Glucose monitoring: state of the art and future possibilities

Med. Eng. P/y. Vol. 18. hio. 4, pp. 273-288, 1996 Copyright 0 1996 Ekvier Science Ltd for IPEMB Printed in Great Britain. All rights reserved I Y&4535/96 $15.00 + 0.00

1350-4533(95)000461 ELSEVIER

Review

Glucose monitoring: state of the art and future E. Wilkins

and P. Atanasov

Department Albuquerque, Received

possibilities

of Chemical and Nuclear NM 87131, USA

6 February

1995,

accepted

19 April

Engineering,

University

of New Mexico,

1995

ABSTRACT This article reviews the development of glucose monitoring techniques and approaches during the last &cade. The predominance of the electrochemical measuring principles reported in the literature makes them a focus of this work. Biosensors are still in the main stream of the research interest of most teams due to their high selectivity fo1 glucose determination. Systematization and classafication of the glucose monitoring @zciples and types of glucose sensors is shown. ‘The review gives a brief &scrip&ion of the basic operational principles of the most popular types of glucose biosensors, providing an enhanced bibliography of the original works of the main groups in establishing or significantly contributing to the deoetopnent of the particular type of glucose biosensor. Different design approaches are overviewed including needletype sensors, sensors biosensors with microdialysis sampling technique. recharging the sensor in situ white implanted is lifespan of the system and ultimately, it could lead ,sensor. Copyright 0 1996 Elsevier Science Ltd for

Keywords:

Glucose mellitus

diabetes Med.

Eng.

Phys.,

monitoring,

1996,

Vol.

glucose

18, 273-288,

for chronical implantation and the combination of the glucose The authors approach for re$.acing of the spent enzyme and thus widely discussed. This approach p-ovides a way to increase the to rare transcutaneous interventions for rehlling of the implanted IPEMB.

sensors,

biosensors,

Diabetes mellitus is a complex endocrine metabolic disorder that results from a total or partial lack of insulin. The earliest manifestation of the disease is the loss of control of the blood glucose level. There was a debate in the 70s as to whether there was a need for a long-term implantable glucase sensor for diabetics’. Recent long-term studies have conclusively demonstrated that if glucose levels can be tightly regulated to be within the normal physiological range, then diabetic complications can be controlled*. Glucose fluctuating within the normal physiological range of 110 mg/dL + 25 mg/dL can be considered an acceptable level of control. The concept of maintaining normal glucose physiological level led to the development of a series of glucose sensing devices suitable for measurCorrespondence

to: Nuclear Mexico,

Ebtisam Engineering, Albuquerque,

Wilkins, Far& NM

Professor, Engineering 87131-1341.

Department Center, USA.

electrodes,

implantable

sensors,

June

INTRODUCTION

Chemical and versity of New

enzyme

of Uni-

ing glucose levels in physiological fluids both in uiuo and in vitro. These sensors were based on electrochemical principles and employed enzymes as biological components for molecular recognition. Recently, many other methods and techniques have been proposed for glucose sensin$. For continuous long-term monitoring in viva, however, the electrochemical biosensor is still the focus of much research4. Several new techniques have evolved for glucose analyses in clinical practice5 as well as in biotechnology6 and in food industry’. This wide field of applications have inspired much of the glucose sensor development and diversification during the last decade. The motivation of providing diabetes control remains a dominant force behind the research efforts of numerous investigators from different academic disciplines and industry, efforts being focused on the implantable shortand long-term glucose sensor . The studies in this field consist of three parts: (a) selection of the measuring principle and basic sensor develop-

Glucose monitoring:

E. Wilkins

and P. Atanasov

ment; (b) evaluation of the sensor in vitro and its refinement; (c) in viva tests of the sensor performanimals and ultimately ance in laboratory human trials. A variety of methods are currently used in analytical practice for glucose determination and concentration measurements3~8. Many of these methods are motivated by the need for automated, fast and accurate clinical, hospital and biomedical research applications. All of these methods must have the following characteristics: l

fast response

l

accuracy

l

sensitivity

0

range

l

stability

change in glucose concentration must be detected within l-5 minutes, depending on the specific application; glucose level must be measured with minimum errors due to the presence of interfering species or changes in physiological parameters (maximum deviation should not exceed 10 mg/dL); the signal to noise ratio must be large, and a detectable signal must result from small (0.1 mM or 2 mg/dL) appr. changes in glucose concentration; all glucose concentrations in the physiological (normoglycemia) and pathophysiological range (hype and hyperglycemia) from 1 to 30 mM (20 to 600 mg/dL) must be measurable; depending on the specific application, the signal due to glucose must not deviate more than +5% of its average value during the operational time of the measuring instrument.

Methods for glucose concentration measurement directly in samples of physiological fluid with minimum pretreatment and separation are preferable. Development of methods and techniques of glucose concentration measurement starts with use of dedicated laboratory equipment requiring specially trained technicians, and evolves toward simplification and user-friendliness. For example, automated glucose analysers have been developed for operation by non-technical personnel. Miniature glucose monitors are becoming popular for self-monitoring by patients”. GLUCOSE

MEASURING

PRINCIPLES

The measuring principles can be classified into two main groups based on the interaction between the patient’s body and analytical devices employed: invasive and non-invasive (Fz@z I). Invasive techniques have intimate mechanical contact of the sensing part of the device with biological tissues or fluids. Non-invasive measure-

274

Glucose

f

measuring

principle

7

Invas Serum/Plasma Interstitial fluid Noninvasive

direct

me

Glucose

Figure

1

Glucose

measuring

monitoring

and monitoring

principles

ments obtain information without mechanical intervention, using characteristic properties (spectral, optical, thermal, electro-magnetic, etc.) of the analyte (glucose) which can be detected remotely. Non-invasive

methods

For glucose concentration measurements and monitoring, the use of the near-infrared (NIR) spectra of the analyte has been proposed”. The NIR region of the electromagnetic spectrum extends from the end of the visible spectrum, at a wavelength of approximately 700 nm, to the beginning of the fundamental infrared (IR) absorption bands at 2500 nm. Absorptions occurring in the NIR are most often associated with overtone and combination bands of the fundamental molecular vibrations of -OH, -NH, and -CH functional groups that are indicated in the mid-IR region. As a result, most biochemical species will exhibit specific absorptions in the NIR. In addition, a few weak electronic transitions of organometallic molecules, such as haemoglobin, myoglobin, and cytochrome, also appear in the NIR. These highly overlapping, weakly absorbing bands were initially perceived to be too complex for interpretation and too weak for practical application. However, improvements in instrumentation and advances in multivariate chemometric data analysis techniques, which can extract vast amounts of chemical information from NIR specul results to be obtained from tra, allow meanin complex spectra’ F. Direct spectroscopic measurements of unmodified body fluids or tissues using the more traditional ultraviolet, visible, and IR regions of the

Gucos~

spectrum are impractical because of limited peninterfering absorptions and etration depths, excessive scattering with inhomogeneous samples. In contrast, the weak absorption of NIR radiation by most biochemicals makes NIR spectroscopy useful because body fluids and soft tissues are relatively transparent at these wavelengths”.“. Glucose NIR measurements are usually carried out in the spectra region from 4250 to 660 cm-l “‘,I. In this region a correlation between optical absorbance and glucose concentration has been established in model fluids’.+” and in body fluids”-I”.I”~” The site of the measurement is usually a finger”’ or other accessible extremity (e.g. the inner lip’” or oral mucosa’“). Devices based on this principle are fast, extremely convenient and easy to use by semi-skilled technicians. However, they are not particularly accurate even in the normal physiological range 2o. Cross-validated standard errors of prediction for glucose concentration range from 39 mg/dL12 to 55 mg/dL18 when normoglycaemic levels are studied. Moreover, a subjectdependent concentration bias has been reported’“. The temperature sensitivity of water absorption bands in the glucose measuring region creates large baseline variations in spectra and can be a significant source of error in practical clinical assays’“. The non-invasiveness of NIR-based methods prevents possible infections and avoids surgical intervention, and devices can be used by many patients in various settings, e.g. in an office, outpatient clinic, home, etc. Ultimately they can possibly be developed for use in self-monitoring applications by the diabetic patient. At present, however, non-invasive devices based on NIR absorption suffer from low sensitivity and thus low accuracy of measurement. From the analytical point of view this method at present is an estimation technique rather than an exact analytic measurement. Poor selectivity of such devices has also been reported, probably caused by NIR absorption by other body chemicals*l. The devices can also be affected by personal characteristics of different patients at the measurement site - skin location, skin and tissue structure with a resulting decrease in the accuracy of measurement. Another type of non-invasive technique for glucose monitoring involves the direct analysis for glucose of samples of physiological fluids obtained from the patient such as saliva, urine, sweat, or tears”2.2”. This approach combines the use of welldeveloped methods for clinical and analytical glucose concentration measurements with the advantages of the non-invasive techniques. However, there are several factors, such as diet and exercise, which can affect glucose levels in the aforementioned fluids, not necessarily due to any pathology. In general there is no strong correlation established between glucose concentration in the blood and in excreted fluids. Thereafter the lagtime between blood and excreted fluid glucose concentrations can be large enough to render such measurements inaccurate.

Invasive

monitoring:

E. Wilkm

and

P. Atanasov

methods

Invasive measurements at the present time are clinical and home glucose popular, dominating monitoring practice 4-‘Lfi. These techniques can be classified by how samples are obtained and analysed; either by invasive insertion of a sensor, or by puncturing tissues to obtain the sample. Most of the glucose measuring devices in routine use in hospitals, clinics and at home are items of equipment external to the patient’s body2“. In all cases a sample of some physiological fluid (blood, serum, interstitial fluid, etc) is obtained by aspiration or by puncture and the analysis performed by the external measuring instrument. Analytical techniques and devices for in vitro glucose measurements have a high level of accuracy (error can be lower than 1%) . Many of these routine methods are accepted as standards for comparison with new devices under develop men?. They are available in a wide range of styles and types, from automated and computerized analytical machines for clinical laboratory use, through non-sophisticated glucose analysers, to pocket-sized devices for domestic use by the diabetic patient’“. Management of diabetes mellitus currently relies on these methods to control the disease and minimize complications. Use of the above in vitro methods for monitoring glucose have two main disadvantages. Firstly, as an invasive method, sampling even a minimal amount of blood on a regular daily basis is associated with risks of infection, nerve and tissue damage, and discomfort to the patients. Secondly, in the case of dynamic changes of glucose concentration, very frequent or even continuous measurement of blood glucose time changes are required. For research purposes these methods are unsuitable, as again continuous measurement of blood glucose time profiles is often required. Continuous

glucose

monitoring

Intensive treatment of diabetes, prevention of complications2, continuous glucose level data”, as well as glycemia research28 all require continuous glucose monitoring”“. This monitoring could be external, as when samples of physiological fluid are continuously withdrawn from the patient, or internal, when the sensing device is in intermediate contact with the fluid in vim (Figure I). Since the pioneering reports on direct measurements of glucose concentration3”,“‘, the ultimate goal has been the development of a device suitable for continuous monitoring in km1.R2-41. The research was inspired by the successful development of the artificial pancreas27,4’ and the promising initial human trials using the im lantable insulin delivery systems (insulin pumps) Ll . The motivation behind this effort was the hypothesis that the maintenance of normoglycaemia could prevent complications due to diabetes mellitus. This hypothesis has been confirmed after years of clinical trials with patients living a normal as possible life”. Therefore, the artificial pancreas with a glucose sensor as a feedback system can provide

275

Glucose monitoring:

E. Wilkins

and P. Atanasov

care and improvement of quality of life for millions of patients, as well as significant reductions in the cost of care delivery . Numerous studies are focussed on the need for an implantable glucose sensor to close the loop ~-p~;;~~s~um~, and obtain a complete arti*28 3 x4* 56. Devices with this function can be classified into three groups, based on the size and site of application4*:

Glucose sensors

bedside units, suitable for hospital or intensive care use only; wearable modules, attachable to the patient’s body (arm, belt, waist or leg) ; and implantable devices (the artificial pancreas). The bedside unit is a big and complex device, used on a cart near the patient. Insulin infusion and blood extraction catheters link the patient to the device. A modified automated laboratory glucose analyser with adequate speed or a miniature sensor system is used for continuous blood glucose monitoring at the required frequency and flow rates4*. Combinations of this bed-side insulin delivery system with miniature implantable sensors have been described”” Wearable units are presently used in research for short term glucose monitoring during glycaemia changes and in the development of the implantable glucose sensor57-60. They employ either direct coupling with a sensor3940, usually subcutaneously implanted44-57, or with a microdialysis sampling device used to supply the sensor with fluid through the skin”‘-““. Implantable devices will be ultimately used in diabetic patients to maintain normoglycaemia through internal insulin infusion, or as alarm devices for continuous monitoring of glucose level. They will be alternatives to the domestic use of insulin injections and glucose self-monitoring kits, strips or portable glucometers70-7”. Development of implantable insulin pumps is at a very advanced stage”-“, with units being implanted for S-10 years in patientsso. However, these implantable pumps lack a continuously functioning implantable glucose sensor with longterm stability. At present, the pumps which are in clinical trials deliver insulin following an external command from the physician or from the patient, or based on some programmed cycle connected with the patient’s diet7”. Closing the loop by means of a glucose sensor which can continuously monitor glucose level, will provide feedback on the exact amount of insulin needed at any moment, and will convert these insulin delivery systems into a complete internal artificial pancreas. Glucose

sensors

Glucose measuring devices (glucose sensors) could be categorized based on the physical principle of the transducer usedsl (see Figure 2) : l l

276

electrochemical sensors; piezoelectric sensors;

I Enzyme sensors

Oxygen electrode Mediated electron

Figure

l l l

2

Types

of glucose

transfer

sensors

thermoelectric sensors; acoustic sensors; and optical sensors.

Electrochemical sensors. In the electrochemical

sensors the electrical signal is a direct consequence of some (chemical) process occurring at the transducer/analyte interface. Research is being carried out on several types of potentially implantable glucose sensors: electrocatalytic sensors based on direct electro-oxidation of glucose on noble metal electrodesand biosensors combining the selectivity of glucose-specific enzymes with the versatility and simplicity of electrochemical electrode systemP. Electrocatalytic sensors have specific problems associated with their low analyte selectivity. The general approach in overcoming these problems has been previously reviewed8s. An approach used to achieve high selectivity is based on the molecular recognition principle, combining biologically active components (enzymes, antibodies, cells, tissues or microorganisms) with some physical transducer. Biosensors used in or proposed for in uiuo experiments are direct enzyme biosensors or affinity sensors based on enzyme labeled immunoassays81.83. In both cases biological catalysts - enzymes are used as a molecular recognition element in glucose sensing. Immunoassays provide the ability of sensing extremely low amounts of analyte. In diabetes related studies, however, this is not absolutely necessary as far as the concentrations of physiological and pathophysiological interest are

achievable measurably direct enzyme assaf’.

by using

less complicated

Piezoekctric, thermoelectric and acoustic sensors. In piezoelectric, thermoelectric and acoustic (surface acoustic wave, SAW) sensors used for glucose measurement an enzyme-catalysed reaction is used to create a measurable change in a physical parameters detected by the transducer. The development of these sensors is at an early laborator) stagex’. Optical sensors. Optical sensors, especially optical biosensors, based on the fast-developing optoelectronic techniques are one of the possible future alternatives for glucose sensing. They are based on changes in some optical parameter due to enzyme reactions or antibody-antigen bonding at the transducer interface. Based on the mam process used they can be divided into two categories: enzyme optrodes and optical immunosensorsX’. Based on the nature of the monitored process they are densitometric, refractometric or colorimetric devices. At present, none of them meets the selectivity requirements to sense and accurately measure glucose in real physiological fluids. The field is presently dominated by the electrochemical sensor, due to the relative simplicity of electrochemical measuring principles and the advanced state of development of this biosensor.

ELECTROCHEMICAL BIOSENSORS

GLUCOSE

The main method for the construction of electrochemical glucose biosensors is the use of enzyme electrodes in which the biological component (enzyme) is incorporated as a part of the transducer design. Glucose biosensors are generally based on the enzyme glucose oxidase (GOD). This enzyme catalyses the oxidation of Pn-glucose by molecular oxygen producing gluconolactone and hydrogen peroxidex4. It is a two stage enzyme process typical for the class of oxidases. The process consists of enzymatic oxidation of glucose by the enzyme in which the co-factor flavin-adenine dinucleotide (FAD) is reduced to FADH, (reaction 1) followed oxidation of the enzyme co-factor bY (regeneration of the bio-catalyst) by molecular oxygen with formation of hydrogen peroxide (reaction 2) : /3-o-glucose

+ GOD(FAD)

+ GOD (FADH,) GOD(FADH,)

+ 0, -

glucono-filactone

-

glucono-Nactone

.

(1)

GOD(FAD)

+ H,O -

gluconic

+ H,O,.

(2)

acid.

(:3)

The gluconolactone produced in the reaction 1 is hydrolysed (reaction 3) in aqueous media to gluconic acid (reaction 3) so the overall reaction is usually expressed as: P-u-glucose

+ 0, + H,O

-

gluconic

acid + H,O,. (4)

Electrochemical biosensors are constructed on the amperometric principle based on the oxidation or reduction of electrochemically active substances involved or produced in reactions l3”. Another possibility is to measure the changes in local pH due to the gluconic acid produced in reaction 3 at a potentiometric sensor, usually a coated wire pH-selective electrode or an ion selective field effect transistor (ISFET)X”. Electrical resistance changes during the overall process (reaction 4) are used as a basis for conductometric biosensor?‘. At present, potentiometric and conductometric glucose biosensors proposed fbr in viva monitoring have vet-v limited applicability due to numerous interfering processes caused by components of the physiological environment other than glucose. Several potentiometric glucose sensors (coated wire sensors) have been proposed for implantable use’:‘. Coated wire sensors are in general easy to fabricate, and are suitable for miniaturization to diameters of 50-200 pm. They could be used in combination with a _ standard cardiographic (EKG) reference electrode’*. The main disadvantage of these sensors is associated with their low sensitivity. The signal of the potentiometric glucose sensors is based on the Nernst equation, which gives a logarithmic dependence of the potential change (slgnal) on analyte concentratior?‘. This could be very useful when the amplitude of analyte concentration changes is several orders of magnitude (as in the case of pH measurements). In physiological fluids, however glucose concentration changes by no more than one order of magnitude, so a sensor with linear signal vs. concentration proportionalit). is preferable to a logarithmic response. Introducing microelectronic techniques in sensor development through the use of ion selective field effect transistors (ISFET) as a basis for the sensor design is promising in the sense of miniaturization and integration of the transducer with some part of the associated electronic circuitry into a single chipxx. Corrosion of the semiconductor material surface in saline (and in physiological) solutions is presently, a problem in introducing these devices for in 711710 trials. This problem may be overcome by surface modification by passivation or by coating”. Being potentiometric measuring devices, biosensors based on enzyme FETs have the same general disadvantages as the coated wire sensors. It appears from the development of biosensors that the ampcrometric techniques are most useful. The preference for this concept for implantable applications is primarily due to the possibility to obtain a signal linearly related to analyte concentration. The linearitv of the sensor response is of great importance ir; the repeated recalibration over time of implantable glucose sensors”‘. Microprocessor evaluation of the nonlinear glucose calibration curve allows recalibration of the sensor in viva if the non-linear function is known. If‘ the limiting process in signal generation are the cnTymatic reactions (I, 2) the dependence of

277

Glucose monitoring:

E. Wilkins

and P. Atanasou

the signal vs. glucose concentration is non-linear according to the Michaelis-Menten kinetics. When the sensor operates in a diffusion limited mode, linear proportionality between the signal and analyte concentration occurs. The amperometric glucose biosensors are usually categorized into three different classes (sometimes called ‘generations’) based on the mechanism of the electrochemical (charge transfer) reaction involved85. In the first class are biosensors based on glucose oxidation by GOD (reactions l-3) in natural conditionssg. In this case the increase in hydrogen peroxide concentration or the decrease in oxygen concentration due to the reaction (4) are detected electrochemitally, both being proportional to the glucose concentration. In the second class of amperometric glucose biosensors low-molecular weight compounds are used as mediators in the process of enzyme oxidation (reaction 2) and signal generation at the electrode surfacego. The third class of amperometric biosensors is based on direct electron transfer between the enzyme and an electrode with special propertiesgl. Oxygen

electrode-based

glucose

biosensors

Use of the oxygen detecting electrode as a basic transducer for glucose monitoring was the first electrochemical method developed30*31. In this case the amperometric signal is a result of the electrochemical reduction of oxygen on a cathodically polarized platinum electrode:

gen concentration, which has to be monitored. This results in complexity of the construction of the sensors, requiring two electrodes and making miniaturization a difficult taskgg-lo’. Further, more complex electronic circuitry is required. Hydrogen

peroxide

electrode-based

biosensors

In the case of the hydrogen peroxide electrode based sensors, the signal is due to the oxidation of the hydrogen peroxide at a catalytic (usually platinum) anode at potential from +0.6 to +0.8 V vs. Ag/AgCPg: H,O,

2 0, + H,O

+ 2e-.

(7) The most important advantage of the hydrogen peroxide electrode based sensors is their ease of fabrication and the possibility of constructing them in small sizes even when simple technology is used. A linear dependence of the amperometric signal is obtained when the mass transfer of both active species (hydrogen electrochemically peroxide) and glucose are the limiting processes. This is achieved by use of a variety of diffusion membranes in the biosensor construction. The main disadvantage of the hydrogen peroxide electrode based amperometric biosensors is that they suffer from poor electrochemical selectivity due to electro-oxidation of species other than lucose present in physiological fluids73,1 #*,ll’. This problem can be corrected to a large extent by the use of semi-permeable membranes having specific transport properties4& 46.107-125

0, + 2H’ + 2e- 2 H,O,. H,O,

+ 2H’ + 2e- 2 2H,O.

(5) (6)

A linear dependence of the amperometric signal due to reactions (5) and (6) is obtained when the mass transfer of oxygen is the limiting process. Thus the polarization of the platinum cathode should be sufficient (usually from -0.6 to -0.9 V vs. a reference electrode, Ag/AgCl). The enzyme catalysed reaction (4) should be limited by glucose diffusion as well. This is achieved by use of a variety of glucose diffusion semipermeable membranes separatin the enzyme layer from the testing solution31,58,g2- %‘. In the oxygen electrode-based glucose biosensors, the signal output is the difference between the base oxygen level and the level attained as a result of oxygen depletion by the enzymatic reaction. In practice, a combination of identical oxygen detectors with and without an active enzyme layer is used so the signal output is the difference between the glucose dependent current and the oxygen dependent background current31. The oxygen electrode based sensors are advantageous in that they require a cathodic potential, at which only a few endogenous chemical species can interfere’%lo’. Also, the gaspermeable hydrophobic layer over the electrode prevents diffusion of electrochemical interferents to the electrode surface’02-‘07. Despite these advantages, the oxygen electrode signal depends on the ambient oxy-

278

Dependence of the sensor signal on oxygen concentration due to reaction (2) is another source of difficulty. The use of membranes or coatings with hydrophobic properties allow restrictions on glucose flux, hence allowing a higher oxygen flux in excess of that of glucose, avoiding oxygen dependence of the signal. Amperometric biosensors electron transfer

with

mediated

In the second class of glucose sensors, oxygen as an electron acceptor (reaction 2) is substituted by an artificial mediatorgo. Ferrocene (Dicyclopentadienil Iron, Fe”+(Cp)*) and its derivatives are the most commonly used mediators. In these sensors instead of reaction (2) a process involving the mediator takes place: GOD(FADH,) + Fe*+(Cp),.

+ Fe3+(Cp),

-

GOD(FAD)

(8)

The oxidation of the ferrocene (signal generating reaction) could be performed at potentials much lower (usually from +0.15 to +0.25 V) than that of hydrogen peroxide (reaction 6) and does not require catalytic or noble metal electrodes: anode

Fe*+(Cp),

-

Fe3+(Cp)*.

At such low potentials the interference effect from other oxidizable compounds is very small. As

oxygen is not involved in the signal generation process the sensor signal becomes independent of oxygen concentration1’7-13fi. Based on the independence of the signal to oxygen concentration and diminished interference effect, the true glucose concentration in the subcutaneous tissue has been measured1”3-1sfi The main problem’associated with in vim use of such sensors is mediator leakage and its toxicity. Immobilization of the mediator by adsorption on the electrode or entrapment of it in an inert polymer 127-‘3” or albumin matrix133-135 are the solutions proposed for in viva use8”. The possibility of lethal or serious side effects from the leakins mediator or damaged sensors limits the use of this advantageous system in viz~. Amperometric transfer

biosensors

with

direct

electron

It has been shown that oxidation of glucose by enzyme glucose oxidase at an electrode constructed of conducting organic salts (charge-transfer organic complexes with electron conductivity) does not depend on the oxygen concentration”‘.‘“‘. The mechanisms of action of such biosensors are not yet clear. Direct electrooxidation of GOD(FADH,) on the electrode surface instead of oxidation by molecular oxygen (reaction 2) is one of the proposed hypotheses’“‘: ~~q+*nic

GOD (FADH,)

salt

-

GOD(FAD)

+ 2H’

+ 2e-. (10)

In this reaction (10) direct electron transfer between the enzyme and the electrode is assumed. Biosensors based on this principle, the third class of biosensors, have low operating potential (usually from +0.2 to +0.4 V) having all the advantages of the mediated sensors. Moreover, they are a typical example of a solid state biosensor and there is no soluble or partially soluble compound involved in the construction. Sensors based on this principle have been successfull tested in monitoring brain glucose levels in rats~~x-1 J1). However, intensive application of such biosensors ilz 71ivo is limited by the lack of knowledge of the biocompatibility of the new electrode materials involved. DESIGN

APPROACHES

The development of implantable glucose biosensors for feed-back control of the insulin pumps diverge into two paths: sensors for long-term implantation”‘!‘:l, 1~0-144 and needle glucose biosenl~~-~f~,107-10!~.1’L1.1”f;.1~r,.157 sors . The first type of implantable glucose biosensors can be relatively big, being an integrated part of the pump itself, and thus suitable for chronic implantation for a time as long as the lifespan of the pump itself. The miniature needle-size sensors are suitable for subcutaneous monitoring of glucose concentration at a site different from the site of the insulin pump implantation or in combination with a wearable artificial pancreas. The latter concept

allows consecutive use of several sensors within the lifetime of the insulin pump. The problems connected with standardization of the sensors and their in viz)0 calibration are of critical importance in this case145-157. Development of the microdialysis technique and commercialization of microsampling devices allow another, third approach to in viuo monitoring of glucose by coupling the biosensor with a wearable microdialysis device in an extra-corporeal circulation lo~p~‘-~‘“~‘“~. The general advantage of such a method is in the separation of sampling and measuring processes. Needle-type

glucose

biosensor

Short-term studies have demonstrated the feasibility of an implanted glucose sensor controlled insulin infusion system in achieving normoglycaemia. Unfortunately, they have also demonstrated the major limitation of this type of sensor - the rapid deterioration of glucose oxidase. Despite much recent research effort, a long-term implantable glucose sknsor is not yet realized. Miniaturization of the sensor to needle-size is an engineering challenge to obtain a sensor suitable for subcutaneous implantation. The general size of the needle-type glucose from sensors ran es 16 to 28 gauge 7 ,‘-l’. A sensor needles”,““*‘” usually consists of three layers-an inner selectively-permeable membrane (which may 01 may not be present), an enzyme layer contaimng glucose oxidase, and an outer semi-permeable mem~raIle’O”,l’“.l~~“.‘““,l~‘” . The sensor generally consists of a platinum working electrode (polarized as anode) and silveP4, stainless steel”“, Ag/AgCl, or p~~~~~~~ll”,l’l,l”” as counter electrode, used as a reference in two-electrode systems. A three-electrode system with a Ag/AgCl wire reference electrode placed inside the needle body has also been reported”“. 1‘,:I. Schichiri rt nl were the first to report success in miniaturizing a glucose sensor and introduced needle enzyme electrodes which had an outer diameter ok 1 mm”‘. The sensor consisted of a platinum working electrode, a silver reference electrode, and an outer polyurethane membrane. The sensor had a short response time of 16 set and a linear concentration range of 2’7 mM of glucose. The sensor was im lanted in dogs and had a lifetime of’ 4 daysJ5.‘ #z. The sensor was also implanted subcutaneously in humans and the sensor output showed a high correlation with blood glucose level~x~l~*“-l’X. Schichiri rt nl also developed a telemetry system for integration with their glucose sensor for monitoring and control of an insulin delivery system5’. Vadgama ct al. presented a similar glucose sensor with an outer polyurethane membrane that extended the linear range of glucose concentration to 70 mM. The sensor also had a fast response time (60 set) 1”7.‘ox. These authors have demonstrated the effect of varying the concentration of polyurethane in the outer membrane on the linearity of the sensor response”‘R.’ *‘.

279

Glucose monitoring:

E. Wilkins

and P. Atanasov

Pfeiffer et al. constructed a glucose sensor using standard stainless steel needles as cathode and a sensor body 42. The sensor had a response time of 100 set and a lifetime of 6 days116. The sensor was implanted subcutaneously in sheep and it was found that the delay in response between the sensor signal and the intravenous glucose level did not exceed 5 min161. Wilson et al. have recently proposed a new configuration, in which the sensing element was located at the sides of the sensor body in the form of a cavity 156. The size was equivalent to that of a 26 gauge needle (0.4 mm outer diameter). The sensor had a silver/silver chloride reference electrode and a linear range of 15 mM of glucose. They used an inner cellulose acetate membrane which reduced the effect of interferents. This sensor was implanted in the subcutaneous tissue of rats and had a lifetime of 10 days111,112,‘62-‘64. They reported the use of the same construction but with lactate oxidase enzyme for lactate monitoring’65. Karube et al. presented a needle glucose sensor which employs an inner Nafion membrane, and an outer cellulose acetate membrane’51-154. This sensor has been reported to show diminished response to interferents, a response time less than 30 set, and a long lifetime of 25 days. Harrison et al. also have used Nafion as a material for outer membrane ‘in the development of their needle glucose sensor llg-121. A cross-sectional view of this miniature sensor, suitable for intravascular glucose monitoring and subcutaneous implantation is shown on Figure 3 (from ref 120). An inner poly(o-phenylenediamine) membrane was electrically deposited on the working electrode’*l. They report that this combination of inner and outer

membranes produced high selectivity and diminished interference effects. The sensor was implanted subcutaneously in dogs and had a lifetime of 14 days. Combination microdialysis

of the glucose sampling

biosensor

Difficulties associated with in viva glucose monitoring in a site close to the pump implantation site (abdominal cavity) or in subcutaneous tissue could be avoided by extracting a subcutaneous dialysate and passing it through a small cell with the sensor in it. In this case, however, the microdialysis system and the sensor should be in a wearable unit. Experimental advantages of this technique are more controllable conditions and the ability to analyse the same sample with different methods62*68. There is a question about biocompatability of such techniques and possibility of infections during long-term microdialysis. Some of the leading groups in implantable glucose biosensors development have turned to this concept of glucose monitoring66,‘58. Figure 4 (from ref 158) illustrates an advanced concept of combination of a needle-type glucose biosensor with a hollow-fibre probe and a flow-cell for microdialysis sampling, developed in Shichiri’s group. There is a recent debate on how superior the microdialysis approach is in comparison to the in viva monitoring with needle sensors166. Following the glycaemia profile with time by using the comA

Polyimide

coated

0=0.15mm

fused-silica

h -

Hollow-fiber

Stainless steel (SUS-316) 0=CL45mm 5-c Outlet caonula

um

Inlet cannula

1 0=4tnm Stainless steel (SUS-316) 0=O.65nun Epoxy

adhesive

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Figure 3 Cross-sectional view of an implantable enzyme electrode for glucose analysis in whole blood (with permission from: Turner RFB, Harrison DJ, Rajotte RV, Baltes HP. A biocompatible sensor for continuous in vivo glucose monitoring in whole blood. Sensors & Aclualm B 1990; 1: 5614, ref. 120)

280

5mm ?1

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FIgure4 Schematic diagram of hollow-fibre probe and flow-cell with a needle glucose sensor for extracorporeal glucose measurement in the miniaturized subcutaneous tissue glucose monitoring system based on microdialysis sampling method. A: Hollow-fibre probe. In an inner cannula, the perfusate is transported to the top of the probe, diffused through the dialysis membrane, and the dialysate then comes out through an outlet cannula. B: Sensor flow-cell (with permission from: Hashiguchi Y, Sakaida M, Nishida K, Uemura T, Kajiwara K, Shihiri M. Development of a miniaturized glucose monitoring system by combining a needle-type glucose sensor with microdialysis sampling method. Diabetes Care 1994; 5: 387-96, ref. 158)

~Xucosu monitoring:

bination of needle sensors and microdialysis sampling is reported to be very close and accurate of the (see Figure 5)15’. One main advantage microdialysis technique is the flexibility of replacing the sensor in case of failure. However the disadvantage will be the use of additional equipment. Employing this technique, however, with additional sophisticated equipment makes it impossible to integrate the measuring device (the sensor) and acting device (the pump) in one unit. The ultimate goal to integrate the electromechanical analogue of the pancreas into one device is achievable through development of chronically implantable sensors. Glucose

biosensors

for chronical

implantation

Chronically implantable glucose biosensors are usually designed as an integrated part of the implantable insulin delivery system (artificial pancreas) 42,167. Another application of such sensors is in an autonomously powered device, often with a radio transmitter - the prototype of a diabetic alarm system40*41. Figure 6 (from ref 41) shows a schematic of a sensor/transmitter system which is reported recently to be implanted and operating in a dog model for three months. Devices based on oxygen and on hydrogen per-

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Fiie 5 Continuous monitoring of subcutaneous tissue glucose concentrations measured by the extracorporeal glucose monitoring system with microdialysis sampling methods after 75 g oral glucose load (A) and after intravenous insulin injection (0.1 U/kg) in five healthy subjects. Subcutaneous tissue glucose concentrations (-) were compared with blood glucose concentration (-) measured by the glucose monitoring system (STG, Nikkiso). Results are analysed at 5-min intervals of the continuous monitoring records and expressed as means fSE (with permission from: Hashiguchi Y, Sakaida M, Nishida K, Uemura T, Kajiwara K, Shihiri M. Develop ment of a miniaturized glucose monitoring system by combining a needle-type glucose sensor with microdialysis sampling method. Diab&s Cm 1994; 17: 387-96, ref. 158)

and I’. Atannsou

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oxide measuring principles have been described. For in viva evaluation of such sensors the usual choice is a large animal such as dog or sheep. The most preferable site of implantation is chosen to be the abdominal cavity and the glucose level is measured in the interstitial fluid” . Calibration of the implantable sensor is proposed to be realized using two glucose levels (twopoint calibration) which is satisfactory in the case of linear dependence of the sensor response’8”~18”. In the development of the integrated implantable sensor, the critical issue is that of long-term operational stability of the sensor. Since the glucose. biosensor will be a part of the implantable insulin delivery system, its lifetime must be at least as long as that of the other parts of the system. The lifetime of the biosensor is limited mainly by the processes of enzyme deactivation105,‘68. The immobilization of the enzyme in gels, on membranes, or on inert dispersed carriers (usually carbon materials) can significantly increase its stability105. However, the lifetime of the biosensors with immobilized enzymes is still limited. Increasing the sensor lifetime has been achieved by using an optimized immobilized enzyme and protective layers 168-184. Successful chronical implantation of such sensors in dogs for a period of three months has been recently reported4+“‘. Rechargeable

0

E. Wkins

glucose

bioseusors

A new approach, developed in the University of New Mexico, makes it possible to extend the sensor lifetime by in situ sensor refilling - replacing immobilized enzyme with fresh spent enZyme14”.144. The enzyme glucose oxidase is immobilized on dispersed carbon powder, which is then held in a liquid suspension. The construction of the biosensor is such that the spent immobilized enzyme can be removed from the sensor body and fresh enzyme suspension injected via a septum, without sensor disassembly. This concept facilitates recharging of the implanted sensor without surgical removal from the patient.

281

Glucose monitoring:

6 Wilkins

and P. Atanmov

Rechargeable sensors based on both hydrogn peroxide’ 5 and oxygen measuring principle ” have been reported. Figure 7 shows a crosssectional schematic of the rechargeable glucose biosensors based on hydrogen peroxide (a) and oxygen measuring principle (b) . The glucose biosensors consist of two parts: an amperometric electrode system and an enzyme micro-bioreactor. A three electrode amperometric scheme is used in both types of the biosensors: a platinum wire (diameter 0.25 mm, length 4 mm) as a working electrode (1)) a silver/silver chloride reference electrode (2) and another platinum wire as a counter electrode (3). The working electrode is polarized +0.6 V for hydrogen peroxide oxidation, or -0.6 V for oxygen reduction vs. the reference electrode. In the oxygen measurement based sensor (Figure 76) these three electrodes are symmetrically assembled and housed in a separate glass tube (4) and cemented by epoxy resin. At the faceend of the housing, a cavity is formed and filled with gelled electrolyte: agarose or poly-hydroxyethylmethacrylate (HEMA) matrix (5) soaked in phosphate buffer solution containing potassium chloride. An oxygen-permeable hydrophobic membrane (6) is used to cover the three electrode system separating it from the enzyme micro-bioreactor. Mechanically attached Teflon membranes or membranes made by dip-coating of the oxygen electrode with silastic latexes are used. In hydrogen peroxide based biosensor (Figure 7~) the three electrode amperometric system is directly inserted in the sensor housing (7)) a plastic tube, face-side closed by a glucose diffusion membrane (8). This enzyme-modified dispersed carbon is used to form a liquid suspension for refilling of the enzyme micro-bioreactor (9), recharging of the sensor. Two capillary plastic tubes (diameter ca 1 mm) - inlet recharge tube (10) and exhaust disB

Figure 7 Schematic of rechargeable glucose biosensors based on hydrogen peroxide (a) and oxygen selective electrodes: Pt-working electrode (1). Ag/AgCI-reference electrode (Z), Pt-auxiliary electrode (3), oxygen-selective electrode housing (4), gel electrolyte in the oxygen-selective electrode (5), oxygen permeable membrane (6), biosensor housing (7), glucose diffusion membrane (8)) micro bioreactor with enzyme suspension (9)) inlet recharge tube (10) and exhaust discharge tube (11)

282

charge tube (11) - are used for replacing spent enzyme from the micro-bioreactor without sensor disassembly. Refilling of the sensor is achieved using two septums via these tubes: one for injecting a fresh enzyme suspension, another for exhausting the spent enzymeg6. Biosensor dimensions are roughly 3 mm in diameter for the hydrogen peroxide based sensor, and 6 mm in diameter when an oxygen-selective electrode is used. Both bigger and smaller sensors have been and are being constructed. This biosensor was integrated with a miniature potentiostat with a signal transmitter into a small independently functioning device suitable for canine or other large animal implantation’86. The amperometric output of the biosensor is linearly proportional to glucose concentration in the entire range of physiological and pathophysiological interest. The upper limit of the linear range of the biosensor varied from 220 to 1000 mg/dL, depending on the type of the sensor and particularly of the t e of glucose diffusion membrane used14%144,1~f Response time of the sensor - time to reach 95% of the steady-state signal value - is from 1 to 10 minutes depending of the glucose concentration step change and the diffusion membrane used. The biosensors are stable while operated continuously in a model phosphate buffer solution (pH 7.4) for a period of 7 months at 25”C, or at least 3 months at 37°C before refilling with fresh enzyme suspension g6,185. The dependence of the steady state sensor signal vs. glucose concentration demonstrated significant reproducibility in their main characteristics: linear range and sensitivity. Standard deviation from linearity for any of the measurements does not exceed +0.002449. Recharging of the sensor with fresh enzyme (refilling the sensor body with new glucose oxidase immobilized on the carbon powder after replacement of the old suspension) can successfully increase the sensor lifetimegGgg,‘8”. The sensitivity of a biosensor (based on the oxygen electrodegg) vs. time of continuous operation through several recharge cycles is presented in Figure 8. It should be noted that during the first days of operation sensor sensitivity decreases from its initial value, probably due to the loss of activity of the glucose oxidase in the solution (not chemically immobilized on the carbon carrier). After this initial period the sensitivity of the biosensors is essentially constant for a period of about 14 days. At the end of this period the sensitivity begins to decrease. When the biosensor sensitivity drops to more than 10% of the previous constant value, the sensor is refilled with fresh immobilized enzyme suspension, using septums (without sensor disassembly). After refilling and initial stabilization the biosensor continued to demonstrate the same constant value of sensitivity as during the first recharge cycle. The biosensor has been refilled seven times with the fluid enzyme-modified carbon suspension, running over eight recharge cycles of about two weeks each, during the four month continuous test (Figure 8). By using the combination of two enzymes (in

Months I

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20

40

100

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120

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Figure 8 Dependence of the biosensor sensitivity on the time of continuous operation in 100 mg/dl (5.6 mM) glucose solution dnring seven recharge cycles. Dates of recharge are shown by arrows. (with permission from: Atanasov P. Wilkins E. Development of biosensor for glucose monitoring Riotcchnol Bioq 1994; 43: 262-6, ref. 99)

oxygen electrode based biosensors) - glucose oxidase and catalase - the recharge cycle (time between biosensor refilling) can be extended up to one month due to the elimination of the hydrogen peroxide as an enzyme deactivatoP8. Glucose concentration in sera samples has been monitored by the biosensor for 16 hours (in vitro) with no noticeable change in the sensor response. After this test a calibration curve of this sensor in phosphate buffer glucose solution is obtained and the results (sensor sensitivity and linearity of the response) coincide with the calibration curve obtained before sera measurements. The glucose biosensor has been implanted and tested in viva in a sheep i8’ . Glucose concentration in the animal interstitial fluids has been monitored by the biosensor and juxtaposed with the blood glucose levels estimated by an external conventional method in the periodically obtained samples (Figure 9). The test includes a step change in glucose concentration by injection of a. glucose dose into the animal. The data shows good correlation between the profiles of the glucose concentration evolution in body fluids obtained by these two different methods. The capability to refill by external means an

0

IO

20

30

sensor without surgical implanted glucose removal would be a major advance in the development of a practical glucose sensor for diabetes control. Our use of enzyme in the form of a liquid suspension of carbon powder with immobilized glucose oxidase on the powder allows enzyme replacement without surgery by injection via a subcutaneous septum using a procedure similar to that developed for refilling of implantable insulin pumps”“. The potential of this approach opens the possibility of constructing sensors suitable for implantation with lifetimes greater than several months in the body.

40

Time. [minutes] Figure 9 Profile of the in viva glucose concentration in the interstitial fluid (cnrve I) measured by the biosensor and blood glucose concentration measured by glucose analyser (curve 2). (with permission from: Wilkins E, Atanasov P, Muggenburg BA. Integrated implantable device for long-term glucose monitoring. Riowrnrors tj Hiorlectronlcs 1995: lo: 485-94, ref. 186)

CONCLUSION Development of glucose monitoring techniques and approaches during the last decade demonstrates the predominance of the electrochemical measuring principles. Biosensors are still the main focus of the research interest of most teams due to their high selectivity for glucose determination. Different design approaches have been proposed and are under current development in order to overcome the problems associated with the sensor characteristic reproducibility and lifetime termination. The approach of small replaceable (disposable) sensors, usually designed as a needletype electrode, is most popular among investigators. Coupling of such a sensor with an insulin pump, usually wearable on the patient’s body (arm), is disadvantaged because of the necessity of frequent transcutaneous invasions when the sensor is replaced. The needle size glucose sensors appear to be most advantageous when used in combination with microdialysis technique for subcutaneous liquid sampling. In this case, the problems of in situ biocompatibility of the sensor interface to the living tissue is avoided. Sensor replacement and other manipulations with it are not associated with surgical or transcutaneous interventions. This kind of design, however, only allows the development of a wearable artificial pancreas. It appears that development of a chronically implantable glucose sensor is the possible solution to achieve an implantable insulin delivery system with an internal feedback control. The immobilization techniques allowed to build sensors with in uiuo lifetime up to three months - much less than the lifetime of the implantable insulin pump itself. The approach for replacing of the spent enzyme and thus recharging the sensor in situ while implanted is promising as a way to increase the lifespan of the system. Ultimately, it could lead to rare transcutaneous interventions for refilling of the sensor. This intervention could be combined with the refilling of the reservoir of the insulin pump itself with fresh insulin. This kind of system seems to be patient friendly enough to enhance the life quality of diabetics. The most promising non-electrochemical approach for glucose monitoring is the remote transcutaneous NIR measurement. Being in an early stage of development, now, this method could provide a desired non-invasive sensing of

283

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monitoring: E. Witkins and P. Atanasov

glucose, especially if a remote feedback from sensor to the implantable pump is realized.

the

ACKNOWLEDGEMENTS The research on rechargeable glucose biosensors at the University of New Mexico is currently supported by a grant from the National Science Foundation and the Whitaker Foundation. The authors would like to express their gratitude to Prof. Michael Wilkins for his help in the preparation of the manuscript.

ing fourier transform through oral mucosa. Med 20. Van Heuvelen A. Blood 4704029 (1987). 21. Guilbault G. Non-invasive ArtiJicial Organs 1988; 13: 22. Guilbault G. Biosensors 23. 24. 25.

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