Single package automated drug delivery system

Abstract

Disclosed herein is a combination automated insulin delivery system and continuous glucose monitor integrated into a single wearable package that can use any combination of delivery and detection methods, including cannula, microneedle array, and transdermal patch, and detection methods including electrochemical, photofluorescence, and spectroscopic methods.

Description

[Technical field]

[0001] Related Applications This application claims the benefit of U.S. Provisional Patent Application No. 63 / 315,179, filed March 1, 2022, the contents of which are incorporated by reference in their entirety herein. [Background technology]

[0002] Many conventional drug delivery systems are known, including, for example, a wearable drug delivery device of the type shown in FIG. 2. The drug delivery device 102 can be designed to deliver any type of liquid drug to a user. In certain embodiments, the drug delivery device 102 can be, for example, an OmniPod® drug delivery device manufactured by Insulet Corporation of Acton, Massachusetts. The drug delivery device 102 can be a drug delivery device such as those described in U.S. Pat. No. 7,303,549, U.S. Pat. No. 7,137,964, and U.S. Pat. No. 6,740,059, the contents of each of which are incorporated herein by reference in their entirety. Summary of the Invention [Problem to be solved by the invention]

[0003] A typical application of drug delivery systems is the delivery of insulin. Many insulin delivery systems can be interfaced with a continuous glucose monitor that communicates blood glucose levels to a controller. Current systems typically require at least two separate units: a continuous glucose monitor and an insulin delivery device, which are worn on the user's body in different locations, in addition to a separate controller device. The need for two separate units can be cumbersome for the user to install and uncomfortable for the user to wear. As such, this configuration is less than ideal for the user's quality of life. Additionally, some users have a small body "area" due to scars, age, build, low body fat, etc., making it difficult to wear multiple devices at once.

[0004] In a wearable on-body device, it is desirable to make the wearable unit as small as possible to minimize the impact on the wearer. It would therefore be desirable to replace prior art insulin delivery systems having separate insulin delivery devices and glucose sensors with a single unit in which the automatic drug delivery device and glucose sensor are adjacent or co-located in a single housing. [Means for solving the problem]

[0005] definition As used herein, the term "liquid drug" should be construed to include any drug in liquid form that can be administered by a drug delivery device via a variety of delivery means, including, for example, insulin, GLP-1, pramlintide, glucagon, morphine, blood pressure medications, chemotherapy drugs, fertility medications, etc., or co-formulations of two or more of GLP-1, pramlintide, and insulin. The term "insulin" as used herein should be construed to mean any liquid drug.

[0006] As used herein, the term "single package" should be taken to mean a product that is preassembled at the time of manufacture and presented to the end user as one product, or a product that is comprised of two or more sub-products that are assembled at the time of use by the user. The product may also be referred to herein as a "single unit" or "single housing."

[0007] overview This Summary is provided to introduce some of the concepts described below in a simplified form in the Detailed Description. This Summary is not intended to identify key features or essential features of the claimed subject matter, nor is it intended as an aid in determining the scope of the claimed subject matter.

[0008] Various embodiments of the present invention are disclosed herein in which an automatic insulin delivery (AID) device and a continuous glucose monitor (CGM) are co-located in one housing footprint. This eliminates the need for multiple devices to be placed in multiple locations on the user's body and the user does not have to engage in extra installation steps. The solution herein describes various continuous glucose sensing methods or embodiments and an insulin delivery system in which any combination of one or more of these methods or embodiments can be combined together in a single or integrated device with a control algorithm to form an AID system. The combination of a CGM and a drug delivery means in a single package is referred to herein as a single package insulin delivery system (SPID) or a single package automated insulin delivery (SPAID) system. As used herein, the term "SPAID" is meant to include non-automated embodiments (i.e., SPID). Any of the exemplary glucose monitoring methods listed in the table below can be combined with any of the exemplary insulin delivery methods. [Table 1]

[0009] Specific details regarding the system architecture and implementation of each technique are described further below. In various embodiments of the invention, the CGM may comprise an electrochemical method using a needle or flexible cannula. In other embodiments, the CGM may comprise a photofluorescence method using a needle or cannula or an implanted target. In yet other embodiments, the CGM may comprise a spectroscopy method using a needle or cannula or proximal monitoring through the user's skin.

[0010] In various embodiments, the insulin infusion method can use one or more of a cannula, a microneedle, a transdermal method, a tethered infusion set, or any other means for insulin infusion, or any combination of the foregoing.

[0011] Any of the above methods may be combined together in a single unit, or in multiple units that are physically combined in use to form a single unit. System architectures are further described herein as well as specific details regarding the implementation of each technique.

[0012] In the drawings, like reference characters generally refer to the same parts throughout the different views.In the following description, various embodiments of the present invention are described with reference to the following drawings, in which: [Brief description of the drawings]

[0013] [Figure 1] FIG. 1 is a functional block diagram of an example system suitable for implementing the systems and methods disclosed herein.

[0014] [Diagram 2] FIG. 2 is a diagram of a prior art wearable drug delivery device.

[0015] [Figure 3A-3B] 3A-3B show an embodiment in which glucose sensing and insulin delivery are combined in a single cannula.

[0016] [Figure 4A-4B] 4A-4B show an embodiment in which glucose sensing and insulin delivery are combined in a single cannula that has a closed end and delivers insulin through a series of fine perforations.

[0017] [Figure 5A-5B] 5A-5B illustrate an embodiment in which insulin delivery is provided through a first cannula configured with a hole or cross hole, and glucose sensing is provided through a second cannula.

[0018] [Figure 6A-6B] 6A-6B show an embodiment in which the electrochemical cell employs a two-electrode configuration consisting of one working electrode and one counter / reference electrode.

[0019] [Figure 7A-7B] 7A-7B show an embodiment in which the electrochemical cell employs a three-electrode configuration consisting of one working electrode, one counter electrode, and one reference electrode.

[0020] [Figure 8] FIG. 8 shows an embodiment in which the electrochemical cell employs a four-electrode configuration consisting of a main working electrode, an additional working electrode, one reference electrode, and one counter electrode.

[0021] [Figure 9] FIG. 9 shows an embodiment employing a first cell consisting of a first working electrode, a first counter electrode, and a first reference electrode, and a second cell consisting of a second working electrode, a second counter electrode, and a second reference electrode.

[0022] [Figure 10] FIG. 10 illustrates an embodiment in which glucose is indirectly measured at the working electrode via the generated hydrogen peroxide.

[0023] [Figure 11]FIG. 11 illustrates an embodiment in which glucose is measured indirectly at the working electrode via a mediator.

[0024] [Figure 12] FIG. 12 illustrates an embodiment in which glucose is measured indirectly at the working electrode via a direct electron transfer (DET) type glucose-reactive enzyme.

[0025] [Figure 13] FIG. 13 illustrates an embodiment in which glucose is measured at the working electrode by direct electron transfer.

[0026] [Figures 14A-14C] 14A-14C show various embodiments of a sensing element in which the sensing electrodes are located on the outside of the cannula.

[0027] [Figures 15A-15C] 15A-15C show various embodiments of a sensing cannula in which the sensing electrodes are configured as bands wrapped around the outside of the cannula.

[0028] [Figure 16] FIG. 16 is a system diagram showing components for use in the electrochemical cell method of glucose detection.

[0029] [Figure 17] FIG. 17 is a diagram illustrating a method for direct glucose-to-light conversion in which a receptor fluorophore is bound to an acceptor molecule and the glucose receptor of the acceptor molecule is not bound to glucose.

[0030] [Figure 18] FIG. 18 shows the energy excitation, energy loss, and energy release states of a receptor fluorophore in the presence of Forster resonance energy transfer (FRET) quenching.

[0031] [Figure 19]Figure 19 shows the direct glucose-to-light conversion method, where a receptor fluorophore and a receptor glucose molecule are bound to an acceptor molecule. *Drawing the direct glucose-to-light conversion method.

[0032] [Figure 20] FIG. 20 shows the energy excitation, energy loss, and energy release states of a receptor fluorophore in the absence of FRET quenching.

[0033] [Figure 21] FIG. 21 shows an indirect glucose-photoconversion method in which a receptor fluorophore and a receptor oxygen molecule are bound to an acceptor molecule.

[0034] [Figure 22] FIG. 22 shows an indirect glucose-photoconversion method in which a receptor fluorophore is bound to an acceptor molecule and the oxygen receptor of the acceptor molecule is not bound to an oxygen molecule.

[0035] [Figure 23] FIG. 23 shows one possible configuration of a glucose sensor using optical fluorescence sensing with the sensing element positioned so that excitation and emission light are transmitted through the skin.

[0036] [Figure 24] FIG. 24 illustrates one possible configuration of a glucose sensor using optical fluorescence sensing, where the sensor coating is placed on the surface of a cannula, and the excitation and emission lights are transmitted through the skin.

[0037] [Figure 25A] FIG. 25A illustrates one possible configuration of a glucose sensor using optical fluorescence sensing, in which a sensor coating is disposed on the surface of a cannula, and excitation and emission light are transmitted through light guides within the cannula. [Figure 25B]FIG. 25B illustrates one possible configuration of a glucose sensor using optical fluorescence sensing, in which a sensor coating is disposed on the surface of a cannula, and excitation and emission light are transmitted through light guides within the cannula. [Figure 25C] FIG. 25C illustrates one possible configuration of a glucose sensor using optical fluorescence sensing, in which a sensor coating is disposed on the surface of a cannula, and excitation and emission light are transmitted through light guides within the cannula.

[0038] [Figure 26A] FIG. 26A illustrates one possible embodiment of a glucose sensor using optical fluorescence sensing, in which multiple sensing areas are provided as a sensor coating on the outer surface of the sensing element to provide measurements of different concentrations of glucose. [Figure 26B] FIG. 26B illustrates one possible embodiment of a glucose sensor using optical fluorescence sensing, in which multiple sensing areas are provided as a sensor coating on the outer surface of the sensing element to provide measurements of different concentrations of glucose. [Figure 26C] FIG. 26C illustrates one possible embodiment of a glucose sensor using optical fluorescence sensing, in which multiple sensing areas are provided as a sensor coating on the outer surface of the sensing element to provide measurements of different concentrations of glucose.

[0039] [Figure 27A-27B] 27A-27B are diagrams of a variation of the embodiment of FIGS. 26A-26C in which an additional sensing area is provided on the outer surface of the sensing element to provide a reference oxygen level.

[0040] [Figure 28] FIG. 28 is a system diagram showing components for use with a photofluorescence method of glucose detection.

[0041] [Figure 29]FIG. 29 is a diagram of an embodiment using implanted photofluorescent seeds as sensing elements, and further where excitation and emission light is transmitted through the skin.

[0042] [Diagram 30] FIG. 30 is a diagram of an embodiment in which glucose sensing is performed by an implantable glucose sensor that wirelessly communicates glucose measurements.

[0043] [Diagram 31] FIG. 31 illustrates one embodiment of a microneedle-based sensing system that includes two three-electrode electrochemical cells.

[0044] [Diagram 32] FIG. 32 illustrates one embodiment of a hollow microneedle-based sensing system including a three-electrode electrochemical cell.

[0045] [Diagram 33] FIG. 33 illustrates one embodiment of a hollow microneedle-based sensing system that includes a photofluorescent sensing array.

[0046] [Diagram 34] FIG. 34 illustrates one embodiment of a hollow microneedle-based sensing system including first and second optical fluorescent sensing arrays.

[0047] [Diagram 35] FIG. 35 illustrates one embodiment of a hollow microneedle-based sensing system that includes a spectroscopy-based sensing array.

[0048] [Diagram 36] FIG. 36 illustrates one embodiment of a hollow microneedle-based sensing system including a spectroscopy-based sensing array that allows for orientation of analyte molecules.

[0049] [Figure 37]FIG. 37 shows a typical plot of resonant glucose energy after glucose is excited with near-infrared light energy.

[0050] [Fig. 38A-38D] 38A-38D show several embodiments using the microneedle sensing method, with different embodiments showing different configurations of the microneedle array.

[0051] [Figure 39] FIG. 39 illustrates an embodiment in which glucose sensing is performed using a spectroscopy-based method in which electromagnetic (EM) energy is directed through the skin.

[0052] [Diagram 40] FIG. 40 illustrates an embodiment in which glucose sensing is performed using a spectroscopy-based method in which secondary molecules are excited with EM energy via a cannula inserted into the skin.

[0053] [Diagram 41] FIG. 41 is a system diagram illustrating one possible limitation of a glucose sensing system using spectroscopy through a cannula.

[0054] [Fig. 42A-42B] 42A-42B show two possible configurations for the cannula tip of the system of FIG.

[0055] [Diagram 43] FIG. 43 illustrates an embodiment utilizing a cannula for delivering liquid medication into the user's body.

[0056] [Figure 44A] FIG. 44A illustrates an embodiment utilizing a microneedle array to deliver a liquid drug into a user's body. [Figure 44B] FIG. 44B illustrates an embodiment utilizing a microneedle array to deliver a liquid drug into the user's body.

[0057] [Diagram 45] FIG. 45 illustrates an embodiment of a delivery system utilizing a microneedle array, where each microneedle is covered with a grid of individual conductive plate segments that are individually connected to control electronics.

[0058] [Diagram 46] FIG. 46 illustrates an embodiment utilizing a hollow microneedle-based delivery system in which diffusion pressure is used to facilitate transport of insulin.

[0059] [Figure 47] FIG. 47 illustrates an embodiment utilizing a hollow microneedle-based delivery system in which iontophoresis is used to enhance the delivery of insulin.

[0060] [Figure 48] FIG. 48 illustrates an embodiment utilizing a hollow microneedle-based delivery system to enhance insulin delivery using iontophoresis and sonophoresis.

[0061] [Figure 49A] FIG. 49A illustrates an embodiment that utilizes a transdermal patch to deliver liquid medication into the user's body. [Figure 49B] FIG. 49B illustrates an embodiment that utilizes a transdermal patch to deliver liquid medication into the user's body.

[0062] [Figure 50] FIG. 50 illustrates an embodiment utilizing a transdermal-based delivery system in which iontophoresis is used to drive the transport of insulin.

[0063] [Fig. 51A-51D]51A-51D show several embodiments utilizing a transdermal-based delivery system, with different embodiments showing different arrangements of multiple polar conductors and glucose sensing regions.

[0064] [Figure 52] FIG. 52 is a schematic diagram of one possible implementation of a circuit used for electroporation.

[0065] [Diagram 53] FIG. 53 is a diagram illustrating an embodiment showing one possible location of electroporation electrodes in an electroporation system.

[0066] [Figure 54A] FIG. 54A illustrates an embodiment utilizing a transdermal-based delivery system in which sonophoresis is used to enhance the transport of insulin.

[0067] [Figure 54B] FIG. 54B illustrates an embodiment utilizing a transdermal-based delivery system in which sonophoresis is used to drive insulin transport and multiple ultrasound transducers are arranged in a ring around the insulin delivery port. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0068] Detailed Description The present disclosure presents various systems, components, and methods for transferring a drug from a drug delivery device to a patient, such as via a needle, cannula, microneedle array, or transdermal patch, in combination with an analyte sensing device, such as a glucose sensor. The embodiments described herein offer one or more advantages over conventional prior art systems, components, and methods, namely, a smaller overall footprint of the combined drug delivery and sensor device.

[0069] Various embodiments of the present invention include systems and methods for delivering an appropriate amount of medication to a user using a drug delivery device (sometimes referred to herein as a "pod"), either autonomously or according to a wireless signal received from an electronic device. In various embodiments, the electronic device may include a user device that comprises a smartphone, a smart watch, a smart necklace, smart jewelry, a module attached to the drug delivery device, or any other type or kind of electronic device that may be carried by or worn on the user's body and that executes an algorithm to calculate the time and dosage of delivery of the medication.

[0070] For example, the user device, or more preferably, the drug delivery device itself, may execute an "artificial pancreas" algorithm that calculates the time of delivery and dosage of insulin based on sensor data, such as an analyte sensor, such as a glucose sensor. The user device, or the drug delivery device itself, may also be in communication with or integrated with one or more additional sensors, such as an accelerometer, a position sensor, a heart rate sensor, a temperature sensor, a heart rate sensor (which can detect heart rate variability), an oxygen saturation sensor, or other sensors that collect data regarding the user's physical attributes or condition. The sensors are placed in or on the user's body and may be part of the drug delivery device or may be separate devices. In an exemplary embodiment, one or more analyte sensors (a glucose sensor as an example) are incorporated into the drug delivery device. The additional analyte sensors may be external to the drug delivery device and in communication with it. The output from the analyte sensor and the additional sensors, if present, may be used as inputs to an algorithm to calculate the time of drug delivery and dosage.

[0071] The drug delivery device may communicate with the sensor instead of or in addition to communication between the sensor and the user device. This communication may be direct (e.g., if the sensor is integrated with or otherwise part of the drug delivery device) or remote / wireless (e.g., if the sensor is located in a different housing than the drug delivery device). In these embodiments, the drug delivery device includes computing hardware (e.g., processor, memory, firmware, etc.) that executes some or all of the algorithms that calculate the time of delivery and dosage of the drug without relying on a remote user device to execute the algorithm or perform such calculations.

[0072] 1 is a functional block diagram of an exemplary closed-loop drug delivery system 100 suitable for implementing the systems and methods described herein. The drug delivery system 100 may implement (and / or provide functionality for) drug delivery algorithms, such as an artificial pancreas (AP) application, to govern or control the automated delivery of drugs or medications, such as insulin, to a user (e.g., to maintain euglycemia, which is a normal level of blood glucose). The drug delivery system 100 may be an automated drug delivery system that may include a drug delivery device 102 (which may be wearable or body mountable), one or more integrated analyte sensors 108, one or more external analyte sensors 109 (also wearable or body mountable) that can communicate with the SPAID 102 via a communication link 195, and a user device 105.

[0073] The drug delivery system 100 may also include one or more accessory devices 106, in optional examples, such as primary and / or secondary smartphones, smart watches, personal assistant devices, inhalers, drug delivery pens, etc., which may communicate with other components of the system 100 via either wired or wireless communication links 192, 193. The accessory device(s) may be used by the user or by another human being, such as, for example, a parent or a medical professional.

[0074] Single Package AID In various exemplary embodiments, the SPAID 102 may be configured to deliver various dosages of medication to the user according to calculations performed by an artificial pancreas (AP) application or drug delivery algorithm (MDA) 129 stored in memory 123 and executed by the controller 121. Alternatively, the SPAID 102 may receive instructions to deliver dosages of liquid medication from an external device, for example, a user app 160 executing on the user device 105 and communicated to the SPAID 102 via a communication link 194.

[0075] The SPAID 102 may use one or more positive displacement pump systems, such that the drive mechanism 125 operates to translate a plunger longitudinally through the reservoir 124, forcing the liquid medication through the outlet fluid port to the patient interface 186. In alternative embodiments, a motor may be used to deliver the liquid medication. In alternative embodiments, the SPAID 102 may include an optional second reservoir 124-2 and a second drive mechanism 125-2, which allows for independent delivery of two different liquid medications. As an example, the reservoir 124 may be filled with insulin, and the reservoir 124-2 may be filled with, for example, pramlintide, GLP-1, or glucagon. In some embodiments, each of the reservoirs 124, 124-2 may be configured with a separate drive mechanism 125, 125-2, which may be separately controllable by the controller 121 under the direction of the MDA 129 or via a signal received from the user device 105. Both reservoirs 124, 124-2 may be connected to a common patient interface 186. The reservoirs 124, 124-2 may be configured to store a drug, agent or therapeutic suitable for automated delivery, such as insulin, pramlintide, GLP-1, co-formulations of insulin and GLP-1, glucagon, morphine, blood pressure medications, chemotherapy medications, fertility medications, and the like.

[0076] The SPAID 102 may be configured with one or more integrated analyte sensors 108 or one or more external analyte sensors 109, which may be, for example, a continuous glucose monitor. The analyte sensors 108, 109 may be configured to detect a number of different analytes, such as glucose, lactate, ketones, uric acid, sodium, potassium, alcohol levels, etc., and output detection results, such as measurements. The analyte sensors 108, 109 may be controllable by the MDA 129 to measure blood glucose levels at predetermined time intervals, such as every 5 minutes, every minute, or in an ad hoc manner, in an exemplary embodiment. Measurements from the analyte sensors 108 may be communicated to the MDA 129 for use in determining the dosage and timing of administration of a liquid medication to the user. Additionally or alternatively, measurements from the analyte sensors 108, 109 may be communicated to the user device 105 via the communication link 194 for use by the user app 160. The analyte sensor 108 is coupled to a sensor patient interface 110, which serves to interface with the patient's body to obtain the information necessary to provide the measurements. Various embodiments of the present invention include various types of analyte sensors 108, 109 and various types of sensor patient interfaces 110, which are described in detail herein. In the description herein, references to an integrated analyte sensor 108 also generally apply to an external analyte sensor 109. In some embodiments, a sensor insertion mechanism 111 can be provided for inserting the sensor patient interface 110. In some embodiments, the sensor patient interface can be, for example, a cannula, needle, or other sensing element. The sensor insertion mechanism 111 can be comprised of an actuator that inserts a sensing element of the sensor patient interface 110 under the skin of the user, in some embodiments.

[0077] 1 as being integral with the drug delivery device 102, in various non-SPAID embodiments, the analyte sensor 108 and drug delivery device 102 may be contained in separate housings and applied to the user's body at different locations. In another embodiment, the one or more analyte sensors 108 and the sensor patient interface 110 may be contained in separate housings that are attachable to the housing of the SPAID 102.

[0078] The SPAID 102 includes a delivery patient interface 186 for interfacing with a user to deliver liquid drug(s) from the reservoir 124 (or in the case of two reservoirs, from either or both of the reservoirs 124, 124-2). The delivery patient interface 186 may be, for example, a needle or cannula for delivering the drug into the user's body (which may be subcutaneous, intraperitoneal, or intravenous). Alternatively, the delivery patient interface 186 may be one or more arrays of microneedles. In yet another alternative embodiment, the delivery patient interface 186 may be a transdermal delivery method. Various types of delivery patient interfaces 186 are described in more detail later in this specification. In the case where the delivery patient interface 186 is a cannula / needle, the SPAID 102 further includes a delivery insertion mechanism 188 for inserting the cannula into the user's body, which may be integral with the SPAID 102 or may be attachable to the SPAID 102. The delivery insertion mechanism 188, in one embodiment, may consist of an actuator that inserts a needle and cannula under the user's skin and then retracts the needle, leaving the cannula in place. The delivery insertion mechanism 188 may be combined with the sensor insertion mechanism 111 such that one mechanism delivers both the sensor and the cannula / needle in one step or multiple steps.

[0079] The SPAID 102 may optionally be configured with a user interface 127 that provides means for receiving input from a user and for outputting information to a user. The user interface 127 may include, for example, a light emitting diode, a button on the housing of the SPAID 102, a sound transducer, a microdisplay, a microphone, an accelerometer for detecting device movement or user gestures (e.g., tapping the housing of the SPAID 102), or any other type of interface device configured to allow a user to input information and / or the SPAID 102 to output information for presentation to a user (e.g., an alarm signal, etc.).

[0080] In one embodiment, the SPAID 102 includes a communication interface 126, which may be a transceiver operating according to one or more radio frequency protocols, such as Bluetooth, Wi-Fi, near field communication, cellular, etc. The controller 121 may communicate with the user device 105, the attached device 106, and the cloud-based service 111, for example, via the communication interface 126.

[0081] In some embodiments, the SPAID 102 may include one or more other integrated sensors 184 and one or more other external sensors 185 that may communicate with the SPAID 102 via a communication link 196. The sensors 184, 185 may include, for example, one or more of an acceleration sensor, a position sensor, a heart rate sensor, a temperature sensor, a heart rate sensor (which may detect heart rate variability), an oxygen saturation sensor, or other sensors, a pressure sensor, a power sensor, etc. that are communicatively coupled to the controller 121 and provide various signals to the MDA 129 or the user app 160. For example, the pressure sensor may be configured to provide an indication of a fluid pressure detected in a fluid path between the patient interface 186 and the reservoir 124 for purposes of identifying an occlusion or improper insertion by the delivery insertion mechanism 188. The pressure sensor may be coupled to or integrated with the insertion mechanism 188 for inserting the patient interface 186 into a user. In one example, the controller 121 may be operable to determine a drug infusion rate based on the indication of the fluid pressure. The drug infusion rate may be compared to an infusion rate reference or threshold, and the comparison results may be usable in determining remaining insulin (IOB), total daily insulin (TDI), or whether an occlusion has occurred.

[0082] The SPAID 102 further includes a power source 128, such as a battery, a piezoelectric device, an energy harvesting device, or the like, for powering all components of the SPAID 102. The power source 128 may be removable or rechargeable.

[0083] The SPAID 102 may be configured to perform the processes necessary to deliver a dose of liquid medication to a user without input from the user device 105 or optional accessory device 106. As will be described in more detail, the MDA 129 may be operable to determine, for example, the amount of insulin to be delivered, residual insulin (IOB), insulin remaining in the reservoir 124, etc., and cause the controller 121 to actuate the drive mechanism 125 (125-2) to deliver the liquid medication from the reservoir 124 (124-2). The MDA 129 may take as input data received from the analyte sensors 108, 109 or other sensors 184, 185.

[0084] The SPAID 102 may be a wearable device and may be attached to the body of a user, such as a diabetic, at a mounting location and may deliver any therapeutic agent, including any drug or medicine, such as insulin, to the user at or near the mounting location. The surface of the SPAID 102 may include an adhesive to facilitate attachment to the user's skin. The adhesive may be a pad that is spot welded to the SPAID 102.

[0085] User Device The user device 105 may be a computing device such as a smartphone, tablet, personal diabetes management (PDM) device, dedicated diabetes care management device, wearable device, etc. In an exemplary embodiment, the user device 105 may include a processor 151, a device memory 153, a user interface 158, and a communication interface 154. The processor 151 may execute software-based processes stored in the device memory 153, such as user applications 160, to manage the user's blood glucose levels and to control the delivery of liquid medication to the user via the SPAID 102, as well as to provide other functions such as calculating carbohydrate correction doses, correction bolus doses, etc., as described below. The user device 105, as well as the optional smart accessory device 106, may be used to program, adjust settings, and / or control the operation of the SPAID 102.

[0086] The user app 160 may be an application operable to calculate dosages of medication to deliver to the user based on information received from the analyte sensor 108 of the SPAID 102, other sensors 184, the cloud-based service 111, and / or the user device 105 or optional accessory device 106. The memory 153 may also store software for operating, for example, the user interface 158 (e.g., touch screen device, camera, etc.), the communication interface 154, etc. When executing the user app 160, the processor 151 may be configured to implement displays and notifications related to food intake, blood glucose measurements, etc. The user interface 158 is under the control of the processor 151 and may be configured to present a graphical user interface that allows for input of food announcements, adjustment of setting selections, etc., as described herein.

[0087] In a specific example, if the user application 160 is an AP application, the processor 151 may also be configured to execute a diabetes treatment plan (which may be stored in the memory) managed by the user application 160. In addition to the above-mentioned functions, if the user application 160 is an AP application, it may further provide functions to determine a carbohydrate correction dose, a correction bolus dose, and determine a basal dose according to the diabetes treatment plan. Furthermore, as an AP application, the user application 160 provides functions to output a signal to the SPAID 102 via the communication interface 194 to deliver the determined bolus dose and basal dose.

[0088] The user device 105 may further comprise one or more output devices 155, which may be, for example, a speaker or a vibration transducer, to provide various signals, e.g., alarms, to the user under the control of the user app 160. The user app 160 may also have access to other equipment of the user device 105, e.g., a GPS unit, a clock, a camera, etc., which may provide the user app 160 with information useful for calculating dosage information of liquids.

[0089] Attached Devices The optional accessory device 106 may be a wearable device, such as, for example, a smart watch (e.g., Apple Watch®), smart glasses, smart jewelry, a GPS-enabled wearable, a wearable fitness device, smart clothing, etc. Additionally, the accessory device 106 may be an inhaler or drug delivery pen that can communicate information regarding the amount of insulin delivered to the user to other components of the SPAID 102. The MDA 129 can use this information to adjust the IOB and then adjust the bolus and basal deliveries accordingly.

[0090] Similar to the user device 105, the accessory device 106 may also be configured to perform various functions, including control of the SPAID 102. For example, the accessory device 106 may include a communication interface 174, a processor 171, a user interface 178, and a memory 173. The user interface 178 may be a graphical user interface displayed on a touch screen display of the smart accessory device 106. The memory 173 may store programming code for operating various functions of the smart accessory device 106, as well as instances of user apps 160, or reduced versions of user apps 160 with reduced functionality. In some embodiments, the accessory device 106 may also include various types of sensors. In some embodiments, the user apps 160 running on the accessory device 106 may work in conjunction with the user apps 160 running on the user device 105 to accept input from a user or provide alerts or alarms to the user.

[0091] Cloud-based services The drug delivery system 100 may communicate with or receive data or services from a cloud-based service 111. Services provided by the cloud-based service 111 may include data storage to store personal or anonymized data, such as blood glucose measurements, past IOB or TDI, previous carbohydrate correction doses, and other data. Additionally, the cloud-based service 111 may process anonymized data from multiple users to provide generalized information related to TDI, insulin sensitivity, IOB, and the like. The communication link 115 coupling the cloud-based service 111 to each device 102, 105, 106 of the system 100 may be a cellular link, a Wi-Fi link, a Bluetooth link, or a combination thereof. In some embodiments, the cloud-based service 111 may include a cloud-based processor that may be used by the drug delivery system 100 to replace or supplement the processor 151 or 171, or the controller 121.

[0092] Communication Links The wireless communication links 115 and 192-196 may be any type of wireless link operating using a known wireless communication standard or a proprietary standard. By way of example, the wireless communication links 192-194 may provide, via their respective communication interfaces 126, 154, 174, a communication link based on Bluetooth, Zigbee, Wi-Fi, a short-range wireless communication standard, a cellular standard, or other wireless protocol.

[0093] calibration A drug delivery system 100 using a combined SPAID 102 combining an insulin / drug delivery system and a CGM may require calibration to maintain accurate monitoring and subsequent control of a user's blood glucose level. The system 100 may be capable of accepting a baseline blood glucose value taken by the user and entered for use by the system 100, for example, via a user app 160 running on the user device 105 or on the accessory device 106.

[0094] In addition to having the user manually input the blood glucose reference value, another possible method for calibration may be an active SPAID 102 to communicate with additional CGM devices that the user is currently using or that may be newly applied to the user's body. If the user is actively wearing an additional CGM device or smartwatch device, such as an accessory device 106, it may be possible for the SPAID 102 to communicate with that device at startup or periodically throughout its usage cycle to collect a reference blood glucose value, establish a startup reference calibration, and / or check or verify its accuracy during runtime.

[0095] An additional use model for SPAID102 is to have the user install a new SPAID102 before the expiration date of the previous SPAID102. The system 100 can warn or notify the user that the previous SPAID102 is nearing the end of its usable period and to install and use the next SPAID102. At this time, the previous SPAID102 can communicate directly by wireless communication or indirectly with the user device 105 or the accessory device 106 acting as an intermediary between the two SPAIDs. The two SPAIDs are then worn and both CGMs in the SPAIDs are active for a short overlap period that is sufficient to calibrate the new CGM. The previous SPAID102 shares its current CGM data with the new SPAID102 to calibrate the CGM of the new SPAID102 and verify its accuracy before terminating its use. The new SPAID102 will only start administering insulin once it receives confirmation that the previous SPAID102 has been removed or disabled and is no longer able to deliver insulin. This is necessary so that both SPAIDs will administer insulin together to prevent the user's blood sugar level from getting too low.

[0096] Example of operation In an operational example, the user application 160 implements a graphical user interface that is the primary interface with the user, collects information required to operate the SPAID 102 (e.g., age, weight, or other information), programs the basal profile, provides use of the bolus calculator for operation in manual mode, starts the SPAID 102, triggers the insertion of the needle / cannula or delivery patient interface 186 using the delivery insertion mechanism 188, triggers the insertion of the sensor patient interface 110 using the sensor insertion mechanism 111, and / or operates in an automatic mode (e.g., a closed-loop mode, or a hybrid of manual and closed-loop where the user can provide some instructions or inputs (e.g., carbohydrate counting or administration of a correction bolus) to the SPAID 102 via the user device 105 or the accessory device 106).

[0097] The user apps 160 provide a graphical user interface 158 that can use large text, graphics, and on-screen instructions to prompt the user through the set-up process and use of the system 100. By way of example, the user apps 160 can also be used to wake up the SPAID 102, deliver the sensor patient interface 110, deliver the delivery patient interface 186, prime the SPAID 102, such as removing air from the fluid delivery path, program the user's custom basal insulin delivery profile, check the status of the SPAID 102, initiate a bolus of insulin, such as a correction bolus or a meal bolus, make changes to the patient's insulin delivery profile, handle system warnings and alarms, and enable the user to switch between automatic and manual modes. The user apps 160 can also be configured to determine and provide dosing information to the SPAID 102, or provide information to the SPAID 102 that enables the MDA 129 to determine the proper dosing of the liquid medication.

[0098] The user app 160 may be configured to operate in a manual mode where the user app 160 may control the delivery of insulin at a programmed basal rate and a programmed or manual bolus amount, with the option to set a temporary basal profile. The controller 121 of the SPAID 102 may also have the capability to function as a sensor-augmented pump, using sensor glucose data provided by the analyte sensor 108 for input into a bolus calculator when the user app 160 is operating in manual mode.

[0099] The user app 160 may be configured to operate in an automated mode where the user app 160 assists in the use of multiple target blood glucose values. For example, in one embodiment, the target blood glucose value may be in the range of 110-150 mg / dL, with increments of 10 mg / dL, 5 mg / dL, or other increments, but preferably in increments of 10 mg / dL. The user experience mirrors the current setup flow, where a healthcare provider assists the user in programming the basal blood glucose value, glucose target value, and bolus calculator settings. These are used by the user app 160 as dosing parameters. The dosing parameters may be adapted over time based on the total daily insulin (TDI) delivered by the SPAID 102. A temporary hypoglycemic protection mode may be implemented by the user in the automated mode for various durations. In the hypoglycemic protection mode, the user app 160 algorithm reduces insulin delivery and is intended for use over temporary periods when insulin sensitivity is expected to be high, such as during exercise.

[0100] The user app 160 (or MDA 129) can control the SPAID 102 to provide periodic insulin microboluses based on past glucose measurements and / or predicted glucose over a prediction interval (e.g., 60 minutes). Optimal post-prandial control may require the user to deliver meal boluses in the same manner as current pump therapy, but normal operation of the user app 160 or MDA 129 compensates for missed meal boluses and reduces persistent hyperglycemia. The user app 160 or MDA 129 uses a goal-directed control strategy that seeks to achieve and maintain set target glucose values, thereby reducing the duration of persistent hyperglycemia and hypoglycemia.

[0101] The user app 160 may provide the ability to calculate a recommended bolus dose through the use of a bolus calculator. The bolus calculator is provided for the user's convenience to aid in determining a suggested bolus dose based on IOB estimated by the user app 160 or MDA 129, taking into account ingested carbohydrates, most recent blood glucose reading (or blood glucose reading if using a fingerstick), programmable correction factors, insulin to carbohydrate ratio, target glucose value, and any manual boluses and insulin delivered by an algorithm.

[0102] Description of Specific Embodiments In various embodiments of the SPAID 102, different types of patient interfaces may be used to deliver the liquid medication and perform the necessary sensing to determine the correct dosing information for the liquid medication. In some embodiments, the delivery mechanism and the sensing mechanism may be combined into a single patient interface, while in other embodiments, the delivery mechanism and the sensing mechanism may use separate patient interfaces. Additionally, different mechanisms may be used to deliver the liquid medication and different methods may be used to perform the necessary sensing.

[0103] In one embodiment, shown generally in FIG. 3A, glucose sensing and insulin delivery are combined in a single cannula 302. Cannula 302 includes a glucose sensing region 304, and liquid medication is delivered through a lumen defined within the cannula and released through an open tip 306 of cannula 302. Glucose sensing region 304 can be, for example, any of the embodiments shown in FIGS. 14A-14C and 15A-15C. FIG. 3B is a conceptual diagram of this embodiment, with an inset showing a close-up view of cannula 302. The distal end of cannula 302 can be tapered to a smaller diameter at its most distal end and can include cross-drilled holes through which liquid medication can be delivered in addition to delivery through open tip 306.

[0104] 4A shows in schematic form one variation of a single cannula configuration, where cannula 402 has a closed end 408 but includes a series of perforations 404 for delivery of liquid medication. In this variation, cannula 402 also includes a glucose sensing region 406. In some embodiments where delivery and sensing are provided on the same cannula, it may be desirable to deliver insulin to the epidermal or dermal layers while the sensing portion of the cannula is positioned subcutaneously.

[0105] Glucose sensing region 406 can be, for example, any of the embodiments shown in Figures 14A-15C and 15A-15C. Figure 4B shows a conceptual image of this embodiment with a close-up of cannula 402 in the inset. Note that cannula 402 shown in Figures 4A, 4B may be at any angle relative to the bottom of SPAID 102. In some embodiments that include both a delivery cannula and a sensing cannula or needle, it may be desirable to maximize the distance between the delivery cannula and the sensing needle, and as such, these needles may be provided at an angle oriented to achieve this goal.

[0106] The variation shown in Figures 3A-4B, which uses a single cannula, has the advantage of reducing the number of times it can be inserted into the patient's body.

[0107] In an alternative embodiment, shown generally in Figure 5A, multiple patient interfaces are inserted into the user. Cannula 502 is configured with an open end 504 for delivering liquid medication, while cannula 506 is configured with a closed end 510 and glucose sensing region 508 for sensing the user's glucose level. The distal end of cannula 502 may be tapered to a smaller diameter at the most distal end and may include cross-drilled holes through which liquid medication may be delivered in addition to delivery through the open tip 504. Figure 5B is a conceptual diagram of this embodiment, showing a close-up of cannula 506 in the inset.

[0108] In the embodiment shown in Figures 3A-5B, a cannula may be housed in the SPAID102 and inserted into the user's body during the installation process of the SPAID102, or at any time throughout the usage cycle, to allow the SPAID102 to remain in the same location on the user for an extended period of time as needed to address possible sensor fouling and site blockage or simply sensor degradation.

[0109] To extend the usable life of the SPAID 102 once attached to the user's body, it may be possible to provide multiple or additional cannulas or sensing element delivery mechanisms to "fire" or self-attach if the originally inserted cannula becomes damaged or inoperable. The insertion mechanism may be identical and may be used to insert additional delivery patient interfaces or sensor patient interfaces. The insertion / delivery mechanisms may be configured in modules within the SPAID 102 that are replaceable.

[0110] Additionally, an optional additional placement mechanism may place an additional sensing element that is either standalone or built directly onto the liquid drug delivery cannula. The additional sensor is placed under the skin at a new site beneath the base of the SPAID 102 and begins to warm up once placed. After a warm-up period, the existing sensing device can be used to calibrate the newly placed device. Once the new sensor and cannula are operational and correctly calibrated, the original electrode and cannula are retracted and unavailable.

[0111] For cannulas that deliver liquid medication, a priming process to ensure there is no air in the cannula upon placement can occur prior to placement of the combined delivery and sensing cannula delivery cannula or prior to adhering the SPAID102 to the user's skin. A pump for the cannula is activated inside the SPAID102 and is activated for a period of time to push a certain amount of liquid medication through the cannula and remove any air therein. The amount pushed should be at least equal to the volume of the cannula tubing from the reservoir to the tip. Prior to placement on the skin, a collection area or absorbent pad can be placed inside the housing of the SPAID102 to collect and absorb any excess liquid medication that may leak out of the cannula tip during this process.

[0112] When additional cannulas and sensors are installed to support additional sensors and cannula placement mechanisms, a retraction mechanism may be used to remove the first installed cannula to prevent inadvertent administration of liquid medication. In one embodiment, a dual spring system may be provided. One spring is used to initially place the device on the skin along the guide track with the sensor and cannula attached to the track using the shuttle carrier device. When the first spring is fired, it is disconnected from the drive path. The second spring applies a reverse force to the shuttle carrier device to retract the sensor and cannula from the body without the need to recompress the initial spring.

[0113] In a variation of the described embodiment, the same deployment system may be capable of being retracted and reinserted into the body. During the retraction process, a wiping action is performed along a drying pad or other chemical-based cleaning surface that can remove bodily fluids and other sensor fouling components, cleaning and restoring the performance of the sensor. Additionally, the cannula may be wiped or cleaned as well to remove any obstructing components. During the retraction of the cannula, the cannula may push additional insulin onto the collection pad to push any obstructing material through the cannula. One possible implementation is to use a dual spring system with a mechanism for enabling and disabling the spring engagement between the deployment spring and the retraction spring to prevent one from exerting a force on the other. Alternatively, a motor-driven screw drive mechanism may be employed to slowly retract the sensor along the guide track. This system also serves to unwind / reload the initial deployment spring so that it can be reused to launch / reload the sensor and cannula.

[0114] Electrochemical Sensing Electrochemical sensing can be used to sense glucose and other analytes in the various drug delivery systems described herein. Furthermore, sensing can be performed with single or multiple electrochemical cells to enhance reliability, accuracy, durability, and / or overall performance. Furthermore, the electrochemical cells can be configured for use in any suitable electrochemical glucose sensing method, including those described in more detail below.

[0115] Various electrode configurations can be employed. In some embodiments, referring to FIGS. 6A-6B, a two-electrode configuration can be employed in which the electrochemical cell includes one working electrode 622 and one counter / reference electrode 624. The electrodes can be operated with either a positive or negative excitation voltage 626 to cause either a reduction or oxidation reaction at the working electrode 622. In a positive excitation mode, a positive voltage is applied to the input terminal 632t of the working amplifier 632 relative to the input terminal 634t of the counter / reference amplifier 634 to generate a positive excitation voltage and drive the reduction reaction. In a negative excitation mode, a negative voltage is applied to the input terminal 632t of the working amplifier 632 relative to the input terminal 634t of the counter / reference amplifier 634 to generate a negative excitation voltage and drive the oxidation reaction. The excitation voltage 626 can be, for example, fixed, variable, or pulsed. Exemplary excitation voltages can range, for example, from +600 mV to -600 mV, with other possible values ​​included. The current generated by the excitation voltage can be measured at either the working electrode terminal or the counter / reference terminal.

[0116] In some embodiments, and with reference to FIGS. 7A-7B, a three-electrode configuration may be employed in which the electrochemical cell is comprised of one working electrode 722, one counter electrode 724, and one reference electrode 728. As with the previous two-electrode configuration, the electrochemical cell may be operated with either a positive or negative excitation voltage 726 to generate a reduction or oxidation reaction at the working electrode 722, respectively. In a positive excitation mode, a positive voltage is applied to the input terminal 732t of the working amplifier 732 relative to the input terminal 734t of the counter / reference amplifier 734 to generate a positive excitation voltage and drive the reduction reaction. In a negative excitation mode, a negative voltage is applied to the input terminal 732t of the working amplifier 732 relative to the input terminal 734t of the counter / reference amplifier 734 to generate a negative excitation voltage and drive the oxidation reaction. The excitation voltage 726 may be, for example, fixed, variable, or pulsed. An excitation voltage 726 may be established between the working electrode 722 and the reference electrode 728 with the counter electrode 724 used to servo positively and negatively to ensure that the excitation voltage 726 is maintained between the working electrode 722 and the reference electrode 728. Exemplary excitation voltages may range, for example, from +600 mV to -600 mV, with other possible values ​​included. The current resulting from the excitation voltage may be measured at either the working electrode terminal or the counter terminal. The reference electrode terminal is only used to "sense" and servo the voltage in the reaction to maintain the appropriate excitation voltage, and does not draw any significant current.

[0117] In some embodiments, and with reference to FIG. 8, a four-electrode configuration may be employed in which the electrochemical cell is comprised of a main working electrode 822a, an additional working electrode 822b, one reference electrode 828, and one counter electrode 824. The electrodes may be operated as described for the three-electrode configuration above with the addition of a fourth electrode configured as the additional working electrode 822b. The additional working electrode 822b may be excited with an excitation voltage 826b that is higher than, lower than, or equal to the excitation voltage 826a associated with the main working electrode 822a. The additional working electrode 822b may have additional chemical moieties added to it, and may serve as a possible "blanking" electrode to measure background interference present in the main working electrode signal. The additional chemical moiety may be any moiety that negates the resulting signal resulting from reaction with the analyte of interest. When glucose is used as the analyte of interest, an enzyme, such as glucose oxidase (GOX) or glucose dehydrogenase (GDH), is used to react with glucose to form hydrogen peroxide, which is then measured at the working electrode. The blanking electrode is sometimes coated with catalase, a natural consumer of hydrogen peroxide, to prevent it from receiving the hydrogen peroxide that is produced when glucose interacts with GOX. With the hydrogen peroxide removed, the blanking electrode is free to respond to and measure only background interference. This "blanking" signal is subtracted by the electronics system, either analog or digitally, or by software, from the main working electrode signal, leaving only the pure glucose / GOX response signal. An additional working electrode current flows through the counter terminal, and its voltage can be established relative to the reference electrode terminal.

[0118] An additional use of the additional working electrode 822b is for "cleaning" the first working electrode when it becomes fouled or coated with unwanted chemical species. Examples of such chemical species include phenols, amino acids, neurotransmitters, proteins, and other biomolecules (including whole or fragmented cells, DNA molecules, and / or RNA molecules), which over time form a layer on the working electrode and reduce its sensitivity and performance. The additional working electrode 822b is initially dormant and not electrically connected, but is then activated as needed (e.g., after contamination or after sufficient time has passed for contamination to occur) and operates at a voltage that can draw unwanted chemical species from the main working electrode 822a to itself. This activation may occur repeatedly at programmed intervals and / or upon detection of reduced sensitivity of the working electrode. In this embodiment, the additional working electrode 822b is close enough to the main working electrode 822a to ensure that cleaning is completed quickly (less than 10 minutes). Generally, the working electrode 822b and the primary working electrode 822a should be as close to each other as the manufacturing process will allow, ideally less than 1 mm. Once the unwanted species have been removed, the system is returned to normal operation. This process can be repeated as necessary.

[0119] In further embodiments, the additional working electrode 822b may be used as a backup electrode for the main working electrode 822a, for example, if the main working electrode 822a becomes contaminated. For example, the additional working electrode 822b may be initially dormant and then activated and used as the main analyte measuring electrode, while the contaminated working electrode 822a is disabled. This may be repeated "N" times for as many additional working electrodes as are added to the system. In such a configuration of the system, the working electrodes should ideally be sufficiently separated so as to be electrically isolated from the other electrodes.

[0120] Any number of possible electrochemical cell configurations can also be established, including those that include multiple electrochemical cells. Such multiple cell configurations can consist of two-electrode, three-electrode, or four-electrode cells. Additionally, multiple cell configurations can be established in which the electrochemical cells use the same electrical ground. Alternatively, multiple electrochemical cells can be provided in which each of the electrochemical cells uses an isolated or floating ground system to prevent "crosstalk" between two adjacent electrochemical cells. This can be particularly useful, for example, when creating blanking cells to prevent a working electrode terminal from attempting to pass current to an adjacent reference or counter electrode terminal.

[0121] In the embodiment shown in FIG. 9, a first cell 920a consisting of a first working electrode 922a, a first counter electrode 924a, and a first reference electrode 928a can be constructed with appropriate chemistry to measure the desired analyte with interference. A second cell 920b consisting of a second working electrode 922b, a second counter electrode 924b, and a second reference electrode 928b can be constructed with appropriate chemistry to measure only the interference. As mentioned above, in the case of a glucose measurement system, an example is to use GOX to react with glucose to form hydrogen peroxide, which is then measured at the working electrode. However, due to unwanted background interference, the first working electrode 922a measures both the glucose signal and interference signals, such as those produced by ascorbic acid and acetaminophen. The second working electrode 922b, which serves as a blanking electrode, can be coated with catalase, a natural consumer of hydrogen peroxide, thus preventing the second working electrode 922b from receiving the hydrogen peroxide that is produced when glucose interacts or reacts with GOX. With the peroxide removed, the second working electrode 922b is free to respond to and measure only the background interference. The electrical signal from the second working electrode 922b is then subtracted from the electrical signal from the first working electrode 922a via analog differential circuitry or using digital electronics and software to obtain the desired analyte signal.

[0122] A feature that may be provided with essentially any electrochemical sensor design is the addition of a spare set of electrochemical electrodes, or in some embodiments, only a spare working electrode that remains dormant until needed. The spare electrode remains disconnected through internal electronics (e.g., through high impedance) to prevent attraction of contaminants. When the currently active electrode shows signs of poor performance or sensitivity, or after a predefined period of time, the spare electrode (e.g., a spare working electrode) can be activated. The spare electrode works in tandem with the originally activated electrode, and once it has "stabilized" and confirmed to be accurate, it becomes the active electrode and the original electrode is deactivated. This process continues until all spare electrodes are consumed.

[0123] Electrode composition / material The working electrodes for the various chemical cells described herein may be comprised of a reactive surface made from one or more of the following materials: platinum, palladium, gold, rhodium, Prussian blue, Prussian blue analogs, carbon, or carbon nanotubes, among others. Additionally, additional lower cost metals or materials, such as carbon, silver, copper, combinations thereof, or other acceptable electrical conductors, may be employed to conduct electrical signals from the electrode surface material to the system electronics.

[0124] The reference electrode used in the various three terminal and three terminal cells described herein, in conjunction with the material selected for the working electrode, determines the oxidation or reduction voltage required to drive the cell to measure the desired analyte. The reference electrode is generally constructed of a mixture of silver and silver chloride, typically used in a 60 / 40 ratio of each due to its electrochemical stability. Other ratios of silver and silver chloride and other materials can be employed. Other materials include platinum, palladium, gold, rhodium, Prussian blue, Prussian blue analogs, carbon or carbon nanotubes, graphene or graphene doped metal alloys, etc. Additionally, less expensive metals or materials, such as carbon, silver, copper, combinations thereof, or other acceptable electrical conductors, can be employed to convey the electrical signal from the surface material of the reference electrode to the system electronics.

[0125] The counter electrode of the various three terminal cells described herein serves to compensate for impedances present in the cell. These impedances may be due to salts or other compounds that result in a voltage drop in the cell as the analyte current being measured increases. The counter electrode voltage is automatically changed so that the excitation voltage in the cell between the reference and working electrodes is maintained. Alternatively, in the various two terminal cells described herein, the reference and counter terminals are generally combined into one counter / reference terminal. The counter or counter / reference electrode is generally composed of a mixture of silver and silver chloride, typically used in a 60 / 40 ratio due to its electrochemical stability. Other ratios of silver to silver chloride and other materials may be employed. Other surface materials for the counter or counter / reference electrode include platinum, palladium, gold, rhodium, Prussian blue, Prussian blue analogs, carbon or carbon nanotubes, graphene, or graphene doped with metal alloys. Additionally, additional less expensive metals or materials such as carbon, silver, copper, combinations thereof, or other acceptable electrical conductors may be employed to conduct the electrical signal from the counter or counter / reference electrode surface material to the system electronics.

[0126] Electrochemical Process Configuration An electrochemical cell for measuring glucose or other analytes can be constructed with electrodes in any of the configurations described above. Additionally, various chemical elements can be added to the cell to facilitate collection and measurement of the desired analyte.

[0127] In various embodiments, and with reference to FIG. 10, glucose can be measured indirectly through the hydrogen peroxide produced. In one possible construction method, glucose oxidase (GOX) enzyme 1015 is cast or dried onto the surface of the working electrode 1022, for example within a matrix of hydrogel or polyethylene glycol surface groups. Upon activation by rehydration, the enzyme 1015 produces gluconate 1014 and hydrogen peroxide 1018 in the presence of endogenous glucose 1012, endogenous coenzymes (which may be present in reduced coenzyme 1016 or oxidized coenzyme 1017), and oxygen 1019 from the air or the body. For measurement, the hydrogen peroxide is then reduced or oxidized depending on the cell's electrodes and voltage configuration, and the reaction produces a certain number of electrons (i.e., 2) per glucose molecule converted. The electrons are then measured as an electrical signal, providing a quantitative value that directly corresponds to the glucose concentration. The described method is just one way of measuring glucose. Other methods may be used.

[0128] In various embodiments, and with reference to FIG. 11, glucose may be measured indirectly via a mediator. In one possible construction method, glucose oxidase enzyme 1115 is bound to a mediator (which may have a reduced mediator 1118 state or an oxidized mediator 1119 state) and cast or dried onto the surface of a working electrode 1122, for example within a matrix of hydrogel or polyethylene glycol surface groups. Upon activation by rehydration in the body, the enzyme with bound mediator, in the presence of glucose 1112 and a coenzyme (which may have a reduced coenzyme 1116 state or an oxidized coenzyme 1117 state), converts glucose 1112 to gluconate 1114 and electrons for measurement, as measured by an electrode, for example electrode 1122. The glucose oxidase with mediator is reduced or oxidized depending on the electrode and voltage configuration, and the reaction produces a number of electrons per glucose molecule converted. The electrons are then measured as an electrical signal that produces a quantitative value that directly corresponds to the glucose concentration. This method of measuring glucose does not require oxygen, which can be limiting in the body.Commonly available mediators suitable for this reaction include potassium ferricyanide, hexammineruthenium(III) chloride, and methoxyphenazine methosulfate.

[0129] In various embodiments, and with reference to FIG. 12, glucose may be measured via a direct electron transfer (DET) type glucose-reactive enzyme. In one possible construction method, a DET glucose-reactive enzyme 1215 is cast or dried onto the surface of a working electrode 1222, for example, within a matrix of hydrogel or polyethylene glycol surface groups. The enzyme 1215 is activated upon rehydration in vivo and, in the presence of glucose 1212 and a coenzyme (having a reduced coenzyme 1216 state and an oxidized coenzyme 1217 state), converts glucose 1212 to gluconate 1214 and electrons for measurement. The enzyme is reduced or oxidized depending on the electrode and voltage configuration, and the reaction produces a certain number of electrons per glucose molecule converted. The electrons are then measured as an electrical signal that produces a quantitative value that directly corresponds to the glucose concentration. This method of glucose measurement has recently become available, and several DET enzymes are currently available.

[0130] In various embodiments, and with reference to FIG. 13, glucose may be measured by direct electron transfer. One possible construction method is to use a working electrode 1322 composed of carbon nanotubes, graphene or graphene doped with metal alloys, and / or alloy nanostructures including platinum, lead, gold, palladium, and / or rhodium. In the presence of glucose 1312, the working electrode material converts glucose 1312 directly to gluconate 1314 and electrons for measurement when an appropriate excitation voltage is applied. This reaction is ideal because it does not use enzymes or mediators to measure glucose concentration. The electrons are then measured as an electrical signal that generates a quantitative value that directly corresponds to the glucose concentration. This method of glucose measurement is emerging with the development of newer nanotechnology and manufacturing systems and has proven to be a valid alternative to traditional glucose sensing methods.

[0131] Next, various configurations of electrochemical cells used as analyte sensors are described. Electrochemical cells for measuring glucose or other analytes can be constructed with electrodes that can be printed, plated, precast, embedded, or fabricated using any other method necessary to construct the electrodes with the correct materials and electrical conduction paths required to send a signal back to the internal electronics for processing, on a plastic or metal needle or cannula. In general, an electrochemical cell requires a working electrode, a counter electrode, and a reference electrode. In some embodiments, the counter electrode and the reference electrode can be combined into one electrode.

[0132] 14A-14C and 15A-15C show various exemplary shapes, patterns, and orientations of two-electrode and three-electrode embodiments of electrochemical cells. It should be noted that these illustrations are exemplary in nature and are not intended to limit the scope of the invention. As will be appreciated by one of ordinary skill in the art, other configurations and arrangements of components and methods of manufacture are possible and are contemplated to be within the scope of the invention.

[0133] For ease of illustration, the electrochemical cell diagrams in Figures 14A-14C and Figures 15A-15C are shown as if a round cannula or needle had been cut open along its length and pressed flat with the outside of the cannula or needle depicted. Figure 14A shows a two-electrode sensor utilizing a working electrode 1402 and a combined counter / reference electrode 1404. Figure 14B shows a three-electrode sensor having a working electrode 1402 with separate counter and reference electrodes shown and designated as reference numbers 1406 and 1408, respectively. Figure 14C also shows a three-electrode sensor having three working electrodes 1402 and separate counter and reference electrodes 406 and 1408, respectively.

[0134] 15A-15C show various embodiments of sensors with electrodes configured as bands wrapped around a needle or cannula. FIG. 15A shows a two-electrode sensor with a working electrode 1502 and a combined counter / reference electrode 1504. Note that the wire 1503 extending from the counter / reference electrode 1504 is insulated. It is located underneath the working electrode 1502. FIG. 15B shows a configuration with a dual two-electrode sensor with working electrode 1502 and counter / reference electrode 1504, respectively. This embodiment can be used, for example, to sense two separate analytes. FIG. 15C shows a three-electrode sensor with working electrode 1502, reference electrode 1508, and counter electrode 1506 configured as a wire mesh on the surface of the cannula.

[0135] Electrochemical cells require reactive chemical components to be placed on the electrode surface for the sensor to function properly. These chemical components are used to diffuse the analyte from the body through the body fluid or a liquid pre-supplied during the manufacturing process so that the analyte reacts completely with the embedded chemical species and / or directly with the electrodes in the chemical cell. For example, a hydrogel, which may contain sterilized components and water from the manufacturing process, can be pre-packed with the sensor to reduce the start-up time of the sensor, increase the diffusion of the analyte into the electrochemical cell, and generate chemicals from the sensor. Alternatively, another example is a sensor with a hydrogel applied to the electrodes to transport glucose molecules directly to the carbon nanotube, graphene, or graphene doped with metal alloy electrodes, where the glucose molecules are directly converted into a quantifiable electrical signal.

[0136] The use of enzymes to convert glucose to a second reactive component or directly to electrons requires manufacturing techniques to deposit and dry the enzyme on the cell's electrodes. In one embodiment, this can be accomplished by mixing the enzyme directly with the electrode ink and drying the combination in such a way that the ink hardens and the enzyme remains effective and active. Other methods of depositing and drying the enzyme on the cell's electrodes can also be used. Alternatively, various gels and polymers can be used to cast and dry the enzyme along with other chemicals such as surfactants and stabilizing compounds to stabilize the enzyme and allow it to receive the analyte and conduct the appropriate chemical component to the electrode to reduce or oxidize, and measure the resulting electron transfer.

[0137] It is not uncommon for the working electrode of an electrochemical cell to become contaminated or damaged by unwanted chemical species that collect on the electrode surface during normal operation. Methods for protecting the electrode include coating it with Nafion®, polysulfone, cellulose, or other blocking membrane materials to keep contaminating molecules away from the sensitive working electrode. Specifically, the coating should have pore size selectivity only for hydrogen peroxide molecules, while excluding larger potential contaminating molecules such as phenolic resins commonly found in insulin.

[0138] In some embodiments, the electrodes are sterilized prior to use. Thus, the materials used to form and / or cover the electrodes may be selected based on their resistance to the sterilization process. Additionally, excess enzyme or sacrificial coatings may be employed to protect the sensor during the sterilization process. For example, an excess of enzyme may be coated onto the electrodes, and the enzyme type and coating thickness may be varied based on the expected amount of enzyme that will be removed from the electrodes during sterilization. The sacrificial coating may include a protective coating, surface finish, or sacrificial protective laminate that is configured based on the expected amount of sacrificial coating that will be removed during the sterilization process. In an exemplary embodiment, the sterilization process may be associated with parameters (temperature, pressure, etc.) that describe how sterilization will occur. The materials forming or covering the electrodes, the enzyme material and thickness, and the properties of the sacrificial coating may be selected based on the sterilization process parameters.

[0139] An added benefit of the membrane is that it acts as a structure for immobilizing the chemical components needed for the electrochemical cell onto the electrodes: enzymes and gels can be cured onto the membrane structure as part of the cell construction process.

[0140] Additionally, the entire sensor can be covered with an additional membrane to limit the influx of glucose. It is important to carefully select the size of the pores and the amount of pores to allow glucose molecules to pass while limiting their amount. The size of the pores must be larger than the size of the glucose molecules (e.g., larger than about 1 nm). The size of the pores can be determined through research and experimentation to be large enough to allow glucose molecules to pass through, but not so large that unwanted molecules also pass through. The amount of pores helps to control the rate of conversion of glucose to hydrogen peroxide, since the reaction requires oxygen, and oxygen is not as abundant as glucose when the sensor is operating in the body. Therefore, a lack of oxygen limits the conversion of glucose to hydrogen peroxide, which directly affects the glucose value calculated by the electronics and software system.

[0141] In one embodiment described above with respect to FIG. 15B, a single sensing device or needle may include multiple electrochemical sensing regions. Each additional sensing region requires its own electronic analog front end to provide and maintain the appropriate excitation voltage. The purpose of the additional sensing elements is to tailor the additional sensing regions to specific sensing characteristics by having separate sensor electrode sizes, materials, coatings, enzymes, chemical enhancers, and membranes. There are many possible combinations and configurations, and one sensor may be used to provide a complete sensing solution that may have a limited range but can respond quickly to rising and falling glucose levels. This is accomplished by miniaturizing the sensor and using thinner hydrogels and filter components to allow glucose molecules to penetrate and diffuse quickly into the sensing region. Such sensors may be easily affected by high glucose levels and contaminants, so placing them away from the insulin source, and possibly high on the skin, naturally protects the sensor electrodes from contamination. Using larger electrodes and thicker gels, coatings, and membranes, additional sensing regions may be employed with slower but averaged responses and a larger overall measurement range.

[0142] FIG. 16 is a block diagram of a possible implementation of electronics and control systems for driving the sensors and collecting quantifiable electrical signal data from various electrochemical cell implementations. In one embodiment, the electronics and control system may be the SPAID 102 electronics, or the control system may be implemented separately. Interpretation of the data collected from the sensors is done by software algorithms, which must be tailored to the individual sensor, taking into account the response characteristics of each sensor and the resulting digitized analog signal. In some embodiments, the software algorithms may be implemented as part of the MDA 129 or user app 160, or may be implemented as separate software stored in memory 123 and executed by controller 121.

[0143] 1 and 16, the sensor patient interface 110 may include one or more electrochemical cells 1602-1...1602-N. In some embodiments, each of the electrochemical cells 1602-1...1602-N may be configured to detect glucose, while in other embodiments, some of the electrochemical cells 1602-1...1602-N may be configured to detect glucose and others may be configured to detect other analytes. The raw signals from the electrochemical cells 1602-1...1602-N in the sensor patient interface 110 are sent to the analyte sensor 108. Within the analyte sensor 108, the signals are amplified at 1604, multiplexed at 1606, and converted to a digital signal at an analog-to-digital converter 1608. The amplifier 1604 is preferably adapted to maintain an appropriate excitation voltage for each electrochemical cell throughout its use cycle. Once converted to a digital signal, a software algorithm, which may be implemented as part of the MDA 129 running on the controller 121, analyzes the raw signal for glucose measurements or other analyte level measurements. The MDA 129 can determine from the resulting glucose measurements whether an alarm or warning is necessary and, if necessary, send user feedback via the user interface 127. Before being used by the MDA 129 and the electrochemical cells 1602-1...1602-N, the signal is converted from a digital to an analog signal by a digital-to-analog converter 1610 and amplified by amplifiers 1604. It should be noted that the electrochemical cells 1602-1...1602-N receive power from amplifiers 1604-1...1604-N to keep the electrochemical cells active.

[0144] Optical fluorescence sensing Photofluorescence sensing may be used to sense glucose in the various drug delivery systems described herein, including the various needle or cannula-based sensing systems described herein. Sensing may be performed by single or multiple photofluorescence sensing regions on or within a given needle or cannula to improve reliability, accuracy, durability, and overall performance. Alternatively, photofluorescence sensing elements may be placed separately at the time of device installation and removed at the end of the device's life cycle, if desired.

[0145] One common method of optical fluorescence sensing is based on Förster (or fluorescence) resonance energy transfer, commonly known as FRET. Referring to Figs. 17-20, the glucose-to-light direct conversion method is shown, in which a receptor fluorophore (generally a fluorescent dye) is bound to an acceptor molecule. The acceptor (also known as a quencher) can bind to both glucose and the receptor fluorophore. Referring to Fig. 17, the acceptor molecule 1702, when it is not also reversibly bound to glucose, abstracts an electron 1704e from the receptor fluorophore 1704 (a glucose receptor without glucose is shown as 1702r). This abstraction of electrons is called FRET quenching. FRET can be considered as a process in which the acceptor molecule 1702 (quencher) suppresses the fluorescence from the receptor fluorophore 1704 (e.g., a fluorescent dye) by removing energy therefrom. This limits the ability of the receptor to emit energy as a photon 1704p after receiving an input excitation photon. The energy extracted from the receptor fluorophore 1704 is in the form of an electron 1704e. This occurs when the physical distance between the acceptor 1702 and the receptor 1704 is small, allowing close-proximity transfer of electrons. In the absence of glucose, the quenching process is strong and the acceptor 1702 is able to extract electronic energy from the emitting fluorophore 1704, resulting in a decrease in light output 1704p.

[0146] FIG. 18 shows the energy excitation 1800p, energy loss 1800l, and energy release 1800e states of a fluorophore when there is FRET quenching. The absorbed excitation light from an external light source is received by the fluorophore, which raises it from the ground energy state S0 to a higher energy state S3. Once raised, the fluorophore's energy quickly dissipates in the form of thermal or vibrational energy 1800l, and the energy level drops to the lowest level S1 of the excited energy state. This energy loss occurs in just a few nanoseconds and is called the excitation energy lifetime. Finally, when FRET quenching is activated, the fluorophore releases energy to the quencher as electrons, rather than as photons. Once most of the energy is transferred to the quencher in the form of electrons, the fluorophore returns to the ground energy state S0 and starts the cycle again if excited.

[0147] 19, upon binding of glucose 1906 to the fluorescent acceptor (quencher) 1902, a conformational change occurs that moves the fluorescent donor (fluorophore) 1904 and the fluorescent acceptor (quencher) 1902 further apart. This reduces electron transfer between the fluorophore 1904 and the acceptor 1902, reducing FRET quenching, allowing the fluorophore 1904 to emit more energy as light rather than being quenched by electronic energy transfer to the acceptor 1902. Thus, as the concentration of glucose 1906 increases, more light is emitted from the fluorophore 1904 when it receives excitation light energy from an external source. Glucose binding to the receptor is a function of the glucose concentration in the solution in which this acceptor / receptor combination operates. Glucose binding is reversible and decreases as the analyte glucose concentration decreases.

[0148] FIG. 20 shows the energy excitation 2000p, energy loss 2000l, and energy release 2000e states of a fluorophore in the absence of FRET quenching. Absorbed excitation light from an external light source is received by the fluorophore, raising it from state S0 to a higher energy state S3. Once raised, the energy is quickly dissipated in the form of heat or vibrational energy, and the energy level drops to a lower level excited state. As mentioned above, this energy loss occurs in just a few nanoseconds and is called the excitation energy lifetime. Finally, in the absence of active FRET quenching, the fluorophore releases energy as photon emission at a lower energy state S1 than the initial input excitation photon energy state S3 due to losses during the energy lifetime transition. This release of energy shifts the color of the emitted light from the originally absorbed light to a longer wavelength / lower energy light. This results in a color shift from the violet end of the light spectrum towards the lower energy red end.

[0149] Photofluorescence sensing can be used to sense glucose in any needle or cannula-based sensing system, and furthermore, sensing can be performed by single or multiple photofluorescence sensing regions on a given needle or cannula to increase reliability, accuracy, durability, and overall performance.

[0150] 21 and 22 show the glucose-photoindirect conversion method, in which the acceptor can bind to both oxygen and the receptor fluorophore. Referring to FIG. 22, the acceptor 2202 can receive electrons (e -) from the fluorophore 2106. This electron abstraction is known as FRET quenching, as explained above. However, in this glucose-light indirect conversion method, the glucose concentration in the analyte solution can be determined based on the amount of oxygen 2108 available in the solution. The concentration of oxygen 2208 decreases as glucose 2206 reacts with glucose oxidase 2210 to produce hydrogen peroxide. This decrease in oxygen with increasing glucose concentration allows for an increase in FRET quenching, resulting in proportionally less light output from the fluorophore, as shown in Figure 22. The opposite scenario occurs when lower glucose concentration results in more oxygen 2108, which binds to the acceptor 2102 and moves the receptor fluorophore 2104 and acceptor 2102 farther apart (see Figure 2). As a result, electron transfer between the receptor fluorophore 2104 and acceptor 2102 is reduced and FRET quenching is reduced, allowing the receptor fluorophore 2104 to emit more energy as light rather than being quenched by electronic energy transfer to the acceptor 2102. Thus, as glucose concentration decreases, more light will be emitted when the receptor fluorophore 2104 receives excitation light energy from an external source. Oxygen binding is reversible and decreases as the analyte glucose concentration increases.

[0151] In all embodiments using optical fluorescence sensing, the sensing material, when excited by a light source, produces a quantifiable light signal of magnitude and duration proportional to the glucose concentration, and this process works for both increasing and decreasing concentrations. Additionally, the length of time that the fluorophore emits light after being excited by a light source is also proportional to the glucose concentration. Both quantifiable measurement methods can be employed to calculate the glucose concentration by the MDA 129 or user app 160.

[0152] 23 shows one possible glucose sensor configuration for the glucose-to-light direct conversion method of photofluorescence sensing using a deployable photofluorescence sensing element 2302. In one embodiment, the sensing element 2302 is deployed through a cannula 2304 or pushed into the skin using a retractable needle or alternative mechanism. The sensing element 2302 is designed to fluoresce in the presence of glucose and should be implanted in the subcutaneous (sub-Q) region of the skin just below the base of the SPAID 102.

[0153] When the sensing element 2302 is placed using a cannula 2304, the cannula placement mechanism presses the element into the skin upon initial insertion, but then the cannula 2304 can be retracted slightly to leave the element in the subcutaneous region of the skin slightly beyond the tip of the cannula 2304. The sensing element 2302 remains mechanically connected to the cannula 2304 via a tether, cord, or cable 2306. At the end of the system's life cycle, the cannula 2304 and sensing element 2302 are pulled back together and the tether 2306 is used to retrieve the sensing element 2302.

[0154] Once implanted, the photofluorescence sensing element 2302 is used to generate a measurable light signal proportional to glucose. An excitation light source 2308 designed to generate the correct wavelength of light required to excite the fluorophores on the sensing element is integrated into the base of the housing of the SPAID 102. Preferably, the deployment mechanism of the SPAID 102 positions the sensing element 2302 directly beneath the excitation light source 2308 at a depth of 2-5 mm such that light emitted from the excitation light source 2308 follows path 2312 to reach the sensing element 2302. The SPAID 102 also incorporates a photodetector 2310 that is mounted at the base of the housing of the SPAID 2310 directly above the sensing element 2302. In some embodiments, the excitation light source 2308 and the photodetector 2301 may be integrated into a single unit. The photodetector 2310 may have selective filtering so that it can detect emitted light having wavelengths produced by the fluorophores in the sensing element. The emission light follows path 2314 to photodetector 2310. Due to the FRET effect, the excitation light and emission light are at different frequencies, which helps to reduce "crosstalk" between the excitation light and the emission light.

[0155] 24 shows an alternative glucose sensor configuration using the glucose-to-light direct conversion method of photofluorescence sensing. In this embodiment, a photofluorescence sensor coating 2404 is disposed on the surface of a cannula 2402 or needle. The cannula 2402 or needle with the photofluorescence sensing region 2404 is embedded with a mechanism that allows the sensing region 2404 to be positioned directly (or nearly directly) beneath an excitation light source 2408 and a detector 2410 that are mounted at the base of the housing of the SPAID 102.

[0156] Once implanted, the photofluorescent sensing region 2404 on the cannula 2402 or needle generates a measurable light signal proportional to glucose when excited with a particular wavelength of light. An excitation light source 2408 designed to generate the correct wavelength of light required to excite the fluorescent dye on the sensor is included in the SPAID 102, preferably integrated into the base or bottom of the SPAID 102. The deployment mechanism of the SPAID 102 is designed to position the sensing region 2404 directly beneath the excitation light source 2408 at a depth of 2-5 mm. The SPAID 102 also incorporates a photodetector 2410 attached to the base or bottom of the SPAID 102 directly above the sensing region 2404. In some embodiments, the excitation light source 2408 and the photodetector 2410 may be integrated into a single unit. The photodetector 2410 has selective filtering that allows it to detect only the wavelengths of light generated by the fluorophores in the sensor 2404. Due to the FRET effect, the excitation light 2408 and the light emitted from the sensor 2404 are at different frequencies, reducing crosstalk between the source light and the detected light.

[0157] 25A shows another possible embodiment of a glucose sensor configuration using a glucose-to-light direct conversion method of photofluorescence sensing using a photofluorescence sensor coating 2504 attached to the surface of a cannula 2502 or needle. The cannula 2502 or needle with the photofluorescence sensing region 2504 is implanted in a subcutaneous region of the skin below the base of the housing of the SPAID 102. In an alternative embodiment, the photofluorescence sensing region may be provided on a separate needle, as shown in FIG.

[0158] Once implanted, the photofluorescence sensing region 2504 on the cannula 2502 or needle generates a measurable light signal proportional to glucose. An excitation light source 2508 designed to generate the appropriate wavelength of light required to excite the fluorophores on the sensor 2504 is integrated inside the housing of the SPAID 102. The cannula 2502 or needle can be constructed with optical conduit or fiber optic light guiding material as shown in FIG. 25B and can be used to guide the excitation light as well as deliver the light emitted by the sensor 2504 from the SPAID 102 to the photofluorescence sensing region 2504. Alternatively, the cannula 2502 or needle can be constructed of alternative materials with light guiding elements embedded in the wall of the device.

[0159] FIG. 25B shows light guiding elements ("S" and "D") embedded in the wall of the cannula 2502 or needle. The amount, ratio and location of light guiding elements may vary. Light sourcing conductors ("S") are interleaved with light detecting conductors ("D") to improve illumination for fluorophore excitation and detection from fluorophore emission.

[0160] The light source 2508 and detector 2510 are mounted inside the SPAID 102 or are part of the SPAID 102 and coupled to the light guide. The light source 2508 and detector 2510 can form a combined device that allows both devices to be in a single package and share the opening of the light guide. Alternatively, a partial mirror device (not shown) or other light directing device (not shown) can be used to couple the light to the opening of the light guide so that the light source 2508 and the light detector 2510 can be mounted separately from each other. In an alternative embodiment, multiple light guides, such as those shown as "S" and "D" in FIG. 25B, can be employed to conduct light to and from the photofluorescence region on the cannula or needle. In this case, a portion of the light guide is dedicated to providing light ("S") and another portion ("D") is used to guide the detected light back to the light detector 2510. This allows the light providing and detection components to be used separately and mounted separately within the housing of the SPAID 102. The inner lumen or cannula 2512 can be used to deliver a liquid medication.

[0161] FIG. 25C is a schematic diagram of this embodiment, showing a close-up view of cannula 2502 in the inset.

[0162] In this embodiment, it is necessary to redirect the excitation light in an orthogonal direction so that it can interact with the phosphor material in the sensing area 2504 attached to the surface of the cannula 2502 or needle. In this embodiment, the excitation light illuminates the phosphor material from the backside of the sensing area 2504 and emission light is further generated and travels in the reverse direction back to the SPAID 102. During the manufacturing process, the walls of the light conduit are "broken" in a controlled manner to allow light to escape the sidewall of the cannula 2502 and interact with the phosphor material in the sensing area 2504. Other methods of mirrors or surface scratches may also be used to allow light to exit and return from the surface of the cannula 2502.

[0163] The photodetector 2510 is configured with selective filtering so that it can detect only the wavelengths of light emitted by the fluorescent material in the sensing region 2504. Due to the FRET effect, the excitation and emission light are at different frequencies, thereby reducing crosstalk between the provided and detected light.

[0164] FIG. 26A shows yet another embodiment of a photofluorescence sensor mounted on the surface of the cannula 2602 or needle, designed to fluoresce in the presence of oxygen. This sensor is used to generate a measurable light signal proportional to glucose by using glucose and glucose oxidase to reduce the local concentration of oxygen through a chemical reaction. The sensor generates a quantifiable light signal whose magnitude is proportional to the oxygen concentration, which is directly proportional to the glucose concentration. Alternatively, a time-dependent glucose signal can be generated, where the length of time the sensor generates a light signal after being excited by the light source 2608 is proportional to the glucose concentration. Either or both methods can be employed to calculate the glucose concentration. Light from the light source 2608 is transmitted from the SPAID 102 and down the cannula 2602 to the sensor area, exciting the photofluorescence sensor to generate return light. The light generated by the photofluorescence sensor travels directly through the wall of the cannula 2602 or needle, which is made of a light-conducting material similar to an optical conduit or optical fiber. The light is then received inside the housing of the SPAID 102 by a photodetector 2610. In an alternative embodiment, as shown in Figure 5A, the photofluorescence sensing region may be provided on a needle or cannula separate from the cannula used to deliver the liquid medication.

[0165] The sensor is composed of three sensing regions 2604, 2605, 2606 of different sizes to generate optical signals proportional to different concentrations of glucose, from low (2604) to medium (2605) to high (2606), allowing for higher resolution in the measurement. Different numbers of sensing regions can be used (e.g., 2, 4, or more). Each of the regions 2604, 2605, 2606 has an individual optical conduit(s), as shown in FIG. 26B, where the optical conduit that conducts the excitation light is labeled "S" and the optical conduit that conducts the light emitted by the sensing regions 2604, 2605, 2606 is labeled "D". The inner cannula 2612 can be used to deliver liquid medication. As with the embodiment shown in FIG. 25B, the amount, ratio, and location of the optical conduits may vary. The light-delivery conductors ("S") may be interleaved with the light-detecting conductors ("D"). Figure 26C is a conceptual diagram of this embodiment, with a close-up of the cannula 2602 in the inset.

[0166] FIG. 27A is a schematic diagram of a variation of the embodiment shown in FIGS. 26A-26C. As with the previous embodiment, the sensor may be comprised of three regions 2704, 2705, 2706 defined on a cannula 2702 of different sizes to generate optical signals proportional to different concentrations of glucose, from low to medium to high. Additionally, a fourth region 2707 directly measures oxygen to establish a baseline oxygen level in the subcutaneous ISF (Interstitial Fluid). Each region may have one or more separate optical conduits conducting excitation light ("S") and emission light ("D"). FIG. 27B is a conceptual diagram of this embodiment, with a close-up of the cannula 2702 shown in the inset. As with the previous embodiment, the photofluorescence sensing region may be provided on a needle or cannula separate from the cannula used to deliver the liquid drug, as shown in FIG. 5A.

[0167] Photoluminescent Materials

[0168] The direct and indirect glucose-to-light conversion methods mentioned above require different chemicals to construct a suitable glucose sensor. For direct glucose-to-light conversion, an acceptor / quencher molecule is needed that can directly bind to both the glucose molecule and the fluorescent dye. A molecule often used for this purpose is boric acid, which, when combined with a fluorescent dye, allows FRET quenching of the dye when a glucose molecule is reversibly bound to it. When glucose is removed, the FRET effect disappears and the fluorophore returns to normal fluorescence emission.

[0169] Furthermore, to complete the glucose-to-light direct conversion method, a photofluorochrome that can be coupled to an acceptor / quencher is required. Ideally, this dye will have a fairly large shift from the wavelength of the excitation light to the wavelength of the emission light. This shift simplifies the construction of the light source and detector system and reduces the possibility of crosstalk, where the emission light is reflected and unintentionally detected. This dye also simplifies the implementation of wavelength-specific filters. Listed below are just a few of the commercially available dyes that are suitable for this type of sensing method. Many alternatives are available and can be implemented. Furthermore, systems can be built with multiple combinations of quenchers and dyes to improve selectivity, precision, and expand the operating range of the sensor.

[0170] Commercially available dyes include, among others: (a) Becton Dickinson Pacific Blue (excitation maximum 401 nm / emission maximum 452 nm), which is based on the 6,8-difluoro-7-hydroxycoumarin fluorophore and is highly fluorescent even at neutral pH, and (b) Becton Dickinson Alexa Fluor 488 (excitation maximum 495 nm / emission maximum 519 nm), whose conjugates are highly photostable and fluoresce over a wide pH range. (c) Becton Dickinson Horizon Brilliant Violet 605 (BV605) (excitation maximum 407 nm / emission maximum 605 nm) is a tandem fluorescent dye that combines BD Horizon BV421 with an acceptor dye that emits at 605 nm, and Becton Dickinson Horizon Brilliant Violet 650 (BV650) (excitation maximum 407 nm / emission maximum 650 nm) is a tandem fluorescent dye of BD Horizon BV421 and an acceptor dye with an emission maximum at 650 nm. For indirect conversion of glucose to light, acceptor / quencher molecules can be used to bind both the oxygen molecule and the fluorescent dye.

[0171] Figure 28 is a block diagram of a possible embodiment of an electronics and control system that can be used to drive the sensor and collect quantifiable electrical signal data from various photofluorescence sensing methods. This system is similar to the system shown in Figure 16 for the electrochemical method, except that in this embodiment, the sensor-patient interface 110 is composed of one or more photofluorescence sensing cells 2802-1...2802-N, each cell being composed of a light source, which in a preferred embodiment may be an LED or laser diode of a specific wavelength to emit excitation light, and a light detection device to detect the light emitted by the fluorescent material that composes the sensor. Otherwise, the raw signals from the photofluorescence sensing cells 2802-1...2802-N are processed in the same manner as the electrochemical cells 1602-1...1602-N, and the results are used in the same manner as previously described.

[0172] Sensing Method Using an Implantable Device Next, various methods of sensing using implantable sensors are described, as well as various methods of delivering liquid medication to a user.

[0173] FIG. 29 shows an embodiment in which glucose sensing is performed by mounting the SPAID 102 directly over the implanted photofluorescent seed 2902. This embodiment operates in a similar manner to the photofluorescence method described earlier herein, except that there is no physical mounting between the SPAID 102 and the photofluorescent seed 2902. In this embodiment, light is transmitted through the user's skin by an excitation light source 2908, and light emitted by the photofluorescent seed 2902 is detected through the skin by a photodetector 2910. This embodiment stimulates the implanted seed to perform a glucose measurement, which then processes the glucose value internally in the MDA 129 or user app 160, which uses the value to deliver the required dose of liquid drug insulin using one of several insulin delivery techniques described later in this specification. This embodiment can be used with the system shown in FIG. 28. 30 illustrates an embodiment in which glucose sensing is performed by mounting the SPAID 102 in proximity to an integrated implantable glucose sensor 3002 configured to communicate with the SPAID 102 via wireless communication 3004 to obtain sensed glucose data. The implantable sensor 3002 includes a power source, a sensing system (photofluorescence, electrochemical, spectroscopic), and the necessary electronics to support data processing, storage, and wireless communication with a body-worn transmitter, or possibly directly to a smart mobile device. The wireless communication module 3004 of the SPAID 102 can communicate with the implantable sensor 3002 via one or more of several different methods, including but not limited to Bluetooth® (BLE) and / or Near Field Communication (NFC).

[0174] Microneedle Sensing Method Glucose sensing can be implemented in the SPAID 102 described herein by using a microneedle array attached to the base of the device separately from the insulin delivery source. In some embodiments, for example, the SPAID 102 may include (a) a microneedle glucose sensing array in combination with (b) an inserted insulin delivery cannula, a microneedle-based insulin delivery system, or a transdermal insulin delivery system. Details regarding each of these options and implementations are described below.

[0175] Electrochemical sensing can be used to sense glucose via a microneedle-based sensing system. In this embodiment, sensing can be accomplished by a microneedle coated with a conductive electrochemical component that allows for the formation of an electrochemical cell circuit when the microneedle is inserted into the skin and interstitial fluid (ISF). The needle may be coated with standard electrochemically conductive materials as described below.

[0176] The electrodes used to form the electrochemical cells of the microneedle-based sensing system can be configured in a number of ways. The electrodes can be configured as individual needles or as clusters of needles to form the desired electrochemical sensing cell. The cells can be configured with two, three, four, or other multi-electrode combinations, and common or separate electrical circuits can be employed to form any number of sensing regions.

[0177] Chemically active materials, including enzymes that may be required to construct the cells, may be coated directly onto the electrodes on the microneedle surface or may be embedded directly into the electrode material itself and deposited during construction of the microneedle array. A variety of methods can be used to apply the conductive and chemically active materials, including vapor deposition, electroplating, spraying, printing, or using microelectronic fabrication techniques.

[0178] Additionally, various fluid-conductive chemical moieties, such as surfactants, can be added to the surface of the needles to improve the flow of bodily fluids (e.g., interstitial fluid) necessary to transport glucose or other analytes to the surface of the microneedle.

[0179] The working electrode (WE), reference electrode (RE), counter electrode (CE), blocking membrane and glucose limiting membrane can be provided as described above.

[0180] 31 illustrates an embodiment of a microneedle-based sensing system that includes two three-electrode electrochemical cells. Each of the electrochemical cells includes a microneedle working electrode 3122, a microneedle counter electrode 3124, and a microneedle reference electrode 3128, each of which is disposed on an electrical substrate 3130 and electrically connected to a control circuit 3132 via electrical conductors 3134 (two of which are numbered). Each microneedle working electrode 3122 includes a layer 3135 of platinum or other working electrode substrate having glucose oxidase bound thereto, onto which a permeation selective or charge selective porous blocking membrane 3136 is applied. Each microneedle counter electrode 3124 includes a layer 3135 of platinum, carbon, or other counter electrode substrate. Each microneedle reference electrode 3128 includes a layer 3135 of silver / silver chloride or other reference electrode substrate.

[0181] Glucose sensing with fluid-conductive microneedles In embodiments where electrochemical sensing is used to sense glucose via a microneedle-based sensing system, sensing may be performed by a tubular or "straw-like" microneedle that allows bodily fluid containing the desired analyte to flow through the microneedle and into the housing of the microneedle array for measurement. The needle may be filled with a fluid-conductive material (e.g., hydrogel) that establishes a fluid-conductive pathway in and out of the microneedle housing through which the analyte may diffuse. Since the diffusion of the analyte is due to changes in the analyte concentration in the skin, the fluid pathway must allow movement in both directions to ensure consistency and accuracy of the concentration in the body and in the microneedle housing. The same fluid-conductive material is also employed inside the housing to deliver the analyte contained in the bodily fluid to the sensing element. Electrochemical, photofluorescent, or spectroscopic based sensing elements may be employed to analyze the analyte once introduced into the body of the microneedle housing.

[0182] In electrochemical-based sensing, the electrodes used to form the electrochemical cells within the sensor housing can be configured in multiple ways. The electrochemical cells can be configured as individual cells or in clusters to form electrochemical sensing arrays to improve performance. The cells can be configured in two-electrode, three-electrode, four-electrode, or other multi-electrode configurations, and any number of sensing regions can be formed using common or separate electrical circuits.

[0183] The chemically active substances, including the enzymes necessary to construct the electrochemical cell, can be coated directly onto the electrode surfaces within the housing or can be mixed into the fluid-conducting hydrogel material.

[0184] A variety of methods can be used to apply the electrode materials and chemically active components, such as by evaporation, electroplating, spraying, printing, or using microelectronic manufacturing techniques.

[0185] Additionally, various chemical moieties, such as surfactants, can be added to the fluid conductive material or hydrogel to enhance the flow of glucose and other analytes to the electrochemical cell within the housing.

[0186] FIG. 32 shows an embodiment of a hollow microneedle-based sensing system using a fluid conductive material that allows diffusion of the analyte to the electrochemical cell electrodes of a three-electrode electrochemical cell. The electrochemical cell consists of a working electrode 3222, a counter electrode 3224, and a reference electrode 3228 electrically connected to a control circuit 3232. The working electrode 3222 consists of a layer of platinum or other working electrode substrate with glucose oxidase bound thereto, onto which a permeable or charge selective porous blocking membrane can be applied as previously described. The counter electrode 3224 may be composed of a layer 3235 of platinum, carbon or another counter electrode substrate, and the reference electrode 3228 may be composed of a layer of silver / silver chloride or another reference electrode substrate. The working electrode 3222, counter electrode 3224, and reference electrode 3228 are disposed within a reservoir defined by a housing. The reservoir includes a fluid conductive material 3242 that allows diffusion of the analyte (represented by dots 3246). The lower wall of the housing 3240 contains microneedles having hollow tips 3248 (numbered three) that open into the reservoirs, allowing analyte 3246 in the surrounding bodily fluid to flow through the microneedles and into the reservoirs for measurement.

[0187] In some embodiments, photofluorescence sensing is used to sense analytes in a microneedle-based sensing system. In these embodiments, a photofluorescence sensing system is employed to measure analytes flowing into the sensor housing. As described above, a fluid conductive material (e.g., a hydrogel) can be used to establish a fluid pathway for glucose or other desired analytes to flow from the body into a reservoir defined by the sensor housing for measurement. In the case of photofluorescence sensing, the fluid conductive material includes fluorophores and FRET acceptor / quencher molecules as described above. The photofluorescent fluorophores and FRET acceptor / quenchers are bound or held in suspension within the fluid conductive material and react directly with glucose or other selected analytes with which the quencher reacts specifically. Inside the housing are light sources and detector elements for exciting the fluorophores and collecting the returning fluorescent light. The light sources and detector elements may be singular or multiple. When multiple light sources and / or detector elements are provided, they can be arranged in a line, array, or any functionally beneficial configuration to maximize measurement performance. Measurements are made when light is used to excite the fluorophores bound to the hydrogel. In the direct glucose-to-light conversion method, an increase in the returned light indicates an increased concentration of the analyte, while a decrease or less light indicates a decreased concentration of the analyte. In addition to the magnitude of the light, the time required for the light emitted from the fluorophore to decay after excitation can also be measured to further quantitatively determine the analyte concentration.

[0188] 33 shows an embodiment of a hollow microneedle-based sensing system using a fluid conductive material that allows diffusion of analytes to a fluorescent sensing array. The optical fluorescent sensing array includes a plurality of excitation light sources 3308 and a plurality of detectors 3310, which are electrically connected to a control circuit 3332. The excitation light sources 3308 and the detectors 3310 are disposed within a reservoir defined by a housing 3340. The reservoir includes a fluid conductive material 3342 (e.g., a hydrogel material) that allows diffusion of the analytes (represented by dots 3346). The fluid conductive material 3342 also includes fluorophores and FRET quencher molecules 3349 that are bound or held in suspension within the fluid conductive material 3342 and directly interact with glucose or other selected analytes to which the quenchers specifically react. The lower wall of the housing 3340 contains a microneedle with a hollow tip 3348 (numbered three) that opens into the reservoir, allowing analyte 3346 in the surrounding bodily fluid to pass through the microneedle and into the reservoir for measurement. If the quencher responds specifically to glucose, the system is a direct sensing system and an increase in returned light indicates an elevated glucose concentration, while a decrease or less light indicates a decreased glucose concentration.

[0189] In another embodiment, an indirect photofluorescence sensing system is employed to measure the analyte flowing into the sensor housing. As described above, a fluid conductive material (e.g., a hydrogel) is used to establish a fluid pathway for glucose or other desired analyte to flow from the body into the sensor housing to be measured. In the case of photofluorescence sensing, the fluid conductive material includes fluorophores and FRET quencher molecules, as described above. The sensor uses glucose and glucose oxidase to generate a measurable optical signal by reducing the local concentration of oxygen through a chemical reaction. The sensor generates a quantifiable optical signal whose magnitude is proportional to the concentration of oxygen, which is proportional to the concentration of glucose. Alternatively, a time-dependent glucose signal can be generated, where the length of time the sensor generates an optical signal after being excited by a light source is proportional to the glucose concentration. Alternatively, both methods can be employed to calculate the glucose concentration. The combination of an oxygen-sensitive FRET quencher and a fluorophore bound in the gel with glucose oxidase produces light that fluoresces in proportion to the concentration of oxygen present in the hydrogel. As mentioned above, the light sources and detector elements can be arranged in a matrix or any suitable pattern to achieve the desired coverage area for illuminating the fluorophore and detecting the response therefrom.

[0190] Additionally, it may be desirable to add a separate set of open-tip microneedles that establish an isolated fluid-conducting pathway to an isolated sensor housing region in order to measure oxygen concentration in bodily fluids without being affected by oxygen concentration reduction due to the reaction between glucose and glucose oxidase (GOX). This new region contains the same light source and sensing element, and the same oxygen-sensitive fluorophore and acceptor / quencher molecules in the hydrogel. The difference is that there is no glucose oxidase in the hydrogel or fluid-conducting material. Measuring the baseline level of oxygen contained in the body improves the measurement accuracy in the analyte measurement region. The baseline offset due to the natural oxygen level in the body can be subtracted from the analyte measurement. In both detection regions, glucose or other analytes are present and oxygen is measured. However, in the glucose oxidase region, the oxygen concentration depends on the glucose concentration.

[0191] One embodiment of such an indirect photofluorescence sensing system is shown in Figure 34. The indirect photofluorescence sensing system includes a first photofluorescence sensing array including a plurality of excitation light sources 3408a and a plurality of detectors 3410a electrically connected to a control circuit 3432a. The excitation light sources 3408a and the detectors 3410a are disposed within a first reservoir defined by a first housing 3440a. The first reservoir includes a fluid conductive material 3442 (e.g., a hydrogel material) that allows diffusion of glucose molecules (represented by dots 3446) and oxygen (represented by dots 3447) and includes glucose oxidase (represented by dots 3444). The lower wall of the housing 3440a includes a microneedle having a hollow tip 3448a that opens into the first reservoir, such that analytes 3446 and oxygen 3447 in the surrounding bodily fluid flow through the microneedle and into the reservoir defined by the housing 3440a. Fluid conductive material 3442 also contains fluorophore and FRET quencher molecules 3449 that are bound or held in suspension within hydrogel material 3442 and directly interact with oxygen 3447 with which the quencher specifically reacts. Thus, an increase in returned light indicates increased oxygen concentration, while less or no light indicates decreased oxygen concentration. In the presence of glucose 3446 and oxygen 3447, glucose oxidase 3444 catalyzes the oxidation of glucose 3446 while consuming oxygen 3447 to form gluconic acid and hydrogen peroxide. Thus, an increase in glucose results in a decrease in oxygen, which results in a decrease in returned light to detector 3410a, while a decrease in glucose results in an increase in oxygen, which results in an increase in returned light to detector 3410a.

[0192] The indirect light fluorescence detection system of Fig. 34 can also include a second fluorescence detection array including a plurality of excitation light sources 3408b and a plurality of detectors 3410b electrically connected to control circuitry 3432b. The excitation light sources 3408b and detectors 3410b are disposed within a second reservoir defined by a second housing 3440b. The second reservoir includes a fluid conductive material (e.g., a hydrogel material 3442) that allows diffusion of glucose molecules (represented by dots 3446) and oxygen molecules (represented by dots 3447). The fluid conductive material 3442 (e.g., a hydrogel) also includes fluorophores and FRET quencher molecules 3449 that are bound or held in suspension within the hydrogel material 3442 and directly interact with oxygen molecules 3447 with which the quencher specifically reacts. The lower wall of the housing 3440b includes a microneedle having a hollow tip 3448b that opens into the second reservoir, allowing glucose 3446 and oxygen 3447 in the surrounding body fluid to flow through the microneedle into the reservoir defined by the housing 3440b. Unlike the first reservoir, the second reservoir does not contain glucose oxidase. In this way, the concentration of oxygen 3447 in the body fluid can be measured without being affected by the decrease in oxygen concentration resulting from the reaction of glucose with glucose oxidase. As explained above, the measurement of the baseline level of oxygen 3447 contained in the body by the second optical fluorescence sensing array serves to increase the accuracy of the measurement of glucose measured by the first optical fluorescence sensing array.

[0193] In other embodiments, a spectroscopy-based sensing system can be employed to measure the analyte flowing into the sensor housing. Spectroscopy uses electromagnetic (EM) radiation of various wavelengths to excite and resonate molecules of the analyte of interest. As discussed above, a fluid conductive material (e.g., hydrogel) can be used to establish a fluid pathway for glucose or other desired analytes to flow from the body into the sensor housing for measurement. In the case of spectroscopy-based sensing, the fluid conductive material (e.g., hydrogel) serves to transport the glucose or other analyte into the sensor housing for measurement by the spectral light source and detector. The spectral light source can be used to generate electromagnetic (EM) radiation at a specific wavelength or set of wavelengths. The wavelengths can range from radio frequency (RF) wavelengths to wavelengths in the visible light spectrum to low energy ultraviolet (UV) wavelengths. The spectral light source can be a tuned RF antenna, an LED light source, a laser light source, or a combination thereof. A spectroscopy-based sensing system may be comprised of multiple spectral light sources. The wavelengths of the light sources are identified and tuned to excite specific analyte molecules. Depending on which analyte is to be measured, the wavelength of electromagnetic radiation from the light source is selected to excite the analyte molecules of interest into a vibrational state. The molecules then produce small amounts of electromagnetic radiation at wavelengths both longer and shorter than the fundamental excitation wavelength. This is known as Raman scattering. The detector can use filtering to mask the EM radiation of the fundamental excitation wavelength to improve detection of the lower intensity Raman scattering wavelengths. The fundamental excitation wavelength can be selected to be most resonant with the desired analyte molecules and to produce the most Raman scattering. Additional sources of EM radiation at other wavelengths can also be used to help further identify the concentration of analytes in the sample, exciting and / or resonating the analyte molecules at other wavelengths to produce other Raman frequencies that can be used to further identify the analyte molecules of interest.

[0194] One embodiment of a hollow microneedle-based sensing system with a fluid conductive material that allows diffusion of analytes to a spectroscopy-based sensing array is shown in FIG. The spectroscopy-based sensing array includes an electromagnetic radiation source 3508 and two electromagnetic radiation detectors 3510, which are electrically connected to a control circuit 3532. The electromagnetic radiation source 3508 and the detectors 3510 are disposed within a reservoir defined by a housing 3540. The reservoir includes a fluid conductive material 3542 (e.g., a hydrogel material) that allows diffusion of the analytes (represented by dots 3546). The lower wall of the housing 3540 includes a microneedle with a hollow tip 3548 (numbered three) that opens into the reservoir, allowing analytes 3546 in the surrounding bodily fluid to flow through the microneedle and into the reservoir defined by the housing 3540 for measurement.

[0195] Additionally, methods can be employed to pre-orient the analyte molecules to further increase the magnitude of the Raman scattered electromagnetic radiation and improve the signal to noise ratio. The system can be constructed using a narrow analyte detection region within the sensor housing in the area where the spectroscopic light source and detection elements are mounted. This narrow region can be fitted with electromagnetic plates, insulating coils, or a combination thereof to create an electromagnetic field that can orient the analyte molecules in a specific, consistent direction prior to excitation by the spectroscopic EM source. This pre-orientation of the analyte molecules serves to rotate the molecules into a more ideal orientation, increasing the amount of energy that the analyte molecules can absorb, thereby enabling them to generate a greater amount of returning Raman energy at the molecule's "signature" wavelength.

[0196] One embodiment of a hollow microneedle-based sensing system that allows for orientation of analyte molecules in conjunction with detection by a spectroscopy-based sensing array is shown in FIG. 36. The spectroscopy-based sensing array includes an electromagnetic radiation source 3608 and two electromagnetic radiation detectors 3610, which are electrically connected to a control circuit 3632. The system also includes an orientation element 3609 (e.g., an electromagnetic plate and / or coil), which forms a sensing region 3611 where the analyte molecules are oriented by polarization alignment. The electromagnetic radiation source 3608, detectors 3610, and orientation element 3609 are disposed within a reservoir defined by a housing 3640. The reservoir includes a fluid-conducting material 3642 (e.g., a hydrogel material) that allows for diffusion and orientation of the analyte (represented by dots 3646). The lower wall of the housing 3640 contains a microneedle with a hollow tip 3648 (two numbered) that opens into a reservoir, allowing analyte 3646 in the surrounding bodily fluid to pass through the microneedle and into the reservoir defined by the housing 3640 for measurement. Additionally or alternatively, the microneedle may have holes on one or both sides to allow bodily fluid to pass through the microneedle and into the reservoir for measurement.

[0197] Any of the microneedle sensing methods described above may be implemented in the SPAID 102. Figures 38A-38D show different possible placements of microneedles that allow for both glucose sensing and insulin delivery. The figures are exemplary in nature, and as will be appreciated, other implementations may be possible and are contemplated to be within the scope of the present invention. Figure 38A shows a first possible embodiment. This figure is a bottom view of the SPAID 102, showing the base 3802 that rests on the user's skin when the SPAID 102 is placed. The adhesive 3804 serves to adhere the SPAID 102 to the user's skin. The liquid drug delivery point is indicated by reference number 3806 and may be any one of a number of delivery methods described later in this specification. A microneedle sensing array 3808 is shown disposed on the base 3802 of the SPAID102 such that when the SPAID102 is placed on the user's skin, the microneedle array 3808 contacts the user's skin and measures the user's glucose level using any of the methods described above. Some microneedle structures suitable for use with the described embodiments are described in U.S. Patent Application No. 63 / 154,003, filed February 26, 2021, the contents of which are incorporated herein in their entirety.

[0198] A first variation of this embodiment is shown in FIG. 38B, where a dual microneedle sensing array 3808 is disposed on the base 3802 of the SPAID 102. FIG. 38C shows a third variation where multiple microneedle sensing arrays 3808 are disposed within the adhesive 3804 (or on the underside of a tray) that adheres the SPAID 102 to the user's body. Finally, FIG. 38D shows yet another variation where the microneedle array 3808 is shown circular and disposed on the base 3802 of the SPAID 102. Note that in the embodiment shown in FIG. 38D, the liquid drug delivery point 3806 is disposed at the center of the circular microneedle sensing array 3808. However, the liquid drug delivery point 3806 may be disposed outside of the circular microneedle sensing array 3808.

[0199] When using an open-tipped fluid-conducting microneedle sensing array 3808, it may be necessary to add a means to filter or block potentially undesirable molecules in order to reduce their ingress within the sensor housing. One possible implementation would be to use a membrane such as Nafion® or other molecularly selective membrane that has a pore size appropriate for glucose (or other desired analyte) while rejecting larger molecules that may be undesirable within the measurement housing.

[0200] Alternatively, a conductive coating or electrodes can be used to create an electric field adjacent to the tips of the microneedles of the microneedle sensing array 3808 to repel or repel unwanted charged molecules from the tips of the microneedles in the body. If the electrodes are coated with a conductive material, they will collect the molecules on the outer surface of the microneedles and prevent them from entering the microneedle sensor housing. Additionally, additional sets of electrodes can be provided within the sensor housing to further filter unwanted charged molecules if they do enter the microneedles.

[0201] The electrodes are driven by the electronics of the SPAID 102 and can generate a fixed DC voltage, or a periodic AC voltage or a pulsed voltage. If necessary, the voltage source can be isolated with respect to other electrodes in the system, especially if an electrochemical measurement system is employed within the measurement housing.

[0202] One possible installation method for the glucose-sensing microneedle sensing array 3808 involves the user firmly placing and pressing the base surface 3802 of the SPAID 102 against the skin at the recommended body location. To further enhance the penetration of the microneedles into the stratum corneum, an ultrasonic transducer (not shown) can be used to temporarily and periodically vibrate the body of the microneedle sensing array 3808. The ultrasonic vibration energy serves to vibrate the microneedle sensing array 3808 and deepen it further to a desired maximum depth controlled by the length of the needles. This is in conjunction with the pressure exerted by the adhesive 3804 that holds the base 3802 of the SPAID 102 firmly against the user's skin and the skin pressure established upon initial donning by the user creating an upward pressure on the microneedle sensing array 3808 towards the base 3802 of the SPAID 102. The adhesive 3804 is preferably disposed adjacent to and around the entire periphery of the microneedle sensing array 3808, which may be mounted so as to protrude slightly from the base. In some embodiments, additional force may be provided from the interior of the housing to the back of the microneedle array to ensure and maintain an adequate depression force of the microneedle array onto the skin surface, such as from a spring, foam, or other force system.

[0203] Spectral sensing method Glucose sensing can be performed on the SPAID 102 by measuring glucose or other analytes directly through the skin using a spectroscopic light source and detector system mounted at the base of the SPAID 102 housing. Alternatively, the SPAID 102 can use a photon-based spectroscopy system to measure glucose or other analytes directly through a light-guiding cannula inserted into the skin that may or may not also be used for insulin delivery. The SPAID 102 can be comprised of a spectroscopy-based sensing system combined with an inserted insulin delivery cannula. Alternatively, the SPAID 102 can be comprised of a spectroscopy-based sensing system combined with a microneedle-based insulin delivery system. Alternatively, the SPAID 102 can be comprised of a spectroscopy-based sensing system combined with a transdermal insulin delivery system.

[0204] As previously mentioned, spectroscopy uses electromagnetic (EM) radiation at multiple wavelengths to excite and resonate molecules of an analyte of interest. In the case of glucose sensing, glucose molecules, when excited by an EM radiation source, then resonate and return EM radiation at wavelengths both longer and shorter than the original excitation wavelength. The return EM energy from molecular resonance is called Raman scattering, and is a key component of what is known as Raman spectroscopy. Glucose molecules, once excited, resonate at specific wavelengths, and this specific pattern of wavelengths and magnitudes forms a unique "fingerprint" for the glucose molecule. The larger the magnitude at a specific wavelength, the greater the number of glucose molecules resonating in the sample. The magnitude information can therefore be used to measure the glucose concentration in a sample solution or in the body. Figure 37 shows a typical "fingerprint" plot of resonating glucose energy after glucose is excited with near-infrared light energy.

[0205] The main challenges associated with in vivo measurements include interference with other molecules present, and resonance and scattering of the electromagnetic radiation source. Additionally, there is the issue of attenuation of both the source EM radiation and the resonant returning EM energy. For in vitro glucose measurements, the EM radiation used to excite the glucose molecules is typically near-field infrared light (IR), which passes through the skin to excite glucose molecules in the interstitial fluid and in whole blood in the capillaries and veins. Water molecules, salts, amino acids, proteins, and other molecules that make up the skin layers attenuate the near-infrared light or generate additional resonant energy. Furthermore, once the glucose molecule is energized, the resonant energy specific to the glucose molecule must return outside the skin to be detected. The magnitude of the returning energy is very small and also experiences similar attenuation by the skin and tissue layers, which may further cause secondary interactions with other molecules at new resonant frequencies. Thus, the closer the system is to sampling a pure analyte with a direct path for both the excitation and resonant energy to travel from the source to the detector and back, the more accurate the system will be.

[0206] FIG. 39 shows an embodiment of the SPAID 102 in which glucose sensing is performed using a spectroscopy-based system that measures glucose through the skin. In this embodiment, a source 3908 and detector 3910 for the spectroscopy system are mounted at the base of the housing of the SPAID 102. EM energy is generated by the source 3910 and directed from the base of the SPAID 102 to the skin. The EM energy can be in the form of ultraviolet, visible, infrared, near infrared, mid infrared, far infrared, and RF energy. Glucose molecules can be vibrated with EM energy across this entire wavelength range. The wavelength used by the source is maximized to deliver the most EM energy to the glucose or other analyte while penetrating the necessary layers of the skin, depending on the mounting location of the SPAID 102. The receiver 3910 is tuned to receive the signature Raman EM energy emitted by glucose molecules when excited by the source EM energy.

[0207] As an example, if 532 nm green visible light is used as the source EM energy, the EM light source 3908 can be a small laser, light emitting diode, or other light source capable of operating in a narrow band of wavelengths around 532 nm. The source EM energy is focused onto the skin using lenses and / or other optics to concentrate and maximize the excitation energy in the area of ​​interest on the skin. Other wavelengths may also be used.

[0208] The returning Raman vibrational energy is collected using a detector 3910, which may be, for example, an array of photodetectors, lenses and mirrors, or other optoelectronic components necessary to collect the returning energy photons. A prism or other optical filtering device is employed to selectively return energy photons of a particular wavelength, or wavelengths corresponding to the Raman spectroscopic signature of a particular excitation energy. The system can be tuned to focus the returning energy at a particular wavelength that matches the peak of energy found in the Raman spectroscopic signature of a given analyte. The selectivity of the filter can improve the signal to noise ratio of the system, thereby ensuring a response at the peak associated with the analyte being measured.

[0209] The electrical signal from the photodetector is amplified if necessary and digitized using an analog-to-digital converter with sufficient resolution and sampling rate to meet the desired system output requirements. For glucose measurements, the system may need to respond in 0.1 mg / dL increments over the range 0-400 mg / dL, requiring a resolution of 4000 steps. Therefore, an A / D converter with a resolution of preferably 12 bits or more can be used. Alternatively, an A / D converter with lower resolution can be used, sampling faster than necessary, and averaging performed in software.

[0210] 40 shows a second embodiment using spectroscopy, where glucose molecules are indirectly measured by exciting secondary molecules with EM energy via a light guide 4002 inserted into the skin. The light guide 4002 may be constructed from fiber optic plastic, glass or other light conducting materials in various embodiments. The secondary molecules may be, for example, hydrogen peroxide, oxygen, or any other molecule that can be produced in direct proportion to glucose concentration in the body.

[0211] When hydrogen peroxide is used, it is generated via a glucose to glucose oxidase reaction. In this embodiment, the glucose oxidase enzyme is coated directly onto the tip of the cannula 4002, bound within a coating or structure of the cannula itself, or bound in suspension in a gel matrix. Once inserted under the skin in the subcutaneous region, glucose in the interstitial fluid reacts with the glucose oxidase to generate hydrogen peroxide directly at the tip of the lightguide 4002. This local generation of hydrogen peroxide is stimulated by generated EM energy generated by the light source 4008, which travels from the SPAID 102 to the tip 4020 of the lightguide 4002. Because the hydrogen peroxide is generated locally at the tip of the lightguide 4002, there is little interference or attenuation of the resonated return Raman EM energy from the source and the hydrogen peroxide molecules. The return Raman EM energy is detected by the detector 4010.

[0212] The block diagram of FIG. 41 shows one possible implementation of a glucose sensing system using spectroscopy through a cannula 4102. The system uses a dual lumen cannula 4102 consisting of light guides 4103, 4104 constructed in a coaxial format. The inner light guide 4103 passes excitation light from an excitation light source 4108 within the housing of the SPAID 102 to the target. The excitation light source 4108 can be a laser, LED, or other light source capable of generating an EM energy source at the wavelength, spectrum, and intensity required to excite the analyte of interest. The excitation light source energy can pass through various lenses, mirrors, prisms, or other elements necessary to properly direct the energy into and through the lumen. The outer light guide 4104 is in the form of a coaxial tube through which the inner light guide 4103 passes. The inner and outer walls of the outer light guide 4104 may be covered with a shielding coating to prevent EM energy from leaking outside the system or between the two light guides 4103, 4104.

[0213] When the excitation EM energy reaches the analyte of interest 4106, the analyte is excited and produces vibrational Raman EM energy at longer and shorter wavelengths than the original excitation. The returning energy returns to the SPAID 102 via the outer light conductor 4104 and is sent to a reflecting mirror (e.g., a 45° reflecting mirror), a selective filter prism, and photodiode detectors 4110-1...4110-N or other detectors sensitive to the returning Raman EM energy. The prisms are useful for selecting and directing specific Raman EM energy to a detector that matches the "fingerprint" energy spectrum of the desired analyte. Alternatively, the prisms can be manipulated or moved to direct energy of other wavelengths that are also part of the "fingerprint" energy spectrum of the desired analyte. Alternatively, some of the prisms can be tuned for a portion of the spectrum of the analyte and other prisms can be tuned to be sensitive to different wavelengths to determine the concentration of the analyte of interest.

[0214] The remaining components are necessary to drive the source and measure the return energy. The programmable amplifier 4108 allows the microprocessor to control the power and intensity of the EM energy source, allowing the output to be on-off controlled and pulsed. The photodiode amplifier 4111 increases the magnitude of the electrical signal from the photodiode which converts the returning Raman EM energy into an electrical current. The multiplexer 4112 is used to select one of the signals from the photodiode amplifier 4110. The ADC 4114 digitizes the analog signal from the MUX 4112 and provides the data to the controller 121 for analysis.

[0215] 41, a light source 4108 for excitation energy may be located inside the housing of the SPAID 102 and may be collimated and directed through a light guide 4103 to the tip 4120 of the cannula 4102 using a combination of mirrors, lenses, and other optical elements. Similarly, the returning resonant Raman light energy is directed through a similar series of mirrors, lenses, prisms, filters, and other optical elements to a number of photodetectors 4110-1...4110-N. The outputs of the photodetectors 4110-1...4110-N are multiplexed by a multiplexer 4114 and sent through amplifiers, analog-to-digital converters, and to a controller 121 running the MDA 129 (and / or user app 160 running on the user device 105) that analyzes the collected signals.

[0216] The inset of Figure 41 shows an arrangement of multiple photodetectors 4110-1...4110-N. In this embodiment, eight photodetectors are shown, but it will be appreciated that any number of photodetectors may be used.

[0217] One important aspect is the structure of the tip 4120 of the cannula 4102 that is inserted into the body and into the interstitial fluid. In one possible embodiment, the tip 4120 is made of a Fresnel lens. All light from the surface of the Fresnel lens is returned to a single point. For the measurement of hydrogen peroxide concentration proportional to glucose, the lens surface is coated with glucose oxidase by binding GOX to the lens material or other binders or coatings, or hydrogels. The generation of hydrogen peroxide at the lens surface allows both the reception of excitation EM light energy and the return of vibrational Raman energy.

[0218] Alternative cannula tip shapes can be used to direct the delivered EM energy and return the Raman resonance energy, as shown in Figures 42A-42B. One possible shape is a cone shape 4206 as shown in Figure 42A, where the light travels down the central tip 4202 of the cone shape 4206 and the return Raman energy travels down a secondary return lumen 4204. Another alternative cannula tip shape 4208 is a hemispherical shape where the light is emitted from the center 4202 and the return Raman energy returns via the secondary return lumen 4204, as shown in Figure 42B.

[0219] In both of the above embodiments of Figures 42A and 42B (cone tip 4206 and hemisphere tip 4208), the inner surface of the return lumen 4204 may be coated with or bonded with GOX. The source light travels down the central lumen 4202 and reflects off the suspended reflector cone 4206 or hemisphere 4208, allowing EM energy to radiate back in all directions towards the GOX coated inner cone 4206 or hemisphere 4208 surface. This energy then reacts with the generated hydrogen peroxide, radiating its Raman return energy back into the coaxial outer lumen 4204 and back to the detector 4010 mounted in the housing of the SPAID 102.

[0220] Insulin Delivery Methods The SPAID 102 includes a delivery patient interface 186 for interfacing with a user to deliver liquid medication from either or both of the reservoirs 124, 124-2 to the user. Various embodiments of the delivery patient interface 186 are now described. Any one of the disclosed embodiments of the delivery patient interface 186 may be combined within the SPAID 102 with any of the glucose detection methods previously discussed (electrochemical sensing, photofluorescence sensing, and spectroscopic sensing) or any other detection method not discussed herein.

[0221] In one embodiment shown in FIG. 43, the delivery patient interface 186 comprises a cannula for delivering liquid medication into the user's body, which may be subcutaneous, intraperitoneal, or intravenous. If the delivery patient interface 186 is a cannula, the SPAID 102 may further include an insertion mechanism 188 for inserting the cannula into the user's body, which may be integral to the SPAID 102 or may be attachable to the SPAID 102. The insertion mechanism 188 may, in one embodiment, comprise an actuator that inserts the needle and cannula 4302 under the user's skin and then retracts the needle leaving the cannula 4302 in place. After use, the insertion mechanism 188 may remain in the SPAID 102 or, if attachable to the SPAID 102, may be detached therefrom. Once the SPAID 102 has expired, the cannula 4302 is retracted from under the user's skin by an actuator (not shown), e.g., a spring-loaded retraction device, before the SPAID is removed from the user's body. This retraction actuator may be part of the insertion mechanism 188 or may be separate therefrom.

[0222] Another embodiment of the delivery patient interface 186 is shown generally in Figure 44A, where one or more microneedle arrays 4402 may be used to deliver liquid medication. Figure 44B is a conceptual diagram of this embodiment, with an inset showing a close-up of the microneedle array 4402. Also shown is a detection means 4404, shown here as a wireless communication antenna such as a near-field antenna, although any detection means may be used in conjunction with the microneedle array 4402 delivery method.

[0223] Insulin delivery can be achieved by using an array of solid-tip microneedles. The microneedle array may be coated with insulin bound to a charged molecule or other binding molecule, and / or an electric field may be used to hold the insulin on the surface of the solid microneedles. In some embodiments, the needles are covered with a grid of electrically conductive plates, each of which may be individually connected to control electronics. Each plate may hold surface-coated insulin molecules that are bound via an electric field. The size of the conductive plate is directly proportional to the amount of insulin molecules bound to its surface. The size of the plate depends on the design and can be sized to meet the minimum amount or dose of insulin to be released. When the microneedle array is placed on the skin and in contact with the interstitial fluid, the insulin molecules remain bound to the microneedle surface via the electric field generated by the charged molecules bound to the conductive plate. Upon a request from the control system to deliver insulin to the skin, the electric field associated with an individual plate, or a portion of an individual plate, is removed, and the insulin is released from the needle surface of that portion or plate, diffuses into the interstitial fluid, and is then delivered into the body. The amount of insulin released from a single conductive plate can be designed to meet the minimum insulin dose required. If a larger dose is required, additional conductive plates or portions can be de-charged, thereby releasing additional insulin payloads. This release process can continue until insulin has been released from all conductive plates.

[0224] Insulin release can be controlled and triggered by an AID control system that receives glucose level information from the glucose sensor and makes subsequent insulin delivery decisions to maintain glycemic control. One embodiment of such a system is shown in Figure 45, which is a schematic diagram of a microneedle array in which each microneedle is covered with a grid of individual conductive plate portions 4509 (three numbered), which are further coated with insulin (not numbered) and individually connected to control electronics 4532.

[0225] Alternatively, to support a charge-based release process, the tip of each needle can be held at system ground potential, or at an opposite charge to the conductive plate, allowing a small current to flow across each needle, drawing the molecules towards the tip in a controlled manner and pushing them deeper into the skin for more rapid diffusion.

[0226] In other embodiments, the system is configured to provide charge only in conjunction with the release of molecules, using a constant charge to hold the molecules and not remove it to release the molecules. A further option is to employ pulsing or duty cycling of the charge to proportionally release the bound molecules. This can be an additional method of release control, rather than total or no release from each plate. For example, each entire needle can correspond to one coated plate, and timed pulses or duty cycling can be used to drive the molecule release process.

[0227] In some embodiments, insulin delivery can be achieved by using an array of hollow-tipped microneedles. The array can be connected to an internal chamber or reservoir in the housing of the device that allows free flow of bodily fluids, such as interstitial fluid. Each of the microneedles and the internal chamber is filled with a hydrogel matrix or other fluid-conductive medium. Inside the housing, one or more tubes or cannulas connect the hydrogel matrix to an insulin delivery control system. The control system can include, for example, a reservoir for insulin, a pump (electrical or mechanical), a fluid control valve, a microprocessor-based control system, a power source, various system feedback mechanisms, and a user interface control and communication system. When insulin is required, for example, by a time-based delivery protocol, by user request, CGM request, or other insulin control criteria, it is released or pumped into the gel matrix in the housing chamber. The insulin then flows into the body through the gel matrix by diffusion. Because the concentration of insulin is higher in the microneedle housing than in the body, the diffusion process serves to equalize the difference, so that insulin flows out of the housing chamber through the microneedles and into the skin.

[0228] The microneedles may be of any suitable size, for example in the range of 100-2000 microns in length, depending on the thickness of the skin and the need to reach into the body to connect with the interstitial fluid layer. The outer diameter may be in the range of, for example, 50-250 microns, and the needles are advantageously conical in shape, narrowest at the tip where they are inserted into the skin. The inner diameter may be in the range of, for example, 10-40 microns at the tip. The microneedle array may be arranged in any suitable pattern. For example, a 10x10 square grid arrangement with needle spacing of 500-2000 microns is conceivable, although many other patterns and spacings are viable options.

[0229] One embodiment of a hollow microneedle-based delivery system is shown in FIG. 46, where the interior housing chamber / reservoir is filled with a hydrogel matrix 4642 or other fluid conductive medium to allow diffusion of insulin (represented by dots 4646). The lower boundary of the housing chamber 4640 contains multiple microneedles with hollow tips 4648 (two are numbered) that open into the interior housing chamber. Multiple insulin supply ducts or tubes 4651 fluidly connect the hydrogel matrix 4642 to an insulin supply control system, which may include other components such as the insulin reservoir and pump system described above. Once delivered into the housing chamber, the insulin 4646 flows through the gel matrix 4642 and out the hollow tips 4648 of the microneedles.

[0230] As previously mentioned, insulin can be delivered by using an array of hollow-tipped microneedles and using diffusion pressure to facilitate transport of insulin from areas of high concentration in the housing chamber to areas of low concentration in the skin or body. Additionally, iontophoresis can be used to enhance or accelerate transport of insulin from the housing chamber into the body by pushing or pulling insulin through the hydrogel from inside the housing chamber to the outside of the microneedles in the skin. This is accomplished by setting up a constant or varying electric field using electrical conductors attached between the area in the housing chamber where the system delivers insulin to the housing chamber and the outer tip of each microneedle. Depending on the pH of the skin and the charge state of the insulin molecules, a particular voltage or set of voltages with different pulse profiles or static voltage levels can be employed. The resulting electric field serves to push or pull the insulin molecules from inside the housing towards the outer tip of the microneedle. Some of the microneedles in the array can also be configured as pH sensors to provide feedback to the system, thereby altering the electric field strength or time to achieve the desired system performance. Once enough insulin has been administered outside the microneedles, the iontophoresis system can be turned off for a period of time to allow diffusion through the body to take over. Iontophoresis of insulin reduces the delay between when an insulin delivery request is received by the SPAID system and when the body receives the insulin. This system does not affect the amount of insulin delivered, but instead controls how quickly a measured dose of insulin enters the body.

[0231] One embodiment of an iontophoresis-based delivery system is shown in FIG. 47, where an interior housing chamber is filled with a hydrogel matrix 4742 or other fluid conductive medium to allow diffusion of insulin (represented by dots 4746). The lower boundary of the housing chamber includes a plurality of microneedles having hollow tips 4748 (two numbered) that open into the interior housing chamber. A plurality of non-conductive insulin delivery ducts or tubes 4751 fluidly connect the hydrogel matrix 4742 to an insulin delivery control system 4752, which may include an insulin reservoir and pump system as described above among other components. Once delivered into the housing chamber, insulin 4746 can flow through the gel matrix 4742 and out of the hollow tips 4748 of the microneedles. The iontophoresis-based delivery system further includes a first electrical conductor 4754 (sometimes referred to herein as an electrode) disposed at the upper boundary of the housing chamber and serving as a first pole of the iontophoresis system. The first electrical conductor 4754 may be in the form of a continuous conductive surface disposed between the non-conductive insulin delivery duct or tube 4751. The remainder of the interior of the housing is similarly non-conductive. A plurality of second electrical conductors 4756 (which may also be referred to as electrodes) are disposed at the outer tip of each microneedle. The plurality of second electrical conductors 4756 may all be electrically connected together to function as the second pole of the iontophoresis system. The remaining outer portion of the microneedle and the bottom (skin side) of the housing may be electrically insulated and non-conductive. Application of an appropriate voltage between the electrical conductor 4754 and the plurality of second electrical conductors 4756 creates an electric field that causes insulin molecules 4746 to migrate out of the needle tip 4748 and collect near the outer electrode. The electric field may be periodically switched off to allow the migrated insulin molecules to diffuse into the body.

[0232] In an alternative configuration of the iontophoresis-based delivery system, a third pole is provided. For example, the electrical conductors 4756 of every other (alternating) microneedle may be electrically connected to form the second and third poles of the iontophoresis-based delivery system. This arrangement can be used to create an electric field between adjacent microneedles, which can then be used to force previously collected insulin molecules through the surface of the microneedles and into the body. In general, the amount of current is controlled so as not to cause discomfort to the patient and not to exceed regulatory requirements.

[0233] As previously indicated, insulin delivery can be achieved using an array of hollow tip microneedles, using diffusion pressure to facilitate transport of insulin from areas of high concentration within the housing to areas of low concentration within the skin and body. Additionally, ultrasonic sonophoresis can be used to enhance or facilitate transport of insulin from the housing into the body. Ultrasonic sonophoresis vibrates or agitates the insulin molecules through the hydrogel from inside the housing to the tip of the microneedle, further driving insulin more quickly into the deeper layers of the skin. This can be achieved by using an ultrasonic transducer attached to the inside of the microneedle housing near the insulin delivery tube and in direct contact with the hydrogel matrix inside. The vibrations produced by the ultrasonic transducer vibrate the microneedle housing, the hydrogel, and the needle itself. This vibration increases transport of insulin molecules from the tip of the hollow microneedle through the hydrogel to the deeper layers of the skin, where the body can absorb and use the insulin, in a process known as sonophoresis. The ultrasound sonophoresis system works by opening up holes in the skin tissue beneath the microneedle array, allowing the physically larger insulin molecules to diffuse more easily and quickly into the skin. The insulin molecule is approximately 5,800 Daltons (approximately 9.63 x 10 -24 For example, a typical low-frequency ultrasound sonophoresis system in the range of 20 kHz to 120 kHz has a capacitance of 25,000 Da (approximately 4.15 × 10-23 kg) ~ 50,000 Da (approx. 8.30 × 10 -23 kg) of molecules can be delivered directly through the epidermis. By using microneedles in conjunction with ultrasonic sonophoresis, the rate and consistency of diffusion can be significantly increased.

[0234] Additionally, a common problem with open-tip microneedles is clogging due to the accumulation of the body's wound-healing components, such as proteins, blood cells, antibodies, and macrophages. Over time, these components can block the openings in the microneedle tips and prevent insulin diffusion. Regular use of an ultrasound sonophoresis system can help prevent the formation of these components, allowing the insulin delivery pathway to remain open and allow for sustainable diffusion.

[0235] An additional embodiment is the combined use of the iontophoresis and sonophoresis systems described above, both mounted within the microneedle housing. These two systems can work together to increase the speed and efficiency of insulin delivery, as well as lengthen the period that the intact SPAID 102 remains functional and effective in place in the body.

[0236] One embodiment of such a delivery system is shown in FIG. 48, where the interior housing chamber is filled with a hydrogel matrix 4842 or other fluid conductive medium to allow diffusion and / or movement of insulin (represented by dots 4846). The lower boundary of the housing chamber includes a plurality of microneedles having hollow tips 4848 (two numbered) that open into the interior housing chamber. A plurality of non-conductive insulin delivery ducts or tubes 4851 fluidly connect the hydrogel matrix 4842 to an insulin delivery control system 4852, which may include an insulin reservoir and pump system as described above among other components. Once delivered into the housing chamber, the insulin 4846 can flow through the gel matrix 4842 and out of the hollow tips 4848 of the microneedles. The delivery system also includes a first electrical conductor 4854 and a plurality of second electrical conductors 4856 disposed at the outer tips of each microneedle, as described above. The delivery system further includes an ultrasonic transducer 4853. Application of an appropriate voltage between electrical conductor 4854 and a plurality of second electrical conductors 4856 creates an electric field that causes insulin molecules 4846 to migrate out of needle tip 4848 and collect near the outer electrode. Transport (movement and / or diffusion) of insulin molecules 4846 is increased by vibrations produced by ultrasonic transducer 4853. Thus, both iontophoresis and sonophoresis support the delivery process in the illustrated embodiment.

[0237] The insulin delivery system described above can be combined with any of the glucose monitoring systems described above. However, if the iontophoretic insulin delivery system described above is combined with an electrochemical glucose monitoring system in the same SPAID unit, the two systems must be electrically isolated from each other. Alternatively, the two systems may run from different power sources or batteries and have isolated electrical ground paths. This prevents "crosstalk" between the electrochemical monitoring system and the iontophoretic-based system.

[0238] A potential problem with using hollow microneedles that penetrate the skin is the wound healing response from the body, which can affect both insulin delivery as well as glucose or other analyte measurements. These effects can be minimized by providing a microneedle that is as traumatic as possible to limit the body's reaction. Trauma can be reduced by proper selection of needle size, shape, and length. However, it can still be difficult to minimize the body's healing response at the insertion site. In addition, it can be difficult to prevent body components from penetrating the tip of the needle and coagulating to form a clot or blockage. Due to these problems, the use of coatings and anticoagulants may be desirable. These coatings can be applied to both the inner and outer surfaces of the needle. The inner surface may be coated with a coating that reduces or prevents the formation of clots or clots, thereby keeping the needle pathway open and free to use for insulin delivery and / or analyte measurement in the case of CGM or other analyte measurement systems. The exterior surface may be coated with a wound healing inhibitor, which may slow the wound healing response and allow the site to remain functional for insulin delivery and analyte measurement between cycles of wearing SPAID 102. Exemplary drugs that act as wound healing inhibitors include cytotoxic anti-tumor and immunosuppressants, corticosteroids, non-steroidal anti-inflammatory drugs (NSAIDs), and anticoagulants.

[0239] In addition to iontophoresis and sonophoresis, additives can be combined with the hydrogel, i.e., the hydrogel that allows insulin to be transported out of the microneedle and glucose or other analytes to be transported into the microneedle. Additives useful for this purpose include surfactants. Surfactants increase the flow or permeation of molecules through the hydrogel in the microneedle by effectively making the hydrogel's path more "slippery". Potential surfactants include docusate (dioctyl sodium sulfosuccinate), alkyl ether phosphates, benzalkonium chloride (BAC), perfluorooctane sulfonate (PFOS), sodium dodecyl sulfate (SDS), cationic surfactants (e.g., a homologous series of alkyl-N,N,N-trimethylammonium compounds with different alkyl chain sizes), zwitterionic surfactants (e.g., N-dodecyl-N,N-dimethylammonium-propanesulfonate, DDPS), and nonionic surfactants (e.g., Triton® X-100).

[0240] In one possible method of installation of the insulin delivery microneedle array, the user firmly places and presses the SPAID102 base against the skin at the recommended body site. To further enhance the penetration of the microneedles into the stratum corneum, an ultrasonic transducer can be used to temporarily and periodically vibrate the body of the microneedle insulin delivery array. The ultrasonic vibration energy can serve to vibrate the array and advance it deeper to a desired maximum depth controlled by the needle length. This effect can be combined with the pressure exerted by the adhesive that can hold the SPAID102 firmly to the skin, and the skin pressure established during initial installation by the user that creates an upward pressure on the microneedle array towards the base of the SPAID102. In this design, the adhesive is placed close to the array, all around the array, and the array is attached to protrude slightly beyond the SPAID base surface. Additionally, the microneedle array 4402 can have a force applied to the back of the microneedle array 4402 from inside the housing, such as from a spring, foam or other force system, to ensure and maintain an appropriate downward force of the microneedle array 4402 on the skin surface.

[0241] Another alternative embodiment of the delivery patient interface 186 is shown in FIG. 49A and consists of a transdermal delivery method consisting of a transdermal patch 4902 attached to the base of the housing of the SPAID 102 to deliver insulin directly through the surface of the skin. Transdermal insulin delivery can be combined with any of the glucose sensing systems previously discussed herein. FIG. 49B is a conceptual diagram of this embodiment, with a close-up of the transdermal patch 4902 in the inset. Also shown is a detection means 4904, shown here as a wireless communication antenna such as a near-field antenna, although any detection means can be used with the transdermal patch 4902 delivery method, such as single needle sensing, microneedle sensing, or spectroscopic sensing.

[0242] As previously mentioned, insulin delivery can be achieved using a technique known as iontophoresis. Iontophoresis can be used to enhance or facilitate transport of insulin from the housing of the SPAID102 into the body by pushing or pulling the insulin molecules from inside the housing through the hydrogel matrix to the external housing of the SPAID102 and into the skin. This is accomplished by setting up a constant or varying electric field using electrical conductors attached between the location of the internal housing where the pump delivers insulin to the internal hydrogel matrix and the external bottom surface of the SPAID102. Depending on the pH of the skin and the charge state of the insulin molecules, a specific voltage or set of voltages with different pulse profiles and / or static levels can be employed. The driven electric field serves to push or pull the insulin molecules from inside the housing through the hydrogel matrix towards the counter conductor. As the counter conductor is attached to the base of the SPAID102 away from the insulin source, the insulin is forced to move through the skin towards the counter conductor, which drives the insulin molecules into the stratum corneum and through the stratum corneum into deeper layers of the skin. Additionally, an additional skin sensing conductor can be employed to sense the pH of the skin to provide feedback to the system so that the electric field strength, waveform, pulse time, or duration of the electric field can be altered to achieve the desired system performance. Once insulin is administered to the skin, the iontophoresis system can be turned off for a period of time for diffusion into the body. The movement of insulin from the housing into the body reduces the delay from the time the insulin delivery request is received by the SPAID 102 to the time the body receives the insulin. This system does not affect the amount of insulin ultimately delivered, but can control how quickly the measured amount of insulin enters the skin.

[0243] FIG. 50 shows the arrangement of electrical conductors as part of the iontophoresis system of the SPAD 102. One of the conductors 5054 is placed on the top of the housing as a continuous conductive surface with a non-conductive insulin delivery duct or tube 5051 placed between them. The insulin delivery duct or tube 5051 fluidly connects the first hydrogel matrix 5042a to an insulin delivery control system 5052, which may include an insulin reservoir and pump system as described above. The remainder of the housing is non-conductive with the exception of the counter pole conductor 5056. The counter pole conductor 5056, which is necessary to create the electric field, is placed on the bottom surface of the SPAID housing to provide a stable low impedance connection to the surface of the skin 5099. The second hydrogel matrix 5042b is placed on top of the counter pole conductor 5056. Application of an appropriate voltage between electrical conductor 5054 and counter conductor 5056 via iontophoresis control circuitry 5053 creates an electric field that causes insulin molecules 5046 to migrate towards counter conductor 5056. The insulin molecules 5046 are also shown flowing out through the hydrogel matrix 5042a and migrating through the skin 5099 towards counter conductor 5056. The system is configured to periodically turn off the electric field to allow the migrated insulin molecules to diffuse deeper into the skin for absorption by the body.

[0244] These patterns may be, for example, sinusoidal, triangular, trapezoidal, square, pulsed, or steady DC levels, or any desired waveform deemed effective for delivering insulin to the skin. Furthermore, the waveforms may be bipolar, alternating, or a combination of AC and DC. The current used to create the electric field may be at any level up to the maximum set by the medical device electrical safety standards. The voltage used may be any voltage limited by the medical device electrical standards required to deliver the desired current. The system may be driven in constant voltage, constant current, constant impedance, or other modes useful for delivering insulin to the skin effectively. Furthermore, to prevent adverse skin reactions, the system may be capable of being turned off completely, with the conductors set to high impedance levels relative to each other.

[0245] Several possible arrangements of the iontophoresis counter electrode conductor and insulin source housing are described below. Designs are not limited to these configurations. Care is taken to not place the CGM sensing region between the insulin source electrode conductor and the counter electrode conductor to avoid migration of insulin molecules across or near the CGM sensing region.

[0246] FIG. 51A is a bottom view of the SPAID 102 showing the base 5102, which rests on the user's skin when the SPAID 102 is placed. The adhesive 5104 serves to adhere the SPAID 102 to the user's skin. The SPAID 102 includes a glucose sensing area 5106. This figure also shows one possible location of the polar conductors that are part of the iontophoresis system. In this implementation, one of the polar conductors (insulin delivery polar conductor 5154) is present on the top of the base 5102. The remainder of the base 5102 is non-conductive, except for the counter polar conductor 5156, which is located on the bottom portion of the base 5102. A first variation of this embodiment is shown in FIG. 51B, where the insulin delivery polar conductor 5154 and the counter polar conductor 5156 are present on the base 5102 in a side-by-side arrangement. Fig. 51C shows a third variation, where an insulin delivery pole conductor 5154 is disposed near the center of the base 5102, and multiple counter pole conductors 5156 are disposed in an adhesive 5804 that adheres the SPAID 102 to the user's skin. Finally, Fig. 51D shows yet another variation, where an insulin delivery pole conductor 5154 and a counter pole conductor 5156 are disposed on the base 3802 of the SPAID 102, the counter pole conductor 5156 is ring-shaped, and the insulin delivery pole conductor 5154 is circular and disposed within the center of the ring-shaped counter pole conductor 5156.

[0247] Insulin delivery can be enhanced using a technique known as electroporation. More specifically, the transport of large molecules such as insulin from the SPAID102 housing into the body can be achieved using electroporation to deliver insulin molecules or molecules of up to 10 kDa (approximately 1.66×10 -23The electroporation can be enhanced by enlarging or opening new pathways to the skin large enough to pass other large molecules, up to 1000 g (1 kg). This is accomplished by using electrodes attached to the base of the SPAID102 that directly contact the skin surface, via a conductive hydrogel matrix or other conductive electrical medium. The electroporation electrodes can generate voltage pulses with a pulse width of up to 2 ms. Pulses can occur as frequently as once every 500 ms. The number of pulses in one electroporation session can be up to 300. The session can be repeatable as many times as necessary during use of the SPAID102 to maintain the pathways for delivering insulin or other molecules to the skin. The electroporation voltage can be a DC voltage of up to 300 V to overcome natural skin resistance. The current is generally limited to meet all medical device electrical safety standards. Alternatively, low frequency AC voltages can be employed to open the pathways as well.

[0248] An additional consideration is the electrical isolation of the glucose sensing system co-located with the insulin electroporation delivery system. The high voltage electrical pulses generated by the electroporation delivery system can damage or at least affect the measurements obtained from the glucose sensing system. Therefore, electrical isolation of the glucose measurement circuitry from the electroporation delivery system can be employed.

[0249] 52 is a schematic diagram showing an exemplary embodiment of a possible implementation of an isolated electroporation circuit. Electrical isolation of the glucose measurement system from the electroporation delivery system can be achieved in one embodiment by isolating the power control signals and poration on / off signals provided by the SPAID processor from the high voltage portion of the poration circuit using optoisolators 5202, 5204, or equivalent circuitry. The high voltage required for poration is generated by a high voltage booster 5206 and delivered to the user via electrodes 5208a and 5208b. In some embodiments, additional electrodes 5208c... may be connected together as needed.

[0250] FIG. 53 shows one possible arrangement of electroporation electrodes as part of the electroporation system of the SPAD 102. In this implementation, one of the conductors 5354 is placed on the top of the housing as a continuous conductive surface with a non-conductive insulin delivery duct or tube 5351 placed between them. The insulin delivery duct or tube 5351 fluidly connects the first hydrogel matrix 5342a to the insulin delivery control system 5352. The remainder of the housing is non-conductive except for the counter pole conductor 5356. The counter pole conductor 5356, which is necessary to create the electric field, is placed on the bottom surface of the SPAID housing to make a stable low impedance connection to the surface of the skin 5399. A second hydrogel matrix 5342b is placed over the counter pole conductor 5356. Application of an appropriate electroporation voltage pulse between electrical conductor 5354 and counter-polarity conductor 5356 via electrophoresis control circuitry 5353 enlarges or opens up new pathways into the skin, allowing insulin molecules 5346 to penetrate the skin 5399. The diagram depicts the flow of insulin molecules 5346 out of the pump housing and into the skin 5399 through the pathways opened by the electroporation system.

[0251] Alternatively, if a cannula or needle-based CGM system is used as part of the SPAID 102, one of the insulated electroporation electrodes may be attached to the cannula or needle. This allows current from the electroporation system to pass from inside the insulin delivery housing through the skin and back through the needle or cannula electrode. This eliminates the need for an external electrode on the skin surface, which may be beneficial in creating a deeper, more conductive pathway in the skin.

[0252] Additionally, it can be seen that possible implementations of an iontophoresis system are very similar to those of an electroporation system, and in this regard, the electrical conductors and counter conductors of an electroporation system can be positioned relative to each other and to the glucose sensing region on the SPAID base along the lines shown in Figures 51A-51D.

[0253] The use of both iontophoresis and electroporation systems can further enhance the delivery performance of the transdermal insulin delivery system. To utilize both systems, the electrodes can be implemented identically to each other. The electronic circuitry used is also very similar. Both circuits can be operated in isolation from the main SPAID control processor through an isolation barrier. The control and timing of which system is active for what period and at what performance level can all be controlled by the SPAID processor. One possible usage scenario is to use the electroporation system after the system is initially installed by the user. After the electroporation process is complete, the electroporation system can be turned off and the iontophoresis system can be operated internally electrically connected to the same electrodes. The iontophoresis system can be used periodically to increase insulin flow. Additionally, the iontophoresis system can be turned off periodically and the electroporation system can be enabled to regenerate or create new insulin conduction pathways. The timing of when these events occur and their duration can be controlled by the SPAID processor. Additionally, the system can be configured to monitor the body's response to insulin delivery based on CGM feedback and apply either or both of the systems with greater frequency, intensity or duration to improve insulin delivery.

[0254] Transdermal delivery of insulin can be improved by using an ultrasonic sonophoresis system. Ultrasonic sonophoresis uses low frequency ultrasound and an acoustic coupling medium to generate cavitation bubbles both under and on the surface of the skin. The radius of the bubbles generated is approximately 150 μm for 20 kHz ultrasound. During the sonophoresis process, cavitation bubbles are formed and then collapse. When bubbles form on the outer surface of the skin in an acoustic coupling medium at a distance equal to the radius from the skin, the collapse force from the bubbles generates microjets into the stratum corneum, causing perturbation of the skin. Such skin perturbation can then be utilized for the delivery of insulin and other macromolecular drugs.

[0255] Ultrasonic sonophoresis can be achieved through the use of an ultrasound transducer that is mounted inside a housing close to the insulin delivery tube and in direct contact with the internal hydrogel matrix. Ultrasonic vibrations generated by the transducer travel through the hydrogel matrix and create cavitation bubbles on the outer surface of the skin. The hydrogel may be formulated to have acoustic properties that match the acoustic impedance of the skin, allowing for maximum energy transfer. The transducer can be operated in the range of 20-100 kHz and has a diameter of 1 cm. 2 The ultrasonic transducer has a power transfer capability of up to 2 W per transducer. The control system of the ultrasonic transducer can include the ability to set the duty cycle, number of cycles, pulse duration, and overall operating time of the device to achieve the desired sonophoretic skin perturbation effect.

[0256] 54A shows a possible embodiment of an ultrasonic sonophoresis skin perturbation system to assist in the delivery of insulin and other large molecules through the skin as part of a sonophoresis system for SPAD 102. In this implementation, an insulin delivery duct or tube 5451 fluidly connects the hydrogel matrix 5442 to an insulin delivery control system 5452. Also shown is an ultrasonic transducer 5453 controlled by a sonophoresis control unit 5455. Upon actuation of the transducer 5453, a skin perturbation occurs, which causes insulin molecules 5446 to penetrate the skin 5499.

[0257] Another possible layout of a sonophoresis system is shown in Figure 54B, which is a bottom view of the SPAID 102 and shows a portion of the base 5402 (not to scale) that rests on the user's skin when the SPAID 102 is deployed. In this view, multiple ultrasound transducers 5453 are arranged in a ring around the insulin delivery port 5451. A thin layer of acoustic impedance matching hydrogel covering the transducers 5453 can also be employed to aid in the transmission of ultrasound energy to the skin.

[0258] The delivery of insulin through the skin can also be improved in any of the above embodiments by the use of chemical enhancers. When used in hydrogels or other delivery mechanisms on the skin surface, chemical enhancers diffuse into the lipid bilayer of the stratum corneum and disrupt the lipids of that layer, allowing nanometer-sized openings or perturbations in that layer, which allow insulin or other large molecules to be transported and delivered to deeper layers of the skin and facilitate absorption by the body. Examples of chemical enhancers, which can be administered topically in wipes or encapsulated in hydrogels or lotions that contact the skin, include trypsin, castor oil, iodine, oleic acid, etc.

[0259] Additionally, nanocarriers can be used as drug transport vehicles in any of the above embodiments to improve the delivery of insulin or other large molecules through the skin. Insulin or large molecules are encapsulated by the transport vehicle and aid in transport through the lipid bilayer of the stratum corneum. Transport vehicles include lipid-based vesicles, nanoparticles (CaCO3), and nanoemulsions, which can significantly improve the transport of insulin through the skin.

[0260] Insulin delivery can be achieved by using any of the enhancement techniques mentioned above in combination. For example, both iontophoresis and electrophoresis systems can utilize the same electrodes and hydrogels to contact the skin. Furthermore, most of the circuitry is the same, with insulation, power, and controls being very similar, with the main differences being in the strength, shape, and duration of the electrical signal. Using both systems can be very beneficial, with electrophoresis opening a pathway in the skin and iontophoresis being used to utilize that pathway to further enhance the transport of insulin and other large molecules. Additionally, with iontophoresis and electrophoresis systems, chemical enhancers can be added to the system to further enhance the transport of insulin through the skin.

[0261] Finally, the addition of sonophoresis to the system can complement the use of iontophoresis, electrophoresis, and chemical enhancers to further open and maintain pathways to the skin.

[0262] Software-related implementations of the techniques described herein may include, but are not limited to, firmware, application specific software, or any other type of computer-readable instructions that may be executed by one or more processors. The computer-readable instructions may be provided via a non-transitory computer-readable medium. Hardware-related implementations of the techniques described herein may include, but are not limited to, integrated circuits (ICs), application specific ICs (ASICs), field programmable arrays (FPGAs), and / or programmable logic devices (PLDs). In some examples, the techniques described herein and / or any system or configuration components described herein may be implemented using a processor executing computer-readable instructions stored on one or more memory components.

[0263] Those skilled in the art to which the present invention pertains may realize many modifications and adaptations of the present invention. The implementations provided herein, including the size, shape, nominal composition and specifications of the various components or component arrangements, as well as the description of certain manufacturing steps, should be considered as illustrative and not limiting of the present invention in any way. As will be appreciated by those skilled in the art, many variations of the implementations described herein are possible and fall within the scope of the present invention. Furthermore, it should be understood that the features of the various embodiments described herein are not mutually exclusive and may exist in various combinations and permutations without departing from the spirit and scope of the present invention, even if such combinations or permutations are not expressly stated herein. Thus, the methods and apparatus disclosed herein should be taken as illustrative of the present invention, and not as limiting thereof. The scope of the present invention is defined by the following claims.

Claims

1. A device, comprising: Housing; one or more reservoirs (124; 124-2) disposed in said housing; a fluid delivery mechanism (125; 125-2) for delivering fluid from said one or more reservoirs (124; 124-2) to a user; a sensing mechanism (108; 109; 2302) for sensing an analyte level of the user; and a controller (121) for analyzing the sensed analyte level, determining the amount and timing of the fluid to be delivered, and controlling the delivery mechanism to deliver the fluid to the user.

2. 10. The device of claim 1, wherein the fluid delivery mechanism comprises a cannula (302; 402; 502; 2512) designed to be inserted subcutaneously into the user, and further wherein the sensing mechanism is integral with the cannula.

3. The device of claim 2 , wherein the fluid is delivered through a lumen of the cannula and expelled from an open tip (306; 504) of the cannula (302; 502).

4. The device of claim 2 , wherein the cannula has a closed tip, and further wherein the fluid is discharged through a set of perforations (404) in a wall of the cannula (402).

5. The device of any one of claims 1 to 4, wherein the sensing mechanism comprises a cannula (506; 2304; 2402; 2502).

6. The device according to any one of claims 1 to 5, wherein the cannula (506; 2304; 2402; 2502) of the sensing mechanism is a separate cannula from the cannula of the fluid delivery mechanism or is the same as the cannula of the fluid delivery mechanism.

7. The device of any one of claims 1 to 6, wherein the sensing mechanism uses one or more electrochemical cells.

8. The device of any one of claims 1 to 7, wherein the sensing mechanism comprises one or more electrochemical cells disposed on all or a portion of an outer surface of the cannula (506; 2304; 2402; 2502).

9. 9. The device of claim 1, wherein the cannula of the sensing mechanism is configured with a two-electrode sensor comprising a working electrode and a counter / reference electrode combination, a three-electrode sensor comprising a working electrode, a reference electrode and a counter electrode, or a three-electrode sensor comprising multiple working electrodes, a reference electrode and a counter electrode.

10. 10. The device of claim 1, wherein the cannula of the sensing mechanism is configured with a two-electrode sensor or a three-electrode sensor, the electrodes of the sensor being located inside the cannula or on an outer surface of the cannula.

11. 11. The device of any one of claims 7 to 10, wherein the electrochemical cell is in the form of one or more microneedle arrays, particularly hollow, open-tipped microneedle arrays, the open tips being in fluid communication with the one or more drug delivery reservoirs disposed within the housing.

12. The device of any one of claims 1 to 11, wherein the fluid delivery mechanism comprises one or more microneedle arrays disposed on a base of the housing of the device.

13. 13. The device of claim 11 or 12, wherein the fluid delivery mechanism comprises one or more transdermal patches disposed on a base of the housing of the device, in particular one or more transdermal patches comprising a hydrogel matrix containing a drug.

14. The device of any one of claims 1 to 13, further comprising: one or more amplifiers (1604; 1604-1; 1604-N) for amplifying signals from the one or more electrochemical cells; one or more multiplexers (4112; 4114) for multiplexing signals from the one or more electrochemical cells; and an analog-to-digital converter for converting analog signals from the one or more electrochemical cells to digital signals; and wherein software running on the controller analyzes the digital signals to determine the analyte level.

15. A device according to any preceding claim, wherein the sensing mechanism comprises one or more photofluorescent cells.

16. Each of the one or more photofluorescent cells comprises: a light source disposed within the housing of the device, the light source emitting excitation light at a first particular wavelength; one or more fluorescent elements that emit light at a second specific wavelength; 16. The device of claim 15, further comprising: a photodetector disposed within the housing of the device for detecting emitted light of the second particular wavelength.

17. 17. The device of claim 16, wherein the fluorescent element is inserted subcutaneously into the user, and further wherein the excitation light and the emission light are transmitted transcutaneously.

18. 18. The device of claim 16 or 17, wherein the one or more fluorescent elements are placed subcutaneously on the user and tethered to a cannula of the sensing mechanism.

19. The one or more fluorescent elements are disposed on an exterior surface of a cannula of the sensing mechanism, the cannula further comprising: a lumen used as the fluid delivery mechanism; one or more light guides coupled to the light source; The device of any one of claims 16 to 18, comprising one or more light guides coupled to the light detector.

20. one or more amplifiers for amplifying signals from the one or more photofluorescence cells; one or more multiplexing devices for multiplexing signals from the one or more photofluorescence cells; an analog-to-digital converter for converting an analog signal from the one or more photofluorescence cells into a digital signal; The device of any one of claims 15 to 19, wherein software running on the controller analyses the digital signal to determine the analyte level.

21. A device according to any preceding claim, wherein the sensing mechanism communicates with the device via wireless communication means.

22. A device according to any one of claims 1 to 21, wherein the sensing mechanism uses Raman shift spectroscopy.

23. The sensing mechanism includes an excitation light source that emits light at a first wavelength; and one or more Raman detectors for detecting Raman shifted light at wavelengths longer and shorter than the first wavelength.

24. 24. The device of claim 23, wherein the excitation light source and the one or more Raman detectors are disposed within the housing of the device.

25. 25. The device of claim 23 or 24, wherein the sensing mechanism cannula has an inner light guide for delivering the excitation light source and an outer light guide for collecting the Raman shifted light.

26. The device of any one of claims 23 to 25, further comprising a cone-shaped reflector at a tip of the cannula of the sensing mechanism for collecting the Raman shifted light.

27. The device of any one of claims 23 to 25, further comprising a hemispherical shaped reflector at the tip of the cannula of the sensing mechanism for collecting the Raman shifted light.

29. A device according to any one of claims 1 to 28, wherein the cannula of the sensing mechanism is also used as the fluid delivery mechanism.

30. one or more amplifiers for amplifying signals from the one or more Raman detectors; one or more multiplexing devices for multiplexing signals from the one or more Raman detectors; one or more analog-to-digital converters for converting analog signals from the one or more Raman detectors to digital signals; The device of any one of claims 23 to 29, wherein software running on the controller analyses the digital signal to determine the analyte level.

31. The device of any one of claims 1 to 30, wherein the one or more drug delivery reservoirs comprise a hydrogel matrix for storing drug.

32. 32. The device of claim 31 , wherein the drug delivery mechanism comprises one or more first electrical conductors disposed within the housing in electrical contact with the hydrogel matrix and one or more second electrical conductors disposed on the exterior of the microneedles proximate an open tip.

33. 33. The device of claim 32, wherein the controller is configured to apply a voltage to the one or more first conductors and the one or more second conductors to cause iontophoretic delivery of the drug to the user.

34. 34. The device of any one of claims 1 to 33, wherein the one or more drug delivery reservoirs are in the form of a coating disposed on an outer surface of the one or more microneedle arrays.

35. 35. The device of claim 34, wherein an outer surface of the one or more microneedle arrays is conductive, and the controller is configured to apply or remove a voltage to the one or more microneedle arrays to cause delivery of a drug to the user.

36. 36. The device of any one of claims 1 to 35, wherein the drug delivery mechanism further comprises one or more first electrical conductors disposed within the housing in electrical contact with a hydrogel matrix, and one or more second electrical conductors disposed on the device in contact with the skin of the user.

37. 37. The device of claim 36, wherein the controller is configured to apply a voltage to the one or more first conductors and the one or more second conductors to cause electrophoretic delivery of a drug to the user.

38. 37. The device of claim 35 or 36, wherein the drug delivery mechanism further comprises one or more ultrasound transducers in electrical contact with the hydrogel matrix.

39. 39. The device of any one of claims 35 to 38, wherein the sensing mechanism comprises one or more electrochemical cells comprising electrodes disposed on the microneedles of the one or more microneedle arrays.

40. 40. The device of any one of claims 35 to 39, wherein the sensing mechanism comprises one or more electrochemical cells comprising electrodes disposed in one or more sensing reservoirs.

41. 41. The device of any one of claims 1 to 40, wherein the sensing mechanism comprises one or more excitation light sources and one or more Raman detectors, the one or more excitation light sources and the one or more Raman detectors disposed within one or more sensing reservoirs.

42. The device of any one of claims 1 to 41, further comprising a component configured to establish an electromagnetic field for orienting analyte molecules.

43. A device according to any preceding claim, wherein the drug delivery mechanism comprises a pump for controlled delivery of a drug.

44. A device according to any preceding claim, wherein the sensing mechanism is designed to periodically detect the analyte level of the user.

45. 45. The device of any one of claims 1 to 44, wherein the one or more drug delivery reservoirs contain a liquid drug.

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