Chronic venous insufficiency of the lower limbs is a disease which is caused by an increased blood pressure inside the veins of the leg and the resulting increase of the contained blood volume. This work focuses on developing a device which uses impedance plethysmography, also known as impedance phlebography, to obtain information about the blood volume in the lower leg and provides the possibility to measure the impedance semi contact-less, e.g. through compression stockings. Furthermore a real-time beat-to-beat interval detection algorithm was implemented. Finally, the function of the developed impedance measuring system and the whole system was verified by comparing it with a gold standard. In comparison to the conductive coupling, the system performed similarly. The analysis showed that the developed system is suitable for semi-capacitive IPG. The algorithm was implemented conservatively since it provided a good false-positive rate of 0 %, but only a moderate sensitivity of about 68 %. Reliable and continuous measurement of the pulse signal was only possible in periods of immobility.
- chronic venous insufficiency
- impedance plethysmography
- capacitive measurement
- real-time detection algorithm
One of the most common diseases is the chronic venous insufficiency (CVI) of the lower extremities, which is caused by a permanent pressure overload of the vein. The venous tree is defective, incapable of removing sufficient amounts of blood from the peripheral veins. This causes pooling of blood and intravascular fluid at the lower gravitational parts of the body - the legs.
Most of the blood in the legs is removed by the contraction of the calf muscle through which pressure is applied to the deep veins, so called muscle pump. If the calf muscles are not used over an extended period of time, the amount of blood in the leg veins increases and a harmful pressure overload occurs. The consequences are, for example, the accumulation of edema or damage of the venous valves, which leads to an increase of the blood pressure, causing additional valves to fail. The most common treatment for CVI is the compression therapy. The compression is usually used by compression stockings where pressure is applied to the deep veins in the legs.
A diagnostic method for detecting CVI is impedance phlebography (IPG), which is a non-invasive medical method that measures small changes of the electrical resistance of the calf. This measurement method was first described by Nyboer  and aims at quantitatively determining blood volume changes of a tissue section. Compared to photoplethysmography, this method allows a measurement of the blood volume variation independently of the arterial blood circulation, which can vary due to physiological processes, for example due to ambient temperature . A disadvantage of the IPG is the technical equipment. In IPG, a high-frequent current (typically between 5 kHz and 100 kHz) is injected into the tissue by two electrodes. Two other electrodes are used to measure the corresponding voltage drop, which depends on the tissue composition. The impedance equals the quotient of the measured voltage and injected current.
The relationship between the measured impedance and changes of blood volume in the measured segment as described by Nyboer  is given by:
Beside the usual conductive coupling of a wearable bioelectrical impedance measurement, textile integrated impedance cardiography or spectroscopy for personal health care applications is currently under research [6, 7, 8]. An alternative contactless method to measure the impedance of the thoracic region is the use of wearable magnetic induction systems..
The aim of this work is to develop a personal healthcare device that can be used while wearing compression stockings and that provides information from impedance plethysmography. The device should be wearable, unobtrusive, should measure the beat-to-beat pulse rate and warn the user when he should use his muscle pump.
The proposed semi-capacitive impedance phlebography should apply an alternating current by commercial conductive electrodes, but should provide a contact-less measurement through an isolating material of the corresponding voltage drop. In Figure 1 the systematic concept of the semi-capacitive impedance phlebography device is shown.
It uses a tetra-polar arrangement of electrodes. The four-electrode technique avoids electrode polarization and minimizes the influence of the skin impedance [10, 11]. The outer pair of conductive electrodes passes a small alternating current
The system samples the corresponding impedance with the sampling-rate
Two specifications are essential to build the device: the sample rate and the resolution. The minimal sample rate can be estimated based on the results of Soares Filho et al. . They examined the time derivative of Δ
The resolution of the measurement device is an issue since the basal impedance
Since the device should measure through an isolating material, a capacitive measurement system was developed. The capacitive measuring was done by active electrodes as an impedance converter, which enables a low-impedance processing of the measured potential
In order to ensure a capacitive coupling in the case of a conductive connection between the electrode surface and the tissue, the capacitor
Eq. (2) describes the Laplace transfer function of the active electrode under the assumption of an ideal operational amplifier.
The active electrodes show high-pass behavior, with a cut-off frequency of
On the one hand, this behavior requires a minimum size of
This is particularly useful when the operational amplifier is not only connected as an impedance converter, but also as an amplifier. In this case, the amplified common mode noise can drive the operational amplifier into saturation quickly. Nevertheless, by using the high-pass behavior, the attenuation of the lower frequencies is dependent on the coupling capacitance
The schematics are shown in Figure 3.
In this illustration
To limit the voltage measurement to
The impedance measuring system consisted mainly of the AFE4300 chip (Texas Instruments, Dallas, USA) and active electrodes for capacitive voltage measurement, as mentioned above. To increase the resolution of the impedance measurement the signal was directly amplified at the active electrodes. An additional amplification was done by a differential amplifier between the electrodes and the AFE4300 chip.
The circuit is shown in Figure 5.
The operational amplifier (TLC252, Texas Instruments, Dallas, USA) and AFE4300 need a supply voltage of 3
Additionally, the oversampling rate was increased by an increase of the clock frequency
The aim of the signal processing was the extraction and separation of the beat-to-beat heart interval. The signal processing was implemented on a MSP430F2618 microcontroller. First, the inverted measured signal was filtered with a bandpass filter to separate the physiological frequency components. A real-time beat-to-beat interval detection algorithm was implemented to the recorded regularities in the occurrence of the maxima over time in order to distinguish measurement artifacts from heart related pulsations. If there was a high regularity of the detected maxima, the presence of a heart related pulse signal was assumed and the signal processing detected a heartbeat and used the time between the new and last recognized maximum as one heart cycle.
The process of the signal processing is illustrated in Figure 6.
For the detection of the maxima, two signals were obtained from the measured signal
Further, this signal was smoothed by a second order low-pass with cut-off frequency at 0
For better illustration, Figure 7 shows an impedance signal measured at the calf as well as the filtered signals
The pulsatile component in
This algorithm showed the best results if flanks rose steadily in
For the evaluation, the impedance of the leg was measured and the data were recorded with a PC. In the further analysis, the current
The first evaluation step was the analysis of the influence of the capacitive measuring method on the signal quality and the detection of the pulse rate. For this, three different configurations were tested: a conductive coupling, a passive capacitive coupling, and an active capacitive coupling. In order to acquire a conductive coupling to the tissue, conventional ECG-adhesive electrodes were used. For passive capacitive coupling, the same electrodes as for the active coupling (described above) were used, but in this case the circuit-boards were unequipped. To provide three comparable conditions, the gain of the active electrode was set to
In the three configurations, the lower leg was stretched on a chair and the signal was recorded for 95 s. For comparison with the gold standard, the PPG-signal was recorded simultaneously on the second toe. Each peak time
During the next test, the subject sat on a chair while the examined lower leg was bent at an angle of about 90° to the thigh and placed perpendicular to the ground to the ground. Additionally to the second test, the leg was stretched and laid on the seat surface of a second chair. To prevent the occurrence of major motion artifacts during the measurements, the subject tried to move as little as possible and was breathing normally. From these measurements the pulsations were separated in correct and wrong detected beats compared to the control results of the gold standard. The accuracy of the detection algorithm, with respect to the peak time and the heart rate, was analyzed in more detail.
Since the measured PPG signal was artifact free, the beats were identified by the local maxima. From the PPG, the peak time
In the first evaluation step no false-positive pulse beats were found by the detection algorithm. Thus, the false-positive rate was 0% for all three experiments. The sensitivity of the conductivity measurement method was approximately 65 %. In this experiment, the capacitive active electrodes could achieve a sensitivity of about 34 %. The lowest sensitivity was found for the capacitive passive measurement, which was only 10 %. By analysis of the raw data pulse curve, it was apparent that the conductive measurement was not so prone to movement artifacts, like breathing, as the capacitive measurements has been. The impedance signal of the capacitive passive measurement was additionally superimposed of strong common-mode interference.
The second evaluation step showed that the waveform of the IPG signal strongly varied over the time. This behavior possibly correlated with the breathing of the subject, which is reflected in form of a low-frequency, additive component in the PPG signal. It also showed that the pulsations of the IPG signal occurred simultaneously with those of the PPG signal and the expected time delay of 78 ms was not recognizable. This was explainable by the time delay of the signal processing of the PPG-meter or it might be a physiologically time delay of the pulse wave transit time between calf and toe.
Every beat of the PPG was assigned to the IPG by the well-known criterion of eq. (17). By this method, each element of
Table 1 shows the measurement period, the amount of heart beats which are detected by the gold standard, and true-positive false-positive and false-negative findings of the algorithm in the IPG signal. The sensitivity of the new system
Measurement period and number of detected pulse waves for the bent and stretched leg.
|bent leg||128 s||150||97||97||53||0|
|stretched leg||100 s||116||81||81||35||0|
These results are comparable to the results of the previous section for the conductive measurement. For the evaluation of the detected heart rate a Bland-Altman plot was used to illustrate the performance, as shown in Figure 8 for the stretched leg and Figure 9 for the bent leg. The horizontal lines are drawn at the mean difference and at the limits of agreement. These are defined as the mean difference plus and minus 2 times the standard deviation
In the case of the bent leg, most differences of the heart rates Δ
This work proposes an integrated solution for monitoring the health status of risk patients as well as the according procedure to design and construct a wearable, semi-capacitive system.
Firstly, the analog front end and the digital back end of the measurement system were described. For this purpose, the frequency response was mathematically derived. Based on this derivation, the analog design of the capacitive electrodes and the amplifier circuit were shown. The system also provides the possibility to replace the capacitive active electrodes with conventional ECG electrode pads in order to increase the signal quality, resulting in a more reliable detection of pulsations. The digital back end is based on the microcontrollers AFE4300 and MSP430. The standard wiring, which was given by the manufacturer, was modified to fulfill the requirements concerning the measurement accuracy. In this work, the measurement system was used to record the time-varying impedance of a leg, which is related to blood.
Secondly, a real-time algorithm for the detection of beat-to-beat pulse waves was presented. The change of the frequencies of the heartbeat from conventional clinical measurement modalities (such as ECG) are a well-studied problem, e.g. for heart rate variability analysis . Furthermore, the real-time analysis of a beat to beat interval by means of IPG gives the advantage that another physiological rhythm can be evaluated contactless, even while wearing compression stockings. Hence, many useful parameters can be extracted in combination with a wearable ECG system, as for example the rise time of the IPG signal or the real-time pulse-wave velocity.
Thirdly, a proof-of-concept trial was conducted to evaluate the whole system in vivo. Two validation tests were performed. The first evaluation step was the analysis of the influence of conductive as well as passive, and active capacitive coupling regarding signal quality and the detection of the pulse rate. The false-positive rate was 0% for all three experiments. The sensitivity was the highest during the conductive measurement and about twice as large as during the active capacitive measurement. The passive capacitive electrodes, compared to the other two, were of insufficient quality, since the measurements were superimposed by many artifacts and had a high common mode noise.
The second evaluation step was done with the highest amplification of the active capacitive electrodes. The analysis showed that the algorithm was implemented conservatively; it provided a low false-positive rate of 0 %, but only a moderate sensitivity of about 68 %. In any case, a reliable and continuous measurement of the pulse signal was only possible in periods of immobility. Furthermore, strong breathing deteriorated the detection rate of the algorithm. The analysis also showed that the results were different for both body postures while the measurment with stretched leg showed slightly better results.
The evaluation suggests that a semi-capacitive impedance phlebography is a promising method to measure blood-dependent time-varying changes of the impedance. Therefore, such a system could be used for personal-healthcare or telemonitoring applications.
The evaluation showed that a semi-capacitive impedance phlebography is a promising method to measure blood-dependent time-varying changes of the impedance. However, different measurement positions on the leg yield differences in the signal quality, which requires further investigation. Even in commercial IPG devices, the electrode positioning is not yet standardized, as it is for impedance cardiography. Nevertheless, the system could assist the physician in continuously diagnosing and monitoring the peripheral vascular status of the patient, who is wearing compression stockings. Furthermore, valuable information about the reaction of the vascular system following its manipulation with, e.g., tilt table tests, exercise , or vasoactive drugs, could be gained. Not only clinical, but also home applications are conceivable. Therefore, such a system could be used for personal healthcare or telemonitoring applications. The system could provide information about the fitting of compression stockings or the current blood flow in the legs. Further investigations in these directions seem appropriate and necessary.
Measurement period and number of detected pulse waves for the bent and stretched leg.
|bent leg||128 s||150||97||97||53||0|
|stretched leg||100 s||116||81||81||35||0|