A 1.87 µW Capacitively Coupled Chopper Instrumentation Amplifier with a 0.36 mV Output Ripple and a 1.8 GΩ Input Impedance for Biomedical Recording ^†

Abstract

Chopper and capacitively coupled techniques are employed in instrumentation amplifiers to create capacitively coupled chopper instrumentation amplifiers (CCIAs) that obtain a high noise power efficiency. However, the CCIA has some disadvantages due to the chopper technique, namely chopper ripple and a low input impedance. The amplifier can easily saturate due to the chopper ripple of the CCIA, especially in extremely low noise problems. Therefore, ripple attenuation is required when designing CCIAs. To record biomedical information, a CCIA with a low power consumption and a low noise, low output ripple, and high input impedance (Z_in) is presented in this paper. By introducing a ripple attenuation loop (RAL) including the chopping offset amplifier and a low pass filter, the chopping ripple can be reduced to 0.36 mV. To increase the Z_in of the CCIA up to 1.8 GΩ, an impedance boost loop (IBL) is added. By using 180 nm CMOS technology, the 0.123 mm² CCIA consumes 1.87 µW at a supply voltage of 1 V. According to the simulation results using Cadance, the proposed CCIA architecture achieves a noise floor of 136 nV/√Hz, an input-referred noise (IRN) of 2.16 µV_rms, a closed-loop gain of 40 dB, a power supply rejection ratio (PSRR) of 108.6 dB, and a common-mode rejection ratio (CMRR) of 118.7 dB. The proposed CCIA is a helpful method for monitoring neural potentials.

1. Introduction

Wireless biomedical sensors (WBSs) are increasingly used to track our daily activities in order to detect cardiovascular diseases at an early stage [1,2,3,4]. Monitoring human biopotential requires the use of low-power sensors deployed in wearable or implantable systems. Researchers are currently developing brain–computer interfaces for numerous applications such as long-term monitoring, sports, rehabilitation, mobile monitoring, and improving the quality of life of patients [5,6]. WBSs typically use an instrumentation amplifier (IA) with low power consumption and low noise to connect with many types of biological sensors. The electrocardiograms (ECGs) of the heart and the electroencephalograms (EEGs) of the brain are examples of these biopotential signals. Neuroscience research and therapy can benefit from the use of biomarkers such as action potentials (APs) and local field potentials (LFPs) [7,8,9]. Biopotential signals often have an extremely small amplitude. For example, the amplitude range of an EEG is from 10 to 100 µV and that of an ECG is about 1 mV. The frequency range of the biopotential signals is 0.5–150 Hz [7,8]. The amplitude of the AP and LFP signals is about 100 µV to 1 mV, with a frequency range of 0.2 to 10 kHz for APs, and 1 to 200 Hz for LFPs [9]. Consequently, before signal processing is applied, these neural signals need to be amplified. A wearable biomedical sensor, constructed as shown in Figure 1, provides these neurological signals.

The first stage measures the amplitude of the small bio signal with a dry or wet electrode. An analog front-end, consisting of an IA at the first stage, a variable gain amplifier (VGA) at the middle stage, and an analog-to-digital converter (ADC) at the last stage, processes the neural signal in the second stage before it is transmitted by RF. The preamplifier of the analog front-end must obtain a high input impedance (Z_in) to reduce the DC input current that could cause tissue damage [8]. The IA must likewise demonstrate a low power, low noise, high-power supply rejection ratio (PSRR), and common-mode rejection ratio (CMRR) to eliminate noise from the power line and environmental factor, which may be important in some cases.

The chopping technique is frequently used in IA [10,11,12,13,14,15,16,17] to create an IA with a high PSRR, CMRR, and noise efficiency. The input capacitance and the switches in the chopper block generate the switched capacitance resistor, which is inversely proportional to the chopper frequency. This leads to a limitation of the Z_in of the amplifier if there are no impedance boosting techniques [18]. The ripple appears as a triangular wave affecting the quality of the signal of interest caused by a modulated intrinsic offset [19,20,21]. Furthermore, for long-term battery life suitable for WBS applications, the power consumption of the IC must be as small as possible, and the noise must also be low so as not to affect the signal quality at the output of the IC. Although a number of biomedical amplifiers with low power consumption have been published, it has not yet been possible to improve the output ripple or Z_in. For example, in 2020, the chopper amplifier in [22] consumed 3.24 µW at a 1.8 voltage supply, the Z_in just reached 440 MΩ, while the ripple suppression technique used an AC coupling capacitor, which caused this design to be affected by the noise folding [19]. In 2021, the current-reuse instrumentation amplifier [23] dissipated 5.94 µW at a voltage supply of 1.8 V to achieve a Z_in of 2.6 GΩ without any ripple suppression approaches. In 2024, although the amplifier [24] consumed only 2.47 µW from a 1.5 V supply, the output ripple and Z_in were not improved.

This paper presents a CCIA for biomedical information recording that is characterized by low noise, high input impedance, low output ripple, and low power consumption. At 1 V, the 0.123 mm² CCIA, which was simulated using a 180 nm CMOS process, consumes 1.87 µW. According to simulation results, it is shown that the output ripple being reduced to 0.36 mV is achieved with an RAL being switched on, and the Z_in of the CCIA increases up to 1.8 GΩ when the impedance boost loop (IBL) is active. When both the RAL and IBL are activated, the proposed CCIA obtains a closed-loop gain of 40 dB, an input referred noise (IRN) of 1.81 µV_rms, a thermal noise of 136 nV/√Hz, a common mode rejection ratio (CMRR) of 118.7 dB, and a power supply rejection ratio (PSRR) of 108.6 dB. Achieving a noise efficiency factor (NEF) of 6.8 and 7.5 with both RAL and IBL turned off and on, respectively, demonstrates that the CCIA records biomedical information successfully.

2. Design

The proposed CCIA for biomedical monitoring applications with a low output ripple and a high Z_in, as shown in Figure 2, consists of the main channel and three auxiliary loops such as a negative feedback loop (FBL), a ripple attenuation loop (RAL), and an impedance boosting loop (IBL) in order to solve the main problems of biopotential amplifiers. The transconductance input stage (G_m1) of the main path is a dual-folded cascode amplifier (DFC) with a biased current of 1.2 µA. In order to attain a working stability and a high swing, the G_m3 used a common source (CS) amplifier, combined with a Miller capacitor of 1.5 pF. A bias current of 1.8 µA is used for the channel and global common mode feedback (CMFB) from a V_DD of 1 V. The CCIA has a closed-loop gain of 40 dB, which is defined by the ratio of the input and negative feedback capacitors. In this design, the input capacitor C_in1,2 is set at 20 pF and the negative feedback capacitor C_fb1,2 is set at 0.2 pF. The PMOS pseudo-resistor R_b1,2 is used to bias DFC using the common mode voltage V_CM = 0.5 V. The capacitors (C_in1,2, C_fb1,2, C_m1,2, and C_LP1,2) are created using the MIM capacitor technique.

As shown in Figure 2, the chopper CH_I is employed to modulate the bio signal input V_bio at low frequency (as shown in Figure 2a) up to a signal at f_CH = 10 kHz, before reducing it with a negative feedback loop at the virtual ground. When V_bio,ω is at virtual ground (as shown in Figure 2b), it can be written as V_bio,ω = V_bio/(1 + βA_V), where A_V is the open-loop gain voltage of the CCIA, and β is the factor of the negative feedback loop based on the ratio of C_fb to C_in. The chopper CH_O converts the V_bio,ω to the essential frequency band after G_m1 has amplified it to produce V_1,Oω. Finally, G_m3 amplifies this signal to generate a bio signal amplification V_O,bio at the CCIA’s output (as shown in Figure 2d). The transfer function of the proposed CCIA can be expressed as follows:

$$H(s)=-\left(\raisebox{1ex}{$g_{\mathrm{m}1}{C}_{\mathrm{in}1,2}$}\!\left/ \!\raisebox{-1ex}{$\left(C_{\mathrm{in}1,2}+C_{\mathrm{fb}1,2}\right)C_{\mathrm{m}1,2}$}\right.\right)/\left(s+\raisebox{1ex}{$g_{\mathrm{m}1}{C}_{\mathrm{fb}1,2}$}\!\left/ \!\raisebox{-1ex}{$\left(C_{\mathrm{in}1,2}+C_{\mathrm{fb}1,2}\right)C_{\mathrm{m}1,2}$}\right.\right)$$

(1)

where g_m1 is the transconductance of the first stage; and C_in1,2, C_fb1,2, and C_m1,2 are the input, negative, and Miller capacitors, respectively. Unfortunately, G_m1 is attached to the offset voltage V_OS1 resulting from the process variation (as shown in Figure 2b). After this is amplified by G_m1 to create V_1,OS at the output of DFC, the V₁,_OS is also chopped to the chopper frequency before being integrated by the Miller integrator. This results in a considerable ripple at the output (as shown in Figure 2d). The amplitude of the output ripple can be described as follows:

$$V_{\mathrm{O},\mathrm{Ripple}}=\frac{V_{\mathrm{OS}1}{g}_{\mathrm{m}1}}{2f_{\mathrm{CH}}{C}_{\mathrm{m}1,2}}$$

(2)

where g_m1 is the transconductance of the first stage; f_CH is the chopping frequency; and C_m1,2 are the phase margin compensation capacitors. For example, V_OS1 = 10 mV, g_m1 = 0.7 µS, f_CH = 10 kHz, C_m1,2 = 1.5 pF, and V_O,Ripple = 233 mV.

The block diagram of an RAL is also shown in Figure 2. Instead of capturing the signal at the output, as in the usual approach, the RAL uses a low-pass filter (LPF) to obtain the signal at the output of the DFC (V_1,O = V_1,OS + V_1,Oω) (as shown in Figure 2c) before the chopper output CH_O. This is because the V_1,OS signal is continuously amplified, while the AC signal V_1,Oω is filtered out by the LPF in this case. To ensure that no AC signal is applied to G_m1b, which has the schemactic shown in next section, the capacitor C_LP2 is added to the output of the RAL, although the LPF has a small low-pass corner controlled by R_LP1,2 and C_LP1. After amplifying V_1,OS, the signal V_O,RAL is connected to G_m1b, creating a negative feedback loop to compensate for V_1,OS (as shown in Figure 2e). This means that the ripple caused by V_OS1 is reduced at the output of the CCIA (as shown in Figure 2f). To increase the loop gain (L_G) of the RAL and achieve a high ripple attenuation factor (RAF), G_m2 is implemented using a two-stage operational amplifier for low noise and low power consumption. We assume that V_OS2, another inherent offset caused by process variations, is similarly associated with G_m2. The modulated offset V_OS2 also generates the ripple at the CCIA’s output and has the same effects as V_OS1, so it needs to be reduced. The ripple at the output of the CCIA is mitigated by a DC loop gain’s factor L_G(0) of the feedback loop RAL. The equation to determine L_G(s) in the technique proposed in this study is as follows:

$$L_G(s)\cong g_{\mathrm{m}1\mathrm{b}}{R}_{\mathrm{LP}}{A}_{\mathrm{vGm}2}=g_{\mathrm{m}1\mathrm{b}}{R}_{\mathrm{LP}}\frac{A_{\mathrm{vGm}2,\mathrm{DC}}}{1+s/{\omega }_{\mathrm{p}}}$$

(3)

$$L_G(0)\cong G_{\mathrm{m}1}{R}_{\mathrm{LP}}{A}_{\mathrm{vGm}2,\mathrm{DC}}$$

(4)

where G_m1b is the auxiliary transconductance of the first stage G_m1; and A_vGm2,DC and f_p (ω_p = 2πf_p) are the DC gain and cut-off frequency of the two-stage amplifier G_m2.

In the chopper biopotential amplifier, the input capacitor and the chopper are combined together, creating a switched capacitor resistor. At the completion of a cycle through the chopper clock f_CH, a charge of Q = 2C_inV_in is delivered [7]. Therefore, Z_in can be determined as Z_in = 1/(2C_inf_CH). For example, Z_in is 2.5 MΩ for biomedical recording applications when the input capacitor C_in = 20 pF and f_CH = 10 kHz. An impedance boost loop (IBL) with a time diagram, as shown in Figure 3, is used to pre-charge Q to the C_in, as the Z_in must be improved by minimizing the charge Q from the input signal V_in. When the IBL is connected to C_in, the connection flowing from the input is interrupted and the V_in is copied by the buffer in IBL and is pre-charged to C_in. Assuming that the pre-charge current from the buffer is high enough, C_in will be fully charged from IBL; thus, when C_in is connected to the input after pre-charging, C_in does not require a charge from the input signal V_in. This results in the fact that the Z_in can be set indefinitely. The Z_in can be represented following the analysis in [7], as follows:

$$Z_{\mathrm{in}}=Z_0/\left(\alpha +\exp \left(\raisebox{1ex}{$-T$}\!\left/ \!\raisebox{-1ex}{$\tau $}\right.\right)\right)$$

(5)

where Z_in is the input impedance, Z₀ is the input impedance without any boosting technique, α is the buffer gain error, T is the pre-charge time, and τ is the actual time constant.

3. Circuit Implementation

Figure 4 shows the dual cascode amplifier (G_m1) combining with the RAL block. A feedback loop is set up comprising G_m1b, a two-stage chopper amplifier, and the RC-LPF in order to mitigate the output ripple. Figure 4 shows that the input stage consumes a power of 1.2 µW from a supply voltage of 1 V. The global common mode feedback circuit (CMFB) [25], which is used and consumed a biased current of 200 nA from 1 V, is employed to control the DC voltage at the output node of the CCIA. The gates of the transistors M₉ and M₁₀ are adjusted using the CMFB circuit (V_CMFB) to control the output DC voltage followed to V_CM = V_DD/2. The power consumption of 1.4 µW of G_m1 including CMFB is used.

Figure 5 shows the architecture of the chopper two-stage amplifier integrating with a common mode feedback circuit (CMFBR) [25] for G_m2. As already mentioned, the inherent offset V_OS2 of G_m2 has the same effect as V_OS1; the offset that is also upmodulated creates the ripple at the output of the CCIA. Consequently, it must be eliminated. In RAL, the chopper CH_R,I is located at the output of the LPF, while the chopper CH_R,O is put between the first and second stage of G_m2. This causes V_OS2 to be modulated up to a high frequency and then modulated down by the chopper CH_O, resulting in an offset voltage in the front of G_m3. Therefore, V_OS2 is considered as an offset at the input of G_m3. Due to the high gain level of G_m1 (about 80 dB), this offset is insignificant compared to the input. Therefore, its influence can be neglected. By using a 1 V supply, the first stage of G_m2 consumes 5 nA, while the second stage of G_m2 is biased to 20 nA for a better swing. The voltage V_CMFBR is generated by the CMFB circuit, which uses a bias current of 5 nA. Therefore, the total power of the RAL is only 30 nW.

The DC gain and the cut-off frequency of G_m2 are decisive factors for determining the ripple attenuation factor and the bandwidth of the loop gain. The Monte Carlo simulation (MCS) method is used to study the fluctuation of these parameters across the chip and the mismatch of the device, including global variation and local mismatch. Figure 6a and Figure 6b show the value of the DC gain and cut-off frequency of G_m2, respectively. These distributions were derived from 300 samples of the MCS. The results show a mean value (MV) of the DC gain of G_m2 of 90.9 dB, with a standard deviation (Std) of 0.167 dB. Furthermore, the MV of the cut-off frequency is 0.042 Hz with a Std of 0.0044 Hz.

Figure 7 shows the schematic of the G_mb in the IBL. The MP0 and MP1 in G_mb are employed to bias and regulate the buffer with the purpose of reducing the impact of process variation on mismatch devices. The boosted input impedance in IBL can be affected by the pre-charge time T, the switches sizing of Φ_1,2, and the buffer design. Figure 8 shows the clock generator for controlling the IBL. The pre-charge time T and C_lk1,2 are generated by dividing the clock signal f_CH, which can be converted into a delayed signal by the use of several MOS capacitors and inverters. Although using a bias current of 700 nA, IBL is only enabled in 6 μs in each cycle of 100 μs (f_CH = 10 kHz); thus, IBL (G_mb) consumes only the current of 700 nA × 6/100 = 42 nA. The size and number of the MOS capacitor are shown in Figure 8. The sizing W/L of the switches SW_1,2 is set at about 0.5 µm/0.25 µm in order to minimize the inherent resistance. Furthermore, the pre-charge time T can be altered by using a two-switch SW_1,2 in order to mitigate the effect of process variability. This research examines the effect of the pre-charge time T and the size of the switches Φ_1,2. Lengthening the pre-charge time T enhances the Z_in of the device, while simultaneously amplifying the noise of the device. The size of switches Φ_1,2 in Figure 3 is another factor that affects the boost in Z_in. A small W/L size can result in a substantial voltage loss between these switches. Figure 9a and Figure 9b show the relationship between the Z_in, the input referred noise, the pre-charge time, and the switch sizing, respectively. As can be seen in Figure 9a, the Z_in improves from 2.5 MΩ to 0.7–1.8 GΩ when IBL is enabled with the pre-charge time T and is increased from 3 to 6 µs; however, the IRN increases sharply from 1.7 to 2.2 µV_rms. As shown in Figure 9b, when the switches sizing (Width-W) of Φ_1,2 increases, the parasitic capacitors of these switches are also increased. When clock control for the pre-charge phase is applied to the gate of the CMOS transistor switches, the charge injection noise and clock feed-through increases [26,27], leading to an increase in the IRN. In this work, when the width of the switches sizing of Φ_1,2 is changed from 1 to 5 µm, the IRN is increased from 1.8 to 2.1 µV_rms, while the noise increases from 2.5 to 3.2 µV_rms when W of Φ_1,2 changes from 6 to 10 µm. According to the simulated results, as shown in Figure 9, in order to optimize Z_in and IRN, the pre-charge time T and the switches sizing Φ_1,2 in this design are therefore set to 6 µs and W/L is set to 5 µm/0.5 µm.

4. Simulation Results

Figure 10 shows the layout of the proposed CCIA in the 180 nm CMOS process. The chip area of the CCIA configuration is 0.123 mm². Table 1 shows the power dissipation of each block in the CCIA. The total power consumption of the proposed CCIA is 1.87 µW from a V_DD of 1 V. G_m1, G_m2, G_m3, and G_mb consume 1400, 30, 400, and 40 nW, corresponding to 74.8%, 1.6%, 21.36%, and 2.24% of the total power consumption, respectively. According to the post-simulation results, it is shown that Figure 11 shows the simulated results of the CCIA’s transfer function—(a) transient; (b) MCs. The CCIA’s Av is 40 dB. The MCS results present that the MV of the closed-loop CCIA gain at 300 samples is 38.9 dB, with a Std of 0.28 dB. Figure 12 shows the MCS results for PSRR and CMRR after running 300 samples. At the 1 V supply, the MV of PSRR and CMRR are 108.6 and 118.7 dB with Stds of 23.7 and 24.4 dB, respectively.

During the simulation, the input of the CCIA is configured so that it is short-circuited in order to test the output spectrum. Both V_OS1 and V_OS2 were set to a voltage of 5 mV. Figure 13 shows the simulated results of the voltage spectrum and the MCS of the output ripple. When the RAL is disabled, the amplitude of the output spectrum at the chopping frequency is about 82.2 mV, as shown in Figure 13a. The amplitude of the output ripple of the CCIA decreases to 0.36 mV when the RAL is enabled, as shown in Figure 13b. The output ripple is verified using an MCS that includes 300 samples and accounts for both local and global process variations. When the RAL level is changed from off to on, the MV of the output ripple decreases from 82.9 mV to 0.36 mV with a Std of 42.6 mV or 93 µV, as shown in Figure 13c,d. The simulation results shown in Figure 14 therefore give an MV of 45.9 for the ripple attenuation, with a Std of 3.68 dB. The proposed feedback network effectively compensates for the offset voltage (V_OS1, V_OS2) caused by mismatches due to process, voltage, and temperature variations, resulting in a significant reduction in the output ripple voltage.

Figure 15 shows the effects of activating the impedance boost loop on the Z_in and the input-related noise. By setting the SW_1,2 parameter, it is possible to achieve different results for the input impedance. When SW_1,2 is set to 01 and 11, corresponding to a pre charge time of 3 and 6 µs, the Z_in in the low frequency range increases to about 0.7 and 1.8 GΩ, respectively, as shown in Figure 15a. Without the presence of IBL, the noise floor is about 119 nV/√Hz, while the 1/f corner frequency is 10 Hz. When IBL is enabled, the noise floor increases to 136 nV/√Hz, resulting in an IRN over a bandwidth of 1 to 200 Hz of 2.16 µV_rms. This increase is observed for different values of SW_1,2, which determines the pulse width of the pre-charge time. Figure 16a shows the variation of IRN for the proposed amplifier across several process corners, ranging from 1.9 to 2.6 µV_rms. On the other hand, Figure 16b shows the MV of IRN, which is 2.16 µV_rms, with a Std of 97.9 nV_rms, verified using an MCS with 300 samples, considering both local and global process variations.

Table 2 shows a brief summary of the primary design specifications, encompassing power consumption, output ripple’s amplitude, Z_in, CMRR, PSRR, and NEF (noise efficient factor). There are several references that show simulation results, such as [24,28,29], in order to guarantee an equitable comparison. Table 2 is employed to evaluate the performance of the proposed design in comparison with the current state-of-the-art studies. The proposed CCIA obtains an NEF of about 7.5 by integrating RAL and IBL. Additionally, it exhibits a minimal output ripple of 363 µV and a significant Z_in of 1.8 GΩ. The CCIA’s dissipation is 1.87 µW from a 1 V supply.

5. Conclusions

The paper presents a 1.87 µW capacitively coupled chopper instrumentation amplifier for biomedical recording. The output ripple is measured at 0.36 mV, and the input impedance is 1.8 GΩ. The CCIA chip occupies a chip area of only 0.123 mm² when implemented in a 0.18 µm CMOS technology. The ripple attenuation loop effectively decreases the output ripple of the proposed CCIA down to 0.36 mV. The CCIA is able to achieve a high input impedance of approximately 1.8 GΩ due to the impedance boosting loop. The low-power chopper amplifier has a power dissipation of 1.87 µW at a V_DD of 1 V. It also obtains a closed-loop gain of 40 dB, a PSRR of 108.6 dB, and a CMRR of 118.7 dB. The noise floor of the CCIA has a magnitude of 136 nV/√Hz, which leads to an IRN of 2.16 µV_rms across a bandwidth of 200 Hz. Thus, an NEF value of 7.5 is attained. This illustrates our ability to evaluate the performance of the proposed CCIA by comparing it to the most recent studies.