IN
DEGREE PROJECT ELECTRICAL ENGINEERING, SECOND CYCLE, 30 CREDITS
STOCKHOLM SWEDEN 2016,
Development and functionalization of stable and selective
miniaturized electrodes for glucose sensing
SARA CAVALLARO
KTH ROYAL INSTITUTE OF TECHNOLOGY
Development and functionalization of stable and selective miniaturized
electrodes for glucose sensing
Sara Cavallaro
Stockholm, September 8
th, 2016 KTH Royal Institute of Technology
School of Electrical Engineering
Department of Micro and Nanosystems
Abstract
Diabetes mellitus is a disease that leads to an unstable blood glucose concentration, with os- cillations and peaks higher than the normal range. It is a worldwide problem that is currently affecting 387 million people. To avoid further complications associated to the disease, patients have to monitor their blood glucose level multiple times per day, making this compound the most commonly tested analyte. Self-monitoring blood glucose strips (SMBG strips) and contin- uous glucose monitoring systems (CGMS) are the most widespread glucose monitoring devices nowadays. CGMS, detecting glucose in interstitial fluid (ISF), even if more advantageous than traditional devices, are still relatively invasive and painful, due to their size and needle-based insertion mechanism.
For the correct functioning of amperometric glucose sensors, having a reference electrode (RE) which is stable in the physiological environment is a crucial requirement. In this work the development of stable and selective miniaturized electrodes to be integrated in biosensors for ISF glucose monitoring is presented. In particular, a method to produce stable miniaturized irid- ium/iridium oxide (Ir/IrOx) quasi-reference electrode (quasi-RE) was firstly developed. Cyclic voltammetry was used for this purpose, and the process allowed the realization of miniaturized electrodes, stable for several days under physiological conditions. By using this technology it is possible to solve some of the common problems arising from the miniaturization of traditional RE.
Secondly, miniaturized working electrodes (WE) for selective glucose detection in presence of interfering species have been developed. In fact, to detect the analyte, amperometric glucose sensors need the application of a potential, which also oxidizes interfering species and results in an overestimation of glucose reading. To solve the problem, anti-interference membranes based on different materials and deposition techniques have been developed. In particular, in this study three different membranes have been studied. The performed techniques allowed to realize reproducible films on top of miniaturized WE. As a future step, the realized electrodes should be integrated and tested as complete miniaturized glucose sensors, to prove the feasibility of further miniaturized CGMS.
Abstrakt
Diabetes mellitus ¨ar en sjukdom som leder till en ostabil koncentration av glukos i blodet och medf¨or att glukoskoncentrationen varierar utanf¨or normala niv˚aer. Denna sjukdom ¨ar ett globalt problem som p˚averkar 387 miljoners m¨anniskors liv. F¨or att undvika vidare komplikationer p˚a grund av sjukdomen s˚a tvingas diabetespatienter att sj¨alva ¨overvaka deras egen blodglukoskon- centration flera g˚anger per dag, vilket g¨or glukos till en av de mest testade analyterna. F¨or dessa kontinuerliga m¨atningar anv¨ands idag fr¨amst s˚a kallade ”Self-monitoring blood glucose strips” (SMBG teststickor) eller ”continuous glucose monitoring systems” (CGMS). CGMS de- tekterar glukos som finns i interstitialv¨atska (ISF) men ¨aven om detta ¨ar mer f¨ordelaktigt ¨an traditionella sensorsystem s˚a ¨ar nuvarande CGMS metoder relativt invasiva och sm¨artfulla.
Detta ¨ar p˚a grund av storleken och den n˚al-baserade inf¨orningsmekanismen hos CGMS.
F¨or att amperometriska glukossensorer ska fungera korrekt s˚a kr¨aver de en referenselektrod (RE) som ¨ar stabil i fysiologiska f¨orh˚allanden. I detta arbete s˚a presenteras utvecklingen av stabila, selektiva och miniatyriserade elektroder vilka ¨ar anpassade f¨or att integreras med biosen- sorer. Dessa biosensorer ska sen anv¨andas f¨or ISF glukos¨overvakning. I f¨orsta delen beskrivs utvecklingen av en tillverkningsmetod f¨or stabila miniatyriserade iridium/iridiumoxid (Ir/IrOx) quasi-RE. I denna metod anv¨andes cyklisk voltammetri vars process m¨ojliggjorde tillverkn- ing av miniatyriserade elektroder som var stabila i flera dagar under fysiologiska f¨orh˚allanden.
Genom att denna teknik anv¨andes s˚a var det m¨ojligt att l¨osa n˚agra av de vanligaste proble- men som uppst˚ar vid miniatyrisering av traditionella RE. I andra delen beskrivs utvecklingen av arbetselektroder (eng. “working electrodes” WE) f¨or selektiv glukosdetektion i n¨arvaro av st¨orande kemiska element. Vid detektion en given analyt med en amperometrisk glukossensor s˚a beh¨ovs en potentialskillnad som ¨aven oxiderar st¨orande kemiska element vilket d˚a resulterar i en ¨overskattning av den utl¨asta glukosm¨atningen. F¨or att l¨osa detta problem utvecklades ¨aven st¨orningsfria membran (s˚akallade ”anti-interference membranes”) som baseras p˚a anv¨andandet av olika material och depositionstekniker. I denna studie unders¨oktes tre olika membraner.
Teknikerna som anv¨andes m¨ojliggjorde en reproducerbar tillverkning av filmer p˚a miniatyris- erade WE.
F¨or att visa f¨ordelarna med att vidare miniatyrisera CGMS s˚a ska i n¨asta steg de tillverkade elektroderna integreras och testas som en komplett miniatyriserad glukossensor.
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Glossary
Ag/AgCl Silver-silver chloride.
BGMDs Blood glucose monitoring devices.
CE Counter electrode.
CGMDs Continuous glucose monitoring devices.
CGMS Continuous glucose monitoring systems.
FAD Flavin adenine dinucleotide.
FADH2 Flavin adenine dinucleotide reduced.
GDH Glucose dehydrogenase.
GOx Glucose oxidase.
IFG Impaired fasting glycaemia.
IGT Impaired glucose tolerance.
Ir/IrOx Iridium/iridium oxide.
PBS Phosphate buffered saline.
PECVD Plasma-enhanced chemical vapor deposition.
PMPD Poly-(m-phenylenediamine).
PPy Polypyrrole.
quasi-REs Quasi-reference electrodes.
RE Reference electrode.
TTF-TCNQ Tetrathiafulvalene-tetracyanoquinodimethane.
WE Working electrode.
WHO World Health Organization.
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Contents
Abstract 2
Abstrakt 3
Glossary 5
1 Overview on blood glucose monitoring devices 10
1.1 Introduction . . . 10
1.2 Continuous glucose monitoring biosensors . . . 11
1.2.1 Basic principles of glucose monitoring biosensors . . . 11
1.2.2 Generations of enzyme-based amperometric glucose biosensors . . . 13
1.2.3 Calibration . . . 14
1.3 State of the art . . . 14
1.3.1 Market . . . 15
1.3.2 Research . . . 18
1.4 Motivation . . . 19
2 Experimental Setup 22 2.1 Equipments . . . 22
2.2 Cyclic voltammetry . . . 23
2.3 Open circuit potential . . . 24
2.4 Amperometry . . . 24
3 Reference electrodes 26 3.1 Introduction . . . 26
3.2 Basic Theory . . . 27
3.3 Reference electrodes . . . 28
3.3.1 Common Reference Electrode Types . . . 29
3.3.2 Pseudo-Reference Electrodes . . . 30
3.4 Results . . . 31
3.4.1 Silver Oxide quasi-REs . . . 31
3.4.2 Silver/Silver Chloride quasi-REs . . . 33
3.4.3 Iridium/Iridium Oxide quasi-REs . . . 38
4 Interfering species 47 4.1 Introduction . . . 47
4.2 Basic Theory . . . 48
4.3 Results . . . 50
4.3.1 Pt response to interfering species . . . 50
4.3.2 Nafion coatings . . . 52 7
4.3.3 Electropolymerized membranes of Diaminobenzene and Resorcinol . . . . 54 4.3.4 Electropolymerized membranes of Diaminobenzene . . . 60 4.3.5 Poly-pyrrole based membranes . . . 63
Conclusions 69
Future perspectives 71
Bibliography 76
Appendix A Table of the techniques used for Ag/AgCl wires 78
Chapter 1
Overview on blood glucose monitoring devices
1.1 Introduction
Glucose is the most widely used sugar and energy source in living organisms, from bacteria to humans. Breakdown of carbohydrates produces mono- and disaccharides, most of which is glucose. In cells, this sugar is used as an energy source during either aerobic respiration, anaerobic respiration, or fermentation. Glucose represents the key energy source for human body through aerobic respiration. It is mainly stored in the liver and muscles as glycogen, and is transported from intestines or liver to cells via the bloodstream.
The human body naturally tightly regulates the blood glucose level. In particular, the normal blood glucose range is 4-8 mM (72-144 mg/dL), whereas the pathophysiological one is 2-30 mM (36-540 mg/dL) [1]. A disease known as diabetes mellitus leads to a blood glucose concentration persistently higher than the normal range.
Figure 1.1: Millions of people and % of total regional population with diabetes by World Health Organization (WHO) in 1980 and 2014. Source: World Health Organization, Global report on diabetes, World Health Day 2016.
Diabetes mellitus is a worldwide chronic health problem, resulting from the failure of blood glucose regulation. As shown by the figure 1.1, the number of people with diabetes has risen
from 108 million in 1980 to 422 million in 2014. According to WHO, in 2012 an estimated 1.5 million deaths were directly caused by this metabolic disorder (Global Report on Diabetes, 2016). The organization declared diabetes as a global epidemic, as well as an economic burden, since about 11.6 % of the total global healthcare costs are spent on this disease. [1].
Diabetes is incurable but manageable, and it can be split in Type I and Type II. Type I diabetes (previously known as insulin-dependent, juvenile or childhood-onset) is due to the ineffectiveness of pancreas to produce sufficient insulin, and requires daily administration of this hormone. Type II diabetes (formerly called non-insulin-dependent or adult-onset) results from the inability of the body to use the produced insulin, and it is often the result of excess body weight or physical inactivity. Moreover, impaired glucose tolerance (IGT) and impaired fasting glycaemia (IFG) can also occur. They are intermediate conditions between normality and diabetes, which may evolve to type II diabetes, even though this is not inevitable.
Keeping the blood glucose level within the physiological range allows diabetic people to avoid further complications associated to the disease. Such complications include blindness, heart diseases, kidney failure, amputations, neuropathy, etc. For these reasons, a tight glycaemic control is needed.
Glucose monitoring technologies have drawn increasing attention over the past decades to help in the management of diabetes [2]. Continuously increasing research efforts have been directed towards the field of self-glucose monitoring, since these devices are greatly used by patients on intensive insulin therapy at least three times per day.
1.2 Continuous glucose monitoring biosensors
Continuous monitoring of glycaemia can significantly improve the quality of diabetes treatment.
In fact, it offers a complete picture of its temporal evolution, besides solving the discomfort of fingers pricks required by conventional methods.
Blood glucose monitoring devices (BGMDs) have evolved tremendously within the last four decades in terms of miniaturization, simplicity, greater specificity and painless sample uptake, making the market of glucose meters very competitive. To accurately function for extended in vivo use, BGMDs must be minimally invasive, painless, user friendly, biocompatible and low cost. Moreover, they must have a simple operating procedure, simple in-vivo calibration, great accuracy and reliability, stability, and enhanced memory [1].
1.2.1 Basic principles of glucose monitoring biosensors
BGMDs are based on a cost-effective electrochemical sensor, that typically employs a two- or three-electrode system composed of working (WE), reference (RE) and counter electrode (CE), manufactured on a substrate. According to the detection mode, electrochemical biosensors can be classified in amperometric, potentiometric, field effect or conductivity sensors [3] [4].
Amperometric biosensors are based on the measurement of the current produced at the WE by the electrochemical oxidation or reduction of an electroactive species. The experiments are performed by maintaining a constant potential between WE and RE, which may also serve as CE in case of low currents. The resulting current is proportional to the bulk analyte concentration, i.e. to the glucose concentration [5].
As regards BGMDs, in most of the cases the WE potential is measured against the one of the RE, and is kept constant over the entire measurement. This allows to activate the electrochemical reaction of the monitored analyte at the WE. A third electrode, known as counter or auxiliary electrode, is then used to close the circuit, in order to avoid a current flow through the RE. There are also possibilities without CE, where the RE also serves as
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CE. Although two-electrode systems are simpler, they may be less reliable. This occurs because when the RE is also used as CE, the current passes through it and this may modify its potential.
As a consequence, the potential difference between WE and RE may change, and the reaction of the analyte at the WE may not happen or it is possible to have the activation of undesired compounds.
Working electrode
The working electrode (WE) is the sensing element of glucose biosensors. In such devices, it usually consists of a Pt film deposited onto a silicon substrate through microfabrication techniques. Its surface is covered by a sensing layer containing active materials, such as enzymes, that react in a specific way with the glucose molecules. The most popular enzymes used within this layer are glucose oxidase (GOx) and glucose dehydrogenase (GDH). They belong to the family of oxidoreductases and contain strongly bound redox cofactors, that allow the electron transfer from glucose towards the WE surfaces. The resulting current then passes through an external circuit, and is converted to the glucose concentration.
A permselective membrane can be deposited on top of the sensing layer, in order to minimize the access of interfering species towards the electrode surface. Common interfering substances, present inside the human body, are acetaminophen, uric acid, ascorbic acid, dopamine, and cholesterol. If the applied potential is relatively high (i.e. +0.6 V vs Ag/AgCl), these species are also electroactive and they are oxidized at the WE. As a result, the selectivity and overall ac- curacy of the device are compromised. To solve this problem, different polymers with transport properties based on charge, size or polarity have been employed. Commonly used electropoly- merized films based on size exclusion include poly-(phenylendiamine) (PMPD), polyphenol, and overoxidized polypyrrole (PPy). Nafion is often used because of its charge exclusion mechanism, since it has a negative charge that rejects negatively charged molecules [6].
A further glucose limiting membrane can also be used, in order to limit the permeation of glucose to the WE surface, which may cause the saturation of the enzyme. Such membrane works as a mechanical filter for the analyte. Since many oxidase-based devices use oxygen as physiological electron acceptor, they are sensitive to changes in the concentration of this com- pound. In the human body, the normal oxygen concentration is about one order of magnitude lower than the physiological level of glucose. This limitation is known as ”oxygen deficit”.
When the glucose concentration is too high, the immobilized enzymes saturate, and they will not detect higher concentrations. This occurs because the stoichiometric 1:1 reaction involving glucose and oxygen can not occur. As a consequence, this deficit reduces the upper limit of glucose detection and thus, the accuracy of the sensor.
Figure 1.2: Illustration of the arrangement of a possible WE configuration.
When mass-transport limiting films are used, all the glucose reaching the WE can be con-
verted by the enzymes. This extends the operational lifetime of the probe, which maintains a linear increase until the glucose upper limits. Materials that are typically used in glucose limiting membranes include polyurethane and polycarbonate.
Figure 1.2 shows a schematic representation of a possible WE configuration.
Reference and counter electrodes
The RE is realized using a material that must provide a stable reference potential over time, and over the range of conditions associated with its intended use. This requirement is essential in order to provide an accurate detection of the analyte.
As already mentioned above, a CE is often used to close the circuit, and it is generally made of the same material of the working electrode. The current passes through it, in order to maintain a constant potential between WE and RE.
1.2.2 Generations of enzyme-based amperometric glucose biosensors
The majority of commercially available glucose biosensors make use of glucose oxidase as enzyme entrapped within the sensing layer. The first step of the reaction of GOx-based glucose sensors is the reduction of the flavin group in the enzyme (GOx(F AD)) by reacting with glucose. This yields to the reduced form of the enzyme (GOx(FADH2)), which is subsequently re-oxidated to (GOx(F AD)) by an electron acceptor (M edox), known as mediator. The regeneration of the initial state of (GOx(F AD)) in the enzymatic cycle is vital in order to let one enzyme molecule work more than once in the reaction cycle. Equations 1.1 and 1.2 show the reactions occurring during the glucose detection.
(GOx(F AD)) + glucose → (GOx(F ADH2)) + gluconic acid (1.1)
(GOx(F ADH2)) + M edox → (GOx(F AD)) + M edred (1.2) According to the kind of M edox used, amperometric glucose biosensors can be classified into three different generations[1.5]. First generation biosensors rely on the use of the physiological mediator O2, and generation and detection of hydrogen peroxide (H202) at the WE (eq. 1.3 and 1.4).
(GOx(F ADH2)) + O2→ (GOx(F AD)) + H202 (1.3)
H202→ 2H++ O2+ 2e− (1.4)
Since O2 is a physiological electron acceptor in the reaction, measurements based on hydrogen peroxide formation are simple, fast and stable, especially when it comes to miniaturized devices.
However, they have the drawbacks of oxygen deficit, and a moderate applied anodic potential (i.e. +0.6 V vs Ag/AgCl). To overcome these issues, first generation glucose biosensors usually use both a permselective membrane and a glucose limiting membrane. In such devices, the concentration of glucose is monitored through the oxidation of hydrogen peroxide at the WE.
The produced current is proportional to the amount of hydrogen peroxide, which in turn is proportional to the blood glucose concentration [6].
Second generation biosensors rely on a nonphysiological mediator to increase the electron transfer rate of such devices. Oxygen deficit represents the most serious problem for the first generation biosensing mode. Therefore, further solutions have been obtained by replacing the
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oxygen with a synthetic electron acceptor that shuttles electrons directly from the redox centre of the enzyme to the electrode surface (eq. 1.5 and 1.6) [6].
(GOx(F ADH2)) + 2M edox→ (GOx(F AD)) + 2M edred+ 2H+ (1.5)
2M edred→ 2M edox+ 2e− (1.6)
The electron acceptor can exist in both oxidized and reduced form. It can be either a solution-state mediator, or an immobilized mediator. The second one can be attached di- rectly to the enzyme, or entrapped within the sensing film. Commonly used mediators for GOx include ferrocene derivatives, conducting organic salts, ferricyanide, quinone compounds, transition-metal complexes, and phenothiazine and phenoxazine compounds [7] [8]. The result of using electron-carrying mediators is that measurements become largely independent of oxy- gen concentration. Moreover, they can be carried out at lower potentials. Interfering species are no more electroactive at the applied potential, thus it is possible to remove the permselective membrane. Nevertheless, these devices may still employ the glucose limiting membrane in order to prevent the saturation of the enzyme.
Ultimately, research is evolving towards third generation biosensors. In such devices, elec- trons are directly transferred from glucose to the electrode via the enzyme active sites. In this way, it would be possible to eliminate the mediator, and use an operating potential that is lower than that used in second generation devices, and closer to the redox potential of the enzyme. A possibility for implementing third generation amperometric glucose sensors is to use conducting organic salt electrodes based on complexes such as tetrathiafulvalene-tetracyanoquinodimethane (TTF-TCNQ) [6].
1.2.3 Calibration
As already mentioned above, when the glucose concentration rises, the hydrogen peroxide output current increases proportionally. The relationship between these values is expressed through an accurate calibration of the glucose sensor. An ideal sensor must provide a reliable real-time glucose monitoring over the entire range of glucose variation. A simple calibration is thus an essential requirement to be fulfilled. Calibration represents the last step of the design of glucose sensors. It ensures that the current signal i(t) reflects the glucose concentration of the patient at time t, Cg(t), and can be performed using one − point or two − point procedures [9].
In both methods, the glucose concentration can be determined by knowing the current output signal and the sensitivity of the sensor. If the relation between these two parameters is both known and unchanged, or predictable after implantation, then a precalibrated sensor does not require recalibration after its implantation. Differently, if this relation changes after the implantation, then the sensor must be recalibrated in vivo.
The facts that calibration may be ultimately performed by the patient, and it should be maintained constant over a period of several days represent two key issues. To overcome such problems, the process must be as user-friendly and efficient as possible. Moreover, it should be repeated during implantation, in order to detect potential variations in sensitivity.
A calibration-free operation is the ultimate goal of glucose sensors, but this would require a detailed understanding ot the sensitivity changes along with highly reproducible devices [6].
1.3 State of the art
The history of glucose biosensors began in 1962 with the development of the first sensor by Clark and Lyons, from the Children’s Hospital of Cincinnati. Their device consisted of an
enzyme-based sensor, where a thin layer containing glucose oxidase was placed over an oxygen electrode via a semipermeable dialysis membrane [10]. After that, glucose sensing technology has been greatly improved, leading to the above-mentioned three generations of glucose sensors.
Efforts and developments have been focused especially in the area of chemistry, reduction of potential interfering species, and miniaturization, in order to decrease the invasiveness of such devices.
1.3.1 Market
At the present time, the market is dominated by two big groups of glucose monitoring devices, including self-monitoring blood glucose strips (SMBG strips), and continuous glucose monitoring devices (CGMDs).
Self-monitoring blood glucose strips
Millions of diabetic people perform glucose assays daily by using self-monitoring blood glucose devices. This is especially important for the most serious patients, who need to know their glucose concentration within 20 minutes, in order to avoid complications associated to hypo- or hyperglycaemia. The majority of self-monitoring blood glucose devices rely on disposable enzyme electrode test strips.
The commercial test strips are considered highly reliable blood glucose sensors, because they are very accurate (5-10 % rms error vs a laboratory standard), fast (5-30 s assay time), reproducible, and require small volume samples (0.3-4.0 µL). They usually comprise a small volume electrochemical cell, a capillary fill, and contain a stable enzyme and a redox mediator.
Typical strips are usually about two and an half centimetre long, half centimetre wide, and they are constituted by a plastic substrate material. At one extremity they have electrical contact pads, to be connected to the meters, whereas at the other extremity they have the sensing area.
The latter shows a WE and a RE/CE, covered by the capillary chamber. The WE, and often the entire site of the capillary chamber, are covered by the enzyme and redox mediator [3].
Figure 1.3a shows a possible configuration of a blood glucose test strip.
(a) (b)
Figure 1.3: F reestyle Optium blood glucose monitoring device. (a) Test strip. (b) Glucose monitoring meter.
The control meter is typically a small, user-friendly, and battery operated device. In order to measure the glucose level, the diabetic patient pricks the finger and places a small droplet of blood on the sensor strip. The strip is then inserted into the meter (figure 1.3b), which measures and displays the glucose concentration within 5-30s.
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Since 1987, when Medisense Inc. launched the first device of this kind, there have been more than 40 commercial test strips. However, over 90% of the market is controlled by four main companies, Abbott, Life Scan, Roche Diagnostics and Bayer [6]. Figure 1.5a shows some SMBG strips actually available on the market.
(a) (b)
Figure 1.4: (a) A sampling of SMBG strips. From left to right, Accuchek Aviva, TRUE Balance, Precision Xtra, FreeStyle Lite, Ascensia Contour, Nova Max, Arkray, One Touch Ultra, One Touch Horizon. (b) FreeStyle Lite SMBG strip and pricking mechanism.
Ferrocene-derivatives, ferrycyanide, and osmium complexes are the most commonly used mediators employed by such companies. In particular, ferrocene-derivatives were used as medi- ators in Medisense ExacTechTMand Precision QIDTMdevices, whereas ferricyanide was used by Bayer and LifeScan [1]. At the present time, Abbott maintains a dominant position in diabetes care, providing osmium-complexes-based monitoring strips [1] [3].
Although SMBG strips have a lot of advantages, their use is being limited by a number or drawbacks. First of all, they are not non-invasive nor painless, since the size of the needle used to prick the finger is not negligible, thus hurting the nerves. Moreover, the finger prick mechanism leads to infection risks for the patients. However, the major inconvenient associated to these devices is that they deter the patient from frequent monitoring, neglecting the detec- tion especially of night-time variations [6]. This means that the measurements do not reflect the real and overall trend of the patient, but result in an approximation of glucose control.
Time and considerable efforts are also required to validate the reliability and accuracy of such SMBG devices. Although being convenient, they are not as reliable as the laboratory measure- ments. According to FDA, to be acceptable their error must be lower than 20 % for glucose concentrations between 1.65 mmol/L and 22 mmol/L, when compared to laboratory methods [1].
Considering these reasons, the research is currently focusing on miniaturized continuous glucose meters. These devices allow for a continuous control, that reflects the real trend of variation of the patient. In addition, their implantation is becoming less and less painful, thanks to their decreased dimensions.
Continuous glucose monitoring systems
Continuous glucose monitoring systems (CGMS) address the drawbacks related to SMBG strips, allowing for a faster and better glucose control. The main advantage is related to the fact that they are implanted for days or weeks, thus permitting a tighter glycemic control, especially in case of sharp changes in the glucose level. Since CGMS also operate during the night and can trigger alarms in case of hypo- or hyperglycaemia, they represent a solution for the major inconvenient of standard finger prick devices. Moreover, they meet the needs of patients. The
reduced sizes make their implantation less invasive, thus reducing the pain. Furthermore, the comfort of the patient is improved thanks to the fact that once implanted, the device remains in its place for hours. In this way the patient can avoid the periodic finger pricks.
Over the past decades, there have been continuous improvements in CGMS. They included surgeon implanted long term glucose monitors, systems with subcutaneous ultrafiltration, mi- crodialysis fibers and externally worn sensors, as well as reverse-iontophoretic systems [3] [11].
Recently, research in the field of real-time glucose monitoring of interstitial fluid (ISF) has gained large attention. Devices that make use of this methodology determine the blood glucose level by monitoring the glucose concentration in ISF. This is feasible because a close relation exists between the concentration of this compound in the ISF and that one in blood. In particular, when the glucose level does not vary, then the glucose concentrations in the two mediums are very similar, or related through a proportionality constant. Instead, when the level increases or drops off, the two concentrations are different, but the difference can be expressed by a grid proposed by Clarke et al. [3]. According to this model, the disparity is a function of the rate at which the glucose concentration changes.
Transcutaneous amperometric sensors based on ISF glucose detection represent the most widely used CGMS. They consist of a thin sensor inserted under the skin, typically composed of three electrodes, WE, RE and CE. These devices are usually placed in arms or abdomen by using adhesive patches, and monitor the glucose level for several days. Figure 1.5 shows some CGMS actually available on the market.
(a) (b) (c)
Figure 1.5: (a) Guardian Real-Time CGMS. (b) FreeStyle Navigator CGMS, (c) FreeStyle Lite CGMS.
Research and Market are especially aiming at further sensor miniaturizations, in order to make the measurement as painless as possible for the patient. Some of the most commonly used systems available on the market, based on H2O2-electrooxidation current (first genera- tion biosensors) include Medtronic’s Guardian REAL-Time System, and DexCom’s continuous glucose monitors.
Guardian REAL-Time System (figure 1.5a) consists of a glucose sensor, a transmitter and a small external monitor that shows the glucose level. The sensor is realized by lithography, and is based on a flexible plastic substrate, less than 1 mm in width, comprising a coplanar WE, RE and CE. GOx is immobilized on the WE surface. The sensor has a small adhesive patch and is usually placed in the abdomen. The system monitors glucose concentrations through the 40-400 mg/dL range, and can alert the patient before the glucose limit is reached. However, it has a recommended usage time of 3 days between replacements, and requires at least 1 finger prick reading every 12 hours to recalibrate the sensor. In particular, it is recommended to calibrate the sensor with fingersticks 3-4 times per day for an optimal glucose sensor accuracy [3].
Dexcom is a two electrode system that reports glucose concentrations within the same range
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of the above-mentioned device. By using Bluetooth technology it is possible to control glucose levels, trends and data from a compatible smart device, such as a smartphone. The sensor has also to be replaced every 3 days, although an improved version called Seven, with a 7 day wear-time is now available and FDA approved. A contraindication for this device is that taking acetaminophen while wearing the sensor may falsely raise the sensor glucose readings.
As regards CGMS based on artificial mediators, FreeStyle Navigator (figure 1.5b) is one of the mostly used sensors of this kind on the market. It was launched by Abbott in 2007 and consists of a disposable 5-day wired GOx-glucose sensor. It employs a vinyl pyridine polymer with pendant osmium complexes, and is implanted by the patient. The device needs a calibration with SMBG strips after 10, 12, 24, and 72h during the 120-h sensor lifetime [11].
At the present time, the most advanced sensor monitoring glucose level in ISF is the Freestyle Libre Flash system (figure 1.5c), developed by Abbott Diabetes Care. The device is intended to be a replacement for the capillary blood glucose measurements, it has a usage time of 14 days of continuous monitoring and once calibrated in the factory, it does not require any further calibration. The described characteristics make this sensor innovative among existing interstitial glucose monitoring technologies. However, its implantation is not painless. In fact, although being smaller than the existing devices, its needle insertion mechanism penetrates the skin to a depth of around 6-7 mm. This makes the sensor placement painful, thus hurting the patient [12].
Considerable time and efforts have also been focused toward non-invasive glucose sensors, in order to eliminate the pain, discomfort and challenges associated with implantable devices.
These newly designed sensors have been directed toward glucose measurements in saliva, tears [2], or sweat. The GlucoWatch sensor is an example of such devices. It was approved by FDA in 2001 as an adjunct to finger sticks, by providing information that could improve therapeutic decisions. The device is worn by the patient as a wrist watch, and monitors the glucose level non-invasively through the skin [6] [11] [13]. However, measurements demonstrated that this sensor was not as reliable as the other commercial devices, thus it was removed from the market.
The fact that non-invasive glucose sensors do not provide very reliable measurements make such devices not successful.
1.3.2 Research
In addition to the sensors described so far and many other devices available on the market, there have been continuously increasing research efforts in the area of glucose sensing. Such studies have been mainly focused on the improvement of single components, such as only the WEs or quasi-REs. In some cases, three-electrode systems have been developed, in order to investigate the performance of the final biosensors. These attempts have considered different electrode arrangements. Natural or artificial mediator-based devices have been proposed as well [6] [3].
Research in the field of REs has been directed towards the implementation of further stable and miniaturized REs for in vivo measurements. In order to realize integrated microsensors, electrodes without filling solutions and in direct contact with the test solutions have been used. Such class of devices is termed as quasi-reference electrodes (quasi-REs). In particu- lar, silver/silver chloride (Ag/AgCl) and iridium/iridium oxide (Ir/IrOx) quasi-REs have been especially investigated.
Efforts have been focused on the possibility to solve the most common problems related to Ag/AgCl quasi-REs. They includ dissolution of the AgCl layer, toxicity of the dissolved film, and filling solution effusion [14]. Such drawbacks have been limited by coating the electrode surfaces with gel or polymeric materials that slow down the dissolution rate of
AgCl. Matsumoto and Ohashi reported a micro-planar Ag/AgCl quasi-RE coated first with γ- aminopropyltriethoxysilane (γ-APTES), and then with perfluorocarbon polymer (PFCP) [14].
Tests demonstrated that the electrode remained stable for a period of 82 days, thus representing a suitable quasi-RE for plasma samples. Huang I, et al. introduced a novel agarose-stabilized KCL-gel membrane, to enhance the endurance of Ag/AgCl quasi-REs against the corrosion caused by test solutions [15].
Alternatively to electrode coatings, other possibilities to solve these problems have been focused on the use of thicker Ag layers. Such films have been deposited on the microelectrode surfaces through electroplating processes. For example, Polk B, et al. fabricated 100 µm2 electrodes by electroplating silver, greater than 1 µm thickness, onto thin-film metal chips [16].
The presence of thicker Ag films allowed to realize glucose biosensors with improved stabilities and lifetimes.
Recently, Ir/IrOx quasi-REs have gained increasing attention, since this material can solve many of the drawbacks encountered with Ag/AgCl quasi-REs. In fact, electrodeposited IrOx films have shown good performance in terms of long-term stability and reproducibility, thus representing suitable quasi-REs in microfabricated biosensors and biochips [17].
Research in the field of WEs has been directed towards the implementation of further selec- tive devices. The aim is to realize electrodes that detect glucose, while rejecting the interfer- ence caused by other molecules present inside the human body. Section 4.2 details the research strategies adopted in this field.
1.4 Motivation
There are many advantages in developing miniaturized biosensors that monitor the glucose level in ISF. First of all, such sensors are less painful than devices measuring glucose concentration in blood. In fact, for their implantation, it is sufficient to penetrate down to the dermis in order to have a direct contact with ISF. Usually, continuous glucose monitoring devices (CGMDs) are placed on the arm, where dermis starts at around 0.16 mm in depth. This means that the electrode sizes can be reduced until a length lower than 1 mm. Moreover, the smaller they are, the less invasive and painless is their implantation [18].
Furthermore, sensors based on this methodology lessen the electrode encapsulation. In fact, when a material is inserted into the human body, the immune system often identifies it as a foreign object. This triggers a cascade of inflammatory and wound healing events that end up with the fibrous encapsulation of the material. A such event would compromise the correct functioning of glucose biosensors, causing the failure of the devices [19]. However, by using biocompatible miniaturized sensors that penetrate just down to the dermis, it is possible to reduce this risk, thus expecting a longer lifetime of the sensor.
Finally, the use of these devices limits the manufacturing costs, since they are realized by using microfabrication techniques at wafer level, such as thin film deposition. This allows for an easy mass-manufacturing with high reproducibility. Thus, miniaturized biosensors that monitor the glucose level in ISF represent a promising avenue in the field of continuous glucose monitoring.
Considering the above mentioned characteristics, this thesis is focused on the study and implementation of miniaturized electrodes for glucose detection in ISF. The aim is to design and realize specific components that can be integrated into a complete miniaturized sensor, consisting of WE, RE and CE, that detects glucose in ISF. The requirements are that the final biosensor should have further miniaturized sizes that can solve the discomfort related to the current available devices. Moreover, it has to detect glucose concentrations through the entire range that is possible to find in the human body. Finally, its usage time should be as long as
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possible, in order to prevent the patient from frequent sensor replacements.
In particular, the first part of the work reports the implementation of miniaturized quasi- REs that show stability for glucose sensing applications. The second part instead focuses on the realization of miniaturized WEs that are highly selective to glucose, rejecting the interference caused by the most common interfering compounds. The sizes of such electrodes are minia- turized enough to allow the potential glucose detection in ISF, while minimizing the possible negative reactions triggered by the human body.
Chapter 2
Experimental Setup
This chapter reports the equipment and experimental procedures used to collect and analyze data during this thesis work. It describes the basic principles of such techniques, even though arrangement and procedures vary slightly from test to test. All the materials used during the experiments were purchased from Sigma Aldrich Chemical Co.
2.1 Equipments
Figure 2.1: From left to right: (a) Equipment used for experiments, including a potentiostat, an electrolytic cell constituted by three electrodes, a magnetic stirrer and a software interface.
(b) Enlargement of a three electrode system immersed in an electrolytic solution.
All the electrochemical experiments mentioned in the following chapters were performed using a DY2000 Series Multichannel Potentiostat (Digi-Ivy, Inc., Austin, TX) in a three-electrode system (WE, RE, and CE). Prior to each experiment, the setup window of the control software was configured as following:
Current polarity: Anodic positive. The displayed current direction is selected as anodic positive.
In all the performed experiments, a commercial Ag/AgCl electrode was used as RE. During this study, the potentiostat was used to carry out cyclic voltammetry, open circuit potential and amperometry measurements.
The electrolytic cell was placed on a LLG magnetic stirrer with heating (uniSTIRRER 3), in order to guarantee a more homogeneous electrolytic solution, and thus a better distribution of the compounds. The stirring was also important to decrease possible noise peaks present during the measurements. The rotating speed of the magnetic bar was adjusted according to each experiment, in order to ensure a proper mixing of the compounds and to avoid solution spills.
Figure 2.1 shows a configuration of the entire system. Each connector of the potentiostat is connected to the CE, RE and WE respectively, through electrical contacts. The system consisting of the three electrodes is immersed in an electrolytic solution, contained within a Petri dish. The dish is in turn placed on the magnetic stirrer.
2.2 Cyclic voltammetry
Cyclic voltammetry (CV) is a type of potentiodynamic electrochemical technique used for the study of electroactive species. It consists of cycling linearly the potential of an electrode, while measuring the resulting current [20]. The adopted convention is that positive-going currents are anodic and negative-going currents are cathodic. The above-mentioned electrode is used as WE, it is immersed in a solution, and its potential is measured against the one of RE, such as a silver/silver chloride electrode. The applied potential has a triangular waveform and it sweeps between two values, also known as switching potentials (fig. 2.2a).
(a) (b)
Figure 2.2: Cyclic voltammetry. (a) A possible waveform of the potential applied between WE and RE. Source: www.expertsmind.com. (b) Typical cyclic voltammogram.
The use of a potentiostat allows to vary the values of switching potentials (V), as well as the scan rate (V/s), the number of cycles, the step (V) and the sensitivity (A/V).
The graph plotting the current (vertical axis) versus potential (horizontal axis) is called cyclic voltammogram. Figure 2.2b shows its typical shape, even though it slightly varies according to the chosen parameters, the electrolytic solution and the electrode material.
Cyclic voltammetry can be used for electrochemical studies of compounds, biological ma- terials or electrode surfaces, as well as for electropolymerization of materials on an electrode surface. In particular, during this study, the technique was used for the realization of miniatur- ized quasi-REs, and the electrodeposition of membranes on top of WEs.
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2.3 Open circuit potential
Open circuit potential (OCP) is a technique that measures the potential difference between WE and RE, using a high input impedance. This means that almost no current flows between the two terminals. When performing this technique, after selecting the sampling time (s) and the total run time (s), the potentiostat monitors and records the potential difference for the desired period of time.
OCP measurements are used when considering reference electrodes, because they count for the stability and reproducibility of such electrodes. In particular, REs must show stable OCPs during glucose measurements. Moreover, to test the reproducibility of a RE, the OCPs must display negligible variations when considering electrodes fabricated using the same procedures and parameters.
2.4 Amperometry
Amperometry is a technique based on the measurement of current, when a constant potential difference is applied between WE and RE for a period of time. The plot displaying the resulting current as a function of time is known as amperogram. The measured current changes when an electroactive analyte is oxidized at the anode, or reduced at the cathod. For this reason, the applied potential can be adjusted in order to maximize the response to the analyte of interest, while minimizing the response to interfering species.
Amperometry is the most important technique in amperometric glucose biosensors. In fact, it represents the method used by such devices to detect the glucose level. To be reliable, a glucose sensor must show a current that is proportional to the glucose concentration, and at the same time, not be affected by interfering species. During this study, amperometric measurements were performed to test the performance of working electrodes, either bare or covered by membranes.
Chapter 3
Reference electrodes
3.1 Introduction
The reference electrode (RE) is a crucial component of electrochemical systems, since it con- tributes to the stability and reliability of the final biosensor. Amperometric biosensors function by applying a constant potential between the working (WE) and reference electrode, thus the role of this electrode is to set a reference potential that is insensitive to any change occurring in the electrolyte solution and that is kept constant over time.
Considering blood glucose sensors, this means that the RE potential has to be stable in the physiological environment and unresponsive to interfering species, as well as glucose, interstitial fluid or cells. In fact, any potential drift could cause the non-activation of the glucose reaction with glucose oxidase, as well as the activation of other compounds present in the surrounding environment. Both cases can lead to an output current that does not correspond to the real concentration of glucose. As mentioned above, a wrong measurement may result in an inap- propriate dose of insulin, which could cause coma, blindness or even death. Such reasons show that having a reliable RE, which allows a tight glycemic control, is an essential requirement for blood glucose sensors.
Several factors are needed to be taken into account when constructing a reference electrode:
The RE must have high exchange current densities and thus be non polarizable (hard to change the electrode’s potential).
The electrode potential must be reproducible.
The potential drift of the electrode, from factors due to filling solution effusion or degra- dation of electrode coating, must be minimized over the entire duration of the device [21].
Continuously increasing research efforts towards the miniaturization of blood glucose monitoring sensors have led the miniaturization of not just the sensing elements, but also of the reference electrodes. This has resulted to several implications and problems, mainly related to the size reduction of these electrodes, in terms of their surface area and thickness.
This chapter first introduces the basic dynamic operation of a reference electrode and some common types of REs used for glucose monitoring applications. It afterwards analyzes different possible microfabricated REs, focusing on Silver Chloride and Iridium Oxide quasi-REs. The central core of the study is focused on the first aim of this thesis, the implementation and in vitro characterization of a miniaturized Iridium/Iridium Oxide (Ir/IrOx) quasi-RE.
3.2 Basic Theory
An electrolytic cell is an electrochemical cell that undergoes a redox reaction when an electrical energy is applied. The flow of charge is accompanied by the electrochemical decomposition of chemical compounds, in a process called electrolysis. This leads the conversion from electrical to chemical energy [22]. Electrolytic cells refer to systems of two electrodes, a cathode and an anode, dipped in an electrolyte, a solution of water or solvent in which ions are dissolved.
According to Faraday’s definition, the electrode to which cations flow within the cell to be reduced is called the cathode. Likewise, the anode is the electrode to which anions flow within the cell to be oxidized. When current is applied to the cell, negatively charged ions migrate to the positive electrode, the anode, whereas positively charged ions migrate to the negative one, the cathode. Figure 3.1 shows a schematic diagram of an electrolytic cell.
Figure 3.1: Schematic diagram of an eletrolytic cell [23].
A system consisting of an electrode (typically metal) and an electrolyte, with a potential associated with the electrode, represents an electrochemical interface. The potential represents the energy per unit charge required to take a charge from one side of the interface to the other one [21]. Nernst equation relates the electrode reduction potential to the standard electrode potential, temperature and activity of the chemical species involved in the electrochemical reaction, as shown in Eq. 3.1.
Ered= Ered0−RT
zF ln ared aox
(3.1) where E is the electrode reduction potential, Ered0 is the standard reduction potential, R is the universal gas constant, T is the temperature in Kelvin, z is the number of electrons transferred per reaction, F is Faraday’s constant, and aredand aoxrepresent the chemical activity of relevant reduced and oxidized species, respectively.
As a cell can be thought as consisting of two half cells joined together by an external circuit, electrode potentials are often expressed in terms of half cell potentials. Moreover, a direct measurement of the potential difference between the electrode and the electrolyte is not possible, since any contact of a measurement probe with the solution phase would involve a second phase boundary between metal and electrolyte somewhere. On this basis, electrode
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potential differences are referred to an arbitrarily chosen reference electrode, the Standard Hydrogen Electrode (SHE), whose potential is set equal to zero. This is due to the fact that SHE achieves its equilibrium potential quickly and maintains it well constant over time, making it well suited for comparisons. Potentials of all other electrode systems are calculated against SHE and are termed standard electrode potentials [22].
When an electrochemical interface is at its standard potential, no reactions occur at the interface because it is at equilibrium. However, although the net current is zero, there is still a faradaic activity expressed in terms of exchange current. When a potential is applied, the flow of charges across the electrode-electrolyte interface induces a change in the electrode potential.
The magnitude of this change is determined by the exchange current density. In particular, the higher is the exchange current density, the lower is the change in potential [21]. Thus, this condition is thought as ideal for reference electrodes.
3.3 Reference electrodes
As mentioned above, the function of a reference electrode is to maintain a stable potential over time that allows the oxidation reaction of glucose to occur at the working electrode. In order to achieve this goal, the following requirements have to be fulfilled:
The RE must have an high exchange current density, must be reversible, non polarizeable, and characterized by an horizontal region in a current-voltage plot, as shown in the figure 3.2. By controlling these properties it is possible to have an exchange of charge between the electrode-electrolyte interface, without any significant change in the electrode potential.
Figure 3.2: Region of interest for a good reference electrode in the current-voltage plot [21].
The electrode reaction should not completely consume the electrode or its surface area in a reasonable time-frame, since this would result in a change in the measured current.
The solution in contact with the RE (filling solution) should be saturated and separated from the test solution. These requirements are necessary for different reasons. An high concentration of reactants allows for an higher exchange current density. Secondly, since the potential has a logarithmic dependence on the concentration, small fluctuations in the
concentrations of the saturated solution will not greatly change the electrode potential.
Moreover, a saturated inner filling solution makes the electrode potential insensitive to any evaporation of the solvent. Finally, the electrode potential is mostly due to the large net dipole at the metal-electrolyte interface. Considering that the composition of the test solution depends on the application, it would be desirable to have the electrode potential as independent as possible from the test solution. Filling solutions separated from test solutions by a porous membrane would represent the ideal solution. In this way, liquid junction potentials do not change much by changing the test solution and would prevent the electrode potential to vary greatly.
The interface between the test and the saturated solution should minimize any kind of convective mixing. Such mixing could affect both solutions by changing the reference potential, and possibly the reactions of interest in the test solution.
The liquid junction potential should be as small as possible and constant over time. This is usually done by using the same solvent in both solutions, in order to avoid net dipole moments at the interface of the porous membrane.
The characteristics of the RE must be reproducible in terms of stability and electrode potential value [22].
A problem related to REs having the test solution separated form the filling solution is that the latter can leak through the porous separator, changing both the RE potential and the liquid junction potential.
The research is currently being focused on REs without filling solution, that are in direct contact with the test solution. Such electrodes are termed as quasi-reference electrodes (quasi- REs) and eliminate the problem of the liquid junction potential. However, the main advantage is related to the fact that they are essential for the realization of integrated microsensors, where the RE is placed on the same chip of the WE and CE. In such systems it would not be possible to maintain the configuration of two separated solutions. Differently from the traditional REs, quasi-REs are in contact with the test solution, which is not saturated, thus they may not provide stable potentials. The potentials can vary significantly with the composition of the test solution. However, quasi-REs can provide a stable reference potential if the electrolyte solution is controlled and kept reasonably stable, as it is in buffered solutions at constant solute concentrations for example.
3.3.1 Common Reference Electrode Types
According to the application and test solution, different systems can be utilized as reference electrodes. Although an ideal reference electrode, suitable for every environment or applica- tion, does not exist, there are few electrode types that are more commonly used in different applications.
Ag/AgCl Electrode
The silver/silver-chloride (Ag/AgCl) electrode is the most commonly used reference electrode in biomedical research and industry. When the filling solution is separated from the test solution by a membrane and a salt bridge, the electrode consists of a silver wire, coated with silver chloride and immersed in a solution rich of chloride ions, such as KCl. Dissolution of the AgCl layer, filling solution effusion and toxicity of the dissolved AgCl represent the main issues for such electrodes, because they compromise the stability and accuracy of the RE [21].
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The redox process for a Ag/AgCl electrode is the following:
AgCl(s) + e−↔ Ag(s) + Cl−(aq) (3.2)
By inserting appropriate quantities in the Nernst equation (Eq. 3.1), it is possible to calcu- late the electrode potential.
EAg,AgCl,Cl− = EAg,Ag0 +,Cl−−RT
F ln aCl− (3.3)
The Eq. 3.3 shows that chloride ions influence the electrode potential and its stability. If the filling solution is saturated of such ions, then the reference electrode potential will minimally vary. The standard electrode potential for Ag/AgCl electrode at 25 °C is +0.22 V (vs. SHE), whereas the electrode potential in saturated Cl− solution is 0.199 V (vs. SHE).
Hydrogen Electrode
The Standard Hydrogen Electrode (SHE) is the electrode to which all other electrodes’ standard potentials are measured, since its value is set at zero Volts. It consists of a platinum surface coated with black platinum, all of which is immersed in an acidic solution. The entire cell is inserted into a glass enclosure and kept at a fixed pressure in order to maintain a stable electrode potential [21]. The needs for a proper gas chamber for hydrogen and for pressure regulators, that keep the pressure constant, make this type of electrode not so much used.
Calomel Electrode
The saturated Calomel Electrode (SCE) consists of mercury covered by a mercury chloride paste, all of which immersed in a saturated KCl solution. Although the electrode is easy to construct and has a stable potential, its toxicity makes it unsuitable for biological applications.
3.3.2 Pseudo-Reference Electrodes
At the present time, a lot of efforts are going towards the implementation of pseudo reference electrodes, especially when it comes to miniaturized devices. Pseudo-reference electrodes rep- resent a class of electrodes whose potentials vary predictably with conditions. Nevertheless, a stable potential can be obtained by fixing some parameters, such as the concentration of the analyte or the test solution pH. Differently from the traditional systems, quasi-REs lack the filling solution and the electrode is in direct contact with the test solution. The absence of the porous membrane makes the design simpler and allows lower drifts due to convective mixing.
On the other hand, the potential of such electrodes is more affected by variations in the test solution, especially when contaminating species are involved.
Most attempts in this field have been directed towards the design and development of Sil- ver/Silver Chloride quasi-reference electrodes [14] [16] [24]. Recently, Iridium/Iridium Oxide quasi-reference electrodes have received increasing attention [25] [17], and have been studied as an alternative of such Ag/AgCl electrodes.
Ag/AgCl quasi-reference electrode
Silver/Silver Chloride (Ag/AgCl) quasi-REs have been widely used in micro electrochemical sensors. To avoid the sensitivity problems of such electrodes to interfering species, coatings of different compounds have been used, even though they do not prevent shifts due to changes in Cl− concentration [21].
In these electrodes, the silver layer can be deposited using microfabrication techniques, such as thin film deposition, electroplating or screen printing. The silver chloride layer can be subsequently grown on it by chemically treating the electrode with different solutions, including KCl or FeCl3.
Dissolution of the AgCl layer represents the main problem for Ag/AgCl quasi-REs, and it is even more accentuated during miniaturization. The problem is especially related to the thickness of the layer, and speeds up as the thickness decreases. When the layer gets depleted, the underlying silver gets exposed to the test solution, creating mixing potentials [21] [25] [16].
Similarly, the poor adhesiveness of the silver layer to a glass substrate, such as SiO2, is a considerable issue. The problem is magnified for miniaturized electrode surfaces, and often causes the layer detachment after few minutes of functioning [24].
Ir/IrOx quasi-reference electrode
Iridium/Iridium Oxide (Ir/IrO2 or Ir/IrOx) represents a promising candidate for miniaturized quasi-REs. It consists of an iridium surface which is modified by using an oxidizing agent to form an oxide coating. Since the Ir/IrOx electrode potential is determined by the test solution pH, the electrode is mostly used as pH sensor. On the other hand, this strong pH dependence makes the Ir/IrOx unsuitable in applications where the pH value changes.
However, when the electrode is used in a phosphate buffered saline (PBS) solution at pH 7, the resistance of the solution itself to change its pH makes possible to use Ir/IrOx as quasi-RE.
Human blood and interstitial fluid can be assimilated to a buffer solution, since their pH stays in a very narrow pH range, between 7.31-7.45. Therefore, in medical and biological applications where pH is regulated, Ir/IrOx can be used as reference electrode thanks to its biocompatibility, mechanical stability and high exchange current [21].
Moreover, many of the drawbacks mentioned for Ag/AgCl quasi-reference electrodes can be solved by using Ir/IrOx. In fact, such electrodes do not show the problem of dissolution of their oxide layer, nor the detachment caused by the poor adhesiveness [25]. A drawback encountered with this material is that the potential is slightly sensitive to interfering species, thus it weakly changes when they are added to the PBS solution. However the problem can be solved by depositing some membranes that have the aim of rejecting such interference [26], and that will be object of the following chapter. All these properties make Ir/IrOx a promising electrode for miniaturized glucose biosensors [17].
3.4 Results
This section shows the results obtained during the implementation of a miniaturized quasi-RE, to be integrated into a continuous glucose monitoring sensor. To monitor the electrode potentials of all the fabricated devices, OCP measurements are performed. Shifts in the electrode potential within the mean value ± 5 mV are considered desirable. However, slightly greater variations can be accepted in particular conditions.
Before working on Ag/AgCl and Ir/IrOx, the possibility to realize a Silver Oxide quasi-RE was investigated.
3.4.1 Silver Oxide quasi-REs
Silver Oxide is a material that has been largely used in a variety of technological applications, including photography, batteries and optical memories. In the field of MEMS, Silver Oxide films have been grown for gas sensing applications using many different techniques. Some researchers
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have directly deposited silver oxide films by RF reactive sputtering, reactive electron beam evaporation and DC magnetron sputtering in an oxygen environment [27] [28] [29]. Others have grown silver films by using the same techniques, and then oxidized them in an oxygen plasma atmosphere or in an electrolyte solution [27] [30]. At the department of Micro and Nanosystems at KTH, Gatty et al. [31] have developed a Silver Oxide quasi-RE for an amperometric gas sensor based on nitric oxide detection.
Considering this wide use, the first step towards the implementation of a miniaturized RE was focused on the possibility to realize a Silver Oxide quasi-RE for glucose amperometric detection. By following the technique used by Gatty [30], a 5%wt H2SO4 electrolyte solution was prepared. Silver wires with a diameter of 50 µm were used as electrodes. To obtain the silver oxide film, the electrode was dipped in the electrolyte solution and a voltage of 1 V was applied between it and a platinum counter electrode. The experiments were carried out considering oxidation times from 2 minutes up to 15 minutes.
After the silver oxide layer was obtained, the electrode was placed in a phosphate buffered solution (PBS) 0.01 M and the OCP was measured over time. The choice of this electrolyte solution is due to the fact that its composition is similar to the one of the interstitial fluid, thus approximating the human body environment. The OCP was measured considering the wire as a WE, a commercial Ag/AgCl as RE and a platinum strip as CE.
The results showed that Silver Oxide electrodes are not suitable quasi-REs for glucose detec- tion applications. In fact, when the electrodes were immersed in PBS, their electrode potential were pretty unstable, ranging from +80 mV (vs. Ag/AgCl) to -50 mV (vs. Ag/AgCl). The potential shifts were thus largely greater than the ones considered acceptable. Moreover, when the OCP of an electrode stabilized, such value changed between two or more consecutive mea- surements, compromising the stability concept. Figures 3.3 a and 3.3 b show this behaviour.
Figure 3.3: (a) OCP of a Silver Oxide wire, first measurement. (b) OCP of the same Silver Oxide wire, second measurement, just after the first one. Silver Oxide layer realized by applying a constant potential of 1 V for 2 minutes. Silver wire used as WE, commercial Ag/AgCl used as RE, platinum strip used as CE.
Finally, the OCPs of different electrodes stabilized to different values, making the entire pro- cess not reproducible. The above mentioned outcomes compromise the essential requirements for a good reference electrode and demonstrate that Silver Oxide can be used as reference electrode for other applications, but the material is not appropriate for glucose sensing mea- surements. The reason is probably related to the environment where the device is supposed to operate, corresponding to PBS during the experiments. This statements can be confirmed by the fact that there are not previous attempts in the literature of using Silver Oxide as quasi-RE for glucose sensing applications.
3.4.2 Silver/Silver Chloride quasi-REs
Considering that Silver/Silver Chloride is by far the most commonly used material for reference electrodes, it seems a good candidate for amperometric glucose sensing applications. Thus, as second step, the possibility of implementing a Ag/AgCl quasi-RE was analysed. Silver wires were tested first and in case of positive outcomes, the processes were scaled to smaller strips and then microelectrodes.
Ag/AgCl wires
Similarly to Silver Oxide, silver wires with a diameter of 50 µm were initially used as reference electrodes. To obtain the AgCl layer, two main techniques, applying a constant current den- sity and a constant potential respectively, were tested. The electrolyte solution was changed according to each method. Table A.1 in Appendix A summarizes the different combinations used.
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The OCP measurements were performed at room temperature in a stirred PBS solution, using the wire as WE, a commercial Ag/AgCl as RE and a platinum strip as CE. Considering that the potential of the RE corresponds to the one for saturated Cl− solutions (+ 0.197 V vs.
SHE), whereas the potential of the WE corresponds to the one for highly diluted solutions (+
0.277 V vs. SHE), the OCP expected for those electrodes ranges between +0.07 V and + 0.08 V.
Figure 3.4: OCP of a Ag/AgCl wire. Chlorinized layer obtained by dipping the electrode in FeCl3 and applying a constant current density of 10 mA/cm2 to a surface of 1.57 mm2 for 8 minutes.
Basing upon the results obtained for the different techniques, the most stable OCP was shown in case of a constant current density of 10 mA/cm2applied for 8 minutes. Figure 3.4 shows the OCP of the Silver/Silver Chloride wire, measured for 16 hours. The measurement shows that after a stabilization for approximately 2 hours, the potential of the Ag/AgCl wire becomes stable at around + 0.082 V, with a drift lower than 7 mV over 16 hours. The negative shift after 8 hours is known to result from the evaporation of PBS during the overnight measurement.
Since the aqueous phase evaporates, the Cl− concentration increases, getting closer to the one of the commercial RE. This leads to a decrease in the potential of the WE and consequently a lower OCP. When fresh PBS is added again to the system (after almost 16 hours), the new diluted solution rises up the OCP to its initial stabilized value. To investigate the reproducibility of the system, three other wires were produced and tested by using the same techniques. Their OCPs showed the same trend.
These results confirm that Ag/AgCl is a suitable material for REs and a potential candidate as quasi-reference electrode for continuous monitoring biosensors. However, silver wires can not be easily applied in miniaturized devices, since they can not be integrated in microfabricated chips. To circumvent this problem, silver strips were used and tested afterwards.
