Electrocardiogram is a major source for doctors who need the information to diagnose patients health condition and pathology. ECG signals are usually small approximate 1mV and carry a lot of noise. This paper will introduce results in designing of electrocardiograph.ECG signals after the instrumentation amplifier will be filtered by a Band pass filter, The instrumentation amplifier is designed using IC AD624 AD of Analog Devices Corp. The work is a part of the framework of the project "Bluetooth enabled Bio signal Acquisition"
Key words: ECG, Instrumentation amplifier, Simulation, Agcl electrode
Introduction
ECG is a recording of the electrical activity on the body surface generated by the heart [1]. ECG measurement information is collected by skin electrodes placed at designated locations on the body. ECG signals are usually small and they may be corrupted by various kinds of noise such as power-line interference, electrode contact noise, motion artefacts [2]. So the measurement of ECG signal is a difficult task. The ECG signals captured by the electrodes are amplified by using the instrumentation amplifier. ECG signals after instrumentation amplifier may still be affected by noise. Therefore, the next step is to apply a low-pass filter for eliminating high frequency noise parts. A high-pass filter is used to get rid of DC (direct current) noise components
In order to record the ECG, we need a transducer capable of converting the ionic potentials generated within the body into electronic potentials which can be measured by conventional electronic instrumentation. Such a transducer consists of a pair of electrodes, which measure the ionic potential difference between their respective points of application on the body surface. Electrodes may be classified either as polarisable, in which case they behave as capacitors, or non−polarisable, in which case they behave as resistors. Common electrodes have characteristics that lie between these extremes; the silver−silver chloride electrode discussed below approximates more closely to a non−polarisable electrode
Silver-silver chloride electrode
Electrodes for recording bio potentials are composed of a metal usually silver and a salt of the metal (usually silver chloride). In addition, some form of electrode paste or jelly is applied between the electrode and the skin. The combination of the ionic electrode paste and the silver metal of the electrode form a local solution of the metal in the paste at the electrode−skin interface. Hence, some of the silver dissolves into solution producing Ag+ ions:
Ag= Ag+ + e-
Ionic equilibrium takes place when the electric field set up by the dissolving ions is balanced by the forces of the concentration gradient. At this point, there is a monomolecular layer of Ag+ ions at the surface of the electrode and a corresponding layer of Cl− ions adjacent to this. This combination is called the electrode double layer and there is a potential drop E across this layer, called the half-cell potential 0.8V in the case of the Ag−AgCl electrode []
C R1 E
Figure 1 Equivalent Circuit of Electrode Interface
The double layer of charges of opposite sign separated by a dielectric constitutes a form a capacitance, say C. However, since the Ag−AgCl electrode behaves mostly as a non−polarisable electrode, the main component of the impedance is resistive, R1. The series model in Figure needs to be modified to account for the fact that the impedance does not increase to infinity as the frequency tends to zero. This is done by adding a parallel resistance R2 as shown in Figure 2 which accounts for the electrochemical processes taking place at the electrode−electrolyte interface. The values of R1, R2 and C depend on the electrode area, surface condition, current density and the type and concentration of electrode paste used. Typical values are R1 = 2kï-, R2=10kï-ï€ and C=10ïF []
Figure 2 Modified Equivalent Circuit Diagram
Movement artefact
If the electrode is moved with respect to the electrolyte, this mechanically disturbs the distribution of charge at the interface and results in a momentary change of the half-cell potential until equilibrium can be re-established. If a pair of electrodes is in contact with an electrolyte and one move while the other remains stationary, a potential difference appears between the two during this motion. This potential is referred to as movement artefact and can be a serious cause of interference in the measurement of ECG
Overall equivalent circuit
Using the simple model of the electrode−electrolyte interface of Figure as well as the even simpler model previously developed for the electrical activity of the heart, we can put together an equivalent circuit which models the impedance seen by the input stage of an ECG system. This overall equivalent circuit is shown in Figure 3. Although C and C, R1 and R1, R2 and R2 may not be exactly equal, E should be equal to E. Hence
V
Figure3 Overall Equivalent Circuit
Figure 3 represents the actual difference of ionic potential between the two points on the body from which the ECG is being recorded.
Designing consideration for ECG Amplifiers
Biomedical parameters are difficult to measure due to their relatively low energies and similarity of signals from organ to organ. This similarity makes discrimination of a particular signal difficult. Not only does the transducer have a certain frequency response and sensitivity to detect a particular signal accurately, but the signal generated by the transducer requires processing (eg, impedance matching or filtering of noise) and amplification.
1) Frequency ranges and amplitude of Bio signal
Amplifier for a particular biological signal must be able to respond to the range of frequencies concerned Interference in the signal can be minimised if the bandwidth of the equipment used is selected to match the signal correctly
Range of frequencies:
EEG: 1-60 Hz
ECG: 0.5-100 Hz
Notice both include potential for 50 Hz power line interference
AHA recommend frequency response of 0.05-100 Hz for ECG analysis
Changes in the ECG waveform that occur:
Rapidly (QRS) contain higher frequency information
More slowly (ST segment) contain lower frequency information
First problem − electric field interference
The ECG voltage V is not the only signal found at the input of the amplifier; one major source of interference is the electrical power system. Capacitance between power lines in the wall, floor and ceiling and nearby equipment couples current into the patient, wires and machine. This current flows through the skin−electrode impedances on the way to ground. The capacitance to these power line sources varies with proximity but is of the order of 50pF which corresponds to an impedance of 64Mï-ï€ at 50Hz. If the right leg is connected to the common ground of the amplifier through an electrode with contact impedance of, say, 5kï-, the mains potential of 240V will appear as a 20mV noise input. This value is well in excess of the ECG signal itself. The key to extracting the desired ECG signal from the 50Hz noise is the fact that the ECG signal is the difference in potential between a pair of electrodes, ie a differential voltage. On the other hand, the 50Hz noise voltage is common to each electrode (it appears equally at both the Right Arm and Left Arm input terminals). Rejection of mains interference therefore depends on the use of a differential amplifier in the input stage of the
ECG machine, the amount of rejection depending on the ability of the amplifier to reject common mode voltages.
Second problem − magnetic induction
Current in magnetic fields induces voltage into the loop formed by the patient leads. The
induced voltage is proportional to the field strength and the coil area. Reducing this
interference requires that the field strength be reduced by moving the equipment and leads
(difficult to do in practice) or that the coil area be reduced by twisting the lead wires
together all along their length.
Third problem − source impedance unbalance
If there is a severe unbalance in the electrode−skin interface impedances (also known as the
Contact impedance), the bodys common−mode potential will be higher at one input than at
the other. Hence a fraction of the common−mode voltage will be seen as a differential
voltage and will be amplified by the differential gain of the amplifier
Hence the output voltage from the differential amplifier consists of 3 components:
The desired output due to amplification of the differential ECG signal.
An unwanted component of the common−mode signal due to the fact that the common−mode rejection is not infinite.
An unwanted component of the common−mode signal due to source impedance unbalance.
Differential Amplifier Design
The characteristics of the discrete differential amplifier may be improved by combining the three individual amplifiers in to a single IC, commonly known as an Instrumentation Amplifier. Instrumentation Amplifiers such as the AD620 and INA22 with input impedances in the range of vastly outperform the discrete component implementation. The AD620 draws a maximum supply current of 1.3mA as compared to a general purpose JFET operational amplifier such as the OP07 which draws 20mA per chip, resulting in a total supply current of 60mA for the configuration shown in Figure 9. Table 1 illustrates some key characteristics of three different instrumentation Amplifiers. Large scale VLSI integration ensures that the resistors within these Instrumentation Amplifier are matched to within 0.1% thus reducing common-mode gain (Equation 9).
ô€€€
Fig 1.1 AD 624 connection diagram
IA Type
AD 620
AMP 01
INA 114
CMR(Gain 100)
130 dB
125 dB
115 dB
Accuracy
0.15%
0.18%
0.18%
Noise
9nV/HZ
10 nV/Hz
11nV/Hz
Power Requirement
1.3 mA
4.8 mA
2.2 mA
Input current
0.5 nA
4 nA
2 nA
Table No.1
FIG:- Gain of AD 624 Ad
FIG
Its differential gain Ad, and common mode gain Ac, are given by the following equations:
The configuration utilizes three op-amps. The circuits optimal performance in terms of stability, differential gain and high common mode rejection ratio (CMRR) are difficult to compete with when using other discrete component implementations.
Accuracy in matching of the resistors is an important consideration in the design of a discrete component differential amplifier. The slightest error in the matching of resistances R3,top R3,bottom, and R4,top, R4,bottom in the discrete implementation (figure 9) results in unwanted common-mode amplification. Assuming we use 1% resistors:
or
with
Substituting the above results into equation 5 yields the unwanted common mode gain which is given by:
Where ε represents the percentage error present in the resistors.
Differential Amplifier Design
The characteristics of the discrete differential amplifier may be improved by combining the three individual amplifiers in to a single IC, commonly known as an Instrumentation Amplifier. Instrumentation Amplifiers such as the AD620 and INA22 with input impedances in the range of 1GΩ/pF, vastly outperform the discrete component implementation discussed in the previous section. The AD620 draws a maximum supply current of 1.3mA as compared to a general purpose JFET operational amplifier such as the OP07 which draws 20mA per chip, resulting in a total supply current of 60mA for the configuration shown in Figure 9. Table 3 illustrates some key characteristics of three different instrumentation Amplifiers. Large scale VLSI integration ensures that the resistors within these Instrumentation Amplifier are matched to within 0.1% thus reducing common-mode gain (Equation 9).
.
The AD620 requires only a single external gain-setting resistor; the gain equation is given by:
Figure illustrates the AD620 connection scheme utilized in the implementation. This method was utilized because two separate amplifiers needed to be interfaced to the right leg common-mode driver. Two 22k resistors are placed in parallel with the gain resistor. The right leg driver interface is at the centre of these two resistors. The two 22k resistor modify the gain equation. The gain of the AD620 is given by:
The AD620 has a maximum gain of ideally we would like to set the gain here to some maximum value thus eliminating the requirements for an additional gain stage prior to microcontroller stage. The gain selection of the AD620 is however critical to the design of the entire system. The gain must be set such that output saturation of the ±5V power supply does not occur. As mentioned above the maximum input is ±5 mV plus a variable normal-mode dc offset of up to ±300 mV. By setting the gain resistor to 6.2k it ensures that the maximum output swing is 3.07V i.e. Gain The gain was verified by grounding the negative input and passing waves of varying amplitude into the amplifier while observing both the input and output waves on an oscilloscope.
Anti-Aliasing Low-pass Filter
An anti-aliasing Low-Pass? filter is required to band limit the incoming ECG signal prior to digitization. Once the converted ECG signal is "Contaminated" with Aliased Noise, it takes time and memory within the processor to eliminate the noise. Furthermore it eliminates the effects of overdriven signals that usually occur beyond the bandwidth of the filter. The anti-aliasing low-pass analogue filter is placed immediately before the ADC.
The three most popular filter types are the Butterworth, Bessel and Tschebyscheff. Listed below are the key characteristics of these filters:
The Butterworth coefficients, optimize the pass-band for maximum flatness
The Tschebyscheff coefficients, sharpening the transition from pass-band into the stop-band
The Bessel coefficients, linearise the phase response up to fC.
Anti-Aliasing Low-pass Filter Design
Figure 12: Sallen key Low-pass Filter
The anti-aliasing filter cut-off frequency (fC) is set to the highest ECG frequency component of interest (fmax) so that fC = fmax. As mentioned previously a maximum frequency component of 250Hz is present in the representation of ECG signals. The filter was hence designed to have a cut-off frequency of 250Hz. The Sallen and Key active filter configuration was preferred to the Multiple Feedback due to its less component count. The Butterworth filter was selected due to its pass-band flatness. Since the ECG signal is so small in magnitude, any ripple in the pass-band would only further attenuate the signal making its recovery more difficult. A simple fourth order low-pass filter was designed for this purpose. Figure 12 shows a second order low-pass filter block. Two of these are cascaded to form the fourth order filter. The calculation of the component values along with the corresponding bode plot is present in the Appendices (A1).
