Journal of Microscopy, Vol. 212, Pt 3 December 2003, pp. 254 –263
Received 13 April 2003; accepted 10 June 2003
Monitoring enzymatic reactions in nanolitre wells
Blackwell Science, Ltd
I . T. YO U N G *, R . M O E R M A N †, L . R . VA N D E N D O E L *,
V. I O R D A N OV ‡, A . K R O O N †, H . R . C . D I E T R I C H †,
G . W. K . VA N D E D E M †, A . B O S S C H E ‡, B . L . G R AY ‡,
L . S A R R O ‡, P. W. VE R B E E K * & L . J. VA N V L I E T *
*Pattern Recognition Group, Faculty of Applied Sciences, Delft University of Technology, Lorentzweg 1,
NL-2628 CJ Delft, The Netherlands
†Kluyver Laboratory for Biotechnology, Faculty of Applied Sciences, Delft University of Technology,
Julianalaan 67, NL-2628 BC Delft, The Netherlands
‡Electronic Instrumentation Laboratory, Faculty of Information Technology Systems and DIMES,
Delft University of Technology, Mekelweg 4, NL-2628 CD Delft, The Netherlands
Key words. Bioluminescence measurements, embedded instrumentation,
fluorescence measurements, microarray, photodiodes, quantitative microscopy.
Summary
We have developed a laboratory-on-a-chip microarray system
based on nanolitre-capacity wells etched in silicon. We have
devised methods for dispensing reagents as well as samples,
for preventing evaporation, for embedding electronics in each
well to measure fluid volume per well in real-time, and for
monitoring the fluorescence associated with the production
or consumption of NADH in enzyme-catalysed reactions.
Such reactions can be found in the glycolytic pathway of yeast.
We describe the design, construction and testing of our
laboratory-on-a-chip. We also describe the use of these chips
to measure both fluorescence (such as that evidenced in NADH)
as well as bioluminescence (such as evidenced in ATP assays).
We show that our detection limit for NADH fluorescence
is 5 µm with a microscope-based system and 100 µm for an
embedded photodiode system. The photodiode system also
provides a detection limit of 2.4 µm for ATP/luciferase bioluminescence.
Received 13 April 2003; accepted 10 June 2003
Introduction
The past decade has seen a rapid growth in the use of microtechnology to produce ever-smaller instrumentation systems
for use in medical diagnostics, studies in cell biology, biochemical process control, and the detection of contaminants and
pathogens in the environment. The general goal of these
systems is to provide high-speed, low-cost, reliable measurements of various biochemical molecules that occur either naturally or as markers.
Correspondence: I. T. Young. Tel.: +31 15 278 1416; fax: +31 15 278 6740;
e-mail: [email protected]
The objective of our research over the past few years has
been to design and build an intelligent analytical system
that measures different molecular analytes simultaneously
using specific molecular interactions in wells on specially
constructed chips. The technology we propose is generic and
could be useful in a variety of applications, such as quality
management in the biotechnology, pharmacology and food
industries, medical diagnostics, and environmental monitoring. The advantage of the proposed technology over existing
methods is that large numbers of different chemical and biochemical analyses can be performed simultaneously in a very
short time, using minimal amounts of reagents and sample.
The large amounts of quantitative data from the measurement system can be evaluated together with qualitative information from experts in the field and historical data for making
decisions in complex environments, thereby increasing speed,
consistency and accuracy, and decreasing costs.
Our work has led to the construction of chips etched in silicon whose lateral dimensions are of the order of 200 µm and
whose depth is of the order of 10 µm for a volume capacity of
about 0.4 nL, and hence are described as nanolitre wells. The
same technology that allows us to produce these wells in
silicon allows us to embed microelectronics in the wells. This
has led to the development of microarrays whose nanowells
contain devices for measuring both liquid volume and light.
Such a chip is shown in Fig. 1.
We have evaluated the performance of our laboratory-on-achip system in the dynamic measurement of liquid volumes
and in the measurement of rhodamine fluorescence, the
NADH fluorescence associated with enzyme-catalysed reactions in the glycolytic pathway of yeast, and the bioluminescence associated with luciferase-catalysed reactions involving
ATP.
© 2003 The Royal Microscopical Society
M O N I TO R I N G E N Z Y M AT I C R E AC T I O N S I N NA N O L I T R E W E L L S
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Fig. 2. Size, shape and capacity of wells etched in silicon. (a) Diameter =
200 µm, depth = 40 µm, volume = 1.26 nL. (b) Diameter = 200 µm,
depth = 23 µm, volume = 0.72 nL. (c) Width = 200 µm, depth = 20 µm,
volume = 0.80 nL. (d) Width = 200 µm, depth = 23 µm, volume = 0.92 nL.
Left, KOH-etched Si; right, RIE-etched Si.
Fig. 1. Laboratory-on-a-chip containing 5 × 5 wells each of dimension
200 µm × 200 µm × 40 µm for a volume capacity of 1.6 nL per well. Each
well contains a photodiode in its ‘floor’ and the external electrical
connections can be seen. The chip is sitting on top of a 96-well microtitre
plate.
Materials and methods
ducted by reactive ion etching (RIE), which results in an equalrate isotropic etching. The RIE-etched wells have a cylindrical or
cubic shape after etching with a circular or square masking pattern, respectively. These wells are also shown in Fig. 2. The wells
are filled according to the methods described in the next section.
Wells in silicon
Reagent and analyte deposition
The microarrays that we use are fabricated in silicon at DIMES
(Delft Institute for Microelectronics and Submicron Technology,
Delft University of Technology, The Netherlands). Different
geometries can be made by using different etching mask
patterns. For our microarrays, one mask with circular patterns has been used and another mask with square patterns.
Typical dimensions of the patterns are 200 µm in diameter or
width, respectively.
The silicon wafers are first covered with a thin layer of
silicon nitride. On top of this layer a thin film of photoresist is
deposited and then exposed through the masking pattern of the
microarrays. After removal of the exposed photoresist, the wafer
is etched using plasma etching. During this process, the silicon is
uncovered at the locations of the wells. The remaining photoresist is removed by rinsing the surface with acetone. Both wet and
dry etching techniques have been employed on the silicon
wafers for the realization of the wells. The wet etching was conducted in potassium hydroxide (KOH). Because of the crystal
structure of silicon, this technique results in an anisotropic
etching of the silicon. The silicon wafers have a <100> surface plane orientation. The etching rate for the <111> crystal
planes is much slower in a KOH solution than for the other crystal planes (typically by a factor of 50). Wet etching of the silicon therefore results in wells with a truncated pyramid shape,
which has <111> planes as side walls for circular as well as for
square etching masks. These two types of wells are shown in
Fig. 2. The angle between the <111> planes and the <100>
plane is approximately 54.7°. The depth of the Si wells for the
largest sizes is typically about 20 µm. The dry etching is con-
We have developed an electrospray mechanism for dispensing
reagents into wells whose volumes range from 60 pL to
500 nL (Fig. 3a) (Moerman et al., 1999b). We have shown that
this mechanism is fast, reliable, reproducible and suitable for
commercial exploitation. In electrospray deposition (ESD),
reagent-bearing fluids are pumped at rates of about 40 pL s−1
through a pipette. A voltage of about 1.2 kV is applied between
the pipette and the wafer and when the electric field strength
exceeds about 3 V µm−1 a mist of very fine droplets is formed
that is deposited in a well. Further details can be found in
Moerman et al. (1999a). A movie of the filling procedure can
be viewed at our web site (www.ph.tn.tudelft.nl/Projects/
DIOC/Progress/DIOC-ESD.html). We have also devised techniques
(illustrated in Fig. 3b) for dispensing samples (analytes) in
such a way that each well is filled with the same volume of
fluid, gas bubbles are eliminated and evaporation is prevented
(Moerman, 2003). Patent applications have been filed for the
techniques that we have developed (Moerman et al., 2001;
Moerman, 2002).
© 2003 The Royal Microscopical Society, Journal of Microscopy, 212, 254–263
Liquid volume measurement
Because it is important to know the amount of fluid that has
been deposited in each and every well, we have developed two
independent methods for assessing the fluid volume per well,
one based upon the use of in situ electrodes and one based
upon the interference fringes observed through a digital imaging microscope system. The former system (Hjelt et al., 1999,
2000a) is ‘real-time’; we can follow fluid evaporation and
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I . T. YO U N G E T A L .
change as a consequence of fluid evaporation. This is shown in
Fig. 5.
If, however, a near-monochromatic light source is used to
illuminate a well filled with water or ethylene glycol, the bottom of the well acts as a reflecting mirror and fringe patterns
are caused by interference between the direct path and the
reflected path of an incident plane wave. The optical path difference (OPD) between the direct and the reflected wave is proportional to the distance to the reflecting bottom surface of the
well. Evaporation decreases the OPD at the meniscus level and
causes alternating constructive and destructive interference
of the incident light, resulting in an interferogram. Imaging
of the space-varying OPD yields a fringe pattern in which the
isophotes correspond to isoheight curves of the meniscus. When
the bottom is flat, the interference pattern allows us to monitor
the liquid meniscus as a function of time during evaporation.
This illustrated in Fig. 6.
This method makes it possible to make an absolute measurement of liquid volume in a well by determining the meniscus height at every point and then integrating this height over
the surface area of the well. We have shown that this technique provides an axial resolution of about 20 nm, which
corresponds to about 70 molecules of water (van den Doel & van
Vliet, 2001). The result of this accurate but computationally
intensive interferometric technique is that we can calibrate
the electrical technique. This is illustrated in Fig. 7, in which
the same evaporation is measured simultaneously by the
two different procedures. The interferometric measurement
technique provides a calibration look-up table for the electrical
procedure.
Fig. 3. Filling of wells. (a) Electrospray deposition of reagent in a well that
is 200 µm × 200 µm and 6 µm deep for a volume of 240 pL. The spray is
from above; the shadow of the nozzle is below. (b) Filling and covering
mechanism for analytes in wells. This process guards against crosscontamination and evaporation and eliminates air bubbles.
filling (see Fig. 4). The latter (van den Doel & van Vliet, 2001)
is absolute in that it provides an absolute measurement of
volume that is accurate in the axial (z) direction to 20 nm,
i.e. about 70 water molecules. The slower interference-based
(optical) technique can be used as a calibration method for the
significantly faster electrical technique (Young et al., 2001).
The electrode-based system works on the basis of the
change in complex impedance that occurs between the electrodes when the fluid volume in the well changes. Using an AC
voltage with frequency 20 kHz, the complex impedance has a
magnitude of about 100 MΩ. The resolution of this measurement system is about 1 pL. Details of this embedded system
for fluid volume measurement can be found in Hjelt et al.
(2000a,b) and Young et al. (2001).
When a solution of rhodamine in ethylene glycol (to slow
evaporation) is observed in a well over an extended period of
time, the fluorescent brightness distribution can be seen to
Fluorescence imaging
The physical signals that we wish to measure in the wells are
based either on fluorescence or – as we shall discuss later – bioluminescence. The fluorescence signals have been primarily
acquired using a digital imaging microscope system (van den
Doel et al., 1999). Used in an ordinary mode the microscope
system consists of an epi-illumination, upright, Zeiss Axioskop
microscope and a Zeiss 2.5× Plan-neofluar microscope objective with an NA of 0.075. This objective allowed us to observe
the entire 5 × 5 array of wells, as shown in Fig. 8. The detection efficiency was measured by imaging the 25 wells and
looking at the total brightness per well when they were filled
with a solution of rhodamine.
Each well in a column of the array was filled with the same
concentration of rhodamine and thus variations of brightness
within a column were an indicator of the inherent variability
of the filling and measurement process. The concentration of
rhodamine was varied from column to column, allowing us to
assess the sensitivity of the measurement system and its detection limit. A KAF 1400 Photometrics camera was used to
acquire the images and one such image, taken with the 2.5×
lens, is shown in Fig. 8(a).
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Fig. 4. Measurement of liquid volume. (a) Construction of an in situ liquid volume sensor using aluminium electrodes. (b) On-chip circuitry to measure
the complex impedance between the aluminium electrodes, Zx, due to the volume of fluid in the well. (c) Voltage change caused by filling a 540-pL well with
fluid. (d) Voltage change caused by fluid evaporation from a 424-pL well. Two independent experiments are superimposed and the time of rupture (100 s)
is indicated.
Fig. 5. Rhodamine evaporation in a 200-µm × 200-µm well as a function of time. Above, fluorescent images; below, fluorescence as a 3D surface. Left,
0 min; middle, 8.5 min; right, 15.5 min.
© 2003 The Royal Microscopical Society, Journal of Microscopy, 212, 254–263
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I . T. YO U N G E T A L .
Fig. 6. Evaporation of fluid from a well produces fringe patterns on the surface of the liquid as observed with an epi-illumination microscope. (Above)
Confocal image in x–z display that shows decrease of meniscus height. (Below) Fringe pattern observed in x–y display of conventional microscope. Time of
observation with the flat meniscus point being considered as t = 0: left, −53 s; middle, 97 s; right, 164 s.
Fig. 7. Evaporation of ethylene glycol in a well is followed simultaneously by (a) the electrical technique and (b) the interferometric technique. Both
techniques give a well-defined reading until the meniscus ruptures. The time of observation is in seconds with the flat meniscus point being considered
as t = 0.
We define the detection limit as that concentration at which
the fluorescence signal is more than 3σ above the background
level associated with a control (blank) solution. The data from
this experiment produced a detection level of about 1 µm concentration. Given the volume of the well, this corresponded to
about 108 fluorescing molecules per well.
We attempted to improve this detection limit – that is reducing it – by first noting that the collection efficiency of light from
the wells is given by NA2 and insensitive to the value of the
magnification. We therefore switched to an infinity-corrected
Zeiss 20× fluar objective with an NA of 0.75, which provided
a 100× improvement in light collection. This step, together
with proper adjustment of the field stop and the use of critical
illumination instead of Köhler illumination, provided a 1000fold improvement in the detection limit to a value of 1 nm,
which corresponds to 100 000 fluorescing molecules per well.
We then replaced the 1× tube lens with a Zeiss infinitycorrected 5× fluar objective with an NA of 0.25 and a working distance of 9.3 mm. This concentrates all the photons
acquired by the primary lens in a smaller region on the CCD
and thus increases the photon flux in that region by 25×. If we
were interested in spatial resolution within a well this would
be a problem. But we are not; we are interested in sensitivity
and lower detection limits (van den Doel et al., 2002). We also
replaced our older camera with a new Princeton Versarray
512B back-illuminated CCD camera. This instrument provides 16 bits of dynamic range and variable readout rates. A
detailed analysis of the issues involved in camera choice and
usage can be found in Mullikin et al. (1994) and van den Doel
et al. (2002). In an important sense, a photomultiplier system
would work as well as our digital camera configuration but the
latter gives us considerably more flexibility. This modification
is shown in Fig. 9.
Measuring enzymatic reactions
We have monitored the reaction rates derived from cell-free
extracts for various enzymatic reactions that are found in
© 2003 The Royal Microscopical Society, Journal of Microscopy, 212, 254 –263
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Fig. 8. Determination of the detection level for fluorescent signal using a
criterion based upon the 3σ level associated with the control. (a) A 5 × 5
array of wells each with width of 200 µm and a pitch of 600 µm. Each
column is filled with a solution of constant concentration. This concentration decreases from left to right until the fifth column, which is the
control column. (b) Measurement of total brightness within each well as a
function of the rhodamine concentration. The 3σ error bars are indicated.
The rhodamine solution is detectable above the 1 µm concentration.
the glycolytic pathway of yeast (Saccharomyces cerevisiae). The
enzymes are glucose-6-phosphate dehydrogenase (G6P-DH),
lactate dehydrogenase (LDH) and alcohol dehydrogenase
(ADH). These reactions are:
G6P-pyruvate + NAD ↔ gluconolactone + NADH (G6P-DH)
(LDH)
pyruvate + NADH ↔ lactate + NAD
(ADH)
ethanol + NAD ↔ acetaldehyde + NADH
For all three enzymes the measurement is based on the conversion of NAD(P) to NAD(P)H or vice versa and the fluorescence
© 2003 The Royal Microscopical Society, Journal of Microscopy, 212, 254–263
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Fig. 9. Microscope system for measurement of fluorescence in nanowells.
(a) A Zeiss Axioskop using an ordinary 20×/0.75 fluar objective, and
inverted 5× objective and a Princeton Versarray CCD camera. (b) The
trinocular tube has been removed and replaced by an inverted 5×/0.25
Zeiss Fluar lens whose working distance is 9.3 mm.
of NADH with an absorption maximum at λ = 340 nm and an
emission maximum at λ = 450 nm. Our detection limit for
NAD(P)H – measured with the method described above – is
about 5 µm, which corresponds to 3 × 109 molecules in a 1-nL
well. Data are automatically incorporated into an analysis
module of our software environment genlab (www.genlab.
tudelft.nl/) (Wessels et al ., 1999), which runs under
matlab. Figure 10(a) shows one image of an array of 5 × 5
wells taken at one point in time. Figure 10(b) shows the
analysis of a sequence of images taken over 2000 s to produce
a profile per well of the consumption of NADH in the LDHcatalysed reaction. The rate constant for each well can be
determined and used to assess enzyme kinetics. The data shown
below are based upon cell-free extracts from Saccharomyces
cerevisiae.
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Fig. 10. Twenty-five images were sequentially scanned and then assembled into the 5 × 5 array display shown in (a). In (b) the dynamic behaviour of
NADH conversion to NAD over an interval of 2000 s. The enzyme concentration LDH is 0.1 U mL−1. (An enzyme concentration of 1 U mL−1 equals an
amount of enzyme molecules that converts 1 µm of substrate in 1 mL solution per minute.) The data follow a first-order exponential decay reaction model.
The italic number in each graph is the decay time τ.
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The measurements shown above were made using the Princeton Versarray 512B CCD camera. As mentioned earlier,
spatial sampling of the region within a well is not necessary.
Furthermore, signal-to-noise considerations show that it is
not desirable.
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A more appropriate sensing technology would involve
the use of one light-sensing device per well and this option
is possible by embedding photodiodes in the ‘floor’ of each
well. The problem associated with this approach, however,
is that the sensor is now ‘looking into’ the excitation source,
a step backward from Ploem’s invention of epi-illumination
fluorescence microscopy in the 1960s (Ploem, 1967). If we
are actually to achieve this goal then filtering will be required
between the liquid sample and the surface of the photodiode.
This is achieved through the use of a polycrystalline silicon
layer.
In situ light measurement
An array of photodiodes can be fabricated in such a way that it
is aligned with our wells and that there is a thin film (∼75 nm)
of re-crystallized silicon between the photodiode and the
liquid sample. The photodiode current can then be sampled
and digitized to measure the photon flux at the sampling time.
The crystalline structure, with a grain size of about 2 µm, is
capable of blocking excitation wavelengths below 400 nm
and thus permitting only the emission wavelengths to pass
(Iordanov et al., 2002). A microarray chip incorporating these
photodiodes and a schematic of the structures are shown in
Fig. 11(a,b).
With this technique a suppression of about 35 dB has been
achieved in practice (see Fig. 11c) and we have reason to
believe that this can be improved to about 50 dB. The results
for measuring NADH fluorescence are shown in Fig. 12.
Using the technique described above to determine the detection limit – 3σ above the control level – the in situ photodiodes
currently achieve a detection limit of about 100 µm. This is
not as good as the NADH detection limit achieved through
the use of the CCD camera, 5 µm. Our poly-Si filter deposition
Fig. 11. In situ photodiodes in a 5 × 5 array of nanowells. The poly-Si thin
film filter blocks excitation light so that it is possible to observe the
fluorescence emission. (a) Micro-array chip with in situ photodiodes and
thin-film Si filters. (b) Schematic representation of a well including the in
situ filter and pohotodiode (not to scale). (c) Comparison of the simulated
to measured filter characterstics for poly-Si thin film filter deposited on glass.
© 2003 The Royal Microscopical Society, Journal of Microscopy, 212, 254–263
Fig. 12. The poly-Si filter clearly suppresses sufficient excitation light to
make fluorescence measurements possible. Note the use of two different
vertical scales. The photodiode current has been sampled at time intervals
of 210 ms.
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I . T. YO U N G E T A L .
technique will have to be improved if we are to achieve the
results obtained with a scientific CCD camera and microscope.
That this problem is primarily associated with the short wavelength blocking effects of the poly-Si filter can be seen in the
following experiments with bioluminescence where no excitation source is used.
Bioluminescence measurements
We have used a Roche ATP Bioluminescence HS II Assay Kit
(Roche Molecular Biochemicals, 1999) to test our filter-covered
photodiode arrays as well as to determine the suitability of
our technology for those applications in which bioluminescence is more appropriate than fluorescence. The specific
reaction is:
ATP + d-luciferin + O2 → oxyluciferin + PPi
+ AMP + CO2 + light
(LUCIFERASE)
where the emission spectrum has a maximum at λ = 562 nm.
We covered a photodiode with 500 nL of solution and
measured the luciferase-catalysed reaction over an interval of
10 min with sampling intervals of 210 ms using varying concentrations of ATP: 0, 10, 20 and 40 µm corresponding to 0,
5, 10 and 20 pmol. We measured the signals from an ‘empty’
photodiode and from the reaction where the ATP concentration was zero as controls. The first 4 min of the results are
shown in Fig. 13.
The empty photodiode produced a signal with an average
value of 3.24 pA and a standard deviation of 0.07 pA whereas
the blank photodiode (ATP concentration = 0) produced a
signal with an average value of 4.32 pA and a standard deviation of 0.08 pA = 80 fA. This indicates that whereas the blank
might have an increased average value, the contribution to
noise in the measurement process is determined by the photodiode itself; the blank liquid does not contribute to the noise.
These statistical measurements were made in the interval
50–200 s, when the electronic transient behaviour had
stabilized. When we subtract the blank average value from the
measurements where ATP was added, we obtain the result
shown in Fig. 14.
When we look at the average value of each of these curves
vs. the ATP concentration, we find that the sensitivity is about
0.1 pA µm –1. The standard deviation (σ) of the noise in the
blank, 80 fA, then translates into 800 nm and a detection level
of 3σ to 2.4 µm. This is very close to the levels that would make
a bioluminescence assay practical. The data are not strictly
linear in the ATP concentration. This can be explained by the
fact that we are operating outside of the linear range of the
HS1II kit as specified by the manufacturer, which is 1 pm to
1 µm (Roche Molecular Biochemicals, 1999).
Summary and conclusions
Fig. 13. Photocurrent of the HS II ATP bioluminescence kit on a single
photodiode. The signals were recorded without any liquid (empty), filled
with HS II reagent and an ATP concentration of zero (blank), and in the
presence of different, non-zero concentrations of ATP.
Fig. 14. The average value of the blank signal has been subtracted from
the photodiode response. The resulting data are displayed in the range
50 ≤ t ≤ 200 s.
We have developed a laboratory-on-a-chip system that is
capable of measuring enzymatic reactions based upon either
fluorescence or bioluminescence. The individual wells in
this microarray have volume capacities of the order of 1 nL.
We have developed and patented delivery systems for both
reagents (using electrospray deposition) and analytes. We have
been able to show that enzyme activity is not affected by the
electrospray deposition; a manuscript describing these experiments is in preparation. Our system for analyte delivery prevents cross-contamination, air bubbles and evaporation.
We have developed systems based on microscopy that permit:
1 measurement of rhodamine fluorescence to concentrations
of 1 nm; and
2 measurement of NADH fluorescence to concentrations of
5 µm.
Further, we have developed in situ (embedded) instrumentation that permits
1 real-time measurement of analyte volume to a precision (σ)
of about 1 pL;
2 measurement of NADH fluorescence to concentrations of
100 µm; and
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3 measurement of ATP-associated bioluminescence to
concentrations of 2.4 µm.
The photodiodes are capable of producing a measurement
of light flux that is both sensitive and quantitative. The sensitivity can be improved in several ways, all of which are being
investigated at this time.
1 Instead of sampling the photodiode current at intervals of
210 ms, we can integrate the current on capacitors that
are embedded in the chip in a manner similar to CCD
camera sensors and then read out the total charge after
a sufficiently long integration period.
2 We can improve the thin-film poly-Si filters to produce
greater wavelength selectivity for fluorescence measurements so that a rejection ratio of some 50 dB instead of
35 dB can be achieved.
3 For bioluminescence, we can use photodiodes that are
not covered by the poly-Si filters. Around the wavelength
of interest, λ = 562 nm, the poly-Si filter has a −10 dB
suppression, meaning that we are losing 90% of the
incoming photons. As there are no excitation photons
to contend with, this means that we are losing emission
signal. In our next series of experiments we will be using
filter-free photodiodes.
In an ideal situation we would be able to measure the ATP
activity in a single (yeast) cell. At the present time the detection level we have achieved (2.4 µm) means that we are measuring the ATP associated with more than 106 cells. This
assumes that a single yeast cell, on average, contains 6.5 µg
ATP per gram (dry weight) cell (J. van Dam, personal communication, 2002). We are clearly a long way from a single-cell
metabolic assay.
Acknowledgements
This work was partially supported by the Physics for Technology programme of the Foundation for Fundamental
Research in Matter (FOM), the Delft Inter-Faculty Research
Center Intelligent Molecular Diagnostic Systems (DIOCIMDS) and the Delft Inter-Faculty Research Center LifeTech
(DIOC- LifeTech).
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