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Miniaturized thermal biosensors
Sensorsand Ariuolors B, J5- 16(1993)443-447
443
miniaturized thermal biosensors Bin Xie and Bengt Danielsson Pure and Applied Biochemistry, Universityof Lund, 22100 Lund (Sweden)
Fredrik Winquist Laburatory of Applied Physics, Linkiiping University,58183 Linkiping (Sweden)
Abstract Miniaturized thermal biosensors based on three different designs have been constructed. Properties relevant to decentralized bioanalysis have been investigated. A short m~~ment period (30 s per sample) and a broad linear mnge (OS mM to 100 mM for glucose and ~nic~lin-V), using I pl sample volume, have been achieved.
Introduction During the past 15 years calorimetric biosensors have been successfully developed and extensively applied in many fields including clinical chemistry [ 1,2], thermometric enzyme-linked immunosorbent assay (TELISA) [3], fermentation analysis and process control [4,5], and environmental analysis f6]. The increasing interest in calorimetric biosensors is derived from their advantages in general applicability, in~nsiti~ty to the optical and electrochemi~l properties of the samples, reusability of the biocatalyst, the possibility for continuous-flow operation, and simple operational procedures [7]. For the applications of biosensors in personal health-monitoring, such as diabetes control, miniaturization of the devices becomes a prerequisite. In addition, the improvement of some properties, such as sensitivity, linear range and measurement period, can be anticipated due to the reduction of the heat capacity and sample volume. Therefore, the investigation of three different types of ~niatur~ed thermal biosensors has been undertaken to determine the degree to which a fun~~oning system for flow-inaction analysis can be scaled down.
Experimental The selection of the constructions below as models was based on the following considerations. As a first step, a version of the conventional enzyme thermistor with considerably reduced dimensions was fabricated. Secondly, a chip sensor based on a plastic construction was made to provide a simple, cheap and flexible copy of micromachined chips. Of potential interest for implantation or in viva use, is the very simple microcolumn sensor that was as a third alternative.
0925-4005/93/$6.00
During the operation, the sensors were located in an aluminium box insulated with polyurethane foam. All experiments concerned here were conducted using sodium phosphate buffer (0.1 M, pH = 7.0) at room temperature (22 “C). Materials and appararus Glucose oxidase (EC. 1.1.3.4) from Afgergillus niger
(type X-S, with a specific activity of 150 U mg-’ solid) and ~-lactose (E.G. 3.5.2.6) from 3u~~i~s cereus (type I, with a specific activity of -2000 U mg-’ solid) were obtained from Sigma (St. Louis, MO, USA). Catalase (B.C. I. 11.1.6, from beef liver) was purchased from Boehringer (Mannheim, Germany). Glutaraldehyde was supplied by Sigma. y-aminopropyltriethoxysilane was purchased from Pierce Eurochemie, (Rotterdam, The Netherlands). Controlled-Pore Glass (particle diameter: loo-160 pm, pore diameter: 50 nm) was obtained from Veb Trisola (Steinach, Germany). Microbead thermistors (MI%-3) were supplied by Wuhan Electronic Component Factory (Wuhan, China). A d.c.-coupled ~ea~tone bridge ~~nst~cted at our institute) was used for ~rn~ra~re m~s~~ents. At ma~mum s~itivity, the signal changes 100 mV per 10m3“C. Besides the peristaltic pump (Alitea pump unit C-4V from Ventur Tekniska AB, Sweden), a perfusion pump (B. Braun Melsungen AG, Germany) and HPLC valves suitable for I ~1 and more than 5 pl samples, (type C14W and C6W, VIGI AG Valco Europe, Switzerland), respectively, were used. Preparation of the enzyme colwnm
Controlled-Pore Glass (CPG), on which enzymes were covalently immobilized, was used as support material of the enzyme columns. The i~ob~zation procedures were Performed as follows. CPG with a particle
@ 1993- Ekevier Sequoia.All rightsmewed
444
size distribution of lOO- 160 pm was first boiled in 5% HNO, for 45 min and was extensively washed with distilled water to clean and hydrate the surface. To a solution of 2 g of y-aminopropyltriethoxysilane in 18 ml of water was added 1 g of clean glass. The pH was adjusted to 3.5 with 6 M HCl and the solution was heated at 75 “C for 3 h with occasional gently stirring. The glass was subsequently collected on a Biichner funnel, washed with distilled water and dried for at least 4 h at 115 “C. To 1 g of alkylamino glass was added 25 ml of 2.5% glutaraldehyde in 0.1 M sodium phosphate buffer adjusted to pH 7.0. The reaction was allowed to take place under reduced pressure for 30 min (using a water aspirator) and then at atmospheric pressure for at least 30 min. The brick-red coloured product was washed exhaustively on a Btichner funnel with distilled water and stored in water in the refrigerator. The enzyme was added to the activated-glass suspension in as small a volume as possible, usually in 0.1 M sodium phosphate buffer, pH 7.0. The protein concentration should preferably be at least l%, i.e., between 50 and 100 mg of enzyme should be used per g of carrier. For glucose oxidase, coimmobilization of catalase was carried out. The coupling was often performed at room temperature and was allowed to proceed for 1 to 4 h with gentle mixing or in the cold overnight. The enzyme preparation was thoroughly washed with buffer prior to use. Constructions Compact sensor
were mounted on the capillaries with a heat-conducting epoxy. Plastic chip sensor
The chip sensor (27 mm x 7 mm x 6 mm) was constructed of Plexiglass (see Fig. 2). The rectangular enzyme cell (5 mm x 3 mm x 0.5 mm) and part of the inlet and outlet were milled into the plastic base to a depth of 0.5 mm. Electrically-insulated thermistors in direct contact with flow stream were placed outside the enzyme cell after the porous polyethylene filters. The enzyme cell was charged with enzyme preparation prior to compaction. Replacement of enzymes is possible due to ready access to the enzyme compartment. Microcolumn sensor
The microcolumn sensor (outer diameter x length: 0.8 mm x 15 mm) was constructed of stainless-steel tubing and free from auxiliary components (Fig. 3). A microbead measurement thermistor was directly mounted on the outer surface at the latter part of the column using heat-conducting epoxy and the reference thermistor was positioned on a piece of the inlet metal tubing. Both the length (about 200 mm) and inner diameter (0.15 mm) of the inlet tubing between the sample valve and the column were minimized in order to reduce sample dispersion during the transportation in
Thermistor
Inlet
The construction of a compact sensor of greatly-reduced dimensions (outer diameter x length: 36 mm x 46 mm) is shown in Fig. 1. In order to conveniently accommodate enzyme columns and to ensure isolation from ambient temperature fluctuations, a cylindrical copper heat sink was included. An outer Delhi jacket further improved the insulation. The enzyme column (outer diameter x length: 4 mm x 3 mm), constructed of Delrin, was held tightly against the inner terminals of the copper core. Short pieces of well-insulated gold capillaries (outer diameter/inner diameter: 0.3 mm/ 0.2 mm) were placed next to the enzyme column as temperature-sensitive elements. Microbead thermistors
Filters
Outlet
Enzyme’cell (a) Enzyme
Covering
ccl ThermiYr
\
Ff”er
/ liim
(b) Enzyme
Fig. 2. (a) Schematic diagram and (b) cross section of the plastic chip thermal biosensor.
column
Inlet
Enzymecolumn
\
Thermistor
Outlet
/ 5 mnl
1
H
cm
Fig. 1. Schematic cross section of the compact biosensor.
Stainless steel tubing
Nylon tubing
Fig. 3. Schematic diagram of the microcolumn sensor.
445
the flow system. Similarly, to reduce heat loss, simple and short connections between the column and the inlet or outlet tubings were required.
Results Effect
and diiussion of miniaturization
on properties
The sensitivity and response process of calorimetricFIA biosensors can be explicitly depicted by their steady-state responses. Flow rates of 108, 75 and 40 nl min-’ were used for the compact sensor, the chip sensor and the microcolumn sensor, respectively. Buffer and penicillin-V (10 mM) were alternatively introduced into the /I-lactamase columns (20 min cycle) using a peristaltic pump. Steady-state response curves corresponding to different degrees of miniaturization of the sensors were obtained and compared in Fig. 4. The height of the curves is governed by the equilibrium of exothermic reactions and heat leaking from the reactive systems. The former depends on the catalytic efficiency of enzymic reactions in the columns, and the latter is dominated by adiabatic properties of the reactors. In general, the greater the amount of enzyme, the higher the catalytic efficiency. Miniaturization necessitates a limited amount of enzymes. Decrease of the enzyme amount could result in decrease of the linear range of the measurement unless flow rate and sample volume are carefully controlled. This was demonstrated in an experiment. Using 20 ul of penicillin-V, the linear range of measurement was limited to 100 mM, 50 mM to 40 mM in correspondence with the volume of the enzyme reactor, which was 21,8 and 4 nl with the compact, the chip and the microcolumn fl-lactamase sensors, respectively. In the latter case, the heat conductivity of the materials
and the heat capacity of the enzyme columns and their surroundings are the major contributory factors. A smaller column may therefore lead to higher sensitivity if a good adiabatic condition can be maintained. This is demonstrated in Fig. 4. Nevertheless, the slope is an exception. It is not directly related to the degree of miniaturization, as with the chip sensor. Larger heat capacity of the enzyme-column surroundings may be the cause of the increased response time, because the reactor bed of the chip sensor acts as both enzyme support and insulating layer with significant heat capacity. The experimental results revealed that a thin-walled metal column surrounded by a layer of air can offer a faster response and much higher sensitivity than plastic columns. For these construction-dependent biosensors, further miniaturization will necessitate the use of smaller enzyme volume and finer carrier materials. This will require a new technology, such as micromachining [8], or more sensitive transducers. Effect
of sample
volume
on properties
One of the purposes of miniaturization is to reduce the sample volume necessary for analysis, which would be of benefit in cases where only limited sample volumes are available. Furthermore, as a reduction of the sample volume increases the linear range, as demonstrated by the conventional enzyme thermistor [I], we anticipated a similar result with the miniaturized glucose oxidase sensors, although at the expense of lower sensitivity. The result with the chip sensor, using different volumes of 1, 5 and 20 ~1, is shown in Fig. 5. The linear range of standard glucose concentrations extended from 8 mM up to 100mM when the sample volume was reduced from 20 ul to 1 ~1. A similar result was also obtained with the microcolumn sensor, although the linear range was limited to 30 mM. The extension of the linear range when reducing the sample volume might be attributed to a relative increase
1
9 0
_
‘Ix____ 0
2
4
6
8
10
0 5 G 12
Time (mins) Fig. 4. Comparison in steady-state responses between the compact (M-l), the plastic chip (M-Z) and the microcolumn (M-3) sensors charged with immobilized fl-lactamase. Penicillin-V samples (10 mM) were pumped with their corresponding optimal flow rates of 108, 75 and 40 pi min-I, respectively.
d
2b
4b
sb
Glucose Cont.
100
(mM)
Fig. 5. Effect of sample volume on the linear range of the chip sensor loaded with immobilized glucose oxidase. Flow rate of 75 11 min-’ was used.
446
in the dispersion of the small sample volume. This factor results in dilution of the sample in the flowinjection analysis (FIA) system [9], particularly in the column. Since reactive processes in an enzyme column are determined by sample volume, a smaller volume should yield a fast response time. In practice, however, this is not always observed due to dispersion of sample between the sample inlet and the column [lo]. Consequently, the response time was not proportional to the sample volume. If the inlet is small in inner diameter and short in flow path (inner diameter x length: 0.15 mm x 200 mm), the measurement period per sample as fast as 30 s (glucose or penicillin-V of 10 mM) was still achieved with the microcolumn biosensor in 1 pl sample volume and a flow rate of 50 pl/min. Effect ofjow rate on properties
In calorimetric FIA biosensors, many properties, apart from the factors discussed above, are subject to the flow rate. The improvement by means of this way is much concerned with response time and linearity, but has less effect on the sensitivity. Too high a flow rate, however, may cause a reduced peak height due to the limited time available for catalytic reactions as shown by varying the flow rate using the chip /?-lactamase sensor. The results in Fig. 6 were obtained by just changing the flow rate instead of the different dimensions of the miniaturized biosensors. It was judged that the flow rate should be less than 100 pl/min in spite of the enlarged response time. The influence of the flow rate on calorimetric FIA biosensors may be characterized by a factor, Q, which is the ratio of an optimal flow rate (FJ to a matrix volume of the enzyme column (V,), such that Q = FJV,. For
TABLE 1. Comparison of characteristic factor Q (min-I) of caloFimetric FIA biosensors: ET, the enzyme thermistor; M- 1, the compact sensor; M-2, the plastic chip sensor; M-3, the microcolumn sensor
type
Column
Column volume (11)
Optimal flow rate (~1 min-‘)
Characteristic factor Q (min-‘)
ET M-l M-2 M-3
200 20 10 4
1000 100 75 20
5 5 1.5 5
a constant column volume, a smaller Q (min-I) value indicates a higher temperature-conversion efficiency of the thermal biosensor. The comparison of characteristic factors between the conventional enzyme thermistor (ET) and the miniaturized thermal biosensors is shown in Table 1. The results reveal that the characteristic factor Q does not become smaller with miniaturization of the sensors. Thus, the miniaturization process does not improve the temperature-conversion efficiency of the thermal biosensors, although some propertie%,such as sensitivity, response time and linear range have been improved. That the chip sensor presents a bigger Q value than others implies that the characteristic factor Q could be also related to the adiabatic situation of the enzyme column, which is less favorable here.
Life time and operational stability
For the chip and the microcolumn sensors, the enzyme reactors are preferably unified with the transducer parts in order to reduce sizes. Therefore, the replacement of enzyme preparations is not as convenient as for easily-disassembled systems. The employment of disposable transducers, e.g. thin-film thermistors [ 111, would facilitate practical operation. Storage of these miniaturized devices must be carefully undertaken. The cells must be filled with buffer solution between experi-
50
40
0 ,"
30
0 =.
20
G 10
Time
(mins)
Fig. 6. Effect of flow rate on sensitivity and response time of steady-state responses. The plastic chip sensor loaded with immobilized fl-lactamase and penicillin-V of IO mM were used: (a) 54 filmin-‘; (b) 105 ~1mix’; (c) 75 fl mix’.
0
100
200
Sample Number Fig. 7. Reproducibility o$ penicillin-G (10 mM) using the plastic chip sensor loaded with /Xxctamase.
447
ments and the systems kept at 4 “C. Under these conditions, using /?-lactamase and glucose oxidase, a life time of more than one month was obtained with the above three sensors. Investigation of the operational stability of these miniaturized biosensors is essential. The stability was tested by injecting more than 200 penicillin-G samples into the chip /?-lactamase sensor. The relative standard deviation (1%) was obtained. Similarly, using the microcolumn glucose oxidase sensor, this value was 6% on 70 glucose samples of 10 mM. Figure 7 shows the reproducibility of the chip sensor for the 200 assays.
Conclusions
Three different types of miniaturized calorimetric FIA biosensors have been developed and studied. Individual advantages are presented from various aspects. The compact sensor, as a version of the conventional enzyme thermistor, retains the same functions, and offers great flexibility in the range of possible applications. The chip sensor is readily constructed and is capable of easy and accurate determinations. The microcolumn sensor presents a fast response time and a higher sensitivity. Miniaturization of thermal biosensors, due to increase of the sensitivity, results in the possible use for such a small sample volume of 1 ~1. This makes it possible to expand the linear range of glucose (from 8 mM to at least 30 mM) and to reduce the measurement time (from more than 1 min down to 30 s). The large linear range possible for glucose determination is of particular interest in diabetes control where the glucose concentration is up to 20mM. Consequently, this method provides a convenient way to directly monitor the blood glucose without the tedious process of sample dilution.
Acknowledgement
This work has been partly supported by a grant from the National Swedish Board for Technical Developments (NUTEK).
References 1 B. Danielsson, K. Gadd, B. Mattiasson and K. Mosbach, Enzyme thermistor determination of glucose in senmr using immobilized glucose oxidase, C/in. Chim. Am, 81 (1977) 163-175. 2 B. Danielsson, Calorimetric biosensors, J. Biotechno/., 15 (1990) 187-200. 3 B. Mattiasson, C. Borrebaeck, B. Sanfridsson and K. Mosbath, Thermometric enzyme-linked immunosorbent assay: TELISA, Biochim. Biophys. Acta, 483 (1977) 221-227. 4 G. Decristoforo and B. Danielsson, Flow injection analysis with enzyme thermistor detector for automated determination of fi-lactamase, Anal. Chem., 56 (1984) 263-268. 5 C. F. Mandenius, L. Billow, B. Danielsson and K. Mosbach, Monitoring and control of enzymic sucrose hydrolysis using on-line biosensors, Appl. Microbial. Eiotechnol., 21 (1985) 135-142. 6 B. Danielsson, B. Mattiasson and K. Mosbach, Enzyme thermistor devices and their analytical applications, Appl. Biochem. Eioeng., 3 (1981) 97-143. 7 B. Danielsson and K. Mosbach, in A. P. F. Turner, I. Karube and G. S. Wilson (eds.), Biosensors: FundamentaLr and Appk cations, Oxford University Press, London, 1987, Ch. 29, pp. 575-595. 8 B. Xie, B. Danielsson, P. Norberg, F. Winquist and I. LundStrom, Development on a micro-thermal biosensor fabricated on a silicon chip, Sensors and Actuators B, 6 (1992) 127- 130. 9 I. Ruzicka and E. H. Hansen (eds.), Flow Injection Analysts, Wiley, New York, 1981. 10 J. Ruzicka and E. H. Hansen. Flow iniection analvsis: Dart X. Theory, techniques and trends, AnaLSChim. Act;, 99’( 1978) 37-76. 11 G. Urban, H. Kamper, A. Jachimowicz, F. Kohl, H. Kuttner, F. Olcaytug, P. Goiser, F. Pittner, T. Schalkhammer and E. Mann-Buxbaum, The construction of microcalorimetric biosensors by use of high resolution thin-film thermistors, Biosensors Bioelectronics, 6 (1991) 275-280.
443
miniaturized thermal biosensors Bin Xie and Bengt Danielsson Pure and Applied Biochemistry, Universityof Lund, 22100 Lund (Sweden)
Fredrik Winquist Laburatory of Applied Physics, Linkiiping University,58183 Linkiping (Sweden)
Abstract Miniaturized thermal biosensors based on three different designs have been constructed. Properties relevant to decentralized bioanalysis have been investigated. A short m~~ment period (30 s per sample) and a broad linear mnge (OS mM to 100 mM for glucose and ~nic~lin-V), using I pl sample volume, have been achieved.
Introduction During the past 15 years calorimetric biosensors have been successfully developed and extensively applied in many fields including clinical chemistry [ 1,2], thermometric enzyme-linked immunosorbent assay (TELISA) [3], fermentation analysis and process control [4,5], and environmental analysis f6]. The increasing interest in calorimetric biosensors is derived from their advantages in general applicability, in~nsiti~ty to the optical and electrochemi~l properties of the samples, reusability of the biocatalyst, the possibility for continuous-flow operation, and simple operational procedures [7]. For the applications of biosensors in personal health-monitoring, such as diabetes control, miniaturization of the devices becomes a prerequisite. In addition, the improvement of some properties, such as sensitivity, linear range and measurement period, can be anticipated due to the reduction of the heat capacity and sample volume. Therefore, the investigation of three different types of ~niatur~ed thermal biosensors has been undertaken to determine the degree to which a fun~~oning system for flow-inaction analysis can be scaled down.
Experimental The selection of the constructions below as models was based on the following considerations. As a first step, a version of the conventional enzyme thermistor with considerably reduced dimensions was fabricated. Secondly, a chip sensor based on a plastic construction was made to provide a simple, cheap and flexible copy of micromachined chips. Of potential interest for implantation or in viva use, is the very simple microcolumn sensor that was as a third alternative.
0925-4005/93/$6.00
During the operation, the sensors were located in an aluminium box insulated with polyurethane foam. All experiments concerned here were conducted using sodium phosphate buffer (0.1 M, pH = 7.0) at room temperature (22 “C). Materials and appararus Glucose oxidase (EC. 1.1.3.4) from Afgergillus niger
(type X-S, with a specific activity of 150 U mg-’ solid) and ~-lactose (E.G. 3.5.2.6) from 3u~~i~s cereus (type I, with a specific activity of -2000 U mg-’ solid) were obtained from Sigma (St. Louis, MO, USA). Catalase (B.C. I. 11.1.6, from beef liver) was purchased from Boehringer (Mannheim, Germany). Glutaraldehyde was supplied by Sigma. y-aminopropyltriethoxysilane was purchased from Pierce Eurochemie, (Rotterdam, The Netherlands). Controlled-Pore Glass (particle diameter: loo-160 pm, pore diameter: 50 nm) was obtained from Veb Trisola (Steinach, Germany). Microbead thermistors (MI%-3) were supplied by Wuhan Electronic Component Factory (Wuhan, China). A d.c.-coupled ~ea~tone bridge ~~nst~cted at our institute) was used for ~rn~ra~re m~s~~ents. At ma~mum s~itivity, the signal changes 100 mV per 10m3“C. Besides the peristaltic pump (Alitea pump unit C-4V from Ventur Tekniska AB, Sweden), a perfusion pump (B. Braun Melsungen AG, Germany) and HPLC valves suitable for I ~1 and more than 5 pl samples, (type C14W and C6W, VIGI AG Valco Europe, Switzerland), respectively, were used. Preparation of the enzyme colwnm
Controlled-Pore Glass (CPG), on which enzymes were covalently immobilized, was used as support material of the enzyme columns. The i~ob~zation procedures were Performed as follows. CPG with a particle
@ 1993- Ekevier Sequoia.All rightsmewed
444
size distribution of lOO- 160 pm was first boiled in 5% HNO, for 45 min and was extensively washed with distilled water to clean and hydrate the surface. To a solution of 2 g of y-aminopropyltriethoxysilane in 18 ml of water was added 1 g of clean glass. The pH was adjusted to 3.5 with 6 M HCl and the solution was heated at 75 “C for 3 h with occasional gently stirring. The glass was subsequently collected on a Biichner funnel, washed with distilled water and dried for at least 4 h at 115 “C. To 1 g of alkylamino glass was added 25 ml of 2.5% glutaraldehyde in 0.1 M sodium phosphate buffer adjusted to pH 7.0. The reaction was allowed to take place under reduced pressure for 30 min (using a water aspirator) and then at atmospheric pressure for at least 30 min. The brick-red coloured product was washed exhaustively on a Btichner funnel with distilled water and stored in water in the refrigerator. The enzyme was added to the activated-glass suspension in as small a volume as possible, usually in 0.1 M sodium phosphate buffer, pH 7.0. The protein concentration should preferably be at least l%, i.e., between 50 and 100 mg of enzyme should be used per g of carrier. For glucose oxidase, coimmobilization of catalase was carried out. The coupling was often performed at room temperature and was allowed to proceed for 1 to 4 h with gentle mixing or in the cold overnight. The enzyme preparation was thoroughly washed with buffer prior to use. Constructions Compact sensor
were mounted on the capillaries with a heat-conducting epoxy. Plastic chip sensor
The chip sensor (27 mm x 7 mm x 6 mm) was constructed of Plexiglass (see Fig. 2). The rectangular enzyme cell (5 mm x 3 mm x 0.5 mm) and part of the inlet and outlet were milled into the plastic base to a depth of 0.5 mm. Electrically-insulated thermistors in direct contact with flow stream were placed outside the enzyme cell after the porous polyethylene filters. The enzyme cell was charged with enzyme preparation prior to compaction. Replacement of enzymes is possible due to ready access to the enzyme compartment. Microcolumn sensor
The microcolumn sensor (outer diameter x length: 0.8 mm x 15 mm) was constructed of stainless-steel tubing and free from auxiliary components (Fig. 3). A microbead measurement thermistor was directly mounted on the outer surface at the latter part of the column using heat-conducting epoxy and the reference thermistor was positioned on a piece of the inlet metal tubing. Both the length (about 200 mm) and inner diameter (0.15 mm) of the inlet tubing between the sample valve and the column were minimized in order to reduce sample dispersion during the transportation in
Thermistor
Inlet
The construction of a compact sensor of greatly-reduced dimensions (outer diameter x length: 36 mm x 46 mm) is shown in Fig. 1. In order to conveniently accommodate enzyme columns and to ensure isolation from ambient temperature fluctuations, a cylindrical copper heat sink was included. An outer Delhi jacket further improved the insulation. The enzyme column (outer diameter x length: 4 mm x 3 mm), constructed of Delrin, was held tightly against the inner terminals of the copper core. Short pieces of well-insulated gold capillaries (outer diameter/inner diameter: 0.3 mm/ 0.2 mm) were placed next to the enzyme column as temperature-sensitive elements. Microbead thermistors
Filters
Outlet
Enzyme’cell (a) Enzyme
Covering
ccl ThermiYr
\
Ff”er
/ liim
(b) Enzyme
Fig. 2. (a) Schematic diagram and (b) cross section of the plastic chip thermal biosensor.
column
Inlet
Enzymecolumn
\
Thermistor
Outlet
/ 5 mnl
1
H
cm
Fig. 1. Schematic cross section of the compact biosensor.
Stainless steel tubing
Nylon tubing
Fig. 3. Schematic diagram of the microcolumn sensor.
445
the flow system. Similarly, to reduce heat loss, simple and short connections between the column and the inlet or outlet tubings were required.
Results Effect
and diiussion of miniaturization
on properties
The sensitivity and response process of calorimetricFIA biosensors can be explicitly depicted by their steady-state responses. Flow rates of 108, 75 and 40 nl min-’ were used for the compact sensor, the chip sensor and the microcolumn sensor, respectively. Buffer and penicillin-V (10 mM) were alternatively introduced into the /I-lactamase columns (20 min cycle) using a peristaltic pump. Steady-state response curves corresponding to different degrees of miniaturization of the sensors were obtained and compared in Fig. 4. The height of the curves is governed by the equilibrium of exothermic reactions and heat leaking from the reactive systems. The former depends on the catalytic efficiency of enzymic reactions in the columns, and the latter is dominated by adiabatic properties of the reactors. In general, the greater the amount of enzyme, the higher the catalytic efficiency. Miniaturization necessitates a limited amount of enzymes. Decrease of the enzyme amount could result in decrease of the linear range of the measurement unless flow rate and sample volume are carefully controlled. This was demonstrated in an experiment. Using 20 ul of penicillin-V, the linear range of measurement was limited to 100 mM, 50 mM to 40 mM in correspondence with the volume of the enzyme reactor, which was 21,8 and 4 nl with the compact, the chip and the microcolumn fl-lactamase sensors, respectively. In the latter case, the heat conductivity of the materials
and the heat capacity of the enzyme columns and their surroundings are the major contributory factors. A smaller column may therefore lead to higher sensitivity if a good adiabatic condition can be maintained. This is demonstrated in Fig. 4. Nevertheless, the slope is an exception. It is not directly related to the degree of miniaturization, as with the chip sensor. Larger heat capacity of the enzyme-column surroundings may be the cause of the increased response time, because the reactor bed of the chip sensor acts as both enzyme support and insulating layer with significant heat capacity. The experimental results revealed that a thin-walled metal column surrounded by a layer of air can offer a faster response and much higher sensitivity than plastic columns. For these construction-dependent biosensors, further miniaturization will necessitate the use of smaller enzyme volume and finer carrier materials. This will require a new technology, such as micromachining [8], or more sensitive transducers. Effect
of sample
volume
on properties
One of the purposes of miniaturization is to reduce the sample volume necessary for analysis, which would be of benefit in cases where only limited sample volumes are available. Furthermore, as a reduction of the sample volume increases the linear range, as demonstrated by the conventional enzyme thermistor [I], we anticipated a similar result with the miniaturized glucose oxidase sensors, although at the expense of lower sensitivity. The result with the chip sensor, using different volumes of 1, 5 and 20 ~1, is shown in Fig. 5. The linear range of standard glucose concentrations extended from 8 mM up to 100mM when the sample volume was reduced from 20 ul to 1 ~1. A similar result was also obtained with the microcolumn sensor, although the linear range was limited to 30 mM. The extension of the linear range when reducing the sample volume might be attributed to a relative increase
1
9 0
_
‘Ix____ 0
2
4
6
8
10
0 5 G 12
Time (mins) Fig. 4. Comparison in steady-state responses between the compact (M-l), the plastic chip (M-Z) and the microcolumn (M-3) sensors charged with immobilized fl-lactamase. Penicillin-V samples (10 mM) were pumped with their corresponding optimal flow rates of 108, 75 and 40 pi min-I, respectively.
d
2b
4b
sb
Glucose Cont.
100
(mM)
Fig. 5. Effect of sample volume on the linear range of the chip sensor loaded with immobilized glucose oxidase. Flow rate of 75 11 min-’ was used.
446
in the dispersion of the small sample volume. This factor results in dilution of the sample in the flowinjection analysis (FIA) system [9], particularly in the column. Since reactive processes in an enzyme column are determined by sample volume, a smaller volume should yield a fast response time. In practice, however, this is not always observed due to dispersion of sample between the sample inlet and the column [lo]. Consequently, the response time was not proportional to the sample volume. If the inlet is small in inner diameter and short in flow path (inner diameter x length: 0.15 mm x 200 mm), the measurement period per sample as fast as 30 s (glucose or penicillin-V of 10 mM) was still achieved with the microcolumn biosensor in 1 pl sample volume and a flow rate of 50 pl/min. Effect ofjow rate on properties
In calorimetric FIA biosensors, many properties, apart from the factors discussed above, are subject to the flow rate. The improvement by means of this way is much concerned with response time and linearity, but has less effect on the sensitivity. Too high a flow rate, however, may cause a reduced peak height due to the limited time available for catalytic reactions as shown by varying the flow rate using the chip /?-lactamase sensor. The results in Fig. 6 were obtained by just changing the flow rate instead of the different dimensions of the miniaturized biosensors. It was judged that the flow rate should be less than 100 pl/min in spite of the enlarged response time. The influence of the flow rate on calorimetric FIA biosensors may be characterized by a factor, Q, which is the ratio of an optimal flow rate (FJ to a matrix volume of the enzyme column (V,), such that Q = FJV,. For
TABLE 1. Comparison of characteristic factor Q (min-I) of caloFimetric FIA biosensors: ET, the enzyme thermistor; M- 1, the compact sensor; M-2, the plastic chip sensor; M-3, the microcolumn sensor
type
Column
Column volume (11)
Optimal flow rate (~1 min-‘)
Characteristic factor Q (min-‘)
ET M-l M-2 M-3
200 20 10 4
1000 100 75 20
5 5 1.5 5
a constant column volume, a smaller Q (min-I) value indicates a higher temperature-conversion efficiency of the thermal biosensor. The comparison of characteristic factors between the conventional enzyme thermistor (ET) and the miniaturized thermal biosensors is shown in Table 1. The results reveal that the characteristic factor Q does not become smaller with miniaturization of the sensors. Thus, the miniaturization process does not improve the temperature-conversion efficiency of the thermal biosensors, although some propertie%,such as sensitivity, response time and linear range have been improved. That the chip sensor presents a bigger Q value than others implies that the characteristic factor Q could be also related to the adiabatic situation of the enzyme column, which is less favorable here.
Life time and operational stability
For the chip and the microcolumn sensors, the enzyme reactors are preferably unified with the transducer parts in order to reduce sizes. Therefore, the replacement of enzyme preparations is not as convenient as for easily-disassembled systems. The employment of disposable transducers, e.g. thin-film thermistors [ 111, would facilitate practical operation. Storage of these miniaturized devices must be carefully undertaken. The cells must be filled with buffer solution between experi-
50
40
0 ,"
30
0 =.
20
G 10
Time
(mins)
Fig. 6. Effect of flow rate on sensitivity and response time of steady-state responses. The plastic chip sensor loaded with immobilized fl-lactamase and penicillin-V of IO mM were used: (a) 54 filmin-‘; (b) 105 ~1mix’; (c) 75 fl mix’.
0
100
200
Sample Number Fig. 7. Reproducibility o$ penicillin-G (10 mM) using the plastic chip sensor loaded with /Xxctamase.
447
ments and the systems kept at 4 “C. Under these conditions, using /?-lactamase and glucose oxidase, a life time of more than one month was obtained with the above three sensors. Investigation of the operational stability of these miniaturized biosensors is essential. The stability was tested by injecting more than 200 penicillin-G samples into the chip /?-lactamase sensor. The relative standard deviation (1%) was obtained. Similarly, using the microcolumn glucose oxidase sensor, this value was 6% on 70 glucose samples of 10 mM. Figure 7 shows the reproducibility of the chip sensor for the 200 assays.
Conclusions
Three different types of miniaturized calorimetric FIA biosensors have been developed and studied. Individual advantages are presented from various aspects. The compact sensor, as a version of the conventional enzyme thermistor, retains the same functions, and offers great flexibility in the range of possible applications. The chip sensor is readily constructed and is capable of easy and accurate determinations. The microcolumn sensor presents a fast response time and a higher sensitivity. Miniaturization of thermal biosensors, due to increase of the sensitivity, results in the possible use for such a small sample volume of 1 ~1. This makes it possible to expand the linear range of glucose (from 8 mM to at least 30 mM) and to reduce the measurement time (from more than 1 min down to 30 s). The large linear range possible for glucose determination is of particular interest in diabetes control where the glucose concentration is up to 20mM. Consequently, this method provides a convenient way to directly monitor the blood glucose without the tedious process of sample dilution.
Acknowledgement
This work has been partly supported by a grant from the National Swedish Board for Technical Developments (NUTEK).
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