The present invention relates to implantable devices for sensing and reporting an O2 level in a subject using ultrasonic backscatter.
Previously known systems for continuous monitoring of regional tissue oxygenation (RTO) provide therapeutic guidance for critical care patients. This allows for a better understanding of health and disease prognosis. For example, blood oxygenation levels are useful in monitoring compartment syndrome, cancer, organ transplants and so forth. However, current technologies for RTO assessment require tethered, wired connections or batteries, creating problems related to implantation and chronic use due to their large volume. What is needed is a smaller implantable device for sensing O2 concentrations.
Described herein are systems and methods for sensing a patient's O2 level with a device implanted in the patient's tissue, and reporting the sensed O2 using ultrasonic backscatter. Further described are systems including one or more implantable devices and an interrogator.
In one aspect, there is provided a mote for measuring an O2 level of a patient, the mote comprising: a mote piezo configured to both send and receive ultrasound (US) waves; a capacitor configured to be powered by the conversion of US waves received by the mote piezo to electrical energy; and a luminescence sensor configured to be powered by the capacitor, wherein at least part of the luminescence sensor is optically isolated by an opaque material.
In some embodiments of the mote, the opaque material is black silicon.
In some embodiments, the optical isolation is optical isolation between the at least part of the luminescence sensor and tissue of a patient.
In some embodiments, the luminescence sensor is entirely optically isolated from tissue of a patient.
In some embodiments of the mote, the luminescence sensor further comprises: a light emitting diode (LED) configured for optical excitation; a biocompatible film configured for encapsulation of O2-sensitive luminescent ruthenium (Ru) dyes; and an optical filter.
In some embodiments of the mote: the capacitor is part of a mote integrated circuit (IC); the mote IC comprises a low dropout (LDO), a voltage doubler, and a light emitting diode (LED) driver; and the mote IC is configured to: in first phase: (i) power the capacitor by the conversion of the US waves received by the mote piezo to electrical energy, and (ii) duty cycle off at least one of the LDO, voltage doubler and LED driver; and in a second phase: receive an US data transmission.
In some embodiments, the luminescence sensor is configured to measure an O2 level of a patient based on the US waves received by the mote piezo.
In some embodiments, the capacitor has a value of less than 100 nF.
In some embodiments, the capacitor has a value of 2.5 nF.
In one aspect, there is provided a method for measuring an O2 level of a patient, the method comprising: in a power up phase, powering a capacitor by receiving an ultrasound (US) signal; and in a data transmission phase, receiving an US data transmission; wherein, during either the power up phase or the data transmission phase, at least one component of a mote is duty cycled off.
In some embodiments of the above-described method, the at least one component of the mote includes at least one of: a low dropout (LDO); a voltage doubler; and a light emitting diode (LED) driver.
In some embodiments of the above-described method, the at least one component of the mote includes all of: a low dropout (LDO); a voltage doubler; and a light emitting diode (LED) driver.
In some embodiments of the above-described method, the method further comprises: transmitting an electrical current generated from the received US data transmission to a luminescence sensor configured to measure the O2 level of the patient; modulating the electrical current based on the measured O2 level; transducing the modulated electrical current into an ultrasonic backscatter that encodes the measured O2 level; and emitting the ultrasonic backscatter to an interrogator.
In some embodiments of the above-described method, the method further comprises, during the data transmission phase: transmitting an electrical current generated from the received US data transmission to a luminescence sensor configured to measure the O2 level of the patient; and modulating the electrical current based on the measured O2 level. In some embodiments of the above-described method, the method further comprises, during a backscatter phase: transducing the modulated electrical current into an ultrasonic backscatter that encodes the measured O2 level; and emitting the ultrasonic backscatter to an interrogator.
In some embodiments of the above-described method, during a backscatter phase: the at least one component of the mote is duty cycled on; and the capacitor discharges to power the at least one component of the mote.
In some embodiments of the above-described method: the mote comprises a luminescence sensor configured to be powered by the capacitor; and at least part of the luminescence sensor is optically isolated by an opaque material.
In some embodiments of the above-described method: the mote comprises a luminescence sensor configured to be powered by the capacitor; and at least part of the luminescence sensor is optically isolated by black silicon.
In some embodiments of the above-described method: the mote comprises a luminescence sensor configured to be powered by the capacitor; the entire luminescence sensor is optically isolated; and at least part of the optical isolation is provided by black silicon.
In some embodiments of the above-described method, the method further comprises exciting an O2-sensitive luminescent ruthenium (Ru) dye based on the received US data transmission.
In yet another aspect, there is a device for sending and receiving ultrasound (US) signals to a mote, the device comprising: a piezo configured to send and receive ultrasound (US) waves; an US interrogator configured to control the piezo to send and receive the US waves such that: in a power up phase: a power US transmission is made to the mote; and in a data transmission phase: a data US transmission is made to the mote.
In some embodiments of the device, the US interrogator is configured to control the piezo to send and receive the US waves such that no data US transmission is made during the power up phase.
In some embodiments of the device: the piezo is further configured to receive US backscatter; and the US interrogator is configured to analyze the US backscatter to determine a measured amount of O2.
In some embodiments of the device, the US interrogator is further configured to charge a capacitor of the mote to a predetermined level by controlling the power US transmission.
In some embodiments of the device, the US interrogator is further configured to bring a voltage level of a low drop out (LDO) of the mote to a predetermined voltage level by controlling the power US transmission.
In some embodiments of the device, the US interrogator is further configured to, by controlling the power US transmission, bring: a voltage level of an analog low drop out (A-LDO) of the mote to a predetermined analog VDD (A-VDD) voltage level; and a voltage level of a digital low drop out (D-LDO) of the mote to a predetermined digital VDD (D-VDD) voltage level.
In some embodiments of the device, a luminescence sensor of the mote is optically isolated from a tissue of a patient.
In some embodiments of the device, the data US transmission is configured to cause a luminescence sensor of the mote to excite an O2-sensitive luminescent ruthenium (Ru) dye.
In yet another aspect, there is a method for measuring an O2 level of a patient using pulse-echo ultrasound (US) communication, the method comprising: dividing data into a first data packet and a second data packet, wherein the first data packet includes most significant bits and the second data packet includes least significant bits; in a first data transmission phase, transmitting the first data packet; in a second data transmission phase, transmitting the second data packet; and measuring the O2 level of the patient according to the transmitted first and second data packets.
In some embodiments of the above-described method, the method further comprises: during a first receive backscatter phase, receiving backscatter of the first data packet; and during a second receive backscatter phase, receiving backscatter from the second data packet.
In some embodiments of the above-described method, the method further comprises, prior to the first data transmission phase: in a power up phase, powering a capacitor by transmitting an US signal.
In some embodiments of the above-described method: a preamble precedes the most significant bits of the first data packet.
In some embodiments of the above-described method: a postamble follows the least significant bits of the second data packet.
In some embodiments of the above-described method: the first data packet and the second data packet are each 15 μs long.
In some embodiments of the above-described method: the most significant bits of the first data packet are five bits; and a one bit preamble precedes the most significant bits of the first data packet.
In some embodiments of the above-described method: the least significant bits of the second data packet are five bits; and a one bit postamble follows the least significant bits of the second data packet.
FIG. 5S shows an example normalized acoustic reflection coefficient (Γ) of the piezo crystal versus load resistance (Rload), measured with ultrasound at 2 MHz. Here Rload simulates Rin.
The present embodiments relate to, inter alia, systems and methods for measuring a patient's O2 level with a device implanted in the patient's tissue. In particular, continuous monitoring of regional tissue oxygenation (RTO) can provide therapeutic guidance for critical care patients. However, current technologies for RTO assessment require tethered, wired connections or batteries, creating problems related to implantation and chronic use due to their large volume.
In this regard, ultrasound (US) has been demonstrated as an efficient way to wirelessly power and communicate with implantable devices deep in tissue, enabling their miniaturization [see T. C. Chang, et al., “A 30.5 mm3 fully packaged implantable device with duplex ultrasonic data and power links achieving 95 kb/s with <10−4 BER at 8.5 cm depth,” IEEE ISSCC, 2017, pp. 460-461; see also M. M. Ghanbari, et al., “A 0.8 mm3 ultrasonic implantable wireless neural recording system with linear am backscattering,” IEEE ISSCC, 2019, pp. 284-286] by eliminating the need for wires or large batteries. The systems and methods disclosed herein present a fully wireless implantable, real-time DO monitoring system that combines a luminescence sensor with US technology. Further presented is the first fully wireless implantable luminescence sensor system for deep tissue O2 monitoring, achieving competitive or better O2 resolution, the lowest power consumption and the smallest volume (4.5 mm3) of any system previously demonstrated.
By way of overview of some of the electrical power aspects, as somewhat discussed in paragraphs 0069-0070 of WO 2018/009905, which is incorporated by reference herein, and as further somewhat discussed in column 8, line 58 to column 9, line 29, of U.S. Pat. No. 10,300,310, which is also incorporated by reference herein, an implantable device (such as a mote) includes a miniaturized ultrasonic transducer (such as a miniaturized piezoelectric transducer) and a physiological sensor (such as a luminescence sensor). The miniaturized ultrasonic transducer receives ultrasonic energy from an interrogator (which may be external or implanted), which powers the implantable device. The interrogator includes a transmitter and a receiver (which may be integrated into a combined transceiver), and the transmitter and the receiver may be on the same component or different components. The physiological sensor detects a physiological condition (such as pressure, temperature, strain, pressure, or an amount of one or more analytes), and generates an analog or digital electrical signal. Mechanical energy from the ultrasonic waves transmitted by the interrogator vibrates the miniaturized ultrasonic transducer on the implantable device, which generates an electrical current. The current flowing through the miniaturized ultrasonic transducer is modulated by the electrical circuitry in the implantable device based on the detected physiological condition. The miniaturized ultrasonic transducer emits an ultrasonic backscatter communicating information indicative of the sensed physiological condition, which is detected by the receiver components of the interrogator.
A significant advantage of the implantable device is the ability to detect one or more physiological conditions in deep tissue while being wirelessly powered, and to have those physiological conditions wirelessly transmitted to an interrogator, which can be external or relay the information to an external component. Thus, the implantable devices can remain in a subject for an extended period of time without needing to charge a battery or retrieve information stored on the device. These advantages, in turn, allow the device to be smaller and less expensive to manufacture. In another advantage, use of ultrasound allows for the relative time for data communication to be related to distance, which can aid in determining location or movement of the implantable device in real time. Further problems with current technologies include O2 consumption, susceptibility to biofouling, long readout time, and inability to operate in deep tissue. The systems and methods described herein avoid these problems and others.
More specifically, with reference to
The mote 110 may also operate at different depths. In some examples, the mote 110 may operate at centimeter scale depths on an anesthetized sheep (e.g., a large animal). In other examples, the mote 110 operates at greater depths ≥5 cm) through ex vivo porcine (anatomically heterogeneous) tissue.
The mote 110, in the example of
The luminescence sensor includes a μLED 150 for optical excitation, a biocompatible film for encapsulation of O2-sensitive luminescent ruthenium (Ru) dyes, an optical filter, and an IC fabricated in a 65 nm LP-CMOS process. In this regard,
Advantageously, the tissue is optically isolated from the luminescence sensor. To accomplish this, in some embodiments, a particular area of the encapsulation 140 (of
The sensor operates on the principle of phase luminometry, wherein the phase shift (Δϕ) between the excitation and emission signals is monitored to detect O2 concentration. In one example implementation, upon light excitation at 460 nm, the Ru-dyes emit light at 618 nm, enabling background/excitation light rejection through an optical filter (see, e.g., the example of
In one example implementation, during operation, the emitted light was square-wave modulated at a fixed operating frequency (fop=20 kHz), exciting the Ru-dyes in the PDMS film with a peak excitation power of ˜1.53 μW, resulting in an average power intensity of ˜4.9 μW/mm2 at the film surface (
Further in this example implementation, photobleaching of the Ru-dye was evaluated using a fully-packaged O2 sensor (
With further reference to the examples of
To further explain, and as discussed by in paragraphs 0131-0133 of WO 2018/009905, which is incorporated by reference herein, in some embodiments, an oxygen sensor comprises a Clark electrode. A Clark electrode measures oxygen on a catalytic surface (such as a platinum surface) surrounded by a membrane, and can be miniaturized to be included on an implantable device (e.g. a mote). The Clark electrode can be attached to an application-specific integrated circuit (ASIC) (e.g. a mote IC) on the implantable device, and variance in the amount of oxygen sensed by the implantable device (which may be blood oxygen or interstitial fluid oxygen) can modulate the ultrasonic backscatter.
In some embodiments, the oxygen sensor includes a light source (such as a light emitting diode or vertical cavity surface emitting laser (VCSEL)) and an optical detector (such as a phototransistor or a photovoltaic cell, or an array of phototransistors or photovoltaic cells). A matrix including an oxygen-sensitive fluorophore is disposed over the light source and the light detector, or in a position bridging the light source and the light detector, and the amount of light detected by the light source depends on the amount of oxygen in the surrounding fluid. Such devices can be referred to as optrodes. The matrix can include, for example, an oxygen-sensitive fluorophore (such as a ruthenium fluorophore), and increased oxygen (depending on the choice of fluorophore) can cause a faster decay of fluorescence and a decrease in intensity. This oxygen-dependent change in intensity and fluorescence decay lifetime can be detected by the optical detector. In some embodiments, the matrix is a hydrogel or polydimethylsiloxane (PDMS) polymer containing a ruthenium fluorophore. In some embodiments, the ruthenium fluorophore is bound to silica particles or silica surfaces contained within the matrix (these can be made by sol-gel processes, for example). The matrix protects the fluorophore from components in the extracellular fluid and inhibits adhesion of proteins, cells and other cellular debris that could affect the diffusion of oxygen into the matrix. Further, encapsulation of the ruthenium metal in the matrix reduces potential toxicity of the ruthenium. The light source and/or optical detector can optionally include a filter to limit emitted or detected light to a narrow bandwidth. The ASIC can drive the light source to emit a pulsed or sinusoidal light signal, which causes the light source to emit the light. The light emitted by the light source causes the fluorophore in the matrix to fluoresce. For example, in some embodiments, the light source emits a blue light or a UV light, and the fluorophore can emit an orange or red light. The fluorescence intensity and/or lifetime (decay) of fluorescence is a function of the oxygen concentration of the matrix, which is influenced by the surrounding fluid (e.g., blood or interstitial fluid). From the fluorescence decay, a fluorescent lifetime decay constant can be determined, which can reflects the oxygen amount.
Use of a light pulse emitted from the light source allows for the observation of fluoresce decay or fluorescence lifetime, which is dependent on oxygen concentration. Thus, in some embodiments, the decay of fluorescence (the fluorescence lifetime) following a light pulse from the light source is used to measure the oxygen concentration surrounding the sensor.
With further reference to the example of
In one example, the total area of the IC die, fabricated in a 65 nm low-power CMOS process, is ˜3.84 mm2. In one example implementation, the minimum electrical input power required for proper operation of the IC was ˜150 μW, generating a rectifier voltage (Vrect) of ˜1.36 V, during the O2-sensing phase. Power-intensive circuits (AFE, LED driver, voltage doubler, and TDC) were duty-cycled off during uplink transmission, reducing IC power consumption to ˜22 μW and thus avoiding the need for a large off-chip Cstore. The average power dissipation of the IC drops to less than 150 μW, including the rectifier's power conversion efficiency, during operation, depending on the O2 sampling rate. The sampling rate (fS) of the system was externally controlled through the external receiver.
In one example, during operation, the external transceiver was switched from TX to RX mode to capture uplink data encoded in the backscatter reflections from the sensor's piezo. The RX path demodulated and decoded the received backscatter, generating real-time O2 data. The data was sent to a computer through a serial link for data storage and further analysis. In order to avoid overlapping TX and RX pulses, the data packet duration (TDM) was kept shorter than the round-trip time-of-flight (2ToF) of a US pulse between the sensor's piezo and the external transducer, (see, e.g.,
It should be understood that in the examples of
In some embodiments, a time-to-digital converter (TDC), operating with a 16 MHz on-chip clock generated by a 5-stage current-starved ring oscillator, converted the time delay (phase difference, Δϕ) between the reference signal (ϕref), used to drive the μLED, and the luminescence signal (ϕPD) into a 10-bit digital data. The 10-bit data may be serialized and divided into two equal 15 μs-long data packets with a preamble and a postamble by a finite state machine; the first packet contains the most significant bits (MSBs). The measured minimum detectable average optical power, yielding a signal-to-noise ratio (SNR) equal to 1 at a 1 Hz bandwidth, was ˜1.3 pW at the peak emission wavelength of ˜621 nm and the operating frequency of 20 kHz for the optical readout, which was dominated by the noise of the AFE. The SNR was ˜53 dB under a typical ˜6.7 nW/mm2 light power after the optical filter, produced by the excited Ru-dyes in the O2-sensing film, for the sensor operated in room air at 37° C. An uplink data transmission began when the on-off-keying (OOK) demodulator detected a “falling edge” in the US data input from the external transceiver, generating a notch (VOOK). The notch served as a reference to time synchronize the sensor IC and the external transceiver during uplink transmission; a data packet was transmitted to the external transceiver after a notch with a duration of shorter than ˜64 μs, equivalent to 127 oscillations of a 2 MHz US carrier. Data packets were encoded in the US reflections from the sensor's piezo and transmitted via digital amplitude modulation of the backscatter. Backscatter amplitude modulation was achieved by modulating the electrical load impedance (Rload), in shunt with the piezo impedance (Zp), through a modulation (transistor) switch controlled by ϕmod, changing the US reflection coefficient at the piezo boundary, and thus the amplitude of the backscatter. When the transistor switch was turned on to transmit the O2 data, the Rload across the piezo was reduced from a resistance value higher than 80 kΩ, depending on the IC power consumption and the amplitude of the rectifier input voltages, to ˜0.5 1Ω (transistor switch on-resistance). The uplink transmission stopped when the notch duration was kept longer than ˜64 μs.
In some embodiments, such as the example of
In one example implementation, the system sampled oxygen 350 times per second with a resolution of <5.8 mmHg/√Hz across the physiologically relevant oxygen range of interest (0-100 mmHg) and a bit error rate of <10−5. The system was first characterized in a water tank setup (e.g., such as in
In one implementation, the example system was operated as described above, with an implant placed at a depth of 5 cm. Measured waveforms and backscatter signal were recorded for the system operated at a sampling rate (fS) of 350 samples per second (Hz) using an acoustic field with a spatial-peak time-average intensity (ISPTA) of 237 mW/cm2 (˜32.9% of the FDA safety limit, ISPTA derated=720 mW/cm2, for diagnostic ultrasound) (
The system response to various O2 concentrations and the Allan deviation of the data are illustrated in the examples of
The luminescence O2 sensor response to changes in O2 concentration is reversible (see, e.g.,
In order to assess the potential effect of sterilization on the O2 sensor functionality, the sensors were first sterilized in ethylene oxide (EtO), and then their response to O2 changes was tested in DI water at 37° C. The sensor response to O2 variation before and after sterilization was nearly identical (
To evaluate the ability of the O2 sensor to resist in vivo biofouling, in vitro experiments were conducted in phosphate-buffered saline (PBS, 1×) and undiluted pooled human serum. Note that human serum, a complex fluid containing hundreds of different proteins, was chosen to simulate in vivo fouling since nonspecific protein adsorption (fouling) on the implant surface is considered the initial step in triggering a foreign-body reaction and one of the critical factors causing failure of many implants. The response of the O2 sensors incubated in PBS and serum at 37° C. was measured at low (6.5 mmHg) and high (156 mmHg) oxygen levels and different times during 10-days (
In one example implementation,
Another example implementation tested the clinical utility of the wireless, direct O2 monitoring system in a physiologically relevant large animal model (a sheep). The sheep model is a standard in fetal, neonatal and adult disease states due to the remarkable similarities in cardiovascular and pulmonary physiology, neurobiology, as well as metabolism. Anesthetized juvenile or adult sheep (n==2 animals) were intubated and mechanically ventilated. The biceps femoris was carefully dissected and the wireless sensor, as well as a commercial wired pO2 sensor, were placed in the plane below the muscle layer; and the muscle, as well as overlying skin, were closed above. The ultrasound transducer, attached to a five-axis micromanipulator for fine alignment, was placed on top of the skin layer on an acoustic standoff pad.
As seen from the data (from animal A) in
More gradual, stepwise reductions in inspired oxygen content resulted in more gradual reductions in tissue pO2, which were accurately determined by a wireless sensor in real-time with similar kinetics as the commercial wired probe (
Regarding the example of
In another example, to further evaluate uplink performance, the system was operated at 350 Hz fS and 10 cm depth in DI water with and without intentional misalignment using the same water tank setup (
In this example ex vivo measurement, the external transducer generated US waves with 706 mW/cm2 ISPTA de-rated by 0.3 dB·cm−1·MHz−1 (FDA-standard US attenuation in soft tissue24), producing an acoustic power of ˜228 mW at the transducer surface, that propagated through approximately 3 mm ultrasound gel, 1 mm skin, 1 mm fat, and 95 mm muscle tissue (
The following presents the first miniaturization of a fully implantable optrode in the mm3 volume range, which is suitable for deep-tissue measurements. Although the focus of the following is a system for measuring oxygen tension in vivo, the fundamental technological achievement opens the door to minimally invasive pulse oximetric sensors, pH sensors, CO2 sensors among others. Each of these, embodied in ultra-miniature and deep tissue systems, would open the door to novel diagnostics.
With regard to the measurement of deep-tissue oxygen tension, the application space of the wireless O2-sensing system is vast. Organ transplantation provides a clear example. The demand for organ transplantation continues to grow. In 2019, there were 94,863 candidates on the waiting list for renal transplants in the U.S. alone. [see National Data—OPTN. https://optn.transplant.hrsa.gov/data/view-data-reports/national-data/]. Monitoring graft oxygenation following organ transplantation is critical but typically relies on indirect methods that require skilled operators and provide only intermittent snapshots of tissue perfusion. Continuous and reliable monitoring of graft oxygenation following orthotopic liver transplantation, for example, may enable early detection of graft ischemia due either to hepatic artery thrombosis or graft vascular disease, allowing timely surgical re-exploration to minimize risk of graft loss, which could be fatal. Notably, these complications can occur months to years following transplant. Minimally invasive wireless modalities, such as those described herein, could enable real-time monitoring of graft oxygenation via wearable applications in the out of hospital setting, providing critical information regarding tissue oxygenation before the emergence of graft dysfunction, allowing timely intervention. This would additionally help differentiate parenchymal rejection from graft vascular disease when organ dysfunction emerges. Furthermore, more than 5.7 million patients are admitted annually to intensive care units (ICUs) in the United States (U.S.). Assessment of tissue oxygenation is a fundamental need in this setting. Although some embodiments require surgical placement, some contemplated embodiments may enable semi-invasive/vascular approaches for probe placement. Depending upon the underlying pathology, the local oxygen supply-demand balance can be distorted during pathological states, such as observed during various forms of shock. Thus, an inadequate delivery—for whatever reason—relative to demand will decrease tissue pO2. On the other hand, a primary reduction in metabolic demand or an inhibition or failure of mitochondrial oxidative phosphorylation will leave oxygen supply largely unaffected, and thus, the tissue pO2 may increase. A close matching of oxygen supply and demand, be it via an overall increase in delivery or decrease in turnover, will result in no net change in tissue pO2. Global measures of cardio-pulmonary performance such as cardiac output, oxygen delivery or blood pressure frequently do not reflect local metabolic demands at the organ and tissue level and can promote excessive fluid loading or inotrope dosing, worsening outcomes. A notable contributor in this setting is a lack of hemodynamic coherence between the microcirculation and the macrocirculation. Given that these changes typically occur over minutes to hours, a slightly longer response time than typically observed for pulse oximetry would still yield important clinical information. Coupling direct measurements of the microcirculation with direct monitoring of tissue pO2 would greatly augment critical care management approaches. Precise measurements of tissue oxygen content are therefore instrumental for the proper management of shock states, but are currently limited to indirect or surface deep methods. Novel non-invasive as well as minimally invasive modalities for monitoring deep-tissue oxygenation, as described herein, are clearly needed to advance our understanding and management of disease states where oxygen delivery or metabolism is compromised.
To clinically adopt the wireless O2-sensing system for the use of chronic, real-time in vivo O2 tracking, a number of technical challenges must be addressed. One of the main challenges is the post-surgical localization of the implant by an external transceiver since the post-surgical in vivo position may drift relative to any external fiducials; such movement can arise due to pressure from outside the body, movements or breathing of the subject, and scar formation. The in vivo localization before each pO2 measurement can be achieved with an external phased-array transceiver that utilizes ultrasound (US) backscatter information first to find and then track the time-dependent position of the implant in the body.
A second challenge arises because acoustic attenuation due to scattering and absorption varies between different US propagation paths to the implant in heterogeneous tissue; a path with higher attenuation in tissue may significantly degrade power transfer efficiency and data transfer reliability of the system. For example, muscle tissue with a more unevenly distributed intramuscular fat content will exhibit greater acoustic attenuation. Here too, an external transceiver with a large-aperture, multi-element transducer array capable of focusing US energy to the implant will allow for steering the US beam along a preferred path. Finally, a phased array could also potentially be used to interrogate multiple O2 sensors implanted in different locations of target tissue in a time-division multiplexing fashion or simultaneously.
In addition to these improvements, chronic in vivo use of the wireless O2 sensor will require hermetic packaging to prevent biofluid penetration into the electronic sensor components (IC and μLED) and the piezo crystal. Traditionally such millimeter-scale implantable hermetic housings make use of ceramic or titanium enclosures brazed or microwelded to achieve the required hermeticity. This is an active area of work both commercially and academically; an extensive review was recently published. [Shen, K. & Maharbiz, M. M. Ceramic Packaging in Neural Implants. bioRxiv 2020.06.26.174144 (2020) doi:10.1101/2020.06.26.174144]. Acoustic windows for efficient ultrasonic energy transfer into ceramic or metallic housings have recently been demonstrated in the academic literature. [Shen, K. & Maharbiz, M. M. Design of Ceramic Packages for Ultrasonically Coupled Implantable Medical Devices. IEEE Trans. Biomed. Eng. 67, 2230-2240 (2020)]. Some embodiments described herein use biocompatible polymer materials (parylene-C, silicone and UV-curable epoxy) to encapsulate the sensor given their ease of use for acute and semi-chronic experiments. It should be noted that polymeric materials at these thicknesses are not suitable for long term in vivo use of the implant due to their high water vapor permeability.
Some implementations of the film fabrication included two steps. First, luminescent dyes, tris-(Bathophenanthroline) Ruthenium (II) Perchlorate (Ru(dpp)3(ClO4)2) (CAS 75213-31-9; GFS Chemicals), were immobilized on the surface of silica particles with a diameter of 10 μm (CAS 7631-86-9; LiChrosorb Si 100 (10 μm); Sigma-Aldrich) at about 1:10 dye:particle ratio by weight. Briefly, 200 mg Ru(dpp)3(ClO4)2 complex was dissolved in 10 ml ethanol (ACS reagent ≥99.5%, CAS 459844; Sigma-Aldrich). Silica gel was prepared by adding 2 g silica particles to 40 ml aqueous NaOH (0.01 N; CAS 1310-73-2; Fisher Scientific) solution and magnetically stirring the mixture at a speed of 1000 rpm for 30 min. Next, the dye-containing ethanol solution was poured into the silica gel solution and stirred at 1000 rpm for 30 min. The dye-containing silica particles were filtered out of the solution through a filter with a pore size of 0.45 μm (Catalog number 165-0045; ThermoFisher Scientific), and then washed once in ethanol and three times in deionized water. All the supernatant was removed, and the dye-loaded silica particles were dried at 70° C. overnight.
Second, the dye-loaded silica particles were incorporated into polydimethylsiloxane (PDMS) to avoid problems related to dye leaching in aqueous media. 2 g dried silica particles were thoroughly mixed with 20 g PDMS prepolymer Part A and 2 g PDMS curing agent Part B (Sylgard 184; Dow Corning). A ˜100 μm-thick film was prepared by spinning a small amount of this mixture at 500 rpm on a microscope slide and then by curing it at 60° C. under dark and vacuum (<10 Torr) for ˜7 days, to remove solvent and air bubbles. The cured film was kept under dark at room temperature for at least 24 h before use and stored under dark at room temperature.
In some implementations, the wireless sensor was built on a 100 μm-thick polyimide, flexible PCB with electroless nickel immersion gold (ENIG) coating (Rigiflex Technology). A 750 μm-thick lead zirconate titanate (PZT) sheet with a 12 μm-thick fired on silver electrodes was diced using a dicing saw with a 300 μm-thick ceramic-cutting blade. A 750 μm3 PZT cube was first attached to a flexible PCB using two-part conductive silver epoxy with 1:1 mix ratio (8331, MG Chemicals), and then the board was cured at 65° C. for 15 min, well below the PZT Curie temperature and the melting temperature of polyimide. The top electrode of the PZT was wire bonded to the PCB using a wedge bonder (747677E; West Bond) to create an electrical connection between the PZT and the IC. The board was then encapsulated with ˜10 μm-thick layer of parylene-C using chemical vapour deposition (Specialty Coating Systems) for insulation due to its biological inertness and resistance to a moisture. The ˜10 μm-thick Parylene-C reduces the power harvesting efficiency of the PZT by ˜49% by damping its vibrations. The metal pads on the PCB for the IC and its wire bonds were carefully exposed by scoring the parylene around the pads using a sharp probe-tip and removing the parylene layer. The IC was attached to the PCB using the same silver epoxy, cured at 65° C. for 15 min, and then wire bonded to the PCB. Next, a ˜250 μm-thick optical long-pass filter with a cut-on wavelength of 550 nm (Edmund optics) were attached to the top of the IC using medical-grade, UV-curable epoxy (OG142; Epotek). The same UV curable epoxy was also used to assemble other sensor components, including a μLED with dimensions of 650 μm×350 μm×200 μm (APG0603PBC; Kingbright) and its 3D-printed holder (Protolabs), to protect the wire bonds of the chip and μLED and provide insulation. After the assembly was completed, the ˜100 μm-thick O2-sensing film was slipped through the gap between the μLED holder and the optical filter. The small residual space between the μLED holder and the film was filled by PDMS (Sylgard 184; Dow Corning). PDMS Sylgard 184 Part A and B were mixed in the ratio of 10:1, degassed, poured between the space, and cured at room temperature for 48 h. Finally, the oxygen-sensing region on the IC was coated with a ˜180 μm-thick layer of biocompatible, highly O2-permeable black silicone. The black silicone consisted of two-part, low-viscosity silicone elastomer (MED4-4220, NuSil Technology, LLC) and black, single component masterbatch (Med-4900-2 NuSil Technology, LLC); the two silicone parts (A and B) were first mixed in a 1:1 weight ratio, and then the masterbatch (4% by weight) was added, thoroughly mixed, degassed for <˜5 min, applied to the sensor surface, and cured at room temperature for 48 h.
In this example, PZT was selected as a piezoelectric material due to its high electromechanical coupling coefficient and high mechanical quality factor, providing high power harvesting efficiency. A lead-free biocompatible barium titanate (BaTiO3) ceramic with a slightly lower electromechanical coupling coefficient can be used in place of PZT.
The volume of a wireless O2 sensor was measured by using a suspension technique. In volume measurements, the sensor without test leads was suspended with a thin, rigid wire below the water surface in a container placed on an electronic balance with a measurement accuracy of 0.1 mg. The volume of the sensor was calculated from the weight difference of a water-filled container before and after submersion of the sensor in water; the weight difference, equal to the buoyant force, was divided by the density of water to determine the actual sensor volume. The volume measurements were performed using two separate sensors; the volume of each sensor was measured five times to determine reproducibility. The data obtained from all the volume measurements were presented by the mean and standard deviation values (mean±2s.d.).
In some implementations, the absorption spectrum of the O2-sensing film and the transmission spectrum of the optical filter were measured with a Jenway 6300 spectrophotometer. The emission spectra of the sensing film and the blue μLED was measured using a fiber-coupled CCD spectrometer (Thorlabs, CCS200/M) operating at an integration time of 1 s and enabled with electric dark correction. The film samples were excited at 450 nm by a laser diode (Osram, PL450B, purchased from Thorlabs) driven with a Keithley 2400 source meter, and its emission was scanned in the range of 515-800 nm. The optical output power level of the μLED was measured using an optical power meter (Thorlabs, PM100D) equipped with a Si photodiode detector (Thorlabs, S121C). The current-voltage curve of the μLED was measured with a Keithley 2400 source meter. The responsivity of the photodiode as a function of wavelength was measured using a halogen lamp coupled to a monochromator, a reference photodiode (Thorlabs, FD11A Si photodiode) and an Agilent B2912A source meter. The same photodiode (FD11A) was also used to measure the output light intensity of the μLED.
Photobleaching of the Ru-dye in the O2-sensing film was evaluated using a fully-packaged O2 sensor (
The external transceiver consisted of transmitter (TX) and receiver (RX) paths. The TX path included a commercial high-voltage pulser with an integrated TX/RX switch (MAX14808; Maxim Integrated) and a digital controller module (NI PXIe-6363; National Instruments). During the TX mode, the high-voltage pulser converted a low-voltage signal from the digital controller module to a high-voltage signal, necessary to drive an external ultrasound transducer to generate ultrasound pulses. The RX path included an ultralow noise amplifier (AD8432; Analog Devices) to receive and amplify the backscatter signal from the external transducer, a gain amplifier to further amplify the signal to a level within the input range the analog-digital converter (ADC), and a digitizer with an antialiasing filter and a 14-bit high-speed ADC (NI PXIe-5122; National Instruments) to filter and digitize the signal after receiving and amplification. In addition to the switch integrated into the pulser, an extra digitally-controlled switch (ADG619; Analog Devices) was used to minimize the electrical coupling (interaction) between the TX and RX paths. The TX and RX paths were synchronized to each other by using the same reference clock integrated into the backplane of the PXI chassis (NI PXIe-10620; National Instruments). Note that the digital controller, digitizer, and NI PXIe-8360 modules were inserted in the chassis, in which the NI PXIe-8360 module was used to connect the chassis to a computer for communication with the other modules and data transfer.
A custom Labview program (Labview 2018; National Instruments) was developed to control the modules and to process the backscatter data in real-time. A (TX and RX) communication protocol was encoded in the program. During real-time data processing, the backscatter data digitized by a 14-bit ADC with a sampling rate of 20 MHz were resampled by a factor of five and then interpolated with a sinc function. The sinc interpolation was followed by a peak detection to extract the envelope of the backscatter signal and linear interpolation to increase the number of data points and hence to improve the accuracy in the determination of an optimal threshold value that minimizes bit error rate (BER). An optimal threshold (that is, the half value of the sum of modulated and unmodulated backscatter signal amplitudes) was determined by taking the mean of the data points from the time intervals where the steady-state backscatter signal was amplitude modulated and unmodulated. The threshold was used to convert the digitized data into digital format: bits (“0” or “1”). The bits were scanned to find a preamble and a postamble and hence to extract data bits. The binary coded data (bits) were converted to numeric data, which was stored on a computer.
An in vitro characterization of the wireless oxygen monitoring system was performed in a custom-built water tank using a 25.4 mm diameter, 2.25 MHz single-element external ultrasonic transducer (V304-SU-F1.88IN-PTF; Olympus) with a focal depth of 47.8 mm, mounted on manual translation stages (Thorlabs) and connected to an external transceiver board, at various alignments and positions of the wireless oxygen sensor with test leads, mounted on top of a steel rod with a diameter of 0.75 mm connected to a manual rotation stage (Thorlabs). In measurements, the external transducer face was covered with a thin sheet of latex by filling the empty space between the transducer face and the latex sheet with castor oil (used as a coupling medium), to protect the matching layer of the transducer from possible damage due to the long-time direct contact with water or ultrasound gel. A hydrophone (HGL-0400; Onda) was used to calibrate the output pressure and hence the acoustic intensity and to characterize the acoustic beam patterns of the external transducer (
In measurements, the water tank was placed on a stirring hotplate (Thermo Scientific Cimarec), to keep the water temperature constant at 37±0.1° C., to simulate physiological temperature, and to stir using a magnetic stirrer to increase the speed of a transition from low to high O2 level and vice versa in distilled water. Water O2 concentration was monitored using a commercial O2 probe with a 300 μm core diameter (NEOFOX-KIT-PROBE; BIFBORO-300-2; Ocean Optics) varied by controlling the ratio of O2 and N2, supplied to the water tank through two pipes, via a matched pair of gas flow controllers (FMA-A2407; Omega) connected to O2 and N2 gas cylinders. A customized Matlab program controlled the gas flow controllers through a digital-to-analog converter board (NI myDAQ; National instruments) connected to a computer.
The measured phase output data from the wireless system were converted to O2 concentration (partial pressure of oxygen, pO2) in mmHg by an exponential equation: pO2 (mmHg)=A·e(B/(ϕ)+C, where ϕ is the phase output, and A, B and C are constant coefficients obtained from curve fitting (see
Ethylene oxide (EtO) sterilization with an exposure time of 4 h at 37±3° C. and an aeration time of 24 h at 37±3 ° C. was performed by a commercial vendor (Blue Line Sterilization Services LLC, Novato, Calif.).
To assess the functionality of the fully-packaged O2 sensors (
Measurements for uplink performance assessment of the system at 10 cm depth in DI water and through a fresh ex vivo porcine specimen were performed with a custom-designed and -built spherically-focused, 2 MHz, 25.4 mm diameter ultrasonic transducer with a focal length of 88.1 mm (Sensor Networks Inc.) (see
Backscatter relative difference is defined as the ratio of the amplitude difference between the modulated and unmodulated backscatter signals to the amplitude of the unmodulated backscatter signal. The modulation depth percentage was calculated by multiplying the backscatter relative difference by 100. Backscatter relative difference plots were obtained by collecting data samples from the time points where the steady-state backscatter signal was amplitude modulated and unmodulated during the O2 measurement.
The US link power transfer efficiency is defined as the ratio of the electrical input power of the IC to the acoustic power emitted from the external transducer, which depends on the beam focusing ability of the external transducer, the frequency-dependent attenuation of US intensity in the propagation media, and the power conversion efficiency of the sensor. The acoustic power at the transducer surface was calculated by integrating the acoustic field intensity data, obtained by a hydrophone at the focal length, over a circular area where the intensity of the side lobes is not negligible. The power conversion efficiency of the sensor, relying on the receive (acoustic-to-electrical conversion) efficiency of the piezo and the impedance matching between the piezo and the IC, is equal to the ratio of the electrical input power of the IC to the acoustic power at the surface of the sensor piezo; the acoustic power at the piezo surface was calculated by integrating the acoustic field intensity data from the hydrophone over the surface of the sensor piezo.
Tissue pO2 measurements were performed with the wireless system operated at a sampling rate of 350 samples per second. In the in vivo measurements, the maximum distance from the external transducer to the wireless O2 sensor, operated with an acoustic field that had a derated ISPTA of 454 mW/cm2 and a mechanical index of 0.08 (both below the FDA regulatory limits of 720 mW/cm2 and 0.19), was ˜26 mm with ˜19 mm consisting of tissue (including skin, fat, and muscle). The distance between the implanted sensor and the external transducer was estimated from the round-trip time-of-flight (that is, the time delay between the received backscatter signal from the sensor piezo and the signal that drove the external transducer). Both the wireless and the wired data, were averaged every 5 s. Two identical wireless O2 sensors were used in in vivo measurements; the first sensor response to various O2 concentrations in water and animal A was shown in
In some example implementations, to assess uplink performance, bit-error rate (BER) measurements are also performed at 50 mm depth and 360 samples per second (sps) fS in deionized (DI) water and a muscle tissue-like phantom (see also
With respect to
[3] indicates measurements from L. Yao, et al., “Sensitivity-enhanced CMOS phase luminometry system using xerogel-based sensors,” IEEE TBioCAS, vol. 3, no. 5, pp. 304-311, October 2009;
[4] indicates measurements from W. P. Chan, et al., “A monolithically integrated pressure/oxygen/temperature sensing SoC for multimodality intracranial neuromonitoring,” IEEE JSSC, vol. 49, no. 11, pp. 2449-2461, November 2014; and
[5] indicates measurements from E. A. Johannessen, et al., “Implementation of multichannel sensors for remote biomedical measurements in a microsystems format,” IEEE Trans. Biomed. Eng., vol. 51, no. 3, pp. 525-535, March 2004.
The following section will describe an example effect of US link misalignment on the system operation. Although the use of acoustic waves, instead of near-field electromagnetic waves, enabled to power and communicate with the mm-scale wireless O2 sensor at great depths ≥5 cm), it made the system sensitive to the US link alignment between the external transceiver and the wireless O2 sensor. Therefore, it was helpful in understanding the impact of US link misalignment on the system operation.
The misalignment sensitivity of the system was evaluated by measuring the sensor Vrect and the uplink bit error rate (BER) (
Alignment measurements were performed in a water tank using a 25.4 mm-diameter transducer with a focal depth of 47.8 mm, generating acoustic pulses at 2 MHz with a fixed ISPTA of 220 mW/cm2 (
The link misalignment measurements showed that the system operation exhibited relatively high tolerance to the depth misalignment compared to the transverse misalignment, as the −3 dB depth-of-field (DOF: ˜12 mm) of the single-element focused transducer is substantially higher than its beam spot size (HPBW: ˜1.2 mm) at the focal depth (
Furthermore, when implemented, any of the methods and techniques described herein or portions thereof may be performed by executing software stored in one or more non-transitory, tangible, computer readable storage media or memories such as magnetic disks, laser disks, optical discs, semiconductor memories, biological memories, other memory devices, or other storage media, in a RAM or ROM of a computer or processor, etc.
Of course, the applications and benefits of the systems, methods and techniques described herein are not limited to only the above examples. Many other applications and benefits are possible by using the systems, methods and techniques described herein.
Embodiment 1. A mote for measuring an O2 level of a patient, the mote comprising:
Embodiment 2. The mote of embodiment 1, wherein the opaque material is black silicon.
Embodiment 3. The mote of any one of embodiments 1-2, wherein the optical isolation is optical isolation between the at least part of the luminescence sensor and tissue of a patient.
Embodiment 4. The mote of any one of embodiments 1-3, wherein the luminescence sensor is entirely optically isolated from tissue of a patient.
Embodiment 5. The mote of any one of embodiments 1-4, wherein the luminescence sensor further comprises:
Embodiment 6. The mote of any one of embodiments 1-5, wherein:
Embodiment 7. The mote of any one of embodiments 1-6, wherein the luminescence sensor is configured to measure an O2 level of a patient based on the US waves received by the mote piezo.
Embodiment 8. The mote of any one of embodiments 1-7, wherein the capacitor has a value of less than 100 nF.
Embodiment 9. The mote of any one of embodiments 1-8, wherein the capacitor has a value of 2.5 nF.
Embodiment 10. A method for measuring an O2 level of a patient, the method comprising:
Embodiment 11. The method of embodiment 10, wherein the at least one component of the mote includes at least one of:
Embodiment 12. The method of any one of embodiments 10-11, wherein the at least one component of the mote includes all of:
Embodiment 13. The method of any one of embodiments 10-12, further comprising:
Embodiment 14. The method of any one of embodiments 10-13, further comprising:
Embodiment 15. The method of any one of embodiments 10-14, wherein during a backscatter phase:
Embodiment 16. The method of any one of embodiments 10-15, wherein:
Embodiment 17. The method of any one of embodiments 10-16, wherein:
Embodiment 18. The method of any one of embodiments 10-17, wherein:
Embodiment 19. The method of any one of embodiments 10-18, further comprising exciting an O2-sensitive luminescent ruthenium (Ru) dye based on the received US data transmission.
Embodiment 20. A device for sending and receiving ultrasound (US) signals to a mote, the device comprising:
Embodiment 21. The device of embodiment 20, wherein the US interrogator is configured to control the piezo to send and receive the US waves such that no data US transmission is made during the power up phase.
Embodiment 22. The device of any one of embodiments 20-21, wherein:
Embodiment 23. The device of any one of embodiments 20-22, wherein the US interrogator is further configured to charge a capacitor of the mote to a predetermined level by controlling the power US transmission.
Embodiment 24. The device of any one of embodiments 20-23, wherein the US interrogator is further configured to bring a voltage level of a low drop out (LDO) of the mote to a predetermined voltage level by controlling the power US transmission.
Embodiment 25. The device of any one of embodiments 20-24, wherein the US interrogator is further configured to, by controlling the power US transmission, bring:
Embodiment 26. The device of any one of embodiments 20-25, wherein a luminescence sensor of the mote is optically isolated from a tissue of a patient.
Embodiment 27. The device of any one of embodiments 20-26, wherein the data US transmission is configured to cause a luminescence sensor of the mote to excite an O2-sensitive luminescent ruthenium (Ru) dye.
Embodiment 28. A method for measuring an O2 level of a patient using pulse-echo ultrasound (US) communication, the method comprising:
Embodiment 29. The method of embodiment 28, further comprising:
Embodiment 30. The method of any one of embodiments 28-29, further comprising, prior to the first data transmission phase:
Embodiment 31. The method of any one of embodiments 28-30, wherein a preamble precedes the most significant bits of the first data packet.
Embodiment 32. The method of any one of embodiments 28-31, wherein a postamble follows the least significant bits of the second data packet.
Embodiment 33. The method of any one of embodiments 28-32, wherein the first data packet and the second data packet are each 15 μs long.
Embodiment 34. The method of any one of embodiments 28-33, wherein:
Embodiment 35. The method of any one of embodiments 28-34, wherein:
Each of the following documents is incorporated by reference in their entirety:
Filing Document | Filing Date | Country | Kind |
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PCT/US2021/018751 | 2/19/2021 | WO |
Number | Date | Country | |
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62978703 | Feb 2020 | US |