Note: Descriptions are shown in the official language in which they were submitted.
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Improved Spacer Membrane for an Enzymatic in-vivo Sensor
Description
The present invention relates to an electrode system for measuring the
concentration of an analyte under in-vivo conditions, comprising an electrode
with immobilized enzyme molecules and an improved diffusion barrier that
controls diffusion of the analyte from body fluid surrounding the electrode
system to the enzyme molecules.
Further, the present invention relates to an electrode system for measuring
the concentration of an analyte under in-vivo conditions, comprising an
electrode with immobilized enzyme molecules, optionally a diffusion barrier
that controls diffusion of the analyte from the exterior of the electrode
system
to the enzyme molecule and an improved spacer membrane which forms at
least a portion of the outer layer of the electrode system.
Sensors with implantable or insertable electrode systems facilitate
measurements of physiologically significant analytes such as, for example,
lactate or glucose in a patient's body. The working electrodes of systems of
this type have electrically conductive enzyme layers in which enzyme
molecules are bound which release charge carriers by catalytic conversion of
analyte molecules. In the process, an electrical current is generated as
measuring signal whose amplitude correlates to the analyte concentration.
Such electrode systems are e.g. known from WO 2007/147475 and WO
2010/028708.
The working electrodes of the electrode system are provided with a diffusion
barrier that controls the diffusion of the analyte to be determined from the
body fluid or tissue surrounding the electrode system to the enzyme
molecules that are immobilized in the enzyme layer. According to WO
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2010/028708, the diffusion barrier of the electrode system is a solid solution
of at least two different polymers, preferably of acrylates. The polymers may
be copolymers, e.g. copolymers of methyl methacrylate and hydroxyethyl
methacrylate or copolymers of butyl methacrylate and hydroxyethyl
methacrylate.
WO 2007/147475 discloses a diffusion barrier made from a polymer having a
zwitterionic structure. An example of such a polymer is poly(2-
methacryloyloxyethyl phosphorylcholine-co-n-butylmethacrylate). The
zwifterionic polymer may be mixed with another polymer, for example
polyurethane.
The use of polymer or copolymer mixtures, however, has drawbacks in that
the preparation of the mixture and its application to the sensor is tedious
and
potentially problematic. Usually, the polymers to be mixed are individually
dissolved and the resulting solutions are thereafter mixed in the desired
ratio.
This, however, may result in precipitation of one of the components and
consequently in workability problems, e.g. in a spraying process. Increased
difficulties occur when the mixture comprises a polymer with ionic
characteristics, i.e. when one of the polymers to be mixed comprises a
monomer having anionic or cationic groups. The presence of such charged
groups, however, has a strong effect on the solubility, making it difficult to
find a solvent suitable for both the charged polymer and an uncharged
polymer.
WO 2006/058779 discloses an enzyme-based sensor with a combined
diffusion and enzyme layer comprising at least one polymer material, and
particles carry an enzyme, wherein the particles are dispersed in the polymer
material. The polymer may comprise hydrophilic as well as hydrophobic
polymer chain sequences, for example, the polymer may be a high or low
water uptake polyether-polyurethane copolymer. The use of block
copolymers having at least one hydrophilic block and at least one
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hydrophobic block as a diffusion layer is not disclosed.
EP-A-2 163 190 describes an electrode system for the measurement of an
analyte concentration in-vivo comprising a counterelectrode with an electric
conductor, and a working electrode with an electric conductor on which an
enzyme layer comprising immobilized enzyme molecules is localized. A
diffusion barrier controls the diffusion of the analyte from surrounding body
fluids to the enzyme molecules. The diffusion barrier may comprise
hydrophilized polyurethanes obtainable by polycondensation of 4,4'-
methylene-bis-(cyclohexylisocyanate) and diol mixtures which may be
polyethyleneglycol and polypropyleneglycol. The hydrophilic polyurethane
layer may be covered with a spacer, e.g. a copolymer of butyl methacrylate
and 2-methacryloyloxyethyl-phosphoryl choline. The use of block copolymers
having at least one hydrophilic block and at least one hydrophobic block as a
diffusion layer is not disclosed. The use of a hydrophilic copolymer of
(meth)acrylic monomers comprising more than 50 mol-% hydrophilic
monomers is not disclosed either.
It is an object of the present invention to provide a diffusion barrier on an
electrode system of an enzymatic in-vivo sensor which provides desirable
physico-chemical characteristics and which can be manufactured easily.
This object is met by providing a diffusion barrier consisting of a single
block
copolymer having at least one hydrophilic block and at least one hydrophobic
block. The hydrophilic and hydrophobic blocks are covalently linked to each
other. Preferably the blocks are (meth)acrylate polymer blocks.
The block-copolymer based diffusion barrier provides excellent physico-
chemical characteristics as follows:
(i) permeability of the diffusion barrier for the analyte to be determined,
(ii) permeability characteristics of the diffusion barrier which are suitable
for
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the short-term behaviour (wettability) and the long-term behaviour
(sensor drift) of the electrode,
(iii) mechanical flexibility of the diffusion barrier,which allows manufacture
of
in-vivo sensors with extended multiple electrodes;
(iv) efficient incorporation of ionic groups into the diffusion layer, i.e.
the
density of cationic or anionic charges within the polymer can be
efficiently adjusted, this is relevant for repulsion or attraction of
charged analytes, and/or control of cell adhesion, e.g. of monocytes from
the surrounding body fluid or tissue.
A subject-matter of the present invention is an electrode system for
measuring the concentration of an analyte under in-vivo conditions,
comprising an electrode with immobilized enzyme molecules and a diffusion
barrier that controls diffusion of the analyte from the exterior of the
electrode
system to the enzyme molecules, characterized in that the diffusion barrier
comprises a block copolymer having at least one hydrophilic block and at
least one hydrophobic block.
Preferably, the diffusion barrier comprises a single, i.e. only one block
copolymer having at least one hydrophilic block and at least one hydrophobic
block, i.e. further polymers or copolymers are absent. More preferably, the
diffusion barrier consists of a single block copolymer having at least one
hydrophilic block and at least one hydrophobic block.
The electrode system of the present invention is suitable for insertion or
implantation into a body, e.g. a mammalian body such as a human body.
The electrode system is adapted for measuring a desired analyte in body
fluid and/or body tissue, e.g. in the extracellular space (interstitium), in
blood
or lymph vessels or in the transcellular space.
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The inserted or implanted electrode system is suitable for short-term
application, e.g. 3-14 days, or for long-term application, e.g. 6-12 months.
During the insertion or implantation period a desired analyte may be
determined by continuous or discontinuous measurements.
The electrode system of the invention is preferably part of an enzymatic,
non-fluidic (ENF) sensor, wherein enzymatic conversion of the analyte is
determined. Preferably, the sensor comprises a working electrode with
immobilized enzyme molecules for the conversion of the analyte which
results in the generation of an electrical signal. The enzymes may be present
in a layer covering the electrode. Additionally, redox mediators and/or
electro-catalysts as well as conductive particles and pore formers may be
present. This type of electrode is described e.g. in WO 2007/147475.
The area of the working electrode is the sensitive area of the sensor. This
sensitive area is provided with a diffusion barrier that controls diffusion of
the
analyte from the exterior, e.g. body fluid and/or tissue surrounding the
electrode system to the enzyme molecules. The diffusion barrier can, for
example, be a cover layer covering the enzyme layer, i.e. an enzyme-free
layer. However, it is feasible just as well that diffusion-controlling
particles
are incorporated into the enzyme layer to serve as a diffusion barrier. For
example, pores of the enzyme layer can be filled with the polymer which
controls the diffusion of analyte molecules. The thickness of the diffusion
barrier is usually from about 2-20 pm, e.g. from about 2-15 pm, or from about
5-20 pm, particularly from about 5-10 pm or from about 10-15 pm (in dry
state).
The diffusion barrier of the electrode system of the present invention
comprises a block copolymer, preferably a single block copolymer having at
least one hydrophilic block and at least one hydrophobic block. The block
copolymer may comprise an alternating sequence of blocks, i.e. a hydrophilic
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block is linked to a hydrophobic block. The hydrophilic and hydrophobic
blocks are covalently linked to each other within a polymer molecule. The
average molecular weight of the polymer (by weight) is usually from 20-70
kD, particularly from 25-60 kD and more particularly from 30-50 kD. The
molar ratio of the hydrophilic to hydrophobic portions in the block copolymer
is usually in the range from about 75% (hydrophilic) : 25% (hydrophobic) to
about 25% (hydrophilic) : 75% (hydrophobic), in the range from about 65%
(hydrophilic) : 35% (hydrophobic) to about 35% (hydrophilic) : 65%
(hydrophobic) or in the range from about 60% (hydrophilic) : 40%
(hydrophobic) to about 40% (hydrophilic) : 60% (hydrophobic).
A hydrophilic block of the block copolymer consists of at least 90%, at least
95% and particularly completely of hydrophilic monomeric units. It usually
has a length of from 50-400, e.g. from 50-200, or from 150-300 particularly
from 100-150, or from 200-250 monomeric molecules. A hydrophobic block
of the copolymer consists of at least 90%, more particularly at least 95% and
even more particularly completely of hydrophobic monomeric units. It has
usually a length of from 50-300, e.g. from 50-200, or from 150-250,
particularly from 80-150, or from 170-200 monomeric units.
The hydrophilic blocks and/or the hydrophobic blocks preferably consist of
(meth)acrylic-based units. More preferably, both the hydrophilic blocks and
the hydrophobic blocks consist of (meth)acrylic-based monomeric units.
The hydrophilic monomeric units of the hydrophilic block are preferably
selected from hydrophilic (meth)acryl esters, i.e. esters with a polar, i.e.
OH,
OCH3 or 0C2H5 group within the alcohol portion of the ester, hydrophilic
(meth)acrylamides with an amide (NH2) or an N-alkyl- or N,N-dialkylamide
group, wherein the alkyl group comprises 1-3 C-atoms and optionally
hydrophilic groups such as OH, OCH3 or 0C2H5, and suitable (meth)acrylic
units having a charged, e.g. an anionic or cationic group, such as acrylic
acid
(acrylate) or methacrylic acid (methacrylate). Further, combinations of
monomeric units may be employed.
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Specific examples of preferred monomeric units for the hydrophilic block are
selected from:
2-hydroxyethyl acrylate,
2-hydroxyethyl methacrylate (HEMA),
2-methoxyethyl acrylate,
2-methoxyethyl methacrylate,
2-ethoxyethyl acrylate,
2-ethoxyethyl methacrylate,
2- or 3-hydroxypropyl acrylate,
2- or 3-hydroxypropyl methacrylate (2- or 3-HPMA),
2- or 3-methoxypropyl acrylate,
2- or 3-methoxypropyl methacrylate,
2- or 3-ethoxypropyl acrylate,
2- or 3-ethoxypropyl methacrylate,
1- or 2-glycerol acrylate,
1- or 2-glycerol methacrylate,
acrylamide,
methacrylamide,
an N-alkyl- or N,N-dialkyl acrylamide, and
an N-alkyl- or N,N-dialkyl methylamide, wherein alkyl comprises
1-3 C-atoms such as methyl, ethyl or propyl,
acrylic acid (acrylate),
methacrylic acid (methacrylate)
and combinations thereof.
Preferred hydrophilic monomers are 2-hydroxyethyl methacrylate (HEMA)
and/or 2- or 3-hydroxypropyl methacrylate (2- or 3-HPMA). More preferably,
the hydrophilic block consists of at least two different hydrophilic monomeric
units. For example, it may be a random copolymer of at least two different
hydrophilic monomeric units such as HEMA and 2-HPMA.
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In order to introduce ionic groups into the monomer, charged monomeric
units such as acrylic acid (acrylate) and/or methacrylic acid (methacrylate)
may be incorporated into the hydrophilic block. Thus, in a particular
embodiment of the present invention, the hydrophilic block can be made
from at least one non-ionic hydrophilic monomeric unit (e.g. as described
above) and from at least one ionic hydrophilic monomeric unit, wherein the
ionic monomeric unit is present in a molar amount of preferably 1-20 mole-
%. In case the hydrophilic block comprises an ionic monomeric unit such as
acrylic acid or methacrylic acid, copolymerization with a hydrophilic monomer
selected from the group of (meth)acrylamides, particularly N,N-dialkyl acryl-
or methacrylamides is preferred.
The hydrophobic monomeric units of the hydrophobic block are preferably
selected from hydrophobic acrylic and/or methacrylic units, styrene-based
monomeric units or combinations thereof. Preferably the hydrophobic
monomeric units are selected from hydrophobic (meth)acryl esters, e.g.
esters having an alcohol portion with 1-3 C-atoms without hydrophilic group.
Specific examples of monomeric units for the hydrophobic block are selected
from:
methyl acrylate,
methyl methacrylate (MMA),
ethyl acrylate
ethyl methacrylate (EMA),
n- or i-propyl acrylate,
n- or i-propyl methacrylate,
n-butyl acrylate,
n-butyl methacrylate (BUMA),
neopentyl acrylate,
neopentyl methacrylate,
and combinations thereof.
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The hydrophobic block preferably comprises at least two different
hydrophobic monomeric units, which are e.g. present as a random
copolymer. In a preferred embodiment, the hydrophobic block comprises
methyl methacrylate (MMA) and n-butyl methacrylate (BUMA). In an
especially preferred embodiment, the hydrophobic block is a random
copolymer of MMA and BUMA. The molar ratio between MMA and BUMA is
preferably about 60% (MMA) : 40% (BUMA) to about 40% (MMA) : 60%
(BUMA), e.g. about 50% (MMA) : 50% (BUMA). The glass transition
temperature of the hydrophobic block is preferably 100 C or less, 90 C or
Do less or 80 C or less, e.g. about 40-80 C. In an alternative embodiment,
the
hydrophobic block may consist of styrenic units, e.g. of polystyrene having a
glass transition temperature of about 95 C.
The block copolymers used in the present invention may be manufactured
according to known methods (Baker et al., Macromolecules 34 (2001), 7477-
7488).
The block copolymers may be applied to the electrode system by usual
techniques, e.g. by providing a solution of the block copolymer in a suitable
solvent or solvent mixture, e.g. an organic solvent, such as ether, which is
applied to the prefabricated electrode system and dried thereon.
When the block copolymer is contacted with water, it shows a water uptake
of preferably about 15%-30% by weight (based on the polymer dry weight) at
a temperature of 37 C and a pH of 7.4 (aqueous phosphate buffer 10 mM
KH2PO4, 10 mM NaH2PO4 and 147 mM NaCI).
In addition to the block copolymer the diffusion barrier may also comprise
further components, particularly non-polymeric components, which may be
dispersed and/or dissolved in the polymer. These further compounds include
plasticizers, particularly biocompatible plasticizers, such as tri-(2-
ethylhexyl)
trimellitate and/or glycerol.
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The diffusion barrier of the invention has a high effective diffusion
coefficient
Deff for glucose which is preferably 10-10 cm2/s, more preferably 5.10-
1
cm2/s, and even more preferably 10-9
cm2/s, and e.g. up to 10-7 or 10-8
cm2/s at a temperature of 37 C and a pH of 7.4. The effective diffusion
coefficient is preferably determined as described in Example 4 according to
the equation:
Deff=SEm/F -Lm.51 82- 10-8
wherein SEm is the sensitivity of the working electrode, F is the area of the
working electrode, and Lm is the layer thickness of the diffusion barrier. SEm
and Lm may be determined as described in the Examples.
The electrode system of the present invention is suitable for measuring the
concentration of an analyte under in-vivo conditions, i.e. when inserted or
implanted into a body. The analyte may be any molecule or ion present in
tissue or body fluid, for example oxygen, carbon dioxide, salts (cations
and/or anions), fats or fat components, carbohydrates or carbohydrate
components, proteins or protein components, or other type of biomolecules.
Especially preferred is the determination of analytes which can be efficiently
transferred between body fluid, e.g. blood and tissue such as oxygen, carbon
dioxide, sodium cations, chloride anions, glucose, urea, glycerol, lactate and
pyruvate.
The electrode system comprises an enzyme immobilized on an electrode.
The enzyme is suitable for the determination of a desired analyte. Preferably,
the enzyme is capable of catalytically converting the analyte and thereby
generating an electric signal detectable by the electric conductor of the
working electrode. The enzyme for measuring the analyte is preferably an
oxidase, for example glucose oxidase or lactate oxidase or a
dehydrogenase, for example a glucose dehydrogenase or a lactate
dehydrogenase. In addition to the enzyme, the enzyme layer may also
comprise an electrocatalyst or a redox mediator which favours the transfer of
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electrons to conductive components of the working electrode, e. g. graphite
particles. Suitable electro-catalysts are metal oxides such as manganese
dioxide or organo-metallic compounds such as cobalt phthalo-cyanine. In a
preferred embodiment the redox mediator is capable of degrading hydrogen
peroxide thereby counteracting depletion of oxygen in the surroundings of
the working electrode. In a different embodiment, a redox mediator may be
covalently bound to the enzyme and thereby effect direct electron transfer to
the working electrode. Suitable redox mediators for direct electron transfer
are prosthetic groups, such as pyrrolo quinoline quinone (PQQ), flavine
adenine dinucleotide (FAD) or other known prosthetic groups. Enzymes
immobilized on electrodes are e.g. described in WO 2007/147475.
A preferred embodiment of the electrode system comprises a
counterelectrode with an electrical conductor and a working electrode with an
electrical conductor on which an enzyme layer and the diffusion barrier are
arranged. The enzyme layer is preferably designed in the form of multiple
fields that are arranged on the conductor of the working electrode at a
distance, e.g. at least 0.3 mm or at least 0.5 mm from each other. The
individual fields of the working electrode may form a series of individual
working electrodes. Between these fields, the conductor of the working
electrode may be covered by an insulation layer. By arranging the fields of
the enzyme layer on the top of openings of an electrically insulating layer,
the
signal-to-noise ratio can be improved. Such an arrangement is disclosed in
WO 2010/028708.
The electrode system of the invention may additionally comprise a reference
electrode capable of supplying a reference potential for the working
electrode, e.g. an Ag/Ag-CI reference electrode. Moreover, an electrode
system according to the invention can have additional counter- and/or
working electrodes.
The electrode system may be part of a sensor, e.g. by being connected to a
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potentiostat and an amplifier for amplification of measuring signals of the
electrode system. The sensor is preferably an enzymatic non-fluidic (ENF)
sensor, more preferably an electrochemical ENF sensor The electrodes of
the electrode system may be arranged on a substrate that carries the
potentiostat or be attached to a circuit board that carries the potentiostat.
A further subject-matter of the invention is related to the use of a block
copolymer having at least one hydrophilic block and at least one hydrophobic
block as a diffusion barrier for an enzymatic electrode. The block copolymer
lo is preferably as described above, e.g. a single block-copolymer. The
diffusion barrier and the enzymatic electrode are preferably also as
described above.
Further details and advantages of the invention are explained based on an
exemplary embodiment making reference to the appended drawings.
Fig. 1 shows an exemplary embodiment of an electrode system according to
the invention.
Fig. 2 shows a detail view of Fig. 1.
Fig. 3 shows another detail view of Fig. 1.
Fig. 4 shows a section along the section line CC of Fig. 2.
Fig. 5 shows the sensitivity (with standard deviation) of four glucose sensors
(at 10 mM glucose) provided with different block polymers (C, F, D, B) as
barrier layers.
Fig. 6 shows the sensor drift of four glucose sensors provided with different
block copolymers (A, C, D, F) as barrier layers.
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Fig. 7 shows the conductivity of block copolymer A dependent on time
(2 experiments).
Fig. 8 shows the conductivity of block copolymer F dependent on time
(3 experiments).
Fig. 9 shows the conductivity of block copolymer H dependent on time for a
layer thickness of 2.77 pm or 4.43 pm, respectively.
Fig. 10 shows the fibrinogen adhesion to different spacer membrane
polymers in-vitro (Adapt and Eudragit E100) with respect to an uncoated
plate (Blank).
Fig. 11 shows the expression of surface protein CD54 by THP-1 cells after
incubation with sensors coated with a spacer membrane (Adapt and
Lipidure CM5206) or uncoated (control = untreated cells).
Fig. 12a and 12b show the secretion of cytokine IL-8 and MCP-1,
respectively, by THP-1 cells after incubation with sensors coated with a
spacer membrane (Adapt and Lipidure CM5206) or uncoated (control =
untreated cells).
Fig. 13 shows the secretion of cytokine IL-8 by THP-1 cells after incubation
with tissue culture plates coated with a spacer membrane (Adapt , Lipidure
CM5206 and Eudragit E100) or uncoated (Polyst.), and an additional layer of
fibrinogen.
Fig. 14 shows the haemolysis after incubation with sensors coated with a
spacer membrane (Adapt and Lipidure CM5206) or uncoated in
comparison to the spacer polymer Adapt without sensor (negative control =
incubation medium only; positive control = 100% osmotic lysis).
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Figure 1 shows an exemplary embodiment of an electrode system for
insertion into body tissue of a human or animal, for example into cutis or
subcutaneous fatty tissue. A magnification of detail view A is shown in Figure
2, a magnification of detail view B is shown in Figure 3. Figure 4 shows a
corresponding sectional view along the section line, CC, of Figure 2.
The electrode system shown has a working electrode 1, a counterelectrode
2, and a reference electrode 3. Electrical conductors of the electrodes la,
2a, 3a are arranged in the form of metallic conductor paths, preferably made
of palladium or gold, on a substrate 4. In the exemplary embodiment shown,
the substrate 4 is a flexible plastic plate, for example made of polyester.
The
substrate 4 is less than 0.5 mm thick, for example 100 to 300 micrometers,
and is therefore easy to bend such that it can adapt to movements of
surrounding body tissue after its insertion. The substrate 4 has a narrow
shaft for insertion into body tissue of a patient and a wide head for con-
nection to an electronic system that is arranged outside the body. The shaft
of the substrate 4 preferably is at least 1 cm in length, in particular 2 cm
to 5
cm.
In the exemplary embodiment shown, one part of the measuring facility,
namely the head of the substrate, projects from the body of a patient during
use. Alternatively, it is feasible just as well, though, to implant the entire
measuring facility and transmit measuring data in a wireless fashion to a
receiver that is arranged outside the body.
The working electrode 1 carries an enzyme layer 5 that contains immobilized
enzyme molecules for catalytic conversion of the analyte. The enzyme layer
5 can be applied, for example, in the form of a curing paste of carbon
particles, a polymeric binding agent, a redox mediator or an electro-catalyst,
and enzyme molecules. Details of the production of an enzyme layer 5 of this
type are disclosed, for example, in WO 2007/147475, reference to which is
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be made in this context.
In the exemplary embodiment shown, the enzyme layer 5 is not applied
continuously on the conductor 1a of the working electrode 1, but rather in the
form of individual fields that are arranged at a distance from each other. The
individual fields of the enzyme layer 5 in the exemplary embodiment shown
are arranged in a series.
The conductor la of the working electrode 1 has narrow sites between the
enzyme layer fields that are seen particularly well in Figure 2. The conductor
2a of the counterelectrode 2 has a contour that follows the course of the
conductor la of the working electrode 1. This means results in an
intercalating or interdigitated arrangement of working electrode 1 and
counterelectrode 2 with advantageously short current paths and low current
density.
In order to increase its effective surface, the counterelectrode 2 can be
provided with a porous electrically conductive layer 6 that is situated in the
form of individual fields on the conductor 2a of the counterelectrode 2. Like
the enzyme layer 5 of the working electrode 1, this layer 6 can be applied in
the form of a curing paste of carbon particles and a polymeric binding agent.
The fields of the layer 6 preferably have the same dimensions as the fields of
the enzyme layer 5, although this is not obligatory. However, measures for
increasing the surface of the counterelectrode can just as well be foregone
and the counterelectrode 2 can just as well be designed to be a linear
conductor path with no coatings of any kind, or with a coating made from the
described block copolymer and optionally a spacer.
The reference electrode 3 is arranged between the conductor la of the
working electrode 1 and the conductor 2a of the counterelectrode 2. The
reference electrode shown in Figure 3 consists of a conductor 3a on which a
field 3b of conductive silver/silver chloride paste is arranged.
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Figure 4 shows a schematic sectional view along the section line, CC, of
Figure 2. The section line, CC, extends through one of the enzyme layer
fields 5 of the working electrode 1 and between the fields of the conductive
layer 6 of the counterelectrode 2. Between the fields of enzyme layer 5, the
conductor la of the working electrode 1 can be covered with an electrically
insulating layer 7, like the conductor 2a of the counterelectrode 2 between
the fields of the conductive layers 6, in order to prevent interfering
reactions
which may otherwise be catalysed by the metal of the conductor paths la,
2a. The fields of the enzyme layer 5 are therefore situated in openings of the
insulation layer 7. Likewise, the fields of the conductive layer 6 of the
counterelectrode 2 may also be placed on top of openings of the insulation
layer 7.
The enzyme layer 5 is covered by a cover layer 8 which presents a diffusion
resistance to the analyte to be measured and therefore acts as a diffusion
barrier. The diffusion barrier 8 consists of a single copolymer with
alternating
hydrophilic and hydrophobic blocks as described above.
A favourable thickness of the cover layer 8 is, for example, 3 to 30 pm,
particularly from about 5-10 pm or from about 10-15 pm. Because of its
diffusion resistance, the cover layer 8 causes fewer analyte molecules to
reach the enzyme layer 5 per unit of time. Accordingly, the cover layer 8
reduces the rate at which analyte molecules are converted, and therefore
counteracts a depletion of the analyte concentration in surroundings of the
working electrode.
The cover layer 8 extends continuously essentially over the entire area of the
conductor la of the working electrode 1. On the cover layer 8, a
biocompatible membrane may be arranged as spacer 9 that establishes a
minimal distance between the enzyme layer 5 and cells of surrounding body
tissue. This means advantageously generates a reservoir for analyte
molecules from which analyte molecules can get to the corresponding
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enzyme layer field 5 in case of a transient disturbance of the fluid exchange
in the surroundings of an enzyme layer field 5. If the exchange of body fluid
in the surroundings of the electrode system is transiently limited or even
prevented, the analyte molecules stored in the spacer 9 keep diffusing to the
enzyme layer 5 of the working electrode 1 where they are converted. The
spacer 9 therefore causes a notable depletion of the analyte concentration
and corresponding falsification of the measuring results to occur only after a
significantly longer period of time. In the exemplary embodiment shown, the
membrane forming the spacer 9 also covers the counterelectrode 2 and the
reference electrode 3.
The spacer membrane 9 can, for example, be a dialysis membrane. In this
context, a dialysis membrane is understood to be a membrane that is
impermeable for molecules larger than a maximal size. The dialysis
membrane can be prefabricated in a separate manufacturing process and
may then be applied during the fabrication of the electrode system. The
maximal size of the molecules for which the dialysis membrane is permeable
is selected such that analyte molecules can pass, while larger molecules are
retained.
Alternatively, instead of a dialysis membrane, a coating made of a polymer
that is highly permeable for the analyte and water, for example on the basis
of polyurethane or of acrylate, can be applied over the electrode system as
spacer membrane 9.
Preferably, the spacer is made from a copolymer of (meth)acrylates.
Preferably, the spacer membrane is a copolymer from at least 2 or 3
(meth)acrylates. More preferably, the spacer membrane comprises more
than 50 mol-%, at least 60 mol-% or at least 70 mol-% hydrophilic monomer
units, e.g. HEMA and/or 2-HPMA, and up to 40 mol-% or up to 30 mol-%
hydrophobic units, e.g. BUMA and/or MMA. The spacer may be a random or
block copolymer. An especially preferred spacer membrane comprises MMA
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or BUMA as hydrophobic moieties and 2-HEMA and/or 2-HPMA as
hydrophilic moieties. Preferably, the amount of the hydrophilic monomers
HEMA and/or HPMA is between 80 mol-% to 85 mol-% and the amount of
hydrophobic component MMA and/or BUMA is between 15 mol-% and
20 mol-%.
The very preferred spacer membrane of the invention is made of the
copolymer Adapt (Biolnteractions Ltd, Reading, England). Adapt
comprises BUMA as hydrophobic moiety and 2-HEMA and 2-HPMA as
hydrophilic moieties, wherein the amount of the 2-HEMA hydrophilic
monomers is about 80 mol-%.
The spacer membrane is highly permeable for the analyte, i.e. it does
significantly lower the sensitivity per area of the working electrode, for
example 20% or less, or 5% or less with a layer thickness of less than about
pm, preferably less than about 5 pm. An especially preferred thickness of
the spacer membrane is from about 1 to about 3 pm.
The enzyme layer 5 of the electrode system can contain metal oxide
20 particles, preferably manganese dioxide particles, as catalytic redox
mediator. Manganese dioxide catalytically converts hydrogen peroxide that is
formed, for example, by enzymatic oxidation of glucose and other
bioanalytes. During the degradation of hydrogen peroxide, the manganese
di-oxide particles transfer electrons to conductive components of the working
electrode 1, for example to graphite particles in the enzyme layer 5. The
catalytic degradation of hydrogen peroxide counteracts any decrease of the
oxygen concentration in the enzyme layer 5. Advantageously this allows the
conversion of the analyte to be detected in the enzyme layer 5 to not be
limited by the local oxygen concentration. The use of the catalytic redox
mediator therefore counteracts a falsification of the measuring result by the
oxygen concentration being low. Another advantage of a catalytic redox
mediator is that it prevents the generation of cell-damaging concentrations of
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hydrogen peroxide.
The preferred spacer membrane polymer described herein may be used as
an outer coating for a diffusion barrier of the present invention, but also as
an outer coating of an electrode system in general, particularly of an
electrode system for measuring the concentration of an analyte under in-vivo
conditions, comprising an electrode with immobilized enzyme molecules and
a diffusion barrier that controls diffusion of the analyte from the exterior
of
the electrode system to the enzyme molecules. Thus, the spacer membrane
can be arranged on the diffusion barrier, however, the spacer membrane can
also be arranged directly on the enzyme layer. In this last context, the
spacer
membrane can also act as a diffusion barrier itself and slow down the
diffusion of analyte molecules to the enzyme layer.
Is When the electrode system of the invention is inserted or implanted into
a
body, the spacer membrane is the interface between the implanted sensor
and the surrounding body fluid or tissue. Consequently, when exposed to the
body fluid or tissue, the spacer membrane of the invention must be
mechanically robust so that it is neither deformed nor shoved off the sensor.
To this end, the spacer membrane copolymer water uptake and the
concomitant swelling of the copolymer must be limited, albeit the inherent
hydrophilicity of the copolymer.
Preferably, the relative water uptake of the spacer membrane copolymer
should not exceed 50 wt-%, preferably 40 wt-%, more preferably 30 wt-%
based on the total rate of the copolymer. Within the context of the present
invention, the measurement of the relative water uptake is performed by
subjecting the dry copolymer to an excess of phosphate-buffer (pH 7.4) for
48 h at a temperature of 37 C. The relative water uptake (WW/0) is
preferably determined according to the equation:
WU% = (m2-rn1)/m1 x 100,
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wherein m1 and m2 represent, respectively, the mass of the dry copolymer
and the copolymer after hydration according to the above measurement
conditions.
The inventors of the present invention determined that the preferred spacer
membrane made of the copolymer Adapt takes up 33 1.8 wt-% of
phosphate-buffer (pH 7.4) relative to its own weight over 48 h at 37 C.
Under the same conditions, a membrane of polymer Lipidure CM5206 (NOF
Corporation, Japan) takes up 157 9.7 wt-% of phosphate-buffer relative to
its own weight. The lower water uptake of the polymer advantageously
increases the mechanical stability of the spacer membrane of the present
invention. By contrast, Lipidure CM5206 shows a higher water uptake and
swells to a hydrogel which is more fragile, easily deformable or shoved off,
particularly when applied on an electrode system of an enzymatic in-vivo
sensor.
Furthermore, during insertion and implementation of the electrode system,
the spacer is in direct contact with the tissue and/or the body fluid, like
interstitial fluid or blood, containing biomolecules like proteins and cells.
Advantageously, the spacer membrane must protect the inserted and
implanted sensor in the tissue and/or body fluid environment and, thus,
minimize the tissue reaction of the body to the implant. In fact, reactions of
the body against implanted material are known as "foreign body response"
(FBR). By FBR, the body tries to destroy the implant or, if it is not
possible, to
create a capsule to separate it from the surrounding tissue (foreign body
granuloma). The first step of the FBR reaction is binding of proteins (e.g.
fibrinogen, albumin, immunoglobulin, complement) to the surface of the
former material, i.e. the implant. This protein coat presents binding sites to
receptors on immune cells. For example, fibrinogen contains a structural
motif that binds to the monocyte receptor MAC-1. When fibrinogen binds to
the surface of the implant, it changes its conformation and exposes the
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binding site for MAC-1. Consequently, immune cells, like monocytes, are
recruited to the implant and activated, secreting enzymes and radicals to
attack the implant. Additionally, immune cells secrete soluble factors, i.e.
cytokines, to recruit and activate other immune cells and thereby amplifying
the immune response. If the implant cannot be removed, a fibrous capsule is
formed by connective tissue cells and proteins. This capsule, however, is a
diffusion barrier for analytes to reach the sensor. Collectively, the events
of
the foreign body response as described above are likely to interfere with the
electrode system function in-vivo and with its life time.
Thus, an improved spacer membrane on an electrode-system of an
enzymatic in-vivo sensor further provides the reduction of the tissue
response to the implant and inhibits the formation of a capsule separating
the sensor from the surrounding tissue and body fluids.
Thus, it is a further object of the present invention to provide an electrode
system for measuring the concentration of an analyte under in-vivo
conditions, comprising an electrode with immobilized enzyme molecules and
preferably a diffusion barrier that controls diffusion of the analyte form the
exterior of the electrode system to the enzyme molecules, characterised in
that a spacer membrane forms at least a portion of the outer layer of the
electrode system, wherein the spacer membrane comprises a hydrophilic
copolymer of acrylic and/or methacrylic monomers, wherein the polymer
comprises more than 50 mol- /0 hydrophilic monomers.
As described above, the spacer membrane of the invention does have
limited protein-binding capacity to protect the electrode system of the sensor
from protein adsorption that might trigger response of immune cells and
might limit or interfere its performance in-vivo. Example 5 and 6 show that
the preferred spacer membrane of the invention provides little binding to
fibrinogen and prevents conformational change of fibrinogen that would lead
to the exposure of the MAC-1 binding motif for monocytes. Advantageously,
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the spacer membrane copolymer material does not activate immune cells
itself. In Example 6, it could be demonstrated that the spacer membrane
copolymer of the invention is able to attenuate the activation of immune cells
by the implanted sensor. Moreover, advantageously, the spacer membrane
is a biocompatible material, in particular, is compatible with body fluids,
e.g.
with blood. Example 7 shows that the spacer membrane copolymer of the
present invention is able to prevent haemolysis and the complement
activation by the implanted sensor. Thus, the spacer membrane of the
invention advantageously not only shows a high mechanical stability, but
io also has optimal biocompatible properties, which is surprising due to
the low
water uptake when wetted.
The features of this embodiment particularly with regard to the structure of
the electrode system, the analyte and the enzyme molecules are as
described herein. The diffusion barrier is preferably as described herein, it
may however also have a different composition or may be absent. According
to one preferred embodiment, the diffusion barrier preferably comprises a
block copolymer having at least one hydrophilic block and at least one
hydrophobic block as described herein.
According to a further preferred embodiment, the diffusion barrier comprises
hydrophilic polyurethanes. The hydrophilic polyurethanes used as diffusion
membrane can be produced by polyaddition of an (poly)diisocyanat,
preferably 4-4-methylene-bis(cyclohexylisocyanate) with a polyalcohol,
preferably a diol mixture.
The components of the diol mixture are preferably polyalkylene glycols, such
as polyethylene glycol (PEG) and polypropylene glycol (PPG) and aliphatic
diols, such as ethylene glycol. Preferably, the hdydrophilic polyurethane
comprises 45-55 mol-`)/0, preferably 50 mol- /0 isocyanate and 25-35 mo1-13/0,
preferably 30 mol-`)/0 ethylene glycol. The degree of hydrophilization is then
adjusted by the ratio of PEG to PPG. Preferably, the polyurethane comprises
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2-3 nnol-`)/0, more preferably 2.5 mol-% PEG and 17-18 mol-%, preferably
17,5 mol-% PPG. In order to increase the hydrophility of the polyurethane,
the proportion of PEG may be increased, for instance, to 4.5 - 5.5 mol-%,
preferably 5 mol-% PEG, in order to obtain an extremely hydrophilic
polyurethane. The different hydrophilic variants of the polyurethanes may
also be mixed in order to optimize the properties of the diffusion barrier.
The preferred acrylic and methacrylic monomers of the spacer membrane
copolymer are as described herein.
The hydrophilic monomeric units are preferably selected from hydrophilic
(meth)acryl esters, i.e. esters with a polar, i.e. OH, OCH3 or 0C2H5 group
within the alcohol portion of the ester, hydrophilic (meth)acrylamides with an
amide (NH2) or an N-alkyl- or N,N-dialkylamide group, wherein the alkyl
group comprises 1-3 C-atoms and optionally hydrophilic groups such as OH,
OCH3 or 0C2H5, and suitable (meth)acrylic units having a charged, e.g. an
anionic or cationic group, such as acrylic acid (acrylate) or methacrylic acid
(methacrylate). Further, combinations of monomeric units may be employed.
Specific examples of preferred monomeric units for the hydrophilic block are
selected from:
2-hydroxyethyl acrylate,
2-hydroxyethyl methacrylate (HEMA),
2-methoxyethyl acrylate,
2-methoxyethyl methacrylate,
2-ethoxyethyl acrylate,
2-ethoxyethyl methacrylate,
2- or 3-hydroxypropyl acrylate,
2- or 3-hydroxypropyl methacrylate (2- or 3-HPMA),
2- or 3-methoxypropyl acrylate,
2- or 3-methoxpropyl methacrylate,
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2- or 3-ethoxypropyl acrylate,
2- or 3-ethoxypropyl methacrylate,
1- or 2-glycerol acrylate,
1- or 2-glycerol methacrylate,
acrylamide,
methacrylamide,
an N-alkyl- or N,N-dialkyl acrylamide, and
an N-alkyl- or N,N-dialkyl methylamide, wherein alkyl comprises
1-3 C-atoms such as methyl, ethyl or propyl,
tO acrylic acid (acrylate),
methacrylic acid (methacrylate)
and combinations thereof.
Preferred hydrophilic monomers are 2-hydroxyethyl methacrylate (HEMA)
and/or 2- or 3-hydroxypropyl methacrylate (2- or 3-HPMA).
The hydrophobic monomeric units are preferably selected from hydrophobic
acrylic and/or methacrylic units or combinations thereof. Preferably the
hydrophobic monomeric units are selected from hydrophobic (meth)acryl
esters, e.g. esters having an alcohol portion with 1-3 C-atoms without
hydrophilic group.
Specific examples of monomeric units for the hydrophobic block are selected
from:
methyl acrylate,
methyl methacrylate (MMA),
ethyl acrylate
ethyl methacrylate (EMA),
n- or i-propyl acrylate,
n- or i-propyl methacrylate,
n-butyl acrylate,
n-butyl methacrylate (BUMA),
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neopentyl acrylate,
neopentyl methacrylate,
and combinations thereof.
In a preferred embodiment, the hydrophobic block comprises methyl
methacrylate (MMA) and n-butyl methacrylate (BUMA).
The outer spacer membrane preferably covers at least the working electrode
portion comprising the enzyme molecules and optionally also other portions,
e.g. the counter electrode. If one is present, the spacer membrane also
covers the reference electrode. The spacer membrane preferably covers the
entire implanted surface of the electrode system. The spacer membrane
preferably covers the working electrode, optionally the counter-electrode and
the reference electrode if present in the form of a continuous layer.
The electrode system comprising the improved spacer membrane of the
invention may be part of a sensor, e.g. by being connected to a potentiostat
and an amplifier for amplification of measuring signals of the electrode
system. The sensor is preferably an enzymatic non-fluidic (ENF) sensor,
more preferably an electrochemical ENF sensor The electrodes of the
electrode system may be arranged on a substrate that carries the
potentiostat or be attached to a circuit board that carries the potentiostat.
Preferably, the sensor is for the measurement of glucose.
A further subject-matter of the invention is related to the use of a
hydrophilic
copolymer of acrylic and/or methacrylic monomers, wherein the hydrophilic
compolymer comprises more than 50 mol-% hydrophilic monomers as a
spacer membrane for an enzymatic electrode. The hydrophilic copolymer is
preferably as described above. Preferably the spacer membrane is used for
minimising the foreign body reaction (FRB) against the enzymatic electrode
when it is inserted or implanted into the body.
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Example 1 Permeability of an enzymatic non-fluidic (ENF) glucose sensor
with distributed electrodes for transcutaneous implantation having a diffusion
layer consisting of one single block copolymer.
The sensor was built on a prefabricated palladium strip conductor structure
on a polyester substrate having a thickness of 250 pm. Working electrode
(WE) and counterelectrode (CE) were arranged distributedly (as shown in
Figs 1-2).
lo The fields of the CE were overprinted with carbon paste, the rest of the
strip
conductor was insulated. The fields of the WE were overprinted with a
mixture of cross-linked glucose oxidase (enzyme), conductive polymer paste
and electric catalyst, here manganese(IV)-oxide (Technipur). The remaining
paths of the strip conductor were again insulated. The reference electrode
(RE) consists of Ag/AgCI paste. The electrodes cover about 1 cm of the
sensor shaft.
The WE-fields were coated with a block copolymer diffusion layer consisting
of a HEMA block and a BUMA block. The thickness of the layer is 7 pm.
Four sensor batches were produced, each provided with a specific block
copolymer as diffusion layer (see list hereinbelow). All block copolymers
were obtained from Polymer Source, Montreal and are listed in the following
Table 1.
Name molecular ratio/% monomeric units molecular weight
Copolymer BUMA/HEMA HEMA Copolymer [kD]
C 73/27 92 47
F 60/40 108 37
D 48/52 162 44
B 62/38 169 61
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The respective block copolymer was dissolved in organic solvent (25%
concentration) and the sensors were coated therewith. After drying by means
of belt driers (2 min, 30 to 50 C), the coated sensors were tested in-vitro in
glucose solutions of different concentrations. Of each sensor batch 10
sensors were measured as random sample. As a measure for the in-vitro
sensitivity, the signal was calculated by the difference of the measured
currents at 10 mM and 0 mM glucose concentration, which then was divided
by 10 mM (cf. Example 4).
All sensors were operated at a polarisation voltage of 350 mV versus
Ag/AgCI, the measured temperature was kept constant at 37 C. The sensors
used for this measurement series did not comprise the spacer described in
WO 2010/028708, which, however, did not make any difference in view of
the tested signal level. Fig. 5 shows the sensor sensitivity with standard
deviations for the four different diffusion layers.
Concerning block copolymers C, D and F, there is a clear connection
between in-vitro sensitivity and molar ratio of hydrophobic block compared to
hydrophilic block. At about identical total chain length of the copolymer, the
sensitivity increases as the amount of hydrophilic block (HEMA) increases.
The sensors having a diffusion layer of block copolymer B are an exception.
Even though polymer B has a relative ratio of hydrophobic to hydrophilic
amount similar to polymer F, the sensitivity and thus the permeability for
glucose is reduced. Empirically it can be stated that in case of polymer B the
total chain length - corresponding to the molecular weight (total) of the
copolymer molecule - is so large that the permeability of the layer is
reduced.
This may also be seen in the gravimetrically determined water uptake of
block copolymer B as compared to the remaining polymers. Polymer B has a
water uptake of 10.6% 1.5% (weight percent referred to the polymer dry
weight). Polymer C lies at 15.6% 0.0%, polymer F at 16.5 3.1% and
polymer D at 27% 1.7%.
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Example 2 Mechanic flexibility of the diffusion layer of an ENF glucose
sensor
The sensor was manufactured as described in W02010/028708, however
having a diffusion layer according to the present invention. It was assumed
that the glass transition temperature (Tg) is a substitute parameter for the
mechanic flexibility. Further, it was assumed that the glass transition
temperature, which may be allocated to the hydrophobic block, determines
the mechanic flexibility in in-vivo applications. It should be noted that
several
Tgs may be identified for one block copolymer, corresponding to the number
of blocks.
The sensors were coated with the same electrode pastes as in Example 1.
Then, some of the sensors were coated with a copolymer selected from
MMA-HEMA (produced by Polymer Source, Montreal). This polymer (called
E) has a total molecular weight of 41 kD, the molar ratio of MMA
(hydrophobic amount) to HEMA is 60%:40%. The glass transition
temperature of the hydrophobic block is 111 C, determined by DSC and a
heating rate of 10 C/min.
Besides, other sensors were provided with a diffusion layer of a block
copolymer of the invention (called A). The hydrophobic block of said
copolymer A contains MMA and BUMA at equal molar amounts in a
randomised sequence. Again, the molar ratio of the hydrophobic part to the
hydrophilic part is 60%:40%. The molecular weight is 36 kD. The Tg of the
hydrophobic block decreases, due to the randomized sequence of MMA and
BUMA (Tg about 45 C), to 73 C.
Both diffusion layers were generated from the respective solution (25%) of
the copolymers in ether and dried as in Example 1. The thickness of the
diffusion layers was 7 pm. A spacer layer was applied subsequently via dip
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coating and dried 24h at room temperature. The spacer layer was made of
Lipidure CM 5206, produced by NOF Japan.
After explantation from the tissue, sensors having a copolymer E diffusion
layer show sporadic cracks in the diffusion layer. This is taken as an effect
of
the mechanic load. In contrast thereto, sensors having a copolymer A
diffusion layer, do not show any cracks under identical load. This is
obviously
due to the reduction of Tg, which increases the mechanic stability of the
copolymer. A physical mixture of two copolymers, as disclosed in
W02010/028708, is no longer required.
Example 3 Optimized permeation behaviour of an ENF glucose sensor with
distributed electrode and diffusion layer according to the invention.
A sensor was manufactured as described in Example 1, but with an
additional spacer layer on the total of the sensor shaft. Sensors with
respective diffusion layer were produced for copolymers A, C, D and F of
Examples 1 and 2. For this purpose, a 24% etheric solution of the
copolymer was generated. Each solution was applied onto a set of sensors
(N=10) and then dried in a band drier. Thereby, diffusion layers having a
thickness of 7pm, were obtained.
Afterwards, the sensors were provided with a spacer layer as described in
Example 2.
The sensor was connected with a measuring system on the sensor head,
which transfers the measured data to a data store. The in-vitro
measurements were carried out as in Example 1, however over a measuring
period of 7 days. From the measured data, the sensitivity drift was calculated
over the respective measuring period for each sensor. Figure 6 shows for
each sensor variant, i.e. sensors of a variant of the diffusion layer, the
mean
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value of the in-vitro drift value for the group. The initial phase of the
measurement - the first 6h, the so-called startup phase - was excluded from
the calculation.
For all copolymers C, D and F having a hydrophobic block of BUMA, there is
a positive drift, i.e. the sensitivity increases according to time. Contrary
thereto, copolymer A with the hydrophobic block of a random copolymer of
MMA and BUMA, leads to a very low, slightly negative, drift.
Do These differences may be explained by the long-time permeability
response
of the respective diffusion layers, which was measured in additional
experiments. Palladium sensors without WE-paste, but with a defined active
surface, i.e. also without an enzyme layer - excluding the influence of its
swelling behaviour on the results - were coated with the above polymer
solutions, and after drying, the thickness of the layer was measured.
Subsequently, conductivity was measured in sodium- and chloride-containing
buffer solution
Fig. 7 shows that the conductivity of copolymer A remained nearly constant
after a short startup phase.
This is not the case for copolymer F, even under identical measurement
conditions, as may be seen in Fig. 8. In this case, a long-term and strong
permeability response of the diffusion layer of copolymer F was observed,
which was practically independent of the layer thickness. For copolymer F -
and also copolymers C and D (not shown) - with a hydrophobic block of
BUMA, an increase of permeability results even over a long time period.
When measured, this leads to a continuous increase of sensitivity if the
diffusion layer is applied onto the sensor with distributed enzyme layer. This
explains the observed positive sensor drift.
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Vice versa, a sensor having block copolymer A, shows a negligible drift,
which is due to a very low permeability alteration in the conductivity
measurement. Directly after starting measurement (until about 1 h
afterwards), however, a strong increase of conductivity is observed in
copolymer A. Here, a very fast startup is observed, which is terminated after
about 1 hour. At this time, the diffusion layer is completely wetted and has
terminated its structural reorganisation due to water uptake. The extent of
the structural change presumably depends on the Tg. It seems plausible that
a copolymer having an increased Tg passes a reorganisation, which is
limited in time and amplitude, as compared to a copolymer having a Tg in the
range of the ambient temperature.
In addition, it has to be stated that sensors with copolymer A show a
comparatively high sensitivity at the start of measurements as compared to
sensors having a copolymer F diffusion layer. This is to be expected due to
the identical relative ratios between hydrophobic and hydrophilic blocks. The
achieved sensitivity range of 1 to 1.5 nA/mM (see Example 1) is deemed
ideal. This sensitivity is likewise obtained for sensors having a diffusion
layer
consisting of copolymer A.
Regarding the sum of the three physico-chemical characteristics -
permeability, mechanic stability and permeability response - an optimal
sensor may preferably be obtained with a diffusion layer of a block
copolymer, having a hydrophobic block with at least two different randomly
arranged hydrophobic monomeric units, such as block copolymer A. None of
the other block copolymers, whose hydrophobic blocks only consist of a
single monomeric unit reaches a quality, which could be compared in all
three parameters with copolymer A.
Example 4 Characterization of block copolymers
A multiple field sensor (10 fields of working electrodes and
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counterelectrodes, respectively) for the continuous measurement of the
glucose was produced and characterized in-vitro.
The sensor was provided with a diffusion layer consisting of a block
copolymer comprising a hydrophobic block of random copolymerized methyl
methacrylate (MMA) and n-butyl methacrylate (BUMA) and a hydrophilic
block of 2-hydroxyethyl methacrylate (HEMA). These polymers (specified G
and H) had been produced by Polymer Source, Montreal, and are more
permeable than polymer A from Examples 1-3, which is included herein by
reference.
In the following Table 2, the copolymers are described:
Polymer G H A
Molecular weights 23.5-b-29 21-b-20.5 21-b-15
Mn [kD]
Weight-% HEMA 55.2 49.4 41.6
Mol-% HEMA 53.5 47.4 40
(stoichiometrically)
Mol-% HEMA 51 46 32.6
(measured by
11-1,13C NMR)
--
Tg [ C] hydropho- 65 68 86
bic block
HEMA monomeric 223 157 115
units
MMA monomeric 194 174 174
units
The molecular weights Mn of each block are separately indicated in the
above Table 2 and represent average values, as polymers are known to
have distributions of molecular chain lengths around a specified mean value.
This also applies to the derived quantities in Table 2.
The indicated glass transition temperatures of the hydrophobic block are
within the desired range in order to guarantee mechanical flexibility.
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The decisive parameter with regard to the permeability of the diffusion
barrier
for the analyte is the sensitivity per area unit of the working electrode
(i.e. the
geometric area). The sensitivity SE was calculated from current (I)
measurements at 10 mM and at 0 mM glucose concentration in phosphate-
buffered solution (pH 7.4) in nA/mM:
SE = [1(10 mM) - 1(0 mM)]/10
for each of the analyzed sensors. From the individual measurement values
(N=8) the mean sensitivity SEm was determined. The obtained sensitivity
values were divided by the microscopically measured geometric total area F
of all working electrode spots on the multi-field sensor. Thereby, a
sensitivity
density SEm/F was obtained.
The linearity Y of the in-vitro function curve is an indication of the
diffusion
control functionality of the polymer cover layer on the working electrode. It
was calculated from current measurements at 20 mM, 10 mM and 0 mM
glucose concentration in VO:
y2OmM = 50.[1(20mM) - 1(0mM)]/[1(10mM) - 1(0mM)]
for each of the analyzed sensors. From the individual measurement values
the mean linearity value and its standard deviation were determined (cf.
Table 3).
Finally, the layer thickness L of the diffusion barrier of the sensors was
determined by optical measurement for each of the polymers. The
corresponding mean values were computed for a sample of > 23 sensors
with the same polymer. Therefrom, the effective diffusion coefficient Deff of
the cover layer may be calculated:
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Deff=SEm/F. Lm.5.182. 10-8
in cm2/s, wherein SEm and Lm are the respective mean values for the
sensitivity and the layer thickness, and F is the total area of all working
electrode spots.
The sensor drift was calculated from repetitions of the glucose concentration
stages over 7 days of in-vitro measurements. The results for polymer H
showing a substantially constant conductivity are depicted in Figure 9.
The following Table 3 shows the results of the functional characterization:
Polymer
SE,õ/F 1.85 1.25
[nA/mNrmm2)]
Drift (%d] -1.5 0.2 0.3 0.1
y20mM [OA 88.2 0.7 88.6 0.3
layer thickness 1_, 11.61 12.69
[pm]
Deff [cm2/s] 1.11305*10-9 8.22019*10-1
For the more hydrophilic polymer G (which is more permeable for glucose)
the diffusion coefficient was also determined with an alternative method, e.g.
permeation of glucose from a chamber with a glucose solution into a
chamber with a glucose-free buffer through a film of the polymer. According
to this method, a similar value for the diffusion coefficient was obtained
(1.17.10-9 cm2/s).
Example 5 Protein binding to spacer layer material
To assess protein binding to spacer layer materials, ethanolic solutions of
AdaptTM (Biointeractions Ltd, Reading, England) or Eudragit E100 (Evonik
Industries) were filled into an incubation plate (FluoroNunc Maxisorp, Thermo
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Scientific). Eudragit E100 is a cationic copolymer based on
dimethylaminoethyl methacrylate, butyl methacrylate and methyl
methacrylate. The polymers were dried over night at 40 C. Thereafter the
spacer materials were overlaid with fibrinogen solution. The solution
contained fibrinogen from human plasma that was conjugated to the
fluorescent dye A1exa488 (purchased from Invitrogen). After 4 h incubation
the fibrinogen solution was aspirated and the spacer layers were washed
eight times with borate buffer. The amount of spacer-bound protein was
analysed by measuring the fluorescence intensity in the incubation plate at
an excitation wavelength of 485 nm and an emission of 528 nm using a
fluorescence reader (Synergy4, BioTek Instruments). Known concentrations
of labelled protein (6.25 ¨ 500 ng) were used to prepare a calibration curve
to convert fluorescence readings to amount of protein.
As expected, fibrinogen attached to the uncoated incubation plate (Blank)
resulting in 390 ng of bound protein (Fig.10). The plate coated with Eudragit
E100 showed reduced protein binding of 60 ng. Hardly any protein binding
was detectable in the AdaptTM coated plate. The readings before incubation
resulted from background fluorescence. These results clearly demonstrate
that surfaces coated with the spacer materials, especially those coated with
AdaptTM, are well protected against fibrinogen adhesion.
Example 6 Cytokine release by cells after contact with spacer layer
Sensors were manufactured as described in Example 2. Afterwards, the
sensors were provided with a spacer layer as described in Example 2. The
spacer layer was made of Lipidure CM 5206 (NOF Corporation, Japan) or
was made of AdaptTM (Biointeractions Ltd, Reading, England).
Sensors without spacer layer, sensors with spacer layer made of Lipidure
CM 5206 or sensors with spacer layer made of AdaptTM were incubated with
monocytic THP-1 cells and induction of inflammatory markers was analysed.
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THP-1 cells were cultured in the presence of the sensors for 24 h at 37 C.
The cells were then collected by centrifugation. The supernatant was used to
determine the release of cytokines whereas the cell pellet was resuspended
in PBS containing 1 % of bovine serum albumin (BSA) and used to analyse
expression of the cell surface protein CD54 (also known as ICAM-1, an
inflammatory biomarker). THP-1 cells were incubated with an anti-CD54
antibody conjugated with the fluorescent dye phycoerythrin (BD Bioscience).
After incubation for 45 min at 4 C, cells were washed in PBS/ 1 % BSA and
the mean fluorescence intensity (MFI) of 10000 cells was determined using a
flow cytometer (excitation wavelength 532 nm, emission wavelength 585 nm)
(BD FACSArray, BD Bioscience). Compared to untreated THP-1 cells,
incubation with sensors without coating resulted in increased relative CD54
expression (6-fold induction) as indicated by high MFI readings (Fig. 11).
Incubating cells with sensors covered with spacer layers of CM 5206 or
AdaptTM resulted in attenuation of CD54 expression by 45 % or 41 %
respectively.
The supernatant was used to determine the amount of the cytokines
Interleukin- 8 (IL-8) and "monocyte chemotactic protein-1" (MCP-1) using a
bead-based immunoassay according to the manufacturer's instructions (Flex
sets, BD Bioscience) and subsequent flow cytometric analysis (BD
FACSArray, BD Bioscience). Data analysis was performed using FCAP array
software v1Ø1 (Soft flow Hungary Ltd.).
Compared to untreated THP-1 cells sensors without coating induced a strong
release of IL-8 (49 vs.197 pg/ml) (Fig. 12a) and MCP-1(6 vs.48 pg/ml) (Fig.
12b). The release of IL-8 and MCP-1 was reduced when sensors were
covered with spacer layers of CM 5206 or AdaptTM. Sensors covered with
AdaptTM induced the release of 100 pg/ml of IL-8 and 25 pg/ml of MCP-1.
Sensors covered with CM 5206 resulted in secretion of 125 pg/ml of IL-8 and
18 pg/ml of MCP-1.
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Collectively, these data indicate that the spacer layers attenuate the
induction of three well known inflammation biomarker, namely CD54, IL-8
and MCP-1.
To analyse the role of protein adsorption for activation of THP-1 cells,
tissue
culture plates were coated with CM 5206, AdaptTM or Eudragit E100. The
spacer was then incubated with human fibrinogen (Sigma-Aldrich). THP-1
cells were incubated with different spacers + fibrinogen layers. After 48 h
incubation at 37 C the cells were sedimented by centrifugation and the
supernatant was analysed for IL-8 release. As a control, cells were grown in
culture plates without spacer but coated with fibrinogen (Polyst. = culture
plate material). As shown in Fig. 13, these cells released 89 pg/ml of IL-8.
Cells cultured on CM 5206 + fibrinogen or on AdaptTM + fibrinogen released
68 or 49 pg/ml, respectively. In contrast, cells grown on Eudragit E100 +
fibrinogen released 206 pg/ml of IL-8. Notably, cells grown on Eudragit E100
without fibrinogen-coating secreted only 59 pg/ml of IL-8. Adsorption of
fibrinogen on the polymer surface and conformational changes in the protein
might expose the MAC-1 binding site. THP-1 cells that are activated via
binding to their MAC-1 receptor release cytokines, like IL-8, and thereby
trigger an inflammatory response. Hence, spacer layers made of AdaptTM or
CM 5206 avoid protein deposition and exposure of structural motifs on
surfaces (like sensors) and thereby minimise inflammatory reactions against
implants.
Example 7 Limited haemolysis of sensors coated with a spacer layer
Sensors were manufactured as described in Example 2. Afterwards, the
sensors were provided with a spacer layer as described in Example 2. The
spacer layer was made of Lipidure CM 5206 (NOF Corporation, Japan) or
was made of AdaptTM (Biointeractions Ltd, Reading, England).
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The haemolytic potential of sensors without spacer layer, sensors with
spacer layer made of Lipid ure CM 5206 or sensors with spacer layer made of
AdaptTM was analysed. Therefore, sensors with a total surface area of 6
cnn2 were incubated with red blood cells and then the lysis was determined
by measuring the release of haemoglobin to the supernatant. Erythrocytes
were isolated from fresh human blood by centrifugation (citrate was used to
avoid coagulation). They were then washed with phosphate buffered saline
(PBS) and thereafter diluted 1:40 in PBS. The erythrocyte suspension was
incubated with the sensors for 24 h at 37 C in the dark on a rotation platform
to (350 rpm). Afterwards, the cells were sedimented by centrifugation and
the
haemoglobin content of the supernatant was determined spectroscopically
by measuring the absorption of the supernatant at a wavelength of 575 nm.
The results are presented as lytic index in %, which is the release of
haemoglobin in a sample divided by haemoglobin release in the positive
control (= complete osmotic lysis of erythrocytes in distilled water). The
results are shown In Fig. 14.
Sensors without a spacer layer significantly caused haemolysis as indicated
by a high haemolytic index of 47.4 %. Coating the sensors with a spacer
layer of Lipidure CM 5206 reduced the haemolytic potential of the sensors as
demonstrated by a lytic index of 14.7 %. Sensors coated with a spacer layer
of AdaptTM marginally caused haemolysis resulting in a lytic index of 7.5 %
which is in the range of the negative control (= erythrocytes in PBS incubated
without any test material) or AdaptTM alone. These results indicated the
protective function of the spacer layer to reduce haemolysis.
