JP2015502785A - Precision flow control in drug pump device - Google Patents

Abstract

The accuracy of drug delivery using a drug pump device can be improved by a combination of pumping at high flow resistance and pump pressure, a pressure relief mechanism, and sensor-based feedback for pump control. The present invention, in various embodiments, provides an electrolysis-driven drug pump device with a control mechanism that facilitates accurate basal and / or bolus delivery. In some embodiments, the pump pressure generated in the electrolysis chamber is controlled by a feedback loop that adjusts the electrolysis rate based on the pump pressure and / or the liquid drug flow rate reading.

Description

(Citation of related application)
This application claims the priority and benefit of US Provisional Patent Application No. 61 / 561,565, filed Nov. 18, 2011, which is hereby incorporated by reference in its entirety. Incorporated inside.

  The present invention relates generally to drug pump devices, and in various embodiments, to electro-driven piston pump devices.

  The treatment of many diseases requires regular subcutaneous skin injections. For example, diabetic patients may require post-meal insulin injections, as well as a low “basal” insulin rate administered continuously. The main technologies currently used for frequent or continuous drug delivery are syringes, prefilled pen injectors, and portable drug pump devices that are filled by the patient. Each of these technologies has problems. For example, a syringe will easily trap air bubbles during the filling process unless it is filled by a well-trained and skilled individual (eg, a medical professional) and pose a risk to patient safety. In addition, certain treatments require injection volumes greater than 1 ml, for example, protein precipitates at high concentrations, so protein solutions often cannot be formulated at high concentrations, instead Requires large injections. However, multiple injection volumes generally make it impossible to use a syringe, as more than 1 ml can cause pain and swelling when injected at a high flow rate using a syringe or the like. A pre-filled pen syringe is advantageous in that it uses a pre-filled bubble-free glass cartridge to facilitate accurate manual insulin administration to simplify the priming process for the patient. However, poor patient compliance (eg, inability to follow improper injection timing and / or dosing regimens) is a major concern since injections are performed manually.

  The portable drug pump device can provide fully controlled drug delivery, thus greatly improving patient compliance. A reduction in the number of injections (eg, once every 3 days) and a programmable dosing schedule can greatly improve the patient's quality of life. In addition, many portable pump devices are provided in the form of patch pumps with a low profile pump profile that can be attached to a patient's skin without interfering with daily activities such as showering, sleeping, and exercising. . The controllable pump device further delivers a drug at a slower rate than syringes and pen injectors, thereby allowing a larger volume of fluid (eg, 1-10 ml) without causing discomfort or damage to the local tissue. Injection can be facilitated, which is particularly important for viscous drug solutions (eg, with viscosities of 35 cSt or greater). However, these pumps are typically filled by the patient, creating a danger during the priming procedure. Improperly prepared reservoirs contain large bubbles and can cause the pump to inject a lot of air into the subcutaneous tissue, which is a serious safety issue. Also, many portable drug pump devices, including commercially available insulin pumps, are driven by a stepper motor (or similar component for rotating gears). Stepping motors are known for their precise rotational pitch control, but their motion is discrete rather than continuous. Thus, the basal delivery flow provided by the stepper motor is also discrete. For example, a basal rate within the range of 5 to 5000 nl / min, a typical dosing schedule for insulin, is much higher with discrete 5 nl delivery at a rate between 1 delivery per hour and 1000 delivery per hour Achieved by the system. This discontinuous drug delivery is a major limitation of stepper motor driven insulin pumps.

  Recently developed electro-driven piston pump devices that utilize pre-filled glass (or polymer) vials in a pen injection configuration solve many of the problems associated with the prior art. They promote steady continuous drug flow at programmed rates and avoid patient compliance issues. In addition, the use of prefilled vials obviates the need for the patient to fill the drug reservoir, makes the pump easier to use, drug leakage due to improper filling, and particulates into the patient's subcutaneous tissue Or eliminate the risk of introducing foreign matter. However, providing controlled and accurate drug delivery remains a challenge for pumps that utilize glass vials as drug reservoirs. This challenge arises largely from the variable static friction / friction force between the glass vial and the piston or plunger that ejects the drug from the vial. The resulting unstable flow resistance makes it difficult to control drug flow. This tends to subject the basal drug delivery to varying flow rates, despite the constant driving pressure, and can make bolus delivery from one bolus to the next bolus unpredictable. Thus, there is a need for an improved flow control scheme and mechanism to ensure a constant and accurate and predictable drug dosage for both basal and bolus delivery.

  The present invention, in various embodiments, provides an electrolysis-driven drug pump device with a control mechanism that facilitates accurate basal and / or bolus delivery. In some embodiments, the pump pressure generated in the electrolysis chamber is controlled by a feedback loop that adjusts the electrolysis rate based on the pump pressure and / or the liquid drug flow rate reading. To reduce the impact of sudden changes in pump conditions (such as fluctuations in frictional force) on the delivery rate, the device has a flow resistance that dominates the overall flow resistance of the device downstream from the drug reservoir, A flow restrictor may be included, for example, a cannula, needle, or other outlet member with a small inner diameter (eg, less than 50 μm) may serve as a flow restrictor, or the flow restrictor may be, for example, May be a separate component located upstream of the outlet member. Further, in order to achieve the desired flow rate (eg, 400-5000 nl / min) despite the high flow resistance, the pump is operated at a high driving pressure (eg, greater than 5 psi and as great as 200 psi). May be. Advantageously, within the high pressure operating plan, the pump pressure and flow rate are directly proportional over a wide range, and the flow rate is accurately and precisely controlled based on the reading value of the pump pressure in conjunction with the knowledge of device flow resistance ( E.g., maintained at a certain constant value).

  For increased accuracy through flow rate control and increased safety through sensor redundancy, flow sensors may be used in combination with pressure sensors, and the flow rate calculated from the measured flow rate and the measured pressure Promote comparison between. In case of a difference between the two, a safety protocol may be initiated, for example, to stop pumping. Alternatively, in some embodiments, the measured flow rate is used for pump control as long as it is within the specified limits of the calculated flow rate; otherwise, the measured flow rate is considered incorrect. And the calculated flow velocity is substituted as a control parameter.

  To achieve bolus delivery, a flow sensor in the outflow cannula, needle, or other outlet member may be used to monitor the real-time flow rate of the drug delivered to the patient. This real-time flow rate signal can be integrated electronically to calculate the cumulatively delivered drug amount. When the volume reaches the target bolus volume, the control circuit in the pump shuts off the power to the pump and the electrolysis stops instantaneously. However, without further measures, the accumulated electrolytic gas in the pump chamber will continue to exert pressure to drive drug delivery. In order to relieve this pressure and prevent unnecessary additional flow, various pump embodiments include, for example, a pressure relief mechanism that causes recombination of the electrolytic gas. This mechanism may work continuously, disadvantageous to electrolysis during active pump operation, and accelerates the pressure drop once the pump is deactivated. Alternatively, the pressure relief mechanism can be controlled either simply to variably achieve full or partial gas recombination with the end pressure between on and off states, or within a target level or continuous range. It may be possible. For single bolus delivery without background basal delivery, the pressure relief mechanism is preferably operated to completely reduce the drive pressure and cause the pump to stop delivery when the pump is deactivated. Is done. On the other hand, for bolus delivery with background basal delivery, the control circuit preferably shuts off the pump power and reduces the pressure in the pump chamber until the driving pressure for background basal delivery is reached. Control the mitigation mechanism. In this delivery scenario, the flow rate drops rapidly and stays at the background basal rate.

  A “flow sensor” as used herein is any sensor that measures flow rate either directly or indirectly through another physical quantity that has a known relationship with the flow rate. is there. For example, the pressure sensor functions as a flow sensor in that the flow velocity can be calculated from the pressure reading value, assuming that the flow resistance is known (as will be explained in more detail below). May be. For clarity, a flow sensor that directly measures the flow rate is referred to herein as a “direct measurement flow sensor”.

Thus, in one aspect, the invention comprises a drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, a flow restrictor for limiting the fluid flow rate through the outlet member, and an intervening An electrolytic pump having a pump chamber in mechanical communication with a drug reservoir via a displacement member (e.g., a piston or diaphragm); and generating pressure in the pump chamber to move the displacement member toward the outlet member And a control circuit for operating the pump to drive and thereby force fluid from the reservoir through the outlet member. The flow resistor has a flow resistance coefficient (further defined below as flow resistance divided by fluid viscosity) of at least 10 ⁶ μl ⁻¹ . The electrolytic pump is operable and the control circuit is configured to operate the pump to exert a pressure of at least 2 psi, preferably at least 5 psi on the displacement member. Thus, the circuit and flow restrictor cooperate to cause continuous fluid flow through the outlet member at a constant flow rate in the range of about 400 nl / min to about 5 μl / min.

  In some embodiments, the pump is operable to exert a pressure of at least 10 psi, at least 50 psi, at least 100 psi, or at least 200 psi. The flow restrictor may have a minimum inner diameter that does not exceed 100 μm, or preferably does not exceed 50 μm. Its length may be in the range of about 1 cm to about 15 cm. In some embodiments, the outlet member includes a flow restrictor. In other embodiments, the flow restrictor is a separate component connected to the outlet member (which can be, for example, a needle or cannula). The flow restrictor coefficient of the flow restrictor may be within a range that provides a substantially linear relationship between pump pressure and fluid flow rate through the outlet member.

  The drug pump device may further include a pressure sensor disposed in the pump chamber to measure pressure therein and / or a flow sensor disposed in the outlet member. The circuit for operating the pump is based on a comparison of the measured pressure or flow rate with the target pressure or target flow rate to cause fluid flow at the target flow rate to cause fluid flow at a constant specific flow rate. It may be configured to adjust the electrolysis current supplied to the electrodes. In certain embodiments, the pump includes both a pressure and a flow sensor, and the circuit calculates a flow rate from the measured pressure, compares the calculated flow rate with the measured flow rate, and responds to the difference between the two. And configured to adjust the electrolysis current supplied to the electrode based on either the measured or calculated flow rate. For example, if the measured flow rate is within 5% of the calculated flow rate, the pump may adjust the electrolysis current based on the measured flow rate, otherwise, based on the calculated flow rate. Current adjustments may be made. This arbitration scheme serves to operate the pump to produce a pressure that causes continuous fluid flow through the outlet member at the target flow rate.

  In certain embodiments, the drug reservoir of the device is formed inside the vial and the displacement member includes a piston movably disposed within the vial. The electrolytic pump may include an electronics module that is mounted at the end of the vial and forms a pump chamber between the piston and the electronics module. The pump chamber may be sealed in the wall formed by the electronics module using an O-ring located on the periphery of the vial and in the circumferential recess of the electronics module. The electronics module may be removable and reusable in a separate drug pump device.

  In another aspect, the present invention provides a drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, and a drug reservoir via an intervening displacement member (eg, a piston or a diaphragm). An electrolytic pump, including a pump chamber in mechanical communication, a pressure sensor disposed within the pump chamber for measuring pressure therein, and an outlet member for measuring fluid flow rate therethrough It is directed to a drug pump device that includes a direct measurement flow sensor disposed and a control circuit. The electrolytic pump is operable to exert pressure to drive the displacement member toward the outlet member, thereby urging the fluid therethrough in the drug reservoir. The control circuit adjusts the electrolysis current supplied to the electrode for electrolysis based on the sensor reading, which calculates the flow rate from the measured pressure, compares the calculated flow rate with the measured flow rate, and calculates The electrolysis current supplied to the electrode is configured to be adjusted based on the measured flow rate if within 5% of the measured flow rate, otherwise based on the calculated flow rate. Thereby, the pump is operated to generate a pressure that causes continuous fluid flow through the outlet member at the target flow rate.

In a further aspect, a method is provided for controlling fluid flow through an outlet member of a drug pump device as described above (ie, including a drug reservoir, an outlet member, and an electrolytic pump). The method measures the pressure in the pump chamber of the electrolysis pump, calculates the flow rate from the pressure, measures the fluid flow rate through the outlet member, and compares the calculated flow rate with the measured flow rate. And adjusting the electrolysis current based on the measured flow rate if within 5% of the calculated flow rate, otherwise based on the calculated flow rate. Thereby, the pump is operated to generate a pressure that causes continuous fluid flow through the outlet member at the target flow rate.

  In yet another aspect, the present invention is in mechanical communication with the drug reservoir via a drug reservoir, an outlet member for fluidly connecting the reservoir to the drug injection site, and an intervening displacement member. An electrolytic pump having a pump chamber and operable to exert pressure and drive a displacement member toward the outlet member, thereby urging fluid therethrough within the drug reservoir. Intended for drug pump devices. The device further includes a pressure sensor for measuring pressure in the electrolysis chamber, a flow sensor for measuring fluid flow through the outlet member, and a plurality of consecutively obtained readings from the sensor over time. Pressure sensor and said flow sensor for manipulating the electrolysis electrode to achieve a target flow rate based at least in part on a plurality of consecutive values of the obtained reading values And a control circuit responsive to The time-aligned reading values may be stored in a circular buffer or may be put in and out of the buffer in a first-in first-out order. The reading value may be obtained at discrete time intervals or at the occurrence of a pump event, regardless of pump status. Such a pump event may be, for example, a pump actuation, and multiple time-separated readings may be obtained as the pump operates following each pump actuation.

  In a further aspect, the present invention provides a method of controlling fluid flow rate through an outlet member of a drug pump device that includes a drug reservoir and an electrolytic pump. The method includes: (i) storing a plurality of consecutively obtained readings of a flow rate through the outlet chamber of the electrolytic pump, and (ii) a flow rate through the outlet member in chronological order; and at least partially And operating the electrolytic pump to achieve the target flow rate based on a plurality of consecutive values of the obtained readings. The time-aligned reading values may be stored in a circular buffer or may be moved in and out of the circular buffer in a first-in first-out order. The reading value may be obtained at discrete time intervals or at the occurrence of a pump event, regardless of pump status. Such a pump event may be, for example, a pump actuation, and multiple time-separated readings may be obtained as the pump operates following each pump actuation.

  In a further aspect, the present invention provides a drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, and a pump chamber in mechanical communication with the drug reservoir via an intervening displacement member. An electrolytic pump, a pressure sensor disposed in the pump chamber for measuring pressure therein, and a direct measurement flow sensor disposed in the outlet member for measuring fluid flow rate therethrough. Including a drug pump device. The pump is operable to exert pressure to drive the displacement member toward the outlet member, thereby urging the fluid in the drug reservoir through the outlet member. The drug pump device further calculates a flow rate from the measured pressure, compares the calculated flow rate with the measured flow rate, and generates a pressure if the measured flow rate is within 5% of the calculated flow rate. Control circuitry for continuing fluid flow through the outlet member (by operating the pump to do so) or otherwise initiating a safety protocol. The safety protocol may include interrupting pump operation and, in some embodiments, resuming pump operation.

  In another aspect, the invention relates to a method of controlling fluid flow rate through an outlet member of a drug pump device comprising a drug reservoir and an electrolytic pump. The method includes measuring a pressure in the pump chamber of the electrolytic pump, measuring a fluid flow rate through the outlet member, calculating a flow rate from the measured pressure, and measuring the calculated flow rate. As compared to the flow rate, if the measured flow rate is within 5% of the calculated flow rate (or in other embodiments, within 3% or within 10%, depending on the required accuracy), a pressure is generated. Continuing fluid flow through the outlet member by operating the pump, otherwise performing a safety protocol (e.g., may involve interrupting pump operation and / or resuming pump operation); including.

  In yet another aspect, the present invention provides a drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, and downstream of the drug reservoir for limiting fluid flow through the outlet member. Having a flow restrictor and a pump chamber in mechanical communication with the drug reservoir via an intervening displacement member to exert pressure and drive the displacement member toward the outlet member, thereby providing a drug chamber An electrolysis pump operable to push fluid therethrough, a pressure sensor for measuring pump pressure in the pump chamber, and a reservoir outlet between the drug reservoir and the flow restrictor Pressure sensor for measuring pressure and monitoring the difference between pump pressure and reservoir outlet pressure, and based on that, operates the electrolytic pump to achieve a substantially constant fluid flow rate through the outlet member in order to And a control circuit, and drug pump devices.

  Another aspect of the present invention is to measure the pump pressure in the pump chamber, measure the reservoir outlet pressure between the drug reservoir and the flow restrictor, and between the pump pressure and the reservoir outlet pressure. Fluid flow rate through the outlet member of the pump as described above by operating the electrolytic pump to achieve a substantially constant fluid flow rate through the outlet member It is intended to control the method.

  In another aspect, the present invention provides a pump chamber and a pair of electrodes in contact with the electrolyte to generate an electrolytic gas from the liquid electrolyte upon application of a voltage between the electrodes, and the liquid contained in the chamber The present invention relates to a pump device with an electrolytic pump, including a pressure relief mechanism for causing recombination of the electrolytic gas into the electrolyte. The device further includes a control mechanism for variably controlling the pressure relief mechanism. In some embodiments, the device is a drug pump device, a drug reservoir in mechanical communication with the pump chamber via an intervening displacement member, and fluidly connecting the reservoir to the drug injection site. Gas generated in the pump chamber exerts pressure to drive the displacement member toward the outlet member, thereby urging the fluid therethrough in the drug reservoir. . The drug pump device further integrates the flow rate sensor to measure the fluid flow rate through the outlet member and the flow rate to determine the fluid dose, and interrupts the application of voltage between the electrodes when a specific dose is reached And / or a control circuit for triggering the pressure relief mechanism. The control circuit may be configured to variably control the pressure release mechanism to facilitate partial pressure relief.

  In some embodiments, the pressure relief mechanism includes a spark ignition circuit having a spark gap disposed within the pump chamber, wherein the creation of a spark in the spark gap causes recombination of at least a portion of the electrolytic gas. The spark ignition circuit may be configured to facilitate interrupting the spark before electrolytic gas recombination is complete. The pump chamber may include first and second compartments and a gas permeable separation therebetween (eg, a membrane, a porous material, a perforated solid wall, or a valve) and the spark gap is a first It may be placed in the compartment, thereby limiting the recombination of the electrolytic gas to the first compartment.

  In some embodiments, the pressure relief mechanism includes a conductive filament disposed within the pump chamber, and heating of the filament upon application of current thereto causes electrolytic gas recombination.

  In some embodiments, the pressure relief mechanism includes a catalyst (eg, including a material such as platinum, palladium, nickel, iridium, or nickel cadmium) that continuously causes recombination of the electrolytic gas. The catalyst may comprise a nanocatalytic material such as, for example, a nanoporous material, a nanowire, or a nanoparticle. The catalyst may cause recombination of at least 99% by volume of the electrolysis gas within only 2 minutes when the application of voltage between the electrodes is stopped.

  In some embodiments, the pressure relief mechanism includes an active pressure relief valve that fluidly connects the interior of the pump chamber with the exterior of the pump chamber. The valve may be or may include, for example, a solenoid valve, a diaphragm valve, a ball valve, or a duckbill valve. The liquid electrolyte may be absorbed into the hydrophilic material that is in contact with the electrode, thereby preventing release of the liquid electrolyte through the valve.

  In some embodiments, the pressure relief mechanism is a gas permeable, liquid impermeable material (eg, porous Teflon, porous sol-gel ceramic, or sintered porous metal) that forms part of the pump chamber enclosure. including. The enclosure portion formed of a gas permeable and liquid impermeable material has a size and / or porosity that corresponds to a predetermined permeability at operating pressure.

  In another aspect, the invention provides a drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, an electrolysis pump, and a flow sensor for measuring fluid flow through the outlet member. A drug pump device including a control circuit. The electrolytic pump includes a pump chamber that is in mechanical communication with the drug reservoir via an intervening displacement member and is in contact with the electrolyte to generate an electrolytic gas from the liquid electrolyte when a voltage is applied between the electrodes. The gas generated in the pump chamber exerts pressure to drive the displacement member toward the outlet member, thereby urging the fluid in the drug reservoir through the outlet member. The electrolytic pump further includes a pressure relief mechanism for causing recombination of the electrolytic gas into the liquid electrolyte. The control circuit is configured to integrate the flow rate to determine the fluid dose and to cease the application of voltage between the electrodes when a specific dose is reached. In some embodiments, the pressure relief mechanism is controllable and the control circuit is configured to trigger the pressure relief mechanism when a specific dose is reached. In some embodiments, the pressure relief mechanism is continuous.

  In yet another aspect, the present invention uses a drug pump device that includes a drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, an electrolytic pump, and a pressure relief mechanism. A method of controlling the dose of fluid delivered to an injection site is provided. The method measures the fluid flow rate through the outlet member, integrates the flow rate to determine the fluid dose, interrupts the generation of electrolytic gas and / or reduces pressure when a specific dose is reached. Triggering the mechanism.

The terms “substantially” and “about” as used herein generally mean within 10%, or preferably within 5%, unless explicitly or otherwise indicated by context. .

  The pump technology described herein includes, but is not limited to, insulin or vaccination for the treatment of diabetes, and other types of proteins for the treatment of obesity, Alzheimer's disease, and / or other diseases. It can be used for a variety of drug therapies, including delivery (eg, including monoclonal antibodies). Various embodiments and features of the invention are particularly advantageous for piston pump devices. In general, however, they are also applicable to other pump device configurations. For example, in some embodiments, a syringe pump is used that can administer the entire reservoir into the tissue within 30 minutes or that can measure the flow rate over several days.

The foregoing contents will be easily understood from the following modes for carrying out the present invention, particularly when interpreted in conjunction with the drawings.
FIG. 1 is a block diagram illustrating various functional components of an electrolytic drug pump device, according to various embodiments. FIG. 2A is a schematic side view of a piston pump device, according to various embodiments. FIG. 2B is a schematic isometric view of an electrolytic pump chamber, according to various embodiments. 3A-3C are an isometric view, a side view, and an exploded view, respectively, of a piston pump device, according to various embodiments. 3A-3C are an isometric view, a side view, and an exploded view, respectively, of a piston pump device, according to various embodiments. 3A-3C are an isometric view, a side view, and an exploded view, respectively, of a piston pump device, according to various embodiments. 4A-4D are schematic side views of assembly into an electrode / electronic / battery module and the pump device of FIGS. 3A and 3B, according to various embodiments. 4A-4D are schematic side views of assembly into an electrode / electronic / battery module and the pump device of FIGS. 3A and 3B, according to various embodiments. 4A-4D are schematic side views of assembly into an electrode / electronic / battery module and the pump device of FIGS. 3A and 3B, according to various embodiments. 4A-4D are schematic side views of assembly into an electrode / electronic / battery module and the pump device of FIGS. 3A and 3B, according to various embodiments. FIG. 5 is an exploded view of another piston pump device, according to various embodiments. 6A and 6B are isometric and exploded views of a pump device with an integrated lancet insertion mechanism, respectively, according to various embodiments. 7A-7C are side views illustrating different stages of lancet insertion using the device of FIGS. 6A and 6B. 7A-7C are side views illustrating different stages of lancet insertion using the device of FIGS. 6A and 6B. 7A-7C are side views illustrating different stages of lancet insertion using the device of FIGS. 6A and 6B. FIG. 8 is a schematic side view of a diaphragm pump device, according to various embodiments. 9A-9C are schematic side views of a flow restrictor, according to various embodiments. 10A and 10B are schematic diagrams illustrating high flow resistance and low flow resistance operating regimes of drug pump devices, according to various embodiments. 10A and 10B are schematic diagrams illustrating high flow resistance and low flow resistance operating regimes of drug pump devices, according to various embodiments. FIG. 11A is a schematic side view of a piston pump electrolysis chamber with a catalyst pressure relief mechanism according to various embodiments. FIG. 11B is a schematic side view of a piston pump electrolysis chamber with a semi-permeable housing, according to various embodiments. FIG. 11C is a schematic side view of a piston pump device with a spark ignition recombination mechanism, according to various embodiments. FIG. 11D is a circuit diagram of a spark ignition circuit that may be used in the embodiment of FIG. 11C. FIG. 11E is a schematic side view of a piston pump electrolysis chamber as depicted in FIG. 11C further including compartment separation, according to various embodiments. FIG. 11F is a schematic side view of a piston pump electrolysis chamber with an active valve pressure relief mechanism, according to various embodiments. 12A-12C are diagrams that illustrate basal, bolus, and basal / bolus drug delivery modes, according to various embodiments. FIG. 13A is a schematic side view of a piston pump device with pressure-based feedback for basal delivery, according to various embodiments. FIG. 13B is a control flow diagram of a pressure-based basal delivery feedback loop, according to various embodiments. FIG. 14A is a schematic side view of a piston pump device with flow-based feedback for basal delivery, according to various embodiments. FIG. 14B is a control flow diagram of a flow-based basal delivery feedback loop according to various embodiments. FIG. 15A is a schematic side view of a piston pump device with dual sensor feedback for basal delivery, according to various embodiments. FIG. 15B is a control flow diagram of basal delivery dual sensor feedback, according to various embodiments. FIG. 16A is a schematic side view of a piston pump device with dual sensor feedback and pressure relief mechanisms for basal / bolus delivery, according to various embodiments. FIG. 16B is a control flow diagram of basal / bolus delivery dual sensor feedback, according to various embodiments. (not listed) FIG. 17 is a schematic diagram illustrating a 24-hour period of basal and bolus delivery using a drug pump device, according to various embodiments. 18A-18D are schematic diagrams illustrating multiple bolus delivery protocols for low basal flow rates of 0.05 μl / hour, 0.1 μl / hour, 0.15 μl / hour, and 0.2 μl / hour, respectively.

  FIG. 1 illustrates in block diagram form the major functional components of a drug pump device 100 according to various embodiments of the invention. In general, the pump device 100 includes a drug reservoir 102 that interlocks with the electrolytic pump 104 via a displaceable member 106. The displaceable member 106 may be, for example, a piston, a diaphragm, a bladder, or a plunger. In use, the drug reservoir 102 is filled with a drug in liquid form, and the pressure generated by the pump 104 moves or expands the displaceable member 106 to push the liquid drug out of the reservoir 102. A cannula, needle, or other outlet member 108 connected to the outlet of the drug reservoir 102 conducts liquid to the infusion set 109. Infusion set 109 may include a catheter that is fluidly connected to cannula 108 to deliver the drug to the subcutaneous tissue region. A lancet and associated insertion mechanism may be used to drive the catheter through the skin. Alternatively, infusion set 109 may include another type of drug delivery vehicle, such as a sponge or other means that facilitates drug absorption through the skin surface.

これらの反応の正味の結果は、薬剤チャンバ内容物の全体的な体積膨張を引き起こす、酸素および水素ガスの産生である。ガス発生は、加圧環境で（報告によれば、約３０，０００ｐｓｉに対応する、最大２００ＭＰａの圧力で）さえも起こる。水の代替案として（または水に加えて）、エタノールが電解質として使用されてもよく、二酸化炭素および水素ガスの発生をもたらす。エタノール電解は、水電解と比較して、そのより優れた効率、結果として、より低い電力消費により、有利である。いくつかの実施形態による電解ポンプは、以下でさらに詳細に説明される。 Electrolytic pump 104 is generally an electrolyte-containing chamber (also referred to herein as a “pump chamber”) and a pair of direct-current power supplies disposed within the chamber to decompose the electrolyte into gaseous products. Including the above electrodes. Suitable electrolytes include water and aqueous solutions of salts, acids or alkalis, and non-aqueous ionic solutions. Water electrolysis is summarized by the following chemical reaction.

The net result of these reactions is the production of oxygen and hydrogen gas that causes the overall volume expansion of the drug chamber contents. Gas evolution occurs even in a pressurized environment (according to reports, at pressures up to 200 MPa, corresponding to about 30,000 psi). As an alternative to water (or in addition to water), ethanol may be used as the electrolyte, resulting in the generation of carbon dioxide and hydrogen gas. Ethanol electrolysis is advantageous due to its superior efficiency and consequently lower power consumption compared to water electrolysis. Electrolytic pumps according to some embodiments are described in further detail below.

  The pressure generated by the drug pump 104 may be adjusted via the pump driver 110 by a system controller 112 (eg, a microcontroller). The controller 112 may set the drive current and thereby control the rate of electrolysis and thus determine the pressure. Specifically, the amount of gas produced is proportional to the drive current integrated over time and can be calculated using Faraday's law of electrolysis. For example, creating two hydrogens and one oxygen molecule from water requires four electrons, so the amount of gas produced by electrolysis of water (measured in moles) is (three Multiply by a factor of 3/4 (since a molecule is generated for every four electrons) and equal to the total charge (ie current x time) divided by the Faraday constant.

  The system controller 112 may execute a drug delivery protocol programmed into the drug pump device 100, such as the pressure within the pump chamber 104, or the flow rate through the cannula 108 (or pressure within the cannula 108). It may be responsive to one or more sensors 113, 114 that measure operating parameters. For example, the controller 112 may adjust the current supplied to the electrolysis electrode based on the pressure inside the pump chamber to achieve the target pressure. In turn, the target pressure may be calculated based on the desired flow rate using a known relationship between flow rate and pressure (eg, as determined by calibration). Due to the low cost of pressure sensors (eg MEMS sensors as used in the automotive industry), this option is particularly advantageous for pumps that are designed for rapid drug delivery. Indeed, two or more pressure sensors 113 may be placed in the pump chamber to simultaneously monitor the pressure therein, and this redundancy provides the controller 112 with additional feedback and information. And serves as a backup in case of failure of one or more of the sensors. Alternatively, the rate of drug outflow from the reservoir 102 may be measured directly in real time using a flow sensor 114 that is incorporated into the outlet member 108 in a conventional manner. The total delivered dose can be calculated by integrating the flow rate over time and may serve as a control parameter for the electrolysis current.

  In some embodiments, the pressure sensor 113 inside the pump chamber is used in combination with the flow sensor 114 in the cannula to increase the accuracy and precision of the feedback control loop. The use of multiple sensors also ensures that if the flow sensor 114 fails, the pressure sensor 113 detects a high drug delivery rate and avoids overdose administration to the patient or damage to the pump device. Ensure that it will be possible to stop the operation. Conversely, the combination of flow and pressure sensors 114, 113 also allows drug storage when pressure is measured in the pump chamber, but no flow is measured in the cannula 108, indicating a potential leak. Violations in the unit 102 can also be detected. In general, the sensors used to measure various pump parameters may be flow rate, heat, time of flight, pressure, or other sensors known in the art and are biocompatible thin film polymers. It may be (at least partially) manufactured from parylene. The cannula 108 may also include a check valve 116 that prevents accidental drug delivery and backflow of liquid into the drug reservoir 112, and like the sensor 114, the check valve 116 is parylene. It may be produced. In other embodiments, silicon or glass is used in part in the construction of flow sensor 114 and valve 116.

The drug pump device 100 adjusts and further processes the sensor signal, optionally using an LED, other visual display, vibration signal, or audio signal to provide electronic status 118 (pump status information to the user). May be integrated with the system controller 112, but is not required). In addition to controlling the drug pump 104, the controller 112 may be used to control other components of the drug pump system, for example, triggering lancet and catheter insertion. The system controller 112 is an integrated circuit that includes a microcontroller, ie, in the form of a processor core, memory (eg, flash memory, read only memory (ROM), and / or random access memory (RAM)), and input / output ports. There may be. The memory may store firmware that directs the operation of the drug pump device. In addition, the device may include a read / write system memory 120. In one alternative embodiment, the system controller 112 is a general purpose microprocessor that communicates with the system memory 120. The system memory 120 (or memory that is part of the microcontroller) can be hard drive, flash drive, or other storage device at the time of manufacture or later, for example via USB, Ethernet, or FireWire port The drug delivery protocol may be stored in the form of instructions executable by the controller 112 that may be loaded into the memory by data transfer from. In an alternative embodiment, the system controller 112 comprises analog circuitry that is designed to perform the intended function.

  Pump driver 110, system controller 112, and electronic circuit 118 may be powered by battery 122 via suitable battery electronics. Suitable batteries 122 include non-rechargeable lithium batteries that approximate the size of the battery used in the watch, as well as rechargeable lithium ion, lithium polymer, thin film (eg, Li-PON), nickel metal hydride, and nickel cadmium batteries. including. Other devices for powering the drug pump device 100 such as a capacitor, solar cell, or motion generated energy system may be used either in place of the battery 122 or complementary to a smaller battery. Good. This can be useful if the patient needs to keep wearing the drug delivery device 100 for several days or more.

  In some embodiments, the drug pump device 100 can be controlled and / or remotely controlled by a wireless handheld device 150, such as a customized remote control or smartphone, as part of the electronic circuit 118 or as a separate component. Or a signal receiver 124 (for one-way telemetry) or a transmitter / receiver 124 (for two-way telemetry) that allows it to be reprogrammed. In certain embodiments, the handheld device 150 and the pump device 100 may utilize one or more inexpensive IR light emitting diodes and phototransistors as transmitters and receivers, respectively, (one or bi-directional) infrared (IR) links. Communicate on. Communication between drug pump device 100 and handheld device 150 may also occur at radio frequency (RF) using, for example, a copper coil antenna as transmitter / receiver component 124.

  The drug delivery device 100 may be manually activated, for example, switched on and off, using a switch incorporated in the pump housing. In some embodiments, using a toggle switch or another mechanical release mechanism, the patient can enclose the drug reservoir 102 enclosure to the needle so as to establish a fluid connection between the reservoir 102 and the cannula 108. (Eg, a drug vial septum as described below with respect to FIG. 2) may be punctured and then pump priming may begin. By coupling the insertion of the needle into the reservoir 102 with the activation of the pump device, the integrity of the reservoir 102 is ensured and thus protects the drug until it is injected, which is pre-filled drug Of particular importance to pump devices. Similarly, the lancet and catheter of infusion set 109 may be inserted by manually releasing the mechanical release mechanism. In some embodiments, insertion of the lancet and catheter automatically triggers electronic activation of the pump, for example by closing the electronic circuit. Alternatively, the pump and / or insertion set may be remotely activated by a wireless command.

  The functional components of a drug pump device as described above may be packaged and configured in various ways. In certain preferred embodiments, the drug pump device is incorporated into a patch that can adhere to the patient's skin. Suitable adhesive patches are generally manufactured from a flexible material that conforms to the contours of the patient's body and adheres via an adhesive on the back surface that contacts the patient's skin. The adhesive may be any material that is suitable and safe for application to and removal from human skin. Many versions of such adhesives are known in the art, but utilizing adhesives with gel-like properties can provide patients with particularly advantageous comfort and flexibility. The adhesive may be covered with a removable layer so as to make premature adhesion impossible before the intended application. Similar to commonly available bandages, the removable layer preferably does not reduce the adhesive properties of the adhesive when removed. In some embodiments, the drug pump device is of a shape and size suitable for implantation.

  Various components of the drug pump device may be held in a housing mounted on the skin patch. The device is completely self-contained or, when implemented as a discrete interconnect module, is entirely within the perimeter of the patch (ie, does not extend beyond any direction) Can be either present. The housing may provide mechanical integrity and protection of the components of the drug pump device 100 and may prevent disruption of the pump's operation from changes in the external environment (eg, pressure changes, etc.). The control system components 110, 112, 118, 120, 122 may be mounted on a circuit board that may be flexible and / or may be an integral part of the pump housing. In some embodiments, the control system component is incorporated into the self-contained unit with the electrode for electrolysis.

  The drug pump device 100 according to the present specification may be designed for single or repeated use. Multi-use pumps generally include a one-way check valve and a flow sensor, as described above, in the cannula. Further, the drug reservoir of the multi-use pump can be refillable via a refill port using, for example, a standard syringe. In some embodiments, the drug pump device 100 is removed from the patient's skin for refilling. The patient may, for example, place the drug pump device 100 and the cartridge containing the new drug in a home refill system, and the pump device and cartridge may be aligned using, for example, a press machine mechanism. Good. The patient may then press a button that triggers automatic insertion of a needle that draws liquid medication from the cartridge into the cannula to activate the electronics and begin priming the pump.

  The electrolysis pump 104 and drug reservoir 102 may be arranged in the device 100 in different ways, the two most common methods being that the pump chamber and reservoir are formed in an elongated vial and attached to the axis of the vial. A piston pump arrangement, separated by a piston movable along, and a reservoir pump arrangement, which is located above the pump chamber and separated therefrom by a flexible diaphragm. Both configurations are described in detail in US patent application Ser. No. 13 / 091,047 filed Apr. 20, 2011, which is hereby incorporated by reference in its entirety.

  FIG. 2A schematically illustrates an example piston pump device 200. Pump device 200 includes a cylindrical (or more generally tubular) vial 202 with a piston 204 movably positioned therein and an electrolysis electrode structure 206 mounted at one end. A septum 208 may be placed at the other end to seal the vial 202. Both piston 204 and diaphragm 208 may be made of an elastic polymer material such as synthetic or natural rubber. In some embodiments, silicone rubber (ie, polydiorganosiloxane, such as polydimethylsiloxane) is used. The piston 204 separates the inside of the vial 202 into a medicine reservoir 210 and a pump chamber 212. In use, the needle 214 punctures the septum 208 to allow fluid entry from the drug reservoir 210, and a cannula (not shown) connected to the needle 214 infuses fluid (not shown). May be conducted. The piston pump device 200 is enclosed in a protective housing 216 made of, for example, hard plastic.

  The electrode 206 may be made of any suitable metal such as, for example, platinum, titanium, gold, or copper, and may form a pair of parallel wires or plates. Alternatively, the electrode can have a non-conventional shape to improve electrolysis efficiency. For example, they may interpenetrate each other or be individually wound into a helical configuration (and oriented to face each other) as illustrated in FIG. 2B. Further, as shown, the electrode 206 may be incorporated into a hydrophilic absorbent material 218 (eg, a cotton ball) that ensures continuous contact with the electrolyte 220. This solves the problem often encountered with conventional electrolytic pumps where the electrode is simply immersed in a liquid electrolyte, and as the gaseous electrolysis product is produced, the piston is moved toward the outlet end of the drug reservoir. Pushing, thereby increasing the volume of the electrolysis chamber and causing a decrease in the level of electrolyte. Depending on the orientation of the device, one or both electrodes may eventually emerge from the electrolyte and be surrounded by gas, eventually forming an open circuit, thus stopping the electrolysis reaction. This problem can be avoided in a variety of ways, one of which involves (but is not limited to) surrounding the electrode with a hydrophilic absorbent material such as a hydrogel, cotton ball, sponge, or superabsorbent polymer. is there. The electrolyte stays inside the hydrophilic absorbent material, effectively discharging the generated gas and keeping the electrode in a state where the electrolyte is replenished.

  Vials 202 may be manufactured from glass, polymers, or other materials that are inert with respect to drug stability and are preferably biocompatible. For example, polymer vials made of polypropylene or parylene are suitable for certain drugs that degrade faster when in contact with glass, such as protein drugs. For many other drugs, glass is the preferred material. Glass is commonly used in commercially available FDA approved drug vials and containers from many different manufacturers. As a result, the drug in the glass container can be aseptically filled, which can accelerate the approval process of the drug pump device, protecting the drug in the glass container and avoiding the need to rebuild expensive aseptic filling production lines There are established and approved procedures for storage. Using glass for the reservoir further allows the drug to come into contact with similar materials during delivery. A suitable glass material for the vial may be selected based on chemical resistance and stability, and the anti-scattering properties of the material. For example, type II or type III soda lime glass, or type I borosilicate material may be used to reduce the risk of container breakage.

  To enhance chemical resistance and maintain the stability of the encapsulated drug formulation, the inner surface of the vial may have a special coating. Examples of such coatings include chemically bonded invisible ultrathin layers of silicon dioxide or medical grade silicone emulsions. In addition to protecting the chemical integrity of the encapsulated drug, a coating such as a silicone emulsion can provide a lower and more uniform friction between the piston and the vial.

  In certain embodiments, the piston pump device 200 is manufactured by fitting a conventional commercially available glass or polymer drug vial, which can already be proven for aseptic filling, along with the piston 204 and electrolytic pump components. A threaded needle cassette may be placed over the diaphragm 208 so that when the patient desires to use the pump, the cassette needle 214 pierces the diaphragm 208 to establish a connection with the cannula. A mechanical actuation mechanism may serve to screw the cassette into the vial 202. To accommodate the electrolysis pump, the vial 202 is longer in some embodiments than conventional commercially available vials, so that proven filling methods and existing aseptic filling line parameters do not need to be changed. And maintain all other properties. The drug pump device may comprise a prefilled vial. When glass vials are used, the drug can be stored in the pump device for a long shelf life without having to change the labeling on the drug.

  3A-3C show a typical piston pump design. In the illustrated embodiment, the pump case 300 integrated with the skin patch 302, ie, attached to the patch 302, eg, with an adhesive, or manufactured with the patch as a single integrated structure is It is adapted to axially receive a pre-filled drug cartridge 304 through 300 back openings. A ready-made lancet set 306 can be mounted at the front end of the case 300 and used to establish a fluid connection between the cartridge 304 and the injection catheter. Electrolysis electrodes, electronics, and batteries are incorporated into a single module 308 that is encapsulated in a plastic housing 310 with electrodes protruding at one end to simplify assembly. A rear screw 312 may be used to secure the electrode / electronic / battery module 308 inside the pump case, with the module housing 310 adjacent to the end of the glass cartridge 304. Use a rubber O-ring 314 to provide a seal between the end of the glass cartridge 304 and the electrode / electronic / battery module 308 to ensure the integrity of the electrolysis chamber formed therebetween. Can do.

  As discussed further below, various pump devices according to the present specification operate at higher pump pressures (eg, pressures above 5 psi, 10 psi, 20 psi, 50 psi, 100 psi, or 200 psi) than conventional drug pump devices. Designed to do. Thus, the case and pump connections for these devices are configured to withstand such high internal pressures without causing any leakage. The main components of the high pressure pump case design according to various embodiments are the electrode / electronic device / battery housing 310. As illustrated in FIGS. 4A and 4B, the housing 310 seals against an O-ring 314 between the drug vial 304 and the housing 310. The bottom wall 320 of the housing 310 may include a circumferential recess 322 in which the O-ring 314 is disposed so that the recess 322 is O when the housing 310 is engaged with the drug vial 304. The ring 314 is sized so that it sits securely on the periphery 324 of the drug vial 304. In this configuration, the O-ring 314 is prevented from sliding inward, i.e., into the interior of the pump chamber, while the rear screw 312 is tightened forward in the assembly process. As a result, a much higher pressure can be generated in the pump chamber without causing leakage. The housing 310 may be secured to the vial 304 via any of a variety of conventional engagement mechanisms. For example, the housing 310 may extend over the vial 304 and may be connected to it around the inner surface by an adhesive, or the vial 304 may have a threaded collar that is threaded or adhered thereto. May be.

  4C and 4D illustrate how the compression nut 404 can be used to secure the vial 304 and the electrode module 400 within the pump case 402 for sealing. Although the electrode module 400 is simply illustrated herein as an electrode structure on a substrate, embodiments instead include an electrode / electronic / battery module 308 as shown in FIGS. 3A-3C. It can be easily corrected. As shown in FIGS. 4A and 4B, the pump chamber 405 is formed at the rear end of the vial 304 between the piston 204 and the electrode module 400 and is disposed between the peripheral edge of the vial 304 and the electrode module 400. Sealed by O-ring 314 (or other gasket). The pump case 402 is sized to fully accommodate the vial 304 and electrode module 400 therein and extend beyond the assembly at the rear end. The overhanging portion 406 of the pump case 402 is threaded into its inner surface, allowing a compression nut 404 to be screwed into the pump case 402 so as to apply a compressive force to the O-ring to seal the pump chamber 405. To.

  In many wearable pump designs, batteries and electronics (which can optionally include a wireless module such as Bluetooth® or ZigBee) are expensive to produce (depending on the materials used). Although representing components that are dangerous to the environment and / or, these drawbacks are improved if the batteries and electronics can be recharged or reprogrammed and thus reused. In certain embodiments, the battery (or batteries) and electronics of the drug pump device can be reused multiple times for multiple infusion pumps, such as a battery and electronics module (BEM) (eg, module 308). Can be packaged. In a preferred embodiment, the BEM can be inserted or snapped into the pump (using an electrical connection) by the patient prior to use and removed at the end of the infusion. In addition to making the use of drug pump devices more economical, this reuse is a material that destroys the environment in landfills (eg, batteries made of nickel metal hydride, liPON, lithium ion, or lithium polymer). Delay or avoid discarding Removable BEM also facilitates sterilizing the pump by methods such as ion beam or gamma irradiation that damage conventional electronic equipment and adding additional electronic equipment after sterilization. In some embodiments, a charging / reprogramming station with a similar insertion or snap-in mechanism is provided to charge / reprogram the BEM between uses. As an alternative to an insertion mechanism in the charging station, the BEM can have additional telemetry coils for inductive charging during injection or for wireless control. Standard micro or normal USB connections may be utilized and water-resistant or waterproof standard packaging can be employed to protect the BEM.

  FIG. 5 illustrates another design of the high pressure pump device. Here, the drug vial 500, the pump assembly 502 (including the electrode structure and absorbent material), and the O-ring 503 are enclosed in a separate two-part housing 504, and the electronics / battery module 506 is attached to the side of the enclosed vial 500. Placed in. The outer case includes a bottom and top portion 508, 510 that is closed around the pump housing 504, electronics / battery module 506, and associated tube 512.

  6A and 6B depict an exemplary piston pump device 600 with an integrated lancet insertion assembly. The insertion assembly includes a shielding housing 602, a needle carrier 604, and a needle / catheter dual spring insertion mechanism 606 and is positioned above a pre-filled cartridge pump 608 (including a drug reservoir and an electrolysis pump). , Fluidly connected to it via a tube 610. The carrier 612 provides a base for the cartridge pump 608 and a connector 614 for the insertion assembly. The shielding housing 602 holds a needle 616 and a catheter 618 connected thereto (eg, made of Teflon), and two springs 620 (for needle and catheter insertion), 622 (for subsequent retraction of the needle). To the catheter hub 624. The cartridge pump 608 may be contained in a pump casing 626 that is enclosed in the outer device shell 630 along with the insertion assembly.

  7A-7C illustrate a mechanism for inserting a catheter into subcutaneous tissue. In the initial position, needle 616 and catheter 618 are located above catheter hub 624. Activation of the trigger button 700 releases the initially compressed insertion spring 620. This causes needle 616, needle carrier 604, and catheter 618 (hereinafter “needle carrier assembly”) to move downward, with needle 616 inserted into the subcutaneous tissue with catheter 618 passing through self-sealing silicone plug 704. (FIG. 7B). Self-sealing silicone plug 704 has two diaphragms (top and bottom layers) that provide an open area between the two layers in which the outlet of fluid tube 610 is in fluid communication. During insertion, the needle carrier assembly is propelled downward by a spring 620 and is stopped when the front (ie, downward) surface of the needle carrier 604 encounters the rear (upward) surface of the catheter hub 624. . The catheter hub 624 may have angled sides that act as a latch to hold the retracting spring 706 in place (still compressed). The retraction spring 622 is at least as rigid as the insertion spring 702, and is typically more rigid than that, and therefore, when released, compresses the insertion spring 702 and returns the needle carrier to its original position. The assembly can be driven back. When the user compresses the side of the catheter hub 624 with the thumb and index finger, the retraction spring 622 is released. As a result, as the needle carrier assembly is driven back into the retracted position, the needle 616 is extracted out of the tissue (FIG. 7C). When the needle is retracted, radial and axial compression on the silicone plug 704 immediately closes the small perforations and provides a tight seal to the fluid path in the infusion set. Following catheter insertion, the lancet insertion assembly and outer shell may be removed leaving only the pump and infusion set on the skin. An alternative catheter insertion mechanism is described in US Provisional Application No. 61 / 704,974.

  In the diaphragm pump 800 schematically illustrated in FIG. 8, the drug reservoir 802 and pump chamber 804 are stacked in a dual chamber configuration where the drug reservoir 802 is separated from the pump chamber 804 by a flexible diaphragm 806. . Typically, the pump chamber 804 is formed between the bottom portion of the housing 808 (which can be attached to the skin patch) and the diaphragm 806, and the drug reservoir 802 is disposed above the pump chamber 804, and the diaphragm It is formed between 806 and the lid-shaped part of the housing 808. The electrolytic gas generated in the pump chamber 804 results in pressure on the expanding diaphragm 806 and releases the liquid drug through the reservoir outlet 810 and into the cannula 812 (or other outlet member). Cannula 812 may be equipped with check valve 814 and flow sensor 816. The control circuit and the battery (not shown) may be mounted on a circuit board incorporated in the bottom portion of the housing 808. In some embodiments, the electrode 818 is etched, printed, or otherwise deposited directly on the circuit board for cost savings and ease of manufacture.

For example, in order to facilitate accurate basal and / or bolus drug delivery using a drug pump device such as that illustrated in FIGS. 2B and 8, various embodiments of the present invention include (1) high pump pressure and Utilizes a combination of flow restriction downstream of the combined reservoir, (2) a pressure relief mechanism that accelerates the pressure drop when electrolysis is stopped, and (3) sensor-based feedback for electrolysis pump control. The following sections describe each of these features in more detail.
(Flow limiter)

  A flow restrictor may be utilized downstream of the reservoir to stabilize the delivery flow rate regardless of friction and / or other pump condition variations. In some embodiments, the outlet member 108 itself serves as a flow restrictor. For example, a cannula or needle that conducts fluid from a drug reservoir to an injection site or a portion thereof may have a small inner diameter, such as, for example, a diameter of less than 100 μm, less than 50 μm, or less than 25 μm. Illustrated in FIG. 9A. Alternatively, a separate flow restrictor element may be integrated with the outlet member, as shown in FIG. 9B, which is a conventional needle whose inner diameter is typically in the range of 100 μm to 500 μm. And allows flow restriction despite the use of a cannula. The flow restrictor may simply constitute part 910 in a fluid path having a small inner diameter, or a valve or similar structure that can variably limit the flow rate (an iris-like valve with a controllable inner diameter). Etc.). In yet another embodiment, shown in FIG. 9C, the flow restrictor 920 is connected between a reservoir outlet 922 and an adapter 924 to which the outlet member is fluidly coupled during use. The flow restrictor 920 and adapter 924 may be formed inside the vial as an extension of and integral with the inner wall of the vial, and the space surrounding the restrictor and adapter configuration inside the vial is mechanically stable. For example, it may be filled with epoxy. In general, any configuration and arrangement in the fluid path between the reservoir and the injection site that intentionally increases downstream flow resistance may serve as a flow restrictor according to the present description.

  In certain embodiments, the flow restrictor is implemented by a micromachined microchannel device. Well-defined microchannels can be produced by either surface or bulk micromachining techniques, as is well known to those skilled in the art. Microchannel depth, width, and length can be machined with high precision, and tolerances can be controlled to nanometers. Furthermore, visual inspection using standard industrial fully automated microscopy systems can fully investigate manufactured microchannel flow restrictors without causing a significant increase in manufacturing costs. Micromachining processes are also very suitable for mass production. Compared to microcapillary flow restrictors made by traditional high-precision protruding processes that require 100% manual flow rate / flow resistance calibration, screening, and quality control, micromachined microchannel devices are manufactured and Cost savings can be provided both in quality inspection.

  The flow restrictor is preferably sized to dominate the overall flow resistance of the drug pump device. As a result, fluctuations in flow resistance imparted by other components have a significantly reduced effect on flow rate. For illustrative purposes, for example, assume that the flow resistance of a conventional piston pump device is due to the outlet member and vial / piston friction in equal amounts. Thus, a sudden drop in friction between the vial and the piston to half of its previous value changes the overall flow resistance by 25%. However, if the outlet member flow resistance is increased by a factor of 10, the same drop in friction results in a flow resistance change of only about 4.5%. Thus, it can be seen that the intentionally introduced high flow resistance reduces the relative effects of any variation in flow resistance of other device components, thereby smoothing and stabilizing the flow rate. Of course, without any equivalent change in drive pressure, increased flow resistance will result in lower flow rates. Because the desired flow rate in a drug pump device is usually determined by medical considerations, devices with high flow resistance embodiments are generally driven at high drive pressures. The pump case and connections are therefore designed to withstand much higher internal pressures without causing any leakage.

  FIG. 10A depicts a typical flow regime of insulin delivery as achieved using a high resistance outlet (eg, an outlet member having an inner diameter of less than 100 μm). The desired basal flow rate is in the range of about 8.3 nl / min to about 5 μl / min. By properly calibrating the flow resistance of the outlet tube by properly selecting its dimensions (ie, length and inner diameter), a flow rate of 400 nl / min can be obtained under a driving pressure of 20 psi. Furthermore, due to the fixed flow resistance at the outlet, the flow rate increases linearly with the driving pressure. Thus, by increasing the drive pressure by 12.5 times to 250 psi, the delivery flow rate can also be increased by 12.5 times to 5000 nl / min. Within this delivery window (400-5000 nl / min), delivery is truly continuous (not as discrete as a commercially available stepper motor pump). Lower flow rates (8.3-400 nl / min) can be achieved via multiple bolus delivery, as further described below. The same flow rate as shown in FIG. 10A is also at a lower pressure when a correspondingly lower resistance is used at the outlet (eg, when the outlet member has an inner diameter substantially greater than 100 μm). FIG. 10B illustrates an embodiment where a flow rate between 400 nl / min and 5 μl / min is generated by a driving pressure ranging from 2 psi to 20 psi, for example. In general, however, the lower the pressure / flow resistance combination, the stronger the effect of any variation in resistance imparted by the piston / vial subsystem. Thus, to ensure high flow rate stability, it is of course desirable to operate at higher flow resistance and pressure if the pump device can withstand such pressure without breakage or leakage. .

For different drug therapies, a different flow regime may be required than that shown in FIGS. 10A and 10. The flow resistance and / or drive pressure can be easily adjusted to achieve these different flow rates while stabilizing the flow rates according to the present description. Further, for different vials and pump systems, different pressure and flow resistance levels may be preferred for any given flow rate and can be determined by one skilled in the art without undue experimentation.

式中、μは、流体の動的粘度であり、

は、流体から独立して流量制限器によって提供される流動抵抗を特徴付けるように、この目的で定義される、流動抵抗係数である。 The relationship between drive pressure and delivery rate is determined by:
P = Q × R
Where P is the driving pressure (excluding any back pressure at the injection site, which is generally negligible), Q is the delivery rate, and R is the flow resistance of the device. As contemplated herein, in embodiments with a high resistance flow restrictor downstream of the reservoir, the contribution of the piston and vial (and any lower resistance portion of the fluid path downstream of the reservoir) Is negligible and R is essentially the flow resistance of the flow restrictor. For a tubular flow restrictor with length l and inner diameter D, the flow resistance can be expressed by the relationship:

Where μ is the dynamic viscosity of the fluid,

Is the flow resistance coefficient defined for this purpose to characterize the flow resistance provided by the flow restrictor independent of the fluid.

In various embodiments, the flow resistance is at least 1 psi / (μl / min), at least 2 psi / (μl / min), at least 4 psi / (μl / min), or at least 10 psi / (μl / min), In embodiments, for example, as high as 50 psi / (μl / min). For normal drug fluid viscosities in the range of about 1 cP to about 35 cP (1 cP = 1 mPa · s is the approximate viscosity of water at room temperature), a length with a diameter of less than 100 μm, preferably less than 50 μm Such a high flow resistance can be achieved using a flow restrictor of 1-15 cm, resulting in a flow resistance coefficient exceeding 10 ⁶ / μl and as high as 2 · 10 ⁹ / μl. For example, a flow restrictor that is 10 cm long and has a diameter of 50 μm results in a flow resistance coefficient of about 6.5 · 10 ⁸ / μl.
(Pressure reduction mechanism)

In order to facilitate accurate bolus delivery, it is important that the pump device be able to block the drug flow as fast as possible. However, while electrolysis can be interrupted almost instantaneously, as the gas constituents (eg, hydrogen and oxygen) gradually recombine with the liquid electrolyte (eg, water), the accumulated electrolytic gas pressure in the pump chamber becomes , Drops over a much longer period. Accordingly, it is desirable to provide an efficient pressure relief mechanism that reduces drive pressure and thus helps to more quickly block drug flow. In preferred embodiments, the pressure can be reduced to substantially zero within 1-2 minutes or less, but excessively rapid pressure relief can cause safety concerns. A suitable mechanism relieves pressure reproducibly, safely, and preferably controllable to allow reliable bolus delivery.

  Generally, the pressure is physically, chemically, mechanically, electrically, electrochemically, or thermally by recombining the electrolytic gas, removing the gas from the pump chamber, or changing the type of gas. Any device or process that can reduce the pressure can be used for pressure relief. Furthermore, with respect to their timing and duration, pressure relief mechanisms generally fall into two categories. The first type of mechanism causes continuous recombination and / or release of electrolytic gas from the pump chamber, and power is supplied to the electrolysis electrode during the period of active pumping operation, i.e., to produce electrolytic gas. When working, it works against the electrolytic pump. In order to achieve the desired pump pressure in this case, the electrolysis rate, and thus the electrolysis current, needs to be greater than without the pressure relief mechanism. Once the electrolysis is stopped, the mechanism operates to quickly reduce the amount of gas in the pump chamber, in other words, accelerates the pressure drop when the gas production is interrupted. The second type of mechanism is actively triggered and operates either for the duration inherent in the mechanism or until interrupted by the control mechanism. Many such active pressure relief mechanisms can be controlled to only partially reduce the pressure in the pump chamber, rather than reducing the pressure to zero. For example, they may be manipulated to achieve a desired end pressure that is selected within a continuous pressure range or between several discrete end pressure levels. A variably controllable pressure relief mechanism facilitates delivery of bolus injections (eg, 10 μl each) in combination with a background basal rate (eg, 500 nl / min).

  An example of a continuously operating pressure relief mechanism is the use of a recombination catalyst in the pump chamber, illustrated in FIG. 11A. Suitable catalyst materials for electrolytic gas recombination include (but are not limited to) metals such as platinum, palladium, nickel, and iridium, and metal alloys such as nickel cadmium, all of which contain gaseous hydrogen and Reduce the activation energy for the formation of liquid water from oxygen. This phase from gas phase hydrogen and oxygen to liquid water when the power to the electrolysis electrode is turned off so that the recombination of hydrogen and oxygen to water is not offset or exceeded by the reverse reaction. The change is accompanied by a significant decrease in volume, a reduction of about 1000 times, and a corresponding significant drop in pump pressure. The catalyst material can increase the rate of recombination by about an order of magnitude or more compared to the base rate of recombination that occurs in the absence of any catalyst or other acceleration mechanism.

  Nanocatalytic materials such as nanoporous materials, nanowires, and nanoparticles provide significantly improved performance compared to conventional scale catalysts. Due to the high surface-to-volume ratio of the nanostructure, recombination velocities of 2 to 3 orders of magnitude (compared to the reference speed) can be obtained. Examples of suitable nanomaterials include (but are not limited to) platinum black, platinum nanowires or nanoparticles, palladium nanowires or nanoparticles, and iridium nanowires or nanoparticles. As shown in FIG. 11A, these nanoparticles 1100 can simply be placed in an electrolysis chamber during the manufacturing and assembly process where they function as is to recombine hydrogen and oxygen constantly. it can. To inject a bolus of drug, power is applied to produce gas at a rate higher than the recombination rate until the desired bolus amount is reached, at which time the power is turned off and the nanoparticles are pressured Recombines the gas rapidly to reduce drug delivery and stop drug delivery.

  Pressure relief based on continuous removal of electrolytic gas from the pump chamber (at a lower rate than that gas is produced during the active pumping cycle) is permeable to gas, as shown in FIG. 11B. It can be achieved with a pump chamber casing that is but impermeable to liquids. Suitable materials include, for example, porous Teflon, porous sol-gel ceramics, and sintered porous metals such as stainless steel, aluminum, and titanium. The gas permeable material need not form the entire portion of the casing around the pump chamber, but may be limited to the lower portion in contact with the interior of the chamber that allows gas to leak out of the chamber. Good. In some embodiments, the inlet of the tube through which the chamber is first filled with liquid electrolyte may be closed, for example, with a gas permeable membrane. In general, the gas permeability through a gas permeable casing part or membrane is determined carefully by the dimensions of the gas permeable part (eg, its thickness and surface area) and the porosity of the material (eg, pore density and size). Can be preset by selection. Through permeability, the rate of pressure decrease after power interruption can be controlled, allowing the device to be manufactured to achieve a specific given bolus amount for a given electrolysis current.

  An example of a controllable pressure relief mechanism is the creation of a discharge spark in the pump chamber that induces a rapid gas recombination ignition process. Spark ignition can be achieved using any of a variety of suitable systems and processes such as, for example, capacitive discharge ignition, inductive discharge ignition, or transistor discharge ignition. For example, as shown in FIG. 11C, a discharge arc can be created simply by applying a high voltage across the gap between the two wires 1110 of the spark plug 1122 disposed in the chamber 1124. it can. Like a chemical catalyst (as described above with respect to FIG. 11A), the spark reduces the activation energy between gas phase hydrogen and oxygen so as to form liquid phase water, and the gas is virtually instantaneous. Recombine. The phase change from gas phase hydrogen and oxygen to liquid phase water can dramatically reduce the volume of the material (eg, about 1000 times), and this rapid volume reduction provides pressure relief. Recombination induced by spark ignition is very fast, usually resulting in almost complete pressure relief (eg, a drop of 1% of the original pressure) in the microsecond to millisecond range.

  Unlike spark ignition in combustion engines, which causes gas expansion, spark ignition, which induces gas recombination, causes volume reduction and consequently no risk of explosion. In addition, only minimal heat is produced during the process and likewise does not pose any safety hazard. However, under certain conditions, a very fast pressure drop induces shock waves inside the pump chamber, potentially damaging certain sensitive components installed in or around the chamber, such as pressure sensors and circuits. Can do. In order to avoid such problems, it may therefore be desirable to reduce the rate of spark ignition recombination. A controllable and adjustable pressure drop is also advantageous for implementing a drug delivery protocol where the pressure during bolus delivery is greater than or equal to zero, ie, a protocol that includes a background basal rate. For bolus basal delivery, the drive pressure is preferably reduced from high pressure to zero and then reduced from high bolus pressure to low basal pressure in a controlled manner, rather than returning to low pressure, Power consumption can be reduced.

  One way to achieve controlled slow recombination is to reduce the spark ignition time. As illustrated in FIG. 11D, a high speed circuit 1115 can be used to quickly turn the spark on and off. The circuit 1115 is basically a spark plug circuit including a direct current source 1117, an instantaneous on / off switch 1119, and a high voltage transformer 1121. In operation, the circuit 1115 operates in the manner of an automobile ignition circuit, with current flowing from the DC current source 1117 through the primary coil winding of the transformer 1121 and the current is interrupted by opening the switch 1119. The primary transformer coil's magnetic field rapidly collapses. The secondary coil induces current in the transformer coil, which is a very high voltage current in the secondary coil, because the number of windings in it is much larger than the number of windings in the primary coil Surrounded by a powerful and changing magnetic field. This voltage causes a breakdown and causes the current to flow in the form of a spark across the spark gap. Therefore, by blocking the spark, the recombination can be intentionally stopped before all the hydrogen and oxygen gases recombine.

  Another way to slow spark spark recombination is to use a separator 1130 that divides the interior of the pump chamber into two compartments 1132, 1134, as shown in FIG. 11E. Separator 1130 can be, for example, a valve, a membrane, a porous material, and / or a solid material with holes. Only the gas in the compartment 1132 that includes the spark gap 1110 will recombine and reduce the compartment pressure to zero (or nearly zero). The gas mixture in the other compartment 1134 adjacent to the piston gradually disperses through the compartment separation (on a time scale much longer than the duration of the spark) and replenishes the first compartment 1132 until pressure equilibrium is reached. Will do. The end pressure can be set via the volume ratio between the two compartments 1132, 1134. For example, the compartment 1132 containing the spark gap will occupy a quarter of the total pump chamber volume and the pressure will drop to about a quarter of its original value. Thus, repetitive spark ignition and pressure balance can be used to relieve pressure incrementally (eg, by a factor of 4 in the example). Of course, a combination of spark timing and compartment separation can also be used to optimize recombination control.

  An alternative approach to controllable pressure relief involves the use of electrical resistance filaments that are heated by the application of current. This mechanism is similar to that used in incandescent lamps, where the filament acts as an electrical resistor, and when sufficient power is applied, the filament temperature rises to several thousand degrees Celsius. The thermal energy of the filament initiates recombination of hydrogen and oxygen gas into the water, thereby reducing the chamber pressure. This mechanism provides a high degree of control because the rate and duration of recombination can be easily adjusted through the magnitude and timing of the current applied to the filament.

  Pressure relief can also be achieved by controllably releasing gas from the pump chamber using an active release valve, as shown in FIG. 11F. This valve 1140 can normally be closed (or have a small controllable leak rate that does not prevent pressure buildup in the chamber) and removes gas from the chamber 1144, thereby reducing pressure. It may be opened as needed to alleviate. Typically, the valve of the pump activates the valve as soon as the correct amount of drug is delivered (eg, as indicated by the reading value from the flow sensor 1146 in the outlet member). Electromechanical or piezoelectric so that it can be controlled by the control circuit. By closing valve 1140 before all gas leaks, the end pressure can be controlled, i.e., does not drain the drug, but reduces the pressure build-up required for the next cycle of drug delivery. Can be held in. The valve 1140 may also facilitate control of the gas release rate by providing different valve opening sizes with associated different air flow resistances. In some embodiments, the valve opening can be continuously adjusted by the pump controller, while in other embodiments, the valve has several discrete size settings (or releases) that the pump controller selects. Or a single setting that is either closed). Suitable active pressure relief valves include, but are not limited to, solenoid valves, diaphragm valves, ball valves, and duckbill valves.

In order to avoid electrolyte drainage during the pressure relief phase, which can ultimately cause the pump to run out of electrolyte so that the electrolytic reaction can no longer take place, the electrolyte can be, for example, hydrogel, cotton fiber, sponge, or super It may be immersed in a highly absorbent material such as an absorbent polymer. The electrolyte will then remain inside the absorbent material, separating from the gas compartment formed in the remainder of the gas chamber. The valve may be incorporated into a portion of the chamber wall adjacent to the gas compartment and / or connected to a tube that opens into the gas compartment, as shown.
(Sensor-based feedback control)

  In various embodiments, drug delivery can be controlled via electrolytically driven current to implement different delivery modes and protocols. Full function insulin pumps, for example, (1) basal delivery at a constant fluid injection rate, (2) bolus delivery (eg, manually by a patient or based on measured blood glucose levels, or throughout the day) Allows three basal delivery modes for the treatment of different types of diabetic patients, such as (3) background basal and (multiple) bolus delivery). 12A-12C schematically illustrate delivery profiles corresponding to all three modes. For example, a supper pump may administer a 150 μl dose of insulin immediately after supper and dispense another 350 μl at a basal rate over 8 hours while the patient is asleep. Different diseases may require different delivery protocols, including complex protocols with intermittent or continuous drug delivery at variable rates, with or without additional bolus injections. In general, a drug delivery protocol may specify the drug delivery time, duration, rate, and dosage depending on the particular application. Referring to FIG. 1, the pump driver 110 may be based on a selected preprogrammed delivery protocol (eg, stored in system memory 120) or based on real-time commands received via, for example, telemetry module 124. And delivery may be controlled. The clinician may change the pump programming in system memory 120 when the patient's symptoms change.

  High precision pump control according to the desired delivery mode or protocol typically utilizes sensor feedback. Different sensor types and feedback systems may be suitable for different modes. Feedback control schemes for basal, bolus, and combined basal / bolus delivery are described below. Although these control schemes will be illustrated with reference to an electrolysis-driven piston pump device (eg, as described with respect to FIG. 2A), many aspects and features of the control scheme are not It should be understood that it is also applicable to drug pump devices.

  13A and 13B illustrate an exemplary piston pump device 1300 and associated feedback loop for basal rate delivery, respectively. In order to obtain a stable flow rate, the device 1300 is preferably operated at a high driving pressure with a high resistance flow restrictor at the outlet as described above. The drive pressure required to achieve a given target flow rate can be calculated based on the known flow resistance. In turn, the flow resistance is determined prior to device deployment, for example, experimentally or by calculation based on known exit dimensions, and is stored in memory (eg, in the memory of system controller 112 or in a separate system memory 120). May be stored. Alternatively, flow resistance may be determined by calibration before drug delivery begins, eg, during the priming phase. If the flow resistance is known prior to device deployment, the target pressure for one or more target flow rates may be similarly calculated and stored in memory, for example in the form of a look-up table. Otherwise, the target pressure may be calculated later by the system controller 112 based on an input indicating the target flow rate. In various embodiments, the pump is operated in a pressure regime where the flow resistance is constant and, as a result, the flow rate is directly proportional to the drive pressure. However, the feedback loop illustrated in FIG. 13B is not subject to such a linear relationship, but can be employed whenever the relationship between flow velocity and pressure is known.

  In order to ensure pump operation at the target pressure, the pump device 1300 includes a pressure sensor 1310 that continuously monitors the drive pressure inside the pump chamber (e.g., inexpensive but accurate as used in the automotive industry). A MEMS sensor). Multiple pressure sensors may be used for increased accuracy and / or to detect sensor failure. As illustrated in FIG. 3B, the pressure sensor creates an output voltage (or other electrical signal) indicative of the measured pressure that is fed back to the electronic circuit. A signal conditioner may amplify the analog voltage signal and convert it to a digital signal. This digital pressure signal is then provided to the system controller 112 which processes the signal according to the control code to compare the drive pressure with the target pressure. Different output digital signals may be sent to the pump driver 110 to provide analog control power to adjust power to the electrolysis pump accordingly.

  In one embodiment, the control logic implemented by the system controller 112 determines whether the measured pressure is within a specified limit of the target pressure (referred to as “bias”) and adjusts the electrolysis current according to the following steps: To do. First, the current is turned on when the measured pressure is below the target pressure minus the bias value. The current is then reduced to an intermediate level when the measured pressure is within the target pressure plus / subtracted bias value. Eventually, once the measured pressure is greater than the biased target pressure, the applied current is interrupted. This process can be repeated continuously to adjust the electrolysis rate to keep the pressure constant at the target pressure. Without limitation, proportional-integral-derivative (PID) control, pulse width modulation (PWM) control, artificial neural network (ANN) control, fuzzy logic control, evolutionary computational control, model predictive control (MPC), and / or Many other control methods may be used, including linear quadrant Gaussian control (LQG). The electrolysis current achieved by the various logic schemes may generally have any waveform, for example a square or triangular wave, or may be pulse width modulated.

  An accurate and constant flow rate for basal delivery can be achieved using a high pressure sealed pump design combined with feedback control of electrolysis rate based on pressure readings in the pump chamber. As an alternative to pressure-based feedback, direct flow control can be implemented. This is illustrated in FIGS. 14A and 14B. In this case, the pump device 1400 includes a flow sensor 1410 at the drug outlet or elsewhere in the outlet member (downstream from the reservoir). The flow sensor monitors the actual delivery rate of the drug flowing out of the outlet member and outputs a voltage signal that feeds back to the control system. Similar to the control method used in pressure feedback, the flow rate voltage signal is compared to the target delivery rate and the supply current to the electrode is adjusted based on the difference between the target delivery rate and the measured delivery rate. The

  In yet another embodiment, the flow rate measurement itself is pressure based. A pressure sensor connected to the drug outlet of the reservoir (eg, an inexpensive but accurate MEMS sensor as used in the automotive industry) is used to measure pressure and the flow rate is collected pressure data. And the flow restrictor dimensions (or alternatively flow resistance as determined by calibration). The flow rate so determined is then processed in the same manner as depicted in FIG. 14B to facilitate stable drug delivery at a constant rate. The use of a pressure sensor for flow rate measurement may generally be preferred over the latter because direct flow rate measurement using a flow sensor is more expensive. In the high flow resistance embodiment as described above, the pressure readings in the pump chamber and drug outlet or anywhere in between (eg, in the drug reservoir) are approximately the same. To provide a value, the pressure drop across the piston and drug reservoir is generally negligible compared to the pressure drop across the flow restrictor.

  For applications requiring high delivery rate accuracy and patient safety, a dual sensor closed loop feedback system can be used to provide redundant control of delivery rate. Advantageously, the use of two sensors allows for safe and controlled drug delivery, even if one of the sensors suffers a strong fluctuation or failure. In a dual sensor feedback system, as illustrated in FIGS. 15A and 15B, both the pressure in the electrolysis chamber and the flow rate at or downstream from the reservoir outlet are measured and used as control signals. . A comparator 1500, typically implemented in programming, may select an action based on signals from sensors 1310, 1410. For example, the measured pressure value may be examined against a pressure / flow rate curve stored in the history module 1510 to find the corresponding calculated flow rate, the curve being, for example, during the priming phase, It may be preloaded or obtained. In some embodiments, any difference between the measured and calculated flow rates that exceed acceptable limits is used as a trigger to initiate the safety protocol. This may involve shutting down the pump and issuing a warning signal, or then restarting the pump to see if the new reading value matches.

  In an alternative embodiment, the comparator 1500 distinguishes or arbitrates between the two signals to adjust the electrolysis current based on one or the other or both. For example, in one embodiment, the measured flow rate matches the calculated flow rate within the error set limits (e.g., the measured flow rate is calculated by only 5%, 10%, or another fixed amount). The measured flow rate is acceptable and current control is based on it (eg, the same type of pressure-based control as described above for PID control, PWM control, etc.) Control mechanism). Otherwise, the calculated flow rate value is accepted and passed as the correct flow rate output value. This mode of arbitration is either more accurate than the calculated flow rate or is far away from the true value, so it is better to discard it completely and substitute it into the calculated flow rate. Based on the assumption of The scheme helps protect against inaccurate readings of the flow rate, which may be due, for example, to bubbles in the outlet member or due to sudden jerking and premature activation or deactivation of the electrolysis gas and pressure generation Can be prevented.

  Alternatively or additionally, the control system may operate the pump based at least in part on recent pump activity via the history module 1510. The pump history is a measurement of pressure and flow readings (as provided by sensors 1310, 1410) measured continuously, over absolute time, or discretely at the time the electrode is energized. It may form. Typically, as the pump operates, a continuous reading value will be stored following each pump operation. The reading value may be obtained during the rest period, or more generally not.

  The reading value is stored in the shift register or circular buffer of the history module 1510, and typically a newer reading value replaces the older reading value in a first-in first-out order. The history module 1510 may implement a hysteresis operation of the pump, and the depth of the register or buffer determines the degree of hysteresis. For example, a recent history of the relationship between pressure and flow rate is used to avoid spastic pumping due to transient events, e.g., momentary blockage of the outlet member 108 by debris carried by blood May be. Rather than increasing the pressure abruptly based on the instantaneous flow reading value, the pressure is increased based on, for example, an average of several previous flow reading values stored in the history buffer. (The number of flow / pressure reading value pairs used to determine the current pumping pressure depends on the application and can be easily determined without undue experimentation. The number of reading value pairs used is itself in the degree of deviation from the previous flow reading value so that the spike is smoothed using more previous reading values. In some embodiments, as few as three data samples are stored and averaged to predict the next.) The historical reading value is a pressure / flow relationship over time (ie, Considering longer term effects on flow rate such as increased blockage due to accumulation of biological material around the outlet port of outlet member 2410 to revise the amount of pressure required to achieve the target flow rate Is done.

  The dual sensor feedback system may also utilize two pressure sensors instead of flow and pressure sensors. Typically, one pressure sensor is placed in the pump chamber and the other sensor is at the outlet of the drug reservoir, i.e., at a location where the reservoir outlet pressure has not yet been appreciably lowered. Arranged downstream. For example, in an embodiment where there is a flow restrictor downstream of the reservoir, as shown in FIG. 9B, the second pressure sensor 930 is disposed in the fluid path between the drug reservoir and the flow restrictor 910. Is done. In embodiments that do not have discrete components downstream of the reservoir and dominate the flow resistance (eg, the entire outlet member or most of it functions as a flow restrictor), the second pressure sensor generally includes: Place as close to the reservoir as possible. When the flow resistance in the reservoir is high, the pressure drop across the drug vial is usually negligible. Thus, the pressure readings from both pressure sensors should be the same or nearly the same. Thus, in some embodiments, any perceptible difference between two pressure readings (eg, a relative difference greater than 5% or 10% of the higher value, or some other predetermined difference) Is interpreted as an indication of an error condition and is used to trigger a pump shutdown or another suitable safety protocol.

  Alternatively, two pressure readings upstream and downstream of the drug reservoir can be used, for example, to characterize the pressure across the reservoir as a function of drive pressure. For example, for a piston pump device with a prefilled drug vial, the static friction / friction profile can be measured and recorded prior to adoption of the device on the patient. Different types of vials from different manufacturers can vary greatly in their quality and the friction generated between the vial inner wall and the piston, for example depending on whether the inner surface is coated with a friction reducing layer. . In some vials, even after the initial static friction force has been overcome, the friction fluctuations will suddenly start when the piston movement is suddenly stopped and the drive pressure is increased to resume drug delivery. , With the risk of overdose to the patient. Rather than requiring and relying on low friction vials and / or accurate a priori knowledge of the effects of friction, a drug pump device with two pressure sensors according to the present specification determines a suitable pump drive pressure, Friction and pressure drop can be calibrated across drug reservoirs for individual vials (and / or vial types) to adjust for known variations in pressure drop during and / or drug delivery. A suitable drive pressure is, for example, a threshold at which the pressure drop and variation therein are measured over a period for different drive pressures, above which the pressure drop and consequently the flow rate at the outlet is sufficiently stable. It may be determined by establishing a pressure. Alternatively or additionally, the pressure readings from the two sensors can be used during drug delivery to compensate for any residual fluctuations in pressure drop across the reservoir to maintain a constant flow rate. .

  In some drug pump devices, flow rates of up to 1 ml / min or even higher are established. In this case, a single pressure sensor may be used to ensure proper pump operation. Safety concerns such as kinking or blockage of the fluid path, or rupture of the pump chamber would be readily detectable via the pressure reading and used to trigger pump shutdown be able to.

  Closed loop feedback for electrolysis control has been described so far for a constant basal flow rate. However, those skilled in the art will appreciate that the control system described above can be readily applied to variable target flow rates. For example, in the manner depicted in FIGS. 13B, 14B, or 15C, measured pressure or flow rate values can be compared against variable time target pressures or flow rates, and electrolysis rates can be adjusted.

  FIGS. 16A and 16B illustrate an exemplary drug pump device 1600 and associated feedback system for accurate bolus delivery. The pump device 1600 includes a pressure relief mechanism 1610 in addition to the flow sensor 1410 and associated control circuitry. Although conceptually illustrated with a valve, the pressure relief mechanism 1610 may be any of the mechanisms described above (eg, recombination spark, catalyst, etc.) or when the electrolytic pump is deactivated. Any other suitable mechanism that helps reduce the pressure may be used. The flow sensor 1410 is used to monitor the real-time flow rate of the drug delivered to the patient. A delivered dose monitor 1620, which can be implemented in programming, integrates this real-time flow rate and compares the amount against the target bolus dose to calculate the delivered bolus amount in real time. When the delivered drug amount reaches the target bolus delivery amount, the system controller causes the pump driver to block the current to the electrolysis electrode so that the pressure in the pump chamber begins to decrease. In pump embodiments that utilize an actively controllable pressure relief mechanism (rather than operating continuously), the system controller triggers the mechanism simultaneously. Thereafter, control of the electrolytic pump and / or pressure relief mechanism depends on the delivery mode.

  For bolus delivery without background basal delivery, the pressure relief mechanism is operated to allow the pressure in the pump chamber to drop to zero, causing the pump to completely stop drug delivery. Furthermore, the electrolysis pump remains deactivated and electrolysis is not resumed until the next bolus delivery is scheduled to take place according to the applicable delivery protocol. In contrast, for bolus delivery with background basal delivery, the pressure relief mechanism is controlled to reduce the drive pressure until the target drive pressure for background basal velocity is reached. Once this pressure is achieved, the pressure relief mechanism is deactivated and power to the electrolysis is restored to maintain the target pressure level. A pressure sensor 1310 and feedback loop as described with respect to FIGS. 13A and 13B may be employed to control the base velocity itself. In addition, flow sensor readings may be used to increase the accuracy and reliability of the pump control scheme. Thus, a fully functional pump system (as shown in FIGS. 16A and 16B) can implement continuous basal delivery, bolus delivery, and highly accurate delivery modes for background basal and multiple bolus delivery. A feedback loop may be combined that is used with base-only and bolus-only flow control. FIG. 17 illustrates the pump pressure and flow rate achieved over a 24-hour period using an exemplary insulin pump system including dual sensor feedback and an active pressure relief mechanism, with a stable background basal delivery rate of 500 nl / min and 3 shows three 10 μL bolus deliveries at 8, 12, and 18 hours, representing breakfast, lunch, and dinner, respectively.

  Bolus control schemes also have very low basal speeds (eg, flow rates below 400 nl / min) that make continuous delivery difficult due to the associated low drive pressure and strong impact resulting from friction between the piston and the vial. Can also be applied. A low basal rate can be achieved by discrete fixed volume (eg, 8.3 nl) bolus delivery at regular time intervals adjusted to the desired average rate. For example, as illustrated in FIG. 18A, an average rate of 0.05 μl / hr can be achieved with a single 8.3 nl injection per hour, while 2 μl / hr is as shown in FIG. 18D Requires an 8.3 nl injection every 15 minutes.

  The various feedback loops described above can generally be implemented in hardware (including analog and / or digital circuitry), software, or a combination of both. For example, the signal voltage supplied by the sensors 1310, 1410 may be first converted to a digital signal and then the system controller based on instructions stored in the system memory 120 to calculate the required electrolysis current. 112 (e.g., could be a microcontroller or a microprocessor). Pump driver 110 may receive the digital current control signal from controller 112, convert it to an analog signal, and amplify the analog signal to provide drive current to the electrodes. The instructions that implement the calculation function compare the measured and target pressure or flow rate, arbitrate between different control parameters, calculate the flow rate from the pressure value based on the stored pump history data, and calculate the delivered dose It may (but need not) be grouped into discrete modules, such as modules to integrate the flow rate to obtain. Modules are generally programmed in any suitable programming language including, but not limited to, a high level language such as C, C ++, C #, Ada, Basic, Cobra, Fortran, or Object Pascal, or a low level assembly language. The language selection may depend on the type of system controller or processor employed.

  The terms and expressions employed herein are used as descriptive terms and expressions, not as limitations, and in using such terms and expressions, the features shown and described and any equivalents thereof. There is no intention to exclude. In addition, after describing certain embodiments of the present invention, it will be appreciated by those skilled in the art that other embodiments incorporating the concepts disclosed herein may be used without departing from the spirit and scope of the present invention. It will be obvious. For example, although the present disclosure specifically relates to electrolytic pumps, certain aspects described herein, such as pumping at high driving pressures or sensor-based feedback, may also include other types of pumps (eg, electric (Chemical, osmotic, piezoelectric, pneumatic, or motor driven pump). Further, embodiments of the invention need not include all of the features or have all of the advantages described herein, but possess any part or combination of features and advantages. Also good. Accordingly, the described embodiments are to be considered in all respects as illustrative but not restrictive.

Claims (69)

A drug reservoir;
An outlet member for fluidly connecting the reservoir to a drug injection site;
A flow restrictor for restricting fluid flow through the outlet member, the flow restrictor having a flow resistance coefficient of at least 10 ⁶ μl ⁻¹ ;
An electrolysis pump comprising a pump chamber in mechanical communication with the drug reservoir via an intervening displacement member, exerting a pressure of at least 5 psi to drive the displacement member toward the outlet member; Thereby, an electrolytic pump operable to urge fluid therethrough in said drug reservoir;
A circuit for operating the pump to produce a pressure of at least 5 psi, the circuit and the flow restrictor configured to move the outlet member at a constant flow rate in the range of about 400 nl / min to about 5 μl / min. A high pressure drug pump device comprising: a circuit cooperating to cause continuous fluid flow therethrough.   The device of claim 1, wherein the pump is operable to exert a pressure of at least 10 psi.   The device of claim 1, wherein the pump is operable to exert a pressure of at least 50 psi.   The device of claim 1, wherein the pump is operable to exert a pressure of at least 100 psi.   The device of claim 1, wherein the pump is operable to exert a pressure of at least 200 psi.   The device of claim 1, wherein a minimum inner diameter of the flow restrictor does not exceed 100 μm.   The device of claim 1, wherein a minimum inner diameter of the flow restrictor does not exceed 50 μm.   The device of claim 1, wherein the flow restrictor has a length in a range of about 1 cm to about 15 cm.   The device of claim 1, wherein the outlet member comprises the flow restrictor.   The device of claim 1, wherein the outlet member comprises a cannula connected to the flow restrictor.   The device of claim 1, wherein the flow resistance coefficient is within a range that provides a substantially linear relationship between the pump pressure and a fluid flow rate through the outlet member.   The device of claim 1, further comprising a pressure sensor disposed in the pump chamber to measure pressure therein.   The circuit for operating the pump is configured to adjust the electrolysis current supplied to the electrolysis electrode based on a comparison of the measured pressure with the target pressure to cause fluid flow at the target flow rate. The device of claim 12.   The device of claim 1, further comprising a flow sensor disposed within the outlet member.   The electrolytic pump is further controlled to adjust the electrolytic current supplied to the electrode based on a comparison of a flow rate reading from the flow sensor with a target flow rate to cause fluid flow at a constant specific flow rate. The device of claim 14, further comprising a circuit.   The device of claim 1, wherein the drug reservoir is formed inside a vial and the displacement member comprises a piston movably disposed within the vial.   The device of claim 16, wherein the electrolytic pump comprises an electronics module mounted on an end of the vial and forming the pump chamber between the piston and the electronics module.   The pump chamber is sealed in a wall formed by the electronics module using an O-ring located on a peripheral edge of the vial and within a circumferential recess of the electronics module. 18. The device according to 17.   The device of claim 17, wherein the electronics module is removable and reusable in a separate drug pump device.   A pressure sensor disposed in the pump chamber for measuring the pressure therein; and a direct measurement flow sensor disposed in the outlet member; and the control circuit further comprises (i) the measured (Ii) compare the calculated flow rate with the measured flow rate, and (iii) based on the measured flow rate if within 5% of the calculated flow rate, so Otherwise configured to adjust the electrolysis current supplied to the electrode based on the calculated flow rate, the pump generating a pressure that causes continuous fluid flow through the outlet member at a target flow rate. The device of claim 1, wherein the device is operated. A drug reservoir;
An outlet member for fluidly connecting the reservoir to a drug injection site;
An electrolysis pump comprising a pump chamber in mechanical communication with the drug reservoir via an intervening displacement member, exerting pressure to drive the displacement member toward the outlet member, thereby An electrolytic pump operable to urge fluid therethrough in the drug reservoir;
A pressure sensor disposed in the pump chamber for measuring pressure therein;
A direct measurement flow sensor disposed in the outlet member to measure fluid flow velocity therethrough;
(I) calculating a flow rate from the measured pressure; (ii) comparing the calculated flow rate with a measured flow rate; and (iii) measuring the flow rate if within 5% of the calculated flow rate. A control circuit for adjusting an electrolysis current supplied to the electrode for electrolysis in the pump chamber based on the flow rate, otherwise based on the calculated flow rate, the pump at a target flow rate Manipulated to generate a pressure that causes a continuous fluid flow through the outlet member;
Drug pump device. A drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, and exerting pressure on the reservoir, thereby urging fluid from the drug reservoir into the outlet member A drug pump device comprising an electrolytic pump operable to control a fluid flow rate through the outlet member, wherein the pressure in the pump chamber of the electrolytic pump is measured and the flow rate is determined therefrom. A calculating step;
Measuring a fluid flow rate through the outlet member;
Comparing the calculated flow rate with the measured flow rate;
Adjusting the electrolysis current based on the measured flow rate if within 5% of the calculated flow rate, and otherwise based on the calculated flow rate, the pump comprising: Operated to generate pressure that causes continuous fluid flow through the outlet member,
Method. A drug reservoir;
An outlet member for fluidly connecting the reservoir to a drug injection site;
An electrolysis pump comprising a pump chamber in mechanical communication with the drug reservoir via an intervening displacement member, exerting pressure to drive the displacement member toward the outlet member, thereby An electrolysis pump operable to force fluid through the drug reservoir and a pressure sensor for measuring the pressure in the electrolysis chamber;
A flow sensor for measuring a fluid flow rate through the outlet member;
(I) storing a plurality of consecutively obtained readings from the sensor in chronological order; (ii) based at least in part on the plurality of consecutive values of the obtained readings. And a control circuit responsive to the pressure sensor and the flow sensor to operate the electrolytic pump to achieve a target flow rate.   24. The device of claim 23, wherein the time aligned readings are stored in a circular buffer.   25. The device of claim 24, wherein the reading value is moved in and out of the circular buffer in a first-in first-out order.   24. The device of claim 23, wherein the reading value is obtained at discrete time intervals regardless of pump status.   24. The device of claim 23, wherein the reading value is obtained at the occurrence of a pump event.   28. The device of claim 27, wherein the pump event is a pump actuation and a plurality of temporally separated readings are obtained when the pump is operated following each pump actuation. A drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, and exerting pressure on the reservoir, thereby urging fluid from the drug reservoir into the outlet member A drug pump device comprising an electrolytic pump operable to control a fluid flow rate through the outlet member, comprising:
(I) storing a plurality of consecutively obtained measurements of the pressure in the pump chamber of the electrolytic pump, and (ii) a flow rate through the outlet member in chronological order;
Manipulating the electrolysis pump to achieve a target flow rate based at least in part on a plurality of consecutive values of the obtained readings.   30. The device of claim 29, wherein the time aligned readings are stored in a circular buffer.   31. The device of claim 30, wherein the reading value is moved in and out of the circular buffer in a first-in first-out order.   30. The method of claim 29, wherein the reading value is obtained at discrete time intervals regardless of pump status.   30. The method of claim 29, wherein the reading value is obtained at the occurrence of a pump event.   34. The device of claim 33, wherein the pump event is a pump actuation and a plurality of temporally separated readings are obtained as the pump operates following each pump actuation. A drug reservoir;
An outlet member for fluidly connecting the reservoir to a drug injection site;
An electrolysis pump comprising a pump chamber in mechanical communication with the drug reservoir via an intervening displacement member, exerting pressure to drive the displacement member toward the outlet member, thereby An electrolysis pump operable to force fluid through the drug chamber;
A pressure sensor disposed in the pump chamber for measuring pressure therein;
A direct measurement flow sensor disposed in the outlet member to measure fluid flow velocity therethrough;
(I) calculating a flow rate from the measured pressure; (ii) comparing the calculated flow rate with a measured flow rate; and (iii) the measured flow rate is within 5% of the calculated flow rate. A drug pump device comprising: a control circuit for continuing fluid flow through the outlet member by operating the pump to generate pressure, if any, otherwise initiating a safety protocol.   36. The device of claim 35, wherein the safety protocol includes interrupting pump operation.   40. The device of claim 36, wherein the safety protocol includes resuming pump operation. A drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, and exerting pressure on the reservoir, thereby urging fluid from the drug reservoir into the outlet member A drug pump device comprising an electrolytic pump operable to control a fluid flow rate through the outlet member, comprising:
Measuring the pressure in the pump chamber of the electrolytic pump;
Measuring a fluid flow rate through the outlet member;
Calculating a flow rate from the measured pressure;
Comparing the calculated flow rate with the measured flow rate and, if the measured flow rate is within 5% of the calculated flow rate, by operating the pump to generate pressure, the outlet Continuing fluid flow through the member, otherwise performing a safety protocol.   38. The method of claim 36, wherein executing the safety protocol includes interrupting pump operation.   40. The device of claim 39, wherein performing the safety protocol further comprises resuming pump operation. A drug reservoir;
An outlet member for fluidly connecting the reservoir to a drug injection site;
A flow restrictor downstream of the drug reservoir for restricting fluid flow through the outlet member;
An electrolysis pump comprising a pump chamber in mechanical communication with the drug reservoir via an intervening displacement member, exerting pressure to drive the displacement member toward the outlet member, thereby An electrolysis pump operable to force fluid through the drug chamber;
A pressure sensor for measuring the pump pressure in the pump chamber;
A pressure sensor for measuring a reservoir outlet pressure between the drug reservoir and the flow restrictor;
A control circuit for operating the electrolytic pump to monitor a difference between the pump pressure and the reservoir outlet pressure and to achieve a substantially constant fluid flow rate through the outlet member based thereon; A drug pump device comprising: A drug reservoir, an outlet member for fluidly connecting the reservoir to a drug injection site, and exerting pressure on the reservoir, thereby urging fluid from the drug reservoir into the outlet member A drug pump device comprising an electrolytic pump operable to control a fluid flow rate through the outlet member, comprising:
Measuring the pump pressure in the pump chamber;
Measuring a reservoir outlet pressure between the drug reservoir and the flow restrictor;
Monitoring the difference between the pump pressure and the reservoir outlet pressure and operating the electrolytic pump to achieve a substantially constant fluid flow rate through the outlet member based on the difference. Method. A pump chamber, and a pair of electrodes in contact with the electrolyte to generate electrolytic gas from the liquid electrolyte upon application of a voltage between the electrodes, contained in the pump chamber;
A pressure relief mechanism for causing recombination of the electrolytic gas to the liquid electrolyte, and comprising an electrolytic pump.
Pump device.   44. The pressure relief mechanism comprises a spark ignition circuit that includes a spark gap disposed within the pump chamber, wherein spark generation in the spark gap causes recombination of at least a portion of the electrolytic gas. The device described.   45. The device of claim 44, wherein the spark ignition circuit is configured to facilitate interrupting the spark before recombination of the electrolytic gas is complete.   The pump chamber comprises first and second compartments and a gas permeable separation therebetween, and a spark plug is disposed in the first compartment, thereby allowing the first compartment to the first compartment. 45. The device of claim 44, which limits recombination of electrolytic gas.   49. The device of claim 46, wherein the separation comprises at least one of a membrane, a porous material, a solid wall with holes, or a valve.   44. The device of claim 43, wherein the pressure relief mechanism comprises a conductive filament disposed within the pump chamber, and heating of the filament upon application of current thereto causes recombination of the electrolytic gas. .   44. The device of claim 43, wherein the pressure relief mechanism comprises a catalyst that continuously causes recombination of the electrolytic gas.   50. The device of claim 49, wherein the catalyst comprises a material selected from the group consisting of platinum, palladium, nickel, iridium, or nickel cadmium.   50. The device of claim 49, wherein the catalyst comprises a nanocatalytic material.   52. The device of claim 51, wherein the nanocatalytic material comprises at least one of a nanoporous material, a nanowire, or a nanoparticle.   50. The device of claim 49, wherein the catalyst causes at least 99% by volume recombination of the electrolytic gas within as little as 2 minutes upon stopping the application of voltage between the electrodes.   44. The device of claim 43, wherein the pressure relief mechanism comprises an active pressure relief valve that fluidly connects the interior of the pump chamber with the exterior of the pump chamber.   55. The device of claim 54, wherein the valve comprises at least one of a solenoid valve, a diaphragm valve, a ball valve, or a duckbill valve.   55. The device of claim 54, wherein the liquid electrolyte is absorbed into a hydrophilic material in contact with the electrode, thereby preventing the release of the liquid electrolyte through the valve.   44. The device of claim 43, wherein the pressure relief mechanism comprises a gas permeable and liquid impermeable material that forms part of the pump chamber enclosure.   58. The device of claim 57, wherein the gas permeable / liquid impermeable material comprises at least one of porous Teflon, porous sol-gel ceramic, or sintered porous metal.   59. The portion of the enclosure formed of the gas permeable and liquid impermeable material has at least one of size or porosity corresponding to a predetermined permeability at operating pressure. Devices.   44. The device of claim 43, further comprising a control mechanism for variably controlling the pressure relief mechanism. A drug reservoir in mechanical communication with the pump chamber via an intervening displacement member;
An outlet member for fluidly connecting the reservoir to a drug injection site, and the gas generated in the pump chamber exerts pressure to drive the displacement member toward the outlet member. , Thereby urging the fluid therethrough in the drug reservoir,
44. The device of claim 43. A flow sensor for measuring a fluid flow rate through the outlet member;
(I) integrating the flow rate to determine a fluid dose, and (ii) a control circuit for interrupting the application of the voltage between the electrodes when a specific dose is reached. 61. The device according to 61.   64. The device of claim 62, wherein the control circuit is further configured to trigger the pressure relief mechanism when the specific dose is reached.   64. The device of claim 62, wherein the control circuit is further configured to variably control the pressure release mechanism to facilitate partial pressure relief. A drug reservoir;
An outlet member for fluidly connecting the reservoir to a drug injection site;
An electrolytic pump,
(I) a pump chamber in mechanical communication with the drug reservoir via an intervening displacement member in contact with the electrolyte to generate an electrolytic gas from the liquid electrolyte when a voltage is applied between the electrodes A pump chamber containing a pair of electrodes, wherein the gas generated in the pump chamber exerts pressure to drive the displacement member toward the outlet member, thereby causing the drug reservoir A pump chamber, forcing fluid through it, and
(Ii) an electrolytic pump comprising: a pressure relief mechanism for causing recombination of the electrolytic gas to the liquid electrolyte;
A flow sensor for measuring a fluid flow rate through the outlet member;
A drug pump device comprising: a control circuit for integrating the flow rate to determine a fluid dose and interrupting the application of the voltage between the electrodes when a specific dose is reached.   66. The device of claim 65, wherein the pressure relief mechanism is controllable and the control circuit is configured to trigger the pressure relief mechanism when the specific dose is reached.   66. The device of claim 65, wherein the pressure relief mechanism is continuous. A drug reservoir, an outlet member for fluidly connecting the reservoir with a drug injection site, and generating an electrolysis gas to exert pressure on the reservoir, thereby from the drug reservoir to the outlet member A drug pump device comprising an electrolytic pump operable to force fluid into the fluid, a flow sensor disposed within the outlet member, and a pressure relief mechanism for causing recombination of the electrolytic gas A method for controlling a dose of fluid delivered to an injection site, wherein the flow sensor is used to measure fluid flow through the outlet member;
Integrating the flow rate to determine a fluid dose and interrupting the generation of the electrolyzed gas when a specific dose is reached.   69. The device of claim 68, further comprising triggering the pressure relief mechanism when the specific dose is reached.

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