WO1993003356A1

WO1993003356A1 - Continuous-flow condenser cell for use as a biosensor for continuous qualitative and quantitative analysis of analytes

Info

Publication number: WO1993003356A1
Application number: PCT/EP1992/001680
Authority: WO
Inventors: Andreas Gebbert; Manuel Alvarez-Icaza
Original assignee: GESELLSCHAFT FüR BIOTECHNOLOGISCHE FORSCHUNG MBH (GBF)
Priority date: 1991-08-07
Filing date: 1992-07-23
Publication date: 1993-02-18
Also published as: DE4126177A1

Abstract

The invention provides a novel analytical condenser measurement cell which gives good results when used as a biosensor. The novel biosensor is based on a tantalum substrate (1) on which is applied a composite insulating film comprising several components and providing a sufficiently thin dielectric. Constant monitoring of the film thickness during the production process ensures very high film-thickness reproducibility. Other improvements with respect to the measurement frequency and the deposition voltage ensure that the instrument gives excellent assay results.

Description

FLOW CAPACITOR CELL FOR USE AS A BIOSENSOR

CONTINUOUS, QUALITATIVE AND QUANTITATIVE DETERMINATION

FROM ANALYZES ___________________________________________________________________________________________________

The present invention relates to a flow condenser cell for use as a biosensor for the continuous, qualitative and quantitative determination of analytes, in particular with a measurement of the capacitance of the flow condenser cell as a function of time, in order thereby to draw conclusions about the analyte concentration of the sample liquid.

Such devices are known in the prior art from the international patent application WO 88/09499, the German published application DE 39 23 420 A1. A decisive disadvantage of these known arrangements is the dependence of the measurement signal on the entire dielectric medium between the electrodes. As a rule, a reference capacitor that compensates for non-specific medium influences (conductivity, gas content) must be added to the measuring system. In recent years, the development of biosensors in the field of research has been strongly promoted, especially in the targeted detection of certain substances from a complex sample matrix, the use of biosensors is becoming increasingly important. Biomolecules (eg enzymes, antibodies, nucleic acids or entire organisms) react specifically with the analyte to be detected. The parameters which change during the reaction of the biological components with the analyte are converted into recordable signals for signal processing via a transducer (eg electrodes, field effect transistors, photometers or glass fiber optics).

Biosensors can e.g. as immersion electrodes for single use determinations or also within a flow system for quasi-continuous determination of the analyte.

In contrast to automated enzymatic analysis, so-called immunoassays, i.e. an antibody specifically binds the substance to be detected, and has only been working with quasi-continuous flow systems since the early 1980s. Immunoassays can generally be classified according to two detection methods; the indirect as well as the direct detection of the analytes.

In the indirect methods, the ligand-analyte binding does not cause a measurable signal. Here, marked analytical analogs (so-called conjugates) must be used in a competitive process. This leads to additional incubation and rinsing steps for conjugates and substrates and thus to more equipment and longer analysis times. However, one advantage of indirect assays is signal amplification experiences the measuring system, for example through the use of enzyme markers.

With the direct measurement methods, the binding of the analyte already causes a measurable electrochemical, optical or electrical signal. Piezoelectric, optical or photometric transducers are mainly used. There is no time difference between analyte binding and signal acquisition, and the binding behavior can be recorded continuously (on-line).

The measurement method of using a capacitor as a transducer in direct immunoassay is the subject of the present invention and will be described in more detail below.

The principle is that a voltage is applied to two insulated plates lying opposite each other so that energy is stored in the form of a static charge. The size of the electrical capacitance of the system can be influenced by changing the dielectric properties of the medium or by varying the geometry of the capacitor that is produced in this way. If, for example, antibodies are immobilized on the capacitor plates, the connection of analytes leads to a change in the electric field between the electrode plates, which can be designed as a pair of comb electrodes in the open capacitor.

Since the reactions of the immobilized antibodies with the analyte only take place in the relatively small surface area, arrangements were sought which react specifically to changes in this area. The development of a transducer based on capacitive measurements is therefore aimed at Application of so-called MIE structures (metal-insulator-electrolyte structure). In the MIE structure, an electrically conductive electrolyte solution serves as the second electrode for the measuring system. The dielectric is formed by the insulation on the metal plate. If antibodies / antigens are bound to the dielectric, the electrical properties change and the layer thickness of the dielectric increases. So far, mainly semiconductor materials have been used as the substrate for the dielectric, on which, for example, silicon oxide or organic polymers have been applied as an insulation layer.

From the above, there is a clear need to manufacture biosensors whose technical outlay is reasonable and which is also inexpensive.

It is therefore the object of the present invention to provide a flow-through capacitor cell as a biosensor, in which the increase in the layer thickness when the analyte is bound, the dielectric of the capacitor can be specifically and reproducibly optimized.

This object is achieved by the features in the characterizing part of claim 1, which are characterized in that the working electrode consists of tantalum, the dielectric is composed of tantalum oxide anodically deposited on the working electrode and an immobilization layer, the counter electrode consists of electrically conductive sample medium and the contacting of the counter electrode is made with a platinum foil. For the production of the device described above for varying the layer thickness of a dielectric using a formation method for applying an oxide layer on a suitable substrate surface (metal surface) and the immobilization of analyte-specific ligands on the oxide layer, a method has been found to be particularly favorable in which a tantalum substrate is used a tantalum oxide layer anodically deposited from the tantalum is applied in 4% boric acid at voltages between 1-250 V using a hydrofluoric acid to etch the surface and then the ligands for specific analytes are immobilized using known covalent binding methods.

When executing the device according to the invention, it is particularly important and advantageous to apply a pear-shaped spacer to the tantalum / tantalum oxide electrode. The diluted part of the pear-shaped spacer must be directed in the direction of flow of the liquid to be analyzed. It is also important with this spacer that there are no sharp-angled corners in the flow area which facilitate the formation of air bubbles.

A decisive advantage of the present invention is the fact that tantalum should be selected as the base electrode in order to be able to use anodic deposition to produce tantalum oxide layers that are smaller than 100 μm and in particular smaller than 5 nm.

A further advantage of the device according to the invention has been found to apply the sensor core, which is composed of the tantalum electrode and the tantalum oxide layer, to an essentially known four-port circuit, around which it witnessed to be able to optimally supply the electrical signal for further processing.

In the method for producing the layer thickness of said dielectric, it has proven to be advantageous to etch the tantalum surface with hydrofluoric acid for about one minute in order to enlarge the surface of the tantalum. The subsequent anodic deposition in 4% boric acid at small maximum voltages between 1-250 V DC is also of particular importance for the manufacture of the device according to the invention.

When forming the tantalum oxide layer on the tantalum, it is also advantageous to measure the capacitance and the forracing voltage simultaneously during the formation of the layers. It is also extremely important and advantageous for the reproducibility of the layer thicknesses of the dielectric to measure and control the layer thicknesses of the dielectric by means of interference spectra during the production process.

Ligands, i. E., Can then be applied to the resulting tantalum oxide layer using known covalent binding methods. Specific binding partners for the analytes to be analyzed are immobilized.

Further advantageous refinements can be found in the subclaims and the description of the exemplary embodiments.

The invention is explained in detail with reference to the drawings. Show it. 1a shows a schematic representation of the flow capacitor cell according to the invention;

Fig. 1b is a schematic representation of an inventive

Spacers;

2a shows a schematic representation of the beam path during the interference measurement for determining the layer thickness of the tantalum oxide;

2b shows an example of an interference spectrum;

Fig. 3 is a graphical representation in which the voltage for

Propagation of the tantalum oxide layer, the thickness of the oxide layer and the duration of a period as a function of time for forming the layer thickness is shown;

4 shows a graphical representation of the period in sec as a function of the thickness of the oxide layer in nm;

5 shows a graphical representation of the change in capacitance in percent as a function of the frequency in kHz of the applied AC voltage;

6 is a graphical representation of the capacitance in nF and the change in capacitance in pF as a function of time in seconds;

Fig. 7 is a graphical representation of the change in capacitance in pF as a function of time in seconds for different sensors, the dotted line one Sensor without immobilized antibody, the solid line is a sensor with immobilized BSA, the broken line is a sensor with immobilized mouse IgG for heated samples of SAM-GAL and the second solid line is sensors with immobilized mouse IgG unheated SAM-GAL;

8a is a graphical representation of the change in capacitance in pF as a function of time in seconds;

8b shows the change in capacitance in pF as a function of

SAM-GAL in ng / ml;

Fig. 9 shows the change in capacitance in pF as a function of

SAS-GAL IgG in ng / ml;

Fig. 10 shows the change in capacitance in pF as a function of

Time in seconds, the dashed line representing an incubation time of 10 minutes and the solid line representing an incubation time of 20 minutes;

11 is a graphical representation of the change in capacity in ppm as a function of time in seconds, the broken line representing GAR IgG (20 ng / ml) and the solid line representing SAM-GAL (20 ng / ml).

Fig. 1 shows a schematic representation of the flow capacitor cell according to the invention with its essential components. The tantalum / tantalum oxide electrode 1, 2 forms the sensor core. The tantalum oxide layer 2 provides layer 2 represents the basis for the dielectric of the capacitor. In general, two types of dielectric can be distinguished; the so-called barrier type and the porous type. The barrier type is thin and closed in structure, while the porous type consists essentially of a double layer, ie an inner thin layer which is closed and an outer thick layer which is porous. The thickness of the barrier type is dependent on the maximum voltage which is applied to the conductive substrate when the layer is formed, while the thickness of the porous type is essentially dependent on the current density, the formation time and the electrolyte temperature.

For the electrolytes, boric acid or citric acid are generally used in the barrier type, while either sulfuric acid, chromic acid or oxalic acid are used in the porous type. The known aluminum electrolytic capacitors, depending on their quality, are a mixture of the types described and are not suitable as a basis for biosensors.

In the present case of the flow-through capacitor cell, dilute boric acid was used to form the tantalum oxide on the tantalum substrate. The maximum voltages for anodic deposition on the substrate were relatively small. In detail, tantalum 1 with a purity of 99.9+% was used, which was etched in hydrofluoric acid (48%) for about one minute. This leads to cleaning and an enlargement of the tantalum surface. The tantalum oxide is then anodically deposited in 4% boric acid on the tantalum substrate at maximum voltages of 1-250 V. The decisive factor in the construction of the flow capacitor cell is the fact che that extremely homogeneous, thin layers can be produced with the dielectric forming method used. These thin layers are extremely important for the production of a biosensor, since the capacitance is, among other things, a function of the inverse layer thickness.

In the present case, the dielectric is formed by the insulation on the metal plate. If antibodies / antigens are bound to the dielectric, this leads to a change in the layer thickness and thus to a change in the dielectric and ultimately to a change in the electrical properties of the capacitor.

Electrically speaking, the immobilization of antibodies to the tantalum oxide represents an additional dielectric layer, which influences the overall capacity of the system. This results as a series connection of the capacitance of the tantalum oxide layer (Ct) and the immobilization layer (Cb) to:

When the analyte, ie the medium to be analyzed, is bound to the immobilization matrix, a further dielectric layer (Cg) is assigned to the system. This results in the overall shape:

The total capacity (Cf) can be calculated from this. This shows that the total capacity is a composition consists of several partial capacities, with the analyte layer (Cg) playing the decisive role in the present case. In other words, this means that the total capacitance is maximum with a minimum tantalum oxide layer thickness, while conversely with a relatively large layer thickness of the dielectric, the sensitivity of the measuring system becomes minimal. Ultimately, for the production of a biosensor, this means that the layer thickness of the dielectric has to be made optimal, ie as thin as possible, in order to achieve the greatest possible capacitance.

For this it is absolutely necessary to ensure the reproducibility and determinability of the layer thickness. The layer thickness can be followed by means of coherent light waves by reflection, refraction or diffraction on thin layers of solid bodies by means of interference spectra and can be measured precisely. The thickness of a homogeneous, thin tantalum oxide layer 2 on tantalum 1 can thus be determined optically by recording interference spectra. In principle, this is shown schematically in FIG. 2a. In the sensors manufactured, the capacitance and the formation voltage were also measured during the formation of the tantalum oxide. At the same time, the interference spectra of the respective layers were recorded by means of spectral analysis and the layers of the dielectric were determined, as shown in FIG. 3.

The capacity increases due to the increasing

Layer thickness in the course of formation, since the capacity is inversely proportional to the layer thickness. In the present case, the period of a charge and discharge was measured. If one looks at the relationship between layer thickness and capacity in the area of short formation times, one learns that at small layer thicknesses (up to approx. 100 nm) there is an increase in the layer thickness by 10 nm and there is a significant decrease in the capacitance. This is shown in FIG. 4. 10 nm is roughly the size range of the antigen binding partner.

The ability to determine the layer thickness and thus the possibility of producing sensors in a reproducible manner and the high sensitivity of the system to minimal changes in layer thickness are the prerequisites for the use of the dielectrics as the basis for the biosensors of the present invention.

Both properties are due to the fact that extremely thin and homogeneous layers of tantalum oxide, which are smaller than 5 nm, could be produced on the substrate tantalum. This has not been possible with other materials of conventional electrolytic capacitors, such as aluminum.

On the sensor core described above, which results from the

Base substrate tantalum 1 and the thin tantalum oxide layer 2 located thereon, a spacer 5 is placed during manufacture. The shape of the spacer was chosen pear-like and plays an important role in avoiding the formation of air bubbles in the flow capacitor cell. The sample liquid 3 to be examined is passed through the spacer at an angle. The sample liquid 3 is supplied at the thickened part and the sample liquid is removed at the tapered part of the spacer 5. This leads to an increasing flow velocity within the flow cell, and existing air bubbles are flushed out of the system.

The sample liquid 3 to be examined serves as a counter electrode. A platinum foil 4 is placed in the spacer as contact to the counter electrode, so that good electrical contact is established. The elements of the sensor described are pressed together by two fixed plates 8, which consist of electrically non-conductive materials, such as Teflon. The tantalum / tantalum oxide electrode and the contacting of the counter electrode with the platinum foil 4 are connected for precision measurement in a known four-port circuit 6 via coaxial cables to the measuring instrument. The entire sensor is connected with its metal jacket 7 to the measuring instrument and thus electrically shielded.

In order to selectively design the surface of the biosensor for the analyte to be determined, the binding partner of the substance to be determined is immobilized on the tantalum oxide.

Common covalent binding methods, such as immobilization via silanization / glutaraldehyde activation or immobilization via carbodiimide, are used in the literature for this purpose.

With the dielectric fabricated in this way, as shown in principle in FIG. 1, sample measurements were carried out which gave results which were not currently achieved with conventional biosensors.

The measurements were carried out in the following way: - The flow condenser was flushed with a working buffer, for which a measured value I was measured;

- Then it was switched to the sample liquid to be examined and flowed through the flow condenser for a certain time by the analyte;

- then it was switched again to the working buffer, which gave a measured value II;

- the biosensor was then rinsed with elution buffer, e.g. Glycine / HCL, regenerated;

- Which finally led to the measurement result that represents the difference in capacitance between the measured value I and measured value II.

The first three steps were carried out automatically. As an alternative to this measurement method, the sample liquid can also be injected into the working buffer.

Furthermore, it was possible to show with the flow capacitor cell described that there is a frequency optimum in the detection of biological materials at which the capacitance measurements should be carried out. To determine this optimum, the changes in capacity for a specific analyte are considered as a function of the measurement frequency. The further measurements are carried out at the optimal frequency for the analyte.

A possible drift in the baseline when flushing the flow condenser with working buffer can help the calculation of the capacity differences can be compensated arithmetically.

In the following, measurement examples are given which are commented on individually.

Example I:

5 shows a frequency optimization for the analyte

Sheep anti-mouse IgG-β-galactosidase (SAM-GAL) is shown.

Mouse IgG were immobilized on the sensor surface of the flow capacitor described using the silanization / carbodiimide method. The change in capacity was then measured in the manner described after 10 minutes of incubation with sheep anti-mouse IgG-β-galactosidase (4 Ig / ml) at measuring frequencies from 100 to 20,000 Hz.

The frequency optimum for the analyte sheep anti-mouse galactosidase IgG was found at 1000 Hz. Subsequent measurements of the ligand analyte behavior were measured at 1000 Hz.

Example II:

FIG. 6 shows the detection of sheep anti-mouse IgG-β-galactosidase (SAM-GAL) (20 ng / ml).

The dielectric of the sensor surface was formed in the way described of 250 V. Mouse IgG was immobilized on the sensor surface and the presence of sheep anti-mouse galactosidase (20 ng / ml) was detected in the sample liquid. Example III

7 shows the selectivity of the sensor surface.

In order to clarify the selectivity of the sensor surface for the analyte to be detected, a) the sensor surface was made with analyte-unspecific protein

(BSA) and b) analyte denatured with analyte-specific protein (mouse IgG) measured by heat (15 min, 100 ° C.).

It turned out that the sensors on which k a. or non-specific protein had been immobilized, gave small measurement signals when measuring undenatured analyte.

The sensor surface specifically designed for the sheep-anti-mouse IgG analyte did not react with the heat-denatured analyte, but did show a significant measurement signal in the undenatured sample.

Example IV:

Figures 8a and 8b show the concentration dependence of the measurement signal.

In order to be used in sample analysis, the flow-through capacitor must supply measurement signals which are dependent on the concentration of the analyte. As an example, mouse IgG was immobilized on a dielectric formatted to 250 V and various concentrations of sheep anti-mouse IgG-β-galactosidase were measured. This showed a signi fictitious dependence of the measurement signal on the sample concentration.

Example V

9 shows the dependence of the sensitivity on the

Layer thickness of the dielectric.

As explained above, the sensitivity and thus the change in capacitance should decrease as the layer thickness, i.e. with smaller formation tension of the tantalum oxide. As an example, measurement tests were carried out analogously to those under Example IV, but with dielectrics which were formed at 5 or 3 V. It turns out that the smaller the layer thickness of the tantalum oxide on the tantalum, the greater the change in capacitance, which was to be expected.

Example VI:

10 shows the time dependence of the measurement signal.

The measurement signal should depend on the incubation time of the analyte with the sensor surface. In one example, mouse IgG was immobilized on the sensor surface and goat anti-mouse IgG was detected. The incubation times were 10 and 20 minutes, respectively. It was found that doubling the incubation time of the sample doubled the measurement signal.

Example VII:

Figure 11 shows the uptake of various analytes.

An important advantage of using direct immunological analysis methods is that no labeled binding Flow condenser rabbit IgG immobilized and goat anti-rabbit galactosidase and sheep anti-rabbit measured.

It can be seen from the examples presented that the dielectric presented has succeeded in characterizing the ligand-analyte behavior directly. In addition to the dependence of the measurement signal on the analyte concentration, the applicability of the measurement system to various analytes could be shown. The dielectric described above appears particularly suitable, for example, in continuous analysis (on-line analysis) of biotechnical processes or as an early warning system for pesticides in drinking water analysis.

In conclusion, it should be noted that tests with aluminum substrates were carried out in a manner similar to that described with the capacitor flow cell described above, which demonstrated the physical properties and the principle suitability of the capacitor cell as a transducer for biosensors. However, the quality of the resulting dielectric was relatively poor, so that no analyte measurements could be carried out.

On the other hand, homogeneous, thin layers of tantalum oxide could be successfully deposited on the tantalum substrate, thus producing high-quality dielectrics. The reproducible manufacture of the sensors was demonstrated by determining the layer thickness using interference spectral analysis.

After testing various immobilization methods on the tantalum oxide surface, the binding of various analytes to the mobilized ligands was continuously characterized. As a detection limit for eg anti-mouse IgG-ß Galactosidase was determined at 1 ng / ml IgG at an incubation time of 10 minutes. Goat anti-mouse IgG could be detected from 10 ng / ml.

All measurements were carried out with a commercially available precision capacitance meter. The measuring instrument was connected to a computer via an interface so that the course of the sample analysis could be followed simultaneously on the computer monitor and automatic data acquisition was ensured.

Claims

PATENT CLAIMS

1.Device for varying the layer thickness of a dielectric on a suitable substrate surface (1) using a formation method for applying an oxide layer (2) and interference spectra of analyte-specific ligands on the oxide layer (2), that means that

- The working electrode (1) consists of tantalum;

- The dielectric is composed of tantalum oxide layer (2) analytically deposited on the working electrode (1) and an immobilization layer composed of ligands and analytes;

- The counter electrode (3) consists of electrically conductive sample medium;

- The contacting of the counter electrode (3) with a platinum foil (4) is made.

2. Device according to claim 1, dadurchgekennzeic hn et that a pear-shaped spacer (5) is placed on the tantalum / tantalum oxide electrode, the diluted part of the bulb pointing in the direction of flow.

3. Device according to claim 1 and 2, d a d u r c h g e k e n n z e i c h n e t that the spacer (5) is angular in the flow region.

4. The device of claim 1, d a d u r c h g e k e n n z e i c h n e t that the tantalum oxide layer (2) on the tantalum (1) is less than 100 nm and in particular less than 5 nm and preferably homogeneously distributed.

Tantalum (1) is less than 5 nm and homogeneously distributed.

5. Apparatus according to claim 1, that the shielding and the connection of the sensor core (1) and (2) to a four-port circuit (6).

6. Device according to one of the preceding claims, ie that the sample liquid to be examined is used as the counter electrode (3).

7. Process for producing a dielectric in one

Apparatus according to claim 1, characterized in that a tantalum substrate (1) is deposited with a tantalum oxide layer (2) anodically deposited from the tantalum in 4% boric acid at voltages between 1-250 V using a hydrofluoric acid to etch the surface and then the Ligands for specific analytes can be immobilized using constant covalent binding methods.

8. The method according to claim 7, characterized in that the tantalum surface (1) is etched for about one minute with hydrofluoric acid (48%).

9. The method of claim 7 and 8, d a d u r c h g e k e nn z e i chn e t that on the etched

Tantalum surface (1) is deposited a tantalum oxide layer (2) smaller than 5 nm anodically in 4% boric acid at voltages between 1 to 250 V DC.

10. The method of claim 7, dadu r c h g e ke nn z e i c h n e t that during the formation of the

Tantalum oxide layer (2) the capacity and the

Forming voltage on the oxide layer is measured simultaneously.

11. The method according to claim 7, so that the layer thicknesses of the dielectric by means of interference spectra during the manufacture of the device described in claim 1

can be measured simultaneously.

12. The method according to claim 1, so that the binding partners (ligands) by means of known covalent binding methods

be immobilized.

13. The method of the device produced according to claims 1 and 2 as a biosensor (flow condenser), with which biological processes are continuously measured in order to directly characterize the ligand-analyte behavior.