permeability that is at
least 2.times. higher than the average permeability of the entire
surface, wherein the hydrogel has an elastic modulus of between 50
and 1000 kPa.
DRAWINGS
[0011] FIG. 1 is a series of images of a hydrogel having its top
part coated and containing methylthioninium chloride (methylene
blue). As can be seen, methylthioninium chloride (which is both a
medication and dye) is only released in the opposite direction.
DETAILED DESCRIPTION OF THE INVENTION
[0012] Hydrogels may be synthesized by cross-linking water-soluble
polymers. Water-soluble polymers such as poly(acrylic acid),
poly(vinyl alcohol), poly(vinylpyrrolidone), poly(ethylene glycol),
polyacrylamide and polysaccharides (e.g. hyaluronic acid) are the
most common systems used to form hydrogels. These water-soluble
polymers are non-toxic and widely used in various pharmaceutical
and biomedical applications. Although there are many different
hydrogels, the present invention focusses on medical hydrogels that
are biocompatible and can be implanted and used in-vivo. Moreover,
they must be biodegradable. For instance, protein-based and/or
polysaccharide-based polymers may be used, such as, hyaluronic
acid, chitosan, and cellulose. Preferably, the hydrogel is based on
gelatin. In addition to, or instead of the protein-based and/or
polysaccharide based polymers, the hydrogel may also comprise other
non-toxic water-soluble synthetic or natural polymers. The other
polymers may compose up to 50% by weight of the entire polymer
content. Given its availability, biocompatibility and cost, the use
of gelatin as sole polymer component is preferred. Of particular
interest is a hydrogel based on gelatin that is functionalized with
a cyclodextrin.
[0013] Although hydrogels for release in-vivo of medication are
known, the present inventors found that existing hydrogels could be
improved in terms of their directional release. As a result, the
new hydrogels of the present invention can be implanted and fixated
to specific locations where medication, in particular to achieve
pain relief, is required. This may be a hydrogel in the form of
e.g. deformable body, whereby the hydrogel conforms to the shape of
a skeletal structure or surgical implant or even organ to which it
is fixated. Of relevance in this respect is that a hydrogel with a
specific elastic modulus in the aforementioned range is used.
Moreover, the hydrogel preferably has a degree of swelling in the
range of 2-20, preferably in the range of 2-6, calculated as
swollen weight (at equilibrium swelling)-dry weight/dry weight.
[0014] The direction of release of medication is achieved by partly
covering the surface of the hydrogel with a coating. As a result,
the hydrogel will have a sub-surface or sub-surfaces with little or
no coating and hence unrestricted permeability of the medication,
and a subsurface or surfaces with coating and therefore a reduced
permeability for the medication. Preferably the nature and
thickness of the coating is selected such as that the permeability
at the desired contact surface, e.g., the bone or organ contact
surface is at least 2.times. higher than the average permeability
of the entire surface. Having the implanted hydrogel affixed
adjacent to the body part that is to be treated, and moreover with
the uncovered surface of the hydrogel adjacent to the body part
that is to be treated, release of medication in other directions is
reduced or even avoided. This has the advantage of reduced-side
effects and the possibility to work with lower concentrations of
medication or, alternatively, with a longer working time due to a
slower release of the regular amount of medication.
[0015] The coating may be composed of the material of the hydrogel,
provided that it contains no medication and is sufficiently thick.
Suitably it is between 10 nm and 200 .mu.m thick. Preferably,
however, the coating is composed of a material that is less
permeable to the medication than the material of the hydrogel
itself. The coating may be flexible or shell-like. Similar to the
hydrogel, the coating must be composed of biocompatible
biopolymers. The biodegradability may be the same or prolonged
compared to the hydrogel. Suitable materials include, but are not
limited to polycaprolactone (hereinafter "PCL"),
poly(lactic-co-glycolic acid) (hereinafter "PLGA"), gelatin, or
alginate. The permeability of the coating may be adjusted, such
that even very small molecules cannot get through. Moreover, the
coating can be made hydrophobic, or hydrophilic, depending on its
intended use.
[0016] The hydrogel may take any particular shape. In a co-pending
application, the use of a hydrogel as carrier for local release of
medication in the form of a ring is described (PCT/NL2018/050832,
incorporated herein by reference) where it is used in combination
with a screw. In another co-pending application the use of a
hydrogel as carrier for local release of medication in the form of
a sleeve, e.g. for a joint prosthesis is described (NL2023208,
incorporated herein by reference). The hydrogel may also be shaped
in the form of a (board) thumb pin for attachment to bone or any
other solid tissue. Finally, the hydrogel may also be shaped to
provide a tight fit in crevices in organs and similar body
structures. In each of these embodiments, the hydrogel is coated
such as to ensure that those parts of its surface that are not in
contact with the body part that is to be treated by direct release
are covered by the coating.
[0017] The coatings may be applied onto the hydrogel by any common
coating process, including dip coating, brush coating, spray
coating and the like. Alternatively, the entire surface of the
hydrogel may be coated, whereas the relevant sub-surfaces intended
for contact with the body part that is to be treated are freed from
coating. Moreover, the coating may be formed and shaped first, as a
shell, whereupon the hydrogel in introduced e.g., as an
non-crosslinked solution. In this case the shell of coating acts as
a mould during the cross-linking and formation of the hydrogel.
Alternative methods include overmolding and the like.
[0018] Using a coating material and method that allows some of the
precursor material to the coating to partially diffuse into the
hydrogel may be particularly beneficial, in particular if this
material is water-soluble. After
polymerization/crosslinking/setting, the coating will be physically
entangled with hydrogel directly underneath the interface, ensuring
a good bond. This method is of particular interest, as it reduces
chances of coating material breaking off, which is detrimental as
it affects the directional release, but which is also detrimental
as it might cause migration of particles of coating that may create
their own problems.
[0019] Preferably between 10 and 90% of the surface of the hydrogel
is covered by a coating. For instance, between 20 and 80% of the
surface is covered by a coating, more preferably between 30 and 70%
of the surface is covered by a coating.
[0020] The present hydrogel is particularly suitable for treatment
of musculoskeletal disorders. These disorders include infection,
inflammation, malignant processes, growth disorders, degenerative
disorders or treatment of pain arising from (surgical treatment of)
these disorders.
[0021] In addition to the medication one or more further
ingredients may be included, preferably further ingredients
selected from co-medication, glycerol and other co-solvents,
colorants, and buffers.
[0022] Methods for making the feedstock for the hydrogel are known.
Thus, it is known to functionalize gelatin and related biopolymers
with tyramine. See Thi et al, 2017 RSC Adv, which has been cited
above, and which is included herein by reference. Of importance,
but common in the field of medical application is to remove all
forms of contamination. By way of example, the hydrogel may be
prepared by the following method: [0023] 1. Solutions of a suitable
cross-linking water-soluble (bio)polymer(s), cross-linker and
medication are prepared. [0024] 2. Solutions of biopolymer(s) and
cross-linker are mixed at pre-determined concentrations to achieve
a cross-linked hydrogel with an elastic modulus in the range of 50
to 1000 kPa. [0025] 3. The obtained hydrogel is then submerged in a
solution of medication to allow for diffusion of the medication
into the hydrogel. Glycerol or similar co-solvent can be added to
the medication solution. Glycerol then also diffuses into the
hydrogel where it acts as a plasticizer, providing additional
robustness and flexibility to the hydrogel. Alternatively, the drug
(e.g. in a nano-/microparticle formulation) can be mixed in with
the polymer solution prior to crosslinking. [0026] 4. The gel is
then dried. [0027] 5. Next, the hydrogel is coated in part, e.g.
with a solution of a biopolymer with a different permeability for
the medication compared to the hydrogel, to ensure directional
release of the encapsulated medication. The coating may also
enhance the mechanical properties of the hydrogel. Alternatively,
it is also possible to form a shell of the coating in a pre-defined
shape, and introduce the solution of step 2, together with the
medication, into this shell, whereby the coating acts as a mould
for the hydrogel.
EXAMPLES
Materials
[0028] Gelatin (porcine skin, type A, 300 g bloom strength),
1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC),
N-hydroxysuccinimide (NHS), tyramine hydrochloride,
2-morpholinoethanesulfonic acid monohydrate (MES), sodium
persulfate (SPS), sodium periodate, 3-cyclodextrin,
phosphate-buffered saline (PBS), riboflavin (RB), ethylene glycol
and glycerol were purchased from Sigma-Aldrich. Cellulose dialysis
membranes (Spectra/Por.TM., 0.5 kDa; 12 kDa molecular weight
cut-off) were purchased from Spectrum Laboratories. Bupivacaine was
obtained from Siegfried, Switzerland.
Synthesis of Gelatin-Tyramine (GTA)
[0029] Gelatin type A (5 g) was dissolved in a 50 mM MES buffer
(300 ml) at 50.degree. C. After dissolution of the gelatin, EDC
(13.7 mmol), NHS (6.85 mmol) and tyramine (15 mmol) were added to
the gelatin solution. The reaction mixture was left to react for 24
h at 40.degree. C. with stirring. After 24 the mixture was dialyzed
against water for 72 h and the product was then obtained by
lyophilization.
Tyramine Content Measurement
[0030] The degree of functionalization of gelatin was determined by
measuring the absorbance of the polymer solution (0.1%, w/v) at 275
nm and calculated from a calibration curve obtained by measuring
the absorbance of known percentages of tyramine in distilled
water.
Oxidation of .beta.-Cyclodextrin
[0031] Oxidized .beta.-cyclodextrin was prepared by reaction with
sodium periodate. Briefly, .beta.-cyclodextrin (5 g) was dispersed
in distilled water followed by addition of sodium periodate (3.77
g) and stirred at room temperature in the dark, overnight. The
reaction was terminated by the addition of ethylene glycol. The
mixture was dialyzed against deionized water using a dialysis
membrane with an MWCO of 500 Da (Spectrum Labs) for 3 days and the
product was collected by lyophilization. The degree of oxidation
was determined by .sup.1H NMR, using either deuterated dimethyl
sulfoxide (DMSO-d6) or deuterium oxide (D2O) as solvent. Whereas
.beta.-cyclodextrin has a ratio of protons at 4.8-4.9 ppm versus 4
ppm of about 2.04, progress of the reaction can be seen by a change
in the ratio, to about 1.49.
Fabrication of GTA/.beta.-Cyclodextrin Hydrogels
[0032] Prior to hydrogel crosslinking, solutions of GTA, op-CD, SPS
and Riboflavin were prepared. Unless indicated otherwise, GTA had a
degree of functionalization of 10-25%, whereas o.beta.-CD with an
oxidation degree of the secondary hydroxyl groups of 15-30% was
used. These solutions were mixed so that final concentrations of 20
wt % GTA, 0-10 wt % o.beta.-CD, 20 mM SPS and 2 mM Riboflavin were
obtained. The obtained solution was exposed to visible light for 30
minutes to enable hydrogel formation. The cross-linked hydrogel had
a degree of swelling of 3-6. Moreover, it had an elastic modulus of
100-600 kPa.
[0033] The obtained hydrogel was then submerged overnight in a
bupivacaine solution with a concentration of bupivacaine of 50
mg/mL to allow for diffusion of bupivacaine into the gel. The
bupivacaine solution contained a concentration glycerol of 30 vol
%. As a result, the concentration of bupivacaine in the hydrogel
was 50 mg/mL (.+-.20).
[0034] Next, the hydrogel was coated with a coating solution
comprising 10% PCL in dichloromethane (DCM) In this case the
hydrogel was dipped into the solution for a number of times to
achieve a coating of about 180 .mu.m. The coating was found to
provide additional strength to the hydrogel.
Drug Loading and In Vitro Drug Release Assay
[0035] For the investigation of drug release properties, the
obtained hydrogels were loaded with bupivacaine by immersion in an
aqueous solution of bupivacaine at 50 mg/mL for 24 hours. The
bupivacaine solution contained a concentration glycerol of 30% vol.
As a result, the concentration of bupivacaine in the hydrogel was
.+-.50 mg/mL (.+-.20).
[0036] The release of bupivacaine from the hydrogels was measured
by placing the hydrogels in a vial containing 1 mL of 0.1M citrate
buffer, pH 6 at 37.degree. C.
[0037] At predetermined time points, aliquots of 100 uL samples
were taken from the release solution and replaced with fresh
buffer. The samples were diluted 1:10. Bupivacaine release was
determined by UPLC using ammonium formate (10 mM, pH 2.4) and a
mixture of acetonitrile/water/formic acid (96:5:0.2, v:v:v) as
mobile phase. This control experiment proves that the hydrogel may
be used for sustained release of medication
Drug Loading and In Vitro Drug Release Assay
[0038] For the investigation of directional release, the obtained
hydrogels were now loaded with methylthioninium chloride. A PCL
shell was acquired by dip-coating of a metal mould. The mould was
dipped twice in 10% PCL solution to obtain a 180 um thick film. A
photocrosslinkable pre-gel solution was then prepared,
methylthioninium chloride was added by mixing a 1 wt % solution in
the pre-gel solution to obtain a final concentration of 0.1 wt %
methylthioninium chloride in the hydrogel. The gel was then
cross-linked on top of the PCL film using exposure to a visible
light-source.
[0039] Release of methylthioninium chloride from the gel was
simulated in a 3% alginate gel, cross-linked with calcium chloride
to obtain a tissue-like consistency. In the images, FIG. 1, the PCL
film is on top of the gel. As shown, release was only visible in
the non-PCL-covered direction. Hydrogels were positioned vertically
to eliminate any effect of gravity on the direction of release.
This experiment proofs that the hydrogel with coating may be used
for sustained directional release of medication.
* * * * *