**4. Description of the main methods**

#### **4.1. Bioreactance**

that used a special breathing circuit extension loop was developed call the NICO (Respironics, Philips Healthcare, USA). The NICO is still produced but its use is restricted to intubated and

**Figure 2.** Elaborate NICO rebreathing loop and circuit attachment that was added to the patient's breathing circuit

In 2004 a device that used the time lags between the ECG and pulse oximetry signals was developed called the FloWave 1000, (Woolsthorpe Technologies, Brentwood, TN, USA). A Japanese group has recently developed a similar device called the esCCO monitor (Nihon Kohden, Tokyo, Japan) [21]. The esCCO also calculates pulse wave transit time from the ECG and pulse oximetry signal which it uses to calibrate the arterial pressure derived cardiac output

**Figure 3.** Illustration of the pulse wave transit time method used by the esCCO monitor. (Image from Nihon Kohden)

when performing the partial carbon dioxide rebreathing method.

ventilated patients (Figure 2).

50 Artery Bypass

(Figure 3).

To understand how the bioreactance method (NICOM, Cheetah Medical) works one first must understanding bioimpedance cardiac output. The older bioimpedance method involved detec‐ tion of electrical resistance changes within the thorax. A high-frequency (50-100 kHz) low am‐ plitude alternating current (<4mA), is passes between skin electrodes placed around the neck and upper abdomen. Inner current sensing skin electrodes detect voltage changes across the thorax and thus the impedance signal produced by the cardiac cycle (Figure 4). Originally, band electrodes were uses, but in the BoMed this was changed to eight dot electrodes. Bioimpe‐ dance is safe electrically because of the high frequency and low amperage of the current. The only report of injury with its use has been a pacemaker malfunction [22].

**Figure 4.** Electrode configurations used by different bioimpedance devices. The BoMed used an eight electrode con‐ figuration with outer current injecting and inner current sensing skin dot electrodes. Some other devices were de‐ signed with fewer but larger patch electrodes on the head and lower torso (current injecting) and neck and lower thorax (current sensing). The bioreactance system (NICOM) also uses a four dual dot electrode configuration with the neck electrodes placed slightly lower at the level of the clavicles.

In the original description of the impedance method the area under the bioimpedance signal curve during systole was used to estimate cardiac output. To simply the method Kubicek et al used the differential signal and its peak reading (dZ/dt(max)) as a surrogate for aortic blood flow [11]. The method also involves measuring the left ventricular ejection time (LVET) from the impedance signal (Figure 5). dZ/dt (max) multiplied by LVET provides stroke volume, but the reading still needs to be calibrated. Cardiac output is calculated by multiplying by heart rate. Other bioimpedance variables measured from the waveforms include: (i) the thoracic im‐ pedance which can be used as an index of lung fluid, (ii) the systolic time intervals, pre ejection period (PEP) and LVET, which can be used to calculate ejection fraction and (iii) the second dif‐ ferential (i.e. d2 Z/dt2 (max)) which can be used as an index of contractility.

**Figure 6.** The steps in deriving bioreactance cardiac output (Images from Cheetah Medical).

may also be influenced by variations in peripheral resistance.

**4.2. Continuous wave Doppler**

an undisclosed algorithm based on age, weight and height is used for calibration.

age of the underlying tissue structure. This is the basis of ultrasound imaging.

Like all surrogate cardiac output methods the bioreactance method needs to be calibrated. When using bioimpedance this requires estimation of the volume of electrically participating tissue (VEPT) lying between the current sensing electrodes. Kubicek et al modeled the thorax on a cylinder [11]. Bernstein later modified the equation to a truncated cone [12]. In the NICOM

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Just like bioimpedance, it is not known precisely what the bioreactance signal truly represents. Rather than the flow of blood, it probably reflects blood volume expansion in the aorta as the vessel distends with the rise in blood pressure generated during systole [16]. Thus readings

When pressure is applied to certain solid materials, notably crystals, they produce an electric charge. Equally, the same crystal will change shape when an electric charge is applied to it. This is known as the piezoelectric effect. If a high frequency current (i.e. 1-10 MHz) is applied the crystal will vibrate producing high frequency sound waves, or ultrasound. If the crystal is place in contact with the skin the ultrasound will be propagated through the underlying tis‐ sues. When the ultrasound beam hits an interface between two tissue structures part of beam is reflected back. If a short burst of ultrasound is used and a second crystal is used as a receiver, then the time delays between transmission and return of this pulse can be used to create an im‐

**Figure 5.** The bioimpedance method uses both the impedance signal (Z – upper waveform) and the differential signal (dZ/dt – lower waveform). From the differential signal the flow variable dZ/dt(max) is measured. The time variable LVET is also measured. A number of other indices that reflect lung fluid and contractility are also measured.

Bioreactance uses a different electrical signal. It detects a property of alternating current called phase. An alternating current has a sinusoidal waveform. As the current flows through different body tissues its passage is delayed by capacitive and inductive tissue effects (X) which cause a shift in its phase. As blood volume in the central thorax region varies with the cardiac cycle so does the phase shift of the current. Like resistance when measuring bioimpedance, a signal of the phase shift (bioreactance signal) can be plotted and from it variables that reflected blood flow (dX/dt(max)) and ventricular ejection time are measured (Figure 6). It is thought that the bioreactance signal is less affected by the factors that troubled the bioimpedance method, such as lung water [15].

**Figure 6.** The steps in deriving bioreactance cardiac output (Images from Cheetah Medical).

Like all surrogate cardiac output methods the bioreactance method needs to be calibrated. When using bioimpedance this requires estimation of the volume of electrically participating tissue (VEPT) lying between the current sensing electrodes. Kubicek et al modeled the thorax on a cylinder [11]. Bernstein later modified the equation to a truncated cone [12]. In the NICOM an undisclosed algorithm based on age, weight and height is used for calibration.

Just like bioimpedance, it is not known precisely what the bioreactance signal truly represents. Rather than the flow of blood, it probably reflects blood volume expansion in the aorta as the vessel distends with the rise in blood pressure generated during systole [16]. Thus readings may also be influenced by variations in peripheral resistance.

### **4.2. Continuous wave Doppler**

In the original description of the impedance method the area under the bioimpedance signal curve during systole was used to estimate cardiac output. To simply the method Kubicek et al used the differential signal and its peak reading (dZ/dt(max)) as a surrogate for aortic blood flow [11]. The method also involves measuring the left ventricular ejection time (LVET) from the impedance signal (Figure 5). dZ/dt (max) multiplied by LVET provides stroke volume, but the reading still needs to be calibrated. Cardiac output is calculated by multiplying by heart rate. Other bioimpedance variables measured from the waveforms include: (i) the thoracic im‐ pedance which can be used as an index of lung fluid, (ii) the systolic time intervals, pre ejection period (PEP) and LVET, which can be used to calculate ejection fraction and (iii) the second dif‐

(max)) which can be used as an index of contractility.

**Figure 5.** The bioimpedance method uses both the impedance signal (Z – upper waveform) and the differential signal (dZ/dt – lower waveform). From the differential signal the flow variable dZ/dt(max) is measured. The time variable

Bioreactance uses a different electrical signal. It detects a property of alternating current called phase. An alternating current has a sinusoidal waveform. As the current flows through different body tissues its passage is delayed by capacitive and inductive tissue effects (X) which cause a shift in its phase. As blood volume in the central thorax region varies with the cardiac cycle so does the phase shift of the current. Like resistance when measuring bioimpedance, a signal of the phase shift (bioreactance signal) can be plotted and from it variables that reflected blood flow (dX/dt(max)) and ventricular ejection time are measured (Figure 6). It is thought that the bioreactance signal is less affected by the factors that troubled the bioimpedance

LVET is also measured. A number of other indices that reflect lung fluid and contractility are also measured.

ferential (i.e. d2

52 Artery Bypass

Z/dt2

method, such as lung water [15].

When pressure is applied to certain solid materials, notably crystals, they produce an electric charge. Equally, the same crystal will change shape when an electric charge is applied to it. This is known as the piezoelectric effect. If a high frequency current (i.e. 1-10 MHz) is applied the crystal will vibrate producing high frequency sound waves, or ultrasound. If the crystal is place in contact with the skin the ultrasound will be propagated through the underlying tis‐ sues. When the ultrasound beam hits an interface between two tissue structures part of beam is reflected back. If a short burst of ultrasound is used and a second crystal is used as a receiver, then the time delays between transmission and return of this pulse can be used to create an im‐ age of the underlying tissue structure. This is the basis of ultrasound imaging.

When a beam of continuous ultrasound encounters moving blood cells flowing In a blood vessel the ultrasound is reflected back at a slightly altered frequency. This phenomenon is known as the Doppler affect. The change or shift in frequency is related to the velocity of the blood cells. The Doppler shift signal can be separated from the ultrasound signal and a profile of the Doppler signal displayed (Figure 7). The angle (theta θ) that the ultrasound beam makes with the direction of blood flow is also important as it affects the magnitude of the Doppler shift frequency. If the direction of the ultrasound beam is parallel to the blood flow the Doppler shift will be maximal, whilst a perpendicular angle of insonation produces no Doppler shift. The angle of insonation (θ) and Doppler shift frequency are related to the cosine of theta (cos(θ)). The velocity of the blood flow is related to the Doppler frequency by the equation *velocity* = *c* × *fD* / 2× *fT* cos *θ*, where fD is the Doppler shift frequency, fT is the ultrasound probe or transmitter frequency, and c is the speed of ultrasound in the tissues, 1540 m/s. The speed of sound in air is around 340 m/s.

Blood flow in the aorta pulsates rather than being continuous, thus a continuous Dop‐ pler ultrasound signal needs to be recorded with sufficient sampling rate to show the de‐ tails of the flow profile (Figure 7). Most ultrasound machines are imaging systems and use pulses of ultrasound to measure distance from the probe or depth like radar or so‐ nar. Doppler is different because it detects change in velocity rather than position and requires a continuous ultrasound beam from a transmitting crystal and a separate receiv‐ ing crystal. From the Doppler profile of blood flow in the aorta the peak velocity of the blood and the duration of flow can be determined. By drawing an envelope around the Doppler flow profile one can calculate the total flow during systole, which is called the

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To convert stroke distance to a volume (i.e. stroke volume) the cross-sectional area of the blood vessel is needed. In conventional echocardiography machines this is measured by ultrasound imaging using the relationship *CSA*= *π* ×*d* <sup>2</sup> / 4, where CSA = cross sectional area and d =

Two Doppler cardiac output systems are currently on the market, the CardioQ (Deltex Medical) (Figure 8) and the USCOM (USCOM Ltd.) (Figure 9). Neither measures the CSA of the aorta directly and both estimate it but in different ways. The CardioQ uses an empirical algorithm based on population data, where the calibration constant is based on the patient's age, gender, height and weight. As the CardioQ measures blood flow from descending aorta where about 30% of the blood flow has left the aorta for the head and arms, its algorithm corrects for this reduction in total flow. The USCOM measures blood flow across the aortic or pulmonary valve. It uses an empirical formula to calculate valve CSA [23] which also requires

The angle of insonation with blood flow of the probe needs to be considered. When the CardioQ is used its probe is in the oesophagus and lies parallel to the descending aorta. The ultrasound crystals at the tip are set to 45-degrees (Figure 8). Therefore, its angle of insonation is 45-degrees. The USCOM probe has a wide beam angle. It is directed at the aortic or pulmonary valves and its beam axis usually lies almost parallel to the direction of flow because of the anatomy. Thus, the angle of insonation (θ) is close to 90-degrees and the cosine of the angle approximates to 1.0. Neither device is corrected for devia‐

Focusing of the probe to obtain the optimal and maximum Doppler signal plays an ex‐ tremely critical role in using these two Doppler devices effectively. Focusing can be per‐ formed both visually by observing the shape of Doppler profiles on the monitor screen or by listening to the quality of the audible Doppler signal. Various numbers of patient examinations are quoted to acquire competence in the focusing technique, 12 for the Car‐ dioQ and 20 for the USCOM [24,25]. However, it takes a much longer time to become sufficiently familiar with the different signal sounds and patterns to recognize when a truly reliable signal has been obtained. Significant experience and psychomotor skill is needed to be able to acquire clinically reliable data, with the CardioQ being easier to

stroke (or minute) distance (Figure 7).

the patient's age, gender, weight and height.

tions in beam angle to blood flow.

learn. Both companies provide training and support.

diameter of blood vessel.

**Figure 7.** Doppler flow profiles from the oesophagus (upper - CardioQ) and the supra-sternal window (lower - US‐ COM). Velocity is shown on the y-axis (m/s) and time along x-axis. The outline of each Doppler signal is automatically detected and drawn. The area of each envelop (stroke distance) is related to stroke volume. A series of cardiac cycles are shown. (Upper image from Deltex Medical)

Blood flow in the aorta pulsates rather than being continuous, thus a continuous Dop‐ pler ultrasound signal needs to be recorded with sufficient sampling rate to show the de‐ tails of the flow profile (Figure 7). Most ultrasound machines are imaging systems and use pulses of ultrasound to measure distance from the probe or depth like radar or so‐ nar. Doppler is different because it detects change in velocity rather than position and requires a continuous ultrasound beam from a transmitting crystal and a separate receiv‐ ing crystal. From the Doppler profile of blood flow in the aorta the peak velocity of the blood and the duration of flow can be determined. By drawing an envelope around the Doppler flow profile one can calculate the total flow during systole, which is called the stroke (or minute) distance (Figure 7).

When a beam of continuous ultrasound encounters moving blood cells flowing In a blood vessel the ultrasound is reflected back at a slightly altered frequency. This phenomenon is known as the Doppler affect. The change or shift in frequency is related to the velocity of the blood cells. The Doppler shift signal can be separated from the ultrasound signal and a profile of the Doppler signal displayed (Figure 7). The angle (theta θ) that the ultrasound beam makes with the direction of blood flow is also important as it affects the magnitude of the Doppler shift frequency. If the direction of the ultrasound beam is parallel to the blood flow the Doppler shift will be maximal, whilst a perpendicular angle of insonation produces no Doppler shift. The angle of insonation (θ) and Doppler shift frequency are related to the cosine of theta (cos(θ)). The velocity of the blood flow is related to the Doppler frequency by the equation *velocity* = *c* × *fD* / 2× *fT* cos *θ*, where fD is the Doppler shift frequency, fT is the ultrasound probe or transmitter frequency, and c is the speed of ultrasound in the tissues, 1540 m/s. The

**Figure 7.** Doppler flow profiles from the oesophagus (upper - CardioQ) and the supra-sternal window (lower - US‐ COM). Velocity is shown on the y-axis (m/s) and time along x-axis. The outline of each Doppler signal is automatically detected and drawn. The area of each envelop (stroke distance) is related to stroke volume. A series of cardiac cycles

speed of sound in air is around 340 m/s.

54 Artery Bypass

are shown. (Upper image from Deltex Medical)

To convert stroke distance to a volume (i.e. stroke volume) the cross-sectional area of the blood vessel is needed. In conventional echocardiography machines this is measured by ultrasound imaging using the relationship *CSA*= *π* ×*d* <sup>2</sup> / 4, where CSA = cross sectional area and d = diameter of blood vessel.

Two Doppler cardiac output systems are currently on the market, the CardioQ (Deltex Medical) (Figure 8) and the USCOM (USCOM Ltd.) (Figure 9). Neither measures the CSA of the aorta directly and both estimate it but in different ways. The CardioQ uses an empirical algorithm based on population data, where the calibration constant is based on the patient's age, gender, height and weight. As the CardioQ measures blood flow from descending aorta where about 30% of the blood flow has left the aorta for the head and arms, its algorithm corrects for this reduction in total flow. The USCOM measures blood flow across the aortic or pulmonary valve. It uses an empirical formula to calculate valve CSA [23] which also requires the patient's age, gender, weight and height.

The angle of insonation with blood flow of the probe needs to be considered. When the CardioQ is used its probe is in the oesophagus and lies parallel to the descending aorta. The ultrasound crystals at the tip are set to 45-degrees (Figure 8). Therefore, its angle of insonation is 45-degrees. The USCOM probe has a wide beam angle. It is directed at the aortic or pulmonary valves and its beam axis usually lies almost parallel to the direction of flow because of the anatomy. Thus, the angle of insonation (θ) is close to 90-degrees and the cosine of the angle approximates to 1.0. Neither device is corrected for devia‐ tions in beam angle to blood flow.

Focusing of the probe to obtain the optimal and maximum Doppler signal plays an ex‐ tremely critical role in using these two Doppler devices effectively. Focusing can be per‐ formed both visually by observing the shape of Doppler profiles on the monitor screen or by listening to the quality of the audible Doppler signal. Various numbers of patient examinations are quoted to acquire competence in the focusing technique, 12 for the Car‐ dioQ and 20 for the USCOM [24,25]. However, it takes a much longer time to become sufficiently familiar with the different signal sounds and patterns to recognize when a truly reliable signal has been obtained. Significant experience and psychomotor skill is needed to be able to acquire clinically reliable data, with the CardioQ being easier to learn. Both companies provide training and support.

In addition to measuring stroke volume and cardiac output, both Doppler devices provide in‐ ternal software to (a) calculate other haemodynamic parameters, (b) display data trends and (c) store data for future reference. One particularly useful parameter measured by these Doppler systems is the flow time corrected (FTc), an index of preload or ventricular filling. It measures the duration of systole corrected for heart rate. More advanced models are sold that calculate

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The arterial pulse contour method in essence is very simple. An arterial catheter is inserted into a peripheral artery, usually the radial or femoral. The catheter is connected to a pressure trans‐ ducer which is zeroed and checked for under or over damping. The analog arterial pressure sig‐ nal is fed into a device that calculates cardiac output from the trace. However, there are at least ten different algorithms that can be used to derive cardiac output from arterial pressure. The theoretical basis to these different algorithms is extremely complicated and involves different mathematical models that describe the circulation and adjust for changes in its impedance and compliance of the peripheral circulation. A brief outline of how these algorithms is given.

**a.** The simplest model that describes the circulation is the pressure = flow x resistance relationship. The area under the arterial pressure curve is directly proportional to cardiac output providing peripheral resistance remains constant. Unfortunately, peripheral resistance does not remain constant. It is constantly changing under the influence of the sympathetic nervous system which helps to maintain blood pressure and the circulation

**b.** Changes in peripheral resistance are reflected in diastolic pressure, so the simplest adjustment to the model is the use of pulse pressure (i.e. systolic-diastolic) rather than the arterial pressure to calculate cardiac output. This method is used in several pulse contour

**c.** The dynamics of the circulation is not as simple as pressure = flow x resistance. The circulation is a pulsitile system and when the heart pumps the arterial system has to expand to accommodate the additional blood. Windkessel compared the arterial system to a capacitor and proposed a two element model of the circulation with both resistive

**d.** The two element model still did not describe the circulation in its entirety. Wesseling et al added a third inductive element to compensate for time lags as blood flowed through the arterial system [19]. Their three-element model was called "Model Flow" and was first

**e.** Although, blood flow in the ascending aorta occurs during systole, as the blood travels more distally a significant proportion of blood flow also occurs in diastole and this component forms part of the peripheral arterial pressure wave. Thus, algorithms that measure cardiac output from a peripheral site such as the radial artery also should compensate for the diastolic component. One method is to identify the dichotic notch in the pulse wave and thus differentiate between the systolic and diastolic components.

used in the Finapres, a finger blood pressure cuff technology.

as body position changes or the person exercises.

inotropy and oxygen delivery from the blood pressure and oxygen saturation readings.

**4.3. Pulse contour analysis**

systems.

and capacitive components.

**Figure 8.** The CardioQ oesophageal Doppler monitor. Monitor and probe tip shown with transmitter and receiver crystals set at a 45-degree angle. Anatomical diagram shows insertion of the probe into the oesophagus via the mouth and insonation of the aorta which lies posterior. (Images from Deltex Medical)

**Figure 9.** USCOM monitor showing Doppler signal data on its screen. The flow profiles are automatically outlined to measure stroke volumes. Below numerical readings are displayed. Lower right is a trend plot of saved cardiac output readings. The hand held USCOM probe is shown in front of the monitor. Ultrasound gel is applied to the probe to improve its acoustic contact.

In addition to measuring stroke volume and cardiac output, both Doppler devices provide in‐ ternal software to (a) calculate other haemodynamic parameters, (b) display data trends and (c) store data for future reference. One particularly useful parameter measured by these Doppler systems is the flow time corrected (FTc), an index of preload or ventricular filling. It measures the duration of systole corrected for heart rate. More advanced models are sold that calculate inotropy and oxygen delivery from the blood pressure and oxygen saturation readings.
