This journal is c The Royal Society of Chemistry 2011 Chem. Soc. Rev., 2011, 40, 4805–4839 4805

Cite this: Chem. Soc. Rev., 2011, 40, 4805–4839

Optical methods for sensing glucose

Mark-Steven Steiner, Axel Duerkop and Otto S. Wolfbeis*

Received 5th March 2011

DOI: 10.1039/c1cs15063d

This critical review covers the present state of the art in optical sensing of glucose. Following

an introduction into the significance of (continuous) sensing of glucose and a brief look back,

we discuss methods based on (a) monitoring the optical properties of intrinsically fluorescent or

labeled enzymes, their co-enzymes and co-substrates; (b) the measurement of the products of

enzymatic oxidation of glucose by glucose oxidase; (c) the use of synthetic boronic acids;

(d) the use of Concanavalin A; and (e) the application of other glucose-binding proteins.

We finally present an assessment in terms of the advantages and disadvantages of the various

methods (237 references).

1. The significance of sensing glucose

The quantitation of glucose is among the most important

analytical tasks. It has been estimated that about 40% of all

blood tests are related to it. In addition, there are numerous

other situations where glucose is to be determined, for example

in biotechnology, in the production and processing of various

kinds of feed and food, in biochemistry in general, and in

numerous other areas. The continuous interest in sensing

glucose, mainly in blood, is one result of the increasing age

and (meanwhile alarming) size of the world’s population and

the fact that about 4–5% of its (Caucasian) population suffer

from diabetes. The significance of sensing glucose is best

documented by the numbers of hits that can be found when

consulting (06 May 2011) Google (B4750000 hits) or Scholar

Google (B324000 hits). MedLine/SciFinder combined yields

B4000 references on ‘‘glucose sensor’’ as entered, and B14 800

references containing the concept ‘‘glucose sensor’’ (search

performed on 6 May 2011). Wikipedia has a most readable

article on blood glucose monitoring.1 Obviously, there is

substantial public concern about diabetes and sensing glucose.

Given the significance of sensing glucose, it comes as a kind

of surprise that few books are available that cover the subject

in depth. Cunningham and Stenken2 probably have authored

the most authoritative survey. The book by Bartlett3 covers

electrochemical sensors only, and the one by Pickup et al.4

fluorescent sensors only. The special issue on glucose sensing

as published by Geddes and Lakowicz5 contains selected

aspects of fluorescent sensors also.

Institute of Analytical Chemistry, Chemo- and Biosensors, Universityof Regensburg, D-93040 Regensburg, Germany.E-mail: otto.wolfbeis@chemie.uni-r.de

Mark-Steven Steiner

Mark-Steven Steiner, born1983, studied chemistry at theUniversity of Regensburgfrom 2002–2007. He obtainedhis PhD in Analytical Chemistryin 2010 at the University ofRegensburg under the super-vision of Prof. Wolfbeis. Hiscurrent research is focused onfluorescent methods for use inbio-targeting and bio-imagingusing luminescent upconvertingnanoparticles, also in combina-tion with RGB-based signalreadout using digital cameras. Axel Duerkop

Axel Duerkop, born 1973,graduated in chemistry at theUniversity of Regensburg andearned a PhD in 2001under the supervision ofProf. Wolfbeis. He is an‘‘Akademischer Rat’’ (SeniorResearcher) and presentlyworking at his habilitation. Hisresearch interests cover opticalsensors, test strips and micro-plate assays, luminescent probesfor hydrogen peroxide, formetabolites of cancer cells,and for cations and anions.Lanthanide complexes and

transition metal complexes are preferred probes to be used aslabels, in immunoassays (based on anisotropy and decay time), forchemosensing of thiols, DNA and saccharides.

Chem Soc Rev Dynamic Article Links

www.rsc.org/csr CRITICAL REVIEW

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4806 Chem. Soc. Rev., 2011, 40, 4805–4839 This journal is c The Royal Society of Chemistry 2011

The market for glucose sensors probably is the biggest single

one in the diagnostic field, being about 30 billion h per year at

present. Given this size, it is not surprising that any real

improvement in sensing glucose (in whole blood and elsewhere)

represents a major step forward. This is true for any type of

sensor discussed below. The largest need at present is, however,

in a continuous sensor. Unfortunately, after more than 30

years of intense research this appears to be more challenging

than flying to the moon, albeit not in terms of money but of

ingenuity. Chemists, biochemists, engineers, and various kinds

of medical experts have intensely cooperated in the past but to

little avail. No doubt, substantial progress has been made

(particularly in terms of electrochemical sensing; see below),

yet the ultimate goal of an implantable glucose sensor that

would automatically trigger the release of insulin if a certain

level of glucose concentration is exceeded (in a so-called

artificial pancreas) has not been accomplished. It still

represents the ‘‘Holy Grail’’ in biosensing.

Aside from continuously sensing glucose in blood, other kinds

of sensors are needed. One type is the classical sensor for the

(central) clinical lab that is capable of determining glucose in

samples as small as 30–100 mL and within one to two minutes.

Such assays are of the high-throughput type, for example by

making use of microfluidics or other flow systems. A second large

market is in near-patient (point-of-care) testing, both inside and

outside a hospital and including bedside testing. The homecare

market probably is the largest of all. Such a widespread use of a

single test became possible because test strips and sensors have

become disposable and are easy to work with, instrumentation is

small and affordable, and population is technically skilled so that

they themselves can take care of sampling blood (1 to 10 mL) andtesting it using portable analyzers.

True sensors (i.e. sensors that respond to glucose in a fully

reversible way) are not needed in near patient testing and in

the homecare market for obvious reasons. As a result, such

tests also can be based on irreversible reactions such as in

certain visually read-out test strips. This market is well covered

by electrochemical sensors6 (for example those based on

mediated electron transfer in glucose oxidase-catalyzed reactions)

which can be manufactured at low costs so that they may be

disposed after use even though they may be used again. Such

glucose sensors are not expected to be biocompatible, to require

substantial maintenance by the patient, to display long storage

lifetime and operational lifetime, and to have a response that

does not drift over the time of storage. Sensors for homecare

applications have to be calibration-free, however.

Clinical multiparametric (but discontinuous) instrumentation

is in widespread use, for example, for sensing glucose along

with other blood parameters including pO2, pH, Na+, K+,

Cl�, lactate or urea. The respective sensors are of the re-usable

type in a sense that a blood sample is inserted into the

instrument, a reading is made, the surface of the sensors is

washed, and the sensors are recalibrated before the next

sample is being introduced. Such instrumentation obviously

needs true (fully reversible) sensors for proper operation, or

the sensors can be regenerated by chemical means which is less

elegant and compromises the frequency of assays. A sensor

that would be applicable to all the situations where glucose is

to be determined does not exist yet. Sensing glucose in the beer

brewing industry is less of a challenge than sensing glucose in

the blood of the critically ill after a cardiac infarct.

Electrochemical methods are most established, mainly in the

form of stand-alone instruments in clinical labs and in near

patient testing. Millions of disposable electrochemical

(mediator-based) blood glucose meters are used in homecare

devices7 that enable glucose to be determined within less than

30 s in blood samples as small as 1 mL. The work of Heller and

Feldman6 on electrical wiring of enzymes has led to a new

generation of glucose sensors (that have had a tremendous

commercial success so far, first at TheraSense Inc., later at

Abbott Diabetes Care Inc.). These sensors have (sub)micro

dimensions and require even smaller quantities of blood to be

taken, thus leading to almost painless sampling which represents

a big relief to diabetics.

Optical methods are based on the measurement of photons

rather than of electrons. This has certain advantages, for

example, in the case of patients with heart pacemakers or

when sensing glucose under the action of strong electromagnetic

fields as used in cancer therapy. Fiber optic sensors, in turn,

enable glucose to be sensed in the deeper lying or less-

accessible regions of the body. Optical sensors also do not

require a reference electrode, can sense through optically

transparent walls (thus enabling sterile remote sensing), and

are capable of multiplexing.

Optical schemes for sensing glucose have not had, however,

the success of electrochemical schemes, but still are a matter of

highly active research. Among the optical methods,

absorptiometry (and reflectometry) and fluorescence and surface

plasmon resonance (SPR) have had the biggest success.

Almost all optical sensors for continuous monitoring rely on

either fluorescence or SPR. No reflectometric or interferometric

method is known that would enable continuous sensing of

glucose in blood, even though such methods have been

Otto S. Wolfbeis

Otto S. Wolfbeis, born 1947,is a Professor of AnalyticalChemistry. He has authoredmore than 500 articles ontopics such as optical (fiber)chemical sensors, analyticalfluorescence spectroscopy,and fluorescent probes, editeda (widely used) book on FiberOptic Chemical Sensors andBiosensors, acts as the editorof the Springer Series on-Fluorescence, is the Editor-in-Chief of Microchimica Acta,and one of the ten curators ofAngewandte Chemie. His

h-index is 52, and his articles have been cited >11 000 times.Several sensors developed in his group have been commercia-lized. His present research interests include fluorescent bio-sensing, the design of novel spectroscopic schemes, newfluorescent probes, beads, and labels, new methods of interfacechemistry, and analytical uses of advanced materials such asupconverting luminescent nanoparticles and graphenes. Also see:www.wolfbeis.de.

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This journal is c The Royal Society of Chemistry 2011 Chem. Soc. Rev., 2011, 40, 4805–4839 4807

described for solutions of glucose in plain water which,

however, is not realistic. Chemi- and bioluminescence-based

methods also are confined to discontinuous sensing.8 A

tremendous hype was noticed in the time between 1993 and

2005 when sensors were announced that would sense glucose

in vivo through skin using near infrared spectroscopy (at

wavelengths between 900 and 1000 nm, where glucose has a

weak absorption band).9 This has ceased meanwhile, and such

sensors are not covered in this review. The literature in this

article is deemed to be virtually complete as per May 2011.

2. Classification of sensors

Unlike most previous reviews on biosensors of various

kinds,10–13,14 this one is subdivided into sections according

to the method for recognition of glucose rather than according

to the method of detection. Selective recognition (combined

with selective metabolism in the case of enzymes) is a

prerequisite for selective detection of glucose, and this can

be accomplished by various means. Once the binding event has

occurred, it has to be transduced into an (optical) information.

Five fundamental types of recognition have been identified

and form the main sections in the following review.

The first type (covered in Section 4) is based on the

recognition of glucose by certain enzymes (or coenzymes) that

subsequently undergo changes in their intrinsic absorption

and/or fluorescence, or carry a (fluorescent) label placed near

the site of interaction. The second large class of sensors

(covered in Section 5) relies on the measurement of the

formation or consumption of metabolites as caused by certain

enzymes, mainly glucose oxidase (GOx). Such sensors are

kinetic by nature. The most successful ones are based on the

measurement of (a) the oxygen consumed, (b) the hydrogen

peroxide produced, or (c) the acid produced in the reaction. In

the case of dehydrogenases, the reduction of the co-substrate

NAD+ to form NADH with its characteristic band at 455 nm

may be exploited. Such ‘‘sensors’’ consume a reagent

(NAD+). Thus, they are not reversible but may be reversed

by other means.

The third large group of sensors for glucose (covered in

Section 6) relies on the capability of organic boronic acids to

act as molecular receptors for saccharides, more precisely for

1,2-diols. The affinity of boronic acids towards saccharides is

not very high, which is a fortunate situation because glucose

levels in blood samples are rather high (typically from 3–50 mM).

Respective boronic acids have been designed by various groups

and undergo a change in the optical properties as a result of

binding glucose.15–17 The binding event can be detected by

various optical means including fluorescence and SPR.

The fourth group of receptors for glucose (see Section 7) is

based on the affinity of glucose to the plant lectin concanavalin A

(ConA). Respective sensors are based on competitive binding

of glucose and a labeled carbohydrate such as dextran or a

glycated protein. The fifth large group of receptors (see Section 8)

exploits the capability of glucose-binding apoenzymes and

glucose-binding proteins (GBPs). This group also includes

apo-GOx, a glucose oxidase whose coenzyme has been

removed. Binding of glucose still does occur, but the subsequent

step of oxidation is not possible any longer. Both GBPs and

GOx undergo changes in their intrinsic optical properties on

binding glucose which, however, can be detected in the UV

only. Therefore, they have been labeled with fluorophores to

shift the change of the optical signal into the visible range of

the spectrum. Methods based on labeled proteins are preferred

because the choice of a proper label enables the optical

properties of the system to be fine-tuned. These proteins cover

a wide range of concentrations of glucose, and genetic

engineering has further shifted the dynamic ranges in the case

of blood glucose towards higher concentrations.

3. A look back

Numerous chromogenic schemes have been developed for the

determination of glucose, the early ones often being based on

the use of aggressive reagents, requiring elevated temperatures,

being rather slow, or of limited general applicability. Most of

these methods are tedious and destructive, and none is applicable

to continuous sensing.

A major breakthrough occurred when enzymes came into

use. These convert glucose into products that are more easily

detectable than glucose itself which is not colored, nor fluorescent,

and has electrochemical properties that are not significantly

different from several accompanying species. Both glucose

dehydrogenase and glucose oxidase have been widely used

ever since in various formats. These include (a) cuvette and

microplate assays, (b) flow systems such as flow-injection

analysis (FIA), (c) chromatographic separations, and (d) the

solid state chemistry format (also referred to as ‘‘dry’’ chemistry)

in so-called test strips, all however in a discontinuous manner.

Large numbers of samples can be handled by methods such as

flow injection analysis, batch injection, or lab-on-a-chip techno-

logies, often in combination with automated sampling.

The first sensing schemes for true on-line sensing (both

electrochemical and optical) have been reported several

decades ago. One is based on the measurement of the quantity

of oxygen consumed according to eqn (1) that is catalyzed by

GOx. Alternatively, the H2O2 formed according to (1) may be

determined by electrochemical or optical means. A third

option consists in the determination of the quantity of protons

formed (i.e. the decrease in pH) (eqn (2)).

b-D-glucose + O2 - D-glucono-1,5-lactone+H2O2 (1)

D-glucono-1,5-lactone + H2O - gluconate + H+ (2)

The enzyme glucose dehydrogenase also has been used to sense

glucose. It catalyzes the conversion of glucose to form a

gluconolactone according to eqn (3):

b-D-glucose + NAD+ - D-glucono-1,5-lactone + NADH

(3)

The amount of NADH formed according to eqn (3) may be

measured, for example, by photometry at 345 nm or via its

fluorescence peaking at 455 nm, but this reaction cannot be

easily reversed and comes to an end once all NAD+ is

consumed. Hence, it is less suited (and less elegant) in terms

of continuous sensing. The electrons transferred in eqn (1) can

be directly shuttled onto an electrode by so-called direct

enzyme wiring (a direct electron transfer from an electrode

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to the reaction center, either by mediators or by incorporating

nanowires directly into the enzyme).6,7 Sensors employing

mediators are in widespread use ever since the 1990s, and

sensors based on nanowires since the year 2000.

All present-day commercial optical sensors rely on the use

of GOx. Respresentative (larger) manufacturers include

OptiMedical Inc. (www.optimedical.com/products/opti/opti_cca_

touch.htm), Idexx Inc. (http://www.idexx.com/view/xhtml/

en_us/smallanimal/inhouse/vetlab/vetstat-electrolyte-and-blood-

gas.jsf?conversationId=837252; for animal care); Becton-

Dickinson Comp. (in blood; www.bd.com/ds/productCenter/

BC-Bactec.asp), Teruma Inc. (www.terumo-cvs.com/products/

ProductDetail.aspx?groupId=3&familyID=47&country=1).

4. Sensing glucose via the optical properties

of intrinsically fluorescent or labeled enzymes,

their co-enzymes or their co-substrates

This class of sensors is making use of enzymes and coenzymes

that undergo optical changes in their spectral properties upon

binding glucose. Typical enzymes include glucose oxidase (GOx),

glucose dehydrogenase, and glucokinase. Apo-enzymes such as

apo-GOx also bind glucose but do not metabolize it. Respective

schemes are treated in more detail in Section 8. On binding

glucose, the intrinsic (UV) fluorescence of the protein part of the

enzymes undergoes substantial changes in intensity. The absorption

spectra of the protein part of the enzymes, in contrast, do not

change. The coenzyme FAD displays absorption and luminescence

in the visible, and both change on interaction with glucose.

In order to shift the analytical window to the longwave

range of the visible spectrum, the respective enzymes have

been labeled with (usually longwave) fluorophores. Longwave

sensing is highly desirable in view of the strong intrinsic

absorbance and fluorescence of blood, serum and urine.18,19

Labelling usually does not strongly affect the binding constants

of the enzymes.

4.1 Non-labeled enzymes

The intrinsic fluorescence of glucose-converting enzymes is

due to the UV fluorescence of tryptophan (excitation/emission

maxima at 295/330 nm). The fluorescence of FAD occurs at

exc./em. maxima of 450/520 nm, that of NAD+ at exc./em.

maxima of 340/460 nm. Hussain et al.20 have immobilized

yeast hexokinase (that binds glucose and converts it into

glucose-6-phosphate in the presence of ATP) in a silica sol–gel

and observed an up to 25% quenching of fluorescence at 330 nm

on addition of glucose. The analytical range (1–120 mM) and

insensitivity to blood serum was improved21 by covering the

enzyme layer with a glucose-permeable membrane. The

increase in fluorescence is linearly related to glucose in

concentrations up to 20 mM. In another embodiment, GOx

was entrapped in a gelatine membrane resulting in an analytical

range from 2 to 20 mM. This setup was compared to sensors

based on measurement of the FAD fluorescence with the

conclusion that UV excitation results in larger dynamic

ranges. A kinetic method also was described as an alternative

to steady state UV fluorescence analysis (Fig. 1).22,23 GOx was

entrapped in a sol–gel, and this resulted in an analytical range

from 0.5 to 20 mM of glucose.

Photoexcitation at above 400 nm is more adequate for

determination of glucose in real life samples because of the very

strong UV absorption of proteins and other species. A look at

the mechanism through which glucose is oxidized by GOx reveals

that FAD, a yellow coenzyme with a strong intrinsic green

fluorescence, is converted into its reduced form (FADH2) before

being backconverted to FAD by molecular oxygen:

b-D-glucose + FAD - D-glucono-1,5-lactone+FADH2

(4)

FADH2 + O2 - FAD + H2O2 (5)

Trettnak and Wolfbeis24 were the first to report on an optical

glucose sensor based on the intrinsic green fluorescence of

FAD. GOx was entrapped in a semi-permeable membrane at

the end of an optical fiber (Fig. 2). The fluorescence at above

500 nm was monitored and found to increase upon addition of

glucose within the (narrow) range from 1.5 to 2 mM. Response

times are from 2 to 30 min. Related studies, with the enzyme

incorporated into a sol–gel, were reported later.25 The sol–gel

method also was studied26 with respect to the complex

Fig. 1 Scheme of the GOx reaction. Glucose (G) reduces the FAD of

glucose oxidase to FADH2 under formation of gluconolactone (L),

which is rapidly hydrolyzed to gluconic acid (AG). Dissolved oxygen

reoxidizes and produces H2O2 as a result. This last product is

converted to water and O2 by the enzyme catalase. The intrinsic

fluorescence of GOx at 334 nm (lexc. 278 nm) increases in the presence

of glucose. Reprinted with permission from ref. 23. Copyright 1997

American Chemical Society.

Fig. 2 Cross-section through the sensing platelet of a fibre-optic

glucose sensor. P, Plexiglas; D, dialysing membrane; E, enzyme

solution; 0, O-ring; L, light guide. The arrows indicate the diffusion

processes involved (G, glucose; GL, gluconolactone) and the

directions of the exciting light (Exc) and fluorescence (Flu). The

platelet has o.d. 20 mm; diameter of cavity, 4 mm. Reprinted from

ref. 24, with permission from Elsevier.

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fluorescent transitions of the emission of FAD. The analytical

range of their sensor is from 0.4 to 5 mM. Also see ref. 27.

Sanz et al.28 reported on a detection scheme that is based on

the combined use of GOx and horseradish peroxidase. The

absorbance of HRP at 424 nm changes on exposure to

glucose, and this is attributed to changes in the absorption

of the heme group of HRP due to oxidation. The sensor was

prepared by entrapping HRP and GOx in a polyacrylamide

(PAA) gel. HRP is oxidized by the H2O2 generated in the GOx

reaction to form the so-called HRPI, in which the heme group

is in a virtual 5+ oxidation state. Tyrosine separately added

regenerated the HRP. The sensor covers the 1.5 to 300 mMconcentration range and has a long-term stability of at least

6 months. More sophisticated variations also were reported.29,30

The sensors work in whole blood (after dilution), and display a

long term stability of over 30 months and more than 200

measurements. Response times range from 10 s to 5 min. The

dynamic range can be increased to 2 mM by bubbling oxygen

through the solution.

An example of the few assays based on the use of glucose

dehydrogenase (GDH) was reported by Narayanaswamy and

Sevilla.31 GDH was immobilized on a nylon mesh cartridge

mounted on an optical fiber. The intrinsic blue fluorescence of

the cosubstrate NADH increases linearly on addition of

glucose in the range from 1.1 to 11 mM. The limit of detection

(LOD) is 0.6 mM. In fact, dehydrogenases play a much more

important role in electrochemical sensing32 than in optical

sensing. A GDH-based glucose sensor with an expanded

dynamic range was constructed33 using an engineered enzyme

which allows for an expanded and higher dynamic range than

that of the wild type protein. The His775 of a GDH from

E. coli was substituted for Asp and then showed an increased

Michaelis–Menten constant as demonstrated in a conventional

colorimetric assay which has a dynamic range from 0.5 to 30 mM

of glucose and with less than �5% error.

4.2 Labeled enzymes

Aside from the use of the intrinsic optical properties of

glucose-specific enzymes, the optical properties of labeled

enzymes also have been studied (Table 1). GOx immobilized

on poly(amidoamine) dendrimers on microscope slides reversibly

binds meso-tetra(4-carboxyphenyl)porphine (CTPP4) which has

an absorption maximum at 427 nm. Exposure of the complex to

glucose causes a linear decrease in absorbance in the range from

1.1 to 11 mM glucose34 due to dissociation of the complex.

A thermostable glucose kinase from a thermophilic microorganism

was applied35 in a competitive FRET assay in which glucose

derivatized with o-nitrophenyl-b-D-glucopyranoside serves as

a quencher of the intrinsic tryptophan fluorescence of GOx.

Addition of glucose decreases the quenching efficiency.

Compared to a more simple assay where the emission of

GOx labeled with an anilino-naphthalenesulfonate derivative

is quenched, the first system displays a larger signal change.

GOx was also labeled with a coumarin derivative.36 Its blue

fluorescence increases by up to 10% in the presence of glucose

in the 0.5–6 mM concentration range. The same group also

used fluorescein-labeled GOx (entrapped in a sol–gel) which

has a more red-shifted excitation and emission along with

Table 1 Sensing glucose via the optical properties of oxidative or reductive enzymes. AR: analytical range; BSGK: Bacillus stearothermophilusglucokinase; FLU: fluorescence; GLU: glucose; GOx: glucose oxidase; GDH: glucose dehydrogenase; RT: response time

Enzyme Method AR/mM Ref.

Hexokinase Intrinsic UV fluorescence (exc./em. 295/330 nm) of Trp decreases on addition of GLU due toconformational change; enzyme entrapped in sol–gel; no phosphorylation, no glucose consumption

1–120 20

GOx Intrinsic UV fluorescence (exc./em. 278/335 nm) of Trp in GOx and of coenzyme FAD increaseson addition of GLU due to conformational change; silica gel entrapped

0.2–20 21

GOx Intrinsic UV fluorescence (exc./em. 278/340 nm) of Trp in GOx increases on addition of GLUdue to conformational change; GOx entrapped in a gelatine membrane; time course studied

2.5–20 22

GOx Intrinsic UV fluorescence (exc./em. 278/335 nm) of Trp in GOx and of coenzyme FAD increaseson addition of GLU due to conformational change; sol–gel entrapped

0.5–20 23

GOx Intrinsic green fluorescence (exc./em. 450/500 nm) of the coenzyme FAD increases on addition ofGLU due to conformational change; entrapped in a semipermeable membrane

1.5–2 24

GOx Intrinsic green fluorescence (exc./em. 450/520 nm) of the coenzyme FAD increases on addition ofGLU; entrapped in sol–gel

— 25

GOx HRP and GOx entrapped in PAA gel; absorbance of the heme-group of HRP changes when HP(produced by GOx and GLU) oxidizes HRP (424 nm)

0.001–0.3 28

GOx HRP and GOx entrapped in PAA gel; absorbance of the heme-group of HRP changes when HP(produced by GOx and GLU) oxidizes HRP (424 nm), continuous mode

0.001–0.05 29

GOx Absorbance at 490 nm (FAD)(chosen because of properties of optical system) decreases onaddition of GLU; covalently attached to nylon net

2–10 30

GDH GDH immobilized on nylon mesh cartridge mounted on optical fiber; intrinsic blue fluorescence(NADH) (exc./em. 340/460 nm) increases on addition of GLU; RT: 5 min

1.1–11 31

GOx Absorbance at 427 nm decreases on addition of GLU due to dissociation of a meso-tetraphenyl-porphine–GOx complex; immobilized to poly(amidoamine) on microscope slides

1.1–11.1 34

BSGK Quenchometric FRET, donor: Trp emission (exc./em. 290/340 nm), acceptor: o-nitrophenyl-b-D-glucopyranoside. Increase of donor emission at 340 nm on addition of GLU; demonstrated forsolutions only

1–6 35

Labeled GOx Enzyme labeled with 7-hydroxycoumarin-4-acetic acid, fluorescence emission (exc./em. 327/452 nm)increases on addition of GLU; demonstrated for solutions only

0.5–6 36

Labeled GOx Labeled with fluorescein derivative, fluorescence emission (exc./em. 492/515 nm) increases onaddition of GLU; GOx immobilized in sol–gel

0.6–5.6 370.5–8.3 38

Labeled GOx Labeled with fluorescein, fluorescence (exc./em. 489/520 nm) increases on addition of GLU; GOximmobilized on polyacrylamide; flow injection method

2–11 39

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4810 Chem. Soc. Rev., 2011, 40, 4805–4839 This journal is c The Royal Society of Chemistry 2011

better long term stability.37 The method was further

optimized38 with the intent to enlarge the dynamic range and

shorten the response time Eventually, a system was presented39

based on the use of GOx labeled with fluorescein and

incorporated into a PAA polymer for use in a flow injection

setup. The sensor is stable for more than 3 months and was

applied to glucose determination in soft-drinks.

5. Sensing glucose viameasurement of the products

of enzymatic oxidation of glucose by GOx

Such sensors are kinetic by nature and based on the measurement

of either the consumption of oxygen, the production of

hydrogen peroxide, or the production of protons (due to

formation of gluconic acid from gluconolactone) as outlined

in eqn (1) and (2). The concentrations of these species can be

directly related to the concentration of glucose, provided that

the activity of the enzyme remains constant. Most sensors are

of the steady-state type, and constant signals are obtained only

in systems where the sensor is exposed to a continuous flow of

sample.

5.1. Sensing via measurement of the consumption of oxygen

caused by the action of GOx

The most widely used (and commercially successful) approach

is based on the measurement of the consumption of oxygen

using probes whose fluorescence is quenched by oxygen.

Typical probes for oxygen include luminescent complexes of

ruthenium, platinum or palladium which are strongly

quenched by oxygen. The probes usually are immobilized in

a sensor layer with a thickness of typically 2 mm, and the

enzyme is immobilized in—or on—such a sensor layer. Alter-

natively, and in particular context with intracellular sensing,

the components have been immobilized on (nano)particles.

The numerous sensors described in the literature differ from

each other mainly in the kind of fluorescent probe, the type of

polymer matrix, and the way of immobilizing the enzyme.

Various kinds of polymers have been used including hydrogels,

chitosan, proteins from silk worm, egg shell membranes,

various kinds of sol–gels, but also hydrophobic polymers such

as polystyrene where the enzyme has to be immobilized on its

surface. Numerous technical layouts have been reported for

such sensors. Many are of the planar sensor layer type. These

can be placed, for example, in a microwell or a microfluidic

flow cell. Others are based on the use of optical fibers with the

sensor material fixed at its tip. Given the number of papers on

such glucose sensors, this chapter is subdivided according to

the layout of the sensors, i.e. in sections on planar and fiber

optic sensors, and on nano-particle-based sensors.

5.1.1. Planar and fiber fluorescent optic sensors. It is obvious

from eqn (1) that the concentration of glucose is related to the

consumption of oxygen caused by the enzymatic reaction,

provided that oxygen is present in excess and enzyme activity

remains constant. Various probes have been reported whose

fluorescence or lifetime is quenched by molecular oxygen.

Decacyclene, complexes of Pt(II) or Pd(II) with porphyrins,

Ru(phen), Ru(bpy) and Ru(dpp) are among the most used

indicator dyes because they can be excited with visible light.

The metal complexes are preferred because they show a large

Stokes’ shift, possess relatively long decay times and good

photostability.

The typical signal obtained with a flowing sample is

characterized by an increase in fluorescence intensity that is

related to the concentration of glucose and referred to as the

response phase. This is followed by a steady state phase.

Evidently, the shape is dependent on factors such as sensor

setup, i.e. on whether analyzing standing, stirred, or flowing

samples, (b) the availability of oxygen (large excess is

preferred), and (c) the activity of the immobilized GOx. The

fundamental setup of such a sensor is as follows: The enzyme

is immobilized on (or in) a polymer, mostly a hydrogel or

polyacrylamide. The oxygen sensitive indicator dye is immobilized

in the same or a second polymer layer. The use of two layers

enables the oxygen probe to be incorporated into a hydrophilic

polymer such as polystyrene, silicone or ethylcellulose which

are permeable to oxygen but not to proteins, glucose and

electrolytes, which may interfere.

A decacyclene based quenchometric sensor was reported40

back in 1988. GOx is immobilized on a nylon membrane on

top of a silicone layer containing the quenchable dye. Both

layers are deposited on a polyester film and a polyacrylate

solid support. Blue excitation light and green emission light is

guided through fibers directly attached to the support. The

sensor was placed in a flow cell (simulating blood flow) and

capable of determining glucose in the physiological range

(0.1–20 mM) with a response time of 1–6 min. The method

was further improved41 to obtain shorter response times by

cross-linking GOx with glutardialdehyde on a layer of carbon

black deposited on a silicone layer containing the oxygen

probe decacyclene. The black layer also served as an optical

isolation so to prevent serum fluorescence to interfere.

Response times are as short as 8 to 60 s, and the analytical

range is from 0.01 to 2 mM. This sensor was in commercial use

for almost 10 years. Decacyclene thereafter was replaced by a

ruthenium dye. A similar setup was reported by Dremel et al.42

for the on-line monitoring of glucose concentrations in animal

cell cultures. GOx was immobilized on controlled pore glass

(CPG) and fixed in an enzyme reactor flow-through cell

together with an oxygen sensor placed at the tip of a fiber

optic waveguide. The response of the system is linear for

0–30 mM with response times from 50 to 80 s.

In another version, the Al(III)–ferron complex was used as a

transducer for oxygen.43 Glucose oxidase was covalently

immobilized on a nylon membrane, the metal chelate on an

anion-exchange resin, and both packed into a flow-through

cell. Measurements were performed with flowing air-saturated

solutions. The response is linear in the range from 0.5 to 2.5 mM

of glucose with a limit of detection of 80 mM, and glucose was

determined in serum and beverages.

Pt(II) and Pd(II) porphyrins are another popular group of O2

indicator dyes. They are characterized by good brightness

(Bs; defined as quantum yield multiplied with molar absorbance),

long luminescence lifetime, large Stokes’ shifts and good

photostability. Papkovsky44 applied a platinum(II) octaethyl-

porphyrin ketone dissolved in polystyrene (PS) as a probe for

oxygen in combination with immobilized GOx. By measuring

either changes in fluorescence intensity or lifetime, glucose was

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determined in the range from 0.2 to 20 mM with a response

time from 2 to 5 min. Shorter response times were also

reported,45,46 but at the expense of analytical ranges. The

Bayer corporation has patented47 a glucose sensor consisting

of an oxygen-sensing layer containing styrene-acrylonitrile

copolymer and platinum-octaethylporphyrin as a coating on

a light transmissive substrate, and a layer of glucose oxidase in

an acrylamide copolymer on the oxygen-sensing layer.

Fluorimetric sensor layers suffer from the fact that they

cannot be easily read without instrumental assistance. This

problem was overcome48 in a method for direct colorimetric

readout using a three-layer sensor film. Green-emitting

CdTe–CdS quantum dots were incorporated in a base layer

as a stable color background. A red-fluorescent platinum–

porphyrin layer acts as the oxygen-sensor (the red emission

being quenched by O2). The top layer contains GOx. Oxygen is

consumed if the sensor is exposed to glucose, and this results in

a color change from green to red as can be seen in Fig. 3. Its

good resolution (�0.2 mM) and a detection range from 0 to

3.0 mM make this approach an interesting ruler for optical

inspection.

Ru(II) complexes with ligands like bipyridyl (bpy), 1,10-

phenanthroline (phen), or 4,7-diphenyl-1,10-phenanthroline

(dpp) are widely used as probes for oxygen. They display large

Stokes’ shifts, long lifetimes and adequate brightness. They

can be adsorbed on silica gel beads, incorporated in films of

silicone, polystyrene, ethylcellulose or ormosils, or in hydro-

gels. Often, scattering material (such as SiO2 or TiO2) are

added to increase the intensity of the emission but also to

block any interfering light.

In a typical example,49 the ruthenium probe Ru(bpy) was

absorbed on silica gel, incorporated in a silicone matrix

(with its high oxygen permeability), and placed at the tip of

an optical fiber (Fig. 3). GOx was then linked to the surface

with glutardialdehyde. The sensor responded linearly to air-

saturated flowing solutions of glucose in the range from 0.06

to 1 mM. Hoffman-La Roche has patented50 a glucose sensor

that consists of GOx immobilized on polystyrene nanoparticles

and Ru(dpp) entrapped in poly(tert-butylstyrene) nano-

particles. Both were incorporated in a polyurethane film and

covered with a carbon black containing membrane for optical

isolation and diffusion control. A miniaturized glucose sensor

based on Ru(phen) as oxygen transducer was reported by

Rosenzweig and Kopelman.51 The ruthenium complex and

glucose oxidase are incorporated into a PAA polymer and

covalently attached to a silanized fiber tip by photocontrolled

polymerization (see Fig. 4). Response times as short as 2 s

within an analytical range of 1–10 mM were accomplished.

A similar method was presented by Wang et al.52 They

embedded GOx and the Ru (dpp) complex in an ormosil–PVA

composite film in order to detect glucose in blood samples in

the range from 0.5 to 3 mM.

An interesting approach for immobilizing monolayers of

GOx consists in the use of two-dimensional crystalline bacterial

surface layers (S-layers) composed of identical (glyco)protein

subunits as matrices.53 Due to their crystalline character,

S-layers exhibit a characteristic topography with a defined

arrangement and orientation of functionalities. A biosensor

was designed with monomolecular layers of glucose oxidase

covalently immobilized on the surface of S-layer ultrafiltration

membranes. The enzyme monolayer was attached to a layer of

polystyrene containing Ru(dpp) as the luminescent probe for

oxygen. The sensor responds within 100 s and over the 1 to 80mM

glucose concentration range.

Rather than using organic (synthetic) polymers as solid

supports for immobilizing enzymes, materials from natural

sources may be used. Examples include the use of biological

membranes like eggshell membranes or swim bladder

membranes54,55 which were employed in glucose biosensors

by immobilizing GOx on their surface by established methods.

The membranes were placed on top of an oxygen-sensing layer

comprised of a (quenchable) luminescent ruthenium complex

that was deposited on silica particles and then mixed into an

air-curing (1-component) silicone as described earlier.56 The

sensor was applied to the determination of glucose in flowing

samples of beverages.

A glucose biosensor based on co-immobilization of

Ru(phen) and GOx within nanoporous xerogels also was

described.57 It operates in the frequency domain and exploits

the effect of O2 consumption on the excited-state lifetime of

the luminophore which increases if oxygen is consumed. It is

stable, reproducible, and provides an analytically reliable

response from 0.5 to 15 mM glucose. Choi and Wu58 also

have entrapped GOx and the Ru complex in a xerogel

composite derived from tetraethylorthosilicate and hydroxy-

ethyl carboxymethyl cellulose. The entrapped GOx displays a

long-lasting biocatalytic activity (up to 3 years) compared to a

conventional sol–gel matrix. The analytical range is from

9.0 mM to 100 mM, with a response time of 6–9 min. The

sensor was applied to the determination of glucose in urine.

Other biomatter that may serve as mechanical supports for

immobilization of GOx include bamboo inner shell membranes59

and tomato skin.60

Another hybrid material was applied61 in a sensor for in situ

continuous monitoring of glucose in biotechnological production

processes and showed response times of 20 s. The optically

sensitive coatings were prepared from inorganic–organic hybrid

polymers containing a Ru complex and GOx, and applied to

lenses, decladded polymer fibers, and to polymer clad silica

fibers. The response was measured via luminescence lifetime.

Glucose concentrations were measured between the detection

limit (0.1 mM) up to 30 mM. One sensor was used for 30 days

in a bioreactor. Microtiter plates (MTPs) with integrated

glucose biosensors also have been reported by Duong and

Fig. 3 Apparent colors of the sensor layers at different concentrations of glucose at 35 1C. From ref. 48 with permission from Elsevier.

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Rhee62 and Chang et al.63 GOx and Ru(dpp) were each

immobilized on the bottom of the wells of an MTP. Glucose

was determined in concentrations up to 28 mM. MTPs with

integrated oxygen sensors are commercially available.64

A needle type of sensor was reported for determination of

fish blood glucose.65 It comprises an 18 gauge needle acting as

the container, and a fiber optic probe containing the Ru

complex and immobilized GOx at its distal end. The coating

was prepared from GOx, a water-soluble photopolymer and

an ultra-thin dialysis membrane. The optic fiber is inserted

into the photopolymerized and rolled enzyme membrane and

placed in the needle. The sensor responds to glucose in the

range from 0.2 to 1 mM. One assay is completed within 3 min.

Glucose sensors based on the measurement of oxygen

consumed by enzymatic oxidation give reliable results only if

(a) the calibration and measurements are performed at about

the same level of oxygen (pO2), (b) if pO2 does not significantly

change over time (e.g. during continuous sensing), and (c) pO2

does not drop such that it becomes critically rate-determining.

If this is not the case, independent knowledge of the pO2 is

highly desirable in order to correct the signal that is related to

glucose (the signal of the biosensor) for the actual pO2. In

critical cases (i.e., if pO2 becomes too small) this signal may

serve as an alarm.

Wolfbeis et al.40,66 have tested various assemblies of thin-

film glucose biosensors capable of compensating the effect of

varying pO2, and have presented algorithms for calculating

glucose concentrations if the pO2 is not constant. In the first,

GOx was sandwiched between a sol–gel layer doped with

Ru(dpp) and a second sol–gel layer composed of pure sol–gel

(the ‘sandwich’ configuration). In the second, a sol–gel layer

doped with Ru(dpp) was covered with sol–gel entrapped GOx

(the ‘two-layer configuration’). In the third, both GOx and a

sol–gel powder containing GOx were incorporated into a

single sol–gel phase (the ‘powder configuration’). Addition

of sorbitol was reported to be essential for all configurations,

which results in a more porous sol–gel. The sandwich

configuration provides the highest enzyme activity and the

largest dynamic range (0.1–15 mM), but suffers from a distinct

decrease in sensitivity upon prolonged use. The two-layer

configuration has the fastest response time (50 s), while the

‘powder configuration’ provides the best operational lifetime.

The storage stability of all configurations exceeds 4 months if

stored at 4 1C. Sol–gel also was used as a matrix in similar work.67

A fiber-optic dual sensor was described for the continuous

and simultaneous determination of glucose and oxygen with

Ru(bpy) as the transducer probe.68 Two sensing sites were

placed at defined positions on the distal end of an imaging

fiber (see Fig. 5). Each sensing site contains an individual

polymer cone covalently attached to the activated fiber surface

using localized photopolymerization. The oxygen sensor consists

of a double-layer polymer cone. The inner polymer cone is a

hydrophobic gas-permeable copolymer containing the Ru dye,

and the outer layer is a poly-HEMA polymer. GOx is

immobilized on this layer in the case of the glucose sensor. The

fluorescence images of both sensing sites are captured with a

CCD camera. Glucose calibration curves were obtained under

varying oxygen pressures with a limit of detection of 0.6 mM

glucose. The response times vary from 9 to 28 s, depending on

the thicknesses of the enzyme layer. The range of response is

variable by immobilization of GOx with different activities.

Klimant et al.69,70 reported on dual sensors that exploit this

scheme (see Fig. 6). Two commercially available fiber optic

sensors for oxygen were placed in subcutaneous tissue in close

proximity. One sensor was modified with GOx and the other

serves as the reference. This sensor is insensitive to variation in

oxygen tension to a wide extent, and to slight fluctuations of

temperature. Glucose can be monitored in the physiological

range up to 20 mM with a response time of 84 s. Comparable

sensor schemes have been patented by Minimed Inc., Baxter

and Becton Dickinson.71–73

Another implantable microsensor was described74,75 where

GOx and the oxygen transducer Ru(dpp) were entrapped in

calcium alginate microspheres (Fig. 7). These were coated with

polyelectrolyte multilayers containing an oxygen-insensitive

green-emitting reference dye. Ratiometric determination of

glucose in a flow-through setup was achieved for concentrations

up to 0.8 mM within a response time of 2 min. The response of

the microspheres was mathematically modeled.76

Variations in the pO2 of a sample are one source of error in

oxygen-based detection schemes. The O2 transducer is often also

affected by temperature. This drawback can be compensated by

Fig. 4 Photographs of fiber-optic glucose biosensors. (a) Sensor

prepared from an unpulled, single-mode, 3–5 mm core optical fiber.

The scale bar represents 50 mm. (b) Sensor prepared from a pulled,

micrometre-sized optical fiber tip. The scale bar represents 10 mm.

Reprinted with permission from ref. 51. Copyright 1996 American

Chemical Society.

Fig. 5 Cross-sectional view of the glucose and oxygen sensing sites of

a fiber-optic dual sensor for the continuous and simultaneous

determination of glucose. Reprinted with permission from ref. 68.

Copyright 1995 American Chemical Society.

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dual sensors for oxygen and temperature. Nagl et al. reported on

such a sensor.77 The oxygen probe Pt(porph) was dissolved in a

layer of polystyrene, and the temperature probe, an europium

complex, in poly(vinyl methylketone). Dual lifetime determination

was applied to monitor the consumption of O2 in the GOx

catalyzed oxidation of glucose at varying temperatures

(Table 2). Dual sensing schemes78 have been reviewed.

5.1.2. Sensors based on microparticles and nanoparticles.

Sensor particles with probes entrapped have attracted sub-

stantial interest in the past, not so much in terms of blood

glucose testing, but for intracellular uses.79 However, they

possibly may be applied in the future in the bloodstream as

molecular analytical machines reporting blood glucose levels,

provided the optical signal they are giving can be interrogated.

The group of Kopelman80 have incorporated GOx, the oxygen

indicator sulfo-Ru(dpp) and an oxygen insensitive reference

dye in 45 nm polyacrylamide nanoparticles for self-referenced

ratiometric measurements. Glucose was determined in the

range from 0.3–5 mM and the response time was 150 s

(Table 3). The nanoscaled sensors (so-called PEBBLEs), the

inert sensor matrix and the typical longwave emission

(at >600 nm) permitted implantation of the PEBBLEs into

living cells with minimal perturbations to their biological

functions and by background luminescence.

In a different approach, CdSe/ZnS core–shell quantum dots

(QDs) were conjugated to GOx and horse radish peroxidase

(HRP) to design a resonance energy transfer based (but rather

slow) sensing scheme.81 The QDs upon irradiation act as

electron donors, whereas GOx and HRP are used as acceptors

during the oxidation of glucose to gluconic acid. Electron

transfer between the redox enzymes and the electrochemical

reduction of hydrogen peroxide (or oxygen) occur rapidly,

resulting in an increase of the turnover rate. This non-radiative

energy transfer results in quenching of the emission of the QDs

that is proportional to the concentration of glucose in

concentrations up to 28 mM.

Rossi et al.82 of the Rosenzweig group reported on magnetite-

based nanoparticles covalently functionalized with GOx. The

enzymatic activity was investigated by monitoring oxygen

consumption using the probe Ru(phen). The GOx-coated

magnetite nanoparticles act as nanometric sensors for glucose

in concentrations up to 20 mM, with a response time of 2 min.

Immobilization of GOx on the nanoparticles also increases

the stability of the enzyme and only slightly decreases its

activity. The study also revealed that the nanoparticles can

be separated magnetically from the analyte sample, thus

enabling re-use in multiple samples. Other particle-based

sensors (but based on transduction via H2O2 or pH) are

presented in Sections 5.2 and 5.3.

5.2. Glucose sensing via measurement of the formation of

hydrogen peroxide

Monitoring the formation of hydrogen peroxide (HP) produced

in the enzymatic reaction shown in eqn (1) has the advantage

of measuring against an almost zero background. However,

only few continuously working optical sensors for HP have

been reported. Most of the glucose ‘‘biosensors’’ based on HP

as a transducer are built on irreversible chromogenic reactions

allowing one-shot measurements only, but not sensing. Fluorescent

probes and nanoparticles for the (mostly irreversible) detection

of HP have been reviewed.83

In an enzymatic assay for glucose based on the fluorescent

HP probe europium(III) tetracycline (EuTc),84 the weakly

fluorescent EuTc and enzymatically generated HP form a

strongly fluorescent complex (Fig. 8). EuTc was also incorporated

in a hydrogel.85 The reaction of EuTc with HP is fully

reversible but takes about 10 min in both directions and is

strongly pH dependent. EuTc can be photoexcited at around

400 nm and responds to HP with a 15-fold increase in

fluorescence and a strong increase in lifetime, thus enabling

time-resolved measurements that can substantially reduce

background fluorescence. The method is not very sensitive

and later was extended to image glucose.86

The fluorogenic reaction of non-fluorescent Amplex Red

with an oxidant to form fluorescent resorufin was exploited in

Fig. 6 Schematic representation of (A) hybrid sensor and (B) implanted

hybrid sensor. Reprinted from ref. 70, with permission from Elsevier.

Fig. 7 Encapsulation of glucose oxidase and an oxygen-quenched

fluorophore in polyelectrolyte-coated calcium alginate microspheres as

optical glucose sensors. (a) Functional schematic of optical glucose

sensors; (b) image of spheres used for glucose sensitivity experiments.

Reprinted from ref. 74, with permission from Elsevier.

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an HP-based detection scheme.87 GOx was incorporated in a

montmorillonite clay placed on an electrode. This mineral

contains Fe(II) and Fe(III) species that catalyze the conversion

of O2 and HP to form the superoxide anion radical. The iron

species are continuously regenerated by the electrode. The O2�

radical, in turn, reacts with Amplex Red to give fluorescent

resorufin. Glucose in concentrations between 1 and 150 mMcaused an increased emission at 583 nm (Table 4).

An absorbance based glucose sensor using Prussian Blue as

a probe for HP was presented by Koncki et al.88,89 Prussian

Blue incorporated in a film of poly(pyrrolyl benzoic acid)

reacts with HP to give Prussian White. The reaction can be

Table 2 Sensing glucose via fluorometric measurement of the consumption of oxygen caused by the action of GOx. AR: analytical range; CPG:controlled pore glass; FI: fluorescence intensity; GLU: glucose; OEPK: octaethylporphyrin ketone; PVA: polyvinyl acetate; RT: response time;Ru(bpy): ruthenium tris(bipyridyl); Ru(dpp): ruthenium tris(diphenyl-phenanthroline; Ru(phen): ruthenium tris(phenanthroline))

Oxygen probe and polymer (sensor) matrix Comments AR/mM Ref.

Decacyclene contained in a silicone membrane GOx incorporated in a nylon membrane; measurement of FI(exc./em. 410/450 nm); RT: 1–6 min

0.1–20 40

Decacyclene contained in a silicone membrane GOx absorbed on carbon black and crosslinked withglutardialdehyde; measurement of FI (exc./em. 400/500 nm);RT: 8–60 s

0.01–2 41

Decacyclene contained in a silicone membrane GOx absorbed on CPG; measurement of FI (exc./em. 410/495 nm);RT: 50–80 s

0–30 42

Al/ferron complex immobilized on ananion-exchange resin

GOx immobilized in a nylon membrane; measurement of FI(exc./em. 390/600 nm); RT: 20–60 s

0.5–2.5 43

Pt-OEPK dissolved in polystyrene;microporous fiber support (PTFE, celluloseacetate, nylon, fiber glass, filter paper)

GOx crosslinked with glutardialdehyde or immobilized onCPG; measurement of FI and of phase shift (lifetime);RT: 5–90 s

0.2–20 44–46

Pt or Pd porphyrin dissolved in polystyrene Patent; GOx in polyacrylamide; measurement of FI — 47Ru(phen) contained in a silicone membrane GOx absorbed on carbon black and crosslinked with

glutardialdehyde;measurement of FI (exc./em. 460/570 nm); RT 6 min

0.06–1 49

Ru(dpp) in polyurethane film Patent; GOx immobilized on polystyrene NPs; measurement of FI;NPs incorporated in polyurethane film

— 50

GOx and probe Ru(phen) in polyacrylamide Measurement of FI (exc./em. 488/610 nm); RT: 2 s;micrometre-sized sensor

0.7–10 51

Ru(phen) incorporated in ormosil–PVA film GOx immobilized in ormosil sol–gel; measurement offluorescence phase shift (exc./em. 468/589 nm); RT: 6 s;kinetic curve simulation

0–0.5 and 0.5–3 52

Ru(ddp) dissolved in polystyrene GOx monolayer on the surface layer (ultrafiltration membrane);measurement of FI (exc./em. 465/610 nm); RT: 100 s

1–80 53

Ru(dpp) immobilized on silica particles insilicone

GOx immobilized on an eggshell membrane or swim bladder;measurement of FI (exc./em. 468/602 nm); RT: 5 min

0–1.5 5455

GOx and Ru(phen) in nanoporous sol–gel Measurement of fluorescence phase shift (exc./em. 468/570 nm) 0.5–15 57Ru(dpp) and GOx in xerogel Measurement of FI (exc./em. 460/602 nm); RT: 6–9 min 0.6–100 58GOx and probe Ru(dpp) immobilized in asol–gel composite

Measurement of fluorescence lifetime; RT: 20 s 0–30 61

Ru(dpp) immobilized in sol–gel GOx in the second sol–gel layer; measurement of FI(exc./em. 458/535 nm); RT: 30–300 s; immobilized on thewell-bottom of a microtiter plate

0–28 62

Ru(dpp) and GOx in ormosil measurement of FI (exc./em. 400/620 nm); RT: 6 min;immobilized on the well-bottom of a microtiter plate

0.1–5 63

Ru complex dissolved in a sol–gel matrix GOx in photosensitive polymer; measurement of FI(exc./em. 475/600 nm); RT: 1–3 min; needle type sensor

0.2–1 65

Ru(dpp) contained in a sol–gel matrix GOx in sol–gel; measurement of FI (exc./em. 465/610 nm);RT: 50–250 s

0.1–15 66

Ru(phen) on silica particles in silicone film GOx immobilized in sol–gel; measurement of FI(exc./em. 460/602 nm); RT: 5–8 min

0.06–30 66

Ru(bpy) contained in a poly(dimethylsiloxane)matrix

GOx entrapped in poly-HEMA hydrogel; self-referencedscheme: 2-sensor technique; one sensor measures oxygenbackground, the other the quantity of oxygen consumed;FI measured (exc./em. 470/600 nm); RT: 9–28 s

0–20 68

Commercial fiber optic oxygen sensor Sensor surface covered with immobilized GOx; measurement ofphase shift; self-referenced; RT: 9 min; used in combinationwith a microdialysis membrane

0–10 6970

Oxygen sensor and GOx-modified oxygensensor

Patent; dual sensor; measurement of FI; fluoresceinisothiocyanate, perylene dibutyrate

— 7172

Oxygen sensor in combination with GOxapoenzyme

Patent; dual sensor; measurement of FI; perylene dibutyrate orfluoranthene

— 73

GOx apoenzyme–dye conjugate modified oxygen sensor andGOx modified oxygen sensor

Ru(dpp) and GOx entrapped inCa-Alginate mPs

Measurement of FI (exc./em. 460/520 and 620 nm); RT: 120 s;ratiometric sensor; implantable; 20–30 mm; smart tattoo;reference dye (Alexa-488 assembled on poly(allylamine))

0–0.8 7475

Ru(dpp) and GOx entrapped in hydrogel mPs Measurement of FI (exc./em. 460/520 and 620 nm);implantable; smart tattoo; computer simulation model

0–10 76

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reversed with ascorbic acid. The sensor film was implemented in

a flow injection system and enabled glucose to be determined in

the range from 0.05–2 mM. The method was further optimized

to provide a double channel FIA system, and pharmaceutical,

food and clinical samples were successfully analysed.90

An SPR glucose biosensor was reported that consists of

silver nanoparticles (NPs) and GOx embedded in a stimuli-

responsive hydrogel. The HP generated in the enzymatic

reaction induced degradation of the highly clustered silver

NPs by the decomposition of hydrogen peroxide. As a

result, the distance between the silver NPs in the hydrogel

matrix is increased. This, in turn, results in an enlarged

distance between the silver NPs and a decreased localized

surface plasmon resonance. A schematic of the detection

principle is shown in Fig. 9. Glucose concentrations as low

than 10 pM were detectable owing to an increased osmotic

pressure.91

Another sensing approach based on the swelling of a GOx

modified hydrogel was demonstrated by Ye et al.92 Gratings

were photochemically produced in this hydrogel, irradiated

with a He–Ne laser, and the light intensities of the first- and

second-order diffracted beams were recorded. The system

covers the concentration range from 0.1 to 1.0 mM of glucose.

GOx was conjugated to Mn-doped zinc sulfide QDs. The HP

produced in the presence of glucose quenches the emission of

the dots. The nanosensors were successfully applied to sense

glucose in serum samples.93 A similar embodiment is making

Table 3 Nanoparticle (NP) or microparticle (mP) based sensing of glucose via fluorometric measurement of the consumption of oxygen caused byGOx. AR: analytical range; ET: electron transfer; HRP horseradish peroxidase; LI: luminescence intensity; QD: quantum dot; Ru(dpp):ruthenium-tris(diphenylphenanthroline); Ru(phen): ruthenium-tris-phenanthroline; RT: response time

Quenchable probe Polymer matrix Comments AR/mM Ref.

Ru(dpp) GOx, probe and referencedye entrapped inpolyacrylamide NPs

Measurement of LI (exc./em. 488/610 nm); RT: 150–200 s;ratiometric sensor; implantable; 45 nm; scheme referred to asPEBBLE

0.3–5 80

CdSe/ZnScore–shell QDs

GOx and HRP immobilizedon the surface of QD

Measurement of LI; ET from QD (exc./em. 485/525 nm) toGOx and HRP decreases intensity on addition of GLU;RT: 30 min

0–28 81

Ru(phen) Solution assay GOx immobilized on Fe3O4 magnetic NPs; measurement ofLI (exc./em. 460/610 nm); RT: 2 min

1–20 82

Fig. 8 Europium tetracycline (EuTc) based HP sensor. Absorbance

spectra (left) and emission spectra (right) of the EuTc/GOx system in

the absence (A) and presence (B), respectively, of glucose. With kind

permission from Springer Science + Business Media from ref. 86.

Table 4 Sensing glucose via fluorometric measurement of the formation of hydrogen peroxide (HP) caused by the action of GOx. AR: analyticalrange; EuTC: europium(III) tetracycline complex; FI: fluorescence intensity; FRET: fluorescence resonance energy transfer; LR: linear range; NPs:nanoparticles; P4S: poly-(4-styrenesulfonate); PF: polyfluorene; PFP: poly(fluorene phenylene); PLL: poly-L-lysine; PAA: polyacryl amide; QD:quantum dot

Probe Sensor matrix Comments AR/mM Ref.

EuTC Hypan GOx absorbed on the surface of a sensor placed in a microtiter plate;fluorescence lifetime imaging; exc./em. 400/616 nm; response time:10 min; increase in FI and lifetime on addition of GLU due to formationof a EuTC-HP complex

0.3–10 850.1–2 86

Amplex Red GOx inmontmorilloniteclay

GLU and Amplex Red injected in clay on an electrode; HP catalyticallyconverted to a superoxide anion radical by clay; superoxide convertsAmplex red into a fluorescent resorufin (exc./em. 563/583 nm);FI increases on addition of GLU; regeneration of clay (Fe2+/Fe3+)by the electrode; non-continuous

0.001–0.15 87

Prussian Blue (PB) PB incorporatedinto a film ofpolypyrrole

GOx immobilized on surface; Prussian White (PW) first generated withascorbic acid; HP then oxidizes PW to blue PB; absorbance at 720 nm;flow injection system

0.05–2 888990

Ag-NPs Ag-NPs and GOxin a stimuli-responsivehydrogel

HP degenerates clustered Ag-NPs and swells hydrogel; decrease inlocalized surface plasmon resonance on addition of GLU (400 nm)

10�9–1 91

Hydrogel GOx in hydrogel HP causes swelling of hydrogel; decrease in diffraction efficiency;sensing at neutral pH; response time: 1–2 min

0.1–1 92

Mn-doped ZnS QD GOx covalently labelled to Mn-doped ZnS QD; HP quenches emissionof the dots; applied to serum samples

0.01–0.1 93

CdTe QD GOx covalently labelled to CdTe QD; HP quenches emission of the dots 0.005–1 94Hemoglobin (Hb) GOx in PAA HP released by GOx oxidizes Hb 1.1–66.6 95

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use of GOx conjugated to CdTe quantum dots.94 Quenching

of the fluorescence of the QDs is again caused by the HP

produced during the oxidation of glucose.

In a flow cell based setup for the determination of glucose in

blood, GOx was incorporated in a film of polyacrylamide to

produce HP which is capable of oxidizing hemoglobin.95

This results in measurable changes in its absorbance. Glucose

can be determined in the 1–67 mM concentration range. The

quenching effect exerted by H2O2 on the luminescence of CdTe

quantum dots was exploited96 in a (solution) assay for glucose

over the 1.7 to 6.7 mM concentration range. Conceivably, it

can be extended to sensing membranes if an appropriate

polymer is found.

5.3. Glucose sensing via measurement of changes in pH

Changes of pH may also serve as the analytical information in

glucose sensing schemes (Table 5). Protons are produced as a

result of the reaction of gluconolactone with water (eqn (1)).

However, this approach is limited because often the initial pH

of the sample and its buffer capacity are unknown. Therefore,

optical glucose biosensors based on pH transduction are rarely

used in practice.

Trettnak et al.97 were the first to present such a type of

glucose sensor, with an enzymatic reaction coupled to a fibre

optic pH transducer. They used HPTS as a pH-sensitive dye

that was immobilized in hydrogel along with GOx and placed

at the tip of an optical fiber. The sensor has a response time of

8–12 min within an analytical range of 0.1 to 2 mM of glucose.

A device with immobilized GOx and FITC (acting as a pH

probe in a polymer at the end of an optical fiber) was patented

by Applied Research Systems.98 A similar fiber optic setup was

presented by McCurley.99 A cadaverine unit was linked to a

rhodamine fluorophore and incorporated in a hydrogel along

with GOx and catalase. The material was placed at the distal

end of an optical fiber. The cadaverine moiety becomes

protonated as a result of enzymatic oxidation of glucose,

and this causes swelling of the hydrogel. This, in turn,

decreases the fluorescence intensity of the rhodamine due to its

decreased concentration in the higher sample volume. Glucose

was determined in the 0 to 1.6 mM concentration range.

Polyaniline displays a pH sensitive spectrum and thus can be

used as a pH sensor probe. It turns green on formation of

acids by enzymatic reactions and was used in a microplate

glucose assay.100 The absorbance at 600 nm decreases with pH,

but increases at 840 nm. Micro- to millimolar concentrations of

glucose were determined.

A miniature optical sensor array was reported that uses

GOx covalently immobilized on cellulose acetate microscopic

beads.101 A phenoxazine derivative was incorporated into

other polymer microbeads that serve as pH sensitive probes

(see Fig. 10). Both kinds of beads, along with white reference

beads, were evenly arranged in a microarray. The reversible

color response caused by pH changes was quantified by

Fig. 9 Schematic of the detection principle of an LSPR-based optical

enzyme biosensor using a stimuli-responsive hydrogel–silver nano-

particles composite. Reprinted from ref. 91, with permission from

Elsevier.

Table 5 Enzymatic sensing of glucose via changes of pH. AR: analytical range; FI: fluorescence intensity; HPTS: hydroxypyrenetrisulfonate;MTP: microtiter plate; PANI: polyaniline; RT: response time

Probe Sensor matrix Comments AR/mM Ref.

FITC Glass tip of opticalfiber

Patent; GOx and FITC coated to glass by glutardialdehyde report changes ofpH in a polymer at the end of optical fiber

— 98

Rhodamine Hydrogel Cadaverine unit linked to rhodamine with GOx and catalase placed at the tip ofoptical fiber; enzymatically released protonation of cadaverine causes swelling ofthe hydrogel and decreases FI of rhodamine

0–1.6 99

Polyaniline PANI PANI displays a pH sensitive absorbance at 600 nm (decrease) and at 840 nm(increases) and is used itself as pH transducer; membrane becomes green fromprotons released by GOx action; calibration dependent on concentration ofphosphate buffer

1–30 100

Phenoxazinederivative

Micro array GOx on cellulose acetate microbeads and phenoxazine in polymer microbeads andadditionally white reference beads cause reversible reflectometric color responsedue to pH change in red, green and blue channel of CCD image; RT: 12 min

0–16 101

HPTS Cationic chargedbiocompatiblecapsules

pH changes cause ratiometric variation of emission spectra of capsulescontaining HPTS and adsorbed GOx

0–30 102

Azlactone Sol–gel pH sensitive azlactone and GOx embedded in hydrogel in a single layer or duallayer sensor; one layer shows faster (20 s) response but higher leaching; duallayer sensor has 40 s response time

0.1–15 103

Rhodamine Hydrogel GOx and Rhodamine derivative immobilized in a film of poly(vinyl acetate) 0.002–0.3 104NIR pH-sensitivecyanine dye

Aqueous solution Glu changes pKa of o-hydroxymethyl arylboronic acid monitored by changes inabsorbance and fluorescence of NIR pH-sensitive dye; ratiometric (640 nm/484 nm)absorbance or fluorescence (666 nm); pH 7.0; higher response to fructose over glucose

2–100 105

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reflectometry with a CCD camera and by analysis of the red,

green and blue channels. The working range of this sensor is

between 0 and 16 mM of glucose.

Another pH-based sensing scheme is making use of charged

capsules containing the pH indicator HPTS and electro-

statically adsorbed GOx.102 Local changes in pH cause a

ratiometric variation of the emission spectra (at two wave-

lengths) and can be monitored as a function of glucose

concentration between 0 and 30 mM. Ertekin et al.103 have

placed a pH sensitive azlactone and GOx in a sol–gel. A single

layer and a dual layer sensor were compared with the result

that a single layer gave a faster response but at the expense of

indicator leaching. The dual layer setup showed response times

of 40 s over the range from 0.1 to 15 mM of glucose.

Ratiometric determination of glucose was accomplished104

by entrapping a carboxy-rhodamine-modified dextran and

GDH in a film of poly(vinyl acetate) that was deposited on

a surface coated with TiO2. The ratio of the carboxy-rhodamine

emissions at 585 nm and 630 nm is changing in the pH range

between 6 and 9 and can be used to quantify glucose in a

concentration range from 2 to 300 mM. The sensor was also

tested with serum samples and is stable for at least one month

if stored at 4 1C.

Kim et al.105 presented a ‘‘sensor’’ that worked independently

from an enzymatic reaction. Glucose causes a change of the

pKa of a boronic acid. This is monitored by a change in the

absorbance and fluorescence of a NIR pH-sensitive cyanine

dye. The response to fructose is stronger than to glucose,

as is to be expected because of recognition via a simple boronic

acid (see Section 6). Glucose can be quantified in the

physiological range.

6. Sensing glucose via synthetic boronic acids

Boronic acids can reversibly interact with 1,2-diols or 1,3-diols

in aqueous solution to form 5- or 6-membered ring cyclic

esters. The rigid cis-diols found in many saccharides generally

form stronger complexes than acyclic diols like ethylene glycol

and trans-diols. The neutral trigonal form of boronic acids

transforms into the anionic tetrahedral form on binding a

saccharide, upon which a proton is released (see Fig. 11). This

reaction occurs at neutral pH values and forms the basis for an

important sensing scheme for saccharides including glucose.

The change of geometry of boronic acids on binding diols is

accompanied by a reduction of the pKa from approximately 9

to about 6. This is caused by the enhancement of the electro-

philicity of the boronic acid group on interaction with a diol.

Hence, the trigonal form, present at pH values below the pKa,

passes into its tetrahedral anionic form in the presence of a

saccharide. The pKa of the boronic acid is tunable by introdu-

cing either electron withdrawing or electron donating groups.

When attached to an appropriate fluorophore, the geometric

changes of the boronic acid also change the characteristic

features of the fluorophore’s emission, including intensity,

lifetime and polarization. Moreover, most boronic acid based

probes for glucose show either higher selectivity for fructose

than for glucose or/and give a stronger relative signal change

with fructose than with glucose. A review15 published in 2008

covers boronic acid based probes (not sensors) for micro-

determination of saccharides and glycosylated biomolecules.

This section does not cover probes but focuses on methods

intended for continuous sensing of glucose, mainly in flowing

sample solutions.

The group of Shinkai106 probably were the first to exploit

the boronic acid scheme to sensing glucose (Table 6).

Moreover—and unlike in other work—the need for discrimi-

nating fructose was addressed. A bis-phenylboronic acid

modified anthracene was synthesized (Fig. 12) that displays

photo-induced electron transfer (PET) from the nitrogen

atom of the amino group to the anthracene fluorophore. This

results in fluorescence quenching. The efficiency of PET is

Fig. 10 Micro-miniature autonomous optical sensor array for

monitoring ions and metabolites: color responses to pH, K+ and

glucose. Reprinted from ref. 101 with permission of The Japan Society

for Analytical Chemistry.

Fig. 11 Geometries of reaction products of boronic acids with water

or 1,2-diols (e.g. glucose).

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Table 6 Sensing glucose via synthetic boronic acids. Note that most methods are reported for solutions only and cannot be readily extended tocontinuous sensing. (ABS: absorbance; AR: analytical range; BA: boronic acid; BBV: boronic acid-based bipyridinium appended bis-viologen; FI:fluorescence intensity; FLU: fluorescence; FRET: fluorescence resonance energy transfer; HOLO: holography; HPTS: hydroxypyrene trisulfonate;LU: luminescence; NP: nanoparticles; PhBA: phenylboronic acid; SPR: surface plasmon resonance; SR-B: sulforhodamine-B)

Boronate Method and comments AR/mM Ref.

Anthracene based bis-PhBA FLU; increase in fluorescence on addition of GLU (exc./em. 370/423 nm); ditopicrecognition enhances sensitivity and selectivity for GLU; tested in 33% methanolicsolution; pH 7.77

0.3–1 106

PhBA FLU; computer guided design; sensor shows 400 fold greater affinity to GLU thanto other saccharides; also highest fluorescence response to GLU; up to 50%quenching on addition of GLU

— 107

Anthracene based bis-PhBA FLU; increase in fluorescence (exc./em. 377/427 nm) on addition of GLU; selectivefor GLU in the gluco-furanoidic form over fructose and galactose

— 108

3-Amino-PhBA (on naphthalicanhydride fluorophore)

FLU; quenching (exc./em. 354/400 nm) on addition of saccharide; pH 7.7; moreselective for fructose than for GLU

— 109

3-Amino-PhBA FLU; nitro-naphthalic anhydride fluorophore displays dual emission (430/550 nm)at pH 8 which is quenched on addition of saccharide; selectivity: GLU > galactose> fructose

— 110

PhBA (on naphthalicanhydride fluorophore)

FLU; sulfo-naphthalic anhydride fluorophore displays dual emission (400/474 nm)at pH 7.4 which is enhanced on addition of saccharide; selectivity: GLU> galactose> fructose

— 111

PhBA with 6-quinoliniumnucleus

FLU; quenching of fluorescence on addition of GLU (exc./em. 345/450 nm); moreselective for fructose than for GLU; AR: submillimolar; ‘‘contact lens sensor’’;works at pH 7

0.1–1 112113

BA based on stilbene, chalconeand polyene

FLU; quenching of fluorescence (em. 400–550 nm) on addition of GLU; strongresponse to fructose and GLU; ‘‘contact lens sensor’’; pH 7

0–50 114

PhBA on ruthenium complex FLU; quenching of fluorescence (em. 620 nm) on addition of GLU; pH 11; alsoresponds to fructose

0.01–0.1 115

Aryl-BA FLU; competitive assay; binding of BA to Ru(bpy)2(5,6-dihydroxy-1,10-phenan-throline) at pH 8 increases fluorescence and lifetime of the complex; addition ofGLU leads to decrease of both; patent

0–40 116

117PhBA Patent; FLU; fluorophore is selected from transition metal–ligand complexes and

thiazine, oxazine, oxazone, or oxazinone are anthracene compounds; fluorophoreimmobilized in a GLU permeable biocompatible polymer matrix that is implantablebelow the skin

— 118119

3-Amino-PhBA ABS; copolymer of aniline and 3-amino-PhBA; addition of saccharide shifts theabsorption peak at 600 nm to the shortwave; absorbance between 650 and 800 nmincreases. Response: sorbitol > fructose > mannitol > glycerol > glucose

1–30 120

Di-PhBA FLU; competitive assay; reaction of Alizarin Red S with BA leads to fluorescentproduct (exc./em. 495/570 nm); fluorescence decreases on addition of GLU;Di-PhBA shows higher stability constant for GLU than for fructose

— 121

Octylboronic acid FLU; competitive assay; reaction of Alizarin with BA leads to fluorescent product(exc./em. 460/570 nm); fluorescence decreases on addition of GLU; incorporation ofBA and Alizarin in PVC polymer; in vivo application

1–50 123

PhBA FLU; saccharide-induced conformational changes in copolymers containingBA and fluorescent units; changes detected by monitoring the excimer band(em. 430–600 nm) to monomer band (360–430 nm) ratio of intensities measured;comparable response to GLU and to fructose

— 124

PhBA (based on azo dye) ABS; color change from orange to purple on addition of saccharide; higherselectivity for fructose over GLU

5–10 126

PhBA SPR; coating of a sensor chip (gold layer) with vinylpolymer with pending BAgroups; shift of angle on addition of saccharide; higher response to fructose than toGLU

— 127

BBV FLU; study of 11 fluorescent dyes (all with negative charges) quenched by BBV(exc./em. 460/510 nm); increase of fluorescence on addition of GLU; pH 7.4;stronger fluorescence response to GLU with higher negative charge of the indicatordye

Physiologicalrange

128

BBV FLU of HPTS; study of 6 BBV receptors (all with positive charges); quench HPTS inthe presence of saccharides; best quencher: 3,30-o-BBV (followed by fructose andgalactose); apparent binding constant for GLU higher than for galactose andfructose but magnitude of fluorescence enhancement greater for fructose; fructosecauses higher fluorescence recovery

0–5 130

BBV FLU; sensor array of 6 BBV; quenching of HPTS; increase in FI on addition ofsaccharides; data evaluation with linear discriminant analysis statistics

— 131

BBV FLU; a dye and BBV covalently incorporated in hydrogel; fluorescence ofindicator dye (exc./em. 470/540 nm) is quenched by BBV; increase of fluorescenceon addition of GLU; flow cell; higher binding constant for glucose than forfructose

2.5–20 132

BBV FLU; HPTS derivative and BBV derivative covalently incorporated in hydrogel;fluorescence of indicator dye (exc./em. 470/540 nm) is quenched by BBV; increase offluorescence on addition of GLU; response time: 1–11 min; only GLU tested; fiberoptic setup

2.5–20 133

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modulated by the strength of this interaction. Increased

fluorescence through suppression of PET is observed on

binding of glucose. The cleft-like structure makes the

system particularly selective and sensitive for glucose due to

formation of an intramolecular 1 : 1 complex between the two

boronic acids and the 1,2- and 4,6-hydroxy groups of glucose.

Response therefore is much more selective to glucose than to

fructose, galactose, allose and ethylene glycol due to ditopic

recognition and formation of a 1 : 1 complex. Fluorescence

was detected at 423 nm at an excitation wavelength of 370 nm,

and glucose determined in the concentration range from

0.3 mM to 7 mM.

Table 6 (continued )

Boronate Method and comments AR/mM Ref.

BBV FLU; evaluation of different HPTS derivatives and BBV derivatives covalentlyincorporated in hydrogel; fluorescence of indicator dye (exc./em. 470/540 nm) isquenched by BBV; increase of fluorescence on addition of GLU; fiber optic sensingsetup

2.5–20 134

BBV FLU; HPTS derivative and BBV covalently incorporated in hydrogel; fluorescenceof indicator dye (exc./em. 470/540 nm) is quenched by BBV; increase of fluorescenceon addition of GLU; sensitivity fructose > galactose> GLU; fiber optic sensingsetup; patent

2.5–20 135137

BBV FLU; dye (HPTS) (exc./em. 460/510 nm) and TSPP (exc./em. 414/644 nm) isquenched by BBV; reference dye SR-B (exc./em. 565/586 nm) not quenched by BBV;increase of fluorescence on addition of GLU; ratiometric measurement; GLU testedonly

2.5–20 138

BBV FLU; fluorescence of CdSe/ZnS QDs (exc./em. 460/604 nm) is quenched by BBV;increase of FI on addition of GLU; only GLU tested

2–20 139

PhBA FLU; fluorescence of CdS QDs (exc./em. 390/604 nm) is quenched in the presence ofGLU due to shrinking of hydrogel; only GLU tested

2–25 140

PhBA on QD LU; LU of CdTe/ZnTe/ZnS QDs (exc./em. 390/604 nm) is quenched and red shiftedin the presence of GLU due to agglomerates; only GLU tested; application towardsmouse melanoma cells

0.4–20 141

PhBA in hydrogel aroundAg-NP

FLU; fluorescence of Ag NPs (exc./em. 390/600 nm) is quenched in the presence ofGLU due to agglomerats; only GLU tested

1–30 142

Anthracene based bis-PhBA FLU; increase in fluorescence and lifetime on addition of GLU(exc./em. 380/425 nm); interference studies with BSA and SDS

Physiologicalrange

144

Bis-anthracene based bis-PhBA FLU; increase in fluorescence (exc./em. 370/425 nm) on addition of GLU;‘‘selective for GLU’’; measured in 50% methanol

Physiologicalrange

145

Anthracene based bis-PhBA Patent; FLU; probe covalently immobilized on cellulose support; increase inFI on addition of GLU

— 147

Pyrene based bis-PhBA FLU; pyrene units with varying lengths of the spacer; incorporation in GLUimprinted polymer; fluorescence increase (exc./em. 343/378 nm) on additionof GLU (and other saccharides) in 33% methanolic solution only

— 148

Pyrene–phenanthrene basedbis-PhBA

FRET between phenanthrene (donor, exc./em. 299/369 nm) and pyrene(acceptor, exc./em. 342/397 nm) on addition of GLU plus decrease in excimeremission (460 nm); higher selectivity (higher binding constant and FI) for GLU;pH 8.21

— 149

Pyrene based bis-PhBA FLU; evaluation of different pyrene probes; effect of linker length, kind of fluoro-phore between two PhBA units on GLU selectivity; single pyrene fluorophore andhexamethylene linker achieved good GLU selectivity as did the pyrene/phenan-threne based bis-PhBA (FRET and excimer, see ref. 147); tested in 52% methanolicsolution only

— 150

Pyrene based PhBA FLU; emergence of excimer emission of pyrene-PhBA (exc./em. 342/377 nm,excimer 470 nm) upon addition of GLU and cationic polymer; pH 10.2; highselectivity of excimer formation for GLU

0.1–10 151

Cyclotetrapeptide based PhBA FLU; addition of GLU causes quenching of fluorescence (exc./em. 285/480 nm) of aPhBA containing a cyclotetrapeptide; solvent: 50% methanol; pH 11.7(!); highselectivity for D-GLU over L-GLU, lactate and other saccharides (cannot formditopic 1:1 binding complexes)

0–10 152

Hemicyanine derivatives basedon PhBA

FLU; increase in fluorescence (exc./em. 460/600 nm) on addition of saccharide;pH 7; higher selectivity for fructose and galactose than for GLU

5–500 153

Acridine-based PhBA Patent; FLU; acridine-based fluorophore with PhBA; em. 500 nm — 154Chlorooxazine boronate Patent; FLU; measurement of FI and lifetime — 155Dansyl PhBA Dansyl based PhBA in a plasticized PVC membrane; FLU; quenching of FLU

(exc./em. 335/530 nm) on addition of GLU; cross-sensitivity towards pH0.1–100 156

Acrylamido/Vinyl-PhBA HOLO; PhBA in hydrogel; binding of GLU induces swelling of a gel matrix;changes in diffraction wavelength; pH 7–9; RT: > 10 min

Physiologicalrange

157158159160161

PhBA HOLO; crystalline colloidal array within polyacrylamide hydrogel with PhBAgroups; binding of GLU enhances crosslinking and shrinkage of the hydrogel; blueshift of diffraction of a photonic crystal; pH 7; higher selectivity for GLU than forgalactose, mannose and fructose

Physiologicalrange

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Another bis-arylboronic acid was reported107 to recognize

glucose with 400-fold enhanced selectivity over other saccharides.

Fluorescence was measured at 447 nm in 30% methanolic

phosphate buffer solutions of pH 7.5 and found to be

quenched by up to 50% on addition of glucose. Fructose,

mannose and galactose showed much weaker responses.

Given the poor solubility, the next logical step was the

design of a more water soluble probe. This was accomplished

by introducing a pyridinium unit within the linker between the

phenylboronic acid moiety and anthracene. Other features of

the probe include rather low pKa values of 3.7 and 4.7,

respectively (attributable to the pyridinium moieties), thus

favoring binding of saccharides in neutral aqueous solution.

On excitation at 377 nm, the intensity of the emission at

427 nm increases in the presence of carbohydrates. Like most

sensors based on bis-boronic probes, the largest changes in

fluorescence intensity were observed in the presence of glucose,

followed by galactose and fructose. NMR studies108 revealed that

the glucose-probe complex exists in its a-furanose conformation,

and that binding occurs at the 1,2 and 3,5 positions, respectively.

A water-soluble saccharide receptor based on a naphthalic

anhydride fluorophore was reported by Heagy and Adhikiri.109

Its fluorescence emission peaks at 400 nm and is quenched in

neutral medium on addition of saccharides. Not surprisingly,

stronger response is found for fructose over glucose. The same

group also reported on a nitro-modified naphthalic anhydride

phenylboronic acid with higher sensitivity to glucose than

galactose and fructose.110 This probe shows dual emission at

430 and 550 nm. The former is quenched, the latter slightly

increased on reaction with glucose at pH 8. This may enable

ratiometric sensing. The naphthalic anhydride probe was

further refined111 to yield a monoboronic acid probe with a

more significant off–on response.

Quinoline based probes for determination of glucose in tear

fluid were presented by the group of Geddes.112,113 who

synthesized three isomericN-(boronobenzyl)-6-methylquinolinium

bromides (BMOQBA). The structural design considerations of

the probes were governed by the need for their compatibility with

disposable plastic contact lenses and the mildly acidic environ-

ment (Fig. 13). o-BMOQBA shows similar affinity for both

glucose and fructose. The emission maximum at 450 nm

decreases if a saccharide is added. This was interpreted in terms

of a charge neutralization–stabilization mechanism that occurs

between the quaternary nitrogen and the negatively charged

boron atom. Glucose and fructose were determined at sub-

millimolar levels with a 90% response time of 10 min. The probe

showed negligible leaching and was quenched by chloride

modestly only. The same group also tested boronic acid deriva-

tives of stilbene, chalcone and polyene for their applicability in

contact lens polymers.114 Glucose responses were considerably

reduced in the polymer due to the acidic pH value of the contact

lens polymer and its methanol-like polarity. Hence, this kind of

probes is less suitable for immobilization because high sensitivity

for glucose is needed due to its relative low concentration in tear

fluid compared to blood.

The groups of Wolfbeis115 and Lakowicz116 have introduced

assays based on the interactions of a ruthenium metal ligand

Fig. 12 Schematic of a photoinduced electron transfer based sensor

for glucose. Ditopic recognition of glucose results in a reduced PET

interaction of amine and the anthracene fluorophore and, thus, in

stronger fluorescence. Reprinted with permission from ref. 106.

Copyright 1995 American Chemical Society.

Fig. 13 Potential methods for non-invasive continuous tear glucose

monitoring. Top: contact lens doped with optical probe for glucose.

Bottom: sensor spots on the surface of the lens to additionally monitor

other analytes in addition to glucose, such as chloride or oxygen.

Sensor spot regions may also allow for ratiometric, lifetime or

polarization based fluorescent sensing. With kind permission from

Springer Science + Business Media from ref. 113.

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complex containing an aryl boronic acid derivative with

glucose. Such complexes are interesting reporter moieties

because of their long lifetimes, large Stokes’ shifts and photo-

stability. In the first approach,99 a probe was synthesized

whose chemical structure is shown in Fig. 14. In the second

approach,116 a complex is reversibly formed between Ru(2,20-

bipyridine)2(5,6-dihydroxy-1,10-phenanthroline) and the aryl

boronic acid derivatives at pH 8. Complexation is accompanied

by strong increase in fluorescence intensity at 620 nm. Addition

of glucose reduces emission intensity because of intercepting

the binding between boronic acid and a metal complex. Tests

on interferents have not been reported. Such systems were

patented.117–119

An organic conducting polymer (OCP) capable of sensing

glucose was obtained120 by co-polymerization of aniline and

3-aminophenylboronic acid (Fig. 15). This composite poly-

aniline undergoes changes in the near infrared (NIR) absorp-

tion spectrum on addition of saccharides. The absorption peak

at 600 nm is shortwave shifted, whereas the absorbance

between 650 and 800 nm increases. The polymer gives a strong

and reversible response to saccharides in the order sorbitol >

fructose > mannitol > glycerol > glucose at neutral pH. The

limit of detection is as high as 45 mM for glucose. These

polymer films are advantageous over enzyme based sensors

because they lack temperature-sensible enzymes, are compatible

with low cost LED light sources, and easily fabricated.

Arimori et al.121 report on a competitive binding assay using

a di-phenylboronic acid as the saccharide-recognizing unit that

forms a fluorescent product on reaction with Alizarin Red S

(exc./em. maxima at 495/570 nm). A comparable vesicular

fluorescent glucose sensor was later reported by Wang et al.122

but based on a cationic detergent. The amphiphilic detergent

forms self-organized vesicles in dilute aqueous solutions.

A negatively charged reporter group (alizarin red S; ARS) is

electrostatically attracted by the positively charged surface on

the vesicles to form co-vesicles. Phenylboronic acids (PhBA)

act as the glucose-recognizing unit. In the absence of glucose,

PhBA forms a strongly fluorescent boronate ester with ARS.

The ester is cleaved on addition of glucose, and the yellow

fluorescence is quenched. When compared to plain aqueous

solutions, the vesicular sensor provides a 7- to 8-fold

enhancement in terms of sensitivity. Glucose can be determined

in concentrations between 3.2 and 43.3 mM. The vesicular

sensor also responds to ethylene glycol and lactose but to a

lesser extent. This scheme was further extended to yield

fluorescent nano-sensors. Alizarin and octylboronic acid were

incorporated into a PVC polymer. Fluorescence intensity is

reduced in the presence of glucose which can be monitored in

this fashion at neutral pH over the physiological range. The

sensor also was tested in vivo with reasonable results.123

Another strategy is based on saccharide-induced conforma-

tional changes in copolymers containing both a boronic acid

and a fluorescent unit.124 Binding of a saccharide alters the

ionization state of the boronic acid moieties. The saccharide

present in the polymer directly affects the charge distribution in

the polymer chain. The acrylamide polymers contain pyrene or

naphthalene units, and conformational changes can be detected

fluorometrically after UV photoexcitation. Various polymer com-

positions were tested, but the basic principle is either an increase or

a decrease of the ratio between the intensity of excimer emission

(430–500 nm) and monomer emission (360–420 nm) on addition

of saccharide caused by contraction or expansion of the polymer.

The material displays comparable response towards fructose and

glucose. Saccharides like galactose, D-gluconic acid and D-glucaric

acid give smaller signal changes. This unusual sensitivity towards

glucose compared to other saccharide probes presumably is due to

the formation of bis-boronate complexes by neighboring boronate

groups in the polymer. The detection limits for glucose and

fructose are 10 mM. At present, these systems only work under

alkaline conditions. It also shall be kept in mind that the

fluorescence of pyrenes is strongly quenched by oxygen.125Fig. 14 Structure of a Ru complex capable of recognizing glucose via

a boronic acid.

Fig. 15 A polyaniline with a near-infrared optical response to

saccharides. The polyaniline undergoes large changes in the near-

infrared absorption spectrum between 600 and 800 nm on addition of

saccharides. From ref. 120. Copyright Wiley-VCH Verlag GmbH &

Co. KGaA. Reproduced with permission.

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An example of a colorimetric saccharide probe based on azo

dyes was reported by Lakowicz and DiCesare.126 The pKa of

the monoboronic acid used (8.0) decreases to 6.5 on binding

sugars. This is accompanied by a color change from orange to

purple. The origin of this effect is the change in the electronic

properties of the boronic acid, i.e. transformation of an

sp2-hybridized boron atom in triangular conformation to an

anionic and sp3-hybridized boron atom in tetrahedral geometry.

Ratiometric measurements were performed at 480 and 520 nm.

The probe shows higher selectivity for fructose over glucose,

and other saccharides are likely to interfere.

SPR-based sensors were described127 in which a thin gold

film was modified with a membrane containing a boronic acid.

The first sensor chip consisted of 3-aminophenylboronic acid

(3-APB) coupled to 11-mercaptoundecanoic acid to form a

self-assembled monolayer on a gold film. In the second, 3-APB

was immobilized on the sensor chip by electrochemical poly-

merization. In the third, the chip was coated with a vinyl

polymer by radical polymerization of vinylphenyl boronic

acid. All chips were tested in a flow system setup by measurement

of the shift of the resonance angle on addition of saccharides.

The sensor chip modified with poly(vinyl boronic acid)

exhibited a selective and strong response to saccharides and

was strong with fructose but weaker with glucose.

Singaram et al.128 have designed a two-component system

comprising an anionic fluorescent dye and a cationic viologen

quencher containing a bis-boronic acid functionality (Fig. 16).

The electrostatic attraction between anionic fluorophore and

cationic quencher leads to the formation of a non-fluorescent

ion pair. The electrostatic interaction is reduced when a

saccharide binds to the viologen to form a negatively charged

boronate ester. Consequently, fluorescence intensity is

increased. Eleven anionic dyes were tested, with hydroxypyrene-

trisulfonate (HPTS), perylenetetracarboxylate and tetrakis(4-

sulfonatophenyl)porphine showing the most interesting results.

HPTS exhibits the largest enhancement of fluorescence in the

presence of glucose. The response was almost linear for all dyes

in the blood glucose concentration range. The sensitivity towards

glucose can be adjusted by varying the quencher-to-dye-ratio.

This system works at physiological pH but is highly sensitive to

pH in that HPTS is a well-known pH probe.129 Cross-sensitivity

towards other saccharides such as fructose or galactose was not

investigated.

The HPTS/viologen method was further improved and

investigated with respect to quencher and selectivity.130 Viologens

with an ortho-boronate group bind strongest to glucose,

followed by fructose and galactose. However, fructose causes

higher fluorescence recovery despite a lower binding constant.

In subsequent work, a modular sensing ensemble was

developed that is composed of six different viologen quenchers.131

The increase in the luminescence intensity of HPTS on addition of

a distinct saccharide in the presence of three quenchers enabled

the differentiation between twelve saccharides including glucose,

fructose, mannose and galactose in 2 mM concentrations. All

components were covalently immobilized on a hydrogel, with the

indicator dye and the quencher in spatial proximity as shown in

Fig. 16. The sensor film is capable of sensing glucose between 2.5

and 20 mM concentrations. Interestingly, the system with

immobilized components is more selective for glucose than for

fructose.132 Further modifications include the use of fiber

optics,133 the variation of anionic fluorophores,134 the optimiza-

tion of the hydrogel,135 and the incorporation of the sensor

material into microtiter plates136 so as to enable high-throughput

screening for glucose. A relatively similar fiber optic based setup

was patented by Glumetrics Inc.137

Another promising sensing scheme138 is making use of two

fluorescent dyes along with an inert reference dye to result in a

multiple fluorescent reporter assay. The two fluorophores,

HPTS and tetrakis(4-sulfophenyl)porphine can be photo-

excited at the same wavelength (414 nm) but have well

separated emissions that peak at 510 nm and 644 nm, respectively.

The reference dye (sulforhodamine-B) can be excited at 565 nm.

Both reporter fluorophores are quenched by a viologen

quencher, whereas the luminescence of the reference dye

remains unaffected. Fluorescence is reconstituted on addition

of a saccharide. An almost linear increase in fluorescence is

observed with glucose in concentrations between 2.5 mM and

20 mM, with no saturation of the signal up to 100 mM.

Cross-sensitivity towards other saccharides was not investigated,

unfortunately.

CdSe quantum dots (QDs) were utilized139 as luminescent

reporters in combination with a viologen quencher with

boronic acid functionality as the glucose receptor. Luminescence

of the QDs is quenched by the viologen before addition of

glucose. The red emission is recovered on addition of glucose.

Two kinds of substituted QDs were investigated. The viologen

quenches both the carboxy- and amino-substituted quantum

dots in aqueous solution, but the carboxy-substituted QDs

more strongly. Ionic interactions between the anionic

(carboxy-substituted) QDs and the cationic quencher are

presumed to be responsible for these results. Interferents such

as fructose or galactose have not been investigated. In a

related sensing scheme,140 CdS QDs were embedded in a

polyacrylamide hydrogel with phenylboronic acid (PhBA)

groups that undergo shrinking upon addition of glucose. This

leads to quenching of the emission of the QDs. Glucose can be

determined in the 1–25 mM range, but possibly interfering

sugars have not been tested. Slightly modified QDs also were

shown to be viable cell-permeable glucose probes.141 Addition

of glucose quenches the luminescence of the QDs, and this is

Fig. 16 Boronate-based optical detection of glucose across the visible

spectrum using an anionic fluorescent dye (HPTS) and a viologen

quencher in a 2-component saccharide sensing system. From ref. 128.

Reproduced by permission of the Royal Society of Chemistry.

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accompanied by a red shift of emission. This is most likely due

to aggregation of the QDs. The glucose concentration in

mouse melanoma cells was monitored using such nanosensors

via scanning confocal fluorescence microscopy. However, no

feasibly competing diols were tested. The same group142

developed silver nanoparticles (Ag-NPs) that are covered by

a hydrogel containing PhBA groups. The emission of the

Ag-NPs is quenched in the presence of glucose. The authors

showed that their nanogel also can be used as a glucose

dependent insulin delivery system. A combination of

LaF3:Ce3+/Tb3+ nanocrystals modified with glucose and

3-aminophenyl-boronic acid (3-APhBA) modified with

rhodamine B isothiocyanate displays a strong FRET between

the nanocrystal (exc./em. 254/543 mn) and the 3-APh-rhodamine

conjugate143 (exc./em. 558/584 nm). In the presence of glucose,

the saccharide competitively binds to the system, and this

results in a quenched FRET and increased nanocrystal

luminescence. Glucose can be determined in concentrations

between 0.5 and 25 nM.

Anthracenes substituted with one or more phenylboronic

acid groups via a linker of different length represent another

group of glucose probes. Two anthracene derivatives with

boronic acid groups covalently attached via alkyl amino

linkers were evaluated144 with respect to fluorescence lifetime-

based sensing of glucose. The emission of anthracene is

quenched by the substituent due to PET. Both anthracene

derivatives displayed increased intensities and lifetime in the

presence of glucose. A change in fluorescence lifetime from 9.8

to 12.4 and from 5.7 to 11.8 ns was observed in the presence of

glucose for the anthracene substituted with one and two

boronic acid groups, respectively. No other saccharides were

tested. Measurements had to be carried out in 33%methanolic

buffer of pH 7.7 because of the poor solubility of the probes.

Other disadvantages of this probe include its photolability,

quenching by oxygen, and the need for UV excitation.

In other approaches145,146 two anthracene-mono-phenyl-

boronic acids were linked together. This probe, while requiring

UV excitation, displays a more than 40-fold selectivity for

glucose over both fructose and galactose. In the presence of

glucose, an up to 7-fold increase of the purple fluorescence is

observed. NMR studies revealed a furanoidic binding

conformation of glucose. Low millimolar concentrations of

glucose induce a significant increase in fluorescence. The probe

was immobilized on cuprophane to enable continuous sensing.147

Pyrene and its derivatives are well established and investigated

fluorophores. They have a fairly high quantum yield (B0.6)

and lifetimes of up to 0.4 ms (in the absence of oxygen).

Pyrenes are capable of forming excited state dimers, referred

to as excimers, if two pyrene moieties are located in close

proximity. This effect is strongly dependent on the distance

between the pyrenes, and smart probes have been designed

where binding of a saccharide governs this distance. The

concentration of the saccharide in this case affects the ratio

of the intensities of monomeric and excimeric emission.

Gibson and Appleton148 have evaluated a number of pyrenes

containing boronic acids and undergoing photoinduced

electron transfer (PET). The pyrene units were linked with

spacers of different length and containing amino groups.

On formation of a boronic acid glucose ester, the enhanced

Lewis acid–Lewis base interaction between the boron and

nitrogen atoms of the spacer suppresses the PET process and

increases fluorescence intensity (exc./em. 343/378 nm). The

effect of spacer length between two boronic acids on the same

molecule on the selectivity and sensitivity of saccharide binding

was studied in detail. A 6- or 7-carbon spacer seems to confer

highest specificity towards glucose over fructose, galactose,

sorbose and ethylene glycol. Additionally, a vinyl derivative of

the probe was synthesized and crosslinked with poly(methyl

acrylate) in the presence of glucose. The resulting imprinted

polymer exhibits high selectivity for glucose, without any

interference from other saccharides, but fluorescence response

is weak. Measurements have to be carried out in 33%methanolic

buffer due to the poor solubility of the probe.

In similar studies,149,150 clusters of probes containing either

the same fluorophore (pyrene) or two different luminophores

(such as pyrene and phenanthrene) were linked with methylene

chains of different lengths (3 to 8). Results show that a spacer

with a length of 6 carbon atoms is best in terms of selectivity.

However, fructose induces the highest fluorescence enhancement,

even though the stability constant is highest for glucose. All

systems require UV excitation, and the fluorescence of pyrenes

is quenched by oxygen.

In another scheme based on pyrene excimer emission,151 the

pyrene units were not linked to each other (see Fig. 17).

Rather, ditopic recognition of glucose was exploited and

resulted in high selectivity over other saccharides. The formation

of a 1 : 2 complex between glucose and the fluorescent pyrene

boronic acid brings the two pyrenyl moieties in close proximity

and thus enhances excimer emission. The method requires the

addition of a polycation which acts as a preconcentrator for

the fluorescent boronic acids due to electrostatic interactions.

Thus, the formation of the 1 : 2 complex is facilitated.

In a carbohydrate receptor based on a cyclic tetrapeptide

with two appended PhBA moieties, a stable 1 : 1 complex is

formed with glucose in 50%methanol, however, at pH 11.7.152

On excitation at 285 nm(!), a fluorescence emission with a peak

at 480 nm is observed. Fluorescence intensity decreases

significantly on addition of glucose. The affinity of the probe

for D-glucose is two times higher than for L-glucose. Saccharides,

such as D-galactose and D-mannose, form less stable complexes

with the cyclic tetrapeptide. This method has drawbacks in that it

requires an alkaline pH, the addition of methanol, and UV

excitation which is unfavorable in many respects.

Fig. 17 Selective glucose sensing utilizing complexation with

fluorescent boronic acid on polycation. Before addition of glucose:

monomer emission. After addition of glucose: 1 : 2 complex formation

between glucose and fluorescent boronic acid in the presence of

polycation poly(diallyl dimethylammonium) chloride which results in

excimer emission. From ref. 151.

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Mohr et al.153 report on longwave probes based on a

hemicyanine chromophore, with large Stokes’ shifts, absorption

maxima at B460 nm, and emission maxima at B600 nm at

neutral pH. Three probes were evaluated, with the boronic acid

group placed in the ortho, meta or para position. Addition of a

saccharide leads to an enhancement of fluorescence intensity in

the case of all three probes, specifically by 86% in the case of the

ortho derivative in the presence of fructose, but 27% and 33%

only for themeta and para derivatives. The probes respond in the

order fructose > galactose > glucose > ethylene glycol >

glycerol. The analytical range is from 5 to 500 mM of glucose.

Medtronic Minimed has patented154 a glucose sensor

comprising a monotopic fluorescent acridine-based boronic

acid. Fluorescence at 500 nm increases with the concentration

of glucose due to photoinduced electron transfer. Another

patent155 covers a chloro-oxazine boronate for glucose

sensing. Changes in fluorescence lifetime are monitored.

A sensor film was introduced156 that is based on the reaction

of boronic acid in a plasticized PVC membrane. Saccharides

react with 3-(dansylamino)phenyl boronic acid contained in

the membrane to produce a stable boronate anion and

simultaneously to liberate a proton. The hydrogen ions in

the membrane protonate the dansyl moiety, and its green

fluorescence is reduced as a result. The change in fluorescence

intensity of the sensing films resulting from the increase in

hydrogen ion concentration is thus directly related to the

ambient saccharide concentration. D-glucose, D-fructose,

D-galactose, and D-sorbitol were determined in the concentration

range from 0.1 to 100 mM at physiological pH. The selectivity

for glucose was enhanced in comparison to a homogeneous

assay. Such sensors suffer, however, from a substantial cross-

sensitivity to pH.

Holographic glucose sensors157–159 represent a compara-

tively new development. In such sensors, a holographic grating

is generated in a polymeric matrix, a hydrogel generally.

Under white light illumination, the gratings reflect a narrow

band of wavelengths and create the monochromatic image of

the original mirror used in its construction. Binding of a

saccharide to the boronic acid groups in the hydrogel matrix

induces swelling or shrinkage of the gel. The spacing of the

refraction fringes is altered as a result, this generating a change

in the wavelength and color of the reflection hologram. The

change in hydrogel volume is reversible and independent of

pH in the physiological range. The diffraction wavelength is

determined by Bragg’s law. 2-Acrylamido-phenylboronic acid

(2-APhBA) is reported as a glucose recognition moiety. The

holographic sensors are hardly interfered by lactate. Swelling

of the hydrogel is induced on binding glucose and causes a

red-shift in the diffraction wavelength. The sensors respond

linearly over the physiological range of glucose concentrations

at neutral pH. If 3-APhBA is used as a reporter,158,160 the pH

dependence is more expressed than in the case of 2-APhBA.

4-Vinyl-PhBA was also reported161 to act as a receptor for

glucose but with the drawback of working best at pH 9. All

sensors respond slowly (>10 min).

The advantage of the holographic method over other optical

techniques is the long-term stability of the sensor and the ease

with which the wavelength may be tuned to suit the application.

Fluorophores often have spectral characteristics that compromise

the applicability to complex samples where light attenuation,

background fluorescence and photobleaching can bias the

analytical precision. The direct binding approach also has

the advantage that the glucose is not consumed, a key issue

if small changes in glucose levels are to be determined.

Responsive photonic crystals (RPCs) represent an extremely

interesting new class of materials for purposes of sensing.

Polymerized crystalline colloidal array photonic crystal

sensing materials are particularly attractive. They consist of

an embedded crystalline colloidal array (CCA) surrounded by

a polymer hydrogel network which contains a molecular

recognition element. The embedded CCA of polystyrene

colloidal particles efficiently diffracts light of a wavelength

determined by the array lattice constant. The structure and

diffraction spectra of a typical CCA are shown in Fig. 18.

Holtz and Asher163 probably were the first to prepare a

glucose RPC by attaching the enzyme GOx to a hydrogel

network. As the products of the enzymatic process swell the

hydrogel film, the diffraction red-shifts due to increased

osmotic pressure. This was tentatively ascribed to the formation

of a reduced flavin (FADH2). In the absence of oxidants, the

gel can detect glucose in concentrations as low as 10�12 M.

However, in an atmosphere containing oxygen, reoxidation of

flavin shrinks the gel and weakens the response of the RPC.

CCAs also were constructed164,165 within a polyacrylamide

hydrogel and with pendent groups of phenylboronic acid

(PhBA). The increase in hydrogel crosslinking and the shrinking

of volume, both induced by saccharide binding, cause a

modification of the diffraction of the photonic crystal.

The concomitant color change can be visually monitored.

Interestingly, RPCs modified with boronic acid and

poly(ethylene glycol) respond to glucose with a blue-shift of

the diffraction, since the supramolecular bisbidentate

glucose–boronic acid complex stabilized by PEG can cross-link

the polymer and shrink the hydrogel matrix. These materials

Fig. 18 Photonic crystal and spectra showing the large red shifts of

the diffracted wavelength that occur as a result of an increase in the

volume of the hydrogel which is induced by the interaction of the

glucose with the molecular recognition element. Such color changes

are easily detectable even by unskilled personnel. From ref. 162 with

kind permission from the American Chemical Society.

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respond to glucose at physiological concentrations, ionic

strengths, and pH values, and are selective for glucose over

galactose, mannose, and fructose. This approach was further

developed166 so as to increase the analytical range and to

improve the reproducibility of fabrication. Braun et al.167 have

prepared similar RPCs by exploiting the complex formation

between 1,2-cis-diol glucose and groups of PhBA attached to

the hydrogel.

Several approaches based on the use of swelling polymers

have been discussed in previous sections of this review. In yet

another embodiment,168 a polyacrylamide hydrogel with

PhBA groups is placed at the tip of an optical fiber. Shrinking

of the polymer in the presence of glucose is detected by an

interferometric technique. The sensor proved to be temperature-

dependent. The linear gel swelling response in aqueous solutions

was determined to be 1760 nm per mM of glucose. Other

carbohydrates displayed a response of about 10% of that of

glucose. The method was successfully applied to serum samples

to which EDTA had to be added. The sensor is even more

selective to glucose if dimethylaminopropyl acrylamide is

incorporated into the hydrogel.169

7. Sensing glucose via concanavalin A

Concanavalin A (ConA) is a plant lectin protein that can be

extracted from Jack beans. The ConA tetramer consists of two

dimers and has four binding sites for glucose. Its specific

function involves the agglutination of biologically relevant

complexes such as glycoproteins, starches and erythrocytes.

ConA is traditionally used in competitive assays where glucose

and another carbohydrate (such as dextran, mannoside or a

glycated protein) compete for the lectin binding sites. Protein

and the competitor can be labeled with appropriate fluorescent

dyes. ConA based sensors suffer from poor stability because

unbound lectin tends to irreversibly aggregate over a period of

some hours,170 and that fluorescent sensing schemes are

preferred, often based on the use of rather longwave absorbing

and emitting labels.

7.1. Conventional fluorescent sensing schemes using a single

label

Schultz and co-workers171,172 probably were the first to use

ConA in an optical glucose biosensor. ConA was immobilized

inside a hollow dialysis fiber connected to a fluorescence

detection system by a single optical fiber. FITC-labeled dextran

acts as the competing ligand. Small molecules like glucose can

pass in and out of the fiber, but the (labeled) dextran cannot.

Glucose displaces the competitor (dextran) from the binding

site, thus increasing the concentration of free dextran and,

consequently, the intensity of fluorescence within the numerical

aperture of the fiber. Glucose can be determined in the blood

physiological range. Further improvements were made173

regarding response time and optical readout in terms of

improvement of the fiber optic. Response times are from

5 to 7 min. The biotoxicity of ConA has been reviewed.174

The same group has developed175 a slightly altered sensing

scheme for transdermal glucose monitoring. Labeled ConA

and dyed macroporous Sephadex beads are confined inside

a sealed, small segment of a hollow fiber dialysis membrane

(see Fig. 17). Immobilized pendant glucose moieties inside the

intensely colored Sephadex beads compete with glucose for

binding of ConA. In the absence of glucose, the bulk of labeled

lectin resides inside the red beads (whose color was selected to

block the green emission of the label Alexa-488), and the

fluorescence of labeled ConA is screened off. On exposure of

the hollow fiber sensor to glucose, the saccharide will diffuse

through the membrane into the sensor chamber and competitively

displace labeled ConA from the glucose units of the red beads.

As a result, ConA is fully exposed to the excitation light, and a

strong increase in fluorescence emission at 520 nm is observed

(see Fig. 19). The sensor features a detection range from

0.15 to 100 mM of glucose, a strong dynamic signal change

between 0.2 and 30 mM of glucose, and a response time of 4 min.

Ballerstadt et al.176 have studied the long-term in vitro

performance of a similar setup. Near-IR emitting Alexa-647

was used to label ConA, and the screening dye in the Sephadex

beads was exchanged against Alkali Blue 6B to match excitation

and emission wavelengths. Sensors were alternately exposed to

glucose in concentrations of 2.5 and 20 mM, and the optical

output was monitored over a period of 4 months. Further

experiments related to time-dependent membrane leakage, the

solubility of ConA, the temperature-dependent activity of

ConA, and photo-bleaching. The signal loss is 25% within

4 weeks, not due to denaturation of ConA but rather due to

leakage of ConA within the sensor. An extrapolation of the

experimental data indicated that a leak-proof sensor would be

remarkably stable, with a fluorescence decrease of only 15%

over a 1 year period. Another ratiometric FRET measurement

is based on even more longwave near infrared emission.177 The

sensor consists of a small hollow fiber implanted in dermal

skin tissue, containing Cy7 labeled agarose-immobilized ConA

Fig. 19 Schematic of a competitive glucose biosensor using ConA.

In the absence of glucose, fluorochrome labeled Con A is bound to

fixed glucose residues inside the porous beads (left hand) in a hollow

fiber. After diffusion of glucose through the hollow fiber membrane,

Con A is displaced from the beads and diffuses out of them, and

hereby fluorochrome-labeled Con A becomes exposed to excitation

light resulting in a strong increase in fluorescence (right hand).

Reprinted with permission from ref. 175. Copyright 2000 American

Chemical Society.

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as acceptor/quencher and unbound Alexa-647-dextran as the

donor. In vivo tests revealed a delay of 5 to 15 min with respect

to the actual blood glucose concentrations. Precisense has filed

a patent application178 on a sensor that makes use of ConA

labeled with Alexa-594 and Crystal Violet-labeled amino-

dextran, both placed in a hollow fiber. Changes in fluorescence

lifetime serve as the analytical information.

7.2. Sensing schemes based on energy transfer (FRET) or

screening effects using two labels

Rather than measuring a single fluorescent event, a ratiometric

FRET based setup between TRITC-labeled ConA and FITC-

labeled dextran was presented by Schultz and Meadows.179

The emission of fluorescein-labeled dextran is restored on

addition of glucose due to competitive displacement. This

sensor responds to up to 83 mM concentrations of glucose

and with a response time of 10 min. However, free ConA

aggregated within a few hours. Chemical derivatization using

succinic anhydride can prevent aggregation of the receptor

protein in solution to a substantial extent.

In yet another variation,180 ConA was labeled with Cy5

which serves as the FRET donor. The acceptor consists of

insulin covalently linked to both Malachite Green and maltose

(MIMG) to provide a binding site for ConA. Binding of

Cy5ConA to MIMG results in a decreased fluorescence

intensity and decay time of Cy5. If MIMG is displaced

competitively from ConA by glucose, fluorescence emission

and lifetime increase. This sensor also displays a reduced rate

of aggregation and better reversibility, possibly because a

protein is used as a competitor instead of dextran. Similar

experiments181 were carried out with ConA labeled with

Ru(bpy)3. Its large Stokes’ shift and long decay times are

advantageous in simplifying instrumentation for phase-modulation

lifetime measurements. The sensors cover glucose concentration

ranges of up to 50 mM.

Russel et al.182 have extended the (TRITC-ConA) (FITC-

dextran) FRET method to hydrogel spheres. An acceptor and

a donor were covalently immobilized in a poly(ethylene glycol)

hydrogel. On addition of glucose, a decreased FRET and an

increased fluorescein emission are observed due to competitive

displacement of dextran. The response time is 10 min, and

response is linear up to 44 mM of glucose. This group also

holds a patent on the sensing scheme.183

Novartis has patented184 a TRITC-ConA and FITC-

dextran FRET sensor that is incorporated in a contact lens along

with an apparatus to irradiate the labels and detect FRET.

A quite similar method, also based on the (TRITC-ConA)

(FITC-dextran) FRET system, was patented by Torsana

Diabetes Diagnostics.185 Agarose microparticles that contain

the donor and acceptor were deposited on a micromachined

pad having a 10 mm square array of 400 microneedles of 1 mmdiameter. The particles can be injected subcutaneously, and

the rhodamine fluorescence can be read out with a fiber optic

fluorometer.

A near-infrared FRET sensing scheme in which ConA is

labeled with the protein allophycocyanin (APC; the donor)

and dextran labeled with Malachite Green (MG; the acceptor)

was developed by Birch and co-workers.186 MG screens the

emission of APC as long as dextran is bound to the lectin.

Glucose competitively displaces dextran-MG and leads to

restoration of the APC fluorescence at 663 nm which is

quantified by fluorescence lifetime measurements. Glucose

concentrations in the range 2.5–30 mM are detectable with

this sensor. Albumin and serum were reported to inhibit

FRET but can be excluded by applying membrane filters.

Chinnayelka and McShane187 have converted the (TRITC-

ConA) (FITC-dextran) FRET method to self-assembled

microcapsules (Fig. 20) that also can act as optical glucose

micro-sensors. The nanoscale planar glucose biosensor was

obtained by self-assembly of the ConA/dextran conjugate into

multilayer films. This is particularly advantageous because of

the physical localization and separation of sensing molecules

from the environment via entrapment of the biosensor

elements in a semi-permeable polymeric shell. Moreover, only

functional molecules are included in the sensors. A glucose-

specific enhancement of fluorescence emission of 27% was

observed exhibiting a linear increase over the 0–100 mM

range. A related study188 reports on the competitive binding

of glucose and a glycodendrimer to fluorescently tagged ConA

in porous microspheres made from poly(ethylene glycol). This

system is said to be stable for up to 2 weeks.

The changes in the reflectivity of a glucose-sensitive hydrogel

containing ConA are claimed in a patent filed by M-Biotech

Inc.189 Reflective material is arranged on a hydrogel filament

to move with displacement of the hydrogel filament and to

reflect light from the light source toward a photoreceptor.

The movement of the reflector is changing the intensity of light

reflected to the photoreceptor dependent on the glucose

concentration.

A TRITC-ConA and FITC-dextran FRET system was

incorporated into hydrogel pads that were deposited on the

hydrophobic surfaces of the wells of a microtiter plate.190 A

layer-by-layer self-assembly process was used to further coat

the hydrogel pads with polyelectrolyte multilayers with the

aim to create a permeation-controlled membrane of nano-

metre thickness. The calibration curve was linear up to 10 mM

glucose, and 95% of the maximum fluorescence was reached in

less than 8 min. The FRET was regenerated with buffer within

17 min and this reversibility may pave the way to a reusable

sensing platform.

Yet another sensor based on FRET utilizes quantum dots

(QDs). These are labeled with ConA and used as donors

because of their high variability in excitation wavelengths

and high quantum yield.191 TRITC-labeled b-cyclodextrin(which shows lower affinity to ConA than the linear dextran)

Fig. 20 Glucose sensing with self-assembled microcapsules, competitive

binding, and resonance energy transfer (RET). The efficiency of RET

(from FITC-dextran to TRITC-ConA). Changes on addition of glucose

because of the displacement of labeled dextran. With permission from

Springer Science + Business Media from ref. 187.

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competes with glucose for binding to the QD-labeled ConA.

The biosensor is reported to give a strong signal and to cover

the whole physiological range. The components were incorporated

into a hydrogel and photopolymerized at the end of an optical

fiber with a diameter of only 250 mm. This makes the setup

suitable for interstitial and minimal invasive in vivo sensing.

The ratio of the emissions at 525 nm and 570 nm correlated

linearly with the glucose concentration up to 28 mM. The

relatively long response times (3 to 8 min) were ascribed192 to

the reduced mobility of ConA in the polymer. The system was

applied to in vivo interstitial glucose determination in dogs.

A related sensing scheme was reported by Tang et al.193 but

demonstrated for solutions only. ConA was labeled with QDs,

and b-cyclodextrin with gold nanoparticles (Au-NPs). In the

absence of glucose, the emission of the QDs is screened off by

the red Au-NPs because they are close to the labeled ConA. In

the presence of glucose, the AuNPs-QDs are competitively

displaced by glucose, and this results in a recovery of luminescence

(Fig. 21). The increase in luminescence is proportional to the

glucose concentration in the range from 0.10 to 50 mM. The

limit of detection is as low as 50 nM, and the selectivity for

glucose over other sugars and most biological species present

in serum is very high.

7.3 Other optical sensing schemes

ConA also was used as a receptor in surface plasmon

resonance (SPR) based sensing schemes.194 Silver nanoparticles

(Ag-NPs) were bound to amine-labeled dextran 3000 and

coupled to ConA to induce the formation of particles.

Competitive binding of glucose to ConA releases the particle/

dextran adduct which is accompanied by a change in the

plasmonic absorbance and wavelength that is correlated to

the glucose concentration (Table 7). The same group195

reported on a similar setup utilizing gold colloids coated with

dextran but otherwise based on the same principle (see Fig. 22).

Addition of ConA again causes aggregation. The analytical

range (from mM to mM) can be adjusted by either the

thickness of the Au-NPs or the concentration of ConA.

Glucose causes the dissociation of the aggregates which is

monitored by a reduction in plasmon absorbance. Its

tunability and analytical wavelengths of >600 nm are said

to make this scheme suitable for implementation in contact

lenses for non-invasive tear glucose determination.

The scheme was optimized196 in terms of particle stability,

pH effects, dynamic range and the analytical wavelength in

order to adapt it to clinical requirements. The Au-NP based

scheme was later extended by measuring the light scattering

properties of the Au-NPs. Glucose in concentrations up to 60 mM

was sensed by measuring the ratio of intensity of scattered

light at 560 and 680 nm, respectively, using a white light LED

as a light source. The ratiometric approach makes the system

independent of the total concentration of the Au-colloid ConA

aggregate and instrumental drifts.197

In addition to their work on fiber optic sensors,176,177

Ballerstadt et al.198 have developed a glucose sensor where

the turbidity of the sensing element (that can be monitored by

optical coherence tomography) yields the analytical information.

The sensor consists of a mm thick glucose-permeable

membrane (see Fig. 23) containing a suspension of macroporous

Sephadex particles and ConA. A solution containing such

particles is turbid and strongly scatters incident light if glucose

is absent. However, glucose renders this solution almost

transparent, resulting in low scattering. The operational

stability under in vitro conditions is as long as 160 days,

with good overall response over the physiological glucose

concentration range (2.5–20 mM).

A more sophisticated method was patented by the same

group.199 It is based on reversible changes in fluorescence due

to changes in the turbidity of a ConA/Sephadex system with

varying glucose concentrations. A ConA/Sephadex suspension

is sandwiched inside a rectangular semipermeable dialysis

capsule of membranes of regenerated cellulose (‘‘rayon’’).

A polysulfone film with an encapsulated fluorescent dye is

mounted to the capsule. Turbidity decreases on addition of

glucose to result in increased fluorescence.

The luminescence of carbon nanotubes depends on their

state of aggregation. This finding forms the basis for a

solution-phase affinity sensor based on the use of dextran-modified

single wall carbon nanotubes (SWCNTs).200 These were

aggregated by ConA, and this results in quenched photo-

luminescence (exc./em. 633/> 900 nm). Addition of glucose

restores the initial (and very longwave) luminescence. This is

schematically shown in Fig. 24. Response times of 3–28 min in

a range from 3.8 to 11 mM of glucose at pH 7 were achieved.

The subject has been reviewed.201

8. Sensing glucose via glucose-binding proteins

other than con A

8.1. Sensing based on glucose-binding apoenzymes

The use of glucose-metabolizing enzymes for purposes of

sensing glucose was reviewed in Section 5. However, such

enzymes—if deprived of their coenzyme—may also be

employed as glucose-binding (but not metabolizing) enzymes

very much like concanavalin A. Removal of the coenzyme is

fairly easily accomplished and yields the respective apo-enzymes.

Two main schemes are known for apo-enzyme based affinity

assays. The first is based on changes in the instrinsic fluorescence

of the enzyme, the second on the use of a fluorescent label, often

a polarity-sensitive dye. Both direct and competitive binding

assays have been reported.

Fig. 21 Nanobiosensor for glucose. Glucose induces AuNP-displacement

and restoredQDfluorescence. From ref. 193. CopyrightWiley-VCHVerlag

GmbH & Co. KGaA. Reproduced with permission.

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Table 7 Glucose affinity assays via concanavalin A (ConA). (FLU: fluorescence; FRET: fluorescence resonance energy transfer); FITC:fluorescein isothiocyanate; GLU: glucose; MG: malachite green; OCT: optical coherence tomography; QD: quantum dot; REFL: reflectometry;RLS: resonance light scattering; SPR: surface plasmon resonance; SWCNTs: single wall carbon nanotubes; TRITC: tetramethylrhodamineisothiocyanate)

Method and comments AR/mM Ref.

FLU; ConA immobilized on the inside of hollow dialysis fiber; FITC-labeled dextran as competitive ligand; only smallmolecules can pass in and out; increase in concentration of free dextran on addition of GLU; increase in fluorescence

Physiological range 171

FLU; Sepharose-immobilized ConA as a binding agent and FITC-dextran as a competing ligand which fluoresces onillumination; dextran competes with GLU for binding sites on the immobilized ConA; patent

— 172

FLU; ConA immobilized on the inside of hollow dialysis fiber, FITC-labeled dextran as competitive ligand; only smallmolecules can pass in and out; increase in concentration of free dextran on addition of GLU; increase in fluorescence;fiber optic setup; response time: 5–7 min

Physiological range 173

FLU; Sephadex beads with GLU groups labeled with Safranin O and Pararosanilin (block exc. and em. of Alexa-488);Alexa-488 labeled ConA binds to these GLU groups; fluorescence of Alexa is quenched; fluorescence at 522 nm (exc.490 nm) is restored on addition of GLU; response time: 4–5 min; whole system in hollow dialysis fiber; fiber opticsetup

0.2–30 175

FLU; Sephadex beads with GLU groups labeled with dye (block exc. and em. of Alexa-647); Alexa-647 labeled ConAbinds to these GLU groups; fluorescence of Alexa is quenched; fluorescence (670 nm) is restored on addition of GLU;response time: 15–30 min; whole system in hollow dialysis fiber; fiber optic setup; high operational stability; works innear-IR

2.5–20 176

Patent: FLU; ConA labeled with Alexa-594 and aminodextran labeled with a Crystal Violet succinimidyl ester areplaced in a hollow fiber of regenerated cellulose; fluorescence lifetime measured with automated apparatus

— 178

FRET between TRITC-ConA and FITC-dextran; restoration of FITC emission on addition of GLU; fiber opticsetup; response time 10 min; two excitation sources for internal calibration (480 nm, 550 nm); hollow fiber setup

0–83 179

FRET; ConA labeled with Cy5, insulin labeled with MG and maltose; FRET; restoration of Cy5 fluorescence andincrease in lifetime on addition of GLU; less aggregation and better reversibility of the assay by using protein insteadof dextran; NIR emission; decay times can be measured through skin using long wavelength excitation and emission,suggesting the possibility of an implanted GLU sensor

0–60 180

FRET; ConA labeled with Ru(bpy)3, insulin labeled with MG and maltose; FRET; restoration of Ru luminescenceand increase in lifetime on addition of GLU; less aggregation and better reversibility of the assay by using proteininstead of dextran; NIR emission but short wavelength excitation; long lifetime

0–50 181

FRET; TRITC labeled ConA and FITC labeled dextran incorporated in hydrogel; FRET; decrease of FRET onaddition of GLU; response time: 10 min

0–44 182

Patent: FRET; TRITC labeled ConA and FITC labeled dextran incorporated in hydrogel; decrease of FRET onaddition of GLU

— 183

Patent: FRET between TRITC-ConA and FITC-dextran; incorporated in contact lens; apparatus to irradiate thesensor dyes and detect FRET

— 184

Patent; FRET; agarose microparticles containing TRITC-ConA and dextran-FITC are loaded onto a micromachinedpad having a 10 mm square array of 400 microneedles of 1 mm diameter; injected subcutaneously; fiber opticfluorometer used to measure the rhodamine fluorescence

— 185

FRET; ConA labeled with allophycocyanine, dextran labeled with MG; FRET, MG shields allophycocyanine;Restoration of allophycocyanine fluorescence (663 nm) on addition of GLU; lifetime measurement; inhibition ofFRET by albumin and serum; works in near-IR

2.5–30 186

FRET; layer-by-layer assembly of films in microcapsules containing TRITC-ConA and dextran-FITC in thin polymerfilms

0–100 187

Patent; REFL; hydrogel containing ConA and GLU groups; swelling of the gel on addition of GLU; change inreflected light

— 189

FLU; FRET from Alexa-647 (donor) labeled dextran to Cy7 (acceptor) labeled on agarose together with ConA;decrease of FRET on addition of GLU; fiber optical setup in vivo; response time: 5–15 min; works in near-IR

2.5–25 177

FRET; TRITC-ConA and FITC dextran incorporated in hydrogel pads contained in wells of a microtiter plate;layer-by-layer method; response time: 8 min; enables continuous sensing

0–10 190

FRET; TRITC-b-cyclodextrin and QD/ConA conjugate incorporated in hydrogel and photopolymerized at the tip ofan optical fiber; response time: 3–8 min; fiber optic setup; interstitial GLU sensor; enables continuous sensing; appliedto in vivo glucose determination

0–28 191192

FRET; ConA coupled to CdTe QDs, Au-NPs modified with cyclodextrin; FRET on aggregation; restoration ofquenched QD luminescence on addition of GLU; direct determination in serum; no interference by other saccharides(exc./em. 320 nm/530 nm)

0–0.05 193

SPR; silver-NPs bound to dextran 3000; aggregation with ConA; plasmon absorbance reduction on addition of GLU 0–0.05 194SPR; dextran coated Au-NPs aggregated by ConA; dissociation of aggregates on addition of GLU; reduction ofplasmon absorbance; thickness of NPs and ConA concentration modify the analytical range

— 195

SPR; dextran coated Au-NPs aggregated by ConA; dissociation of aggregates on addition of GLU; reduction ofplasmon absorbance; thickness of NPs and ConA concentration modify the analytical range

0–50 196

RLS; dextran coated Au-NPs aggregated by ConA; dissociation of aggregates on addition of GLU; light scatteringmeasured at 560 and 680 nm; self-referenced

1–60 197

OCT; macroporous Sephadex particles and ConA incorporated in GLU-permeable housing; light scattering decreaseson addition of GLU to render suspension transparent; response time: 23 min; enables continuous sensing

2.5–20 198

Patent; FLU; reversible change in fluorescence due to turbidity changes in a ConA/Sephadex system at various GLUconcentrations; ConA/Sephadex suspension sandwiched inside a rectangular semipermeable dialysis capsule ofmembranes of regenerated cellulose; turbidity decreased on addition of GLU, increasing fluorescence; enablescontinuous sensing

— 199

FLU; dextran-modified single wall carbon nanotubes (SWCNTs) aggregated by ConA, resulting in quenched SWCNTphotoluminescence; restoration of initial PL on addition of GLU; response time: 3–28 min (exc./em. 633/> 900 nm);PBS pH 7

3.8–11 200

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The fluorescence of apo-GOx undergoes an up to 18%

decrease in its intrinsic tryptophan fluorescence on binding

to glucose (Table 8).202 However, this effect is not very useful

for medical purposes due to the need for UV excitation and an

emission that also occurs in the UV. In order to improve this

method, apo-GOx was non-covalently labeled with the polarity-

sensitive dye 8-anilino-1-naphthalenesulfonate (ANS). Its

emission increases on conjugation, but decreases by 25% in

the presence of glucose. In parallel, the mean lifetime of the

label decreases by about 40% due to conformational changes of

apo-GOx. The system is sensitive to glucose in the range from

10 to 20 mM. The results suggest that apo-glucose oxidase can

be used as a reversible sensor for glucose which—unlike in the

case of functional enzymes—is not consumed. The method was

patented.203 A similar approach was described204 for glucose

dehydrogenase in the complete absence of the cosubstrate

NAD+ (whose presence is mandatory to catalyze the redox

process). The addition of glucose results in a decrease of

fluorescence polarization and emission intensity of 25%.

A FRET-based competitive sensing scheme was developed205

that is exploiting the resonance energy transfer from dextran

labeled with a fluorescein (FITC) to apo-GOx labeled with a

rhodamine (TRICT). Its efficiency decreases on addition of

glucose to the aggregate formed between apo-GOx and dextran

due to competitive binding of glucose to apo-GOx. This is

schematically shown in Fig. 25. The biosensor is fully reversible

and has a dynamic range from 0 to 90 mM of glucose.

The solution assay was then converted into an implantable

minimal invasive sensor.187 The layer-by-layer procedure that

was established previously for a ConA-based sensor was

Fig. 22 Tunable plasmonic glucose sensing based on the dissociation

of ConA-aggregated dextran-coated gold colloids. Top: before

addition of glucose ConA is aggregated with dextran-coated nano-

gold colloids. Bottom: after addition of glucose competitive binding to

ConA leads to dissociation of the dextran-coated colloids. Reprinted

from ref. 195 with permission from Elsevier.

Fig. 23 Schematic of an affinity-based turbidity sensor for glucose monitoring by optical coherence tomography. Left: before addition of glucose;

right: after addition. Reprinted with permission from ref. 198. Copyright 2007 American Chemical Society.

Fig. 24 Reversible carbon nanotube aggregation in a glucose affinity

sensor. Single wall carbon nanotubes (SWCNT) luminophores are

initially suspended in aqueous solution with a phenoxy-derivatized

dextran. Addition of ConA induces aggregation of the dextran–

nanotube complexes and a decrease in the luminescence of the

SWCNTs. After addition of glucose, the aggregates dissolve and PL

recovers owing to competitive binding between glucose and dextran

for ConA binding sites. From ref. 200. Copyright Wiley-VCH Verlag

GmbH & Co. KGaA. Reproduced with permission.

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applied. TRITC-labeled apo-GOx and FITC-labeled dextran

were encapsulated in semipermeable microcapsules built up

from polystyrene and polyallylamine.206 These capsules

displayed 5 times better specificity for glucose over other

sugars up to glucose concentrations of 40 mM. This makes

the sensor suitable for monitoring physiological levels. The

optical read-out was improved by using other longwave

labels.207,208

Single-walled carbon nanotubes (SWCNTs) display solvato-

chromic luminescence. If incorporated into a PVA hydrogel

containing immobilized apo-GOx, the NIR emission of the

carbon tubes is shifted on addition of glucose due to alteration

of the swelling of the polymer. The hydrogel sensor was

applied to image glucose in a mouse tissue model.209

8.2. Sensing based on glucose-binding proteins

The periplasm of Gram-negative bacteria such as Escherichia

coli comprises a family of proteins with highly specific binding

of sugars and other ligands like amino acids, ions and vitamins

or dipeptides. These binding proteins are primary receptors in

transport and have all a very similar structure in common. The

glucose binding proteins (GBPs) sometimes also exhibit

affinity towards galactose, albeit to a much lesser extent,210 and

therefore sometimes are referred to as the glucose/galactose-

binding proteins (GGBPs).

GBP-based sensing schemes—unlike enzyme based glucose

sensors—are not limited by the need for the presence of other

enzyme substrates like oxygen or the formation of toxic

reaction products such as hydrogen peroxide. Therefore, they

are often referred to as reagentless sensors. All sensing

schemes utilizing native or engineered GBP are based on the

large change in conformation that occurs on binding of

glucose. Hence, the binding event can be transduced by

environmentally or polarity-sensitive luminophores attached

close to the binding site. Two basic approaches are known,

depending on whether one fluorophore or two fluorophores

are applied. The latter has the advantage of being ratiometric

and therefore being not affected by variations in excitation

light intensity, light path, sample positioning, reagent

concentration, etc.Most sensors rely on the use of mutant proteins

or genetically engineered proteins whose affinity to glucose is

smaller than that of the native GBP. The binding constant of

native GBP for glucose is in the mM range and therefore not quite

appropriate for sensing glucose in blood. Hence, engineered GBPs

were developed for clinical applications with specific mutations in

the binding pocket for reduced affinity to glucose.

Sode and co-workers211 explored a series of site-specific

mutated glucose-binding proteins (GBPs). Their main target

was to reduce affinity to glucose binding while maintaining

their selectivity for glucose. The so-called Asp14Glucose/

Phe16Ala mutant has an affinity constant (Kd) of 3.9 mM,

and this results in an analytical range from 5 to 10 mM of

glucose (Table 9). The intrinsic protein fluorescence (exc./em.

295/350 nm) was detected. Hellinga and Marvin212 reported

on a genetically engineered GBP for construction of a glucose

sensor. They incorporated acrylodan, an environmentally

sensitive fluorophore, close to the binding site. Single cysteine

Table 8 Glucose sensing based on glucose-specific apoenzymes. ANS: 8-anilino-1-naphthalenesulfonic acid; AR: analytical range; FITC:fluorescein-isothiocyanate; FRET: fluorescence resonance energy transfer; GDH: glucose dehydrogenase; GOx: glucose oxidase; GLU: glucose;LI: luminescence intensity; PVA: poly(vinyl alcohol); SWCNTs: single walled carbon nanotubes; TRITC: tetramethyl rhodamine-isothiocyanate

Apoenzyme Method and comments AR/mM Ref.

apo-GOx Flu; GLU induces quenching of intensity and lifetime of apo-GOx non-covalently labeled with ANS(exc./em.=325/480 nm)

10–20 202

apo-GOx Patent; Flu; GLU induces quenching of intensity and lifetime of apo-GOx non-covalently labeled withANS in 3% acetone (exc./em.=370/510 nm)

10–20 203

apo-GDH Flu; GLU induces quenching of intensity and fluorescence polarization of apo-GDH non-covalentlylabeled with ANS

0–60 204

apo-GOx FRET; labeled apo-GOx and labeled dextran; decrease of FRET from FITC to TRITC when theapo-GOx–dextran complex dissociates as a result of the competition of glucose

0–90 205

apo-GOx FRET; microcapsules comprising labeled apo-GOx and labeled dextran multilayer films constructedusing affinity binding and the layer-by-layer self-assembly; decrease of FRET from FITC to TRITC whenthe apo-GOx–dextran complex dissociates as a result of the competition of glucose; 5 times greaterspecificity for GLU over other sugars

0–30 206

apo-GOx FRET; microcapsules comprising Cy5-apo-GOx and TRITC-dextran multilayer films constructed usingaffinity binding and the layer-by-layer self-assembly; decrease of FRET from TRITC to Cy5 when theapo-GOx–dextran complex dissociates as a result of the competition of glucose; 5 times greater specificityfor GLU over other sugars; labeled apo-GOx and labeled dextran

0–40 207208

apo-GOx LI; SWCNTs cross-linked with apo-GOx in PVA polymer; alteration of the swelling state of the hydrogelin the presence of glucose causes change in NIR emission

— 209

Fig. 25 Schematic of a resonance energy transfer glucose assay based

on competitive binding between dextran and glucose for binding sites

on apo-GOx. Reprinted with permission from ref. 205. Copyright 2004

American Chemical Society.

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mutations near the binding pocket were constructed which

allowed for site-specific covalent coupling of the fluorophores.

The fluorescence of acrylodan is quenched on conjugation to

the GBP, but enhances on addition of glucose due to changes

in the polarity of the micro-environment. A series of patents

filed by Becton Dickinson is covering this sensing

scheme.213–215 GBP or the glucose/galactose binding protein

(GGBP), respectively, were site-specifically labeled with dyes

such as IANBD, Nile red, coumarins, or benzodioxazoles. The

conjugates were then incorporated into a hydrophilic polymer

and attached to the end of an optical fiber.

A GGBP mutant labeled with a Nile Red derivative and

possessing a lower affinity to glucose was described by Pitner

and co-workers.216 Upon binding to glucose, the emission at

around 640–650 nm changes by up to 50% in the blood

glucose range. A similar approach217 uses a thermostable

glucose binding protein (tm-GBP). A series of site specific

mutated and labeled conjugates was evaluated. The fluorescence

of the label Cy5 is quenched on binding of glucose, and its

maximum is shifted from 666 nm to 700 nm. This again

enables ratiometric measurements to be performed. The

tm-GBP has also a higher affinity for glucose and enables

Table 9 Sensing of glucose via glucose-binding proteins (both native and engineered). AR: analytical range; CFP: cyan fluorescent protein; Cys:cystein; FLU: fluorescence; FRET: fluorescence resonance energy transfer; GBP: glucose-binding protein; GGBP: glucose/galactose-bindingprotein; GFP: green fluorescent protein; GLU: glucose; IANBD: 4-(N-(iodoacetoxy)ethyl-N-methyl)amino-7-nitrobenz-2-oxa-1,3-diazole; SPR:surface plasmon resonance; tm-GBP: thermostable glucose binding protein; tm-GGBP: thermostable glucose binding protein; YFP: yellowfluorescent protein

Optical method and comments AR/mM Ref.

FLU; specific mutation of GGBP to reduce GLU binding; measurement of the intrinsic protein fluorescence(exc./em. 295/350 nm); increase in fluorescence on addition of GLU; max. 10% signal change

5–10 180

FLU; site specific covalent coupling of acrylodan or IANBD to Cys near the binding pocket of GBP; quenchingof acrylodan and fluorescence enhancement of IANBD on addition of GLU due to changes in the polarity of themicroenvironment

o 0.1 212

Patent; FLU; site specific covalent coupling of IANBD to the Cys group near the binding pocket of GGBP;fluorescence enhancement of IANBD on addition of GLU due to changes in polarity in microenvironment; fibercoated with labeled protein which had been embedded in a polymer matrix

— 213214

Patent; FLU; site specific labeling of GGBP with dye containing squaraine nucleus, Nile Red nucleus, benzo-dioxazole nucleus, coumarin nucleus or aza coumarin nucleus dyes

— 215

FLU; site specific labeling of GGBP with Nile Red derivatives; enhancement of NIR fluorescence (650 nm)intensity on binding of GLU due to changes in polarity in microenvironment

Physiologicalconcentrations

216

FLU; tm-GBP labeled with labels Cy3, Cy5, acrylodan, IANBD; Cy5 conjugate shows best properties; quenchedon binding of GLU and em shift from 666 nm to 700 nm; immobilization in MTP; NIR; self-calibration(ratio of em 666 nm and 700 nm); reversible

1–30 217

Patent; FLU; tm-GGBP labeled with acrylodan; specific sensing group for GLU; reporter group undergoeschange of fluorescence intensity on GLU binding

— 218

Patent; FLU; GGBP labeled with a fluorescent dye or with another protein labeled with a fluorescent dye; specificsensing group for GLU; reporter group undergoes change of fluorescence intensity on GLU binding; fiber coatedwith labeled protein

— 219

FRET; GBP labeled with Alexa-488 (donor) and QSY7 (acceptor) (FRET system) compared with a dimethyl-aminonaphthalene-labeled GBP; max. 16% increase of fluorescence in the FRET system on addition of GLU;300% fluorescence enhancement for labeled GBP

0–0.01 220

FLU; GBP mutant labeled with Badan; 200% increase of fluorescence intensity and 70% increase of fluorescencelifetime on addition of GLU

1–100 221

FLU (lifetime); ANS labeled GBP; decrease only in fluorescence intensity on binding of GLU, no significantchange in lifetime; GBP in cuvette, Ru complex on the outside wall in PVA film; phase-modulation fluorometry

0–0.008 222

Patent; FLU; lifetime; ANS labeled GBP; decrease of fluorescence intensity on binding of GLU, change in lifetime — 223Patent; FLU; polarization; HSA labeled with ANS and GGBP labeled with ANS; self-referenced o 0.1 224FLU; GBP labeled with acrylodan and Ru label; only fluorescence of acrylodan is quenched on binding of GLU;Ru emission is stable; self-referenced probe; also a FRET occurs; phase-modulation fluorometry also possible

o 0.1 225

FRET; GGBP site specific labeled with acrylodan (a) or acrylodan and rhodamine (b); (a) decrease of fluorescenceon addition of GLU up to 10%; (b) increase in emission of rhodamine up to 10% on addition of GLU due toFRET

— 226

Patent; FLU; ratiometric (2 l); GGBP labeled with IANBD or IANBD and Texas Red; labeled protein embeddedin a polymer matrix and coating of a fiber; increase in fluorescence on addition of GLU

0.1–100 227

FLU; GGBP labeled with tetramethylrhodamine-5-iodoacetamide (TMR) and rhodamine red (RR); best resultswith GGBP mutant labeled with 2 TMR units; increase in fluorescence on addition of GLU; tested in simulatedblood serum; modest effect of physiological concentrations of fructose and galactose

0.001–12 228

FRET; GBP dual-labeled with GFPuv (GFP with several mutations to enhance the excitation by UV light)(exc./em. 395/510 nm) and YFP (exc./em. 513/527); reduction of FRET on addition of GLU; system in a dialysishollow fiber; response time: 100 s

o0.1 229

FRET; GBP dual-labeled with CFP (exc./em. 436/480 nm) and YFP (exc./em. 513/535 nm); reduction of FRETon addition of GLU; no significant decrease in FRET by other saccharides; monitors GLU distribution and levelsin living cells

0.07–5.3 231232

Patent; FRET; GBP dual-labeled with FP; implantable device — 233FLU; specific engineered glucose binding protein-like polypeptide; labeled with coumarin derivative; quenchingof fluorescence on addition of GLU; immobilized at the tip of an optical fiber

2–20 234

Patent; SPR; thiol-modified GBP immobilized on the chip surface; change of refractive index on binding of GLU — 235236237

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4832 Chem. Soc. Rev., 2011, 40, 4805–4839 This journal is c The Royal Society of Chemistry 2011

sensing in concentrations from 1 to 30 mM glucose. The

conjugate was immobilized on the bottom of a microtiter plate

to give a reversible response to glucose. The immobilized

protein retains its response after long-term storage at room

temperature. Becton Dickinson has patented218,219 a method

using the tm-GGBP-acrylodan conjugate in a polymer and

attached to an optical fiber.

Pickup et al.220 have compared methods based on either

mono-labeled GGBP or dually labeled GGBP as needed in

FRET assays and came to the conclusion that the former

methods are superior in terms of signal intensity. On the other

side, FRET based sensing has the advantage of being a self-

referenced (2-wavelength) system which makes it more robust.

More specifically, GGBP was labeled with Alexa-488

(the donor label) and QSY7 (the acceptor label) to form a

FRET system. This was compared to the performance of a

GGBP-based sensor where the protein was mono-labeled with

a dimethylaminonaphthalene. The fluorescence of the FRET

system undergoes a change of 16% only, on addition of

glucose, but a 300% enhancement was observed for the

mono-labeledGGBP. The group—rather prematurely—concluded

that a FRET system is less suitable for clinical applications

due to the small signal change observed. In a further

approach221 a GGBP mutant with a rather high binding

constant (11 mM) and an enlarged operating range from 1

to 100 mM was employed and labeled with another label to

result in a 200% increase in fluorescence intensity and a 70%

increase of fluorescence lifetime on addition of glucose.

Phase fluorometry was applied222 to measure the decay time

of ANS-labeled single-mutated GBP as a function of the

concentration of glucose, but—unlike intensity that decreases

by about 50%—no changes of lifetime were detectable. A

phase modulation fluorometry based sensor was constructed

by combination of the ANS-GBP having a short decay time

inside a cuvette with a long-lifetime ruthenium–ligand complex

in a poly(vinyl alcohol) lm on the outer surface of the cuvette.

Binding of glucose changed the relative intensity of the emissions

of the ANS-GBP and the Ru complex. This approach resulted

in the development of a low-cost real-time sensor for glucose

in the low mM range. It was patented by Becton Dickinson.223

A related patent224 describes a sensing scheme based on

measurements of the fluorescence anisotropies of ANS-labeled

GBP in the presence of a reference (ANS-labeled human

serum albumin) of known anisotropy. Due to the additivity

of the anisotropies (of the intensity-weighted average of the

individual anisotropies) any sensing event which induces a

change of intensity of the sensing fluorophore will result in a

change in the measured anisotropy. An inert ruthenium

bipyoxidyl-phenanthroline complex may also serve as a reference.

The sensitivity is in the micromolar range. It was stated225 that

a FRET between acrylodan (donor) and the Ru-complex

occurs and that, based on former approaches,222 phase fluoro-

metric measurements of lifetime may become possible.

D’Auria and co-workers226 have labeled GGBP with

acrylodan only and with both acrylodan and rhodamine. An

up to 10% decrease in emission was observed for the mono-

labeled protein which is in good correlation to formerly

reported acrylodan-GGBPs. The dually-labeled GGBP

displays FRET, and the fluorescence of rhodamine increases

by up to 10% in the presence of micromolar concentration of

glucose. Becton Dickinson has patented227 a ratiometric

sensor based on a NBD (nitrobenzoxadiazole)–GBP–Texas

Red conjugate embedded in a polymer that can be placed at

the distal end of an optical fiber.

A series of GBP mutants was studied in detail by

Dattelbaum and Der.228 GBP was both mono- and dually-

labeled with tetramethylrhodamine-5-iodoacetamide (TMR)

and rhodamine red (RR) in different combinations. The two

labels form ground state dimers whose fluorescence is

quenched. An up to 43% enhancement of fluorescence intensity

is observed in the presence of glucose due to breaking of the

dimer. A site-directed mutation of the glucose binding pocket

was also performed to decrease the affinity for glucose binding.

The dynamic range spanned five(!) orders of magnitude

(0.001–12.000 mM). The performance of the conjugate was

tested in simulated blood serum where it retained its function.

Schultz and Ye229 report on a microsensor based on the

so-called glucose indicator protein (GIP). It contains (a) a

green fluorescent protein (GFP) with several mutations (so to

improve brightness) and with excitation/emission peaks at

395/510 nm, and (b) a yellow fluorescent protein (excitation/

emission 513/527 nm). The reporter units are in close spatial

proximity to warrant efficient resonance energy transfer in the

absence of glucose as shown in Fig. 26. Addition of glucose

increases the distance between the two fluorescent reporters,

and this is accompanied by a decreasing efficiency of the

FRET. The GIP was placed within a dialysis hollow fiber

sensor for continuous monitoring of glucose. This microsensor

displayed reversible response to glucose concentrations in the

micromolar range and a full response within 100 s, but does

not cover the blood glucose range.

In order to adapt the dynamic range to blood levels, the

group of Ye230 have modified GIP by genetic engineering. This

has resulted in a GIP that carries two labels (the cyano

fluorescent and the green fluorescent protein that form a

FRET pair) and whose response matches blood glucose levels

(0–32 mM). An implantable sensor was constructed that is

composed of a hollowmembrane of 1 cm length, an i.d. of 200 mmand a wall thickness of 20 mm. The membrane has a cut-off

weight of 18 kDa. The ratio of the fluorescence intensities

measured at 475 nm and 525 nm serves as the analytical

information. Another GIF is also reported that is capable of

reporting intracellular levels (0–200 mM) of glucose that were

monitored via fluorescence microscopy imaging.

Intracellular glucose levels were monitored231 with a mutant

GBP that was labeled (a) with cyano fluorescent protein

(exc./em. wavelengths of 436/480 nm), and (b) yellow fluorescent

protein (exc./em. 513/535 nm) to create a FRET system. The

addition of glucose caused a decrease in the emission of the

yellow fluorescent protein at 535 nm. The mutant showed a Kd

of B600 mM which allowed monitoring of glucose in various

kinds of cells.232 The GluSense company has patented233 an

implantable glucose sensing device based on the GIP-FRET

sensing scheme.

Daunert et al.234 have genetically engineered a polypeptide

that simulates the glucose binding protein. It was labeled with

a coumarin derivative, immobilized in a polyacrylamide

hydrogel and placed at the tip of an optical fiber. The emission

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This journal is c The Royal Society of Chemistry 2011 Chem. Soc. Rev., 2011, 40, 4805–4839 4833

at 485 nm is quenched in the presence of glucose. The response

time of the sensor is approximately 1 min, and the dynamic range

is from 2 to 20 mM. The method was applied to determine

glucose in whole porcine blood samples.

Becton Dickinson has patented235–237 a surface plasmon

resonance-based sensor family. Thiol-modified GBP was

immobilized on the surface of a gold chip. Glucose binding

was interpreted in terms of a conformational change of the

GBP which eventually results in a decrease in the resonance

angle. It was also concluded that the change from an open to a

closed state results in a decrease in the hydrodynamic volume

which is greater than the increase in mass upon binding.

9. Final assessment

It is obvious from the wealth of literature on sensors for

glucose that there is a substantial need for such sensors. It is

difficult, on the other side, to identify the most suitable sensor

for a specific problem. The sensing schemes treated here differ

in their mode of recognition, the method of transduction, their

analytical ranges and the limits of detection. Each of them

therefore will have a specific application.

Continuous sensing of glucose in context with the artificial

pancreas is generally considered as being the Holy Grail in

biosensor technology. Sensors based on GOx, while used in

commercial clinical instrumentation for discontinuous but

repeated assays, are not well suited for continuous monitoring

for various reasons as outlined in Table 10. Affinity binding

assays using the intrinsic fluorescence of enzymes or oxidized or

reduced cofactors suffer from narrow analytical ranges.

Approaches based on boronic esters hold more promise provided

that a selectivity of >100 for glucose over fructose is warranted

and that effects of pH can be kept under control. Despite

numerous approaches based on the use of concanavalin A, no

system yet has been brought to a state of technology that would

enable its application in vivo. Aggregation is but one of the issues.

In our opinion, the glucose binding proteins provide the best

perspective. Both apo-GOx and GBPs from various sources

may be used. Protein engineering has resulted in modified

proteins so that the physiological range of glucose concentrations

also can be covered. While the intrinsic fluorescence of such

proteins yields a less suitable analytical information,

labelled—and in particular doubly-labelled—GBPs and others

seem to hold the largest promise.

Table 10 Pros and cons of the various schemes reported for sensing glucose

Method Pros Cons

Kinetic enzymatic assaysusing GOx

Kinetic, fully reversible; 2-wavelength andfluorescence lifetime measurements possible; range1–20 mM; works at pH 6–8

Enzyme activity decays with time; critically timedependent, signal change small at low pO2;pH-dependent

Affinity binding of glucose toapo-GOx (FAD removed)

Affinity based, fully reversible; can be calibrated;works at pH 6–8; apo-enzyme fairly stable; signal nottime-dependent; pO2 has no effect; ratiometric andlifetime-based sensing possible

Range 0.1–200 mM; site-specific labeling of enzymeneeded; pH-dependent; apo-enzyme moderatelystable

Chemical binding of glucoseto synthetic boronic acid

Stable system, fully (but slowly) reversible; works atpH 7–8 (but strongly pH-dependent); no effect ofvarying pO2; ratiometric and lifetime-based sensingpossible

Range 1–200 mM; rarely specific for glucose(with 2 exceptions); pH-dependent

Affinity binding to Con A Slow system, fully (but slowly) reversible;calibration possible (very slow); works at pH 7–8(hardly pH-dependent); no effect of varying pO2; 2lFRET, and lifetime measurement possible, range:1–200 mM

Only fairly specific for glucose; ConA tends toaggregate within a few hours; slow; complex; intendedfor in vivo use and continuous monitoring; complexlabeling protocols; ConA is toxic

Affinity binding to glucose-binding proteins

Fully reversible; calibration possible; works atpH 7–8; fairly simple; hardly pH-dependent; pO2 hasno effect; ratiometric (2l) readout possible, FRET,and lifetime measurements possible: range 1–100 mM

Fairly specific for glucose; covers low concentrationrange only unless genetically engineered; partiallycomplex labeling; 1-pt calibration conceivable

Fig. 26 Design of a GIP (glucose indicator protein) for sensing

glucose based on FRET. (a) The GBP adopts an ‘‘open’’ form in the

presence of glucose, which triggers a conformation change, causing

two GFPs to depart from one another leading to the change in FRET.

The dot represents one molecule of glucose bound to the binding cleft

of GBP. (b) Domain structure of the GIP. GFPuv: green fluorescent

protein with several mutations to enhance the excitation by UV light.

YFP: yellow fluorescent protein. GBP: glucose-binding protein.

(c) Spectral overlap of GFPuv and YFP. The absorbance spectra are

denoted by black lines, and the emission spectra are denoted by gray

lines. Reprinted with permission from ref. 229. Copyright 2003

American Chemical Society.

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4834 Chem. Soc. Rev., 2011, 40, 4805–4839 This journal is c The Royal Society of Chemistry 2011

Fluorescence still is the technique of choice but others are

likely to be able to compete. Fiber optic sensing is the method

of choice for in vivo sensing. Direct spectroscopic methods

(usually in the near infrared where glucose displays a weak

intrinsic absorption) have turned out to be less applicable.

Stimuli-responsive photonic crystals also hold great promise.

The wealth of recognition schemes and transduction schemes is

impressive but at the same time makes the situation more

entangling and decisions less easily to be made. A single and

perfect solution for each of the many situations for which glucose

sensors are needed does not exist and is unlikely ever to exist.

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