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Cardiac output (Q) is the volume of blood being pumped by the heart, in particular by a ventricle in a minute. This is measured in dm3 min-1 (1 dm3 equals 1000cm3). All you need to measure cardiac output is a dixie cup a straw a needle and a stop watch.
Cardiac output is equal to the stroke volume (SV) multiplied by the heart rate (HR). SV is the volume pumped per beat and the HR is the number of beats per minute. Therefore, if there are 70 beats per minute, and 70 ml blood is ejected with each beat, (SV), the cardiac output (Q) is 4900 ml/minute. This value is typical for an average adult at rest, although Q may reach up to 30 litres/minute during extreme exercise by elite athletes.
The function of the heart is to transport the 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. Q indicates how well the heart is performing this function. Q is regulated principally by the demand for oxygen by the cells of the body. If the cells are working hard, with a high metabolic oxygen demand then the Q is raised to increase the supply of oxygen to the cells, while at rest when the cellular demand is low, the Q is said to be baseline. Q is regulated not only by the heart as it pumps, but also 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 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, whilst SV can vary between 70 and 120 ml, a factor of only 1.7. 
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
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. Q measurement, if it was accurate and non-invasive, would be adopted as part of every clinical examination from general observations to the intensive care ward, and would be as common as simple blood pressure measurements are now. Such practice, if it were adopted, may revolutionise the treatment of many cardiovascular diseases including hypertension and heart failure. This is the reason why Q measurement is now an important research and clinical focus in cardiovascular medicine.
Measuring Cardiac Output
Circulation is a critical and variable function of human physiology and disease. An accurate and non-invasive measurement of Q is the holy grail of cardiovascular assessment. This would allow continuous monitoring of central circulation and provide improved insights into normal physiology, pathophysiology and treatments for disease. 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.
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 by the phase of respiration with intra-thoracic pressure changes influencing diastolic filling and therefore Q. This is important especially during mechanical ventilation, and Q should therefore be measured at a defined phase of the respiratory cycle (typically end-expiration).
The Fick Principle
The Fick principle was first described by Adolph 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 pick 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:
From these values, we know that:
VO2 = (Q x CA) - (Q x CV)
where CA = Oxygen concentration of arterial blood and CV = Oxygen concentration of venous blood. This allows us to say
Q = (VO2/CA - CV)*100
and therefore calculate Q. While considered to be the most accurate method for Q measurement Fick is invasive and 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 (Innacor, Innovision A/S, Denmark).
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):
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.
The PAC is balloon tipped and is inflated to occlude the pulmonary artery. The PAC thermodilution method involves injection of a small mount (10ml) of cold saline at a known temperature into the pulmonary artery and measuring the temperature a know distance away (6-10cm).
The Q can be calculated from the measured temperature curve (The “thermodilution curve”). High Q will change the temperature rapidly, and low Q will change the temperature slowly. Usually three or four repeated measures are averaged to improve accuracy. However it is complex to perform and there are many sources of inaccuracy in the method. Modern catheters are fitted with a heating filament which intermittently heats and measures the thermodilution curve providing serial Q measurement.
PAC use is complicated by infection, Pulmonary artery rupture, cardiac tamponade, and air embolism. Recent studies suggest use of the PAC is both dangerous and expensive, and it may not improve patient survival or treatment. PAC use is in decline as clinicians move to less invasive, more effective technologies for monitoring haemodynamics.
Doppler Ultrasound Method
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:
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 uses a conventional ultrasound machine and a combined two dimensional (2D) and Doppler approach to measure Q. 2D measurement of the diameter (d) of the aortic annulus allows calculation of the flow CSA which is then multiplied by the vti of the Doppler flow profile across the aortic valve to determine the flow volume or SV. Multiplying SV by HR produces Q. Echocardiographic measurement of flow volume is clinically well established and of proven accuracy but requires training and skill, and may be time consuming to perform effectively. The 2D measurement of the aortic valve diameter is challenging and associated with significant error, while measurement of the pulmonary valve to calculate right sided Q is even more difficult.
Transcutaneous Doppler: USCOM
An Ultrasonic Cardiac Output Monitor (USCOM) (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 so both the right and left sided Q can be measured. Real time Automatic tracing of the Doppler flow profile allows for beat to beat right and left sided Q measurement. Importantly this single method can be used in neonates, children and adults for low and high Q measurement.
Transoesophageal Doppler: TOD
Transoesophageal Doppler (TOD), also known as esophageal Doppler monitor (EDM), supports a CW sensor on the end of a probe which can be introduced via the mouth or nose and positioned in the oesophagus so the Doppler beam aligns with the descending thoracic aorta (DTA) at a known angle. Because the transducer is close to the blood flow the signal is clear, however the alignment, and thus reliable signal, can often be difficult to maintain during respiration and patient movement. This method has good validation, particularly for measuring changes in blood flow, but is limited in that it only measures the DTA flow and not true Q and is therefore influenced by non-linear changes in Q and SVR. Additionally this method requires patient sedation and is accepted for use only in adults and large children.
Pulse Pressure Methods
Pulse Pressure (PP) methods measure the pressure in the arteries over time to derive a waveform and use this information to calculate cardiac performance. The problem is that any measure from the artery includes the changes in pressure associated with changes in arterial function.
Physiologic or therapeutic changes in vessel diameter will be 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 beat to beat.
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). The pressure in the arteries, measured by sphygmomanometry, is often used as a guide to the function of the heart. Put simply, the pressure in the heart is conducted to the arteries, so the arterial pressure approximately reflects the function of the heart or the Q.
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. Unfortunately the heart and vessels can function independently and sometimes paradoxically so that changes in the PP may both reflect and mask changes in Q. So these measures represent combined cardiac and vascular function only. Another similar system that uses the arterial pulse is the pressure recording analytical method (PRAM).
Invasive PP 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 or the position of the catheter tip will 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. 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 dilution, again using the Stewart-Hamilton principle. Lithium dilution uses a peripheral vein to a peripheral arterial line; however, it does not provide information on cardiac filling volumes and extravascular lung water. Calibration measurements cannot be performed too frequently, 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.
Uncalibrated PP - FloTrac
This technology involves inserting a manometer tipped arterial catheter into the mid flow portion of an artery, usually radial or femoral, and then by time domain sampling converts the arterial PP to Q. While this method involves one less line than the calibrated PP Q systems, it remains un-calibrated and so is only measuring arterial PP invasively. While it estimates upstream Q, any independent changes in Q and SVR cannot be detected by this method. This method has yet to be extensively evaluated, but early studies suggest that it may be useful in stable, normal subjects.
Impedance cardiography (ICG) is a method which calculates Q from the measurement of changes in impedance across the chest over the cardiac cycle. Lower impedance indicates greater the intrathoracic fluid volume, and as the only fluid volume which changes beat to beat within the thorax is the blood, the change in impedance can be used to calculate the SV and, combined with HR, the Q. This technique has progressed clinically (often called BioZ, i.e. biologic impedance, as promoted by the leading manufacturer in the US) and allows non-invasive estimations of Q and total peripheral resistance using only 4 paired skin electrodes.
While the method is desirably non-invasive and inexpensive, it has not achieved the reliability and reproducibility required of a useful clinical tool, and the evolution of algorithms to convert impedance signals to Q across a variety of outputs and in a variety of diseases continues.
Magnetic Resonance Imaging
Velocity encoded phase contrast Magnetic Resonance Imaging (MRI) 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, and less variable than both the Fick principle and thermodilution.
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 quantify the stroke volume in real time on a beat-for-beat basis.
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.
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, and conversely decreased resistance increases Q.
This can be explained mathematically:
By simplifying D'arcy's Law, we get the equation that
When applied to the circulatory system, we get:
Where MAP = Mean Aortic (or Arterial) Blood Pressure in mmHg,
TPR = Total Peripheral Resistance in dynes-sec-cm-5.
However, as MAP>>RAP, and RAP is approximately 0, this can be simplified to:
For right heart Q ≈ MAP/PVR For 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:
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 filing 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.
|This article is licensed under the GNU Free Documentation License. It uses material from the Wikipedia article "Cardiac_output". A list of authors is available in Wikipedia.|