sensitive to the condition of the skin and are highly susceptible to motion artifacts. An easier to use and less obtrusive technology is called for to match the advancements made in wireless body sensor networks.
In contrast to wet and dry contact sensors, non-contact capacitive electrodes do not require an ohmic connection to the body. For body sensor applications, this offers numerous advantages since non-contact electrodes require zero preparation, are completely insensitive to skin conditions and can be embedded within comfortable layers of fabric. While the concept of non-contact biopotential sensors is not new, with the first working device reported decades ago, a practical device for patient use has yet to be realized. More recently, several authors have presented results from designs utilizing the latest in commercially available discrete low noise amplifiers, including some wireless designs. In all cases, the challenges in noncontact sensing have lead to many clever, and often-times, proprietary circuit designs in an effort stabilize the electrode’s input.
In this paper, we expand on the work previously presented by building a sensor with much improved noise performance. In addition, the full design and schematics for a wireless, non-contact EEG/ECG system with features designed for specifically for practical body sensor networks is described.
System design
The system contains a set of non-contact biopotential electrodes connected along a single common wire. The sensors can be either in direct contact with the skin or embedded within fabric and clothing. A small base unit powers the entire system and contains a wireless transmitter to send data to a computer or other external device. Near the base unit, a single adhesive or dry contact sensor placed anywhere convenient is used to establish the ground reference for the system.
А. Electrode Construction
Each electrode is constructed from two, US quarter sized, PCBs. The upper PCB contains a low noise differential amplifier and a 16-bit ADC. Rather than outputting a single analog signal, the electrode outputs the digitized value, which can be carried in serial daisy chain to drastically reduce the number of wires needed. A miniature 10-wire ribbon cable carries the power supply, digital control as well as analog common mode reference from electrode-to-electrode.
The lower PCB contains the INA116 configured as an ultra-high input impedance amplifier. The bottom surface of the PCB is a solid copper fill, insulated by solder mask, that functions as the electrode. This surface forms a coupling capacitor with the body. An active shield formed by in a solid inner plane protects the
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electrode from external noise pick-up. To minimize the shield capacitance, an extra thick PCB is used for the electrode.
B. Front End Amplifier
Designing an ultra-high input impedance amplifier with low noise levels is the main challenge in implementing non-contact electrodes. Signal sources from the body (EEG/ECG) can be thought of as a voltage source, Vs, connected to the input of an amplifier via a small coupling capacitance, Cs. All real amplifiers will also have some finite resistance, Rb, and input capacitance. A small amount of positive feedback can be applied through, Cn, to neutralize the effect of the input capacitance for better channel matching and CMRR.
Important noise sources include the input referred voltage noise of the amplifier, Vna, the input current noise, Ina and the additional current noise, In b, due to the leakage and conductance of the biasing element. The current noise contribution will either 4kTR thermal noise for a resistive device or 2qI shot noise for a PN junction. Bootstrapping can be used to electronically boost the effective impedance of the biasing element, but the noise contribution depends only on the physical resistance or leakage current, illustrating the challenge in finding suitable components for a non-contact sensor.
The equation clearly shows the effect of the parasitic input capacitances and leakage currents on the noise performance of the amplifier and the difficulty in designing a non-contact electrode. Any excess input capacitance will directly multiply the effect of the amplifier’s input voltage noise as Cin+Cn> Cs.
Furthermore, since biopotential signals are at low frequencies (.1-100 Hz), even small amounts of current noise become integrated into large amounts of input voltage noise. This necessitates an amplifier with both very low input and guard capacitance as well as almost zero leakage currents.
The INA116 by Burr-Brown is an amplifier that is well-known for ultra-high input impedance applications by virtue of its extremely low current noise. However, any circuit introduced to bias the inputs will significantly degrade the noise performance of the amplifier. An extremely difficult to obtain resistor would be required to match the current noise specification of the INA116. Fortunately, it was found during experiments that the INA116 would reliably charge a floating input to a point inside the allowable input range shortly after power-up purely through leakage currents, removing the need for any external bias network. To remove drift and DC offsets, a low-passed version of the input signal was taken from the noninverting input’s guard and connected to the inverting input. This effectively performs
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AC coupling without degrading the input impedance and centers the output to mid-rail for maximum signal swing.
It is also worth noting that a similar ’bias-free’ technique can also be applied successfully to rail-to-rail input/output operational amplifiers configured in unity gain by simply leaving the non-inverting input floating and AC coupling the output to the next stage. As long as DC measurements are not needed (as in EEG/ECG applications), the amplifier is guaranteed to operate somewhere inside the supply rails. We have successfully tested this with the LMC6081 and LMP7702 operational amplifiers. The overall performance is comparable to the INA116 circuit. More detailed characterization of the design will be explored in a future paper.
To ensure that the electrode’s gain is constant over a wide range of coupling distances, a small amount of positive feedback, adjusted byR3, is applied back to the input throughC2. Each electrode is carefully calibrated at a test-bench byadjustingR3until the gain is constant for different coupling distances. In practice, however, the capacitance neutralization circuit is not needed since the system is wireless with a floating ground – no significant 50/60 Hz mains interference is observed even with mismatched electrodes.
C. Differential Amplifier and ADC
The output of the INA116 is coupled to a differential gain amplifier through an additional high-pass filter with a cutoff of :1 Hz to remove the relatively high DC offset of the INA116. The LTC6078 micro power operational amplifier was chosen for it’s excellent low noise and low offset characteristics. A simple noninverting differential gain stage of 40:1 dB was implemented by connecting the electrodes together through the node Vcm, which is carried in the daisy chain. This constructs what is essentially a multi-channel instrumentation amplifier to remove the common-mode noise while amplifying the local biopotential signal.
Here it is shown the full measured transfer function of the front-end and differential amplifier. A gain of 46 dB and cut off frequencies of 0:7 Hz and 100 Hz were obtained as expected.
The total in-band input referred noise was measured to be 3.8µ VRMS. At present, it appears that the noise pickup from external sources is as problematic, if not more so, as the intrinsic noise sources in the circuit, even with the active shield layer. This is not surprising due to the sensitivity of the ultra-high impedance input node. Future versions of the electrode will incorporate more comprehensive shielding strategies than a simple inner PCB plane.
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Conclusion
We present a wireless body sensor network for high quality EEG/ECG recordings utilizing non-contact electrodes. The full schematics for building the simple, low noise capacitive electrode are presented. Future work will focus on miniaturizing and better packaging the electrode as well as reducing the power consumption of the digital and wireless transmitter components.
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Тема 2. Кардиограф
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Blood Pressure Monitor. Fundamentals and Design
This application note demonstrates the implementation of a basic blood pressure monitor using Freescale products. The blood pressure monitor can be implemented using any of the Freescale medical oriented MCUs: Kinetis MK53N512 and Flexis MM members MC9S08MM128 and MCF51MM256 embedding a 16bit ADC, 12-bit DAC, 2 Programmable-Gain Op-Amps, 2 TRIAMPS, Analog Comparators, and Vref generator. The K50 family can also perform DSP instructions for signal treatment and MCF51MM can perform multiply and accumulate (MAC) instructions. This document is intended to be used by biomedical engineers, medical equipment developers, or any person related with the practice of medicine and interested in understanding the operation of blood pressure monitors. Nevertheless, it is necessary to know fundamentals of electronic, analog, and digital circuits.
Arterial pressure
Arterial pressure is defined as the hydrostatic pressure exerted by the blood over the arteries as a result of the heart left ventricle contraction. Systolic arterial pressure is the higher blood pressure reached by the arteries during systole (ventricular contraction), and diastolic arterial pressure is the lowest blood pressure reached during diastole (ventricular relaxation). In a healthy young adult at rest, systolic arterial pressure is around 110 mmHg and diastolic arterial pressure is around 70 mmHg. Blood flow is the blood volume that flows through any tissue in a determined period of time (typically represented as ml/min) in order to bring tissue oxygen and nutrients transported in blood. Blood flow is directly affected by the blood pressure as blood flows from the area with more pressure to the area with less pressure. Greater the pressure difference, higher is the blood flow. Blood is pumped from the left ventricle of the heart out to the aorta where it reaches its higher pressure levels. Blood pressure falls as blood moves away from the left ventricle until it reaches 0 mm Hg, when it returns to the heart’s right atrium.
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