IMPLANTABLE ANALYTE SENSOR
First Claim
1. A method for implanting an analyte sensor in a host, the method comprising:
- forming a precisely dimensioned pocket in a subcutaneous space of a host, wherein the pocket is dimensioned no greater than dimensions of an analyte sensor; and
inserting the analyte sensor into the precisely dimensioned pocket so as to minimize movement of the sensor within the pocket.
1 Assignment
0 Petitions

Accused Products

Abstract
An implantable analyte sensor including a sensing region for measuring the analyte and a non-sensing region for immobilizing the sensor body in the host. The sensor is implanted in a precisely dimensioned pocket to stabilize the analyte sensor in vivo and enable measurement of the concentration of the analyte in the host before and after formation of a foreign body capsule around the sensor. The sensor further provides a transmitter for RF transmission through the sensor body, electronic circuitry, and a power source optimized for long-term use in the miniaturized sensor body.
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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DexCom Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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DexCom Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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DexCom Incorporated
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DexCom Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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Abbott Diabetes Care Incorporated
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10 Claims
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1. A method for implanting an analyte sensor in a host, the method comprising:
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forming a precisely dimensioned pocket in a subcutaneous space of a host, wherein the pocket is dimensioned no greater than dimensions of an analyte sensor; and inserting the analyte sensor into the precisely dimensioned pocket so as to minimize movement of the sensor within the pocket. - View Dependent Claims (2, 3, 4, 5, 6, 7, 8, 9, 10)
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1 Specification
This application is a division of U.S. application Ser. No. 10/838,912, filed May 3, 2004, the contents of which is hereby incorporated by reference herein in its entirety.
The present invention relates generally to systems and methods for making and using an implantable analyte sensor.
Diabetes mellitus is a disorder in which the pancreas cannot create sufficient insulin (Type I or insulin dependent) and/or in which insulin is not effective (Type 2 or non-insulin dependent). In the diabetic state, the victim suffers from high blood sugar, which may cause an array of physiological derangements (for example, kidney failure, skin ulcers, or bleeding into the vitreous of the eye) associated with the deterioration of small blood vessels. A hypoglycemic reaction (low blood sugar) may be induced by an inadvertent overdose of insulin, or after a normal dose of insulin or glucose-lowering agent accompanied by extraordinary exercise or insufficient food intake.
Conventionally, a diabetic person carries a self-monitoring blood glucose (SMBG) monitor, which typically comprises uncomfortable finger pricking methods. Due to the lack of comfort and convenience, a diabetic will normally only measure his or her glucose level two to four times per day. Unfortunately, these time intervals are so far spread apart that the diabetic will likely find out too late, sometimes incurring dangerous side effects, of a hyper- or hypo-glycemic condition. In fact, it is not only unlikely that a diabetic will take a timely SMBG value, but the diabetic will not know if their blood glucose value is going up (higher) or down (lower) based on conventional methods, inhibiting their ability to make educated insulin therapy decisions.
The prior art discloses a variety of analyte sensors that provide complex, short-term, transcutaneous, or partially implantable analyte sensors. Unfortunately, each of these sensors suffers from various disadvantages, such as lack of continuous care (short-term sensors), discomfort (transcutaneous and partially implantable sensors), and inconvenience (sensors with multiple components).
There is a need for a device a long-term, implantable analyte sensor that functions accurately and reliably, to provide improved patient convenience and care.
Accordingly, in a first embodiment, an implantable analyte sensor for measuring an analyte in a host is provided, the sensor including: a sensor body including a sensing region for measuring the analyte and a non-sensing region for immobilizing the sensor body in the host; a first biointerface material adjacent to the sensing region, wherein the first biointerface material includes a porous architecture that promotes vascularized tissue ingrowth and interferes with barrier cell layer formation, for allowing analyte transport to the sensing region in vivo; and a second biointerface material adjacent to at least a portion of the non-sensing region, wherein the second biointerface material includes a porous architecture that promotes tissue ingrowth for anchoring the sensor in vivo.
In an aspect of the first embodiment, the first biointerface material further includes a domain proximal to the sensing region that is impermeable to cells or cell processes and is permeable to the passage of the analyte.
In an aspect of the first embodiment, the second biointerface material and the first biointerface material include porous silicone, and wherein first biomaterial material includes porous silicone with a through porosity substantially across the entire material.
In an aspect of the first embodiment, the sensing region is on a first surface of the sensor body, and wherein the sensor body includes a second surface opposite the first surfaces, and wherein the second biointerface material is disposed on a substantial portion of the first and second surfaces of the sensor body.
In a second embodiment, an analyte sensor for short-term and long-term immobilization in a host'"'"'s soft tissue is provided, the sensor including: a short-term anchoring mechanism for providing immobilization of the sensor in the soft tissue prior to substantial formation of the foreign body capsule; and a long-term anchoring mechanism for providing immobilization of the sensor in the soft tissue during and after substantial formation of the foreign body capsule.
In an aspect of the second embodiment, the short-term anchoring mechanism includes a suture tab on the sensor body. In an aspect of the second embodiment, the short-term anchoring mechanism includes a suture. In an aspect of the second embodiment, the short-term anchoring mechanism includes at least one of prongs, spines, barbs, wings, and hooks. In an aspect of the second embodiment, the short-term anchoring mechanism includes a geometric configuration of the sensor body. In an aspect of the second embodiment, the geometric configuration includes at least one of a helical, tapered, and winged configuration.
In an aspect of the second embodiment, the long-term anchoring mechanism includes an anchoring material disposed on the sensor body. In an aspect of the second embodiment, the anchoring material includes a fibrous material. In an aspect of the second embodiment, the anchoring material includes a porous material. In an aspect of the second embodiment, the anchoring material includes a material with a surface topography. In an aspect of the second embodiment, the long-term anchoring mechanism includes a surface topography formed on an outer surface of the sensor body.
In a third embodiment, a method for immobilizing an analyte sensor in soft tissue is provided, the method including: implanting the analyte sensor in a host; anchoring the sensor in the host prior to formation of a foreign body capsule for at least short-term immobilization of the sensor within the soft tissue of the host; and anchoring the sensor within the foreign body capsule for long-term immobilization of the sensor within the soft tissue of the host.
In an aspect of the third embodiment, the short-term immobilization step includes suturing the sensor to the host'"'"'s tissue. In an aspect of the third embodiment, the suturing step includes suturing the sensor such that the sensor is in compression. In an aspect of the third embodiment, the short term immobilization step includes utilizing at least one of prongs, spines, barbs, wings, and hooks on the sensor to anchor the sensor into the host'"'"'s tissue upon implantation.
In an aspect of the third embodiment, the long-term immobilization step includes disposing an anchoring material on the sensor body that allows host tissue ingrowth into the material. In an aspect of the third embodiment, the anchoring material includes a fibrous material. In an aspect of the third embodiment, the anchoring material includes a porous material. In an aspect of the third embodiment, the anchoring material includes a material with a surface topography. In an aspect of the third embodiment, the long-term immobilization step includes utilizing a surface topography formed on an outer surface of the sensor body that allows tissue ingrowth into the surface of the sensor body.
In a fourth embodiment, a method for continuous measurement of an analyte in a host is provided, the method including: implanting an analyte sensor in a host; and measuring the concentration of the analyte in the host before and after formation of a foreign body capsule around the sensor.
In an aspect of the fourth embodiment, the analyte sensor is wholly implanted into the host. In an aspect of the fourth embodiment, the method further includes explanting the analyte sensor from the host. In an aspect of the fourth embodiment, the method further includes implanting another analyte sensor in the host. In an aspect of the fourth embodiment, the another analyte sensor is wholly implanted into the host.
In an aspect of the fourth embodiment, the method further includes anchoring the analyte sensor in the host prior to formation of a foreign body capsule for at least short-term immobilization of the sensor within the soft tissue of the host. In an aspect of the fourth embodiment, the method further includes anchoring the sensor within the foreign body capsule for long-term immobilization of the sensor within the soft tissue of the host.
In a fifth embodiment, a method for implantation of an analyte sensor in a host is provided, the method including: forming a precisely dimensioned pocket in the subcutaneous space of the host, wherein the pocket is dimensioned no greater than the dimensions of the analyte sensor; inserting the analyte sensor into the precisely-dimensioned pocket so as to minimize movement of the sensor within the pocket.
In an aspect of the fifth embodiment, the step of forming a pocket includes using a tool that allows precise dimensioning of the pocket. In an aspect of the fifth embodiment, the tool includes a head dimensioned substantially similar to the dimensions of the analyte sensor and a handle for guiding the head into the pocket. In an aspect of the fifth embodiment, the tool includes a head dimensioned smaller than the dimensions of the analyte sensor and a handle for guiding the head into the pocket.
In an aspect of the fifth embodiment, the method further includes suturing the analyte sensor to the host tissue.
In an aspect of the fifth embodiment, the pocket is formed adjacent the fascia of the host. In an aspect of the fifth embodiment, the analyte sensor includes a sensing region for measuring an analyte concentration, and wherein the analyte sensor is placed within the pocket such that the sensing region is located adjacent to the fascia.
In an aspect of the fifth embodiment, the method further includes a step of forming a vertical incision prior to the step of forming a pocket. In an aspect of the fifth embodiment, the method further includes a step of forming a horizontal incision prior to the step of forming a pocket. In an aspect of the fifth embodiment, the pocket is formed in the abdominal region of the host.
In a sixth embodiment, an implantable analyte sensor for measuring an analyte concentration in a host is provided, the sensor including: a sensor body substantially formed from a water vapor permeable material; and electrical components encapsulated within the sensor body, wherein the electrical components include RF circuitry and an antenna adapted for RF transmission from the sensor in vivo to a receiver ex vivo, wherein the RF circuitry is spaced a fixed distance from the sensor body so as to support a dielectric constant that enables RF transmission between the sensor in vivo to the receiver ex vivo.
In an aspect of the sixth embodiment, the fixed distance includes a configuration that reduces water permeability therein. In an aspect of the sixth embodiment, the configuration includes conformal coating. In an aspect of the sixth embodiment, the conformal coating includes Parylene.
In an aspect of the sixth embodiment, the configuration includes epoxy. In an aspect of the sixth embodiment, the configuration includes glass. In an aspect of the sixth embodiment, the configuration includes one or more hermetic containers.
In a seventh embodiment, an analyte sensor for RF transmission between the analyte sensor in vivo and a receiver ex vivo is provided, the sensor including: a sensor body including RF circuitry encapsulated within a substantially water vapor permeable body that enables RF transmission therethrough; a sensing region located on an outer surface of the sensor body for measuring an analyte in soft tissue; a biointerface material disposed adjacent to the sensing region that supports vascularized tissue ingrowth for transport of the analyte to the sensing region; and an anchoring material on a non-sensing outer surface of the sensor body that supports tissue ingrowth for immobilization of the sensor body in soft tissue.
In an aspect of the seventh embodiment, the sensor further includes an antenna encapsulated within the sensor body. In an aspect of the seventh embodiment, the sensor further includes a power source encapsulated within the sensor body.
In an aspect of the seventh embodiment, the sensor body is formed from plastic. In an aspect of the seventh embodiment, the plastic includes epoxy.
In an aspect of the seventh embodiment, the sensor body is molded around the RF circuitry. In an aspect of the seventh embodiment, the sensor further includes an electrode system exposed at the sensing region. In an aspect of the seventh embodiment, the electrode system extends through the water vapor permeable body and is operably connected to the RF circuitry.
In an aspect of the seventh embodiment, the biointerface material includes a solid portion with a plurality of interconnected cavities. In an aspect of the seventh embodiment, the biointerface material further includes a domain proximal to the sensing region that is impermeable to cells or cell processes and is permeable to the passage of the analyte.
In an eighth embodiment, an electrochemical analyte sensor for measuring an analyte concentration is provided, the sensor including: a sensor body including electronic circuitry encapsulated within the sensor body; and a plurality of electrodes that extend from an outer surface of the sensor body to the encapsulated electronic circuitry, wherein the electrodes are mechanically and electrically connected and aligned to the electronic circuitry prior to encapsulation within the sensor body.
In an aspect of the eighth embodiment, the electrodes are swaged to the electronic circuitry. In an aspect of the eighth embodiment, the electrodes are welded using a technique selected from the group consisting of spot welding, ultrasonic welding, and laser welding.
In an aspect of the eighth embodiment, the electrodes and electronic circuitry are encapsulated in the sensor body by a molding process. In an aspect of the eighth embodiment, the sensor body includes a water vapor permeable material.
In an aspect of the eighth embodiment, the electronic circuitry is spaced from the water vapor permeable sensor body, such that water vapor penetration within a fixed distance from the electronic circuitry is inhibited. In an aspect of the eighth embodiment, the electronic circuitry is spaced from the water vapor permeable sensor body by epoxy. In an aspect of the eighth embodiment, the electronic circuitry is spaced from the water vapor permeable sensor body by a glass tube. In an aspect of the eighth embodiment, the electronic circuitry is spaced from the water vapor permeable sensor body by Parylene. In an aspect of the eighth embodiment, the electronic circuitry is spaced from the water vapor permeable sensor body by one or more hermetic containers.
In an aspect of the eighth embodiment, the sensor body includes a substantially seamless exterior with the electrodes extending through the sensor body to the outer surface thereof.
In a ninth embodiment, a method for manufacturing an electrochemical analyte sensor is provided, the method including: providing electronic circuitry designed to process signals from the sensor; swaging a plurality of electrodes to the electronic circuitry; and molding a plastic material around the electronic circuitry to form the sensor body.
In a tenth embodiment, a method for manufacturing an analyte sensor is provided, the method including: providing sensor electronics designed to process signals from the sensor; conformally coating the sensor electronics with a material that has a first water permeability rate; and molding a water vapor permeable material that has a second water permeability rate around the coated sensor electronics to form a substantially seamless sensor body, wherein the second water permeability rate is greater than the first water penetration rate.
In an aspect of the tenth embodiment, the molding includes a two-step molding process to form the substantially seamless sensor body.
In an aspect of the tenth embodiment, the two-step molding process includes: holding a first portion of the coated sensor electronics and molding around a second portion of the coated sensor electronics; and holding a portion of the cured sensor body and molding around the first portion of the coated sensor electronics.
In an eleventh embodiment, a method for manufacturing a multilayer membrane for an analyte sensor is provided, the method including: serially casting and subsequently curing each of a plurality of layers to form the multilayer membrane onto a liner, wherein the layers include a resistance layer for limiting the passage of an analyte and an enzyme layer including an enzyme for reacting with the analyte; and releasing the multilayer membrane from the liner for application onto the analyte sensor.
In an aspect of the eleventh embodiment, the multilayer membrane further includes an interference layer that substantially prevents passage of potentially electrochemically interfering substances.
In an aspect of the eleventh embodiment, the multilayer membrane further includes an electrolyte layer including a hydrogel for maintaining hydrophilicity at electrochemically reactive surfaces of the analyte sensor.
In a twelfth embodiment, a method for casting a membrane that regulates the transport of glucose, the method including: forming a solvent solution including a polymer blend and a solvent, wherein the polymer blend includes hydrophilic and hydrophobic components; maintaining the solution at a first elevated temperature for a predetermined time period in order to mix the hydrophilic and hydrophobic components with each other and the solvent, wherein the elevated temperature is above room temperature; applying the composition to a liner to form a film thereon; and curing the film, wherein the curing is accomplished while ramping the temperature at a predetermined ramp rate to a second elevated temperature that is above the first temperature.
In an aspect of the twelfth embodiment, the first elevated temperature is between about 60° C. and about 100° C. In an aspect of the twelfth embodiment, the first elevated temperature is about 80° C.
In an aspect of the twelfth embodiment, the predetermined time period is at least about 24 hours. In an aspect of the twelfth embodiment, the predetermined time period is at least about 44 hours.
In an aspect of the twelfth embodiment, the predetermined ramp rate is between about 3° C. per minute and 12° C. per minute. In an aspect of the twelfth embodiment, the predetermined ramp rate is about 7° C. per minute.
In an aspect of the twelfth embodiment, the second elevated temperature is at least about 100° C.
In a thirteenth embodiment, a method for casting a membrane for use with an electrochemical glucose sensor is provided, wherein the membrane substantially prevents passage of potentially electrochemically interfering substances, the method including: forming a sufficiently diluted solvent solution including a polymer and a solvent, wherein sufficiently diluted solvent solution includes a ratio of polymer to solvent of about 1 to 10 wt. % polymer to about 90 to 99 wt. % solvent; and applying the solvent solution at a sufficiently fast casting speed that substantially avoids film thickness inhomogeneities due to evaporation during casting of the sufficiently diluted solvent solution.
In an aspect of the thirteenth embodiment, the membrane limits the diffusion of hydrophilic species and large molecular weight species.
In an aspect of the thirteenth embodiment, the membrane includes a thickness between about 0.1 and 5 microns. In an aspect of the thirteenth embodiment, the membrane includes a thickness between about 0.5 and 3 microns.
In an aspect of the thirteenth embodiment, the polymer includes polyurethane.
In an aspect of the thirteenth embodiment, the sufficiently fast casting speed is between about 8 to about 15 inches/second. In an aspect of the thirteenth embodiment, the sufficiently fast casting speed is about 11.5 inches/second.
In a fourteenth embodiment, an implantable analyte sensor is provided, the sensor including: a body including a material which is permeable to water vapor, the body further including a sensing region for measuring levels of an analyte; and a transmitter within the body for transmitting the measurements obtained by the sensing region, wherein at least a portion of the transmitter is spaced from the body by a material adapted to reduce water from penetrating therein.
In an aspect of the fourteenth embodiment, the transmitter includes an oscillator and at least a portion of the oscillator is spaced from the body by the material adapted to inhibit fluid from penetrating therein. In an aspect of the fourteenth embodiment, the oscillator includes an inductor and wherein the inductor is spaced from the body by the material adapted to inhibit fluid from penetrating therein. In an aspect of the fourteenth embodiment, the oscillator includes a voltage controlled oscillator.
In a fifteenth embodiment, an implantable analyte sensor is provided, including: electronics encapsulated within a water vapor permeable body, wherein the electronics include a microprocessor module and an RF module that has an RF transceiver with a phase-locked loop, and wherein the microprocessor module is programmed to initiate re-calibration of the PLL responsive to detection of off-frequency shift.
In an aspect of the fifteenth embodiment, an electrochemical glucose sensor including a three-electrode system, the sensor including: an electrochemical cell including a working electrode, reference electrode, and counter electrode; and a potentiostat that controls the potential between the working and reference electrodes, wherein an allowable range for the counter electrode voltage is set sufficiently wide such that the glucose sensor can react with other reducible species when oxygen becomes limited and sufficiently narrow to ensure the circuitry does not allow excessive current draw or bubble formation to occur.
In an aspect of the fifteenth embodiment, limiting the current of at least one of the working or counter electrode amplifiers to a preset current value configures the allowable range. In an aspect of the fifteenth embodiment, setting the op-amp to be offset from battery ground configures the allowable range. In an aspect of the fifteenth embodiment, a reference voltage setting between about +0.6V and +0.8V with respect to battery ground configures the allowable range. In an aspect of the fifteenth embodiment, a reference voltage setting of about +0.7V with respect to battery ground configures the allowable range.
In a sixteenth embodiment, a method for manufacturing an analyte sensor is provided, the method including: providing a sensor body, wherein the sensor body includes a sensing region for measuring the analyte; forming a multilayer membrane on a liner; releasing the multilayer membrane from the liner and onto the sensor body; and attaching the multilayer membrane to the analyte sensor body proximal to the sensing region.
In an aspect of the sixteenth embodiment, the attaching step includes a mechanical attachment. In an aspect of the sixteenth embodiment, the mechanical attachment includes a metal or plastic O-ring adapted to fit around a raised sensing region. In an aspect of the sixteenth embodiment, the mechanical attachment includes a metal or plastic disc adapted to be press-fit into the sensor body. In an aspect of the sixteenth embodiment, the mechanical attachment includes a metal or plastic clip adapted to be snap-fit into the sensor body.
In a seventeenth embodiment, a method for manufacturing an analyte sensor is provided, the method including: providing a sensor body, wherein the sensor body includes a sensing region for measuring the analyte; forming a multilayer membrane on a liner; releasing the multilayer membrane from the liner and placing onto the sensor body; and attaching the multilayer membrane to the analyte sensor body proximal to the sensing region.
In an aspect of the seventeenth embodiment, the attaching step includes a mechanical attachment. In an aspect of the seventeenth embodiment, the mechanical attachment includes a metal or plastic O-ring adapted to fit around a raised sensing region. In an aspect of the seventeenth embodiment, the mechanical attachment includes a metal or plastic disc adapted to be press-fit into the sensor body. In an aspect of the seventeenth embodiment, the mechanical attachment includes a metal or plastic clip adapted to be snap-fit into the sensor body.
The following description and examples illustrate some exemplary embodiments of the disclosed invention in detail. Those of skill in the art will recognize that there are numerous variations and modifications of this invention that are encompassed by its scope. Accordingly, the description of a certain exemplary embodiment should not be deemed to limit the scope of the present invention.
In order to facilitate an understanding of the disclosed invention, a number of terms are defined below.
The term “ROM,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, read-only memory. The term is inclusive of various types of ROM, including EEPROM, rewritable ROMs, flash memory, or the like.
The term “RAM,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, random access memory. The term is inclusive of various types of RAM, including dynamic-RAM, static-RAM, non-static RAM, or the like.
The term “A/D Converter,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, hardware and/or software that converts analog electrical signals into corresponding digital signals.
The term “microprocessor,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation a computer system or processor designed to perform arithmetic and logic operations using logic circuitry that responds to and processes the basic instructions that drive a computer.
The term “RF transceiver,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, a radio frequency transmitter and/or receiver for transmitting and/or receiving signals.
The terms “raw data stream” and “data stream,” as used herein, are broad terms and are used in their ordinary sense, including, without limitation, an analog or digital signal directly related to the measured glucose from the glucose sensor. In one example, the raw data stream is digital data in “counts” converted by an A/D converter from an analog signal (for example, voltage or amps) representative of a glucose concentration. The terms broadly encompass a plurality of time spaced data points from a substantially continuous glucose sensor, which comprises individual measurements taken at time intervals ranging from fractions of a second up to, for example, 1, 2, or 5 minutes or longer.
The term “counts,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, a unit of measurement of a digital signal. In one example, a raw data stream measured in counts is directly related to a voltage (for example, converted by an A/D converter), which is directly related to current from the working electrode. In another example, counter electrode voltage measured in counts is directly related to a voltage.
The term “potentiostat,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, an electrical system that controls the potential between the working and reference electrodes of a three-electrode cell at a preset value. It forces whatever current is necessary to flow between the working and counter electrodes to keep the desired potential, as long as the needed cell voltage and current do not exceed the compliance limits of the potentiostat.
The term “electrical potential,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, the electrical potential difference between two points in a circuit which is the cause of the flow of a current.
The term “physiologically feasible,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, the physiological parameters obtained from continuous studies of glucose data in humans and/or animals. For example, a maximal sustained rate of change of glucose in humans of about 4 to 5 mg/dL/min and a maximum acceleration of the rate of change of about 0.1 to 0.2 mg/dL/min/min are deemed physiologically feasible limits. Values outside of these limits would be considered non-physiological and likely a result of signal error, for example. As another example, the rate of change of glucose is lowest at the maxima and minima of the daily glucose range, which are the areas of greatest risk in patient treatment, thus a physiologically feasible rate of change can be set at the maxima and minima based on continuous studies of glucose data.
The term “ischemia,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, local and temporary deficiency of blood supply due to obstruction of circulation to a part (for example, sensor). Ischemia can be caused by mechanical obstruction (for example, arterial narrowing or disruption) of the blood supply, for example.
The term “system noise,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, unwanted electronic or diffusion-related noise which can include Gaussian, motion-related, flicker, kinetic, or other white noise, for example.
The term “biointerface membrane” as used herein is a broad term and is used in its ordinary sense, including, without limitation, a permeable membrane that functions as a device-tissue interface comprised of two or more domains. In some embodiments, the biointerface membrane is composed of two domains. The first domain supports tissue ingrowth, interferes with barrier cell layer formation, and includes an open cell configuration having cavities and a solid portion. The second domain is resistant to cellular attachment and impermeable to cells (for example, macrophages). The biointerface membrane is made of biostable materials and can be constructed in layers, uniform or non-uniform gradients (i.e., anisotropic), or in a uniform or non-uniform cavity size configuration.
The term “sensing membrane,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, a permeable or semi-permeable membrane that can be comprised of two or more domains and is typically constructed of materials of a few microns thickness or more, which are permeable to oxygen and may or may not be permeable to glucose. In one example, the sensing membrane comprises an enzyme, for example immobilized glucose oxidase enzyme, which enables an electrochemical reaction to occur to measure a concentration of analyte.
The term “domain” as used herein is a broad term and is used in its ordinary sense, including, without limitation, regions of a membrane that can be layers, uniform or non-uniform gradients (for example, anisotropic) or provided as portions of the membrane. The term is broad enough to include one or more functions one or more (combined) domains, or a plurality of layers or regions that each provide one or more of the functions of each of the various domains.
The term “barrier cell layer” as used herein is a broad term and is used in its ordinary sense, including, without limitation, a cohesive monolayer of cells (for example, macrophages and foreign body giant cells) that substantially block the transport of at least some molecules across the second domain and/or membrane.
The term “cellular attachment,” as used herein is a broad term and is used in its ordinary sense, including, without limitation, adhesion of cells and/or mechanical attachment of cell processes to a material at the molecular level, and/or attachment of cells and/or cell processes to micro- (or macro-) porous material surfaces. One example of a material used in the prior art that allows cellular attachment due to porous surfaces is the BIOPORE™ cell culture support marketed by Millipore (Bedford, Mass.) (see Brauker '"'"'330, supra).
The phrase “distal to” as used herein is a broad term and is used in its ordinary sense, including, without limitation, the spatial relationship between various elements in comparison to a particular point of reference. For example, some embodiments of a device include a biointerface membrane having a cell disruptive domain and a cell impermeable domain. If the sensor is deemed to be the point of reference and the cell disruptive domain is positioned farther from the sensor, then that domain is distal to the sensor.
The term “proximal to” as used herein is a broad term and is used in its ordinary sense, including, without limitation, the spatial relationship between various elements in comparison to a particular point of reference. For example, some embodiments of a device include a biointerface membrane having a cell disruptive domain and a cell impermeable domain. If the sensor is deemed to be the point of reference and the cell impermeable domain is positioned nearer to the sensor, then that domain is proximal to the sensor.
The term “cell processes” as used herein is a broad term and is used in its ordinary sense, including, without limitation, pseudopodia of a cell.
The term “solid portions” as used herein is a broad term and is used in its ordinary sense, including, without limitation, a solid material having a mechanical structure that demarcates the cavities, voids, or other non-solid portions.
The term “substantial” as used herein is a broad term and is used in its ordinary sense, including, without limitation, a sufficient amount that provides a desired function. For example, in the micro-architecture of the preferred embodiments, a substantial number of cavities have a size that allows a substantial number of inflammatory cells to enter therein, which may include an amount greater than 50 percent, an amount greater than 60 percent, an amount greater than 70 percent, an amount greater than 80 percent, and an amount greater than 90 percent of cavities within a preferred nominal pore size range.
The term “co-continuous” as used herein is a broad term and is used in its ordinary sense, including, without limitation, a solid portion wherein an unbroken curved line in three dimensions exists between any two points of the solid portion.
The term “biostable” as used herein is a broad term and is used in its ordinary sense, including, without limitation, materials that are relatively resistant to degradation by processes that are encountered in vivo.
The term “analyte” as used herein is a broad term and is used in its ordinary sense, including, without limitation, to refer to a substance or chemical constituent in a biological fluid (for example, blood, interstitial fluid, cerebral spinal fluid, lymph fluid or urine) that can be analyzed. Analytes can include naturally occurring substances, artificial substances, metabolites, and/or reaction products. In some embodiments, the analyte for measurement by the sensing regions, devices, and methods is glucose. However, other analytes are contemplated as well, including but not limited to acarboxyprothrombin; acylcarnitine; adenine phosphoribosyl transferase; adenosine deaminase; albumin; alpha-fetoprotein; amino acid profiles (arginine (Krebs cycle), histidine/urocanic acid, homocysteine, phenylalanine/tyrosine, tryptophan); andrenostenedione; antipyrine; arabinitol enantiomers; arginase; benzoylecgonine (cocaine); biotinidase; biopterin; c-reactive protein; carnitine; carnosinase; CD4; ceruloplasmin; chenodeoxycholic acid; chloroquine; cholesterol; cholinesterase; conjugated 1-β hydroxy-cholic acid; cortisol; creatine kinase; creatine kinase MM isoenzyme; cyclosporin A; d-penicillamine; de-ethylchloroquine; dehydroepiandrosterone sulfate; DNA (acetylator polymorphism, alcohol dehydrogenase, alpha 1-antitrypsin, cystic fibrosis, Duchenne/Becker muscular dystrophy, glucose-6-phosphate dehydrogenase, hemoglobinopathies, A, S, C, E, D-Punjab, beta-thalassemia, hepatitis B virus, HCMV, HIV-1, HTLV-1, Leber hereditary optic neuropathy, MCAD, RNA, PKU, Plasmodium vivax, sexual differentiation, 21-deoxycortisol); desbutylhalofantrine; dihydropteridine reductase; diphtheria/tetanus antitoxin; erythrocyte arginase; erythrocyte protoporphyrin; esterase D; fatty acids/acylglycines; free β-human chorionic gonadotropin; free erythrocyte porphyrin; free thyroxine (FT4); free tri-iodothyronine (FT3); fumarylacetoacetase; galactose/gal-1-phosphate; galactose-1-phosphate uridyltransferase; gentamicin; glucose-6-phosphate dehydrogenase; glutathione; glutathione perioxidase; glycocholic acid; glycosylated hemoglobin; halofantrine; hemoglobin variants; hexosaminidase A; human erythrocyte carbonic anhydrase I; 17 alpha-hydroxyprogesterone; hypoxanthine phosphoribosyl transferase; immunoreactive trypsin; lactate; lead; lipoproteins ((a), B/A-1, β); lysozyme; mefloquine; netilmicin; phenobarbitone; phenytoin; phytanic/pristanic acid; progesterone; prolactin; prolidase; purine nucleoside phosphorylase; quinine; reverse tri-iodothyronine (rT3); selenium; serum pancreatic lipase; sissomicin; somatomedin C; specific antibodies (adenovirus, anti-nuclear antibody, anti-zeta antibody, arbovirus, Aujeszky'"'"'s disease virus, dengue virus, Dracunculus medinensis, Echinococcus granulosus, Entamoeba histolytica, enterovirus, Giardia duodenalisa, Helicobacter pylori, hepatitis B virus, herpes virus, HIV-1, IgE (atopic disease), influenza virus, Leishmania donovani, leptospira, measles/mumps/rubella, Mycobacterium leprae, Mycoplasma pneumoniae, Myoglobin, Onchocerca volvulus, parainfluenza virus, Plasmodium falciparum, poliovirus, Pseudomonas aeruginosa, respiratory syncytial virus, rickettsia (scrub typhus), Schistosoma mansoni, Toxoplasma gondii, Trepenoma pallidium, Trypanosoma cruzi/rangeli, vesicular stomatis virus, Wuchereria bancrofti, yellow fever virus); specific antigens (hepatitis B virus, HIV-1); succinylacetone; sulfadoxine; theophylline; thyrotropin (TSH); thyroxine (T4); thyroxine-binding globulin; trace elements; transferrin; UDP-galactose-4-epimerase; urea; uroporphyrinogen I synthase; vitamin A; white blood cells; and zinc protoporphyrin. Salts, sugar, protein, fat, vitamins and hormones naturally occurring in blood or interstitial fluids can also constitute analytes in certain embodiments. The analyte can be naturally present in the biological fluid, for example, a metabolic product, a hormone, an antigen, an antibody, and the like. Alternatively, the analyte can be introduced into the body, for example, a contrast agent for imaging, a radioisotope, a chemical agent, a fluorocarbon-based synthetic blood, or a drug or pharmaceutical composition, including but not limited to insulin; ethanol; cannabis (marijuana, tetrahydrocannabinol, hashish); inhalants (nitrous oxide, amyl nitrite, butyl nitrite, chlorohydrocarbons, hydrocarbons); cocaine (crack cocaine); stimulants (amphetamines, methamphetamines, Ritalin, Cylert, Preludin, Didrex, PreState, Voranil, Sandrex, Plegine); depressants (barbituates, methaqualone, tranquilizers such as Valium, Librium, Miltown, Serax, Equanil, Tranxene); hallucinogens (phencyclidine, lysergic acid, mescaline, peyote, psilocybin); narcotics (heroin, codeine, morphine, opium, meperidine, Percocet, Percodan, Tussionex, Fentanyl, Darvon, Talwin, Lomotil); designer drugs (analogs of fentanyl, meperidine, amphetamines, methamphetamines, and phencyclidine, for example, Ecstasy); anabolic steroids; and nicotine. The metabolic products of drugs and pharmaceutical compositions are also contemplated analytes. Analytes such as neurochemicals and other chemicals generated within the body can also be analyzed, such as, for example, ascorbic acid, uric acid, dopamine, noradrenaline, 3-methoxytyramine (3MT), 3,4-dihydroxyphenylacetic acid (DOPAC), homovanillic acid (HVA), 5-hydroxytryptamine (5HT), and 5-hydroxyindoleacetic acid (FHIAA).
The terms “operably connected” and “operably linked” as used herein are broad terms and are used in their ordinary sense, including, without limitation, one or more components being linked to another component(s) in a manner that allows transmission of signals between the components. For example, one or more electrodes can be used to detect the amount of analyte in a sample and convert that information into a signal; the signal can then be transmitted to a circuit. In this case, the electrode is “operably linked” to the electronic circuitry.
The term “host” as used herein is a broad term and is used in its ordinary sense, including, without limitation, mammals, particularly humans.
The phrase “continuous (or continual) analyte sensing” as used herein is a broad term and is used in its ordinary sense, including, without limitation, the period in which monitoring of analyte concentration is continuously, continually, and or intermittently (regularly or irregularly) performed, for example, about every 5 to 10 minutes.
The term “sensing region” as used herein is a broad term and is used in its ordinary sense, including, without limitation, the region of a monitoring device responsible for the detection of a particular analyte. The sensing region generally comprises a non-conductive body, a working electrode (anode), a reference electrode and a counter electrode (cathode) passing through and secured within the body forming an electrochemically reactive surface at one location on the body and an electronic connective means at another location on the body, and a multi-region membrane affixed to the body and covering the electrochemically reactive surface. The counter electrode has a greater electrochemically reactive surface area than the working electrode. During general operation of the sensor a biological sample (for example, blood or interstitial fluid) or a portion thereof contacts (directly or after passage through one or more membranes or domains) an enzyme (for example, glucose oxidase); the reaction of the biological sample (or portion thereof) results in the formation of reaction products that allow a determination of the analyte (for example, glucose) level in the biological sample. In some embodiments, the multi-region membrane further comprises an enzyme domain (for example, and enzyme layer), and an electrolyte phase (i.e., a free-flowing liquid phase comprising an electrolyte-containing fluid described further below).
The term “electrochemically reactive surface” as used herein is a broad term and is used in its ordinary sense, including, without limitation, the surface of an electrode where an electrochemical reaction takes place. In the case of the working electrode, the hydrogen peroxide produced by the enzyme catalyzed reaction of the analyte being detected reacts creating a measurable electric current (for example, detection of glucose analyte utilizing glucose oxidase produces H2O2 peroxide as a by product, H2O2 reacts with the surface of the working electrode producing two protons (2H+), two electrons (2e−) and one molecule of oxygen (O2) which produces the electronic current being detected). In the case of the counter electrode, a reducible species, for example, O2 is reduced at the electrode surface in order to balance the current being generated by the working electrode.
The term “electronic connection” as used herein is a broad term and is used in its ordinary sense, including, without limitation, any electronic connection known to those in the art that can be utilized to interface the sensing region electrodes with the electronic circuitry of a device such as mechanical (for example, pin and socket) or soldered.
The term “oxygen antenna domain” as used herein is a broad term and is used in its ordinary sense, including, without limitation, a domain composed of a material that has higher oxygen solubility than aqueous media so that it concentrates oxygen from the biological fluid surrounding the biointerface membrane. The domain can then act as an oxygen reservoir during times of minimal oxygen need and has the capacity to provide on demand a higher oxygen gradient to facilitate oxygen transport across the membrane. This enhances function in the enzyme reaction domain and at the counter electrode surface when glucose conversion to hydrogen peroxide in the enzyme domain consumes oxygen from the surrounding domains. Thus, this ability of the oxygen antenna domain to apply a higher flux of oxygen to critical domains when needed improves overall sensor function.
The term “casting” as used herein is a broad term and is used in its ordinary sense, including, without limitation, a process where a fluid material is applied to a surface or surfaces and allowed to cure. The term is broad enough to encompass a variety of coating techniques, for example, using a draw-down machine, dip coating, or the like.
The term “water vapor permeable” as used herein is a broad term and is used in its ordinary sense, including, without limitation, characterized by permitting water vapor to permeate therethrough.
The following abbreviations apply herein: Eq and Eqs (equivalents); mEq (milliequivalents); M (molar); mM (millimolar) μM (micromolar); N (Normal); mol (moles); mmol (millimoles); μmol (micromoles); nmol (nanomoles); g (grams); mg (milligrams); μg (micrograms); Kg (kilograms); L (liters); mL (milliliters); dL (deciliters); μL (microliters); cm (centimeters); mm (millimeters); μm (micrometers); nm (nanometers); h and hr (hours); min. (minutes); s and sec. (seconds); and ° C. (degrees Centigrade).
The continuous analyte sensor 12 measures a concentration of an analyte or a substance indicative of the concentration or presence of the analyte. Although some of the following description is drawn to a glucose sensor, the analyte sensor 12 may be any sensor capable of determining the level of any analyte in the body, for example oxygen, lactase, insulin, hormones, cholesterol, medicaments, viruses, or the like. Additionally, although much of the description of the analyte sensor is focused on electrochemical detection methods, the systems and methods may be applied to analyte sensors that utilize other measurement techniques, including enzymatic, chemical, physical, spectrophotometric, polarimetric, calorimetric, radiometric, or the like.
Reference is now made to
In one preferred embodiment, the analyte sensor is a glucose sensor, wherein the sensing region 16 comprises electrode system including a platinum working electrode, a platinum counter electrode, and a silver/silver chloride reference electrode, for example. However a variety of electrode materials and configurations may be used with the implantable analyte sensor of the preferred embodiments. The top ends of the electrodes are in contact with an electrolyte phase (not shown), which is a free-flowing fluid phase disposed between a sensing membrane and the electrodes. In one embodiment, the counter electrode is provided to balance the current generated by the species being measured at the working electrode. In some embodiments, the sensing membrane includes an enzyme, for example, glucose oxidase, and covers the electrolyte phase. In the case of a glucose oxidase based glucose sensor, the species being measured at the working electrode is H2O2. Glucose oxidase catalyzes the conversion of oxygen and glucose to hydrogen peroxide and gluconate according to the following reaction:
Glucose+O2→Gluconate+H2O2
The change in H2O2 can be monitored to determine glucose concentration because for each glucose molecule metabolized, there is a proportional change in the product H2O2. Oxidation of H2O2 by the working electrode is balanced by reduction of ambient oxygen, enzyme generated H2O2, or other reducible species at the counter electrode. The H2O2 produced from the glucose oxidase reaction further reacts at the surface of working electrode and produces two protons (2H+), two electrons (2e−), and one oxygen molecule (O2).
A potentiostat (
A microprocessor module 28 includes the central control unit that houses ROM 30 and RAM 32 and controls the processing of the sensor electronics 22. It is noted that certain alternative embodiments can utilize a computer system other than a microprocessor to process data as described herein. In some alternative embodiments, an application-specific integrated circuit (ASIC) can be used for some or all the sensor'"'"'s central processing as is appreciated by one skilled in the art. The ROM 30 provides semi-permanent storage of data, for example, storing data such as sensor identifier (ID) and programming to process data streams (for example, programming for data smoothing and/or replacement of signal artifacts such as described in copending U.S. patent application Ser. No. 10/648,849 and entitled, “SYSTEMS AND METHODS FOR REPLACING SIGNAL ARTIFACTS IN A GLUCOSE SENSOR DATA STREAM,” filed Aug. 22, 2003, which is incorporated herein by reference in its entirety). The RAM 32 can be used for the system'"'"'s cache memory, for example for temporarily storing recent sensor data. In some alternative embodiments, memory storage components comparable to ROM 30 and RAM 32 may be used instead of or in addition to the preferred hardware, such as dynamic-RAM, static-RAM, non-static RAM, EEPROM, rewritable ROMs, flash memory, or the like.
A battery 34 is operably connected to the sensor electronics 22 and provides the necessary power for the sensor 12. In one embodiment, the battery is a Lithium Manganese Dioxide battery, however any appropriately sized and powered battery can be used (for example, AAA, Nickel-cadmium, Zinc-carbon, Alkaline, Lithium, Nickel-metal hydride, Lithium-ion, Zinc-air, Zinc-mercury oxide, Silver-zinc, and/or hermetically-sealed). In some embodiments the battery is rechargeable. In some embodiments, a plurality of batteries can be used to power the system. In yet other embodiments, the sensor can be transcutaneously powered via an inductive coupling, for example. In some embodiments, a quartz crystal 36 is operably connected to the microprocessor 28 and maintains system time for the computer system as a whole.
An RF module 38 is operably connected to the microprocessor 28 and transmits the sensor data from the sensor 12 to a receiver 14 within a wireless transmission 40 via antenna 42. In some embodiments, a second quartz crystal 44 provides the system time for synchronizing the data transmissions from the RF transceiver. The RF transceiver generally includes a register and a phase-locked loop (PLL) with an oscillator, phase discriminator (PD), loop filter (LPF), and a voltage-controlled oscillator (VCO) as is appreciated by one skilled in the art. In some alternative embodiments, however, other mechanisms such as optical, infrared radiation (IR), ultrasonic, or the like may be used to transmit and/or receive data.
It is noted that the preferred embodiments advantageously encapsulate the electronics in a water vapor permeable material, such as described in more detail with reference to
The electronics subassembly 46 generally includes hardware and software designed to support the functions described above; additionally, the electronics subassembly of the preferred embodiments is configured to accommodate certain preferred design parameters described herein, which facilitate analyte sensor immobilization within the subcutaneous pocket. Immobilization of the sensor within the host tissue is advantageous because motion (for example, acute and/or chronic movement of the sensor in the host tissue) has been found to produce acute and/or chronic inflammation, which has been shown to result in poor short-term and/or long-term sensor performance. For example, during an experiment wherein larger, bulkier versions of the analyte sensor were implanted in humans for an average of 44 days+/−14 days [See Garg, S.; Schwartz, S.; Edelman, S. “Improved Glucose Excursions Using an Implantable Real-Time Continuous Glucose Sensor in Adults with Type I Diabetes.” Diabetes Care 2004, 27, 734-738], it was discovered that movement of the sensor resulted in thicker foreign body capsule formation, which correlated with decreased sensor performance. While not wishing to be bound by theory, it is believed that size optimization (for example, miniaturization) of the analyte sensor enables more discrete and secure implantation, and is believed to reduce macro-motion of the sensor induced by the patient and micro-motion caused by movement of the sensor within the subcutaneous pocket, and thereby improve sensor performance.
Additionally, in contrast to devices made from hermetic materials, the preferred embodiments of the present invention advantageously encapsulate the electronics in a material that is water vapor permeable. The use of a water vapor permeable material, for example moldable plastic, is advantageous for a variety of reasons described elsewhere herein, for example, ease of design changes, security and alignment of the electronics during and after the molding process, and the ability to machine the device with precise curvatures. In some embodiments, the electronic subassembly possesses features that maintain the frequency of the voltage controlled oscillator (VCO) if water vapor penetrates into the water vapor permeable sensor body. For example, in some embodiments, the electronics subassembly reduces changes in the inductor parameters, which may otherwise occur as a result of water vapor within the electromagnetic field of the inductor (for example, which may result in a shift in the carrier frequency of the VCO). Such field effects may cause the VCO to transmit off of its tuned carrier frequency, which degrades the RF telemetry capabilities of the sensor. Furthermore, the carrier frequency of the sensor of the preferred embodiments is optimized for sensor longevity.
Additionally, the analyte sensor of the preferred embodiments supports a high frequency, low power operation, which supports miniaturization of the analyte sensor with optimized functionality. Accordingly, the design of the electronics subassembly 46 of the preferred embodiments provides a discrete, efficient configuration, while maintaining long-term power supply and functional RF telemetry within the water vapor permeable body in vivo.
The PCB 48 supports the components, for example microprocessor module 28, including ROM 30 and RAM 32, the potentiostat 24, A/D converter 26, RF module 38, two crystals 36, 44, and a variety of other supporting components, deposited bonding pads, and conductors, which provide the necessary functionality described above. Additionally, the electronics subassembly 46 supports an electrode system 54 including a working electrode 54a, a reference electrode 54b, and a counter electrode 54c in one embodiment, such as described in more detail above in the Overview section, however alternative electrode systems and/or measurement techniques may be implemented.
The antenna board 50, on which the antenna (42 in
Reference is now made to the electrode system 54 of the preferred embodiments, including the working electrode (anode) 54a, the reference electrode 54b, and the counter electrode (cathode) 54c, such as shown in
The working electrode 54a and counter-electrode 54c of a glucose oxidase-based glucose sensor 12 require oxygen in different capacities. Within the enzyme layer above the working electrode 54a, oxygen is required for the production of H2O2 from glucose. The H2O2 produced from the glucose oxidase reaction further reacts at the surface of the working electrode 54a and produces two electrons. The products of this reaction are two protons (2H+), two electrons (2e−), and one oxygen molecule (O2) (See Fraser, D. M. “An Introduction to In Vivo Biosensing: Progress and problems.” In “Biosensors and the Body,” D. M. Fraser, ed., 1997, pp. 1-56 John P. Wiley and Sons, New York). In theory, the oxygen concentration near the working electrode 54a, which is consumed during the glucose oxidase reaction, is replenished by the second reaction at the working electrode 54a; therefore, the net consumption of oxygen is zero. In practice, however, not all of the H2O2 produced by the enzyme diffuses to the working electrode surface nor does all of the oxygen produced at the electrode diffuse to the enzyme domain.
Additionally, the counter electrode 54c utilizes oxygen as an electron acceptor. The most likely reducible species for this system are oxygen or enzyme generated peroxide (Fraser, D. M. supra). There are two main pathways by which oxygen may be consumed at the counter electrode 54c. These are a four-electron pathway to produce hydroxide and a two-electron pathway to produce hydrogen peroxide. Oxygen is further consumed above the counter electrode by the glucose oxidase. Due to the oxygen consumption by both the enzyme and the counter electrode, there is a net consumption of oxygen at the surface of the counter electrode 54c. Thus, in the domain of the working electrode 54a there may be significantly less net loss of oxygen than in the region of the counter electrode 54c. Furthermore, it is noted that there is a close correlation between the ability of the counter electrode 54c to maintain current balance and sensor performance. Taken together, it is believed that counter electrode 54c function becomes limited before the enzyme reaction becomes limited when oxygen concentration is lowered. When this occurs, the counter electrode limitation begins to manifest itself as this electrode moves to increasingly negative voltages in the search for reducible species. Thus, when a sufficient supply of reducible species, such as oxygen, is not available to a conventional sensor, the counter electrode voltage reaches a circuitry limit, resulting in compromised sensor performance.
In order to overcome the above-described limitations, the diameter of the counter electrode is at least twice the diameter of the working electrode, resulting in an approximately 6-fold increase in the exposed surface area of the counter electrode of the preferred embodiment. Preferably, the surface area of the electrochemically reactive surface of the counter electrode is not less than about 2 times the surface area of the working electrode. More preferably, the surface area of the electrochemically reactive surface of the counter electrode is between about 2 and about 50, between about 2 and about 25, or between about 2 and about 10 times the surface area of the working electrode.
Reference is now made to
Thus, potentiostat 24 creates current in the counter electrode by controlling the voltage applied between the reference and the working electrode. The reaction that occurs on the counter electrode is determined by how much voltage is applied to the counter electrode. By increasing the voltage applied to the counter electrode, increased amount and type of species may react in order to create the necessary current, which may be advantageous for the same reasons as described above with reference to the electrode configuration of the preferred embodiments.
In addition to the net oxygen loss described above, implantable glucose sensors face an additional challenge in maintaining sensor output during ischemic conditions, which may occur either as short-term transient events (for example, compression caused by postural effects on the device) or as long-term low oxygen conditions (for example, caused by a thickened FBC or barrier cells). When the sensor is in a low oxygen environment, the potentiostat will react by decreasing the voltage relative to the reference electrode voltage applied to the counter electrode, which may result in other less electro-active species reacting at the counter electrode.
In some embodiments, the potentiostat settings are configured to allow the counter electrode to react with other reducible species when oxygen concentration is low. In some circumstances, glucose sensors may suffer from a negative voltage setting that is too low, particularly in low oxygen environments. For example, as the voltage on the counter electrode becomes more negative, it will begin to create current, by reacting with other reducible species, a byproduct of this reaction is H2. Two potential problems can occur because of the production of Hydrogen at the counter electrode: 1) bubble formation, which disconnects the counter from the current carrying buffer and causes the sensor to lose function and 2) an interfering signal at the working electrode.
In order to overcome the potential problems, the preferred embodiments optimize the potentiostat settings to enable functionality of the potentiostat even in low oxygen conditions while limiting the counter electrode to ensure that the sensor does not create conditions that could damage it. Namely, such that when oxygen concentration decreases, the counter electrode is pushed negative enough to allow it to react with the next most abundant reducible species, for example water, which is not typically limited (for example, in vivo).
In one embodiment, the potentiostat settings are optimized by setting the allowable range for the counter electrode voltage sufficiently wide such that the sensor can react with other reducible species when oxygen becomes limited, while setting the range sufficiently narrow to ensure the circuitry does not allow excessive current draw or bubble formation to occur. The counter electrode is preferably restricted such that the species that will react at the counter electrode electroactive surface do not restrict the contact of the counter electrode, while causing excess current flow and potential current damage. Thus, the negative voltage range is preferably wide enough to function in low oxygen environments, while being limited enough to prevent the sensor from applying a voltage to the counter electrode that would cause bubble formation. The limit also provides a fail-safe mechanism for prevention of a H2 feedback loop. If hydrogen diffuses to the working electrode and creates current the counter electrode would be pushed to its electronic limit. When the potentiostat reaches the electronic limit it is no longer able to maintain the potential applied between the working and the reference electrode and the applied potential is decreased. At this point, a maximum limiting electrode current condition is attained. Additionally, the optimized potentiostat settings of the preferred embodiments provide a failsafe mechanism that prevents a cascade reaction that could cause damage to the sensor.
In one implementation of an implantable glucose sensor, a reference voltage between about +0.6V and +0.8V, and preferably about +0.7V, with respect to battery ground is chosen to ensure functionality even in low oxygen conditions yet limiting the counter electrode to a minimum of potential equal to ground potential to ensure that the sensor does not create conditions that could damage it. However, one skilled in the art appreciates that the ratio of the electroactive surface areas of the working and counter electrodes will influence the voltage operating point of the counter electrode with larger counter electrode areas requiring a less negative voltage relative to the reference electrode voltage for the same working electrode current. Additionally, one skilled in the art appreciates that optimization of the potentiostat to produce the above-described results can be attained by limitations other than on the reference voltage, for example, by limiting the current of the working or counter electrode amplifiers to a preset current limit or by setting or the op-amp offset (VOFFSET) from battery ground (see
Reference is now made to
Therefore, the preferred embodiments provide for a mechanical and electrical connection of the electrodes 54 to the PCB 48 that is extremely secure, robust, and easily reproducible in manufacture. In preferred embodiments, the electrodes 54 are swaged to the PCB 48 prior to assembling the electronics subassembly. Referring to
Swaging is a process whereby metal is shaped by hammering or pressure with the aid of a form or anvil called a swage block, substantially without heating. Notably, swaging is a solderless attachment with good mechanical accuracy, stability, orientation, and provides a quick and clean method of manufacture. The resulting swaged electrode-to-PCB connection 60 at least in part enables the reliable encapsulation of the electronics subassembly in a molded material and longevity of the device due to reliability and reproducibility of a stable electrode system 54.
In some alternative embodiments the electrodes 54 are welded to the PCB 48, which may include, for example, spot welding, laser welding, ultrasonic welding, or the like. Although these techniques include some heating, they are typically cleaner than conventional soldering techniques, for example, and may be advantageous in the some embodiments.
In an alternative embodiment, the inductor 62 may be shielded from water vapor by a metal dome, box, or cover 68, for example, such as shown in
Referring now to the configuration of the RF telemetry module of the preferred embodiments, the hardware and software are designed for low power requirements to increase the longevity of the device (for example, to enable a life of 3 to 24 months) with maximum RF transmittance from the in vivo environment to the ex vivo environment (for example, about one to ten meters). Preferably, a high frequency carrier signal in the range of 402 to 405 MHz is employed in order to maintain lower power requirements. Additionally, the carrier frequency is adapted for physiological attenuation levels, which is accomplished by tuning the RF module in a simulated in vivo environment to ensure RF functionality after implantation. Accordingly, it is believed that the preferred glucose sensor can sustain sensor function for greater than 3 months, greater than 6 months, greater than 12 months, and greater than 24 months.
In some alternative embodiments, hermetic packaging encompasses some parts of the implantable analyte sensor, while water vapor permeable packaging encompasses other parts of the implantable analyte sensor. For example, the implantable analyte sensor body may be formed from a hermetic material (such as Titanium), which encompasses the RF circuitry and/or other water vapor-sensitive components; and a water vapor permeable insert or piece may be incorporated onto the hermetic body, which encompasses the antenna and/or other non-water vapor-sensitive components operably connected thereto. In this way, the RF circuitry and other sensitive components are protected from negative effects that water vapor causes, while allowing unobstructed transmissions and receiving via the antenna.
In the preferred embodiments, a substantial portion of the electronics 46 is coated with a conformal coating 66. This conformal coating preferably has a water permeability rate that is less than the water permeability rate of the sensor body and enables sufficient spacing for the electromagnetic field as described in more detail above with reference to
In one preferred embodiment, one or more conformal Parylene coatings are applied prior to encapsulation in the sensor body. Parylene is known to have a slow water vapor permeability rate and is suitable for biomedical applications. The Parylene coating process exposes product to the gas-phase monomer at low pressure. Through vacuum deposition, Parylene condenses on the object'"'"'s surface in a polycrystalline fashion, providing a coating that is truly conformal and pinhole free. Compared to liquid processes, the effects of gravity and surface tension are negligible so there is no bridging, thin-out, pinholes, puddling, run-off or sagging; additionally, the process takes place at room temperature so there is no thermal or mechanical stress on the product. Parylene is physically stable and chemically inert within its usable temperature range. Parylene provides excellent protection from water vapor, corrosive vapors, and solvents, for example. In alternative embodiments, other conformal coatings (for example, HumiSeal®, Woodside, N.Y.), spray coatings, or the like, may be used for the less- or minimally-water vapor penetrable layer, which protect the PCB 48 and electronics subassembly 46 from damage during the molding process (
In one alternative embodiment, a coating of a secondary material, such as silicone, is applied after the protective coating 66. The secondary coating is preferably made from a material that is able to absorb mechanical stresses that may be translated from the molding process to the sensitive electrical components beneath the protective coating. Thus, silicone, or other similar material with sufficient elasticity or ductility, may be applied to the coated electronics subassembly prior to forming the sensor body, which is described in more detail below.
In one embodiment, the body of the sensor is preferably formed from a plastic material molded around the sensor electronics, however in alternative embodiments, the body may be formed from a variety of materials, including metals, ceramics, plastics, resins, or composites thereof.
It is noted that conventional prior art implantable sensors that have electronics therein generally use a hermetic material for at least a portion of the body that houses the sensitive electronics. However conventional hermetic implantable devices suffer from numerous disadvantages including: difficulty in RF transmissions through the hermetic material, seams that may allow water vapor penetration if not perfectly sealed, minimal design or shape changes without major manufacturing changes (inability to rapidly iterate on design), and need to mechanically hold and reinforce the electronics inside, increased weight and density, for example.
To overcome the disadvantages of the prior art, the preferred embodiments mold a plastic material around the electronics subassembly 46 (
Referring now to the molding process for forming the body, a two-step process is preferably used in order to provide a total seal of the components within the device and for forming a seamless device with a desired curvature thereon. The two-step molding process decreases micro-fissures that may form during the initial molding step, while mechanically securing and protecting the electrical components. However, some alternative embodiments may utilize a one-step molding process, for example by injection-molding the device.
During the primary casting process, the primary mold 70 is preferably pre-filled with a predetermined amount of a selected plastic material in order to substantially cover the lower portion 72 of the primary mold 70. The electronics subassembly 46 is then pressed into the pre-filled primary mold 70, after which the material 74 is filled around the electronics subassembly 46 to a predetermined fill line or weight amount, ensuring minimal or no air bubbles exist within the material (
During the secondary casting process, the secondary mold 80 is pre-filled with a predetermined amount of the selected material in order to substantially cover the lower portion of the secondary mold 80. The primary potted device 78 is then pressed into the pre-filled secondary mold 80, after which the material 74 is filled around the primary potted device to a predetermined fill line or weight amount, ensuring minimal or no air bubbles exist within the material. Finally a secondary hold down fixture is secured over the device. The material is cured using standard techniques known in the art, for example the plastic material may be placed into a pressure vessel and heated, or the like. It is noted that this secondary casting is advantageous because it creates additional strength over a primary casting alone, fills in micro-fissures or other damage that may have occurred during or after the primary process, and provides a seamless exterior to prevent leakage into the device.
Reference is now made to
It is noted that additional outer coatings may be advantageously applied to the secondary potted device 82, such as one or more Parylene coatings, in order to decrease water vapor penetration, for example. As another example of an outer coating, a silicone layer may be applied to the non-sensing region of the device, which serves to fill in any microfissures or micropores, for example, within the molded material, to provide a smooth outer surface, and/or to enable attachment of additional materials (for example, a silicone anchoring material). Other coatings may be applied as is appreciated by one skilled in the art.
In some alternative embodiments, the electronics subassembly 46 is placed within a pre-formed shell, rather than molding the body around the subassembly. In these alternative embodiments, the shell configuration advantageously provides air space surrounding the electronics, which aids in maintaining a stable dielectric constant surrounding the VCO circuitry, such as described in more detail with reference to
In this illustration, the analyte sensor 12 is shown without subsequent membrane and anchoring material thereon and is used to illustrate sensor geometry. The analyte sensor 12 includes the sensing region 16 located on a curved portion of the sensor body, and including no abrupt edge or discontinuous surface in the proximity of the sensing region. Additionally, the overall curvature of the surface on which the sensing region is located, including rounded edges, invokes a generally uniform FBC around that surface, decreasing inflammatory response and increasing analyte transport at the device-tissue interface.
Perpendicular forces 84, depicted in
In preferred embodiments, the sensing membrane is constructed of two or more domains and is disposed adjacent to the electroactive surfaces of the sensing region 16. The sensing membrane provides functional domains that enable measurement of the analyte at the electroactive surfaces. For example, the sensing membrane includes an enzyme, which catalyzes the reaction of the analyte being measured with a co-reactant (for example, glucose and oxygen) in order to produce a species that in turn generates a current value at the working electrode, such as described in more detail above in the Overview section. The sensing membrane can be formed from one or more distinct layers and can comprise the same or different materials.
In some embodiments, the sensing membrane 88 includes an enzyme, for example, glucose oxidase, and covers the electrolyte phase. In one embodiment, the sensing membrane 88 generally includes a resistance domain 90 most distal from the electrochemically reactive surfaces, an enzyme domain 92 less distal from the electrochemically reactive surfaces than the resistance domain, and an electrolyte domain 96 adjacent to the electrochemically reactive surfaces. However, it is understood that a sensing membrane modified for other devices, for example, by including fewer or additional domains, is within the scope of the preferred embodiments. Co-pending U.S. patent application Ser. No. 09/916,711, entitled, “SENSOR HEAD FOR USE WITH IMPLANTABLE DEVICES” and U.S. patent application Ser. No. 10/153,356 entitled, “TECHNIQUES TO IMPROVE POLYURETHANE MEMBRANES FOR IMPLANTABLE GLUCOSE SENSORS,” which are incorporated herein by reference in their entirety, describe membranes that can be used in the preferred embodiments. It is noted that in some embodiments, the sensing membrane 88 may additionally include an interference domain 94 that limits some interfering species; such as described in the above-cited co-pending patent application. Co-pending U.S. patent application Ser. No. 10/695,636, entitled, “SILICONE COMPOSITION FOR BIOCOMPATIBLE MEMBRANE” also describes membranes that may be used for the sensing membrane of the preferred embodiments, and is incorporated herein by reference in its entirety.
In some embodiments, the domains of the sensing membrane are formed from materials such as silicone, polytetrafluoroethylene, polyethylene-co-tetrafluoroethylene, polyolefin, polyester, polycarbonate, biostable polytetrafluoroethylene, homopolymers, copolymers, terpolymers of polyurethanes, polypropylene (PP), polyvinylchloride (PVC), polyvinylidene difluoride (PVDF), polybutylene terephthalate (PBT), polymethylmethacrylate (PMMA), polyether ether ketone (PEEK), polyurethanes, cellulosic polymers, polysulfones and block copolymers thereof including, for example, di-block, tri-block, alternating, random and graft copolymers, or the like.
Referring now to the function of the resistance domain 90, it is noted that there exists a molar excess of glucose relative to the amount of oxygen in blood; that is, for every free oxygen molecule in extracellular fluid, there are typically more than 100 glucose molecules present (see Updike et al., Diabetes Care 5:207-21 (1982)). However, an immobilized enzyme-based sensor employing oxygen as cofactor should be supplied with oxygen in non-rate-limiting excess in order to respond linearly to changes in glucose concentration, while not responding to changes in oxygen tension. More specifically, when a glucose-monitoring reaction is oxygen-limited, linearity is not achieved above minimal concentrations of glucose. Without a semipermeable membrane situated over the enzyme domain to control the flux of glucose and oxygen, a linear response to glucose levels can be obtained only up to about 40 mg/dL. However, in a clinical setting, a linear response to glucose levels is desirable up to at least about 500 mg/dL.
The resistance domain 90 includes a semipermeable membrane that controls the flux of oxygen and glucose to the underlying