Hydrophilic, swellable coatings for biosensors

Details Hydrophilic, swellable coatings for biosensors Hydrophilic, swellable coatings for biosensors

2002-08-06

2004-08-31

Cooney, Jr., John M. (Department: 1711)

Synthetic resins or natural rubbers -- part of the class 520 ser

Synthetic resins

From reactant having at least one -n=c=x group as well as...

C427S002120, C427S058000, C428S423100, C428S426000, C528S085000

Reexamination Certificate

active

06784274

ABSTRACT:

BACKGROUND OF THE INVENTION
1. Field of the Invention
This invention lies in the field of polymer chemistry in which the polymers formed are suitable for coating biosensors. The coatings act to decrease the impedance at the sensor's electrode and thereby enhance the signal during in vivo placement of the sensor.
2. Description of Related Art
Biosensors are small devices that use biological recognition properties for selective analysis of various analytes or biomolecules. Typically, the sensor will produce a signal that is quantitatively related to the concentration of the analyte. To achieve a quantitative signal, a recognition molecule or combination of molecules is often immobilized at a suitable transducer which converts the biological recognition event into a quantitative response.
A variety of biosensors have been developed for use with numerous analytes. Electroenzymatic biosensors use enzymes to convert a concentration of analyte to an electrical signal. Immunological biosensors rely on molecular recognition of an analyte by, for example, antibodies. Chemoreceptor biosensors use chemoreceptor arrays such as those of the olfactory system or nerve fibers from the antennules of the blue crab
Callinectes sapidus
to detect the presence of amino acids in concentrations as low as 10
−9
M. For a review of some of the operating principles of biosensors, see Bergveld, et al., ADVANCES IN BIOSENSORS, Supplement 1, p. 31-91, Turner ed., and Collison, et al., Anal. Chem 62:425-437 (1990).
Regardless of the type of biosensor, each must possess certain properties to function in vivo and provide an adequate signal. First, the elements of the biosensor must be compatible with the tissue to which it is attached and be adequately shielded from adjacent tissues such that allergic or toxic effects are not exerted. Further, the sensor should be shielded from the environment to control drift in the generated signal. Finally, the sensor should accurately measure the analyte in the presence of proteins, electrolytes and medications which may interfere.
The prototype biosensor is the amperometric glucose sensor. There are several reasons for the wide ranging interest in glucose sensors. In the healthcare arena, glucose sensors are useful for glucose monitoring of patients with diabetes mellitus. Additionally, a working glucose sensor is required for the development of a closed loop artificial pancreas with an implanted insulin pump. A commercial interest focuses on sensors that can be used to monitor fermentation reactions in the biotechnology arena. From a scientific standpoint, interest is driven by the availability of a very robust enzyme, glucose oxidase, which can be used to monitor glucose, as well as the desire to develop model sensors for a wide variety of analytes.
Any amperometric glucose sensor or any oxido-reductase enzyme that uses O
2
as a co-substrate and is designed for subcutaneous or intravenous use requires both an outer membrane and an anti-interference membrane. The requirement of two distinct membranes is due to the fundamental nature of the sensor as well as the environment in which the measurement is made.
A glucose sensor works according to the following chemical reaction (Equation 1):
In this reaction, glucose reacts with oxygen in the presence of glucose oxidase (GOX) to form gluconolactone and hydrogen peroxide. The gluconolactone further reacts with water to hydrolyze the lactone ring and produce gluconic acid. The H
2
O
2
reacts electrochemically as shown below (Equation 2):
H
2
O
2
→O
2
+2
e
−
2H
+
  (II)
The current measured by the sensor/potentiostat (+0.5 to +0.7 v oxidation at Pt black electrode) is due to the two electrons generated by the oxidation of the H
2
O
2
. Alternatively, one can measure the decrease in the oxygen by amperometric measurement (−0.5 to −1 V reduction at a Pt black electrode).
The stoichiometry of Equation 1 clearly demonstrates some of the problems with an implantable glucose sensor. If there is excess oxygen for Equation 1, then the H
2
O
2
is stoichiometrically related to the amount of glucose that reacts at the enzyme. In this case, the ultimate current is also proportional to the amount of glucose that reacts with the enzyme. If there is insufficient oxygen for all of the glucose to react with the enzyme, then the current will be proportional to the oxygen concentration, not the glucose concentration. For the sensor to be a true glucose sensor, glucose must be the limiting reagent, i.e. the O
2
concentration must be in excess for all potential glucose concentrations. For a number of conditions, this requirement is not easily achieved. For example, the glucose concentration in the body of a diabetic patient can vary from 2 to 30 mM (millimoles per liter or 36 to 540 mg/dl), whereas the typical oxygen concentration in the tissue is 0.02 to 0.2 mM (see, Fisher, et al., Biomed. Biochem. Acta. 48:965-971 (1989). This ratio in the body means that the sensor would be running in the Michaelis Menten limited regime and would be very insensitive to small changes in the glucose concentration. This problem has been called the “oxygen deficit problem”. Accordingly, a method or system must be devised to either increase the O
2
in the GOX membrane, decrease the glucose concentration, or devise a sensor that does not use O
2
.
Several approaches to solving the deficit problem have been attempted in the past. The simplest approach is to make a membrane that is fully O
2
permeable, with no glucose permeability and mechanically perforate it with a small hole that allows glucose to pass. Here the differential permeability is defined by the ratio of the small hole area to the total membrane area. Two significant problems with this method are first that reproducibly making small holes is difficult and second and more serious, the O
2
permeability is a strong function of the thickness of the membrane and thickness is difficult to control in mass production. Microporous membranes (U.S. Pat. No. 4,759,828 to Young et al., incorporated herein by reference) have also been tried with limited success. Another problem with both the perforated membrane approach and the microporous membrane approach is that the sensor electrodes and the enzyme layer are exposed to body fluids. Body fluids contain proteins that coat the electrodes leading to decreased sensitivity of the sensor and enzymes (proteases) that can digest or degrade the sensor active enzyme.
Another approach to the oxygen deficit problem is described by Gough (U.S. Pat. No. 4,484,987, incorporated herein by reference). The approach uses a combination membrane with discrete domains of a hydrophilic material embedded in a hydrophobic membrane. In this case, the membrane is not homogenous and manufacturing reproducibility is difficult. Physical properties of the membrane are also compromised. In a similar manner, Gough (U.S. Pat. No. 4,890,620, incorporated herein by reference) describes a “two dimensional” system where glucose diffusion is limited to one dimension while the oxygen diffusion is from both dimensions. This sensor is extremely complicated and manufacturing on a large scale is expected to be difficult.
Several other groups have used a homogenous membrane of a relatively hydrophobic polyurethane and reported good results. See, for example, Shaw, et al., Biosensors and Bioelectronics, 6:401-406 (1991); Bindra, et al., Anal. Chem 63:1692 (1991); and Schichiri, et al., Horm. Metab. Resl. Suppl. Ser., 20:17 (1988). In classical diffusion experiments with these membranes, however, the glucose diffusion is extremely small. It is believed that the ability of these polyurethane layers to allow glucose diffusion is due to micro cracks or micro holes in these materials when applied as membranes.
Still others have developed homogeneous membranes with both hydrophilic and hydrophobic regions to circumvent the oxygen deficit problem. See, Allen et al., U.S. Pat. Nos. 5,284,140 and 5,322,063, the disclosures of each being incorporate

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