Publications

D. D. He, I. A. Nausieda, K. K. Ryu, A. I. Akinwande, V. Bulovic, C. G. Sodini, “An Integrated Organic Circuit Array for Flexible Large-area Temperature Sensing,” International Solid-State Circuits Conference, 7-11 Feb. 2010.

K. Ryu, I. Nausieda, D. He, A. Akinwande, V. Bulovic, C. G. Sodini, “Bias Stress Effect in Pentacene Organic Thin-Film Transistors,” IEEE Transactions on Electron Devices, vol. 57, pp. 1003-1008, May 2010.

K. M. Nguyen, A. Accardi, H. Kim*, G. W. Wornell, C. G. Sodini “Digital Phase Tightening for Millimeter-wave Imaging,” 2010 Custom Integrated Circuits Conference, Sept. 2010.

D. D. He, E. S. Winokur, T. Heldt, and C. G. Sodini, “The Ear as a Location for Wearable Vital Signs Monitoring,” 2010 Engineering in Medicine and Biology Conference, Sept. 2010

I. Nausieda, K. Ryu, D. He, A. Akinwande, V. Bulovic, C. G. Sodini, “Dual Threshold Voltage Organic Thin-Film Transistor Technology,” IEEE Transactions on Electron Devices, Nov. 2010.

Optically transparent, wide band gap metal oxide semiconductors are a promising candidate for large area flexible electronics. Because most commercially available flexible substrates, particularly polymer substrates, cannot withstand the high temperature processing (>400°C) required for traditional silicon device fabrication, the development of new materials and devices that can be processed at low temperatures in a scalable manner is needed. Metal oxide semiconductors have been demonstrated to retain high carrier mobilities even in the disordered, amorphous state obtained when processed at near-room temperatures ^{[1] [2]} . Compared to amorphous silicon field effect transistors (FETs), which are the dominant technology used in display backplanes, metal-oxide-based FETs have been demonstrated with higher charge carrier mobilities, higher current densities, and faster response performance ^{[3] [4]} .

It has been shown both in simulation and by experiment that FET threshold voltage (V_T) can be modified simply by changing the channel layer thickness, without requiring the additional complexity of multiple channel materials or different dopings. In this project we have developed a low temperature (~100°C), scalable lithographic process for top-gate, bottom-contact amorphous metal oxide-based FETs using parylene, a room temperature-deposited CVD polymer, as gate dielectric. Figure 1 shows a micrograph of an array of FETs fabricated with different channel lengths. The baseline process was extended to enable the integration of FETs with different threshold voltages on the same substrate. The availability of FETs with different threshold voltages enables the implementation of enhancement/depletion (E/D) logic circuits that have faster speeds and smaller device areas than single-V_T topologies. Using the two-V_T lithographic process, we fabricated and characterized integrated E/D inverters and ring oscillators that operate rail-to-rail at supply voltages as low as V_DD = 3V. An example inverter characteristic is plotted in Figure 2. These results demonstrate the potential for low V_DD metal oxide-based integrated circuits fabricated in a low temperature budget, fully lithographic process for large area electronics.

1. K. Nomura, H. Ohta, A. Takagi, T. Kamiya, M. Hirano, and H. Hosono, “Room-temperature fabrication of transparent flexible thin-film transistors using amorphous oxide semiconductors,” Nature, vol. 432, pp. 488-492, Nov. 2004.
2. J. Robertson, “Disorder and instability processes in amorphous conducting oxides,” Physica Status Solidi B-Basic Solid State Physics, vol. 245, pp. 1026-1032, June 2008.
3. R. L. Hoffman, B. J. Norris, and J. F. Wager, “ZnO-based transparent thin-film transistors,” Applied Physics Letters, vol. 82, pp. 733-735, Feb. 2003.
4. E. Fortunato, P. Barquinha, G. Goncalves, L. Pereira, and R. Martins, “High mobility and low threshold voltage transparent thin film transistors based on amorphous indium zinc oxide semiconductors,” Solid-State Electronics, vol. 52, pp. 443-448, Mar. 2008.

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.

Figure 1

Most current ultrasound imaging systems use piezoelectric materials for the ultrasound transducer. The recent development of micro-electromechanical systems (MEMS) allowed fabrication of capacitive micromachined ultrasound transducers (CMUTs).  A CMUT is a micromachined capacitor whose value changes according to the DC bias voltage or external pressure due to the physical deformation of the top plate by electrostatic force or external pressure. The major advantages of this transducer technology are the potential for integration with supporting electronic circuits, ease of fabrication, higher resolution due to small transducer size, and improved bandwidth and sensitivity ^[1] .

This project focuses on the front-end design of portable ultrasound systems using CMUTs. Figure 1 presents a conceptual block diagram of the system. Implementing an ADC at each channel input makes possible digital beam-forming in the receive (Rx) path, which enhances ultrasound image quality. To implement as many ADCs as the number of transducer channels, each ADC must consume as little power as possible, and each should be implemented in a small area. Considering the required performance, a zero-crossing-based (ZCB) pipelined ADC is a suitable architecture ^[2] .  For the first part of this project, a 50-MHz 12-bit ZCB pipelined ADC is designed. The highly digital implementation characteristic of the zero-crossing detection technique enables energy-efficient operation and voltage scaling. Supply voltage scaling based on the required sampling frequency and resolution provides constant energy efficiency over a wide range of sampling frequencies and resolutions.

Recently, a few 2D imaging systems using CMUT as ultrasound transducers have been reported, but they do not use real-time imaging ^[1] . The digital image processing block will be considered in the system level for real-time imaging.  After completing the 2D ultrasound image system using a 1D transducer, we will examine the feasibility of the 3D ultrasound image system using 2D transducers.

1. O. Oralkan, “Acoustical imaging using capacitive micromachined ultrasonic transducer arrays: Devices, circuits, and systems,” Ph.D. thesis, Stanford University, Palo Alto, CA, 2004.
2. L. Brooks and H.-S. Lee, “A zero-crossing-based 8b 200MS/s pipelined ADC,” IEEE International Solid-State Circuits Conference, 2007. Digest of Technical Papers, pp. 460-615.