Cardiac output

Cardiac output (Q or $\dot Q_{c}$ or CO ) is the volume of blood being pumped by the heart, in particular by a left or right ventricle in the time interval of one minute. CO may be measured in many ways, for example dm3/min (1 dm3 equals 1 litre). Q is furthermore the combined sum of output from the right ventricle and the output from the left ventricle during the phase of systole of the heart. An average resting cardiac output would be 5.6 L/min for a human male and 4.9 L/min for a female.[1]

Clinical uses

The function of the heart is to transport blood to deliver oxygen, nutrients and chemicals to the cells of the body to ensure their survival and proper function and to remove the cellular wastes. Since the heart is a 'demand pump', that pumps out whatever blood comes back into it from the venous system, it is effectively the amount of blood returning to the heart that determines how much blood the heart pumps out (Q). This, in turn, is controlled principally by the demand for oxygen by the cells of the body and the capacitance of the arterio-venous system. If the body has a high metabolic oxygen demand then the metabolically controlled flow through the tissues is increased, leading to a greater flow of blood back to the heart. This is also modified by the function of the vessels of the body as they actively relax and contract thereby increasing and decreasing the resistance to flow.

When Q increases in a healthy but untrained individual, most of the increase can be attributed to an increase in heart rate (HR). Change of posture, increased sympathetic nervous system activity, and decreased parasympathetic nervous system activity can also increase cardiac output. HR can vary by a factor of approximately 3, between 60 and 180 beats per minute, while stroke volume (SV) can vary between 70 and 120 ml, a factor of only 1.7.[2][3][4]

A parameter related to SV is Ejection Fraction (EF). EF is the fraction of blood ejected by the Left Ventricle (LV) during the contraction or ejection phase of the cardiac cycle or Systole. Prior to the start of Systole, the LV is filled with blood to the capacity known as End Diastolic Volume (EDV) during the filling phase or diastole. During Systole, the LV contracts and ejects blood until it reaches its minimum capacity known as End Systolic Volume (ESV), it does not empty completely. Clearly the EF is dependent on the ventricular EDV which may vary with ventricular disease associated with ventricular dilatation. Even with LV dilatation and impaired contraction the Q may remain constant due to an increase in EDV.

Stroke Volume (SV) = EDV – ESV
Ejection Fraction (EF) = (SV / EDV) × 100%
Cardiac Output (Q) = SV × HR
Cardiac Index (CI) = Q / Body Surface Area (BSA) = SV × HR/BSA
HR is Heart Rate, expressed as BPM (Beats Per Minute)
BSA is Body Surface Area in square metres.

Diseases of the cardiovascular system are often associated with changes in Q, particularly the pandemic diseases of hypertension and heart failure. Cardiovascular disease can be associated with increased Q as occurs during infection and sepsis, or decreased Q, as in cardiomyopathy and heart failure. The ability to accurately measure Q is important in clinical medicine as it provides for improved diagnosis of abnormalities, and can be used to guide appropriate management.

Measuring cardiac output

There are a number of clinical methods for measurement of Q ranging from direct intracardiac catheterisation to non-invasive measurement of the arterial pulse. Each method has unique strengths and weaknesses and relative comparison is limited by the absence of a widely accepted "gold standard" measurement. Q can also be affected significantly by the phase of respiration; intra-thoracic pressure changes influence diastolic filling and therefore Q. This is especially important during mechanical ventilation where Q can vary by up to 50%[citation needed] across a single respiratory cycle. Q should therefore be measured at evenly spaced points over a single cycle or averaged over several cycles.

Invasive methods are well accepted, but there is increasing evidence that these methods are neither accurate nor effective in guiding therapy, so there is an increasing focus on development of non-invasive methods.[5][6][7]

The Fick principle

The Fick principle was first described by Adolf Eugen Fick in 1870 and assumes that the rate at which oxygen is consumed is a function of the rate of blood flows and the rate of oxygen picked up by the red blood cells. The Fick principle involves calculating the oxygen consumed over a given period of time from measurement of the oxygen concentration of the venous blood and the arterial blood. Q can be calculated from these measurements:

• VO2 consumption per minute using a spirometer (with the subject re-breathing air) and a CO2 absorber
• the oxygen content of blood taken from the pulmonary artery (representing mixed venous blood)
• the oxygen content of blood from a cannula in a peripheral artery (representing arterial blood)

From these values, we know that:

VO2 = (Q×CA) - (Q×CV)

where

• CA = Oxygen content of arterial blood
• CV = Oxygen content of venous blood.

This allows us to say

$Q\ = \frac{{{V}_O}_2}{{C}_A - {C}_V}$

and therefore calculate Q. Note that (CA – CV) is also known as the arteriovenous oxygen difference.[8]

While considered to be the most accurate method for Q measurement, Fick is invasive, requires time for the sample analysis, and accurate oxygen consumption samples are difficult to acquire. There have also been modifications to the Fick method where respiratory oxygen content is measured as part of a closed system and the consumed Oxygen calculated using an assumed oxygen consumption index which is then used to calculate Q. Other modifications use inert gas as tracers and measure the change in inspired and expired gas concentrations to calculate Q (Innocor, Innovision A/S, Denmark).

Additionally, the calculation of the arterial and venous oxygen content of the blood is a straightforward process. Almost all oxygen in the blood is bound to hemoglobin molecules in the red blood cells. Measuring the content of hemoglobin in the blood and the percentage of saturation of hemoglobin (the oxygen saturation of the blood) is a simple process and is readily available to physicians. Using the fact that each gram of hemoglobin can carry 1.34 ml of O2, the oxygen content of the blood (either arterial or venous) can be estimated by the following formula:

$Oxygen\ content\ of\ blood = \left [hemoglobin \right] \left ( g/dl \right ) \ \times\ 1.34 \left ( ml\ \mathrm{O}_2 /g\ of\ hemoglobin \right ) \times\ saturation\ of\ blood\ \left ( percent \right )\ +\ 0.0032\times\ partial\ pressure\ of\ oxygen \left ( torr \right )$

Dilution methods

The output of heart is equal to the amount of indicator injected divided by its average concentration in the arterial blood after a single circulation through the heart.

This method was initially described using an indicator dye and assumes that the rate at which the indicator is diluted reflects the Q. The method measures the concentration of a dye at different points in the circulation, usually from an intravenous injection and then at a downstream sampling site, usually in a systemic artery. More specifically, the Q is equal to the quantity of indicator dye injected divided by the area under the dilution curve measured downstream (the Stewart (1897)-Hamilton (1932) equation):

$Cardiac\ output = \frac{Quantity\ of\ Indicator}{\int_0^\infty Concentration\ of\ Indicator\cdot {dt}}$

The trapezoid rule is often used as an approximation of this integral.

Pulmonary artery thermodilution (trans-right-heart thermodilution)

The indicator method was further developed with replacement of the indicator dye by heated or cooled fluid and temperature change measured at different sites in the circulation rather than dye concentration; this method is known as thermodilution. The pulmonary artery catheter (PAC), also known as the Swan-Ganz catheter, was introduced to clinical practice in 1970 and provides direct access to the right heart for thermodilution measurements. Continuous invasive cardiac monitoring in the Intensive Care Unit has been all but phased out in an age of hospital acquired infection. Use of the PAC is still useful in right heart study in the cardiac catheterization laboratory today.

The PAC is balloon tipped and is inflated, which helps "sail" the catheter balloon through the right ventricle to occlude a smaller branch of the pulmonary artery system. The balloon is deflated. The PAC thermodilution method involves injection of a small amount (10ml) of cold glucose at a known temperature into the pulmonary artery and measuring the temperature a known distance away (6–10 cm) using the same catheter with temperature sensors set apart at a known distance.

The historically significant Swan-Ganz multi-lumen catheter allows reproducible calculation of Cardiac Output from a measured time/temperature curve (The "thermodilution curve"). Enabled Thermistor technology allowed the observation that low CO registers temperature change slowly, and inversely, high CO registers temperature change rapidly. The degree of change in temperature is directly proportional to the cardiac output. Under this unique method, three or four repeated measurements or passes are usually averaged to improve accuracy.[9][10] Modern catheters are fitted with a heating filament which intermittently heats and measures the thermodilution curve providing serial Q measurement. However, these take an average of measurements made over 2–9 minutes, depending on the stability of the circulation, and thus do not provide continuous monitoring.

PAC use is complicated by arrhythmias, infection, pulmonary artery rupture, and right heart valve damage. Recent studies in patients with critical illness, sepsis, acute respiratory failure and heart failure suggest use of the PAC does not improve patient outcomes.[5][6][7] This clinical ineffectiveness may relate to its poor accuracy and sensitivity which has been demonstrated by comparison with flow probes across a sixfold range of Qs.[11] PAC use is in decline as clinicians move to less invasive and more accurate technologies for monitoring hemodynamics.

Doppler ultrasound method

Doppler signal in the left ventricular outflow tract: Velocity Time Integral (VTI)

This method uses ultrasound and the Doppler effect to measure Q. The blood velocity through the heart causes a 'Doppler shift' in the frequency of the returning ultrasound waves. This Doppler shift can then be used to calculate flow velocity and volume and effectively Q using the following equations:

• Q = SV × HR
• SV = VTI × CSA

where:

• CSA = valve orifice cross sectional area; use pr²
• VTI = the velocity time integral of the trace of the Doppler flow profile

Doppler ultrasound is non-invasive, accurate and inexpensive and is a routine part of clinical ultrasound with high levels of reliability and reproducibility having been in clinical use since the 1960s.

Echocardiography

Echocardiography is a noninvasive method of quantifying cardiac output using Ultrasound. Two dimensional (2D) ultrasound with Doppler measurements are used together to calculate Cardiac Output. 2D measurement of the diameter (d) of the aortic annulus allows calculation of the flow CSA (cross-sectional area) which is then multiplied by the VTI of the Doppler flow profile across the aortic valve to determine the flow volume per beat (Stroke Volume, SV) which is then multiplied by the Heart Rate to obtain cardiac output. Although used in clinical medicine, it has a wide test-retest variability.[12] It is said to require extensive training and skill, but the exact steps needed to achieve clinically adequate precision have never been disclosed. 2D measurement of the aortic valve diameter is one source of noise, and beat-to-beat variation in stroke volume and subtle differences in probe position are the others. Measurement of the pulmonary valve to calculate right sided CO is an alternative, that is not necessarily more reproducible. While in wide general use the technique is time consuming and limited by the reproducibility of its component elements. In the manner used in clinical practice, precision of SV and CO is of the order of ±20%.

Transcutaneous Doppler: USCOM

The Ultrasonic Cardiac Output Monitor (USCOM,[11] Uscom Ltd, Sydney, Australia) uses Continuous Wave Doppler (CW) to measure the Doppler flow profile vti, as in echocardiography, but uses anthropometry to calculate aortic and pulmonary valve diameters and CSA's allowing both right and left sided Q measurements. This also significantly improves reproducibility compared with the echocardiographic method and therefore increases sensitivity for detection of changes in flow. Real time Automatic tracing of the Doppler flow profile allows for beat to beat right and left sided Q measurements significantly simplifying operation and reducing the time of acquisition compared with the conventional echocardiographic method. USCOM has been validated from 0.12 l/min to 18.7 l/min[13] in neonates,[14] children[15] and adults.[16] This means the method can be applied with equal accuracy to neonates, children and adults for the development of physiologically rational haemodynamic protocols. USCOM is the only method of cardiac output measurement to have achieved equivalent accuracy to the gold standard implantable flow probe.[11] This accuracy has ensured high levels of clinical utility across a range of applications including sepsis, heart failure and hypertension.[17][18][19]

Transoesophageal Doppler: TOD

Transoesophageal Doppler (TOD), is a term encompassing two main technologies: Transoesophageal Echocardiogram (TOE/TEE) which is primarily used for diagnostic purposes, and (what is commonly termed) oesophageal Doppler (ODM/EDM), primarily used for the clinical monitoring of cardiac output. The latter utilises CW ultrasound and the Doppler effect to measure blood velocity in the descending thoracic aorta. An ultrasound probe is inserted either orally or nasally into the oesophagus to mid-thoracic level, at which point the oesophagus lies alongside the descending thoracic aorta. Because the transducer is close to the blood flow the signal is clear, however the probe may require re-focussing to ensure an optimal signal. This method has good validation, is widely used for fluid management during surgery with evidence for improved patient outcome,[20][21][22][23][24][25][26][27] and has been recommended by the UK's National Institute for Health and Clinical Excellence (NICE).[28] One limitation is that ODM measures the velocity of blood and not true Q, therefore relies on a nomogram[29] based on patient age, height, and weight to convert the measured velocity into Stroke Volume and Cardiac Output. This method generally requires patient sedation and is accepted for use in both adults and paediatrics.

Pulse Pressure methods

Pulse Pressure (PP) methods measure the pressure in an artery over time to derive a waveform and use this information to calculate cardiac performance. However any measure from the artery includes the changes in pressure associated with changes in arterial function (compliance, impedance, etc..).

Physiologic or therapeutic changes in vessel diameter are assumed to reflect changes in Q. Put simply, PP methods measure the combined performance of the heart and the vessels thus limiting the application of PP methods for measurement of Q. This can be partially compensated for by intermittent calibration of the waveform to another Q measurement method and then monitoring the PP waveform. Ideally, the PP waveform should be calibrated on a beat to beat basis.

There are invasive and non-invasive methods of measuring PP:

Non-invasive PP – Sphygmomanometry and Tonometry

The sphygmomanometer or cuff blood pressure device was introduced to clinical practice in 1903 allowing non-invasive measurements of blood pressure and providing the common PP waveform values of peak systolic and diastolic pressure which can be used to calculate mean arterial pressure (MAP) and pulse pressure (PP). The pressure in the arteries, measured by sphygmomanometry, is often used as an indicator of the function of the heart. The pressure pulses in the heart are conducted to the arteries, and the arterial pressure is assumed to reflect the function of the heart or the Q. However no account is made of the elasticity of the arterial bed or its impact on the pressure signal.

• The pressure in the heart rises as blood is forced into the aorta
• The more stretched the aorta, the greater the pulse pressure (PP)
• In healthy young subjects, each additional 2 ml of blood results in a 1 mmHg rise in pressure
• Therefore:
SV = 2 ml × Pulse Pressure
Q = 2 ml × Pulse Pressure × HR

By resting a more sophisticated pressure sensing device, a tonometer, against the skin surface and sensing the pulsatile artery, continuous PP wave forms can be acquired non-invasively and analysis made of these pressure signals. However as the heart and vessels function independently and sometimes paradoxically the changes in the PP both reflect and mask changes in Q. So these measures represent combined cardiac and vascular function only. A similar system that uses the arterial pulse is the pressure recording analytical method (PRAM).

Finapres methodology

In 1967 the Czech physiologist Jan Peñáz invented and patented the volume clamp method to measure continuous blood pressure. The principle of the volume clamp method is to provide equal pressures dynamically on either side of the wall of an artery: inside pressure (= intra-arterial pressure) equals outside pressure (= finger cuff pressure) by clamping the artery to a certain volume. He decided that the finger was the optimal site to apply this volume clamp method. The use of finger cuffs excludes the device from application in patients without vasoconstriction, such as in sepsis, or patients on vasopressors.

In 1978 scientists at BMI-TNO, the research unit of Netherlands Organization for Applied Scientific Research at The University of Amsterdam, invented and patented a series of additional key elements to make the volume clamp work in clinical practice, among them: the use of modulated infra-red light in the optical system inside the sensor, the light-weight, easy to wrap finger cuff with Velcro fixation, a new pneumatic proportional control valve principle and last but not least the invention of a setpoint strategy for the determination and tracking of the correct volume at which to clamp the finger arteries – the Physiocal system. An acronym for PHYSIOlogical CALibration of the finger arteries, this Physiocal tracker turned out to be surprisingly accurate, robust and reliable and was never changed since its invention.

The Finapres methodology was developed to use this information to accurately calculate arterial pressure from the finger cuff pressure data. A generalized algorithm to correct for the pressure level difference between the finger and brachial sites within an individual patient was developed and this correction worked under all circumstances that it was tested, even when it was not designed for it, since it applied general physiological principles. The first implementation of this innovative brachial pressure waveform reconstruction was in the Finometer, the successor of Finapres that BMI-TNO introduced in the market in 2000.

The availability of a continuous, high-fidelity, calibrated blood pressure waveform opened up the perspective of beat-to-beat computation of integrated hemodynamics, based on two notions:

1. That pressure and flow are inter-related at each site in the arterial system by their so-called characteristic impedance and
2. That at the proximal aortic site, the 3-element Windkessel model of this impedance can be modeled with sufficient accuracy in an individual patient when age, gender, height and weight are known.

Recent work comparing nonivasive peripheral vascular monitors suggests modest clinical utility restricted to patients with normal and invariant circulation.[30]

Invasive PP

Invasive PP monitoring involves inserting a manometer (pressure sensor) into an artery, usually the radial or femoral artery and continuously measuring the PP waveform. This is usually done by connecting the catheter to a signal processing and display device. The PP waveform can then be analysed to provide measurements of cardiovascular performance. Changes in vascular function, the position of the catheter tip, or damping of the pressure waveform signal will all affect the accuracy of the readings. Invasive PP measurements can be calibrated or uncalibrated.

Calibrated PP – PiCCO, LiDCO

PiCCO (PULSION Medical Systems AG, Munich, Germany) and PulseCO (LiDCO Ltd, London, England) generate continuous Q by analysis of the arterial PP waveform. In both cases, an independent technique is required to provide calibration of the continuous Q analysis, as arterial PP analysis cannot account for unmeasured variables such as the changing compliance of the vascular bed. Recalibration is recommended after changes in patient position, therapy or condition.

In the case of PiCCO, transpulmonary thermodilution is used as the calibrating technique. Transpulmonary thermodilution uses the Stewart-Hamilton principle, but measures temperatures changes from central venous line to a central arterial line (i.e. femoral or axillary) arterial line. The Q derived from this cold-saline thermodilution is used to calibrate the arterial PP contour, which can then provide continuous Q monitoring. The PiCCO algorithm is dependent on blood pressure waveform morphology (i.e. mathematical analysis of the PP waveform) and calculates continuous Q as described by Wesseling and co-workers.[31] Transpulmonary thermodilution spans right heart, pulmonary circulation and left heart; this allows further mathematical analysis of the thermodilution curve, giving measurements of cardiac filling volumes (GEDV), intrathoracic blood volume, and extravascular lung water. While transpulmonary thermodilution allows for less invasive Q calibration, the method is also less accurate than PA thermodilution and still requires a central venous and arterial line with the attendant infection risks.

In the case of LiDCO, the independent calibration technique is lithium chloride dilution using the Stewart-Hamilton principle. Lithium chloride dilution uses a peripheral vein to a peripheral arterial line. Like PiCCO frequent calibration is recommended when there is a change in Q.[32] Calibration events are limited in frequency because it involves injection of Lithium Chloride, and can be subject to error in the presence of certain muscle relaxants. The PulseCO algorithm used by LiDCO is based on pulse power derivation and is not dependent on waveform morphology.

Statistical analysis of Arterial Pressure — FloTrac/Vigileo

FloTrac/Vigileo (Edwards Lifesciences LLC, U.S.A.) is an uncalibrated pulse contour analysis-based hemodynamic monitor that estimates cardiac output (Q) utilizing a standard arterial catheter with a manometer located in the femoral or radial artery. The device consists of a special high fidelity pressure transducer which, when used with a supporting monitor (Vigileo or EV1000 monitor), derives left-sided cardiac output (Q) from a sample of arterial pulsations. The device utilises an algorithm that is based on the principle that pulse pressure (PP) is proportional to stroke volume (SV).[33] The algorithm calculates the product of the standard deviation of the arterial pressure wave (AP) (over a sampled period of time of 20 seconds) and a vascular tone factor (Khi) to generate stroke volume. The equation in simplified form is as follows: SV=std(AP) * Khi or BP x k(constant). Khi is conceived to reflect arterial resistance, and compliance is a multivariate polynomial equation that continuously quantifies arterial compliance and vascular resistance. Khi does so by analyzing the morphologic change of the arterial pressure waveforms on a bit by bit basis (based on the principle that changes in compliance or resistance affect the shape of the arterial pressure waveform). By analyzing the shape of the arterial pressure waveform, the effect of vascular tone is assessed allowing calculation of SV. Cardiac Output (Q) is then derived utilizing the equation Q=HR*SV. Only perfused beats that generate an arterial waveform are counted for HR.

This system estimates Q using an existing arterial catheter with variable accuracy and precision. While these invasive arterial monitors do not require intracardiac catheterisation from a pulmonary artery catheter, they do require an arterial line and are invasive. As with the other arterial waveform systems the short time required for set up and data acquisition are additional benefits of this technology. Disadvantages include its inability to provide data regarding right-sided heart pressures, or mixed venous oxygen saturation.[34][35][36] Further arterial monitoring systems are unable to predict changes in vascular tone and so estimate changes in vascular compliance. The measurement of pressure in the artery to calculate the flow in the heart is physiologically irrational and of questionable accuracy,[37] and of unproven benefit [38] Arterial pressure monitoring is limited in patients off ventilation, in atrial fibrillation, in patients on vasopressors and in patients with a dynamic autonomic system such as in sepsis.[32]

Uncalibrated, pre-estimated demographic data-free — PRAM

Pressure Recording Analytical Method (PRAM), estimates Q from the analysis of the pressure wave profile obtained from an arterial catheter (radial or femoral access). This PP waveform can then be used to determine Q similarly to FloTrac. As the waveform is sampled at 1000 Hz, the detected pressure curve can be measured to calculate the real (relative to the patient under examination) and actual (beat-to-beat) Stroke Volume. Unlike FloTrac, no constant values of impedance deriving from an external calibration neither form pre-estimated in vivo/in vitro data are needed.
PRAM has been validated against the considered gold standard methods in stable condition[39] and in various hemodynamic states;[40] it can be used to monitor pediatric[41] and mechanically supported[42] patients.
A part to generally monitored hemodynamic values and to fluid responsiveness parameters, an exclusive reference is also provided by PRAM: Cardiac Cycle Efficiency (CCE). Expressed by a pure number ranging from 1 (the best) and -1 (the worse) it indicates the overall heart-vascular response coupling; the ratio between the heart performed and consumed energy, represented as CCE "stress index", can be of paramount importance in understanding patient present and next future course.[43]

Impedance cardiography

Impedance cardiography (often related as ICG or TEB) is a method that measures changes in impedance across the thoracic region over the cardiac cycle. Lower impedance indicates greater the intrathoracic fluid volume and blood flow. Therefore, by synchronizing fluid volume changes with heartbeat, the change in impedance can be used to calculate stroke volume, cardiac output, and systemic vascular resistance.[44]

Both invasive and non-invasive approaches are being used.[45] The noninvasive approach has achieved some acceptance with respect to its reliability and validity.[46][47][48][49] although there is not complete agreement on this point.[50] The clinical use of this approach in a variety of diseases continues.[51]

Noninvasive ICG equipment includes the Bio-Z Dx[52][verification needed] (Sonosite Inc, Bothell, WA) and the niccomo[53][verification needed] (medis GmbH, Ilmenau, Germany).

Ultrasound dilution method

Ultrasound dilution (UD) uses body temperature normal saline (NS) as an indicator introduced into an extracorporeal loop to create an AV circulation, with an ultrasound sensor used to measure the dilution and then calculate cardiac output using a proprietary algorithm. A number of other hemodynamic variables can also be calculated such as total end-diastole volume (TEDV), central blood volume (CBV) and active circulation volume (ACVI). The UD method was firstly introduced in 1995.,[54] and it was used extensively to measure flow and volumes with extracorporeal circuits condition such as ECMO[55][56] and Hemodialysis,[57][58] leading more than 150 peer reviewed publications, and now it has adapted to Intensive Care Units (ICU) settings as COstatus (Transonic System Inc. Ithaca, NY). The UD method is based on ultrasound indicator dilution.[59] Blood ultrasound velocity (1560–1585 m/s) is a function of total blood protein concentration (sums of proteins in plasma and in red blood red cells), temperature etc. Injection of body temperature normal saline (ultrasound velocity of saline is 1533 m/s) into a unique AV loop decreases blood ultrasound velocity, and produce dilution curves. UD requires establishment of an extracorporeal circulation through its unique AV loop with two preexisting arterial and central venous lines in ICU patients. When the saline indicator is injected into the A-V loop, it is detected by the venous clamp-on sensor on the AV loop before it enters the patient’s right heart atrium. After the indicator traverses the heart and lung, the concentration curve in the arterial line is recorded and displayed on the COstatus HCM101 Monitor. Cardiac output is calculated from the area of the concentration curve by the classic Stewart-Hamilton equation. It is a non-invasive procedure only by connection the AV loop and two lines of a patient. UD has been specialised for application in pediatric ICU patients, and has been demonstrated to be a relatively safe, although invasive, and reproducible tool.

Electrical Cardiometry

Electrical Cardiometry is a non-invasive method similar to Impedance cardiography, in the fact that both methods measure thoracic electrical bioimpedance (TEB). The underlying model is what differs, being that Electrical Cardiometry attributes the steep increase of TEB beat to beat to the change in orientation of red blood cells. Four standard ECG electrodes are required for measurement of cardiac output. Electrical Cardiometry is a method trademarked by Cardiotronic, Inc., and shows promising results in a wide range or patients (is currently US market approved for use in adults, pediatrics, and neonates). Electrical Cardiometry monitors have shown promise in postoperative cardiac surgical patients (both hemodynamicially stable and unstable).[60]

Magnetic Resonance Imaging

Velocity encoded phase contrast Magnetic Resonance Imaging (MRI)[61] is the most accurate technique for measuring flow in large vessels in mammals. MRI flow measurements have been shown to be highly accurate compared to measurements with a beaker and timer[62] and less variable than both the Fick principle[63] and thermodilution.[64]

Velocity encoded MRI is based on detection of changes in the phase of proton precession. These changes are proportional to the velocity of the movement of those protons through a magnetic field with a known gradient. When using velocity encoded MRI, the result of the MRI scan is two sets of images for each time point in the cardiac cycle. One is an anatomical image and the other is an image where the signal intensity in each pixel is directly proportional to the through-plane velocity. The average velocity in a vessel, i.e. the aorta or the pulmonary artery, is hence quantified by measuring the average signal intensity of the pixels in the cross section of the vessel, and then multiplying by a known constant. The flow is calculated by multiplying the mean velocity by the cross-sectional area of the vessel. This flow data can be used to graph flow versus time. The area under the flow versus time curve for one cardiac cycle is the stroke volume. The length of the cardiac cycle is known and determines heart rate, and thereby Q can be calculated as the product of stroke volume and heart rate. MRI is typically used to quantify the flow over one cardiac cycle as the average of several heart beats, but it is also possible to quantify the stroke volume in real time on a beat-for-beat basis.[65]

While MRI is an important research tool for accurately measuring Q, it is currently not clinically used for hemodynamic monitoring in the emergency or intensive care setting. Cardiac output measurement by MRI is currently routinely used as a part of clinical cardiac MRI examinations.[66]

Cardiac output and vascular resistance

The vascular beds are a dynamic and connected part of the circulatory system against which the heart must pump to transport the blood. Q is influenced by the resistance of the vascular bed against which the heart is pumping. For the right heart this is the pulmonary vascular bed, creating Pulmonary Vascular Resistance (PVR), while for the systemic circulation this is the systemic vascular bed, creating Systemic Vascular Resistance (SVR). The vessels actively change diameter under the influence of physiology or therapy, vasoconstrictors decrease vessel diameter and increase resistance, while vasodilators increase vessel diameter and decrease resistance. Put simply, increasing resistance decreases Q; conversely, decreased resistance increases Q.

This can be explained mathematically:

By simplifying Darcy's Law, we get the equation that

Flow = Pressure/Resistance

When applied to the circulatory system, we get:

Q = (MAP – RAP)/TPR

Where MAP = Mean Aortic (or Arterial) Blood Pressure in mmHg,

RAP = Mean Right Atrial Pressure in mmHg and

TPR = Total Peripheral Resistance in dynes-sec-cm-5.

However, as MAP>>RAP, and RAP is approximately 0, this can be simplified to:

Q ≈ MAP/TPR

For the right heart Q ≈ MAP/PVR, while for the left heart Q ≈ MAP/SVR.

Physiologists will often re-arrange this equation, making MAP the subject, to study the body's responses.

As has already been stated, Q is also the product of the heart rate (HR) and the stroke volume (SV), which allows us to say:

Q ≈ (HR × SV) ≈ MAP / TPR

Cardiac output and respiration

Q is affected by the phase of respiration with intra-thoracic pressure changes influencing diastolic heart filling and therefore Q. Breathing in reduces intra-thoracic pressure, filling the heart and increasing Q, while breathing out increases intra-thoracic pressure, reduces heart filling and Q. This respiratory response is called stroke volume variation and can be used as an indicator of cardiovascular health and disease. These respiratory changes are important, particularly during mechanical ventilation, and Q should therefore be measured at a defined phase of the respiratory cycle, usually end-expiration.

Combined cardiac output

Combined cardiac output (CCO) is the sum of the outputs of the right and left side of the heart. It is useful in fetal circulation, where cardiac outputs from both sides of the heart partly work in parallel by the foramen ovale and ductus arteriosus, both directly supplying the systemic circulation.[67]

Cardiac input

Cardiac input (CI) is the inverse operation of cardiac output. As cardiac output implies the volumetric expression of ejection fraction, cardiac input implies the volumetric injection fraction (IF).

IF = end diastolic volume (EDV) / end systolic volume (ESV)

Cardiac input is a readily imaged mathematical model of diastole.

Example values

Measure Typical value Normal range
end-diastolic volume (EDV) 120 mL[68][non-primary source needed] 65–240 mL[68][non-primary source needed]
end-systolic volume (ESV) 50 mL[68][non-primary source needed] 16–143 mL[68][non-primary source needed]
stroke volume (SV) 70 mL 55–100 mL
ejection fraction (Ef) 58% 55–70%[69]
heart rate (HR) 75 bpm 60–100 bpm[70]
cardiac output (CO) 5.25 L/minute 4.0–8.0 L/min[71]

References

1. ^ Guyton, Arthur C.; John E. (John Edward) (2006). Textbook Of Medical Physiology (11th ed.). Philadelphia: Elsevier Inc. ISBN 0-7216-0240-1.
2. ^ Levy, Matthew N.; Berne, Robert M. (1997). Cardiovascular physiology (7th ed.). St. Louis: Mosby. ISBN 0-8151-0901-6.
3. ^ Rowell, Loring B. (1993). Human cardiovascular control. Oxford [Oxfordshire]: Oxford University Press. ISBN 0-19-507362-2.
4. ^ Braunwald, Eugene (1997). Heart disease: a textbook of cardiovascular medicine (5th ed.). Philadelphia: Saunders. ISBN 0-7216-5666-8.
5. ^ a b Binanay C, Califf RM, Hasselblad V, et al. (October 2005). "Evaluation study of congestive heart failure and pulmonary artery catheterization effectiveness: the ESCAPE trial". JAMA 294 (13): 1625–33. doi:10.1001/jama.294.13.1625. PMID 16204662.
6. ^ a b Pasche B, Knobloch TJ, Bian Y, et al. (October 2005). "Somatic acquisition and signaling of TGFBR1*6A in cancer". JAMA 294 (13): 1634–46. doi:10.1001/jama.294.13.1634. PMID 16204663.
7. ^ a b Hall JB (October 2005). "Searching for evidence to support pulmonary artery catheter use in critically ill patients". JAMA 294 (13): 1693–4. doi:10.1001/jama.294.13.1693. PMID 16204671.
8. ^ "Arteriovenous oxygen difference". Sports Medicine, Sports Science and Kinesiology. Net Industries and its Licensors. 2011. Retrieved 30 April 2011.
9. ^ Iberti TJ, Fischer EP, Leibowitz AB, Panacek EA, Silverstein JH, Albertson TE (December 1990). "A multicenter study of physicians' knowledge of the pulmonary artery catheter. Pulmonary Artery Catheter Study Group". JAMA 264 (22): 2928–32. doi:10.1001/jama.264.22.2928. PMID 2232089.
10. ^ Johnston IG, Jane R, Fraser JF, Kruger P, Hickling K (August 2004). "Survey of intensive care nurses' knowledge relating to the pulmonary artery catheter". Anaesth Intensive Care 32 (4): 564–8. PMID 15675218.
11. ^ a b c Phillips RA, Hood SG, Jacobson BM, West MJ, Wan L, May CN. Pulmonary artery catheter (PAC) accuracy and efficacy compared with flow probe and transcutaneous Doppler (USCOM): An ovine validation. Crit Care Res Prac 2012; doi:10.1155/2012/621496
12. ^ Finegold JA, Manisty CH, Cecaro F, Sutaria N, Mayet J, Francis DP. Choosing between velocity-time-integral ratio and peak velocity ratio for calculation of the dimensionless index (or aortic valve area) in serial follow-up of aortic stenosis. Int J Cardiol 2013;167(4):1524-31. doi: 10.1016/j.ijcard.2012.04.105
13. ^ Su BC, Yu HP, Yang MW, Lin CC, Kao MC, Chang CH, Lee WC. Reliability of A New Ultrasonic Cardiac Output Monitor in Recipients of Living Donor Liver Transplantation. Liver Transpl 2008;14:1029-1037
14. ^ Phillips RA, Paradisis M, Evans NJ, Southwell DL, Burstow DJ, West MJ. Validation of USCOM CO Measurements in Preterm Neonates by Comparison with Echocardiography. Crit Care 2006; 10(Supl1) 144
15. ^ Cattermole GN, Leung M, Mak PSK, Chan SSW, Graham CA, Rainer TH. The normal ranges of cardiovascular parameters in children measured using the Ultrasonic Cardiac Output Monitor. Crit Care Med 2011;38(9) 1875-1881
16. ^ Jain S, Vafa A, Margulies DR, Liu W, Wilson MT, Allins AD. Non-invasive Doppler ultrasonography for assessing cardiac function: can it replace the Swan-Ganz catheter? Am J Surgery 2008;196(Dec):961-968.
17. ^ Horster S, Stemmler HJ, Strecker N, Brettner F, Hausmann A, Cnossen J, Parhofer KG, Nickel T, Geiger S. Cardiac Output Measurements in Septic Patients: Comparing the Accuracy of USCOM and PiCCO. Crit Care Res Prac 2012; doi:10.1155/2012/270631
18. ^ Phillips RA, Lichtenthal PR, Sloniger JA, Burstow DJ, West MJ, Copeland JG. Noninvasive Cardiac Output Measurement in Heart Failure Subjects on Circulatory Support. Anesth Analg 2009;108:881-6
19. ^ Kager CCM, Dekker GA, Stam MC. Measurement of cardiac output in normal pregnancy by a non-invasive two-dimensional independent Doppler device. ANZ J Obs Gyn 2009 doi:10.1111/j.1479-828X.2009.00948.x
20. ^ Mythen, M.G.W., A. R., Perioperative plasma volume expansion reduces the incidence of gut mucosal hypoperfusion during cardiac surgery. Arch Surg, 1995. 130(4) p. 423-9.
21. ^ Sinclair, S.J., S.; Singer, M., Intraoperative intravascular volume optimisation and length of hospital stay after repair of proximal femoral fracture: randomised controlled trial. BMJ, 1997. 315(7113) p. 909-12.
22. ^ Conway, D.H.M., R.; Abdul-Latif, M. S.; Gilligan, S.; Tackaberry, C., Randomised controlled trial investigating the influence of intravenous fluid titration using oesophageal Doppler monitoring during bowel surgery. Anaesthesia, 2002. 57(9) p. 845-9.
23. ^ Gan, T.J.S., A.; Maroof, M.; el-Moalem, H.; Robertson, K. M.; Moretti, E.; Dwane, P.; Glass, P. S., Goal-directed intraoperative fluid administration reduces length of hospital stay after major surgery. Anesthesiology, 2002. 97(4) p. 820-6.
24. ^ Venn, R.S., A.; Richardson, P.; Poloniecki, J.; Grounds, M.; Newman, P., Randomized controlled trial to investigate influence of the fluid challenge on duration of hospital stay and perioperative morbidity in patients with hip fractures. Br J Anaesth, 2002. 88(1) p. 65-71.
25. ^ Wakeling, H.G.M., M. R.; Jenkins, C. S.; Woods, W. G.; Miles, W. F.; Barclay, G. R.; Fleming, S. C., Intraoperative oesophageal Doppler guided fluid management shortens postoperative hospital stay after major bowel surgery. Br J Anaesth, 2005. 95(5) p. 634-42.
26. ^ Noblett, S.E.S., C. P.; Shenton, B. K.; Horgan, A. F., Randomized clinical trial assessing the effect of Doppler-optimized fluid management on outcome after elective colorectal resection. Br J Surg, 2006. 93(9) p. 1069-76.
27. ^ Pillai, P.M., I.; Gaughan, M.; Snowden, C.; Nesbitt, I.; Durkan, G.; Johnson, M.; Cosgrove, J.; Thorpe, A., A double-blind randomized controlled clinical trial to assess the effect of Doppler optimized intraoperative fluid management on outcome following radical cystectomy. J Urol, 2011. 186(6) p. 2201-6.
28. ^ http://www.nice.org.uk/mtg3
29. ^ Lowe, G.D.C., B. M.; Philpot,E. J. ; Willshire, R. J., Oesophageal Doppler Monitor (ODM) guided individualised goal directed fluid management (iGDFM) in surgery - a technical review. http://www.deltexmedical.com, 2010.
30. ^ de Wilde RBP, Schreudder JJ, van den Berg PCM, Jansen JRC. An evaluation of cardiac output by five arterial pulse contour techniques during cardiac surgery" Anaesthesia 2007;62:760-768
31. ^ Wesseling KH, Jansen JR, Settels JJ, Schreuder JJ (May 1993). "Computation of aortic flow from pressure in humans using a nonlinear, three-element model". J. Appl. Physiol. 74 (5): 2566–73. PMID 8335593.
32. ^ a b Bein B, Meybohn P, Cavus E, Renner J, Tonner PH, Steinfath M, Scholz J, Doerges V. The Reliability of Pulse Contour-Derived Cardiac Output During Hemorrhage and after Vasopressor Administration. Anesth Analg 2007;105:107-13
33. ^ (Or Starling's Law)
34. ^ Singh and Taylor: The FloTrac Device should not be used to follow cardiac output in cardiac surgical patients. J Cardiothor Vasc Anes 2010;24(4) 709-711
35. ^ Manecke G: Intrinsic to all arterial waveform technologies is the measurement of Stroke Volume Variation (SVV) which predicts volume responsiveness and is used for managing fluid optimization in high risk surgical or critically ill patients. A Physiologic Optimization Program based on hemodynamic principles that incorporates the data pairs SV and SVV has been published.
36. ^ McGee WT: A simple physiologic algorithm for managing hemodynamics using stroke volume and stroke volume variations: physiologic optimization program. J Inten Care Med 2009;24(6) 352-360.
37. ^ Su BC, Tsai YF, Chen CY, et al: Cardiac output derived from arterial pressure waveform analysis in patients undergoing liver transplantation: Validity of a third generation device. Transplant Proc 44: 424-428, 2012
38. ^ Takala J, Ruokonen E, Tenhunen JJ, Parviainen I and Jakob SM. Early non-invasive cardiac output monitoring in hemodynamically unstable intensive care patients: A multi-center randomized controlled trial. Crit Care 2011;15:R148,doi:10.1186/cc10273
39. ^ Romano SM, Pistolesi M (August 2002). "Assessment of cardiac output from systemic arterial pressure in humans". Crit. Care Med. 30 (8): 1834–41. doi:10.1097/00003246-200208000-00027. PMID 12163802.
40. ^ Scolletta S, Romano SM, Biagioli B, Capannini G, Giomarelli P (August 2005). "Pressure recording analytical method (PRAM) for measurement of cardiac output during various haemodynamic states". Br J Anaesth 95 (2): 159–65. doi:10.1093/bja/aei154. PMID 15894561.
41. ^ Calamandrei M, Mirabile L, Muschetta S, Gensini GF, De Simone L, Romano SM (May 2008). "Assessment of cardiac output in children: a comparison between the pressure recording analytical method and Doppler echocardiography". Pediatr Crit Care Med 9 (3): 310–2. doi:10.1097/PCC.0b013e31816c7151. PMID 18446106.
42. ^ Scolletta S, Gregoric ID, Muzzi L, Radovancevic B, Frazier OH (21 November 2006). "Pulse wave analytical to assess systemic blood flow during mechanical biventricular support". Perfusion.
43. ^ Scolletta S, Romano SM, Maglioni E, et al. (2005). "Left ventricular performance by PRAM during cardiac surgery". Intensive Care Med 31 (Suppl 1): S157.
44. ^ [Bernstein, D. P. (2010). Impledance cardiography: Pulsatile blood flow and the biophysical and electrodynamic basis for the stroke volume equations. Journal of Electrical Bioimpedance, 1(1), 2-17.]
45. ^ [Costa, P. D., Rodrigues, P. P., Reis, A. H., & Costa-Pereira, A. (2010). A review on remote monitoring technology applied to implantable electronic cardiovascular devices. Telemedicine Journal and e-Health: The Official Journal of the American Telemedicine Association, ] doi:10.1089/tmj.2010.0082
46. ^ [Tang, W. H., & Tong, W. (2009). Measuring impedance in congestive heart failure: Current options and clinical applications. American Heart Journal, 157(3), 402-411.] doi:10.1016/j.ahj.2008.10.016
47. ^ [Ferrario, C. M., Flack, J. M., Strobeck, J. E., Smits, G., & Peters, C. (2010). Individualizing hypertension treatment with impedance cardiography: A meta-analysis of published trials. Therapeutic Advances in Cardiovascular Disease, 4(1), 5-16. ] doi:10.1177/1753944709348236
48. ^ [Moshkovitz, Y., Kaluski, E., Milo, O., Vered, Z., & Cotter, G. (2004). Recent developments in cardiac output determination by bioimpedance: Comparison with invasive cardiac output and potential cardiovascular applications. Current Opinion in Cardiology, 19(3), 229-237.]
49. ^ [Parry, M. J., & McFetridge-Durdle, J. (2006). Ambulatory impedance cardiography: A systematic review. Nursing Research, 55(4), 283-291]
50. ^ [Wang, D. J., & Gottlieb, S. S. (2006). Impedance cardiography: More questions than answers. Current Heart Failure Reports, 3(3), 107-113]
51. ^ [Ventura, H. O., Taler, S. J., & Strobeck, J. E. (2005). Hypertension as a hemodynamic disease: The role of impedance cardiography in diagnostic, prognostic, and therapeutic decision making. American Journal of Hypertension, 18(2 Pt 2), 26S-43S. ] doi:10.1016/j.amjhyper.2004.11.002
52. ^ http://www.sonosite.com/products/bioz-dx
53. ^ http://www.niccomo.com/en/
54. ^ Krivitski NM (July 1995). "Theory and validation of access flow measurement by dilution technique during hemodialysis". Kidney Int. 48 (1): 244–50. doi:10.1038/ki.1995.290. PMID 7564085.
55. ^ Tanke RB, van Heijst AF, Klaessens JH, Daniels O, Festen C (January 2004). "Measurement of the ductal L-R shunt during extracorporeal membrane oxygenation in the lamb". J. Pediatr. Surg. 39 (1): 43–7. doi:10.1016/j.jpedsurg.2003.09.017. PMID 14694369.
56. ^ Casas F, Reeves A, Dudzinski D, et al. (2005). "Performance and reliability of the CPB/ECMO Initiative Forward Lines Casualty Management System". Asaio J. 51 (6): 681–5. doi:10.1097/01.mat.0000182472.63808.b9. PMID 16340350.
57. ^ Tessitore N, Bedogna V, Poli A, et al. (November 2008). "Adding access blood flow surveillance to clinical monitoring reduces thrombosis rates and costs, and improves fistula patency in the short term: a controlled cohort study". Nephrol. Dial. Transplant. 23 (11): 3578–84. doi:10.1093/ndt/gfn275. PMID 18511608.
58. ^ van Loon M, van der Mark W, Beukers N, et al. (June 2007). "Implementation of a vascular access quality programme improves vascular access care". Nephrol. Dial. Transplant. 22 (6): 1628–32. doi:10.1093/ndt/gfm076. PMID 17400567.
59. ^ Krivitski NM, Kislukhin VV, Thuramalla NV (July 2008). "Theory and in vitro validation of a new extracorporeal arteriovenous loop approach for hemodynamic assessment in pediatric and neonatal intensive care unit patients". Pediatr Crit Care Med 9 (4): 423–8. doi:10.1097/01.PCC.0b013e31816c71bc. PMC 2574659. PMID 18496416.
60. ^ Funk DJ, Moretti EW, Gan TJ (March 2009). "Minimally invasive cardiac output monitoring in the perioperative setting". Anesth. Analg. 108 (3): 887–97. doi:10.1213/ane.0b013e31818ffd99. PMID 19224798.
61. ^ Arheden H, Stahlberg F (2006). "Blood flow measurements". In Roos, Albert de; Higgins, Charles B. MRI and CT of the cardiovascular system (2nd ed.). Hagerstwon, MD: Lippincott Williams & Wilkins. pp. 71–90. ISBN 0-7817-6271-5.
62. ^ Arheden H, Holmqvist C, Thilen U, et al. (May 1999). "Left-to-right cardiac shunts: comparison of measurements obtained with MR velocity mapping and with radionuclide angiography". Radiology 211 (2): 453–8. PMID 10228528.
63. ^ Razavi R, Hill DL, Keevil SF, et al. (December 2003). "Cardiac catheterisation guided by MRI in children and adults with congenital heart disease". Lancet 362 (9399): 1877–82. doi:10.1016/S0140-6736(03)14956-2. PMID 14667742.
64. ^ Kuehne T, Yilmaz S, Schulze-Neick I, et al. (August 2005). "Magnetic resonance imaging guided catheterisation for assessment of pulmonary vascular resistance: in vivo validation and clinical application in patients with pulmonary hypertension". Heart 91 (8): 1064–9. doi:10.1136/hrt.2004.038265. PMC 1769055. PMID 16020598.
65. ^ Petzina R, Ugander M, Gustafsson L, et al. (May 2007). "Hemodynamic effects of vacuum-assisted closure therapy in cardiac surgery: assessment using magnetic resonance imaging". J. Thorac. Cardiovasc. Surg. 133 (5): 1154–62. doi:10.1016/j.jtcvs.2007.01.011. PMID 17467423.
66. ^ Pennell DJ, Sechtem UP, Higgins CB, et al. (November 2004). "Clinical indications for cardiovascular magnetic resonance (CMR): Consensus Panel report". Eur. Heart J. 25 (21): 1940–65. doi:10.1016/j.ehj.2004.06.040. PMID 15522474.
67. ^ Walter F., PhD. Boron (2003). Medical Physiology: A Cellular And Molecular Approaoch. Elsevier/Saunders. p. 1197. ISBN 1-4160-2328-3.
68. ^ a b c d Schlosser, Thomas; Pagonidis, Konstantin; Herborn, Christoph U.; Hunold, Peter; Waltering, Kai-Uwe; Lauenstein, Thomas C.; Barkhausen, Jörg (2005). "Assessment of Left Ventricular Parameters Using 16-MDCT and New Software for Endocardial and Epicardial Border Delineation". Am J Roentgenol 184 (3): 765–773. doi:10.2214/ajr.184.3.01840765. Values:
• End-diastolic volume (left ventricular) – average 118 and a range of 68 – 239mL and
• End-systolic volume (left ventricular) – average 50.1 and range, 16 – 143 mL:
• Also, ejection fraction was estimated in this study to be average 59.9% ± 14.4%; range, 18 – 76%, but secondary source (see above) is used in this article instead.
69. ^ O'Connor, Simon (2009). Examination Medicine (The Examination). Edinburgh: Churchill Livingstone. p. 41. ISBN 0-7295-3911-3.
70. ^ Normal ranges for heart rate are among the narrowest limits between bradycardia and tachycardia. See the Bradycardia and Tachycardia articles for more detailed limits.
71. ^