Patent Application
20140088555
The present invention relates to various aspects of drug pump devices and their operation, in particular, in various embodiments, to electrolysis pumps and associated control methods as well as to needle-insertion systems for subcutaneous drug infusions.
Subcutaneous drug delivery is the delivery of drugs to the subcutaneous tissue (i.e., the layer below the skin), and is a very common route of drug administration. In cases where slow absorption of a drug (such as, e.g., insulin) is required, subcutaneous delivery may even be the only allowable drug-administration method. Advantages of subcutaneous delivery over other routes of administration, such as intravenous delivery, include its suitability for at-home injections, increased convenience, and the resulting improvement in the patient's quality of life.
Conventionally, subcutaneous injections have been performed manually with syringes or pre-filled pen injectors via a needle inserted into the patient. Manual drug injection into subcutaneous tissue can, however, cause discomfort or even severe pain to the patient if drug volumes larger than 1 mL are injected, and may leave undesired trauma at the injection site. Furthermore, it is generally difficult for the patient or medical staff to hold a syringe plunger for a longer time to administer a large dosage. Therefore, drug volumes delivered subcutaneously through a single syringe injection are typically kept below 1 mL or, at most, 1.2 mL. This is a severe limitation for protein-based drug therapies (including, e.g., monoclonal antibody therapies), which frequently require larger volumes to be delivered, requiring patients to undergo multiple injections during a single dosing cycle to receive the appropriate amount. To obviate the need for large-volume injections, drug companies have spent a significant amount of time and money to make more concentrated formulations. This is, however, difficult and costly, and introduces other issues, such as increased viscosity. Therefore, there is a need for (preferably portable) drug-delivery devices capable of delivering larger drug volumes (e.g., volumes in excess of 1 mL) into subcutaneous regions over a time span longer than that achieved by conventional syringe injection.
Previously, mechanical spring-loaded or motor-driven drug-delivery devices have been used to substitute for manual drug administration. Both types of devices have certain problems. Spring-loaded devices, for example, are limited in the driving forces they can provide; thus, when a small outlet needle (e.g., a 30- to 34-gauge needle) is used (e.g., to minimize the pain and inconvenience associated with needle insertion) or a large flow resistance is encountered for other reasons (e.g., due to high viscosity of the drug), drug delivery is slow, which prolongs the injection time (thereby inconveniencing the patient). In addition, since the spring force is constant, the drug delivery rate varies when the plunger inside the drug vial of the device traverses a surface having inconsistent frictional properties; further, the plunger may slip or even completely stop when it encounters, respectively, a very wide or a very narrow vial cross-section. A motor-driven pump, on the other hand, can provide larger driving force with accurate plunger-displacement control. However, to facilitate the larger driving forces, the motor has high power and voltage requirements. A big battery is therefore often used to power the motor, making the device larger, heavier, and thus less suitable as a portable device.
Electrolysis provides an alternative pump mechanism for automated drug delivery that facilitates small devices, large driving forces, and accurate dosage control. However, the lifetime of electrolysis pumps may be limited due to corrosion and other wear on the electrolysis electrodes during pump operation. In addition, continued electrolysis when the plunger stops, e.g., due to an obstacle in its path, poses a considerable safety risk, as it can quickly over-pressurize the pump chamber. Accordingly, improved electrolysis pump devices that address these deficiencies are highly desirable.
In various embodiments, the present invention provides portable electrolysis-driven drug pump devices capable of delivering in excess of 1 mL (preferably 3 mL or more) of drug without refilling the drug reservoir, as well as methods for monitoring the devices and controlling pump operation to avoid safety hazards. Typically, a drug pump device in accordance herewith includes a pre-filled vial forming the drug reservoir therein, a plunger movable within the vial, and an electrolysis pump for driving the plunger so as to expel drug from the reservoir. The pump, in turn, generally includes a pump housing forming a pump chamber for containing the electrolyte, and one or more electrode pairs within the pump chamber in contact with the electrolyte. The electrodes may be formed on a substrate, and may be separated from each other by a trench in the substrate, a mask material, or sufficient spacing to reduce metal oxidation and metal re-deposition.
Pump control may be achieved, in various embodiments, based on pressure measurements in the pump chamber in conjunction with measurements of the drive current supplied to the electrodes: using the ideal gas law (or a variation thereof), the volume of electrolysis gas generated in the pump chamber may be computed from the pump-chamber pressure and the amount of gas (in moles) as determined from the drive current integrated over time. From the changes in the pump-chamber volume, the cumulative dispensed drug volume and/or the drug flow rate may be inferred. Adjustments to the delivery rate and/or volume may be made, if necessary, by changing the electrolysis drive current. Pump-pressure measurements may, moreover, be used to detect when the plunger stops moving, e.g., at the end of the dose or due to an occlusion in the fluid path, which entails characteristic pressure signatures. When an end-of-dose or occlusion signature is detected, pump operation may be halted, or some other appropriate action be taken.
In one aspect, the invention pertains to a method for controlling a drug-delivery device including a drug vial, a plunger movably disposed within the vial, and an electrolysis pump for causing movement of the plunger so as to expel liquid drug from an outlet of the vial. The method includes causing generation of gas in an electrolysis pump chamber by supplying a drive current to electrolysis electrodes contained in the chamber; continuously measuring (i) a drive current supplied to the electrolysis electrodes, and (ii) a gas pressure in the electrolysis pump chamber; computing a volume of electrolysis gas generated in the pump chamber based on the measured drive current and the measured gas pressure; inferring a drug flow rate and/or a cumulative delivered volume of drug from the computed volume of electrolysis gas; and controlling the drive current based, at least in part, on the inferred drug flow rate or cumulative delivered volume of drug. In some embodiments, the volume of electrolysis gas generated in the pump chamber is computed from the measured pressure and a known volume of electrolysis gas generated at a second pressure (e.g., as determined experimentally) using an ideal gas law (e.g., Boyle's law). The drug flow rate may be inferred from the cumulative delivered volume by differentiation with respect to time. The cumulative delivered volume may be based, further, on a dead volume of the pump chamber. The method may also include computationally filtering the measured drive current and/or the measured gas pressure.
In another aspect, the invention provides a drug-delivery device including a drug vial; a plunger movably disposed within the vial; an electrolysis pump comprising a pump chamber and, contained therein, electrolysis electrodes for generating gas in the chamber upon application of a drive current thereto for causing movement of the plunger so as to expel liquid drug from an outlet of the vial; sensors for continuously measuring the drive current supplied to the electrolysis electrodes and a gas pressure in the pump chamber; and a control system for (i) computing, based on the measured drive current and the measured gas pressure, a volume of electrolysis gas generated in the pump chamber, (ii) inferring at least one of a drug flow rate or a cumulative delivered volume of drug from the computed volume of electrolysis gas, and (iii) controlling the drive current based, at least in part, on the inferred drug flow rate or cumulative delivered volume of drug.
In a further aspect, the invention relates to a method for monitoring an electrolysis pump by causing generation of gas in an electrolysis pump chamber; continuously measuring a gas pressure in the electrolysis pump chamber; monitoring the measured gas pressure for at least one of an end-of-dose signature or an occlusion signature; and upon detection of an end-of-dose signature or an occlusion signature, initiating a control action (such as, e.g., shutting down pump operation, relieving pump pressure, issuing a warning signal, initiating a diagnostic to determine a cause of an occlusion, or initiating an occlusion-elimination action). The end-of-dose signature or occlusion signature may include a pressure in excess of a set pressure threshold or a rate of pressure increase above a set rate threshold. The method may further include discriminating between end-of-dose and occlusion events based on at least one of an elapsed pumping time, a volume of pumped liquid, a value of the rate of pressure increase, or a value of a time derivative of the rate of pressure increase.
In a further aspect, the invention provides a drug-delivery device including a drug vial; a plunger movably disposed within the vial; an electrolysis pump comprising a pump chamber and, contained therein, electrolysis electrodes for generating gas in the chamber upon application of a drive current thereto for causing movement of the plunger so as to expel liquid drug from an outlet of the vial; a sensor for continuously measuring a gas pressure in the pump chamber; and a control system for (i) monitoring the measured gas pressure for at least one of an end-of-dose signature or an occlusion signature, and (ii) upon detection of an end-of-dose signature or an occlusion signature, initiating a control action.
In yet another aspect, embodiments of the invention provide an electrolysis pump configured for prolonged electrode lifetime. The pump may include a pump housing forming a pump chamber for containing electrolyte therein, and, formed on a substrate and contained within the pump chamber so as to be in contact with the electrolyte, electrolysis electrodes made of metal and comprising an anode and a cathode separated from the anode by at least one of a trench in the substrate or an electrically insulating mask material (such as, e.g., epoxy or a soldering mask); separation of the anode and cathode serves to reduce metal oxidation and metal re-deposition. The electrolysis electrodes may be made, e.g., of platinum, palladium, gold, silver, iridium, aluminum, copper, chromium, titanium, or a combination thereof. In some embodiments, the electrolysis electrodes comprise a metal or metal alloy plated on a surface of a core electrode material having greater resistance to degradation under electrolysis conditions than the plated metal or metal alloy. The cathode and anode may be interdigitated, and separation of the anode from the cathode may involve separation of neighboring cathode and anode portions from each other. Alternatively, the cathode and anode may include multiple cathode-anode pairs, cathodes and anodes of the multiple pairs being arranged in an alternating fashion. The multiple cathode-anode-pairs may be electrically connected in parallel to a single power source, or each of the multiple cathode-anode-pairs may be electrically connected to its own power source.
The foregoing will be more readily understood from the following detailed description, in particular, when taken in conjunction with the drawings, in which:
Various embodiments of the present invention provide portable electrolysis-driven drug pump devices capable of delivering in excess of 1 mL (preferably 3 mL or more) of drug without refilling the drug reservoir. Such devices are particularly useful for drug therapies that require large dose volumes, such as many biologics (i.e., drugs created by biological processes rather than chemically), including, for instance, monoclonal antibodies, insulin, or other protein-based drugs. Advantageously, portable electrolysis pump devices facilitate accurately dosed automatic injections over longer periods of time (as compared with traditional syringe injections), which reduces pain and discomfort and increases the convenience of drug administration. In some embodiments, the drug pump devices are supplied with pre-filled drug containers, e.g., in the form of standard glass or polymer drug vials or cartridges, eliminating the need for the patient to fill the reservoir prior to use; typically, such pre-filled pump devices are designed for one-time use and subsequent disposal.
The pump housing 112 forms a pump chamber 114, which contains one or more pairs of electrodes 116 and, in contact therewith, an electrolyte. In some embodiments, the electrolyte is provided in liquid form. In other embodiments, the electrolyte is absorbed within a matrix (e.g., a hydrogel, cotton ball, or other absorbent material) placed within the pump chamber adjacent or surrounding the electrodes; the matrix may serve to ensure contact between the electrodes and the electrolyte regardless of the device orientation (see, e.g., U.S. patent application Ser. No. 13/091,047, entitled “Electrolytically Driven Drug Pump Devices”, filed on Apr. 20, 2011, which is hereby incorporated herein by reference in its entirety). The electrodes 116 may be, e.g., wires, pins, foil segments, wire loops, and/or microelectrodes, and may extend far into interior of the pump chamber 114 or, as shown, be disposed on a flat substrate 118, which may also form the back wall of the pump housing 112.
During pump operation, the electrodes 116 are driven by a battery or other power source (not shown) to break the electrolyte into gaseous products. Suitable electrolytes include water and aqueous solutions of salts, acids, or alkali, as well as non-aqueous ionic solutions. The electrolysis of water, for example, involves the following chemical reactions:
The net result of these reactions is the production of oxygen and hydrogen gas, which causes an overall volume expansion of the pump chamber 114, pushing the plunger 106 forward and thereby expelling drug from the reservoir 108. As an alternative (or in addition) to water, ethanol may be used as an electrolyte, resulting in the evolution of carbon dioxide and hydrogen gas. Ethanol electrolysis is advantageous due to its greater efficiency and, consequently, lower power consumption, compared with water electrolysis. Regardless of the particular electrolyte used, electrolysis provides a simple and efficient means of generating pump pressure, and can be implemented in a small, low-cost pump. For example, in some embodiments, a compact micro-electrolysis pump with an initial pump-chamber volume of only about 100 μL (or less) can deliver the entire contents of a 3-mL standard drug vial. The pump may be capable of providing a large driving force (up to 200 psi) with low power consumption (on the order of a few mW) and a low drive voltage (in the range from 3 to 5 V).
In various embodiments, the electrolysis pump uses microelectrodes disposed on a flat substrate.
Microelectrodes in accordance herewith may be manufactured, e.g., by micromachining, printed-circuit-board techniques, screen printing, electrical-discharge machining (EDM), computer-numerical-control (CNC) machining, or electroplating. Among these techniques, all of which are well-known to persons of skill in the art, PCB methods allow for low-cost electrode manufacturing, and micromachining provides good electrode performance (i.e., high gas-generation efficiency) and mechanical stability. A performance and stability comparable to that of micromachined microelectrodes can also be achieved by electroplating electrodes manufactured on a printed circuit board (PCB). PCBs typically use copper or aluminum for the electronic circuitry; these materials generally corrode quickly during electrolysis and, consequently, tend to peel off or dissolve. Electroplating a chemically more stable material onto the copper or aluminum metal core may prevent corrosion and prolong electrode life (while retaining the lower cost of PCB manufacturing as compared to, e.g., microelectromechanical-systems-(MEMS)-based manufacturing processes). Many plating metals can be used for this purpose, including, e.g., platinum, palladium, iridium, gold, silver, other noble materials, or certain alloys. The particular metal used may depend on a number of pump conditions and pump-performance criteria, including, e.g., the desired gas-generation rate, the electrolyte used, any limits on acceptable voltage and/or current requirements (as imposed, e.g., by size constraints for the power source), the required pump (and, thus, electrode) lifetime, and the requisite structural integrity of the electrode structure to achieve a desired uniformity and/or controllability of electrolysis. Iridium, for example, facilitates electrolysis at low drive voltages, and palladium, platinum, and (to a lesser extent) gold offer high long-term stability due to slow oxidation rates.
High electrode stability is particularly important for high-flow-rate applications (e.g., drug injection at rates of a few hundred μL/min to a few mL/min), since the high electrical drive currents generally required to sustain the high flow rates tend to enhance mechanisms that negatively affect the structural integrity and performance of the microelectrodes; these mechanisms include, in particular, metal delamination, metal oxidation, metal re-deposition, and metal dissolution. The risk for delamination can be reduced by selecting a metal that ensures good adhesion between the bottom seed layer and the top reactive layer. Platinum, for instance, adheres better to gold than copper or aluminum adhere to gold. To combat metal oxidation, re-deposition, and dissolution, an inert metal such as gold, palladium, iridium, and platinum may be used. Platinum is the preferred choice since it provides the most inert surface for electrochemical reactions and is characterized by the lowest oxidation and dissolution rates among metals suitable as electrode materials. There are many plating solutions that can be used for electroplating platinum onto electrodes: dinitroplatinite sulfate, chloroplatinic acid, P salt [(NH3)2Pt(NO2)2], and Q salt [(NH3)4Pt(HPO4)], to name just a few. P-salt-plated platinum (which is widely used in the aircraft industry for engine-blade coating) provides electrodes with particularly good performance and stability.
However, if platinum is not available (e.g., because its use is cost-prohibitive) and oxidation and re-deposition can be expected occur, other techniques may be used to extend the lifetime of the electrodes. In one approach, shown in
In another approach, illustrated in
In a third approach, shown in
In electrolysis pumps, pump pressure can be straightforwardly controlled via the drive current supplied to the electrolysis electrodes: the amount of gas generated is proportional to the drive current integrated over time, and can be calculated using Faraday's law of electrolysis. For example, creating two hydrogen molecules and one oxygen molecule from water requires four electrons; thus, the amount (measured in moles) of gas generated by electrolysis of water equals the total electrical charge (i.e., current times time), multiplied by a factor of ¾ (because three molecules are generated per four electrons), divided by Faraday's constant. As more gas is generated in the pump chamber, the volume of the chamber, the pressure in the chamber, or—generally—both increase; typically, the gas within the chamber can be treated, in good approximation, with the ideal-gas law, and the product of pump-chamber volume and pressure is, thus, proportional to the amount of gas (measured in moles). Any volume increase of the chamber translates directly into an equal volume of drug expelled from the drug reservoir. Accordingly, the drug flow rate equals the time derivative of the chamber volume.
The ratio of pump pressure to flow rate is given by the flow resistance of the downstream fluid path and any back pressure at the injection site. The flow resistance, in turn, is generally a function primarily of the fluid-path dimensions (e.g., the inner diameter of a needle or cannula downstream the reservoir), but can be affected significantly by occlusions in that path as well as by any friction between vial and plunger. Thus, while the flow resistance can, in principle, be derived from the device specifications and/or calibrated, it is generally desirable to also monitor the pump pressure and/or the flow rate during drug delivery. The measured parameters may then be fed into a feedback loop to control pump operation to achieve a desired flow rate and/or a desired dosage.
The system memory 308 (or memory that is part of a microcontroller) may store a drug-delivery protocol (specifying, e.g., drug delivery times, durations, rates, and dosages) in the form of instructions executable by the controller 306, which may be loaded into the memory at the time of manufacturing, or at a later time by data transfer from a hard drive, flash drive, or other storage device, e.g., via a USB, Ethernet, or firewire port. In alternative embodiments, the system controller 306 comprises analog circuitry designed to perform the intended function (e.g., to deliver the entire bolus upon manual activation by the patient). In certain embodiments, the control system 300 include a signal receiver (for uni-directional telemetry) or a transmitter/receiver 310 (for bi-directional telemetry) that allows the device to be controlled and/or re-programmed remotely by a wireless handheld device, such as a customized personal digital assistant (PDA) or a smartphone 312. The smartphone 312 may communicate with the control system 300 using a connection already built into the phone, such as a Wi-Fi, Bluetooth, or near-field communication (NFC) connection. Alternatively, a smartphone dongle 314 (i.e., a special hardware component, typically equipped with a microcontroller, designed to mate with a corresponding connector on the smartphone) may be used to customize the data-transfer protocol between the smartphone and the control system 300.
The control system 300 may be responsive to sensor measurements from one or more sensors embedded in various components of the drug pump device, such as a pressure sensor 316 in the pump chamber, one or more flow sensors 318 and/or pressure sensors in the drug reservoir and/or the cannula or other fluid-path component 319 downstream of the reservoir, and/or an Ampere-meter 320 measuring the current through the electrodes. Accurate micro-flow sensors and micro-pressure sensors suitable for use in small, portable pump devices are commercially available. With present micro-technology, micro-flow sensors generally cost much more than micro-pressure sensors, rendering control methods that rely on pressure-sensor measurements for monitoring pressure, volume, and flow rate an attractive low-cost solution.
The sensor feedback may be processed by dedicated sensor circuitry 322 and/or by the system controller 306 and used in conjunction with the pre-programmed drug-delivery protocol and/or real-time control input to monitor drug delivery and control pump operation based thereon; various suitable control mechanisms are described in detail in U.S. patent application Ser. No. 13/680,828 (entitled “Accurate Flow Control in Drug Pump Devices”), filed on Nov. 19, 2012, which is incorporated herein by reference in its entirety. When rapid infusion is the main goal and flow accuracy is less important, a simple pressure-based feedback loop may be used as a safety mechanism to monitor the drug pump device for over-delivery, under-delivery, occlusions within the fluid path, and/or the end of pumping (which occurs when the plunger 106 reaches the end of the vial). For many applications involving the dispensing of liquid, however, it is desirable to monitor and/or accurately control the flow rate (e.g., within the range from μL/min down to pL/min) and/or the cumulative volume of drug dispensed during a time period. In some embodiments, this is achieved by directly measuring the flow rate using a flow sensor downstream of the drug reservoir, and calculating the volume from the flow rate. Flow sensors, however, are not only expensive, but generally ought to be made of a biocompatible material as they come into contact with the drug. Therefore, in advantageous alternative embodiments, the delivered volume is calculated indirectly using a low-cost pressure sensor upstream of the reservoir in the electrolysis chamber of the device, where the sensor does not come in contact with the drug.
The flow rate and cumulative volumes delivered by an electrolysis pump device can be determined based on continuous measurements of the pressure in the electrolysis pump chamber and the electrolysis current supplied to the electrodes, under the following assumptions: (1) all gases in the pump chamber, including air, hydrogen, and oxygen produced by electrolysis, are ideal gases; (2) the environment sustains constant temperature and atmospheric pressure (i.e., 14.7 psia) throughout drug delivery; (3) there is no leakage of gas from the electrolysis chamber (or any non-zero leakage is negligibly small); and (4) all relevant device components, including the drug vial, pump seal (e.g., O-ring), and electrolysis chamber are rigid, i.e., do not exhibit any compliance (or change in volume) throughout delivery. If these assumptions are satisfied to a good approximation, the flow rate and/or delivered volume can be computed with significant accuracy. Even if an assumption does not hold, calibrations can typically compensate for any deviation therefrom (allowing the assumptions to be considered valid for purposes of the computations).
From the measured electrolysis current, the rate of electrolysis-gas generation can be derived as explained above; integrating the gas-generation rate over time yields the amount of electrolysis gas generated (in moles). Using the ideal gas law,
p·V=n·R·T,
the computed amount nel of electrolysis gas can be converted into the volume Vel occupied by the gas for any given pressure p (and temperature T). Thus, based on measurements of the pressure p in the pump chamber, the volume Vel of electrolysis gas in the chamber may be determined. A change in the volume of the pump chamber corresponds to a volume of liquid drug delivered from the reservoir. In practice, however, the pump-chamber volume changes not only as a result of electrolysis-gas generation, but also due to the compression of any air that is in the pump chamber even prior to pump operation. The “dead volume” V0 occupied by the air initially, i.e., before electrolysis begins, is generally known from pump design (and is typically small compared with the volume of the drug vial). Assuming that the pump chamber is initially under atmospheric pressure pat, the contribution Vair of the air to the overall volume of the pump chamber is, at a later time for a pressure p in the pump chamber:
V
air
=V
0
·p
at
/p.
Thus, the pump-chamber volume for pressure p can be computed as:
V=1/p·(V0·pat+nel·R·T).
The cumulative delivered drug volume may then be computed as the difference between the current pump-chamber volume V and the dead volume V0, and the flow rate may be derived from the delivered volume by differentiation. In some embodiments, the electrolysis pump chamber is evacuated prior to deployment of the device, dispensing with the need to account for the dead volume.
In some embodiments, the electrolysis-gas volume is not computed based on theoretical considerations, but is determined empirically. For example, in an open-benchtop test setup, the electrolysis pump may be operated under atmospheric pressure Pat (or some other known, constant pressure p1) and the gas-generation Qgg rate be measured as a function of drive current I. In one test setup, the gas-generation rate was determined to be about Qgg=(12/(1A)·I−23.7) (μL/min). From the experimentally determined gas-generation rate Qgg, the cumulative volume Vel1 of electrolysis gas may then be computed by integrating over time, or, for a constant drive current, by multiplying the gas-generation rate with the pumping time:
V
ell
=Q
gg
·t,
For a given volume V1 of gas at pressure p1, the corresponding volume V2 of the same amount (in moles) of gas at pressure p2 can be calculated using Boyle's law (which is derived from the ideal gas law):
p
1
V
1
=p
2
V
2.
Taking the dead volume V0 of air in the electrolysis chamber into consideration, the pump-chamber volume at pressure p can, thus, be computed as:
V=p
at
/p·(V0+Qgg·t)
As with the above-described approach using a computationally determined electrolysis gas volume, the cumulative delivered drug volume can be computed as the difference between the current pump-chamber volume V and the dead volume V0, and the flow rate may be derived from the delivered volume by differentiation.
Yet another way of determining the delivered drug volume and flow rate involves monitoring (and integrating over) incremental changes in the pump pressure and the amount of electrolysis gas. The incremental increase dV in the volume of the gas chamber may be computed, based on the ideal gas law, directly from the incremental changes in the electrolysis gas amount (dnel) and the pump-chamber pressure (dp):
dV=−(nel+n0)·R·T/p2·dp+R·T/p·dnel.
Herein, n0 is the amount of air (in moles) initially in the chamber (corresponding to the dead volume V0); after the electrolysis pump has operated for a while, this amount may become negligibly small. Note that, unlike the amount of electrolysis gas nel, the amount of air does not change. From the above equation, it follows that the flow rate can be determined based on the rate of change in the pump pressure and the electrolysis gas generation rate dn/dt (both of which can be derived straightforwardly from continuously acquired pressure and electrolysis-current data) according to:
dV/dt=−n·R·T/p
2
·dp/dt+R·T/p·dn/dt.
In practice, the pressure-sensor signal contains noise that, without further processing, affects the delivery-volume computations according to any of the methods described above. Therefore, in various embodiments, a real-time filter, such as, e.g., a Kalman filter, moving-average filter, running-average filter, or band-pass filter is used to eliminate or at least decrease the noise (thereby increasing the signal-to-noise ratio (SNR)); these filters are well-known to persons of skill in the art. The filter (or multiple filters) may be implemented, e.g., as a hardware module within the sensor circuitry 322. Alternatively, the filter may be provided in a software module stored, for example, in the system memory 308.
In more detail, a Kalman filter computes the monitored quantity (such as the pressure in the pump chamber) based on the modeling assumption that the true state of the quantity at time k has evolved from the state at time (k−1) according to
x
k
=F
k
·x
k-1
+B
k
·u
k-1
+w
k,
where Fk is the state-transition model, which is applied to the previous state xk-1; Bk is a control-input model, which is applied to a control vector uk; wk is the process noise, which is assumed to be drawn from a zero-mean multivariate normal distribution with covariance Qk (i.e., wk˜N(0,Qk)). At time k, an observation (or measurement) zk of the true state xk is made according to
z
k
=H·x
k
+v
k,
where Hk is the observation model, which maps the true-state space into the observed space, and vk is the observation noise, which is assumed to be zero-mean Gaussian white noise with covariance Rk (i.e., vk˜N(0,Rk)). The initial state and the noise vectors at each step {x0, w1, . . . , wk, v1 . . . vk} are all assumed to be mutually independent. Many real dynamical systems do not exactly fit this model. In fact, unmodeled dynamics can seriously degrade the filter performance, even when it is supposed to work with unknown stochastic signals as inputs, because the effect of unmodeled dynamics depends on the input, and can, therefore, cause instability of the estimation algorithm (i.e., divergence). On the other hand, independent white-noise signals do not make the algorithm diverge. Separating between measurement noise and unmodeled dynamics is a problem addressed by control theory under the framework of robust control (as known to persons of skill in the art).
It is important that a drug pump device (or other automatic liquid-dispensing device) be able to detect if the fluid path is occluded or if delivery of the reservoir contents is complete (i.e., an “end-of-dose” event has occurred for a drug pump designed for one-time use). For traditional motor-driven infusion systems, end-of-dose and occlusion events can be detected simply by monitoring the motor-gear system displacement. Such displacement monitoring for a gear system is, however, not applicable to electrolysis pump systems, which do not have a gear mechanism. Alternative techniques for detecting end-of-dose and occlusion events are therefore needed. Various embodiments of the present invention exploit the fact that, for an electrolysis engine (or other gas-generation engine), the pressure in the pump chamber exhibits a characteristic “pressure signature” when an occlusion or end-of-dose event occurs. Accordingly, monitoring the drug pump device for end-of-dose and/or occlusion events in accordance with various embodiments involves continuously measuring the pressure in the pump chamber, and monitoring the measured gas pressure for end-of-dose and/or occlusion signatures. When one of these types of events is detected, the system controller may initiate an appropriate response, such as shutting down the pump (by interrupting the drive current supplied to the electrolysis electrodes) to avoid over-pressurization of the pump chamber (which would potentially constitute a safety hazard). Alternatively or additionally, further processing of the detected signatures may be performed to discriminate between the two types of events.
In various embodiments, an occlusion or end-of-dose event is detected based on a pressure increase above a certain threshold and/or a significant change in the slope of pressure as a function of time (dp/dt).
To enable the pump to respond appropriately to a sudden pressure increase, it may be advantageous to discriminate between end-of-dose and occlusion events. In the case of the end of dose delivery, it generally suffices to shut down the electrolysis pump, and possibly to withdraw the needle, cannula, or other delivery vehicle from the patient's subcutaneous tissue. In the event of an occlusion, however, it is desirable to inform the patient thereof (in addition to interrupting pump operation for safety reasons) so that the cause of the occlusion can be found and eliminated and drug-delivery can resume. For example, the pump may perform an automatic diagnostic or occlusion-removal sequence, which may include increasing or decreasing the driving pressure and monitoring the response of the system through one or more sensors, such as the pressure sensor located in the electrolysis chamber. In case of either an end-of-dose or an occlusion event, it may be desirable to relieve the pressure within the electrolysis chamber, e.g., by means of recombination or release of gas through a valve.
A number of different methods are available for distinguishing between end-of-dose events and occlusions. In some embodiments, the cumulatively delivered volume is monitored, and when the sudden pressure increase (or other signature) is detected as the cumulative volume reaches the total volume of drug provided in the vial (or is close), the pressure increase is deemed due to the end of dose. Similarly, in some embodiments, an expected dose-completion time is computed (e.g., based on the total volume to be delivered in conjunction with the gas-generation rate), and if the event occurs close to the expected completion time, an end-of-dose event is declared. Both of these methods depend, in general, on the repeatability of drug delivery, i.e., on repeatable dispensing volumes and times. Of course, these methods can result in false negatives for the detection of occlusions, as occlusions can, in principle, occur near the end of dose. However, it is generally highly probable that a pressure increase that occurs close to the completion time and/or volume signifies an end-of-dose event.
A different means of distinguishing between occlusions and the end of dose may be provided by the pressure signature itself. In devices with a fixed overall volume (i.e., sum of pump-chamber and drug-reservoir volumes), the slope of the pressure change with respect to time, dp/dt, is typically less steep for an end-of-dose event than for an occlusion event. This is because the change in pressure dp/dt is inversely proportional to the amount of gas in the pump chamber, which is greater at the end of the dose when the plunger has reached the front-end of the vial. The slope associated with the end of dose may be calibrated in a bench-top test and stored in memory on the pump device (e.g., in the system memory 308, or in memory of the system controller 306). A comparison of the measured pressure change with the calibrated value may then be used to determine whether an end-of-dose event has happened. Furthermore, in devices that utilize an elastomer plunger and, for the reservoir, a rigid cartridge with a tapered shoulder, the concavity of the pressure with respect to time, i.e., the second derivative d2p/dt2 is typically less pronounced (i.e., has a lower value) for an end-of-dose event. The concavity of the pressure results, in this case, from compression of the (cylindrical) plunger into an approximately conically shaped space at the shoulder of the cartridge, and may likewise be calibrated. Accordingly, the second time derivative of the measured pressure may be used discriminate between end-of-dose and occlusion events.
If a single method of detection or differentiation is not sufficiently accurate (as may be the case in some applications), the results of multiple methods may be combined to produce a more accurate result. For example, in some embodiments, the results from multiple methods (e.g., analysis of the second time derivative and comparison of the time and/or cumulative volume at which the event occurs with the expected completion time and/or volume) are combined using a point system in which each result is assigned a certain number of points (or a certain weight), depending on the confidence, associated with the respective method, of accurately predicting or differentiating between events. Once the sum of the points passes a pre-set threshold, the overall confidence may be deemed sufficient to make an accurate determination.
Drug pump devices in accordance with various embodiments may be used to deliver an individual drug dosage over several minutes, or to inject drug continuously or in multiple boluses over an even longer period of time (e.g., tens of minutes, hours, or even days). At the shorter end of this range, the drug may be delivered subcutaneously directly through a needle piercing the skin. For prolonged delivery, a soft cannula preferably establishes the fluid path between the drug reservoir and the injection site; in this case, a needle may be used to insert the cannula through the skin, but the needle is thereafter retracted. In either case, the entire volume of drug in the reservoir is generally delivered with only a single needle insertion. To ensure injection at the proper subcutaneous location and prevent trauma to the injection site, it is important to avoid movement of the needle and/or cannula relative to the patient as much as possible. In some embodiments, this is achieved by integrating the needle and/or cannula into the drug pump device and placing the device securely against the patient's skin, typically with a relatively large flat or suitably curved (e.g., to conform to the patient's anatomy) surface, hereinafter the “bottom surface,” of the device housing. In certain embodiments, the drug pump device is attached to the skin via an adhesive skin patch, obviating the need to hold the device against the skin by hand and rendering it wearable. The skin patch may be adhered to the bottom surface of the device, or integrated with the device, e.g., so as to form an envelope around the other device components.
The needle may be inserted through the skin by an automatic, albeit typically manually triggered, mechanism. In certain embodiments, this mechanism, as well as the needle and/or cannula themselves, are integrated with the drug pump device to facilitate ease of use by reducing the number of system components that need to be handled and allowing the patient, once the device is properly placed against the skin, to insert the needle by the push of a button or some other simple and convenient trigger mechanism. To avoid accidental firing of the needle, the device may include a safety interlock, such as, e.g., a second, separate activation mechanism, or a skin sensor that deactivates the needle-insertion mechanism unless and until contact with the skin is detected. (A skin-sensor-based safety interlock is described, for example, in U.S. patent application Ser. No. 13/960,470, filed on Aug. 6, 2013, the entire disclosure of which is incorporated herein by reference.)
As illustrated, the needle 502 and cannula 504 ride above, with their axes parallel to, the drug vial 506. The drug vial 506, in turn, is placed in the device parallel to the bottom wall 508 of the housing. The outer surface of the bottom wall 508 is designed for placement against the patient's skin; thus, the vial 506 and portions of the needle 502 and cannula 504 are, during use, oriented parallel to the skin. Positioning the needle 502 and/or cannula 504 above or, in alternative embodiments, below the vial 506, as opposed to side-by-side with the vial, may have several advantages. For example, it allows the flexible tubing connecting the drug vial 506 to the needle 502 (as described further below) to be routed on the opposing side of the vial 506, preventing entanglement with the components of the insertion mechanism (such as the needle and cannula carriers, track, and springs, also described further below). Further, it can improve human factors by aligning the needle 502 with the vial 506, thereby providing cues that make the location of the injection site relative to the device more intuitive. It may also allow the needle 502 to be visualized through the same window in the device housing as the vial 506. In addition, needle/cannula positioning above or below the vial 506 may improve ease of assembly by moving the fluid path away from the insertion mechanism, and may provide improved access of assembly equipment for easier welding of components required for the needle-cannula interface seal. In
To deploy the device 500 for drug delivery, a fluid path needs to be established between the drug vial 506 and the injection site. As shown in
The septum needle 522 may be connected via flexible tubing 524 (made, e.g., of parylene, phthalate-free (in particular, Di(2-ethylhexyl)phthalate-(DEHP)-free) PVC, ethylene-vinyl acetate (EVA), low-density polyethylene (LDPE), polyurethane, tygone, silicone, or another soft polymer) to the injection needle 502 and cannula 504. The tubing may also be multi-layered, and may have a surface coating for drug compatibility. The connection between the septum needle 522 and the flexible tubing 524, in turn, may be established by a fluid-path connector 526 that forms a channel section into whose ends the needle 522 and tubing 524 can be seal-fitted. The fluid-path connector 526 may be contained within or be part of a septum-needle hub (not shown), which may also provide structural support for securing the vial 506 in the pump housing (including, e.g., during device manufacture while a force is applied to the vial to seal it against the electrolysis chamber). In general, the hub may be fixed or moveable. For example, the hub may be moveable through user action, such as pushing of a button on the external envelope of the device, which may correspond directly to an internal movement (e.g., a sliding action) of the hub or indirectly cause movement of the hub, e.g., via release of a spring; hub movement may result in the puncturing of the vial septum 520 with the septum needle 522. Alternatively, if the hub is fixed, the housing may be movable.
The fluid-path connector 526 and/or hub may include a vent or filter 528 that allows air, but not liquid, to pass through; such a vent/filter may be made, for example, of a fluoropolymer (e.g., of PTFE or polyvinylidene fluoride (PVDF)) membrane or of unsaturated polyester (UPE) of a mofied acrylic. The vent or filter 528 may serve two purposes: On the one hand, it may allow air bubbles that have been trapped in the drug vial 506 to escape through the vent/filter 528, preventing the injection of air into the patient. On the other hand, under circumstances in which a below-atmospheric pressure or vacuum pressure is created in the drug vial 506, the vent/filter 528 may allow air to fill the void, preventing suction in the fluid path from pulling bodily fluids or tissue from the injection site (e.g., blood and/or drug) into the device. Low pressure in the drug vial 506 can result, for instance, from the recombination of electrolysis gases in the pump chamber; such recombination may be induced deliberately (e.g., by introduction of a catalyst or via spark- or heat-ignited combustion) upon emptying of the drug vial 506 to allow for the safe handling and/or disposal of the pump device. However, recombination may cause a vacuum in the pump chamber which, via retraction of the plunger, may create suction in the vial 506 and can, if the pump device is still in fluid communication with the patient's body, draw matter from the site where liquid drug was just dispensed. This undesirable effect is undermined by the vent or filter 528.
In some applications, the fluid path downstream of the drug vial 506 (or a segment of the fluid path) needs to be primed prior to infusion, i.e., liquid drug needs to be pumped into the fluid path before a fluid connection with the injection site in the patient is established. This can be achieved by briefly operating the electrolysis pump following insertion of the septum needle 522 through the vial septum 520. Alternatively, priming may be accomplished by a separate, e.g., manual, mechanism. For example, the user may actuate a button, lever, or other mechanical element on the housing to move the plunger slightly to dispense the liquid drug.
With renewed reference to
The pump device 500 may include a needle guide 542 that, upon advancement of the injection needle 502 and cannula 504, deflects the needle 502 and cannula 504 from their horizontal axis (i.e., the axis parallel to the bottom wall 508) and guides them downward at a pre-determined angle toward an opening 544 in the bottom wall 508 of the device housing (which, if an adhesive patch is used, is generally aligned with an opening in the patch). Significant needle bending is possible, despite the rigid material from which the needle 502 is typically made, if the needle wall is sufficiently thin (e.g., between 0.002 and 0.006 inches). The needle guide 542 may define a curved guide channel aligned, at one end, with the axis of the needle 502 and, at the other end, with the opening 544, and may be formed integrally with the housing or a part thereof (e.g., with the bottom wall 508 or the device cover (not shown)), or provided as a separate component secured to the housing or other part of the device. In some embodiments, the needle guide 542 also provides support for and/or aids in the alignment of the septum needle 522; for example, the septum-needle hub may be integrated with the needle guide.
For many subcutaneous drug-delivery applications, it is desirable to insert the needle 502 and/or soft cannula 504 into the patient's skin at a 45° angle, which typically helps minimize pain and discomfort during the injection. However, other angles (e.g., 30° or 90°) are possible and may, in some circumstances, be desirable. The needle guide 542 can generally be designed to direct the needle 502 and/or cannula 504 into the skin at the desired angle. In various embodiments, as shown in
The needle insertion mechanism of the pump device 500 is illustrated in more detail in
The lever 536 pivots about a horizontal axis 708 that is well-supported by a structural element of the housing or a scaffold therein. Upon rotation of the lever 536 around this axis, the slot in the end stop 540 is unblocked, releasing the cannula carrier 532. The cannula carrier 532 advances, urged forward by the insertion spring 704. At the same time, the needle carrier 530 advances, pulled by the cannula carrier 532, e.g., via a latch mechanism. The carriers 530, 532 may travel along a guiding rod 710 that is support only on one end to prevent binding (i.e., over-constraining). The latch mechanism, shown in a close-up view in
The advancing motion ceases once the cannula carrier 532 reaches its physical end of travel when hitting the end stop 540;
Of course, the device 500 described with respect to
In the illustrated configuration, the latches 714 are integrated with the cannula carrier 532 and symmetrically placed around the latch head 708. In other embodiments, however, the latches 714 may instead be located on the needle carrier 530 or another component. Further, the latches 714 need not be symmetrical about a cross-sectional plane, and may be oriented differently from the manner illustrated (e.g., at an angle or vertically, in relation to the axis of carrier travel). Alternatively or in addition to being constrained by side walls, the latches 714 may be constrained from the top and/or bottom. The latches 714 may also be mechanically assisted to release the needle carrier 530 (at the appropriate time) by physically contacting other elements. The physical ends of travel may be cushioned by, e.g., an elastomer, shock absorber, integral rigid spring, etc. to reduce device shock and sound level.
The insertion mechanism may be designed to permit assembly of the carriers, springs, and other components from one end of the housing, providing for simple assembly that may be done manually, semi-automatically, or automatically. The carriers and springs may form a subassembly that is put together separately and then introduced into the device housing. The cannula and/or needle may be guided at different angles with respect to the insertion/retraction mechanism and may follow different paths, according to the design preferences. The cannula/needle may also be offset from the points at which they are attached to the carriers, and guided to protrude from the device at a desired location (e.g., centered in relation to the body of the device).
The terms and expressions employed herein are used as terms and expressions of description and not of limitation, and there is no intention, in the use of such terms and expressions, of excluding any equivalents of the features shown and described or portions thereof. Rather, having described certain embodiments of the invention, it will be apparent to those of ordinary skill in the art that other embodiments incorporating the concepts disclosed herein may be used without departing from the spirit and scope of the invention. Further, embodiments of the invention may possess any or all of the features and advantages described herein, in any suitable combination, even if such combinations were not made express herein. For example, while the embodiments described above generally refer to electrolytically driven drug pump devices, certain inventive aspects, such as needle-insertion mechanisms integrated into drug pump devices, may be applicable to devices with other pump mechanisms as well. Conversely, the above-described features of electrolysis pumps may be used in drug-delivery devices with or without integrated needle-insertion mechanisms. Accordingly, the described embodiments are to be considered in all respects as only illustrative and not restrictive.
The present application claims priority to and the benefit of, and incorporates herein by reference in their entireties, U.S. Provisional Applications No. 61/704,974, filed on Sep. 24, 2012, and No. 61/875,470, filed on Sep. 9, 2013.
| Number | Date | Country | |
|---|---|---|---|
| 61704974 | Sep 2012 | US | |
| 61875470 | Sep 2013 | US |
