Patent Grant
10827954
There are a number of instances when it is desirable or necessary to monitor the concentration of an analyte, such as glucose, lactate, or oxygen, for example, in bodily fluid of a body. For example, it may be desirable to monitor high or low levels of glucose in blood or other bodily fluid that may be detrimental to a human. In a healthy human, the concentration of glucose in the blood is maintained between about 0.8 and about 1.2 mg/mL by a variety of hormones, such as insulin and glucagons, for example. If the blood glucose level is raised above its normal level, hyperglycemia develops and attendant symptoms may result. If the blood glucose concentration falls below its normal level, hypoglycemia develops and attendant symptoms, such as neurological and other symptoms, may result. Both hyperglycemia and hypoglycemia may result in death if untreated. Maintaining blood glucose at an appropriate concentration is thus a desirable or necessary part of treating a person who is physiologically unable to do so unaided, such as a person who is afflicted with diabetes mellitus.
Certain compounds may be administered to increase or decrease the concentration of blood glucose in a body. By way of example, insulin can be administered to a person in a variety of ways, such as through injection, for example, to decrease that person's blood glucose concentration. Further by way of example, glucose may be administered to a person in a variety of ways, such as directly, through injection or administration of an intravenous solution, for example, or indirectly, through ingestion of certain foods or drinks, for example, to increase that person's blood glucose level.
Regardless of the type of adjustment used, it is typically desirable or necessary to determine a person's blood glucose concentration before making an appropriate adjustment. Typically, blood glucose concentration is monitored by a person or sometimes by a physician using an in vitro test that requires a blood sample. The person may obtain the blood sample by withdrawing blood from a blood source in his or her body, such as a vein, using a needle and syringe, for example, or by lancing a portion of his or her skin, using a lancing device, for example, to make blood available external to the skin, to obtain the necessary sample volume for in vitro testing. The fresh blood sample is then applied to an in vitro testing device such as an analyte test strip, whereupon suitable detection methods, such as colorimetric, electrochemical, or photometric detection methods, for example, may be used to determine the person's actual blood glucose level. The foregoing procedure provides a blood glucose concentration for a particular or discrete point in time, and thus, must be repeated periodically, in order to monitor blood glucose over a longer period.
Conventionally, a “finger stick” is generally performed to extract an adequate volume of blood from a finger for in vitro glucose testing since the tissue of the fingertip is highly perfused with blood vessels. These tests monitor glucose at discrete periods of time when an individual affirmatively initiates a test at a given point in time, and therefore may be characterized as “discrete” tests. Unfortunately, the fingertip is also densely supplied with pain receptors, which can lead to significant discomfort during the blood extraction process. Unfortunately, the consistency with which the level of glucose is checked varies widely among individuals. Many diabetics find the periodic testing inconvenient and they sometimes forget to test their glucose level or do not have time for a proper test. Further, as the fingertip is densely supplied with pain receptors which causes significant discomfort during the blood extraction process, some individuals will not be inclined to test their glucose levels as frequently as they should. These situations may result in hyperglycemic or hypoglycemic episodes.
Glucose monitoring systems that allow for sample extraction from sites other than the finger and/or that can operate using small samples of blood, have been developed. (See, e.g., U.S. Pat. Nos. 6,120,676, 6,591,125 and 7,299,082, the disclosures of each of which are incorporated herein by reference for all purposes). Typically, about one μL or less of sample may be required for the proper operation of these devices, which enables glucose testing with a sample of blood obtained from the surface of a palm, a hand, an arm, a thigh, a leg, the torso, or the abdomen. Even though less painful than the finger stick approach, these other sample extraction methods are still inconvenient and may also be somewhat painful.
In addition to the discrete, in vitro, blood glucose monitoring systems described above, at least partially implantable, or in vivo, blood glucose monitoring systems, which are designed to provide continuous or semi-continuous in vivo measurement of an individual's glucose concentration, have been described. See, e.g., U.S. Pat. Nos. 6,175,752, 6,284,478, 6,134,461, 6,560,471, 6,746,582, 6,579,690, 6,932,892 and 7,299,082, the disclosures of each of which are incorporated herein by reference for all purposes.
A number of these in vivo systems are based on “enzyme electrode” technology, whereby an enzymatic reaction involving an enzyme such as glucose oxidase, glucose dehydrogenase, or the like, is combined with an electrochemical sensor for the determination of an individual's glucose level in a sample of the individual's biological fluid. By way of example, the electrochemical sensor may be placed in substantially continuous contact with a blood source, e.g., may be inserted into a blood source, such as a vein or other blood vessel, for example, such that the sensor is in continuous contact with blood and can effectively monitor blood glucose levels. Further by way of example, the electrochemical sensor may be placed in substantially continuous contact with bodily fluid other than blood, such as dermal or subcutaneous fluid, for example, for effective monitoring of glucose levels in such bodily fluid, such as interstitial fluid.
Relative to discrete or periodic monitoring using analyte test strips, continuous monitoring is generally more desirable in that it may provide a more comprehensive assessment of glucose levels and more useful information, including predictive trend information, for example. Subcutaneous continuous glucose monitoring is also desirable as it is typically less invasive than continuous glucose monitoring in blood accessed from a blood vessel.
Regardless of the type of implantable analyte monitoring device employed, it has been observed that transient, low sensor readings which result in clinically significant sensor related errors may occur for a period of time. For example, it has been found that during the initial 12-24 hours of sensor operation (after implantation), a glucose sensor's sensitivity (defined as the ratio between the analyte sensor current level and the blood glucose level) may be relatively low—a phenomenon sometimes referred to as “early signal attenuation” (ESA). Additionally, low sensor readings may be more likely to occur at certain predictable times such as during night time use—commonly referred to as “night time drop outs”. An in vivo analyte sensor with lower than normal sensitivity may report blood glucose values lower than the actual values, thus potentially underestimating hyperglycemia, and triggering false hypoglycemia alarms.
While these transient, low readings are infrequent and, in many instances, resolve after a period of time, the negative deviations in sensor readings impose constraints upon analyte monitoring during the period in which the deviations are observed. One manner of addressing this problem is to configure the analyte monitoring system so as to delay reporting readings to the user until after this period of negative deviations passes. However, this leaves the user vulnerable and relying on alternate means of analyte measuring, e.g., in vitro testing, during this time. Another way of addressing negative deviations in sensor sensitivity is to require frequent calibration of the sensor during the time period in which the sensor is used. This is often accomplished in the context of continuous glucose monitoring devices by using a reference value after the sensor has been positioned in the body, where the reference value most often employed is obtained by a finger stick and use of a blood glucose test strip. However, these multiple calibrations are not desirable for at least the reasons that they are inconvenient and painful, as described above.
One cause of spurious low readings or drop outs by these implantable sensors is thought to be the presence of blood clots, also known as “thrombi”, formed as a result of insertion of the sensor in vivo. Such clots exist in close proximity to a subcutaneous glucose sensor and have a tendency to “consume” glucose at a high rate, thereby lowering the local glucose concentration. It may also be that the implanted sensor constricts adjacent blood vessels thereby restricting glucose delivery to the sensor site.
One approach to addressing the problem of drop outs is to reduce the size of the sensor, thereby reducing the likelihood of thrombus formation upon implantation and impingement of the sensor structure on adjacent blood vessels, and thus, maximizing fluid flow to the sensor. One manner of reducing the size or surface area of at least the implantable portion of a sensor is to provide a sensor in which the sensor's electrodes and other sensing components and/or layers are distributed over both sides of the sensor, thereby necessitating a narrow sensor profile. Examples of such double-sided sensors are disclosed in U.S. Pat. No. 6,175,752, U.S. Patent Application Publication No. 2007/0203407, now U.S. Pat. No. 7,826,879, and U.S. Provisional Application No. 61/165,499 filed Mar. 31, 2009, the disclosures of each of which are incorporated herein by reference for all purposes.
It would also be desirable to provide sensors for use in a continuous analyte monitoring system that have negligible variations in sensitivity, including no variations or at least no statistically significant and/or clinically significant variations, from sensor to sensor. Such sensors would have to lend themselves to being highly reproducible and would necessarily involve the use of extremely accurate fabrication processes.
It would also be highly advantageous to provide continuous analyte monitoring systems that are substantially impervious to, or at least minimize, spurious low readings due to the in vivo environmental effects of subcutaneous implantation, such as ESA and night-time dropouts. Of particular interest are analyte monitoring devices and systems that are capable of substantially immediate and accurate analyte reporting to the user so that spurious low readings, or frequent calibrations, are minimized or are non existent.
It would also be highly advantageous if such sensors had a construct which makes them even less invasive than currently available sensors and which further minimizes pain and discomfort to the user.
Embodiments of the present disclosure include continuous analyte monitoring systems utilizing implantable or partially implantable analyte sensors which have a relatively small profile (as compared to currently available implantable sensors). The relatively small size of the subject sensors reduce the likelihood of bleeding and, therefore, minimize thrombus formation upon implantation and the impingement of the sensor structure on adjacent blood vessels, and thus, maximizing fluid flow to the sensor and reducing the probability of ESA or low sensor readings.
In certain embodiments, the sensors are double-sided, meaning that both sides of the sensor's substrate are electrochemically functional, i.e., each side provides at least one electrode, thereby reducing the necessary surface area of the sensor. This enables the sensors to have a relatively smaller insertable distal or tail portion which reduces the in vivo environmental effects to which they are subjected. Further, the non-insertable proximal or external portion of the sensor may also have a relatively reduced size.
The subject continuous analyte monitoring systems include a skin-mounted portion or assembly and a remote portion or assembly. The skin-mounted portion includes at least the data transmitter, the transmitter battery, a portion of the sensor electronics, and electrical contacts for electrically coupling the implanted sensor with the transmitter. The remote portion of the system includes at least a data receiver and a user interface which may also be configured for test strip-based glucose monitoring. The skin-mounted portion of the system has a housing or base which is constructed to externally mount to the patient's skin and to mechanically and electrically couple the implanted sensor with the transmitter. Removably held or positioned within the housing/base structure is a connector piece having an electrical contact configuration which, when used with a double-sided sensor, enables coupling of the sensor to the transmitter in a low-profile, space-efficient manner. The skin-mounted components of the system, including the associated mounting/coupling structure, have complementary diminutive structures which, along with the very small sensor, which maximize patient usability and comfort.
Embodiments further include systems and devices for implanting the subject analyte sensors within a patient's skin and simultaneously coupling the analyte monitoring system's external, skin-mounted unit to the implanted sensor. Certain insertion systems include at least a manually-held and/or manually-operated inserter device and an insertion needle which is carried by and removably coupled to the inserter. In certain of these embodiments, only the insertion needle is disposable with the inserter or insertion gun being reusable, reducing the overall cost of the system and providing environmental advantages. In other embodiments, the skin-mounted unit and sensor are inserted manually without the use of an insertion device.
Embodiments of the subject continuous analyte monitoring systems may include additional features and advantages. For example, certain embodiments do not require individual-specific calibration by the user, and, in certain of these embodiments, require no factory-based calibration as well. Certain other embodiments of the continuous analyte monitoring systems are capable of substantially immediate and accurate analyte reporting to the user so that spurious low readings, or frequent calibrations, are minimized or are non-existent.
The subject analyte sensors usable with the subject continuous analyte monitoring systems are highly reproducible with negligible or virtually non-existent sensor-to-sensor variations with respect to sensitivity to the analyte, eliminating the need for user-based calibration. Furthermore, in certain embodiments, the analyte sensors have a predictable sensitivity drift on the shelf and/or during in vivo use are provided. Computer programmable products including devices and/or systems that include programming for a given sensor drift profile may also be provided. The programming may use the drift profile to apply a correction factor to the system to eliminate the need for user-based calibration.
These and other features, objects and advantages of the present disclosure will become apparent to those persons skilled in the art upon reading the details of the present disclosure as more fully described below.
The following patents, applications and/or publications are incorporated herein by reference for all purposes: U.S. Pat. Nos. 4,545,382; 4,711,245; 5,262,035; 5,262,305; 5,264,104; 5,320,715; 5,356,786; 5,509,410; 5,543,326; 5,593,852; 5,601,435; 5,628,890; 5,820,551; 5,822,715; 5,899,855; 5,918,603; 6,071,391; 6,103,033; 6,120,676; 6,121,009; 6,134,461; 6,143,164; 6,144,837; 6,161,095; 6,175,752; 6,270,455; 6,284,478; 6,299,757; 6,338,790; 6,377,894; 6,461,496; 6,503,381; 6,514,460; 6,514,718; 6,540,891; 6,560,471; 6,579,690; 6,591,125; 6,592,745; 6,600,997; 6,605,200; 6,605,201; 6,616,819; 6,618,934; 6,650,471; 6,654,625; 6,676,816; 6,730,200; 6,736,957; 6,746,582; 6,749,740; 6,764,581; 6,773,671; 6,881,551; 6,893,545; 6,932,892; 6,932,894; 6,942,518; 7,041,468; 7,167,818; and 7,299,082; U.S. Published Application Nos. 2004/0186365, now U.S. Pat. No. 7,811,231; 2005/0182306, now U.S. Pat. No. 8,771,183; 2006/0025662, now U.S. Pat. No. 7,740,581; 2006/0091006; 2007/0056858, now U.S. Pat. No. 8,298,389; 2007/0068807, now U.S. Pat. No. 7,846,311; 2007/0095661; 2007/0108048, now U.S. Pat. No. 7,918,975; 2007/0199818, now U.S. Pat. No. 7,811,430; 2007/0227911, now U.S. Pat. No. 7,887,682; 2007/0233013; 2008/0066305, now U.S. Pat. No. 7,895,740; 2008/0081977, now U.S. Pat. No. 7,618,369; 2008/0102441, now U.S. Pat. No. 7,822,557; 2008/0148873, now U.S. Pat. No. 7,802,467; 2008/0161666; 2008/0267823; and 2009/0054748, now U.S. Pat. No. 7,885,698; U.S. patent application Ser. No. 11/461,725, now U.S. Pat. No. 7,866,026; Ser. No. 12/131,012; 12/242,823, now U.S. Pat. No. 8,219,173; Ser. No. 12/363,712, now U.S. Pat. No. 8,346,335; Ser. No. 12/495,709; 12/698,124; and Ser. No. 12/714,439; U.S. Provisional Application Ser. Nos. 61/184,234; 61/230,686; and 61/347,754.
A detailed description of various aspects, features and embodiments of the present disclosure is provided herein with reference to the accompanying drawings, which are briefly described below. The drawings are illustrative and are not necessarily drawn to scale, with some components and features being exaggerated for clarity. The drawings illustrate various aspects and features of the present disclosure and may illustrate one or more embodiment(s) or example(s) of the present disclosure in whole or in part. A reference numeral, letter, and/or symbol that is used in one drawing to refer to a particular element or feature maybe used in another drawing to refer to a like element or feature. Included in the drawings are the following:
Before the embodiments of the present disclosure are described, it is to be understood that the present disclosure is not limited to particular embodiments described, as such may, of course, vary. It is also to be understood that the terminology used herein is for the purpose of describing particular embodiments only, and is not intended to be limiting, since the scope of the present disclosure will be limited only by the appended claims.
Where a range of values is provided, it is understood that each intervening value, to the tenth of the unit of the lower limit unless the context clearly dictates otherwise, between the upper and lower limit of that range and any other stated or intervening value in that stated range, is encompassed within embodiments of the present disclosure. The upper and lower limits of these smaller ranges may independently be included in the smaller ranges as also encompassed within embodiments of the present disclosure, subject to any specifically excluded limit in the stated range. Where the stated range includes one or both of the limits, ranges excluding either or both of those included limits are also included in the present disclosure.
Generally, embodiments of the present disclosure relate to methods and devices for detecting at least one analyte, such as glucose, in body fluid. Embodiments relate to the continuous and/or automatic in vivo monitoring of the level of one or more analytes using a continuous analyte monitoring system that includes an analyte sensor for the in vivo detection, of an analyte, such as glucose, lactate, and the like, in a body fluid. Embodiments include wholly implantable analyte sensors and analyte sensors in which only a portion of the sensor is positioned under the skin and a portion of the sensor resides above the skin, e.g., for contact to a control unit, transmitter, receiver, transceiver, processor, etc. At least a portion of a sensor may be, for example, subcutaneously positionable in a patient for the continuous or semi-continuous monitoring of a level of an analyte in a patient's interstitial fluid. For the purposes of this description, semi-continuous monitoring and continuous monitoring will be used interchangeably, unless noted otherwise. The sensor response may be correlated and/or converted to analyte levels in blood or other fluids. In certain embodiments, an analyte sensor may be positioned in contact with interstitial fluid to detect the level of glucose, which detected glucose may be used to infer the glucose level in the patient's bloodstream. Analyte sensors may be insertable into a vein, artery, or other portion of the body containing fluid. Embodiments of the analyte sensors of the subject disclosure may be configured for monitoring the level of the analyte over a time period which may range from minutes, hours, days, weeks, or longer.
Analytes that may be monitored include, but are not limited to, acetyl choline, amylase, bilirubin, cholesterol, chorionic gonadotropin, creatine kinase (e.g., CK-MB), creatine, creatinine, DNA, fructosamine, glucose, glutamine, growth hormones, hormones, ketone bodies, lactate, oxygen, peroxide, prostate-specific antigen, prothrombin, RNA, thyroid stimulating hormone, and troponin. The concentration of drugs, such as, for example, antibiotics (e.g., gentamicin, vancomycin, and the like), digitoxin, digoxin, drugs of abuse, theophylline, and warfarin, may also be monitored. In those embodiments that monitor more than one analyte, the analytes may be monitored at the same or different times.
The analyte monitoring system 100 includes a sensor 101, a data processing unit 102 connectable to the sensor 101, and a primary receiver unit 104 which is configured to communicate with the data processing unit 102 via a communication link 103. In certain embodiments, the primary receiver unit 104 may be further configured to transmit data to a data processing terminal 105 to evaluate or otherwise process or format data received by the primary receiver unit 104. The data processing terminal 105 may be configured to receive data directly from the data processing unit 102 via a communication link which may optionally be configured for bi-directional communication. Further, the data processing unit 102 may include a transmitter or a transceiver to transmit and/or receive data to and/or from the primary receiver unit 104 and/or the data processing terminal 105 and/or optionally the secondary receiver unit 106.
Also shown in
Only one sensor 101, data processing unit 102 and data processing terminal 105 are shown in the embodiment of the analyte monitoring system 100 illustrated in
The analyte monitoring system 100 may be a continuous monitoring system or semi-continuous. In a multi-component environment, each component may be configured to be uniquely identified by one or more of the other components in the system so that communication conflict may be readily resolved between the various components within the analyte monitoring system 100. For example, unique IDs, communication channels, and the like, may be used.
In certain embodiments, the sensor 101 is physically positioned in and/or on the body of a user whose analyte level is being monitored. The sensor 101 may be configured to continuously or semi-continuously sample the analyte level of the user automatically (without the user initiating the sampling), based on a programmed intervals such as, for example, but not limited to, once every minute, once every five minutes and so on, and convert the sampled analyte level into a corresponding signal for transmission by the data processing unit 102. The data processing unit 102 is coupleable to the sensor 101 so that both devices are positioned in or on the user's body, with at least a portion of the analyte sensor 101 positioned transcutaneously. The data processing unit 102 may include a fixation element such as adhesive or the like to secure it to the user's body. A mount (not shown) attachable to the user and mateable with the unit 102 may be used. For example, a mount may include an adhesive surface. The data processing unit 102 performs data processing functions, where such functions may include, but are not limited to, filtering and encoding of data signals, each of which corresponds to a sampled analyte level of the user, for transmission to the primary receiver unit 104 via the communication link 103. In one embodiment, the sensor 101 or the data processing unit 102 or a combined sensor/data processing unit may be wholly implantable under the skin layer of the user.
In certain embodiments, the primary receiver unit 104 may include a signal interface section including an radio frequency (RF) receiver and an antenna that is configured to communicate with the data processing unit 102 via the communication link 103, and a data processing section for processing the received data from the data processing unit 102 such as data decoding, error detection and correction, data clock generation, data bit recovery, etc., or any combination thereof.
In operation, the primary receiver unit 104 in certain embodiments is configured to synchronize with the data processing unit 102 to uniquely identify the data processing unit 102, based on, for example, an identification information of the data processing unit 102, and thereafter, to continuously or semi-continuously receive signals transmitted from the data processing unit 102 associated with the monitored analyte levels detected by the sensor 101. Referring again to
The data processing terminal 105 may include an infusion device such as an insulin infusion pump or the like, which may be configured to administer insulin to patients, and which may be configured to communicate with the primary receiver unit 104 for receiving, among others, the measured analyte level. Alternatively, the primary receiver unit 104 may be configured to integrate an infusion device therein so that the primary receiver unit 104 is configured to administer insulin (or other appropriate drug) therapy to patients, for example, for administering and modifying basal profiles, as well as for determining appropriate boluses for administration based on, among others, the detected analyte levels received from the data processing unit 102. An infusion device may be an external device or an internal device (wholly implantable in a user).
In certain embodiments, the data processing terminal 105, which may include an insulin pump, may be configured to receive the analyte signals from the data processing unit 102, and thus, incorporate the functions of the primary receiver unit 104 including data processing for managing the patient's insulin therapy and analyte monitoring. In certain embodiments, the communication link 103 as well as one or more of the other communication interfaces shown in
As can be seen in the embodiment of
In certain embodiments having a test strip interface 301, the interface includes a glucose level testing portion to receive a blood (or other body fluid sample) glucose test or information related thereto. For example, the interface may include a test strip port to receive a glucose test strip. The device may determine the glucose level of the test strip, and optionally display (or otherwise notice) the glucose level on the output 310 of the primary receiver unit 104. Any suitable test strip may be employed, e.g., test strips that only require a very small amount (e.g., one microliter or less, e.g., 0.5 microliter or less, e.g., 0.1 microliter or less), of applied sample to the strip in order to obtain accurate glucose information, e.g. Freestyle® and Precision® blood glucose test strips from Abbott Diabetes Care Inc. Glucose information obtained by the in vitro glucose testing device may be used for a variety of purposes, computations, etc. For example, the information may be used to calibrate sensor 101 (however, calibration of the subject sensors may not be necessary), confirm results of the sensor 101 to increase the confidence thereof (e.g., in instances in which information obtained by sensor 101 is employed in therapy related decisions), etc. Exemplary blood glucose monitoring systems are described, e.g., in U.S. Pat. Nos. 6,071,391, 6,120,676, 6,338,790 and 6,616,819, and in U.S. application Ser. No. 11/282,001, now U.S. Pat. No. 7,918,975 and Ser. No. 11/225,659, now U.S. Pat. No. 8,298,389, the disclosures of each of which are incorporated herein by reference for all purposes.
In further embodiments, the data processing unit 102 and/or the primary receiver unit 104 and/or the secondary receiver unit 106, and/or the data processing terminal/infusion section 105 may be configured to receive the blood glucose value from a wired connection or wirelessly over a communication link from, for example, a blood glucose meter. In further embodiments, a user manipulating or using the analyte monitoring system 100 (
Additional detailed descriptions are provided in U.S. Pat. Nos. 5,262,035, 5,262,305, 5,264,104, 5,320,715, 5,593,852, 6,103,033, 6,134,461, 6,175,752, 6,560,471, 6,579,690, 6,605,200, 6,654,625, 6,746,582 and 6,932,894, and in U.S. Published Patent Application Nos. 2004/0186365, now U.S. Pat. No. 7,811,231 and 2005/0182306, now U.S. Pat. No. 8,771,183, the disclosures of each of which are incorporated herein by reference for all purposes.
The sensor may be wholly implantable in a user or may be configured so that only a portion is positioned within (internal) a user and another portion outside (external) a user. For example, the sensor 400 may include a portion positionable above a surface of the skin 410, and a portion positioned below the skin. In such embodiments, the external portion may include contacts (connected to respective electrodes of the second portion by traces) to connect to another device also external to the user such as a transmitter unit. While the embodiment of
In addition to the electrodes, sensing layer and dielectric layers, sensor 500 may also include a temperature probe, a mass transport limiting layer, a biocompatible layer, and/or other optional components (none of which are illustrated). Each of these components enhances the functioning of and/or results from the sensor.
Substrate 504 may be formed using a variety of non-conducting materials, including, for example, polymeric or plastic materials and ceramic materials. (It is to be understood that substrate includes any dielectric material of a sensor, e.g., around and/or in between electrodes of a sensor such as a sensor in the form of a wire wherein the electrodes of the sensor are wires that are spaced-apart by a substrate). In some embodiments, the substrate is flexible. For example, if the sensor is configured for implantation into a patient, then the sensor may be made flexible (although rigid sensors may also be used for implantable sensors) to reduce pain to the patient and damage to the tissue caused by the implantation of and/or the wearing of the sensor. A flexible substrate often increases the patient's comfort and allows a wider range of activities. Suitable materials for a flexible substrate include, for example, non-conducting plastic or polymeric materials and other non-conducting, flexible, deformable materials. Examples of useful plastic or polymeric materials include thermoplastics such as polycarbonates, polyesters (e.g., Mylar and polyethylene terephthalate (PET)), polyvinyl chloride (PVC), polyurethanes, polyethers, polyamides, polyimides, or copolymers of these thermoplastics, such as PETG (glycol-modified polyethylene terephthalate).
In other embodiments, the sensors, or at least a portion of the sensors, are made using a relatively rigid substrate, for example, to provide structural support against bending or breaking. Examples of rigid materials that may be used as the substrate include poorly conducting ceramics, such as aluminum oxide and silicon dioxide. One advantage of an implantable sensor having a rigid substrate is that the sensor 500 may have a sharp point and/or a sharp edge to aid in implantation of a sensor without an additional insertion device. It will be appreciated that for many sensors and sensor applications, both rigid and flexible sensors will operate adequately. The flexibility of the sensor may also be controlled and varied along a continuum by changing, for example, the composition and/or thickness and/or width of the substrate (and/or the composition and/or thickness and/or width of one or more electrodes or other material of a sensor).
In addition to considerations regarding flexibility, it is often desirable that implantable sensors should have a substrate which is non-toxic. For example, the substrate may be approved by one or more appropriate governmental agencies or private groups for in vivo use.
Although the sensor substrate, in at least some embodiments, has uniform dimensions along the entire length of the sensor, in other embodiments, the substrate has a distal end or tail portion and a proximal end or body portion with different widths, respectively, as illustrated in
For subcutaneously implantable sensors which are designed for continuous or semi-continuous monitoring of the analyte during normal activities of the patient, a tail portion or distal end of the sensor which is to be implanted into the patient may have a width of about 2 mm or less, e.g., about 1 mm or less, e.g., about 0.5 mm or less, e.g., about 0.25 mm or less, e.g., about 0.15 mm or less. However, wider or narrower sensors may be used. The proximal end of the sensor may have a width larger than the distal end to facilitate the connection between the electrode contacts and contacts on a control unit, or the width may be substantially the same as the distal portion.
The thickness of the substrate may be determined by the mechanical properties of the substrate material (e.g., the strength, modulus, and/or flexibility of the material), the desired use of the sensor including stresses on the substrate arising from that use, as well as the depth of any channels or indentations that may be formed in the substrate, as discussed below. The substrate of a subcutaneously implantable sensor for continuous or semi-continuous monitoring of the level of an analyte while the patient engages in normal activities may have a thickness that ranges from about 50 μm to about 500 μm, e.g., from about 100 μm to about 300 μm. However, thicker and thinner substrates may be used.
The length of the sensor may have a wide range of values depending on a variety of factors. Factors which influence the length of an implantable sensor may include the depth of implantation into the patient and the ability of the patient to manipulate a small flexible sensor and make connections between the sensor and the sensor control unit/transmitter. A subcutaneously implantable sensor of
Electrodes 501, 502 and 503 are formed using conductive traces disposed on the substrate 504. These conductive traces may be formed over a smooth surface of the substrate or within channels formed by, for example, embossing, indenting or otherwise creating a depression in the substrate. The conductive traces may extend most of the distance along a length of the sensor, as illustrated in
The conductive traces may be formed on the substrate by a variety of techniques, including, for example, photolithography, screen printing, or other impact or non-impact printing techniques. The conductive traces may also be formed by carbonizing conductive traces in an organic (e.g., polymeric or plastic) substrate using a laser. A description of some exemplary methods for forming the sensor is provided in U.S. patents and applications noted herein, including U.S. Pat. Nos. 5,262,035, 6,103,033, 6,175,752 and 6,284,478, the disclosures of each of which are incorporated herein by reference for all purposes.
Another method for disposing the conductive traces on the substrate includes the formation of recessed channels in one or more surfaces of the substrate and the subsequent filling of these recessed channels with a conductive material. The recessed channels may be formed by indenting, embossing, or otherwise creating a depression in the surface of the substrate. Exemplary methods for forming channels and electrodes in a surface of a substrate can be found in U.S. Pat. No. 6,103,033, the disclosure of which is incorporated herein by reference for all purposes. The depth of the channels is typically related to the thickness of the substrate. In one embodiment, the channels have depths in the range of about 12.5 μm to about 75 μm, e.g., about 25 μm to about 50 μm.
The conductive traces are typically formed using a conductive material such as carbon (e.g., graphite), a conductive polymer, a metal or alloy (e.g., gold or gold alloy), or a metallic compound (e.g., ruthenium dioxide or titanium dioxide). The formation of films of carbon, conductive polymer, metal, alloy, or metallic compound are well-known and include, for example, chemical vapor deposition (CVD), physical vapor deposition, sputtering, reactive sputtering, printing, coating, and painting. In embodiments in which the conductive material is filled into channels formed in the substrate, the conductive material is often formed using a precursor material, such as a conductive ink or paste. In these embodiments, the conductive material is deposited on the substrate using methods such as coating, painting, or applying the material using a spreading instrument, such as a coating blade. Excess conductive material between the channels is then removed by, for example, running a blade along the substrate surface.
In certain embodiments, some or all of the electrodes 501, 502, 503 may be provided on the same side of the substrate 504 in the layered construction as described above, or alternatively, may be provided in a co-planar manner such that two or more electrodes may be positioned on the same plane (e.g., side-by side (e.g., parallel) or angled relative to each other) on the substrate 504. For example, co-planar electrodes may include a suitable spacing there between and/or include dielectric material or insulation material disposed between the conducting layers/electrodes. Furthermore, in certain embodiments, one or more of the electrodes 501, 502, 503 may be disposed on opposing sides of the substrate 504. Variations of such double-sided sensors are illustrated in
As noted above, analyte sensors include an analyte-responsive enzyme to provide a sensing component or sensing layer 508 proximate to or on a surface of a working electrode in order to electrooxidize or electroreduce the target analyte on the working electrode. Some analytes, such as oxygen, can be directly electrooxidized or electroreduced, while other analytes, such as glucose and lactate, require the presence of at least one component designed to facilitate the electrochemical oxidation or reduction of the analyte. The sensing layer may include, for example, a catalyst to catalyze a reaction of the analyte and produce a response at the working electrode, an electron transfer agent to transfer electrons between the analyte and the working electrode (or other component), or both.
In certain embodiments, the sensing layer includes one or more electron transfer agents. Electron transfer agents that may be employed are electroreducible and electrooxidizable ions or molecules having redox potentials that are a few hundred millivolts above or below the redox potential of the standard calomel electrode (SCE). The electron transfer agent may be organic, organometallic, or inorganic. Examples of organic redox species are quinones and species that in their oxidized state have quinoid structures, such as Nile blue and indophenol. Examples of organometallic redox species are metallocenes such as ferrocene. Examples of inorganic redox species are hexacyanoferrate (III), ruthenium hexamine etc.
In certain embodiments, electron transfer agents have structures or charges which prevent or substantially reduce the diffusional loss of the electron transfer agent during the period of time that the sample is being analyzed. For example, electron transfer agents include, but are not limited to, a redox species, e.g., bound to a polymer which can in turn be disposed on or near the working electrode. The bond between the redox species and the polymer may be covalent, coordinative, or ionic. Although any organic, organometallic or inorganic redox species may be bound to a polymer and used as an electron transfer agent, in certain embodiments the redox species is a transition metal compound or complex, e.g., osmium, ruthenium, iron, and cobalt compounds or complexes. It will be recognized that many redox species described for use with a polymeric component may also be used, without a polymeric component.
One type of polymeric electron transfer agent contains a redox species covalently bound in a polymeric composition. An example of this type of mediator is poly(vinylferrocene). Another type of electron transfer agent contains an ionically-bound redox species. This type of mediator may include a charged polymer coupled to an oppositely charged redox species. Examples of this type of mediator include a negatively charged polymer coupled to a positively charged redox species such as an osmium or ruthenium polypyridyl cation. Another example of an ionically-bound mediator is a positively charged polymer such as quaternized poly(4-vinyl pyridine) or poly(l-vinyl imidazole) coupled to a negatively charged redox species such as ferricyanide or ferrocyanide. In other embodiments, electron transfer agents include a redox species coordinatively bound to a polymer. For example, the mediator may be formed by coordination of an osmium or cobalt 2,2′-bipyridyl complex to poly(l-vinyl imidazole) or poly(4-vinyl pyridine).
Suitable electron transfer agents are osmium transition metal complexes with one or more ligands, each ligand having a nitrogen-containing heterocycle such as 2,2′-bipyridine, 1,10-phenanthroline, 1-methyl, 2-pyridyl biimidazole, or derivatives thereof. The electron transfer agents may also have one or more ligands covalently bound in a polymer, each ligand having at least one nitrogen-containing heterocycle, such as pyridine, imidazole, or derivatives thereof. One example of an electron transfer agent includes (a) a polymer or copolymer having pyridine or imidazole functional groups and (b) osmium cations complexed with two ligands, each ligand containing 2,2′-bipyridine, 1,10-phenanthroline, or derivatives thereof, the two ligands not necessarily being the same. Some derivatives of 2,2′-bipyridine for complexation with the osmium cation include, but are not limited to, 4,4′-dimethyl-2,2′-bipyridine and mono-, di-, and polyalkoxy-2,2′-bipyridines, such as 4,4′-dimethoxy-2,2′-bipyridine. Derivatives of 1,10-phenanthroline for complexation with the osmium cation include, but are not limited to, 4,7-dimethyl-1,10-phenanthroline and mono, di-, and polyalkoxy-1,10-phenanthrolines, such as 4,7-dimethoxy-1,10-phenanthroline. Polymers for complexation with the osmium cation include, but are not limited to, polymers and copolymers of poly(l-vinyl imidazole) (referred to as “PVI”) and poly(4-vinyl pyridine) (referred to as “PVP”). Suitable copolymer substituents of poly(l-vinyl imidazole) include acrylonitrile, acrylamide, and substituted or quaternized N-vinyl imidazole, e.g., electron transfer agents with osmium complexed to a polymer or copolymer of poly(l-vinyl imidazole).
Embodiments may employ electron transfer agents having a redox potential ranging from about −200 mV to about +200 mV versus the standard calomel electrode (SCE).
As mentioned above, the sensing layer may also include a catalyst which is capable of catalyzing a reaction of the analyte. The catalyst may also, in some embodiments, act as an electron transfer agent. When the analyte of interest is glucose, a catalyst such as a glucose oxidase, glucose dehydrogenase (e.g., pyrroloquinoline quinone (PQQ), dependent glucose dehydrogenase or oligosaccharide dehydrogenase, flavine adenine dinucleotide (FAD) dependent glucose dehydrogenase, or nicotinamide adenine dinucleotide (NAD) dependent glucose dehydrogenase) may be used. A lactate oxidase or lactate dehydrogenase may be used when the analyte of interest is lactate. Laccase may be used when the analyte of interest is oxygen or when oxygen is generated or consumed in response to a reaction of the analyte.
In certain embodiments, a catalyst may be attached to a polymer, cross linking the catalyst with another electron transfer agent which, as described above, may be polymeric. A second catalyst may also be used in certain embodiments. This second catalyst may be used to catalyze a reaction of a product compound resulting from the catalyzed reaction of the analyte. The second catalyst may operate with an electron transfer agent to electrolyze the product compound to generate a signal at the working electrode. Alternatively, a second catalyst may be provided in an interferent-eliminating layer to catalyze reactions that remove interferents.
Certain embodiments include a Wired Enzyme™ sensing layer (such as used in the FreeStyle Navigator® continuous glucose monitoring system by Abbott Diabetes Care Inc.) that works at a gentle oxidizing potential, e.g., a potential of about +40 mV. This sensing layer uses an osmium (Os)-based mediator designed for low potential operation and is stably anchored in a polymeric layer. Accordingly, in certain embodiments the sensing element is redox active component that includes (1) Osmium-based mediator molecules attached by stable (bidente) ligands anchored to a polymeric backbone, and (2) glucose oxidase enzyme molecules. These two constituents are crosslinked together.
In certain embodiments, the sensing system detects hydrogen peroxide to infer glucose levels. For example, a hydrogen peroxide-detecting sensor may be constructed in which a sensing layer includes enzymes such as glucose oxidase, glucose dehydrogenase, or the like, and is positioned proximate to the working electrode. The sensing layer may be covered by one or more layers, e.g., a membrane that is selectively permeable to glucose. Once the glucose passes through the membrane, it may be oxidized by the enzyme and reduced glucose oxidase can then be oxidized by reacting with molecular oxygen to produce hydrogen peroxide.
Certain embodiments include a hydrogen peroxide-detecting sensor constructed from a sensing layer prepared by crosslinking two components together, for example: (1) a redox compound such as a redox polymer containing pendent Os polypyridyl complexes with oxidation potentials of about +200 mV vs. SCE, and (2) periodate oxidized horseradish peroxidase (HRP). Such a sensor functions in a reductive mode; the working electrode is controlled at a potential negative to that of the Os complex, resulting in mediated reduction of hydrogen peroxide through the HRP catalyst.
In another example, a potentiometric sensor can be constructed as follows. A glucose-sensing layer is constructed by crosslinking together (1) a redox polymer containing pendent Os polypyridyl complexes with oxidation potentials from about −200 mV to +200 mV vs. SCE, and (2) glucose oxidase. This sensor can then be used in a potentiometric mode, by exposing the sensor to a glucose containing solution, under conditions of zero current flow, and allowing the ratio of reduced/oxidized Os to reach an equilibrium value. The reduced/oxidized Os ratio varies in a reproducible way with the glucose concentration, and will cause the electrode's potential to vary in a similar way.
The components of the sensing layer may be in a fluid or gel that is proximate to or in contact with the working electrode. Alternatively, the components of the sensing layer may be disposed in a polymeric or sol-gel matrix that is proximate to or on the working electrode. Preferably, the components of the sensing layer are non-leachably disposed within the sensor. More preferably, the components of the sensor are immobilized within the sensor.
Examples of sensing layers that may be employed are described in U.S. patents and applications noted herein, including, e.g., in U.S. Pat. Nos. 5,262,035, 5,264,104, 5,543,326, 6,605,200, 6,605,201, 6,676,819 and 7,299,082, the disclosures of each of which are incorporated herein by reference for all purposes.
Regardless of the particular components that make up a given sensing layer, a variety of different sensing layer configurations may be used. In certain embodiments, the sensing layer covers the entire working electrode surface, e.g., the entire width of the working electrode surface. In other embodiments, only a portion of the working electrode surface is covered by the sensing layer, e.g., only a portion of the width of the working electrode surface. Alternatively, the sensing layer may extend beyond the conductive material of the working electrode. In some cases, the sensing layer may also extend over other electrodes, e.g., over the counter electrode and/or reference electrode (or counter/reference is provided), and may cover all or only a portion thereof.
In other embodiments the sensing layer is not deposited directly on the working electrode. Instead, the sensing layer may be spaced apart from the working electrode, and separated from the working electrode, e.g., by a separation layer. A separation layer may include one or more membranes or films or a physical distance. In addition to separating the working electrode from the sensing layer the separation layer may also act as a mass transport limiting layer, and/or an interferent eliminating layer, and/or a biocompatible layer.
In certain embodiments which include more than one working electrode, one or more of the working electrodes may not have a corresponding sensing layer, or may have a sensing layer which does not contain one or more components (e.g., an electron transfer agent and/or catalyst) needed to electrolyze the analyte. Thus, the signal at this working electrode may correspond to background signal which may be removed from the analyte signal obtained from one or more other working electrodes that are associated with fully-functional sensing layers by, for example, subtracting the signal.
Whichever configuration of the sensing component or layer is employed, at least one factor in minimizing variations in sensor sensitivity, at least within the same sensor batch or lot (or all sensors made according to the same specification), is by strictly maintaining the dimensions (width, length, diameter and thickness) of the active area, i.e., the area of the working electrode in contact with the sensing component or layer, from sensor to sensor. Optimizing sensitivity, including reproducing substantially the same sensitivity for sensors within a lot or batch of sensors, reduces and in certain embodiments eliminates the need for sensor calibration, by the user. Accordingly, sensors that do not require a user to calibrate, using for example an in vitro test strip or the like after insertion of the sensor into the body for testing, are achieved. Examples of sensors for use in one or more embodiments of the present disclosure can be found in, among others, U.S. patent application Ser. No. 12/714,439, the disclosure of which is incorporated herein by reference for all purposes.
Calibration, when an electrochemical glucose sensor is used, generally involves converting the raw current signal (nA) into a glucose concentration (mg/dL). One way in which this conversion is done is by relating or equating the raw analyte signal with a calibration measurement (i.e., with a reference measurement), and obtaining a conversion factor (raw analyte signal/reference measurement value). This relationship is often referred to as the sensitivity of the sensor, which, once determined, may then be used to convert sensor signals to calibrated analyte concentration values, e.g., via simple division (raw analyte signal/sensitivity=calibrated analyte concentration). For example, a raw analyte signal of 10 nA could be associated with a calibration analyte concentration of 100 mg/dL, and thus, a subsequent raw analyte signal of 20 nA could be converted to an analyte concentration of 200 mg/dL, as may be appropriate for a given analyte, such as glucose, for example.
There are many ways in which the conversion factor may be obtained. For example, the sensitivity factor can be derived from a simple average of multiple analyte signal/calibration measurement data pairs, or from a weighted average of multiple analyte signal/calibration measurement data pairs. Further by way of example, the sensitivity may be modified based on an empirically derived weighting factor, or the sensitivity may be modified based on the value of another measurement, such as temperature. It will be appreciated that any combination of such approaches, and/or other suitable approaches, is contemplated herein.
For subcutaneous glucose sensors, calibration at the site of manufacture, that may be relied upon to calibrate sensor signal for the useful life of a sensor, presents numerous challenges to the feasibility. This infeasibility may be based on any of a number of factors. For example, variations in the within-lot sensitivity of the analyte sensors and/or variations in sensor drift may be too great.
The present disclosure provides sensor embodiments which attempt to address both the in vivo environmental effects and the manufacturing-based inconsistencies which can lead to variation in sensor sensitivity, and/or which obviate the need for any form of calibration, whether at the factory or by the user, at anytime prior to or during operative use of the sensor.
Certain of these sensor embodiments are double-sided, i.e., both sides of the sensor's substrate are electrochemically functional, with each side providing at least one electrode. Because both sides of the sensor are utilized, the smaller the necessary surface area required per side to host the electrodes. This space-efficient construct allows the sensor to be miniaturized and much smaller than conventional sensors, and, in particular, have a relatively narrower tail portion, i.e., at least the portion of a sensor that is constructed to be positioned beneath a skin surface of a user is miniaturized. A narrower structure reduces trauma to tissue at the implantation site, thereby reducing bleeding and the production of thrombi around the sensor. The smaller structure also minimizes impingement upon adjacent blood vessels. The smaller width of the sensor allows, in addition to perpendicular diffusion of the analyte (e.g., glucose), for the lateral diffusion of analyte molecules towards the active sensing area. These effects substantially if not completely eliminate spurious low readings.
In addition to providing micro tail sections, these double-sided sensors are designed and configured to be highly reproducible. Further, they may be fabricated by methods, techniques and equipment which minimize inconsistencies in the registration, deposition and resolution of the sensor components, as described herein.
Referring now to
Sensor tail portion 600 includes a substrate 602 (see
One of the top or bottom conductive layers, here, top conductive layer 604a, serves as the sensor's working electrode. The opposing conductive layer, here, bottom conductive layer 604b, serves as a reference and/or counter electrode. Where conductive layer 604b serves as either a reference or counter electrode, but not both, a third electrode may optionally be provided on a surface area of the proximal portion of the sensor (not shown). For example, conductive layer 604b may serve as reference electrode and a third conductive trace (not shown), present only on the non-implantable proximal portion of the sensor, may serve as the sensor's counter electrode.
Disposed over a distal portion of the length of conducting layer/working electrode 604a is sensing component or layer 606. Providing the sensing layer closer to the distal tip of the sensor places the sensing material in the best position for contact with the analyte-containing fluid. As only a small amount of sensing material is required to facilitate electrooxidization or electroreduction of the analyte, positioning the sensing layer 606 at or near the distal tip of the sensor tail reduces the amount of material needed. Sensing layer 606 may be provided in a continuous stripe/band between and substantially orthogonal to the substrate's side edges 614a, 614b with the overlap or intersection of working electrode 604a and the sensing layer 606 defining the sensor's active area. Due to the orthogonal relationship between sensing layer 606 and conductive layer 604a, the active area has a rectilinear polygon configuration; however, any suitable shape may be provided. The dimensions of the active area may be varied by varying either or both of the respective width dimensions of the sensing and conducting layers. The width WS of the sensing layer 606 may cover the entire length of the working electrode or only a portion thereof. As the width WC of the conductive layer is dictated by the width of the tail portion's substrate in this embodiment, any registration or resolution inconsistencies between the conductive layer and the substrate are obviated. In certain embodiments, the width of the sensing layer WS is in the range from about 0.05 mm to about 5 mm, e.g., from about 0.1 mm to about 3 mm; the width of the conductive layer WC is in the range from about 0.05 mm to about 0.6 mm, e.g., from about 0.1 mm to about 0.3 mm, with the resulting active area in the range from about 0.0025 mm2 to about 3 mm2, e.g., from about 0.01 mm2 to about 0.9 mm2.
Referring again to the electrodes, the same materials and methods may be used to make the top and bottom electrodes, although different materials and methods may also be used. With the working and reference electrodes positioned on opposing sides of the substrate as in the illustrated embodiment of
Selection of the conductive materials for the respective electrodes is based in part on the desired rate of reaction of the sensing layer's mediator at an electrode. In some instances the rate of reaction for the redox mediator at the counter/reference electrode is controlled by, for example, choosing a material for the counter/reference electrode that would require an overpotential or a potential higher than the applied potential to increase the reaction rate at the counter/reference electrode. For example, some redox mediators may react faster at a carbon electrode than at a silver/silver chloride (Ag/AgCl) or gold electrode. However, as Ag/AgCl and gold are more expensive than carbon, it may be desirous to use the former materials judiciously.
The sensor embodiment of
Referring again to sensor 600, an insulation/dielectric layer 608a, 608b is disposed on each side 600a, 600b of the sensor, over at least the sensor's body portion (not shown), to insulate the proximal portion of the electrodes, i.e., the portion of the electrodes which in part remains external to the skin upon implantation. The top dielectric layer 608a disposed on the working electrode 604a may extend distally to but preferably not over any portion of sensing layer 606. Alternatively, as illustrated in
As for the dielectric layer on the bottom/reference electrode side of the sensor, it may extend any suitable length of the sensor's tail section, i.e., it may extend the entire length of both of the primary and secondary conductive layers or portions thereof. For example, as illustrated in
Finally, one or more membranes, which may function as one or more of an analyte flux modulating layer and/or an interferent-eliminating layer and/or biocompatible layer, discussed in greater detail below, may be provided about the sensor, e.g., as one or more of the outermost layer(s). In certain embodiments, as illustrated in
Based on current sensor fabrication techniques, provision of the sensor's conductive layers can be accomplished more accurately than provision of the sensing layers. As such, improving upon the accuracy of providing the sensing component on the sensor, and thus, the accuracy of the resulting active area, may significantly decrease any sensor to sensor sensitivity variability and obviate the need for calibration of the sensor. Accordingly, the present disclosure also includes methods for fabricating such analyte sensors having accurately defined active areas. Additionally, the methods provide finished sensors which are smaller than currently available sensors with micro-dimensioned tail portions which are far less susceptible to the in situ environmental conditions which can cause spurious low readings.
In one variation of the subject methods, web-based manufacturing techniques are used to perform one or more steps in fabricating the subject sensors, many of the steps of which are disclosed in U.S. Pat. No. 6,103,033. To initiate the fabrication process, a continuous film or web of substrate material is provided and heat treated as necessary. The web may have precuts or perforations defining the individual sensor precursors. The various conductive layers are then formed on the substrate web by one or more of a variety of techniques as described above, with the working and reference (or counter/reference) electrode traces provided on opposite sides of the web. As mentioned previously, the electrode traces may be provided in channels formed in the surface of the substrate material; however, with the desire to provide a sensor having a tail portion that has the smallest functional profile possible, and particularly with the sensor tail having two functional sides, the use of channels may not be optimal as it requires a thicker substrate material. Also, as mentioned previously, a third, optional electrode trace (which may function as a counter electrode, for example) may be provided on the proximal body portion of the sensor precursors. The “primary” conductive traces provided on the area of the tail portions of the precursor sensors have a width dimension greater than the intended width dimension of the tail portions of the finalized sensors. The precursor widths of these conductive traces may range from about 0.3 mm to about 10 mm including widths in range from about 0.5 mm to about 3 mm, or may be even narrower. The primary conductive layers are formed extending distally along the tail section of the sensor precursors to any suitable length, but preferably extend at least to the intended distal edge of the finalized sensors to minimize the necessary sensor tail length.
Next, the sensing layer and secondary conductive layers, if employed, are formed on the primary conductive layers on the respective sides of the substrates or substrate web. As discussed, each of these layers is preferably formed in a stripe or band of the respective material disposed orthogonally to the length of the primary conductive layer/sensor tail. With a single, continuous deposition process, the mean width of the sensing strip is substantially constant along the substrate webbing, and ultimately, from sensor to sensor. The secondary conductive layer (e.g., Ag/AgCl on the reference electrode), if provided, may also be formed in a continuous orthogonal stripe/band with similar techniques. One particular method of providing the various stripes/band of material on the sensors is by depositing, printing or coating the sensing component/material by means of an inkjet printing process (e.g., piezoelectric inkjet as manufactured by Scienion Inc. and distributed by BioDot Inc.). Another way of applying these materials is by means of a high precision pump (e.g., those which are piston driven or driven by peristaltic motion) and/or footed needle. The respective stripes/bands may be provided over a webbing of sequentially aligned sensor precursors prior to singulation of the sensors or over a plurality of sensors/electrodes where the sensors have been singulated from each other prior to provision of the one or more stripes/bands.
With both the sensing and conductive layers/strips having substantially constant widths and provided substantially orthogonal to each other, the active area which their intersection forms is also substantially constant along both the length and width of the sensor. In such embodiments, the active area (as well as the intersecting area of the primary and secondary conductive layers which form the reference electrode) has a rectilinear polygonal shape which may be easier to provide in a reproducible manner from sensor to sensor; however, any relative arrangement of the layers resulting in any suitable active area geometry may be employed.
The sensor precursors, i.e., the template of substrate material (as well as the conductive and sensing materials if provided on the substrate at the time of singulation), may be singulated from each other using any convenient cutting or separation protocol, including slitting, shearing, punching, laser singulation, etc. These cutting methods are also very precise, further ensuring that the sensor's active area, when dependent in part on the width of the sensor (i.e., the tail portion of the substrate), has very accurate dimensions from sensor to sensor. Moreover, with each of the materials (i.e., the primary and secondary conductive materials, sensing component, dielectric material, membrane, etc.) provided with width and/or length dimensions extending beyond the intended dimensions or boundaries of the final sensors, issues with resolution and registration of the materials is minimized if not obviated altogether.
The final, singulated, double-sided sensor structures have dimensions in the following ranges: widths from about 500 μm to about 100 including widths in range from about 300 μm to about 150 μm; tail lengths from about 10 mm to about 3 mm, including lengths in range from about 6 mm to about 4 mm; and thicknesses from about 500 μm to about 100 including thicknesses in range from about 300 μm to about 150 μm. As such, the implantable portions of the sensors are reduced in size from conventional sensors by approximately 20% to about 80% in width as well as in cross-section. The reduced size minimizes bleeding and thrombus formation upon implantation of the sensor and impingement on adjacent tissue and vessels, thereby minimizing impediment to lateral diffusion of the analyte to the sensor's sensing component or sensing layer.
The substrate web may have precuts or perforations that provide guidance for the final cut employed to singulate the precursors. Depending on the layout and orientation of the sensor precursors, the singulation lines may be at fixed or varying intervals. For example, if the orientation and spacing of the sensor precursors are serial and constant over the area of the substrate material, the singulation lines will typically be at fixed intervals in all directions. However, where the sensors having irregular or asymmetrical shapes (e.g., as illustrated in
Embodiments include sensor lots having very low variations in sensitivity of sensors within the lot. Low sensitivity variation enables sensors that do not require calibration by a user after a sensor is positioned in the body. Accordingly, in certain embodiments, sensor lots are provided that have a coefficient of variation (CV) of about 5% or less, e.g., about 4.5% or less, e.g., about 4% or less, e.g., about 3% or less.
Sensors having predictable sensor in vivo sensitivity and signal are provided. For example, sensors having predictable shelf life sensitivity drift (the period of time between manufacture and use) and predictable in vivo sensitivity drift, including substantially no shelf and in vivo sensitivity drift, are also provided. In embodiments in which sensors have drift (e.g., where the sensor sensitivity drifts an expected percentage over a certain time), a drift profile is contemplated. This drift profile may be contemplated by an algorithm of the monitoring system to determine a drift correction factor that may be applied to sensor signal to obtain a glucose measurement (mg/dL). Due, at least in part, to the high reproducibility of the manufacturing process that results in low manufacturing coefficient of variation (CV), a single drift correction factor may be used for all sensors of a given sensor manufacturing lot or batch.
In certain embodiments, sensor sensitivity may be determined post-fabrication by the manufacturer at the site of manufacture. This “factory-determined” sensitivity may then be used in an algorithm to calibrate sensor signal for the useable lifetime of the sensor, negating the need for a user to obtain a reference value, e.g., from a test strip, for calibration. Sensitivity may include determining the relationship of sensor signal to a reference such as an in vitro reference (a known glucose level to which one or more sensors of a sensor lot may be compared). Sensitivity may include determining a conversion factor as described herein. In certain embodiments, the determined sensitivity may be further augmented. For example, one or more additional factors (e.g., to account for the relationship of blood to subcutaneous tissue glucose, effect of oxygen, temperature, etc.) may be contemplated. In any event, a sensitivity value is determined. Exemplary calibration protocols are described, e.g., in U.S. Pat. No. 7,299,082, the disclosure of which is incorporated herein by reference for all purposes.
Because the sensitivities of each sensor of a given manufacturing lot are substantially the same according to the embodiments herein, the factory-determined sensitivity may be applied to all sensors of such a lot, i.e., a single calibration algorithm may be used for all the sensors of a given lot. In one embodiment, the information is programmed or is programmable into software of the monitoring system, e.g., into one or more processors. For example, the factory-determined sensitivity may be provided to a user with a sensor(s) and uploaded to a calibration algorithm manually or automatically (e.g., via bar code and reader, or the like). Calibration of sensor signal may then be implemented using suitable hardware/software of the system.
A mass transport limiting layer or membrane, e.g., an analyte flux modulating layer, may be included with the sensor to act as a diffusion-limiting barrier to reduce the rate of mass transport of the analyte, for example, glucose or lactate, into the region around the working electrodes. The mass transport limiting layers are useful in limiting the flux of an analyte to a working electrode in an electrochemical sensor so that the sensor is linearly responsive over a large range of analyte concentrations. Mass transport limiting layers may include polymers and may be biocompatible. A mass transport limiting layer may provide many functions, e.g., biocompatibility and/or interferent-eliminating, etc.
A membrane may be formed by crosslinking in situ a polymer, modified with a zwitterionic moiety and a non-pyridine copolymer component. The modified polymer may be made from a precursor polymer containing heterocyclic nitrogen groups. For example, a precursor polymer may be polyvinylpyridine or polyvinylimidazole. Embodiments also include membranes that are made of a polyurethane, or polyether urethane, or chemically related material, or membranes that are made of silicone, and the like.
Optionally, another moiety or modifier that is either hydrophilic or hydrophobic, and/or has other desirable properties, may be used to “fine-tune” the permeability of the resulting membrane to an analyte of interest. Optional hydrophilic modifiers, such as poly(ethylene glycol), hydroxyl or polyhydroxyl modifiers, may be used to enhance the biocompatibility of the polymer or the resulting membrane.
The membrane may also be formed in situ by applying an alcohol-buffer solution of a crosslinker and a modified polymer over an enzyme-containing sensing layer and allowing the solution to cure for about one to two days or other appropriate time period. The crosslinker-polymer solution may be applied to the sensing layer by placing a droplet or droplets of the solution on the sensor, by dipping the sensor into the solution, or the like. Generally, the thickness of the membrane is controlled by the concentration of the solution, by the number of droplets of the solution applied, by the number of times the sensor is dipped in the solution, or by any combination of these factors. A membrane applied in this manner may have any combination of the following functions: (1) mass transport limitation, i.e., reduction of the flux of analyte that can reach the sensing layer, (2) biocompatibility enhancement, or (3) interferent reduction. Exemplary mass transport layers are described in U.S. patents and applications noted herein, including, e.g., in U.S. Pat. Nos. 5,593,852, 6,881,551 and 6,932,894, the disclosures of each of which are incorporated herein by reference for all purposes.
A sensor may also include an active agent such as an anticlotting and/or antiglycolytic agent(s) disposed on at least a portion a sensor that is positioned in a user. An anticlotting agent may reduce or eliminate the clotting of blood or other body fluid around the sensor, particularly after insertion of the sensor. Examples of useful anticlotting agents include heparin and tissue plasminogen activator (TPA), as well as other known anticlotting agents. Embodiments may include an antiglycolytic agent or precursor thereof. Examples of antiglycolytic agents are glyceraldehyde, fluoride ion, and mannose.
The electrochemical sensors of the present disclosure may employ any suitable measurement technique, e.g., may detect current, may employ potentiometry, etc. Techniques may include, but are not limited to, amperometry, coulometry, and voltammetry. In some embodiments, sensing systems may be optical, colorimetric, and the like.
The subject analyte measurement systems may include an optional alarm system that, e.g., based on information from a processor, warns the patient of a potentially detrimental condition of the analyte. For example, if glucose is the analyte, an alarm system may warn a user of conditions such as hypoglycemia and/or hyperglycemia and/or impending hypoglycemia, and/or impending hyperglycemia. An alarm system may be triggered when analyte levels approach, reach or exceed a threshold value. An alarm system may also, or alternatively, be activated when the rate of change, or acceleration of the rate of change, in analyte level increase or decrease approaches, reaches or exceeds a threshold rate or acceleration. A system may also include system alarms that notify a user of system information such as battery condition, calibration, sensor dislodgment, sensor malfunction, etc. Alarms may be, for example, auditory and/or visual. Other sensory-stimulating alarm systems may be used including alarm systems which heat, cool, vibrate, or produce a mild electrical shock when activated.
The subject disclosure also includes sensors used in sensor-based drug delivery systems. The system may provide a drug to counteract the high or low level of the analyte in response to the signals from one or more sensors. Alternatively, the system may monitor the drug concentration to ensure that the drug remains within a desired therapeutic range. The drug delivery system may include one or more (e.g., two or more) sensors, a processing unit such as a transmitter, a receiver/display unit, and a drug administration system. In some cases, some or all components may be integrated in a single unit. A sensor-based drug delivery system may use data from the one or more sensors to provide necessary input for a control algorithm/mechanism to adjust the administration of drugs, e.g., automatically or semi-automatically. As an example, a glucose sensor may be used to control and adjust the administration of insulin from an external or implanted insulin pump.
Referring now to
As illustrated in
Insertion device 1000 comprises a body 1002 having a distal base portion 1008 having a bottom surface configured for placement on the skin surface 1005. It is noted that the figures show, with solid drawing lines, components of the insertion device and the analyte monitoring system that would otherwise not be visible when positioned or housed within device body 1002 for purposes of illustration and ease of description. For example, in
With reference to
To commence the sensor insertion/transmitter mounting procedure, the front edge 906e (see
As mentioned previously, a battery may be provided within the transmitter housing to power the transmitter 902 as well as to provide the necessary electrical signals to sensor 906. The battery may be rechargeable/replaceable through a door (not shown) provided in the transmitter housing. To minimize the size of the on-skin unit, the battery may be relatively small, having only a moderately-lasting charge, e.g., about 3-14 days more or less. In another variation, the battery is not rechargeable or replaceable, but is disposed of along with the transmitter upon expiration of the battery charge. As this arrangement is more expensive, having a battery/transmitter that has a longer-lasting charge, e.g., about 6 months to a year may be necessary; of course, the tradeoff being a larger unit. Still yet, the transmitter may be extensively reusable with the battery being disposable along with the sensor upon expiration of the sensor's useful life, typically, between about 3 to about 14 days, in which case, the battery may be very small to last only as long as the sensor.
All of the on-skin portions of the subject continuous monitoring systems have a very low-profile configuration. While certain embodiments have at least one dimension that is extremely small, other dimensions may be slightly greater to provide the necessary volume to house the various components of the on-skin units. For example, an on-skin unit may have a very low height dimension, but have relatively greater width and length dimensions. On the other hand, the width/length dimensions may be very small with the height being relatively greater. The optimal dimensions of a particular on-skin unit may depend on where on the body the unit is intended to be mounted. One exemplary set of dimensions for an on-skin unit of the present disclosure includes a width from about 7.5 to about 8.5 mm, a length from about 10 to about 11 mm, and a height from about 2.5 to about 3.3 mm.
Exemplary analyte monitoring systems are described in, for example, U.S. patent application Ser. No. 12/698,124 entitled “Compact On-Body Physiological Monitoring Devices and Methods Thereof” and in U.S. patent application Ser. No. 12/730,193 entitled “Methods of Treatment and Monitoring Systems for Same”, the disclosures of each of which are incorporated herein by reference for all purposes. Exemplary methods and systems for inserting a an analyte sensor are described in, for example, U.S. Pat. No. 6,990,366, U.S. patent application Ser. Nos. 12/698,124, 12/698,129, now U.S. Pat. No. 9,402,544, and U.S. Provisional Application Nos. 61/238,159, 61/238,483 and 61/249,535, the disclosures of each of which are incorporated herein by reference for all purposes.
Although the subject sensors may be inserted anywhere in the body, it is often desirable that the insertion site be positioned so that the on-skin sensor control unit can be concealed. In addition, it is often desirable that the insertion site be at a place on the body with a low density of nerve endings to reduce the pain to the patient. Examples of preferred sites for insertion of the sensor and positioning of the on-skin sensor control unit include the abdomen, thigh, leg, upper arm, and shoulder.
In one embodiment, the subject sensors are injected between 2 to 12 mm into the interstitial tissue of the patient for subcutaneous implantation. Preferably, the sensor is injected 3 to 9 mm, and more preferably 5 to 7 mm, into the interstitial tissue. Other embodiments of the present disclosure may include sensors implanted in other portions of the patient, including, for example, in an artery, vein, or organ. The depth of implantation varies depending on the desired implantation target. Sensor insertion angles usually range from about 10° to about 90°, typically from about 15° to about 60°, and often from about 30° to about 45°. The construct of the insertion device, of course, will vary depending on the desired angle of insertion.
In one embodiment, a continuous analyte measurement system may include a base unit configured for mounting on a skin surface, an analyte sensor comprising two functional sides, a proximal portion configured for positioning within the base unit and a distal portion configured for insertion into the skin surface, and a conductive member positionable within the base unit and in electrical contact with the two functional sides of analyte sensor.
The proximal portion of the analyte sensor may have a planar configuration and the conductive member may be mechanically and electrically coupled to the two functional sides of the proximal portion of the analyte sensor.
The base unit may be compressible on opposing sides at least about the conductive member.
Furthermore, the system may include a component for compressing the opposing ends of the conductive member.
In one aspect, the component for compressing may be flexures on opposing sides of the base unit about the conductive member.
In another aspect, the component for compressing may be a clamping fixture positionable on opposing sides of the base unit about the conductive member.
In one aspect, the system may include an alignment pin extending through the proximal portion of the analyte sensor.
The base unit may be a non-conductive compressible material.
The non-conductive compressible material may be silicone.
The conductive connector may be a conductive compressible material.
The conductive compressible material may be carbon-doped silicone.
In a further aspect, the system may include a transmitter configured for mounting to the base unit in a low-profile manner, wherein the base unit includes a pair of receptacles for receiving a corresponding pair of conductive pins of the transmitter, and the conductive pins contact the conductive member when the transmitter is operatively mounted to the base unit.
The transmitter may mount with the base unit in a side-by-side configuration.
The transmitter may mount atop the base unit.
The transmitter may house a battery.
The base unit may house a battery.
Moreover, the base unit may include a second pair of receptacles for receiving a corresponding second pair of conductive pins of the transmitter, wherein the conductive pins contact the battery when the transmitter is operatively mounted to the base unit.
The base unit may include a cradle therein for receiving and holding the conductive member.
The cradle may compress opposing ends of the conductive member when held within the cradle.
The conductive member may include a conductive core and an insulating shell covering the conductive core.
In one aspect, the conductive member may include a non-conductive inner member within the conductive core, wherein the non-conductive inner member extends through an opening in the analyte sensor.
The base unit may include an adhesive bottom for adhering to the skin surface.
The base unit may include an opening therein through which the distal end of the analyte sensor extends.
The distal end of the analyte sensor may extend along a sidewall of the base unit.
Regarding methodology, the subject methods may include each of the mechanical and/or activities associated with use of the devices described. As such, methodology implicit to the use of the devices described forms part of the present disclosure. Other methods may focus on fabrication of such devices. The methods that may be performed according to embodiments herein and that may have been described above and/or claimed below, the operations have been described in selected typographical sequences. However, the sequences have been selected and so ordered for typographical convenience and are not intended to imply any particular order for performing the operations.
As for other details of the present disclosure, materials and alternate related configurations may be employed as within the level of those with skill in the relevant art. The same may hold true with respect to method-based aspects of the present disclosure in terms of additional acts as commonly or logically employed. In addition, though embodiments of the present disclosure have been described in reference to several examples, optionally incorporating various features, the present disclosure is not to be limited to that which is described or indicated as contemplated with respect to each variation of the present embodiments. Various changes may be made to the embodiments described and equivalents (whether recited herein or not included for the sake of some brevity) may be substituted without departing from the true spirit and scope of the present disclosure. Any number of the individual parts or subassemblies shown may be integrated in their design. Such changes or others may be undertaken or guided by the principles of design for assembly.
Also, it is contemplated that any optional feature of the inventive variations described may be set forth and claimed independently, or in combination with any one or more of the features described herein. Reference to a singular item, includes the possibility that there are plural of the same items present. More specifically, as used herein and in the appended claims, the singular forms “a,” “an,” “said,” and “the” include plural referents unless the specifically stated otherwise. In other words, use of the articles allow for “at least one” of the subject item in the description above as well as the claims below. It is further noted that the claims may be drafted to exclude any optional element. As such, this statement is intended to serve as antecedent basis for use of such exclusive terminology as “solely,” “only” and the like in connection with the recitation of claim elements, or use of a “negative” limitation. Without the use of such exclusive terminology, the term “comprising” in the claims shall allow for the inclusion of any additional element—irrespective of whether a given number of elements are enumerated in the claim, or the addition of a feature could be regarded as transforming the nature of an element set forth in the claims. Stated otherwise, unless specifically defined herein, all technical and scientific terms used herein are to be given as broad a commonly understood meaning as possible while maintaining claim validity.
In all, the breadth of the present disclosure is not to be limited by the examples provided.
The present application is a continuation of U.S. patent application Ser. No. 12/842,013 filed Jul. 22, 2010, now U.S. Pat. No. 9,795,326, which claims priority under 35 U.S.C. § 119(e) to U.S. Provisional Application No. 61/227,967 filed Jul. 23, 2009, entitled “Continuous Analyte Measurement Systems and Systems and Methods for Implanting Them”, the disclosures of each of which are incorporated herein by reference for all purposes.
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| Number | Date | Country | |
|---|---|---|---|
| 20180035931 A1 | Feb 2018 | US |
| Number | Date | Country | |
|---|---|---|---|
| 61227967 | Jul 2009 | US |
| Number | Date | Country | |
|---|---|---|---|
| Parent | 12842013 | Jul 2010 | US |
| Child | 15789942 | US |
