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Fundamentals of Glucose Sensors U. Fischer Institute of Diabetes ’Cerhardt Katsch’, Karlsburg, Germany

KEY WORDS Glucose Biosensor Glucose sensors

The Need for lntracorporal Glucose Sensing During the last decade, various new tools in metabolic management such as structured education programmes, multiple injection regimens, insulin pumps, and specially designed pocket computers have contributed to the progress towards near-normoglycaemic blood glucose control in Type 1 diabetic patients. This has only been possible since patients have been able to monitor their own blood glucose levels under ambulatory conditions by means of reagent strips,’ either at regular intervals or sporadically. By these means the feedback loop of metabolic management is closed a little further. However, several problems in this field are still far from being solved. Many patients consider self blood glucose monitoring laborious and inconvenient, and wish to be free of it. The closer metabolic control is to permanent normoglycaemia the greater i s the danger of hypoglycaemia2, which in turn prevents perfect control. To protect patients from resultant adverse effects would require hypoglycaemia warning device^.^ There are a number of patients who cannot be adequately controlled using any of the open loop modes of insulin admini~tration,~ and even those who can be adequately controlled by no means remain within the near-normoglycaemic range over the longer term.3 Thus early data from the DCCT study have shown that during intensive insulin therapy the glycosylated haemoglobin level was significantly improved but not n ~ r m a l i z e d . ~ the future, for less adequately controlled patients we might have to consider transplantation of insulinproducing tissue either directly6 or processed to give a bioartificial ~ y s t e mor , ~the insertion of the bioengineered proinsulin gene into somatic cells.8 At the moment, however, for maintaining permanent euglycaemia these patients would need a miniaturized device for external insulin delivery in a closed-loop controlled system, the

Correspondence to: Professor Dr U. Fischer, Division of Pathophysiology, Institute of Diabetes ‘Gerhardt Katsch‘, Creifswalder Str. 1 la, Karlsburg 0-2201, Germany

Electrode

Implantation

Diabetes care

artificial B - ~ e l l . lo ~ , As an alternative in the meantime frequently sampled or continuously monitored intracorporal blood glucose values are required to support more safely the decisions on insulin doses.”, l 2

Biosensors All these requirements necessitate the development of a biosensor able to monitor intracorporal glucose concentration. Biosensors (Figure 1) are small-probe analytical devices comprising a biological sensing element, such as enzymes or antibodies, in connection with a physicochemical tran~ducer.’~-”They allow continuous, rapid, and specific measurement of concentrations of chemical species in unprocessed biological material. Iri principle, biosensors may be applied to extra- or para-corporal samples or to intracorporal measuring compartments.’a In the latter case, the functional biostability of sensor materials in contact with tissue, blood or other body fluids, is of particular importance. Usually these are synthetic membranes, such as various polyurethanes, in order to protect the inner compartment from adverse effects of the biological environment, and also to guarantee the differential but reproducible mass transfer both of the analyte and its co-substrateb) on the one side, and of the reaction product(s) to be removed on the other. This property, namely the permeability of sensor membranes, affects the specificity of the sensor and is crucial for its linear range in calibration and response kinetics. Permeability depends on the hydrophobicity of biomaterials used which may be higher or lower, down to nearly total h y d r ~ p h i l i c i t y . ’In ~ the former case, for example, using polyethylene, gases such as oxygen and carbon dioxide are easily transported but low molecularsize solutes such as glucose are not. These, however, are transported at high rates if hydrophilic membranes are used, including various derivatives of cellulose. Thus it i s important that during intracorporal residence neither changes in membrane swelling nor deposits from the biological environment would alter these permeability features. It must also be taken into consideration that the functional biostability of sensors largely depends on their biocompatibility,18, 2 1 that i s on their ability to perform within the body without adverse reactions of the host22 due to toxicity, local foreign body reactions, complement

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DIABETIC MEDICINE, 1991; 8: 309-321

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ANALYTE

biological sensing element (enzyme, antibody

...I

BIOMAT ERI AL

physico-chemical transducer (electrochemical, calorimetric

...I

Figure 1 . The main components of a biosensor

possible. But there would still be the need for a membrane as is the case both in the pancreatic B-cell and in other artificial systems. This is necessary both to separate (to protect), and to connect (to provide substrate) the metabolic and sensing compartments. Thus, the main drawback of intracorporal glucose sensors, namely their bioincompatibility,22 would not be circumvented by external replacement. Therefore, achievements to date in glucose sensor research for diabetes management are exclusively based or1 established physical and (bio)chemical principles (Table 1). With the exception of some electrochemical approaches, all analytical principles investigated for medical purposes are still the subject of basic research. Working Principles This applies particularly to optical methods which are With the exception of optical rotation as utilized in all fast-response analytical devices, but are either not laboratory-style polarimetry, and of certain infra-red specific (infra-red spectroscopy) or exhibit other unsolved absorption maxima as utilized in infra-red spectroscopy, problems such as instability even in vitro, and the 2a30 there is no specific property of the glucose molecule impossibility of cheap mass production (fluorescence sensors). In the latter case (Figure 2) specificity has been that could be made use of as a physical or chemical principle in quantitative analysis. Thus a biological guaranteed either by means of monosaccharide-specific molecular receptors or by enzyme catal ysed glucose sensing element is required to generate a measurable signal.”, 3 ’ oxidation, which provides glucose-related alterations of At a first glance it might look promising simply to induced fluorescence via the stoichiometrically varying imitate the most efficient glucose sensing system, as oxygen concentration in the sensing compartment. realized naturally in pancreatic B-cells. This i s obviously Micro-calorimetric methods are also considered potenthe pancreatic isoform of the enzyme g l u ~ o k i n a s e ~ ~ tially intracorporally applicable for medical purposes. These approaches are based on the energy balance of which very rapidly, on the basis of glucose-controlled enzyme-mediated oxidation of glucose (and hydrogen glucose utilization within the B-cells, controls their ATP peroxide). Thus, either the electrical conductivity of the resources and related phosphorylation processes. This i s followed by an increase in intracellular free calcium sensing element is altered in a temperature-dependent accompanied by alterations in membrane transport of fashion (thermistor), or an energy-dependent voltage is other cations and thereby membrane depolarization, generated within a micro-chip, the so-called thermopile. finally leading to p-granule exocytosis. In principle, Both methods are very fast and extraordinarily sensitive imitating this very complex chain of events appears but also easily affected by a number of disturbing

activation, thrombogenicity causing haemo-incompatibility in the case of intravascular application^,^^ or infection which may be harmless but results from the impossibiIity of sterile hand Iing. 24 There are well-established laboratory, bedside” or of glucose sensors even p a t i e n t - ~ o r n26 ~ ~applications , on the basis of an electrochemical principle. lntracorporal glucose sensors, however, are still in the experimental stage.*’ The following parts of this text mainly focus on the relevance for medical purposes of sensor signals from intracorporal measuri ng sites.

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Table 1. Investigated working principles of glucose sensors for potential paracorporal or intracorporal application Technique

Basic principle

Non-electrochemical approaches Optical Infrared spectroscopy

Microcalorimetry

Specific features Advantages Disadvantages

No chemical reaction, fast response

Specificity and device miniaturization not guaranteed; expensive Instability, device miniaturization hard

State of the art

Selected references

Established physical principle

28-30

In vitro examination and preliminary in vivo data

33-34

Fluorescence sensor: affinity systems oxygen-dependent fluorescence quenching

Highly specific, fast response

Enzyme thermistor

Precision

Problems in insulation/miniaturization, need for reference thermosensor

Applications in biotechnology

35

Enzyme thermopile

No external

Miniaturization, drift, disturbance variables

Prelimi nary paracorporal approach Preliminary animal trials Physical chemistry of appropriate membranes characterized

36

power supply

37-39

N o power supply required, no transduction element

No data on glucose concentration provided, Iimited capacity requires frequent reimplantations

Potentiometry, anodic glucose oxidation (metal catalysis)

Reactionlimited function 02independent

Not specific, need for sensor regeneration

Preliminary in vivo experiments

14, 42-45

Fuel cell

Amperometry after catalytic glucose oxidation

No current supply

Diffusion-dependent, no miniaturization

Established in vitro

Coated-wire electrode

Dipped glucose-selective membranes on platinum electrodes

Sensitivity, no reproducible manufacturing, large reference electrode

In vitro verification

42, 4 7-5 3 50, 51

Reference electrodes required (in p H and PO2 sensors), various disturbance variables (e.g. Po ), diffusionlimited function (particular dependence on biocompatibility)

Application in biotec hnology

15, 42

Established in vitro, anecdotal animal experiments Serial application in laboratory analysers and in animals, anecdotal human trials

24, 42, 52-59

Instability in vivo

Serial application in wearable analysers, anecdotal in vivo experiments

31, 63, 74-80

Shape (planar chips), drift in sensitivity (strong influence of pH and protein).

Preliminary in vitro studies

‘Chemical feedback’ insulin delivery (pharmaceutical drug delivery systems)

Electrochemical devices Catalytic electrode

Enzyme electrode

Affinity systems

GIucose-dependent variation of insulin permeability of biomaterial membrane

independent

p H alteration caused by generated gluconic acid

Precision

Glucose oxidaseloxygen (amperometry)

Stoichiometry

C Iucose oxidaselH,O, (amperometry)

Stoichiometry

Glucose oxidase/electron mediator (amperometry)

Enfet (potentiometry on enzyme-spin-coated ionselective field effect transistor)

GLUCOSE SENSORS

0 2 -

0 2 -

independent

Reproducible mass production

40, 41

10, 1 1 , 50, 60-73

81

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glucose - permeable mem brane

u

data processing @ glucose molecule

Concanavalin A Dext ran - FITC-complex Figure 2. Working principle of affinity fluorescence glucose sensor: in the sensing compartment glycated fluorescein isothiocyanate (FITC) i s prevented from fluorescence because of reversible coupling onto glucoreceptors (e.g. Concanavalin A). Glucose, via the sensor membrane, diffuses in a reversible, concentration-dependent manner from body fluids into the sensing compartment where it competes with glycated FlTC at the Concanavalin A binding sites. Freed glycated FlTC produces fluorescence if stimulated at the appropriate UV wavelength3’

variables, and therefore need expensive insulation which glucose, or on the basis of control of permeability of reservoir membranes for small proteins such as insulin reduces the chance of reasonable miniaturization. by the pH-alteration which is caused by gluconic acid Since it is the ultimate goal of a glucose sensor to close the loop between the intracorporal glucose produced during enzyme-mediated glucose oxidation. concentration and the mechanism of insulin d e l i ~ e r y , ~ As yet, there are particular drawbacks in these systems because of adverse reactions to implants, including the fascinating principle of a direct chemical feedback immune responses and the need for frequent reimplanhas been established as a research tool (Table 1, Figure 3). In this, insulin i s released from implanted pelleted drug tation. Until now, among the various working principles of delivery systems in a glucose concentration-dependent glucose biosensors only the electrochemical approach manner, working either on glycated insulin which is has gained practical use, and even this is still experimental displaced from specific binding sites by intracorporal

biodegradable polymer matrix /

ronosacchar tde receptor (Con. A 1

m

glycated insulin displaced from receptor ext mcellular glucose gl ycat ed insulin bound to receptor Figure 3. Working principle of an implanted drug delivery system exhibiting glucosecontrolled insulin release by means of a chemical feedback: glycated but biologically active insulin i s reversibly immobilized onto an implantable, biodegradable matrix via gluco-receptors (e.g. Concanavalin A). Glucose diffuses from the surrounding body fluids in a concentration-dependent manner onto the binding sites of glycated insulin where the latter i s irreversibly displaced and released to exert its biological effects. Note: the number of binding sites is limited, e.g. in diabetic rats the capacity of one pellet may provide the required insulin doses over approximately 3 weeks”, 38

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(Table l).50 With the exception of the coated-wire electrode, they all depend on glucose oxidation (Figure 4). The coated-wire electrode, however, achieves specificity of glucose-derived ions through impregnation during the manufacturing processes of sensor membranes with glucose salts which are later removed. As yet, they are apparently not suitable for practical applications. The main drawback of most other electrodes comes from the fact that stable, diffusion-limited membranes are required to guarantee appropriate concentration ratios of the two substrates, glucose and oxygen, within the reaction ~ o m p a r t m e n t . 5~2~, 6o , This dependence results in a crucial influence of bio-incompatibility reactions which, for example, can deposit layers on the sensor membrane and thereby alter its permeability properties and cause unstable function of implanted electrodes. In general, however, these electrodes are superior to other principles of glucose sensing because of (1) their stoichiometric function, (2) the general technological progress in polarographic-amperometric sensors, and (3) the possibility of easy miniaturization and mass production. In practice, some prototypes are well established in extracorporal analytical applications such as the amperometric oxygen-electrode in the Beckman analyzer, and the amperometric glucose oxidase/hydrogen peroxide electrode in the Yellow Springs analyzer and the Biostator bedside-type artificial B-cell.” 61 Other examples include the wearable miniaturized Elco Direct 30/30 device, and the disposable ‘glucose-pen‘ electrodefz5,6 2 using electron-mediated glucose oxidase/ferrocene strips as in the Exactech system. As to in vivo application, each type of enzyme electrode is subject to specific drawbacks (Table 1) which, until now, could best be compensated for in the glucose oxidase/hydrogen peroxide type electrode. Unlike the oxygen measuring sensor this does not require a third electrode as an enzyme-free reference, and unlike the electron-mediated sensor it is relatively stable.

Furthermore, unlike the Enfet (Table 1) it can be manufactured in the form of a miniaturized needle. Since we have been able to validate various types of glucose oxidaseihydrogen peroxide sensors, the rest of this review will mainly consider the intracorporal application of this type of electrode, which has been extensively characterized under in vitro and in vivo

condition^.^^-^^

The Glucose Oxidase/Hydrogen Peroxide Electrode: Design and Functional Characteristics This type of sensor can easily be manufactured as a single needle (for references see Table 1) either with or without the cathode as a separate entity that may lie within or outside the body. Basically, the sensor compartment is protected from the biological environment by means of an outer covering membrane (Figure 5). This is hydrophilic and thus permeable to the analyte (glucose) and the co-substrates (oxygen, water) and the reaction products (gluconic acid, hydrogen peroxide). Appropriate concentrations of the two reactants in the enzyme compartment are provided by an inner hydrophobic membrane which is easily permeable to oxygen but can only be passed by glucose (and water) via a small hole that is mechanically placed in front of the anode.64 In the inner enzyme compartment, glucose is oxidized with the help of glucose oxidase, and the hydrogen peroxide produced during this reaction i s measured by means of a polarographic electrode. The resulting current is linearly dependent upon the concentration of hydrogen peroxide which in turn is stoichiometrically related to the glucose concentration within the sensor compartment. The latter depends finally upon the glucose concentration in the surrounding biological medium (blood, tissue). The whole process functions linearly as long as sufficient

(instead of 02).whose regeneration

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Figure 4. Established principles of employing various aspects of enzyme- or non-enzyme mediated oxidation of the glucose molecule for its electrochemical detection GLUCOSE SENSORS

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* hydrophilic diffusion membrane perforated hydrophobic membmne react ion compartment enzyme react ion electrochemical reaction

polarograp hic electrode cathode IAg /AgCl) Figure 5. The amperometric glucose oxidaseihydrogen peroxide electrode; glc, glucose; GOD, glucose oxidase

oxygen supply is guaranteed in relation to the amount of glucose entering the enzymatic reaction.42 The electrodes used in our study were cylindrical Clark-type systems. Details of their manufacture and application have been previously described.64, 67, 82 Briefly, the membrane i s a sandwich of regenerated cellulose (PT 150 Cuprophane), polyethylene with a mechanically produced central perforation of approximately 10 p m diameter, and glucose oxidase immobilized onto sepharose 8B. The diameter of the anode was 0.3 mm and the overall outer diameter was approximately 2 mm. These sensors were used at a polarization voltage of 700 mV, and the current outputs were either amplified and recorded by means of a wearable monitor designed for this system, or transmitted by means of a commercially available telemetry system. The amount of enzyme layered onto one sensor was approximately 10 IJ.I suspension representing an enzyme activity of about I iu. Because of the laboratory-style manufacturing process, this number is a rough estimate only, but it is interesting that so far the stability of enzyme activity has never been a crucial factor in sensor function. We were, for instance, able to maintain a sensor preparation over 4 months working in a bedside-type analyser alternating between room temperature and thermostated at 37 “C. Under an oxygen partial pressure above 5 kPa, linear calibration of this type of electrode preparation could be established to > 15 mmol I-l, residual current at zero glucose concentration was 1.09 ? 0.02 (+SD) nA, and sensitivity was 1.05 2 0.17 nA mmol-’l ( n = 10 sequentially used electrodes). In vitro, there was no interfering reaction of sensor current with physiological concentrations of urea, amino acids, sodium chloride, and potassium ~ h l o r i d e . ~ ’ However, we and others found current-generating effects of dialysed albumin, which are still unexplained but might be due to an influence of the viscosity of the solution at the sensor surface. Furthermore we have seen an interfering effect of therapeutically relevant oxidants 314

such as ascorbic acid and paracetamol, as has been repeatedly reported by other authors.l5. 1 7 , 42, 69, 83 Th is must be considered in any clinical application of such sensors. However, there are also approaches to eliminating these influences, such as by means of socalled lipid-modified electrode^.^^

Where to Sense Glucose From the patient’s and physician’s viewpoint it would be ideal to obtain figures on the intracorporal glucose concentration through a non-invasive route. Some attempts at this have been made30r3’ but due to difficulties in construction of devices and reproducibility of signal kinetics they have failed so far to attain practical relevance. Using invasive techniques, such as the insertion of sensors into tissue or natural cavities, the intracorporal glucose concentration can be analysed either in blood, in transcellular fluid, or in interstitial fluid of tissues (Table 2). The paracorporal monitoring of glucose in diluted blood over short periods of time by means of doublelumen catheters has been established for nearly 20 years.61 This is, however, not useful in the long-term metabolic management of diabetic patients. From the clinical viewpoint and from analogies to cardiac pacemakers, placement of glucose sensors intravenously appears feasiblelo5 and, because of its well-established kinetics, most useful in metabolic control on the basis of proved algorithms. In glucose oxidase electrodes, this would be particularly advantageous, since the oxygen tension in any blood vessel is higher than that in tissue. Furthermore, immunological host reactions may be expected to be less pronounced than those in connective tissue. Thrombogenicity, however, appears to be the main reason for the lack of any substantial progress in this field over the last few years, despite some anecdotal reports (Table 2 ) . U FISCHER

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Table 2. Possible sites for sensing the intracorporal glucose concentration Approach

Advantage

Disadvantage

State of the art

Selected references

Non-invasive Optical (conjunctival: liquid media of the eye) 'Transbuccal' (mouth mucosa)

Sensor-related intracorporal complications excluded; easy removal and recalibration

Casual study, patents

29, 30

Anecdotal animal ex per iments

05, 85-88

Not reproducible and not predictable delay and attenuation of estimated glucose levels in relation to actual levels

Transcellular fluids (Cerebrospinal, urine, tears, aqueous and vitreous humour of the eye, sweat, saliva, peritoneal)

Potentially reduced local host reactions

Microdialysed/microfiltrated subcutaneous tissue (paracorporal devices) Interstitial fluid (mainly subcutaneous)

Easy access, established kinetics

Skin penetration, no long-term applicatian

Animal studies, casual human trials

8 1 , 89-93

Easy access, established kinetics, possibility of total implantation

Host reactions (not fully understood)

Serial experimental applications in man and animals

82

Easy access, ample c Ii n ica I ex per ience (instrumentation,data hand I ing)

Com plicated handI ing, no long-term application

Established clinical applications

1 1 , 26, 61

Thrombogenicity

Casual report

57, 94-96, 105

Circulating blood Paracorporal (bedside or portable mostly via doublelumen catheter) lntracorporal (intravascular sensor)

Several attempts have been made to monitor glucose in selected compartments of transcellular fluid, such as peritoneal fluid, saliva, tears, sweat, and cerebrospinal f l ~ i d . ~There ~ - ~is,~ however, in general a considerable but non-reproducible delay and attenuation of glucose concentration changes compared with blood. Best characterized and least different from blood are probably the kinetics of transcellular fluid glucose concentration in the peritoneal 86, 97 However, more data must be collected before conclusions can be drawn with respect to clinical application. Therefore, most authors have focused their attention on the measurement of glucose in non-glucose-producing tissues. Because of easy access, the subcutaneous tissue i s widely preferred as a site of glucose sensor insertion.65, 99 The data discussed below were obtained in the subcutaneous tissue of the neck, the back, and the abdominal wall of normal and chronically diabetic dogs, either with or without general anaesthesia. For calibration and therapeutic applications the subcutaneous glucose data of an implanted sensor's outputs must be validated against the 'true' glucose concentration in subcutaneous interstitial fluid. This is important because apparent subcutaneous glucose concentrations of between 20 and 85 % of blood glucose have been reported GLUCOSE SENSORS

using in vitro calibrated enzyme electrodes,'0* 64, 74 a finding in contrast to acknowledged models of microcirculation and mass exchange across the capillary wall. Employing sampling procedures such as the wick techniquelS5,98, 99 m i c r ~ d i a l y s i sor~ ~microfiItration,8' it has been demonstrated by means of various independent analytical methods that the 'true' glucose concentration of subcutaneous tissue is essentially identical to that in blood. It should be noted, however, that this applies only to steady state conditions and to tissue which is not inflamed or otherwise damaged. Furthermore after glucose or insulin loads we have seen a distinc:t influence through the physiological delay between alterations in circulating glucose concentrations and those in the tissue compartment, probably due to the glucose kinetics of the microcirculation, microconvection, and diffusion at the tissue level (Figure 6). This delay probably has more influence on the dynamic output pattern of a subcutaneously inserted glucose electrode than the in vitro properties of the sensor per se. From these and other data it was concluded that an appropriate method for in situ calibration of an implanted sensor is mandatory in the event of clinical application of such devices. Various such protocols have been designed, such as the so-called two-point ~ a l i b r a t i o nor ~ ~continuous non-

31 5

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Time Figure 6. In vivo response characteristics of peripheral venous blood glucose (paracorporal monitoring, thin continuous line), of current output of an amperometric glucose sensor inserted subcutaneously in the neck (thick line), and of plasma glucose concentration taken at intervals ( 0 - 0 ) in a normal dog before, during and after a square-wave intravenous infusion of glucose (GIT). Sensor insertion and beginning of run-in at 13.00 h. Note the lack of real dead time in all three curves, and no appreciable difference in the kinetics between the different measuring sites (all time constants approximately 60 min in non-linear regression analysis). In vitro Tq5of the sensor was 6.5 min both before and after in situ residence and at increasing and decreasing steps in glucose concentration. with permission) (Taken from Fischer et

linear regression analysis of blood glucose/sensor output data.68 Their practicability, however, remains to be proved.

Dependence of Sensor Function on the Availability of Oxygen Basically, the described glucose electrodes were designed to measure the reaction equilibrium or the end-point,

respectively, of the glucose oxidase reaction. Its dependence on the prevailing oxygen tension was even observed in reagent strip analytical systems.'00, In contrast to reaction-limited electrodes42 and to electronmediated sensors,b3 the diffusion-dependent sensor principle described here requires the permanent availability of oxygen within the reaction compartment (Figure 5) in an appropriate relation to the amount of glucose to be 64, l o o Therefore, the linear range of this type of sensor is narrow (far below 1 mmol 1-l) when they are covered only with a hydrophilic membrane which would be highly permeable to the two substrates of the glucose oxidase reaction (see above). In extracorporal blood glucose monitoring this is usually compensated for by means of sample dilution via double-lumen 316

catheters."' 5 8 , 6 1 , 64 In intracorporal glucose measurement, however, membranes with relatively low permeability to glucose relative to oxygen must be generated. Different approaches have been reported to solve this problem.56,59, 60, 64, 67, 73, 96 In the glucose electrodes described above we have dealt with this by means of the hydrophobic membrane which must be perforated to allow some limited passage of The diameter of the perforation and its position in relation to that of the platinum anode and to that of the entire membrane are major determinants of sensitivity, response time, and of linearity of these sensors. Taking in vitro calibration curves at different oxygen partial pressures, we were able to demonstrate these sensors as linear over the euglycaemic range even at hypoxic levels. However, elevated blood glucose concentrations (such as 20 mmol 1-l) were not reliably monitored if the Po, was lower than it was normally in arterial blood, namely 13 kPa.27 Considering that the half-saturation constant (K,) of glucose oxidase for the co-substrate oxygen is approximately 0.2 mmol I-', 42, 46 and that the oxygen concentration in blood is between 0.04 and 0.12 mmol 1-l (physically dissolved) and between 6 and 8 mmol I-' U. FISCHER

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(total concentration), respectively, it i s important to ensure that a given type of oxygen-dependent sensor would function properly in the tissue environment. In our laboratory, this was accomplished by means of simultaneous monitoring of Po, and the output of the electrochemical glucose sensor by implanting two appropriate electrodes, both in normal and in experimentally diabetic dogs.82 No difference could be ascertained in tissue Po, between the two groups of animals: normal (n = 8) 7.2 k 1 . 7 (? SD) vs diabetic (n = 5) 7.3 1.8 kPa when the arterial Po, was 12.6 0.8 and 12.7 2 0. 7 kPa, respectively. Furthermore it was demonstrated in both groups of animals when they were anaesthetized that induced alterations in tissue Po, over a range expected under physiological or pathological conditions (between 2 and 20 kPa) do not alter the output of glucose sensors under hypoglycaemic, normoglycaernic, or hyperglycaemic conditions. One result from this study is shown in Figure 7.

*

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Time (min) Figure 7. Effect of alterations in tissue Pc)2 on the current of subcutaneously inserted glucose oxidase/H,O, glucose sensor (upper panels open symbols) in an anaesthetized diabetic dog kept intentionally near-hypoglycaemic by means of a bedsidetype artifical B-cell which was run by means of external settings on the basis of the peripheral-venous plasma glucose levels (upper panel, filled circles). Bottom panel: arterial (filled circles) and subcutaneous (open symbols) Po, in response to variations in O2 content of inspired air. Intravenous glucose injections (arrows in top panel) were performed at the beginning and at the end of the experiment to check the performance of the glucose sensor in situ (After Fischer et with permission) GLUCOSE SENSORS

Taken together, no significant influence of fluctuation in tissue Pc’, on the functional stability of implanted glucose electrodes is to be expected as long as the sensor i s working within its linear calibration range. Similar conditions can be expected in peritoneal fluid where oxygen tensions of between 40 and 70 mmHg have been reported. l o o

Current Achievements and Pitfalls As yet, animal experiments have been reported on subcutaneous glucose measurement, using so-called needle-type amperometric sensors in rats,2 1, ;” sheep,’ pigs,74 and dog^.^"-^^. Only very few preliminary attempts in man have come to our attention.4h, I o 2 , I U 3 In most cases, stable sensor function could be guaranteed over a few hours only. We have been using self-manufactured electrodes and membrane-sandwich formulations which were processed under clean-room conditions but were not sterilized. The typical features and drawbacks of implanted sensor function are demonstrated in Figure 8 where the telemetered sensor current is related to frequently measured plasma glucose values in a free-moving dog before, during, and after identical glucose loads on different days after sensor implantation. During the first 4 days, the pattern of blood glucose is well mirrored by that in sensor current. But there i s a progressive decrease in the estimated sensitivity of the implanted sensor, and unexplained oscillations appear in the basal current. In the end the sensor was removed because of definite loss in sensitivity. This result was accompanied by some perisensor oedema and exudative reaction when the animal was still in good general condition and did not show any febrile reaction or leucocytosis. Most importantly, in a group of experiments like this, lasting between 12 and 96 h, there was a remarkable difference in glucose concentration between the inflammatory exudate sampled after removal of the sensor and the simultaneously measured peripheral venous plasma glucose concentration: 1.6 2 1.4 vs 5.8 -+ 2.0 mmol I-’ (n=18). The individual findings were related neither to the duration of implantation nor to the existence of diabetes nor to the biomaterial applied to the sensor surface (regenerated cellulose vs polyurethane), but they were dependent on the local conditions at the implantation site and on the outer diameter of individual sensors which was between 0.8 and 3.0 mm. From this observation it i s concluded that the in situ malfunction of glucose sensors is a function of the entire system (’interstitial glucose concentration plus implanted electrode’) through the low and irregular glucose concentration in the exudate. This in turn is mainly an expression of tissue-incompatibility of the sensor. It should also be noted that recently the accumulation of catalase enzyme activity has been reported to be another potential cause of decrease in sensors’ ~ensitivity,~’since it interferes with the 31 7

REVIEW

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nA nAmmol-' I B 0.45 0.17 0.34 0.24 0.16 Figure 8. Subcutaneous glucose monitoring in a non-diabetic dog over 9 days. The sensor was permanently connected to a wearable voltage sourcehonitor. Readings were taken once a day over the indicated intervals when an intravenous glucose infusion was applied. The numbers in the bottom panel give the apparent basal current (A) and sensitivity (6)as assessed by means of linear regression analysis of sensor current (interrupted lines) vs plasma glucose estimates (dots). In vitro calibration: A = 1.1 nA, B = 0.98 nA mmolF'l. For methods see von Woedtke et 0.37

-0.19

reproducible stoichiometric relation between generated H 2 0 2 and prevailing glucose concentration.

Automated Feedback Control of Insulin Administration with the Help of Subcutaneous Glucose Electrodes One of the crucial points in insulin administration is the optimal timing of individual insulin doses. This applies even to automated feedback control of insulin dose, when the infusion rate must respond as early as possible to any alteration in the intracorporal glucose concentration,6' in order to prevent unphysiological fluctuations in blood glucose levels and to reduce required insulin dosage and thus related hyperinsulinaemia. From this point of view, the response time of glucose sensors appears important. Thus, the kinetic properties of the sensing system should, if possible, be incorporated into an artificial B-cell algorithm. In vitro characteristics of sensors in response to square-wave alterations in ambient glucose level in both stirred and unstirred samples have shown that:

1 . there was practically no 'dead time' in sensor response after alteration of glucose concentration; as evaluated by nonlinear regression analysis may be rather long (up to 8 min with a change from 5 to 10 mmol I-' glucose concentration); sensor kinetics are in most cases practically identical before and after an in vivo application, but T,, dependence on glucose concentration increases.

2 . the response time T,,

3.

Even more important is an analogous in vivo design for sensor testing. Thus Figure 6 shows that even the output of sensors which are characterized by an in vitro T,, of approximately 7 min do excellently reflect circulating glucose concentration as followed by two independent reference methods. From observations like

31 8

0.16

this it i s concluded that, for medical applications, the 'natural' lag time, namely the delay in the system 'implanted sensor plus subcutaneous glucose compartment') is a more important determinant of patterns of sensor current than the 'technical' delay. To 'answer the question as to whether differences in the response times of an artificial B-cell might severely influence the quality of metabolic control, a model experiment was designed for diabetic dogs.27 Intravenous glucose-controlled insulin infusions were applied on the basis of blood glucose measurements, employing the same standard proportional-differential algorithm for insulin dosage. But the dosage intervals were set at either 1 or 5 min which means that either each minute or every 5 min a newly adjusted micro-bolus of insulin was infused, dependent on the actual blood glucose level and on its rate of change. There was no difference in the blood glucose curves between the two protocols. Main conclusions from this experiment were that basically the artificial control of intracorporal glucose concentration in diabetes should be possible by means of subcutaneously implanted glucose sensors of a delay within the range described, and that glucose monitoring at intervals of some minutes might be a useful approach. This could be important both for data processing (for example sensor calibration) and for the regeneration of the implanted sensor, if required. This was confirmed by a study in diabetic dogs.66 The an imals were receiving feedback-control led intravenous insulin infusions on the basis of monitoring of either the apparent subcutaneous or the peripheral venous glucose concentration. The same proportional-differential algorithm was used in all trials, the dosage interval was 1 min, and calibration of both sensors (paracorporal blood glucose and subcutaneous glucose monitoring) was achieved according to simultaneous plasma glucose values. Stable blood glucose control was obtained both under basal conditions and during an oral glucose U. FISCHER

Dm tolerance test, and nearly identical insulin doses were required in the t w o situations. This is in agreement w i t h other observation^.'^ It raises the hope that automated intracorporal glucose control under ambulatory conditions will b e genuinely closer t o reality once problems of clinical sensor application over longer intervals are solved.

REV1EW 3.

4.

Conclusion As yet, practically relevant applications of intracorporal glucose sensors have only been reported w i t h subcutaneous amperometric glucose oxidase/H,O, electrodes. When appropriate sensor design and manufacture are used, it can be shown that neither the availability of oxygen within the sensor compartment nor the specific kinetic properties of such electrodes are major drawbacks in their application. Furthermore, the necessary calibration of electrodes in situ can be achieved b y using reference information on the actual blood glucose concentration. The sensitivity of such sensors per se as checked in vitro, remains stable in situ as long as their membranes are mechanically intact. Thus future research has t o tackle the following points: The principal problem comes from the lack of functional biostability of the system ’implanted sensor plus its biological environment’. This i s a bioi ncompati bi Iity reaction mediated b y irregu larl y low glucose concentration in the surrounding exudate even if the amount of this exudate i s very small. The selection of more appropriate biomaterials for outer sensor membranes, the further miniaturization of sensors, and the development of appropriate sterilization procedures are the m a i n tasks i n this field. Research should also b e directed t o the development of devices that might b e used without permanent skin penetration. Finally, those more advanced n e w biosensor approaches that are not yet ready for any intracorporal application, such as enzyme ISFETs, SPA (scintillation proximity assay), o r bioluminescence devices, should be developed further as potential tools for intracorporal use.lo4

Acknowledgement Some of the results discussed here were obtained i n studies w h i c h were part of the research project HFR M 2 8 of the Ministry of Health of the former GDR. The collaboration of P. Abet, E. Brunstein, K. Rebrin, T. von Woedtke, E.-J. Freyse, and E. Salzsieder is gratefully acknowledged. W e are grateful to G. Reach, Paris, for critical discussions over many years.

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