Optical sensor for in situ measurement of analytes

An implantable sensor for use in the detection or quantitative measurement of an analyte in subcutaneous fluid, the sensor being biodegradable or hydrolysable in vivo. The sensor incorporates an assay for the analyte, the readout of which is a detectable or measurable optical signal which can, when the sensor is in operation in a subcutaneous location, be interrogated transcutaneously by external optical means.

The present invention relates to a sensor for use in the measurement or 
monitoring of analytes in subcutaneous fluid using optical techniques and 
to an analyte monitoring system using this sensor. The sensor is 
particularly suitable for use in situations in which analyte levels must 
be closely monitored, for example with drugs that must be maintained 
within a narrow therapeutic window or where analyte measurements must be 
taken repeatedly, such as in long term diabetes. 
In the management of diabetes, the regular measurement of glucose in the 
blood is essential in order to ensure correct insulin dosing. Furthermore, 
it has been demonstrated that in the long term care of the diabetic 
patient better control of blood glucose levels can delay, if not prevent, 
the onset of retinopathy, circulatory problems and other degenerative 
diseases often associated with diabetes. Thus there is a need for reliable 
and accurate self-monitoring of blood glucose levels by diabetic patients. 
Currently, blood glucose is monitored by diabetic patients with the use of 
commercially available colorimetric test strips or electrochemical 
biosensors (e.g. enzyme electrodes), both of which require the regular use 
of a lancet-type instrument to withdraw a suitable amount of blood each 
time a measurement is made. On average, the majority of diabetic patients 
would use such instruments to take a measurement of blood glucose twice a 
day. However, the US National Institutes of Health recently recommended 
that blood glucose testing should be carried out at least four times a 
day, a recommendation that has been endorsed by the American Diabetes 
Association. This increase in the frequency of blood glucose testing 
imposes a considerable burden on the diabetic patient, both in terms of 
financial cost and in terms of pain and discomfort, particularly in the 
long term diabetic who has to make regular use of a lancet to draw blood 
from the fingertips. Thus, there is clearly a need for a better long term 
glucose monitoring system that does not involve drawing blood from the 
patient. 
There have been a number of recent proposals for glucose measurement 
techniques that do not require blood to be withdrawn from the patient. 
Various attempts have been made to construct devices in which an enzyme 
electrode biosensor is placed on the end of a needle or catheter which is 
inserted into a blood vessel (Wilkins E and Atanasov P, Med. Eng. Phys 
(1996) 18: 273-288). Whilst the sensing device itself is located within a 
blood vessel, the needle or catheter retains connection to the external 
environment. In practice, such devices are not suitable for use in human 
patients firstly because the insertion of a needle or catheter into a 
blood vessel poses an infection risk and is also uncomfortable for the 
patient and hence not suitable for continuous use. Secondly, devices of 
this type have not gained approval for use in human patients because it 
has been suggested that the device itself, on the end of a needle or 
catheter, may be responsible for the shedding of thromboses into the 
patient's circulation. This obviously poses a very serious risk to the 
patient's health. 
Mansouri and Schultz (Biotechnology 1984), Meadows and Schultz (Anal. Chim. 
Acta. (1993) 280: pp 21-30) and U.S. Pat. No. 4,344,438 all describe 
devices for the in situ monitoring of low molecular weight compounds in 
the blood by optical means. These devices are designed to be inserted into 
a blood vessel or placed subcutaneously but require fibre-optic connection 
to an external light source and an external detector. Again the location 
of these devices in a blood vessel carries an associated risk of promoting 
thromboses and in addition, in one embodiment the need to retain a 
fibre-optic connection to the external environment is impractical for long 
term use and carries a risk of infection. 
In the search for a less invasive glucose monitoring technique some 
attention has also been focused on the use of infra-red spectroscopy to 
directly measure blood glucose concentration in blood vessels in tissues 
such as the ear lobe or finger tip which are relatively `light 
transparent` and have blood vessels sited close to the surface of the skin 
(Jaremko J and Rorstad O, Diabetes Care 1998 21:, 444-450 and Fogt E J, 
Clin. Chem. (1990) 36:, 1573-80) This approach is obviously minimally 
invasive, but has proven to be of little practical value due to the fact 
that the infra-red spectrum of glucose in blood is so similar to that of 
the surrounding tissue that in practical terms it is virtually impossible 
to resolve the two spectra. 
It has been observed that the concentration of analytes in subcutaneous 
fluid correlates with the concentration of said analytes in the blood, 
consequently there have been several reports of the use of glucose 
monitoring devices which are sited in a subcutaneous location. In 
particular, Atanasov et al. (Med. Eng. Phys. (1996) 18: pp 632-640) 
describe the use of an implantable glucose sensing device (dimensions 
5.0.times.7.0.times.1.5 cm) to monitor glucose in the subcutaneous fluid 
of a dog. The device consists of an amperometric glucose sensor, a 
miniature potentiostat, an FM signal transmitter and a power supply and 
can be interrogated remotely, via antenna and receiver linked to a 
computer-based data acquisition system, with no need for a connection to 
the external environment. However, the large dimensions of this device 
would obviously make it impractical for use in a human patient. 
In WO 91/09312 a subcutaneous method and device is described that employs 
an affinity assay for glucose that is interrogated remotely by optical 
means. In WO 97/19188 a further example of an implantable assay system for 
glucose is described which produces an optical signal that can be read 
remotely. The devices described in WO 91/09312 and WO 97/19188 will 
persist in the body for extended periods after the assay chemistry has 
failed to operate correctly and this is a major disadvantage for chronic 
applications. Removal of the devices will require a surgical procedure. 
There remains a clear need for sensitive and accurate blood glucose 
monitoring techniques which do not require the regular withdrawal of blood 
from the patient, which do not carry a risk of infection or discomfort and 
which do not suffer from the practical disadvantages of the previously 
described implantable devices. 
Accordingly, in a first aspect the present invention provides a sensor for 
the detection or quantitative measurement of an analyte in subcutaneous 
fluid, the sensor being characterised in that it can function in a 
subcutaneous location with no physical connection to the external 
environment, said sensor incorporating an assay for said analyte the 
readout of which is a detectable or measurable optical signal, which 
optical signal can, when the sensor is in operation in a subcutaneous 
location, be interrogated transcutaneously by external optical means and 
said sensor being biodegradable or hydrolysable in vivo. 
The sensor of the invention incorporates assay means for detecting an 
analyte or for measuring the amount of an analyte, the readout of the 
assay being an optical signal. Because the sensor is located just under 
the skin, an optical signal generated in the sensor can be detected 
transcutaneously (i.e. through the skin) thus obviating the need for any 
direct connection between the sensor and the external environment. Once 
the sensor is in place in a subcutaneous location analyte measurements can 
be taken as often as is necessary with no adverse effects. This is a 
particular advantage in relation to the long term care of diabetic 
patients because if glucose measurements are taken more frequently, 
tighter control can be maintained over the level of glucose in the blood 
and the risk of developing conditions related to poorly regulated blood 
glucose, such as retinopathy, arthritis and poor circulation, will be 
reduced. 
Because the sensor of the invention does not itself contain any of the 
optical components required to interrogate the readout of the assay (these 
being provided separately and located outside the body) the sensor can 
easily be provided in a form which is injectable with minimal discomfort 
to the patient. In a preferred embodiment the components of the assay are 
incorporated into a matrix material which is permeable to subcutaneous 
fluid thereby allowing analytes such as glucose to enter the sensor by 
diffusion and to interact with the components of the assay. The matrix 
material may be an injectable formulation that forms a gel at the point of 
injection under the skin of the patient. Alternatively, the sensor may be 
formed from a solid polymeric matrix material incorporating the components 
of the assay which is again injected subcutaneously, the polymeric 
material being of a size suitable for injection through a narrow gauge 
needle to minimise the discomfort to the patient. When placed 
subcutaneously the solid polymeric material absorbs water and expands to 
form a gel thus hydrating the components of the assay. 
The device of the present invention is biodegradable or hydrolysable in 
vivo. In operation, this sensor is placed in a subcutaneous location but 
then degrades slowly over a period of time. Once the sensor has degraded 
to an extent that it has ceased to be functionally effective in the 
monitoring of analytes a fresh sensor can be simply injected or implanted 
and there is no need for the old sensor to be surgically removed. For 
reasons of safety, it is desirable that the sensor should degrade into 
material which is completely eliminated from the body. In practice, this 
requires that the sensor degrade into materials capable of passing through 
human kidney membrane to be excreted in urine or which can be metabolised 
by the body. 
As used herein the term `biodegradable` should be taken to mean that the 
sensor device of the invention is degraded within the body into materials 
which can be substantially completely eliminated from the body leaving no 
residue and that once placed in the body the sensor of the invention 
becomes degraded to extent that it is substantially completely eliminated 
from the body within a reasonable timescale relative to the active 
life-span of the reactive components incorporated into the device. In 
other words, the sensor device of the invention should ideally be 
completely eliminated from the body shortly after it has ceased to be 
effective in accurately measuring/monitoring analyte. Thus, once the 
sensor ceases to be effective a fresh sensor can simply be put in place 
and the problem of accumulating old or spent devices within the body will 
be avoided. Preferably the sensor of the invention will become degraded 
such that it is substantially completely eliminated from the body over a 
period of less than one year, more preferably over a period of several 
months. 
Materials suitable for the construction of such a biodegradable sensor 
include biodegradable block copolymers such as those described by Jeong et 
al., Nature 388: pp 860-862. Aqueous solutions of these materials are 
thermosensitive, exhibiting temperature-dependent reversible gel-sol 
transitions. The polymer material can be loaded with the components of the 
assay at an elevated temperature where the material forms a sol. In this 
form the material is injectable and on subcutaneous injection and 
subsequent rapid cooling to body temperature the material forms a gel 
matrix. The components of the assay are suspended within this gel matrix 
which thus constitutes a sensor suitable for detecting or measuring 
analytes in subcutaneous fluid. Low molecular weight analytes, such as 
glucose, can freely diffuse into the gel matrix from the surrounding 
subcutaneous fluid. This particular embodiment of the invention has the 
advantage that there is no requirement for a surgical procedure for 
implantation of the sensor. Subcutaneous injection of the sol phase 
material causes neither significant pain nor tissue damage. 
As an alternative to the gel based sensor described above the sensor may be 
constructed from a solid or gel-like biodegradable polymer matrix material 
within which the assay components are distributed. When injected or 
implanted subcutaneously this solid polymer sensor hydrates, swells and 
analyte penetrates through the structure to encounter the assay 
components. 
In a further embodiment the sensor may be constructed in the form of a 
hollow chamber, the walls of the chamber being constructed of solid 
biodegradable polymer material or of a soluble glass and defining a 
central space in which the components of the assay are contained Low 
molecular weight analytes are able to diffuse through the chamber walls or 
through a permeable end plug into the central space and thus come into 
contact with the components of the assay. The walls of the hollow chamber 
would gradually degrade over time. 
Both the solid polymer sensors and the hollow chamber sensors may be 
introduced into a subcutaneous location by implantation or injection, 
injection being preferred for sensors with a diameter of less than 2 mm. 
Both types of sensors can be formed in a wide variety of geometric shapes 
as desired prior to injection or implantation, cylindrical sensors being 
particularly preferred. 
Biodegradable materials suitable for use in the construction of the hollow 
chamber and solid polymer sensors include cross-linked proteins such as 
human albumin, fibrin gels, polysaccharides such as starch or agarose, 
poly (DL-lactide) and poly (DL-glycolide), polyanhydrides, fatty 
acid/cholesterol mixtures that form semi-solid derivates, hyaluronates and 
liquid crystals of monooliein and water. 
In a still further embodiment, the sensor may be formed as a suspension of 
microparticles of preferred dimeter &lt;100 .mu.m each of which contains the 
assay components either encapsulated inside a hollow microparticle, or 
dispersed within the material of a solid microparticle. Such a suspension 
of microparticles is readily injected subcutaneously. Preferably the 
microparticles are formed from a material which is biodegradable or 
hydrolysable in vivo. Alternatively, liposomes containing the assay 
components can be used. Liposomes of diameter 0.3 to 2.0 .mu.m have been 
shown to remain at the site of injection (Jackson A J., Drug Metab. 
Dispos. 1981 9, 535-540) so they would be suitable for use in the sensor. 
In a further embodiment the sensor comprises a plurality of empty 
erythrocytes which have been loaded with assay components and then 
injected subcutaneously. Empty erythrocytes, also known as erythrocyte 
ghosts, can be prepared by exposing intact erythrocytes to a hypotonic 
solution so that they swell and burst to release their cytoplasmic 
contents. The empty erythrocytes can then be loaded with assay components 
before allowing the plasma membranes to re-seal. 
In the preferred embodiments of the sensor (i.e. gel, solid polymer, hollow 
chamber, or microparticles) it is advantageous for the assay components to 
have a restricted diffusion in order to minimise their loss from the 
sensor. This can be achieved by ensuring that the gel or the biodegradable 
material has a pore size that permits the diffusion of low molecular 
weight analytes but not the assay components themselves. These would only 
be lost as the material or gel degrades over time. The assay components 
are preferably of high molecular weight, such as proteins or polymers, in 
order to restrict their loss from the sensor. 
There are several different mechanisms by which a biodegradable sensor can 
be degraded into materials which can be eliminated from the body, 
including hydrolysis, dissolution and cleavage of susceptible bonds by 
enzymic action, including the action of components of the immune system. 
The mechanism of degradation is ultimately dependent on the nature of the 
material from which the sensor is constructed. For example, most 
biodegradable polymer materials, such as polylactides and polyglycolides, 
are hydrolysable by water. A sensor device comprising a matrix of 
hydrolysable polymer in which the reactive components of the assay are 
embedded therefore degrades as a result of hydrolysis of the matrix, 
gradually releasing the reactive components. The overall time taken for 
the matrix to degrade will generally be dependent on the concentration of 
polymer in the matrix material and on the overall dimensions of the 
sensor. Once released from the matrix material, the reactive components of 
the assay, being protein or carbohydrate based, are either taken up by the 
liver and thereby removed or broken down in situ by the action of enzymes. 
Assays suitable for use in the sensor include reactions such as hydrolysis 
and oxidation leading to a detectable optical change i.e. fluorescence 
enhancement or quenching which can be observed transcutaneously. A 
preferred assay for use in the sensor of the invention is a binding assay, 
the readout of which is a detectable or measurable optical signal which 
can be interrogated transcutaneously using optical means. The binding 
assay generating the optical signal should preferably be reversible such 
that a continuous monitoring of fluctuating levels of the analyte can be 
achieved. This reversibility is a particular advantage of the use of a 
binding assay format in which the components of the assay are not 
consumed. Binding assays are also preferred for use in the sensor of the 
invention for reasons of safety as they cannot generate any unwanted 
products as might be generated by an enzymatic or electrochemical 
reaction. 
Preferred binding assay configurations for use in the sensor of the 
invention include a reversible competitive, reagent limited, binding 
assay, the components of which include an analyte analog and an analyte 
binding agent capable of reversibly binding both the analyte of interest 
and the analyte analog. The analyte of interest and the analyte analog 
compete for binding to the same binding site on the analyte binding agent. 
Such competitive binding assay configurations are well known in the art of 
clinical diagnostics and are described, by way of example, in The 
Immunoassay Handbook, ed. David Wild, Macmillan Press 1994. Suitable 
analyte binding agents for use in the assay would include antibodies or 
antibody fragments which retain an analyte binding site (e.g Fab 
fragments), lectins (e.g. concanavalin A), hormone receptors, drug 
receptors, aptamers and molecularly-imprinted polymers. Preferably the 
analyte analog should be a substance of higher molecular weight than the 
analyte such that it cannot freely diffuse out of the sensor. For example, 
an assay for glucose might employ a high molecular weight glucose polymer 
such as dextran as the analyte analog. 
Suitable optical signals which can be used as an assay readout in 
accordance with the invention include any optical signal which can be 
generated by a proximity assay, such as those generated by fluorescence 
energy transfer, fluorescence polarisation, fluorescence quenching, 
phosphorescence techniques, luminescence enhancement, luminescence 
quenching, diffraction or plasmon resonance, all of which are known per se 
in the art. 
The most preferred embodiment of the sensor of the invention incorporates a 
competitive, reagent limited binding assay which generates an optical 
readout using the technique of fluorescence energy transfer. In this assay 
format the analyte analog is labelled with a first chromophore 
(hereinafter referred to a the donor chromophore) and the analyte binding 
agent is labelled with a second chromophore (hereinafter referred to as 
the acceptor chromophore). It is an essential feature of the assay that 
the fluorescence emission spectrum of the donor chromophore overlaps with 
the absorption spectrum of the acceptor chromophore, such that when the 
donor and acceptor chromophores are brought into close proximity by the 
binding of the analyte analog to the analyte binding agent a proportion of 
the fluorescent signal emitted by the donor chromophore (following 
irradiation with incident radiation of a wavelength absorbed by the donor 
chromophore) will be absorbed by the proximal acceptor chromophore, a 
process known in the art as fluorescence energy transfer, with the result 
that a proportion of the fluorescent signal emitted by the donor 
chromophore is quenched and, in some instances, that the acceptor 
chromophore emits fluorescence. Fluorescence energy transfer will only 
occur when the donor and acceptor chromophores are brought into close 
proximity by the binding of analyte analog to analyte binding agent. Thus, 
in the presence of analyte, which competes with the analyte analog for 
binding to the analyte binding agent, the amount of quenching is reduced 
(resulting in a measurable increase in the intensity of the fluorescent 
signal emitted by the donor chromophore or a fall in the intensity of the 
signal emitted by the acceptor chromophore) as labelled analyte analog is 
displaced from binding to the analyte binding agent. The intensity of the 
fluorescent signal emitted from the donor chromophore thus correlates with 
the concentration of analyte in the subcutaneous fluid bathing the sensor. 
An additional advantageous feature of the fluorescence energy transfer 
assay format arises from the fact that any fluorescent signal emitted by 
the acceptor chromophore following excitation with a beam of incident 
radiation at a wavelength within the absorption spectrum of the acceptor 
chromophore is unaffected by the fluorescence energy transfer process. It 
is therefore possible to use the intensity of the fluorescent signal 
emitted by the acceptor chromophore as an internal reference signal, for 
example in continuous calibration of the sensor or to monitor the extent 
to which the sensor has degraded and thus indicate the need to implant or 
inject a fresh sensor. As the sensor degrades, the amount of acceptor 
chromophore present in the sensor will decrease and hence the intensity of 
fluorescent signal detected upon excitation of the acceptor chromophore 
will also decrease. The fall of this signal below an acceptable baseline 
level would indicate the need to implant or inject a fresh sensor. 
Competitive binding assays using the fluorescence energy transfer technique 
which are capable of being adapted for use in the sensor of the invention 
are known in the art. U.S. Pat. No. 3,996,345 describes immunoassays 
employing antibodies and fluorescence energy transfer between a 
fluorescer-quencher chromophoric pair. Meadows and Schultz (Anal. Chim. 
Acta (1993) 280: pp 21-30) describe a homogeneous assay method for the 
measurement of glucose based on fluorescence energy transfer between a 
labelled glucose analog (FITC labelled dextran) and a labelled glucose 
binding agent (rhodamine labelled concanavalin A). In all of these 
configurations the acceptor and donor chromophores/quenchers can be linked 
to either the binding agent or the analyte analog. 
An alternative to the fluorescence energy transfer is the fluorescence 
quenching technique. In this case a compound with fluorescence quenching 
capability is used instead of the specific acceptor chromophore and the 
optical signal in a competitive binding assay will increase with 
increasing analyte. An example of a powerful and non-specific fluorescence 
quencher is given by Tyagi et al. Nature Biotechnology (1998) 18: p 49. 
The sensor of the invention can be adapted for the detection or 
quantitative measurement of any analyte present in subcutaneous fluid. 
Preferred analytes include glucose (in connection with the long-term 
monitoring of diabetics), urea (in connection with kidney disease or 
disfunction), lactate (in connection with assessment of muscle performance 
in sports medicine), ions such as sodium, calcium or potassium and 
therapeutic drugs whose concentration in the blood must be closely 
monitored, such as, for example, digoxin, theophylline or 
immunosuppressant drugs. The above analytes are listed by way of example 
only and it is to be understood that the precise nature of the analyte to 
be measured is not material to the invention. 
The sensor is interrogated transcutaneously using optical means i.e. no 
physical connection is required between the sensor and the optical means. 
When the sensor incorporates a competitive, reagent limited, binding assay 
employing the technique of fluorescent energy transfer, the optical means 
should supply a first beam of incident radiation at a wavelength within 
the absorption spectrum of the donor chromophore and preferably a second 
beam of incident radiation at a wavelength within the absorption spectrum 
of the acceptor chromophore. In addition, the optical means should be 
capable of measuring optical signals generated in the sensor at two 
different wavelengths; wavelength 1 within the emission spectrum of the 
donor chromophore (the signal generated in connection with the measurement 
of analyte and wavelength 2 in the emission spectrum of the acceptor 
chromophore (which could be the analyte signal or the internal reference 
or calibration signal). 
Optical means suitable for use in remote interrogation of the device of the 
invention include a simple high-throughput fluorimeter comprising an 
excitation light source such as, for example, a light-emitting diode 
(blue, green or red &gt;1000 mCa), an excitation light filter (dichroic 
filter), a fluorescent light filter (dichroic or dye filter) and a 
fluorescent light detector (PIN diode configuration). A fluorimeter with 
these characteristics may exhibit a sensitivity of between picomolar to 
femtomolar fluorophore concentration. 
A suitable fluorimeter set-up is shown in the accompanying FIG. 1 and 
described in the Examples included herein. The fluorimeter separately 
measures the following parameters: 
At wavelength 1 (donor chromophore) 
Excitation light intensity, I(1,0) 
Ambient light intensity, I(1,1) 
Intensity of combined fluorescent and ambient light, I(1,2) 
At wavelength 2 (acceptor chromophore) 
Excitation light intensity, I(2,0) 
Ambient light intensity, I(2,1) 
Intensity of combined fluorescent and ambient light, I(2,2) 
Measurements are taken by holding the fluorimeter close to the skin and in 
alignment with the sensor. When making transcutaneous measurements of the 
fluorescent signals generated in the sensor it is necessary to take 
account of the absorption of signal by the skin, the absorptivity of human 
skin is found by experiment to be lowest in the range from 400 nm to 900 
nm. The final output provided is the normalized ratio between the 
fluorescent intensity from the two fluorophores, defined by the following 
relation (Equation 1): 
EQU (I(1,2)-I(1,1)) I(2,0) 
EQU (I(2,2)-I(2,1)) I(1,0) (1) 
In a third aspect the invention provides an analytical system suitable for 
the detection or quantitative measurement of an analyte in subcutaneous 
fluid, said analytical system comprising, 
(i) a sensor for the detection or quantitative measurement of an analyte in 
subcutaneous fluid in accordance with the first aspect of the invention; 
(ii) optical means suitable for the transcutaneous interrogation of the 
sensor of (i). 
In a fourth aspect the invention provides a method of detecting or 
quantitatively measuring an analyte in the subcutaneous fluid of a mammal, 
which method comprises the steps of, 
(a) injecting or implanting a sensor for the detection or quantitative 
measurement of an analyte in subcutaneous fluid in accordance with the 
first aspect of the invention; 
(b) allowing the assay of said sensor to reach thermodynamic equilibrium; 
(c) interrogating the readout of said assay using optical means; and 
(d) relating the measurement obtaining in (c) to the concentration of 
analyte. 
The final output from the optical means (e.g. the fluorimeter) as given by 
Equation 1 above is converted to analyte concentration preferably by means 
of a computer using calibration data which can be obtained based on the 
principles set out below. 
A calibration curve can be established empirically by measuring response 
versus analyte concentration for a physiologically relevant range of 
analyte concentrations. Preferably this takes place in vitro as part of 
the production of the sensor device. The calibration procedure can be 
simplified considerably by using the mathematical relation between 
response and analyte concentration in a competitive affinity sensor which 
is derived as follows: 
The response of a competitive affinity sensor is governed by the reactions: 
EQU RC-R+C 
EQU RL-R+L 
designating the dissociation of the complexes RC and RL, formed by the 
combination of analyte binding agent (R) with analyte (L) or analyte 
analog (C). 
The corresponding dissociation equilibrium constants are: 
##EQU1## 
where C designates the number of moles of the species in the sensor 
divided by the sensor volume. Using this measure of concentration both 
immobilized species and species in solution are treated alike. 
The mass balance equations are: 
EQU T.sub.C =C.sub.C +C.sub.RC 
for total analyte analog concentration and, 
EQU T.sub.R =C.sub.R +C.sub.RC +C.sub.RL 
for total analyte binding agent concentration. 
Using the expression above, the relation between response and analyte 
concentration is derived: 
##EQU2## 
By using this relation the amount of data necessary for the calibration 
can be reduced to two key parameters: Total analyte binding agent 
concentration and total analyte analog concentration. The calibration 
curve is thus determined by two points on the curve.

EXAMPLE 1 
A glucose assay according to Meadows and Schultz (Talanta, 35, 145-150, 
1988) was developed using concanavalin A-rhodamine and dextran-FITC (both 
from Molecular Probes Inc, Oregon, USA). The principle of the assay is 
fluorescence resonance energy transfer between the two fluorophores when 
they are in close proximity; in the presence of glucose the resonance 
energy transfer is inhibited and the fluorescent signal from FITC 
(fluorescein) increases. Thus increasing fluorescence correlates with 
increasing glucose. The glucose assay was found to respond to glucose, as 
reported by Schultz, with approximately 50% recovery of the fluorescein 
fluorescence signal at 20 mg/dL glucose. Fluorescence was measured in a 
Perkin Elmer fluorimeter, adapted for flow-through measurement using a 
sipping device. 
EXAMPLE 2 
The glucose assay components of Example 1 were added to stirred solutions 
(1 ml) of 1%, 1.5% and 2% w/v of a low melting temperature agarose (Type 
IX, Sigma, St. Louis, USA) at 45.degree. C. After dispersal, the 
temperature was reduced to 20.degree. C. and the stirring was stopped. 
When the gel had formed (after approximately 3 hours) it was placed in a 
ceramic mortar and ground to a particle size of 50 to 100 .mu.m, by visual 
reference to a polystyrene bead preparation with the same diameter. The 
particle preparation was suspended in 0.9% w/v saline and filtered through 
a nylon mesh to remove the larger particles. The particles that passed 
through the mesh were then centrifuged in a bench centrifuge at 500 g and 
the supernatant containing fines was discarded. During the process the 
particles retained their fluorescence by visual inspection and by 
measurement of the rhodamine fluorescence in the Perkin Elmer fluorimeter. 
Adding glucose at 20 mg/dL to a sample of the suspended particles resulted 
in a rise in the fluorescein fluorescence signal over a 30 minute period. 
Thus the assay components contained within the agarose gel were responsive 
to glucose. 
The glucose assay chemistry components are inherently biodegradable (or 
excretable) in the human body since they are based on peptide and 
carbohydrate materials. Degradation by enzyme digestion and/or simple 
transport to the liver or kidney will ensure that the assay components 
will be removed from the body following implantation or subcutaneous 
injection. 
EXAMPLE 3 
1 ml samples of the agarose particle suspensions containing glucose assay 
reagents (described in Example 2) were placed in a water bath at 
37.degree. C. to simulate human body temperature. Over the following six 
weeks the particle structure was lost as the structures degraded, with the 
higher concentration gels exhibiting the slowest degradation. The light 
scattering signal measured using the fluorimeter also fell as the 
particles dispersed, indicating degradation of the particles. This 
experiment simulates conditions in the human body--it is expected that the 
particles will eventually be cleared from the site of injection and the 
assay chemistry components will be degraded either in situ or transported 
to the liver for further processing, or excreted in the urine. 
EXAMPLE 4 
A fibre optic spectrometer was assembled as follows: 
The optical part of a fibre optic fluorimeter was made from standard 
components on a micro bench. The setup, comprising a red LED as light 
source, lenses, dichroic beamsplitter and filters and detector diodes, was 
as shown in FIG. 1. Briefly, the fluorimeter comprises a light-emitting 
diode (1) providing an excitation light beam which passes through a 
condenser (2) containing an excitation filter (3) and is incident upon a 
beamsplitter (4). Part of the excitatory beam is thereby deflected into 
launching optics (5) and enters an optical fibre (6). When the fluorimeter 
is in use in the interrogation of a subcutaneously located sensor the end 
of optical fibre (6) is positioned close to the surface of the skin, in 
alignment with the subcutaneous sensor, so that beam of excitatory light 
is incident upon the sensor. A portion of the optical signal emitted from 
the sensor following excitation enters the optical fibre (6) and is 
thereby conveyed into the fluorimeter where it passes through a blocking 
filter (8) and is measured by a signal detector diode (7). The fluorimeter 
also contains a reference detector diode (9) which provides a reference 
measurement of the excitatory light emitted from the LED (1). The ends of 
a 1 m long Ensign Beckford optical fibre, 0.5 mm in diameter, numerical 
aperture of 0.65, were around to a mirror finish using diamond paste on 
glass paste. One end of the fibre was mounted in an X Y Z holder in front 
of a 20 x microscope objective. The diodes (LED (1) and detector diodes 
(7) and (9)) were connected to a custom made driver/amplifier circuit as 
shown in FIG. 2. The circuit comprises a sender (10), current amplifiers 
(11) and (12), multiplexers (13) and (14), integrators (15) and (16) and 
analog divider (17). The driver circuit was set to drive the LED (1) at 
238 Hz and the signals from the detector diodes (7) and (9) were switched 
between ground and the storage capacitors (integrator with a time constant 
of 1 second) synchronised with the driver signal. The two integrated 
signals correspond to background-corrected fluorescent signal and 
background corrected excitation light level (LED intensity). The former 
divided by the latter was supported by an analog divider as shown in FIG. 
2. For test purposes, the distal end of the fibre (6) was dipped into 
dilute solutions of rhodamine and the optics were adjusted for maximum 
signal from the analog divider. 
The fluorimeter/spectrometer is battery operated (typical power consumption 
150 mA at 9V) and for convenience can be constructed in the shape and 
dimensions of a pen. 
EXAMPLE 5 
1.5% w/v agarose particles of approximately 50 .mu.m diameter containing 
the assay components (as described in Example 2) were washed several times 
by centrifuging and resuspending in 0.9% w/v saline solution. This washing 
procedure removed excess reagents that were not trapped within the gel 
structure. The particles remained highly fluorescent during this process. 
Then the particle suspension was loaded into a standard disposable syringe 
(Becton Dickinson, USA) and injected subcutaneously under the skin on the 
back of the hand of a human volunteer. A fibre optic spectrometer (see 
Example 4) was directed at the skin and a rhodamine fluorescence signal 
was obtained, indicating that transdermal measurements can be made on 
implanted biodegradable sensors. 
EXAMPLE 6 
A glass capillary of dimensions 10 mm.times.2 mm composed of a phosphate 
based soluble glass (Pilkington CRS, Wrexham, UK) was part-filled with a 
solution of glucose assay reagents. The ends of the capillary were then 
sealed by dipping in a solution of 1% w/v low melting temperature agarose 
heated to 45.degree. C. and then cooling to room temperature (22.degree. 
C.). Both ends of the capillary were sealed in this manner. The sealed 
capillary was placed on a flat surface and the fibre optic spectrometer 
(see Example 4) was positioned to take a fluoresence reading of the 
contents. The strong rhodamine fluorescence indicated that reagents had 
been incorporated into the capillary. The capillary was then placed in a 
top-stirred solution of 20 mg/dL glucose in 0.9% w/v saline for 16 hours 
at room temperature (22.degree. C.). The capillary was then removed from 
the glucose solution, the external surface was dried and the capillary was 
placed on a flat surface. A reading with the fibre optic spectrometer 
indicated that an increase in fluorescein fluorescence had occurred, 
showing that glucose had penetrated the capillary by diffusion through the 
agarose gel caps. The capillary was then replaced in the top-stirred 0.9% 
w/v saline solution at 37.degree. C. After 200 hours the structure of the 
capillary had begun to collapse as the phosphate glass wall was breached 
and the contents were released into the surrounding fluid. After a further 
20 days no residue of the glass capillary structure remained and the 
agarose gel caps had also dissolved. Thus the entire capillary based 
sensor was shown to be fully degradable under physiological conditions.