Optogenetics is commonly used for precision modulation of the activity of specific neurons within neural circuits ^[1] , but assessing the impact of optogenetic neural modulation on millisecond-timescale local and global circuit neural activity remains difficult.  We have developed a novel strategy for designing and fabricating silicon-based microelectrode arrays with customizable electrode locations, targetable to defined neural substrates distributed in a 3-D pattern throughout a neural network in the mammalian brain, and compatible with simultaneous use of a diversity of existing light delivery devices.  Our design of these 3-D electrode arrays provides for both easy electrical and mechanical assembly, and provides for scaling of arrays to up to 1000 neural recording channels and beyond.

Our approach relies upon a number of innovations at the material, structural, electrical, and data acquisition levels.  First, typical silicon-based electrodes that are arranged in a 1-dimensional linear array, or 2-dimensional comb-like fashion, often use linear or tetrode-style electrode locations along the comb’s fingers, with stereotyped spacing and pad sizes.  Our software-driven approach enables variable spacing and pad sizes, so that electrode geometries can be customized to the cellular properties of the brain circuits under investigation.  Second, to support the assembly of such electrode arrays into a 3-dimensional array, we have developed novel electrical and mechanical connector strategies to make the assembly as automated and reliable as possible.  Third, we have developed strategies for amplifying and acquiring data that simplify the use of these probes in an intact, in vivo, mammalian context.  Fourth, we have implemented hybrid electrodes that contain both a low-impedance metallic pad for recording of spike activity, as well as an indium tin oxide (ITO) pad that can report local field potentials (LFPs) without the photo-electrochemical artifacts common in optogenetics.  Finally, these 3-D probes are designed to be easy to use, from design to surgery.  We have developed a user-friendly interface that enables neuroscientists to specify probe geometries based upon neural target geometries and coordinates, and are developing supporting surgical and behavioral strategies for use of such arrays in vivo.

1. X. Han and E. S. Boyden, “Multiple-color optical activation, silencing, and desynchronization of neural activity, with single-spike temporal resolution,” PLoS ONE, vol. 2, no. 3, p. e299, Mar. 2007.

An electrospray emitter ionizes polar liquids using high electrostatic fields. The electric field produces suction on the free surface (meniscus) of an electrically conductive liquid, and the surface tension of the liquid tends to counteract the effect of the electrostatic suction. If the electric field is larger than a certain threshold, the meniscus snaps into a conic shape called a Taylor cone ^[1] (Figure 1). A Taylor cone emits charged particles from its apex due to the high electrostatic fields present there; these particles can be ions, droplets, fibers, etc., depending on the working liquid and the emitter flowrate ^[2] . In particular, electrospray in cone-jet mode ^[3] creates near-monodispersed charged droplets that can be used for many applications including mass spectrometry ^[4] , etching ^[5] , and nanosatellite propulsion ^[6] . In this project we are exploring electrospray in cone-jet mode as a technology to create controlled coating of electrospun nanofiber mats (Figure 2) with liquids such as fluorescent dye and nanoparticles solutions, as an alternative technology to nano-pipetting or ink jet printing. The long-term goal of the project is to investigate the design space of the technology to make low-cost and low false-positive biochemical detectors through the exploration of the multiplexing and scaling-down limits of cone-jet mode electrospray sources using batch micro- and nanofabrication ^[7] .

1. G. I. Taylor, “Disintegration of water drops in an electric field,” Proc. R. Soc. London A vol. 280, pp. 383 – 397, 1964.
2. J. Fernandez de la Mora, “The fluid dynamics of Taylor cones,” Ann. Rev. of Fluid Mech., vol. 39: pp. 217– 243, 2007.
3. J. Fernandez de la Mora, “The current emitted by highly conductive Taylor cones,” J. Fluid Mech., vol. 260, pp 155 – 184, 1994.
4. J. B. Fenn, M. Mann, C. K. Meng, S. F. Wong, and C. M. Whitehouse, “Electrospray ionization for mass spectrometry of large biomolecules,” Science, vol. 246, no. 4926, pp. 64-71, 1989.
5. M. Gamero-Castaño and M. Mahadevan, “Sputtering of silicon by a beamlet of electrosprayed nanodroplets,” Appl. Surf. Sci., vol. 255, pp. 8556-8561, 2009.
6. L. F. Velásquez-García, A. I. Akinwande, and M. Martinez-Sanchez, “A planar array of micro-fabricated electrospray emitters for thruster applications,” J. of Microelectromech. Syst., vol. 15, no. 5, pp. 1272–1280, 2006.
7. B. Gassend, L. F. Velásquez-García, A. I. Akinwande, and M. Martinez-Sanchez, “A microfabricated planar electrospray array ionic liquid ion source with integrated extractor,” J. of Microelectromech. Syst., vol. 18, no. 3, pp. 679 – 694, 2009.

This project aims to develop the technology for field-enabled low-power portable vacuum sources that can be made cheaply and reliably, opening the doors to exciting applications such as portable mass spectrometers and high-performance sensors for inertial navigation. Our micropump uses arrays of isolated vertically aligned carbon nanotubes (VA-CNTs) to field-ionize the background gas, that is, to quantum tunnel electrons from the outer shell of neutral gas molecules due to the presence of a very high electrostatic field near the VA-CNT tip (Figure 1). Field strength of at least 10⁸V/cm is needed to field-ionize gases ^[1] .  The ions are then implanted in a non-evaporative getter structure biased at a high negative voltage, hence obtaining vacuum. The field ionization micropump that we are developing is designed to work at pressures as high as 30 Torr.  Our fabricated field ionizer, shown in Figure 2, is composed of arrays of VA-CNTs surrounded by a ring of VA-CNTs. The central VA-CNT of each unit enhances the electric field to achieve field ionization, while the high-transparency ring increases the flux of neutral molecules to the ionization region. VA-CNTs are ideal for field ionization because of their high aspect ratio, which enables low-voltage field ionization and their inherent chemical and mechanical robustness.  Unlike electron impact ionizers ^[2] , the field enhancer of a field ionizer is biased at a higher voltage than the gate.  Therefore, the ions it creates do not stream back to the field enhancers, which results in enhanced reliability.  The getter will be biased at a lower potential with respect to the gate to attract and implant the positive ions. Current research efforts include optimization of the fabrication of the devices and experimental characterization as ionizers and pumps.

1. R. Gomer, Field Emissions and Field Ionization, New York: Springer-Verlag, Dec. 1992.
2. L.-Y. Chen and A. I. Akinwande, “Aperture-collimated double-gated silicon field emitter arrays,” IEEE Transactions on Electron Devices, vol. 54,  no. 3, pp. 601-608, Mar. 2007.

With the escalating costs of hospital visits, clinicians are opting to use at-home monitoring devices to diagnose patients.  Current ECG Holter monitoring devices typically have 24-48 hour memory and battery capacity ^[1] .  With many patients experiencing intermittent heart problems that can occur once every week or month, the Holter monitor is not a good solution; an event recorder or loop recorder is required ^[2] .  However, each of these recorders can save only up to a few minutes of ECG recordings.  This constraint leads to the loss of most of the data, which could be very important in alerting the user to the onset of future episodes.  Therefore, we have developed a Holter monitor prototype with the goal of battery and memory capacity of two weeks.  Figure 1 shows a block diagram of the system.

We based the long-term monitor prototype around a Texas Instruments MSP430 low-power microcontroller that enables high computing power with very low power consumption.  The prototype monitor is mounted on standard 3M 2560 Red Dot electrodes and fabricated on a flexible PCB substrate.  Mounting the PCB directly on the electrodes improves the SNR by an estimated 40 dB compared to using wired leads ^[3] .  The monitor is “L”-shaped with rounded corners and placed on the patient’s chest (Figure 2).  The “L” shape enables several different ECG vectors to be recorded, depending on what the cardiologist wants to observe.  The monitor has 1 GBit of FLASH memory, which is enough to store 6 days of data sampled at 250 Hz continuously without compression.  Total power consumption of the system is approximately 2 mW.

1. D. Jabaudon, J. Sztajzel, K. Sievert, T. Landis, and R. Sztajzel, “Usefulness of ambulatory 7-day ECG monitoring for the detection of atrial fibrillation and flutter after acute stroke and transient ischemic attack,” Stroke, J. Amer. Heart Assoc., vol. 35, pp. 1647–1651, May 2004.
2. M. A. Rockx, J. S. Hoch, G. J. Klein, R. Yee, A. C. Skanes, L. J. Gula, and A. D. Krahn, “Is ambulatory monitoring for “Community-acquired” syncope economically attractive? A cost-effective analysis of a randomized trial of external loop recorders versus Holter monitoring,” AHJ 150(5), pp. e1.1065 – e5.1075, Nov. 2005.
3. A. Searle and L. Kirkup, “A direct comparison of wet, dry and insulating bioelectric recording electrodes,” Physiol. Meas., vol. 21, pp. 271-283, 2000.

Intracranial pressure (ICP) is a key factor in monitoring a patient’s cerebrovascular state.  However, current ICP measurement modalities are highly invasive, relying on surgical penetration of the skull.  Recent developments in model-based physiology allow ICP to be estimated using arterial blood pressure and cerebral blood flow velocity (CBFV) measurements ^[1] .  CBFV can be obtained non-invasively using transcranial Doppler (TCD) ultrasonography, but requires bulky capital equipment and an expert operator to manually focus the ultrasound beam on a particular intracranial blood vessel.  Therefore, TCD measurements of CBFV are currently restricted to clinical environments in which such technology and expertise are available (typically neurocritical care units).

This project seeks to develop a low-power, miniaturized TCD ultrasonography system for measuring CBFV in the middle cerebral artery (MCA) in support of continuous monitoring of ICP.  The MCA is typically about 3 mm in diameter and is insonated through the temporal bony window at a distance of 40 to 60 mm from the ultrasonic transducer array.  These anatomic considerations place significant constraints on the focal length and spatial resolution requirements of the transducer array.  Adjusting the transmit amplitude and phase of each element in the 2D transducer array via a digital beamformer and high voltage (HV) pulser achieves electronic beam steering in three spatial dimensions.  Figure 1 shows relative acoustic power density for a 15° off-axis focus using a 2D transducer array with electronic beam steering.

Development of a beam steering algorithm will allow for autonomous location of the MCA, eliminating the need for a skilled operator. TCD ultrasonography focusing is further complicated by the highly non-homogenous acoustic propagating medium (i.e., presence of high-density cranium).   This issue can be mitigated, however, using calibration methods ^[2] . An HV multiplexer (MUX) is utilized so that a single transmit/receive (T/R) channel can connect to multiple transducer elements and thus greatly reduce the necessary electronics and power requirements.  This system architecture, as illustrated in Figure 2, will allow for a self-contained system for continuous CBFV measurement in a low-power and wearable form-factor.

1. F. M. Kashif, T. Heldt, and G. C. Verghese. “Model-based estimation of intracranial pressure and cerebrovascular autoregulation,” Computers in Cardiology, vol. 35, pp. 369-372, Sep. 2008.
2. G.T. Clement and K. Hynynen, “A non-invasive method for focusing ultrasound through the human skull,” Physics in Medicine and Biology, vol. 47, pp. 1219-1236, Apr. 2002.

Vital signs such as heart rate, blood pressure, blood oxygenation, cardiac output, and respiratory rate are necessary in determining the overall health of a patient.  Continuous monitoring of these vital signs can help assess the wearer’s overall state of health and identify risks for cardiovascular diseases ^[1] .

We propose the site behind the ear as a location for an integrated wearable vital signs monitor ^[2] . This location offers physiological signals such as the electrocardiogram (ECG), the photoplethysmogram (PPG), and the head ballistocardiogram (hBCG). The ECG measures the electrical activity from the heart and offers information such as continuous heart rate, blood pressure (when coupled with PPG or hBCG), and respiratory rate. The PPG measures the blood volume and color under the skin using optical illumination. The PPG offers information such as continuous heart rate and blood oxygenation. The hBCG measures the head’s mechanical reaction to the blood expelled by the heart and offers information about continuous heart rate, cardiac output, and respiratory rate.

A simultaneous measurement of ECG, PPG, and hBCG is shown in Figure 1. Using the peak timing data from ECG, PPG, and hBCG, blood pressure can be estimated. Figure 2 compares the estimated blood pressure with a commercial blood pressure measurement during a Valsalva breath-hold maneuver.

To make the monitor wearable, the electrodes must be small, comfortable, and gel-less to avoid skin irritation. We use 1-cm² capacitive and dry electrodes made of wearable fabric materials. The device is designed to use the ear as a discreet and a natural anchor that reduces device visibility and the need for skin adhesives.

1. S. D. Pierdomenico, M. Di Nicola, A. L. Esposito, R. Di Mascio, E. Ballone, D. Lapenna, F. Cuccurullo, “Prognostic value of different indices of blood pressure variability in hypertensive patients,” American Journal of Hypertension, vol. 22(8), pp. 842-847, June 2009.
2. D. He, E. S. Winokur, T. Heldt, C. G. Sodini, “The ear as a location for wearable vital signs monitoring,” Proc. of the IEEE Engineering in Medicine and Biology Conference, Sept. 2010, pp. 6389-6392.

Figure 1: Simplified system block diagram.

Epilepsy is a common chronic neurological disorder that affects about 1% of the world population ^[1] . It is characterized by repeated seizures, which are caused by an abnormal neuronal firing rate of the affected brain area. One way to detect a seizure is through an electroencephalogram (EEG), which is the recording of the electrical activity produced by the firing of neurons in the brain. Continuous EEG recording is extremely important for patients with epilepsy because doctors cannot only track the number of seizures the patient has, but also have access to the recordings during the attacks which allows doctors to determine the efficacy of the treatment.

The most common EEG recording place is at the scalp; however, it can also be obtained at the skull (subdermal) or at the brain (subdural). Continuous EEG recordings through measurements at the scalp require that the patient wears an external medical device at all times, which can be extremely inconvenient. One way to solve this problem is to implant the medical device. A subdural EEG implant would require a very complex surgery, so we have decided to do a subdermal EEG and avoid such complications. Our minimally invasive implant will be placed between the scalp and the skull behind the right ear, and the electrodes will run to the front part of the skull.

Our system consists of 2 EEG channels sampled at approximately 250 Hz with a 12-bit resolution. Figure 1 shows the simplified system block diagram.

1. W. C. Stacey and B. Litt, “Technology insight: neuroengineering and epilepsy – designing devices for seizure control,” Nature Clinical Practice Neurology, vol. 4, pp. 190-201, 2008.

The Capacitive Micromachined Ultrasound Transducer (CMUT) is an alternative to traditional piezoelectric transducers. The CMUT technology provides an opportunity for highly integrated ultrasound-imaging system solutions because of its CMOS compatibility, ease of large array fabrication, and improved bandwidth and sensitivity performance ^[1] .

This project aims to provide a highly flexible platform for 3D ultrasound imaging. Figure 1 presents the system architecture. The CMUT device is flip-chip bonded to the supporting electronic circuits, which eliminates the cables that are usually required by traditional systems between the piezoelectric transducers and circuits. As a result, the channel count of the imaging system is increased and the capacitive loading due to cables is greatly reduced.

The first prototype chip of the transmitter and receiver analog front-end for a 1D CMUT array is fabricated and is under testing. The block diagram of the implemented chip is shown in Figure 2. It contains four channels of independent transmitters and receiver chains. External control can be implemented for beamforming and Time-Gain Compensation (TGC). In each channel, the transmitter generates high voltage electric pulses to drive the CMUT device. A 3-level pulse shaping transmitter is designed to increase the transmitted signal power within the transducer bandwidth. The design uses MOS high voltage transistors for a pulse magnitude as large as 32 Vpp. The pulse frequency is programmable between 1~10 MHz and the pulse duration is programmable between about 0.5~20 us.

On the receiver side, a Low Noise Amplifier (LNA) implemented with a trans-impedance amplifier interfaces to the CMUT. The LNA is optimized for noise, power, and bandwidth trade-offs. The LNA can also be switched from “on” and “off” within 5 us. This switching saves system power when LNA is not needed. A Variable Gain Amplifier (VGA) follows the LNA to realize the Time-Gain Compensation function. Instead of a linear TGC profile, this VGA implements the TGC in a low power way, with discrete gain steps to compensate signal attenuation with coarse resolution. The VGA consumes 300 uA, and the gain setting can be changed in 6 dB per step with a tunable range of about 54 dB.

The prototyped chip is 3 mm X 3 mm in size. The simulated performance shows that each receive channel consumes 18.1 mW in normal mode and 1.7 mW in sleep mode. The programmable Rx gain range is from 152 dB to 99 dB at 3 MHz, with gain steps of 6 dB per step. The Rx Bandwidth is 6.0 MHz and the Rx Noise Figure is 11.3 dB within the signal bandwidth. The Tx pulsing energy efficiency is 38.2 nJ / pulse for an assumed 60 pF load from one CMUT element.

1. O. Oralkan, “Acoustical imaging using capacitive micromachined ultrasonic transducer arrays: Devices, circuits, and systems,” Ph.D. dissertation, Stanford, Palo Alto, 2004.

To achieve comfortable form factors for wireless medical devices, battery size, and thus power consumption, must be curtailed.  Often the largest power consumption for wireless medical devices is in storing or transmitting acquired data.  Body area networks (BAN) can alleviate power demands by using low-power transmitters to send data “locally” around the body to receivers that are around areas of the body that allow for larger form factors, like the wrist or the waist.  These receivers, which have larger power budgets, can then process and store the data or send it elsewhere using higher power transmitters.

Body coupled communication (BCC) shows great potential for forming a BAN.  Two-node BCC works by forming two capacitive links between a transmitter and a receiver, creating a circuit loop.  One of these links is created by both the transmitter and receiver capacitively coupling to the body, effectively using the body as a low resistance channel between the respective capacitors.  The second link is created by both the transmitter and receiver coupling to the environment, or “earth ground,” and using it as a return path.  Larger BANs can be made by coupling additional nodes to both the body and the environment.

A model for implantable BCC has been developed in which the electrodes couple to the body and use the body as a lossy transmission line.  This model was tested by placing all four electrodes on the skin of a human, simulating how the electrodes would couple to body if they were implanted.  Signals were still able to be transmitted and received.

The BCC channel’s attenuation varies with body position and distance between receiver and transmitter.  Thus communication schemes that encode data with frequency or phase modulation work best.  Both FSK and PM binary signals have been successfully sent and received using both the traditional BCC and the implantable model.

1. T. G. Zimmerman, “Personal Area Networks (PAN): Near-field intrabody communication,” M.S. thesis, Massachusetts Institute of Technology, Cambridge, 1995.
2. S.-J. Song, N. Cho, S. Kim, J. Yoo, S. Choi, H-J Yoo ., “A 0.9V 2.6mW Body-Coupled Scalable PHY Transceiver for Body Sensor Applications,” ISSCC Dig. Tech. Papers, Feb. 2007, pp. 366-367.
3. A. Fazzi, S. Ouzounov, J. Homberg., “A 2.75mW wideband correlation-based transceiver for body-coupled communication,” ISSCC Dig. Tech. Papers, Feb. 2009, pp. 204-205.

We have designed an ultra-low-power 32-channel neural recording system in a 0.18-µm CMOS technology. The system consists of eight neural recording modules; each module contains four neural amplifiers, an analog multiplexer, an A/D converter, and a serial programming interface. Each amplifier can be programmed to record either spikes or LFPs with a programmable gain from 49-66 dB. To minimize the total power consumption, an adaptive-biasing scheme is utilized to adjust each amplifier’s input-referred noise to suit the background noise at the recording site. The amplifier’s input-referred noise can be adjusted from 11.2 µV_rms (total power of 5.4 µW) down to 5.4 µV_rms (total power of 20 µW) in the spike recording setting. The ADC in each recording module digitizes the signal from each amplifier at 8-bit precision with a sampling rate of 31.25 kS/s per channel and an average power consumption of 483 nW per channel. It achieves an ENOB of 7.65, resulting in a net efficiency of 77 fJ/State, making it one of the most energy-efficient designs for neural recording applications. The presented system was successfully tested in an in-vivo wireless recording experiment from a behaving primate with an average power dissipation per channel of 10.1 µW. The neural amplifier and the ADC occupy the areas of 0.03 mm² and 0.02 mm², respectively, making our design simultaneously area-efficient and power-efficient.

1. W. Wattanapanitch and R. Sarpeshkar, “A low-power 32-channel digitally-programmable neural recording system,”  IEEE Transaction on Biomedical Circuits and Systems, 2011, accepted for publication.