1. The system is a closed circle rather than being open-ended and linear.

    

2. The system is elastic rather than rigid.

    

3. The system is filled with liquid at a positive mean pressure ("mean cardiovascular pressure"), which exists independent of the pumping action of the heart.

    

4. The right and left ventricles, which pump into the same system that they pump out of, are in series with two interposed vascular beds (systemic and pulmonary).

    

5. The heart fills passively, rather than by actively sucking.

    

6. As a consequence of the heart's passive filling, the circulation rate is normally regulated by peripheral-vascular factors, rather than by cardiac variables.

    

7. The flow from the heart is intermittent, while the flow to it is continuous.

    

8. Normally, there is an excess expenditure of energy by the heart needed for the circulation rate imposed by peripheral vascular regulators ("pump energy excess").

    

9. Normally, ventricular capacity is in excess of the diastolic filling volume ("pump capacity excess").

    

10. The slowing effect of any vascular resistance on flow rate depends on its location, with reference to upstream compliance, as well as its magnitude.

     

In order to understand circulatory phenomena in the elastic circular hydraulic system (Fig. 1), where every point is both upstream and downstream from every other point, and where the non-sucking ventricles pump out of the same system they pump into, we need to examine the ten unique factors individually, before we can amalgamate them into a meaningful whole.

To clarify and facilitate understanding of the features peculiar to the heart, it is helpful to compare three major types of pumps. The heart�s unique characteristics as a pump are of paramount importance in understanding how the cardiovascular system works.

PUMP TYPE #1: This type of pump both sucks and forcibly ejects fluid. This pump uses energy both to actively fill at its inlet and to empty its contents at its outlet. Two examples of this type of pump are: (1) a piston pump, which expends energy to suck in the stroke volume which is then forcibly ejected; and (2) a roller pump, which sucks at its inlet by the recoil of the resilient tubing that has been compressed by the roller as it moves forward, ejecting the fluid in front of it. With type #1 pumps, the output in a hydraulic system is determined exclusively by two pump variables: the stroke rate and the stroke volume. The reason for discussing this is that the heart is a different type of pump, and those two variables are frequently and erroneously projected onto cardiovascular function, in a way in which they do not apply.

PUMP TYPE #2: This pump sucks and blows but, instead of producing a specific flow rate, creates a specific pressure gradient between its inlet and outlet. Centrifugal pumps fall into this category. With this type, two pump factors (rate and power), as well as two non-pump factors (pressure and resistance in the system), effect output. As the heart is not such a pump, there is danger in borrowing explanations of cardiovascular function from hydraulic systems containing centrifugal pumps.

PUMP TYPE #3: This type of pump is passive filling, and does not suck at its inlet. It expends no energy to fill, it only expends energy to empty. An example of this type of pump is the urinary bladder. It is a flaccid, hollow organ that does not create any negative pressure or suck on the ureters or kidneys to fill. The bladder merely exerts energy to empty. To calculate the flow of urine for any given period, you can obtain the answer by multiplying the stroke rate times the stroke volume. However, it is important to underline here that those two things, stroke rate and volume, are not determinants of bladder output. You cannot increase the output merely by changing the rate of bladder contraction (stroke rate). The urinary bladder cannot expend energy to increase its filling and thus its stroke volume. This type of pump, even though it does the work and thus produces the flow, is totally dependent upon external factors (e.g., renal function) to determine the output. At a given rate of urinary production, stroke rate and stroke volume are reciprocals of one another. If the bladder is emptied twice as often, the stroke volume will be one half as much.

THE RIGHT AND LEFT VENTRICLES ARE TWO TYPE #3 PUMPS:

The heart, like the urinary bladder, is a hollow muscular organ that does not suck to fill, but produces circulation by ejecting whatever fluid enters it at diastole. During normal function, the heart not only doesn't develop a pressure negative to the intrathoracic pressure, but it offers an impediment to filling because of its limited volume-pressure compliance.

The evidence that the heart fills passively, and does not suck to fill, is found in the data from innumerable heart catheterizations, which all show a positive diastolic pressure in the ventricles (Fig. 2, a to c). In fact, as ventricles fill, they not only do not suck, but they offer an increasing impediment to filling, as noted by the progressive increase in pressure toward the end of diastole (Fig. 2, c). Negative ventricular pressure, in relation to intrathoracic pressure, has not been found in physiologic states. The heart, like all other muscles in the body, expends energy and does work by contracting. It cannot expend energy to do mechanical work by forcibly elongating its fibers to suck like the type #1 pump, which uses energy to suck in a stroke volume. Do not confuse the negative pressure created in the chest by inspiration as negative heart pressure (Fig. 2, b). The chest can suck, the heart cannot.

One can calculate cardiac output in the same way as in the urinary bladder example: by multiplying stroke rate times stroke volume. Also, just as in that example, while stroke rate multiplied by stroke volume measures the amount of output, those variables are not determinants of that output. The heart, by filling passively, pumps out blood at a rate determined by the rate of blood coming to it. Given that the heart is a pump that produces the flow but has a flow rate determined by extra-cardiac factors, let us examine those factors that determine circulation rate.

With pumps that cannot suck to fill, there must be a positive pressure at the inlets for any blood to run into the ventricles, in order for there to be any pump output. If there is no pressure in the cardiovascular system, no blood can run into the ventricles and there can be no flow. Normally, there is a mean cardiovascular pressure above zero, which the heart distributes. The heart, rather than being responsible for the pressure in the vascular system, is a circulating device. It lowers the pressure at the ventricular inlets and raises it at the ventricular outlets. With a positive pressure in the cardiovascular system, when blood is ejected into the arterial side of the circle, a pressure gradient is created between the arteries and veins. This gradient causes blood to flow around the circle back to the ventricular inlets. Therefore, the output rate varies directly with the magnitude of that mean cardiovascular pressure. The higher the pressure, the higher the gradient, the greater the flow rate. The circular system being elastic, and having resistance and other impediments to flow, the energy from ventricular contraction does not transfer instantaneously around the circle after each heart beat, as would occur in a rigid system. The energy of venous flow is several heart beats behind that of ventricular ejection. The greater the elasticity and impediments to flow, the slower the flow rate.

Therefore, during normal function, cardiac output varies directly with the mean cardiovascular pressure and inversely with the impedance to blood flow to the heart.

I. MEAN CARDIOVASCULAR PRESSURE

Definition: The mean cardiovascular pressure is the pressure in the cardiovascular system with the circulation stopped, after the pressure has equalized between the arteries, capillaries, veins, and cardiac chambers. Do not confuse this pressure with central venous pressure, venous filling pressure, or mean-arterial pressure. Mean cardiovascular pressure is the pressure related to the blood volume and the compliance of the entire elastic cardiovascular compartment.

Measurement: Mean cardiovascular pressure is expressed in centimeters of water above ambient pressure, with zero being at mid-heart level. The mean cardiovascular pressure can be approximated during cardiac arrest. After arrest, a pressure equalization occurs between the various cardiovascular compartments in about thirty seconds. The arterial pressure falls and the venous pressure rises as some of the arterial blood moves into the veins during pressure equalization. Therefore, the mean cardiovascular pressure is always above venous pressure and below arterial pressure. Normally, mean cardiovascular pressure is between 15 and 18 cm. of water above mid-heart level. We have some approximation of its magnitude from fortuitous records of the arterial and venous pressures equalizing, obtained during short periods of cardiac arrest of patients in coronary care settings, in emergency rooms, or in operating rooms during heart surgery. Even in these situations, the resulting pressure can be regarded as only an approximation, as the shift in fluid, from hypoxia caused by lack of circulation, may have altered the vascular compliance. (See Appendix for clinical measurement technique)

Significance: Without mean cardiovascular pressure there would be no circulation. The heart doesn't generate the pressure in the vascular system, it merely distributes the mean cardiovascular pressure. The cardiac ventricles take the mean cardiovascular pressure and distribute it by raising the pressure on the arterial sides while lowering it on the venous sides. The two ventricles, being passively filling pumps, cannot suck, so they lower the inlet pressures toward — but never below — zero, in relation to the ambient pressure in the chest.

The higher the mean cardiovascular pressure, if the ventricles are not failing, the higher the ventricles can elevate the arterial pressure while reducing the venous pressure toward zero, and thus the greater the cardiac output. Conversely, the lower the mean cardiovascular pressure, the less the heart can raise the arterial pressure by lowering the venous pressure, and thus the lower the cardiac output.

Origin of the mean cardiovascular pressure: The mean cardiovascular pressure results from the volume of blood and the compliance of the cardiovascular system. The volume in the cardiovascular system results from an equilibrium between the rate of water, electrolytes, and other blood constituents entering the body by way of the gastrointestinal tract, and leaving the body primarily by the kidneys (Fig. 3). The mean cardiovascular pressure is the result of a continuing dynamic process.

Homeostatic maintenance of normal mean cardiovascular pressure:

(1) Slow feedback mechanism:

A slow homeostatic feedback mechanism tends to keep the mean cardiovascular pressure at a constant level: Elevation of the mean cardiovascular pressure above normal —› causes increase in cardiac output —› causes increased renal blood flow —› results in increased renal output —› thereby lowering blood volume and mean cardiovascular pressure back to normal. Conversely, low mean cardiovascular pressure —› low cardiac output —› low renal blood flow —› decreased renal output until the mean cardiovascular pressure is restored to normal by continuing fluid intake. With elevated mean cardiovascular pressure, the rate of return to normal is dependent on renal function, whereas, with low mean cardiovascular pressure the rate of return can vary greatly, depending on the rate of restoration of blood volume.

Clinical evidence of the homeostatic mechanism:

 

• Response to weightlessness by going into orbit: In the absence of gravity, normal blood volume shifts centrally from the lower part of the body, thereby, increasing the mean cardiovascular pressure at heart level —› resulting in increased cardiac output —› causing greater renal blood flow —› leading to greater urinary output —› causing a decrease in blood volume and, thus, a decrease in mean cardiovascular pressure back to normal. The converse is found when astronauts return to gravity. It takes a few hours to restore normal mean cardiovascular pressure by intake of fluid and electrolytes after returning to earth, during which transition time they conserve the fluid they take in by putting out very little urine.

   

• During any hypovolemic shock state, such as massive hemorrhage or severe dehydration, the urinary output goes abruptly down and remains low until restoration of normal mean cardiovascular pressure.

(2) Rapid mean cardiovascular pressure buffer mechanisms:

(a) Elasticity: The elasticity of the vascular system prevents sudden blood volume loss or gain from causing a linear, temporary change in mean cardiovascular pressure. Elasticity has, of course, an instantaneous buffer effect. Evidence of this is found in one's ability to give a pint of blood at the blood bank without going into severe low cardiac output. This buffer effect bolsters circulation while the blood volume — and, thus, mean cardiovascular pressure — is restored by subsequent oral intake of fluid.

(b) Vascular/extravascular equilibrium: There is a pressure equilibrium between the various extravascular compartments of the body and the cardiovascular space (Fig. 3). Changes in mean cardiovascular pressure result in shifts of fluid back and forth which tend to buffer sudden changes. This buffer system is fairly rapid but not instantaneous, as evidenced by the observation that a person going into severe shock from sudden loss of blood would not have had the same severe shock state if the loss had occurred more gradually over a period of an hour or so.

(3) Humeral and neuro-muscular-vascular reflexes:

These responses from stimuli, which alter vascular compliance, act as buffer systems. They prevent sudden changes in mean cardiovascular pressure from sudden position changes, such as going from lying to standing. They also buffer the effect of sudden loss of blood volume from hemorrhage.

II. IMPEDANCE TO THE FLOW OF BLOOD FROM THE OUTLETS AROUND TO THE INLETS OF THE VENTRICLES

Four factors tend to impede the flow of blood in the cardiovascular circle. Therefore, they are inverse determinants of cardiac output: (1) resistance, (2) elasticity, (3) limited compliance of the ventricles to filling, and (4) inertia of intermittent blood flow to the ventricles. The interrelationship of these four factors makes flow determination much more complicated than plain resistance, which is the single impediment in rigid, open ended, linear hydraulic systems.

(1) Resistance and (2) Elasticity

Unlike rigid linear hydraulic systems, where a given resistance may have the same effect on flow, irrespective of its location, the elasticity of the vascular system makes location of the resistance a significant parameter. Because of the elasticity of the circle, the location of any particular resistance determines to what extent that resistance has on impeding blood flow. A given resistance may have little or no effect in determining cardiac output if it is near the outlet of the ventricles, yet the same magnitude of resistance may have tremendous slowing effect on circulation if located near the inlets of the heart. Resistance points that have little compliant vascular bed "upstream" (arteriolar resistance), increase pump work but may not affect cardiac output significantly. The heart, except during failure, exerts enough energy to force the blood past any resistance near its outlet, with no hold up in flow. On the other hand, resistance located with a large compliant bed upstream has a tremendous effect on flow rate by slowing blood return to the ventricles. Venous sided resistance, with the compliance of the whole vascular bed upstream, is a major inhibitor of circulation rate. Thus, venous sided resistance is a major determinant of cardiac output, but arterial resistance is not. Interpolation of the additive effect of all resistance points in the circle on cardiac output must include the amount of compliant bed upstream from each resistance point.

All resistance factors — including blood viscosity, cross-section area of any vascular bed, margination of blood constituents, etc. — play roles in flow rate determination only when linked with their location to compliance upstream. Adding venous sided resistance of a magnitude that results in only a few mm. water pressure gradient may cause a marked reduction of flow. Whereas, increasing arteriolar resistance to the point of severe arterial hypertension may not appreciably change cardiac output.

In human cardiovascular physiology, the significance of resistance can only be understood when coupled with upstream compliance. It is not just the magnitude of resistance but the location of that resistance that determines its effect on circulation rate.

(3) Impediment to Ventricular Filling

The end-diastolic pressure: Ventricles not only do not suck to fill, they offer an impediment to filling. The left ventricle has thicker and stiffer walls than the right, so it tends to retard filling more than the right ventricle. Catheterization data shows end-diastolic pressure of five to ten centimeters of water above intrathoracic pressure. Don't confuse the negative pressure of inspiration, transmitted to the heart, as the heart sucking. The intracardiac pressure is always above that intrathoracic pressure (Fig. 2 at C).

(4) Inertia of Intermittent Blood Flow Offset by "The Atrial Effect"

Atrial function facilitates circulation by preventing the retarding effect that would otherwise occur from the intermittent inflow to the intermittent outflow ventricles. By being partially empty and distensible, atria prevent the interruption of venous flow to the heart that would occur during ventricular systole if the veins ended at the inlet valves of the heart.

Atria have four essential characteristics that cause them to promote continuous venous flow. (1) There are no atrial inlet valves to interrupt blood flow during atrial systole. (2) The atrial systole contractions are incomplete and thus do not contract to the extent that would block flow from the veins through the atria into the ventricles. During atrial systole, blood not only empties from the atria to the ventricles, but blood continues to flow uninterrupted from the veins right through the atria into the ventricles. (3) The atrial contractions must be gentle enough so that the force of contraction does not exert significant back pressure that would impede venous flow. (4) The "let go" of the atria must be timed so that they relax before the start of ventricular contraction, to be able to accept venous flow without interruption.

By preventing the inertia of interrupted venous flow that would otherwise occur at each ventricular systole, atria allow approximately 75% more cardiac output than would otherwise occur. The fact that atrial contraction is 15% of the amount of the succeeding ventricular ejection has led to the false conclusion that atria have their benefit by pumping up the ventricles (the so-called "atrial kick"). The real benefit is in preventing inertia and allowing uninterrupted venous flow.

The 20% to 25% increase in cardiac output from synchronized atrial function over that of atrial fibrillation doesn't belie the 75% contribution of the atrial effect, as atrial fibrillation eliminates only part of that effect. Atrial compliance, elasticity, and gravity help in emptying the atria at ventricular diastole during atrial fibrillation. Also, cardiac output during atrial dysfunction is buffered: the initial fall in circulation rate during atrial fibrillation reduces renal flow, thereby causing retention of water and the subsequent rise in mean cardiovascular pressure, which then partially offsets the slowing effect on circulation.

Thus, four factors impede the flow of blood around the cardiovascular circle: (1) resistance with (2) upstream compliance, (3) ventricular non-compliance, and (4) inertia if there is intermittent venous flow. The combined effect of these factors that impede the flow of blood to the inlet of the ventricles and, therefore, determine cardiac output in a negative way, will be referred to as: inlet impedance.

The passive filling characteristic of the ventricles is the feature that accounts for the automatic maintenance of blood volume equilibrium between the pulmonary and systemic vascular beds. With passively filled pumps, the relative blood volume in the two circuits is determined by their relative size and elasticity. On the other hand, if the ventricles were type #1 pumps, which actively fill, any discrepancy in the output of the two pumps, or flow through normal physiological shunts between the two vascular systems, would very quickly shift all of the blood volume into one circuit at the expense of the other, resulting in disaster.

The right and left ventricles, by filling passively, pump out whatever amount comes to them, determined by extracardiac factors. The blood volume equilibrium, therefore, is determined by the relative size and compliance of the two circuits and, to a minor extent, by the relative impediment of flow to the two pumps. It has been noted that the output of the two sides of the heart is never equal because there are physiological shunts, which connect one vascular bed to the other. The largest of these shunts, in normal physiology, are the bronchial arteries, which go from the systemic circuit to the lungs. The bronchial blood flow is a left-to-right shunt that accounts for the left ventricular output normally being at least 10% larger than the right. Because of the shunt, more blood returns to the left ventricle than the right, the left ventricle passively fills more than the right, thereby causing it to produce a greater output and thus the equilibrium is maintained. With other types of pumps this would not occur.

A dramatic illustration of this volume equilibrium, automatically being maintained during a massive discrepancy in output of the two pumps, is seen in atrial septal defects. In this case, the right ventricular output may be four or five times that of the left. Yet, the volume equilibrium is maintained. A large atrial septal defect virtually results in a single atrium above the two ventricles. The shunt occurs during ventricular diastole, because the right, thin walled ventricle is more distensible than the non-compliant thicker walled left ventricle. The blood in the common atrium goes the way of least resistance. The greater filling into the more compliant right ventricle results in greater right ventricular output. The greater right output goes to the lungs and then directly back to the right ventricle, returning again through the septal defect. Because of the passive filling, this results in no progressive increase in the blood volume in the lungs, and no disturbance in the maintenance of the blood volume equilibrium. After closure of the septal defect, resulting in much smaller right heart output, the volume equilibrium remains. This equilibrium, which persists after such sudden, massive changes in right heart output, occurs automatically because of the passive filling characteristic of the ventricles. It is the physical characteristics of the two vascular beds (e.g., relative size, compliance, and impediment to flow), that determines the volume balance with passive filling ventricles.

Normally, the passively filling ventricles are not maximally filled at diastole. Also, they are exerting excess energy over that needed to eject blood at the flow rate of blood entering the ventricles. With such volume and energy reserve, if blood enters the heart faster, the flow rate can go up without any change in strength of contraction or heart rate, unless the heart is being filled maximally (heart failure). Conversely, if blood enters the ventricles at a slower flow rate, the output will go down irrespective of any lowering of the rate or decrease in strength of cardiac contraction.

Thus, at a given flow rate, with the normal excess pump power, in both strength and rate, the heart rate and stroke volume become reciprocals of one another, as long as the heart doesn�t go into failure. At a given flow rate, decreasing the heart rate increases the stroke volume; while increasing the heart rate decreases the stroke volume. The observation — that cardiac output usually parallels the cardiac rate and strength of contraction — has resulted in the fallacious conclusion, based on post hoc ergo propter hoc reasoning ("following this, therefore because of this"), that they are cause and effect.

Even though these two variables do not normally control cardiac output, their variability has physiological benefit in energy conservation. During low circulation rates, when the heart is not filling to its maximal volume during diastole, the heart down shifts its rate and strength of contraction, thus conserving energy. Also, during high output states, it increases its rate and strength of contraction, thus preventing failure from any limitation of output that might be imposed by ventricular chamber size. If the heart functioned at its maximum strength of contraction and a rate of 150 beats a minute, the cardiac output would go up and down just as it does normally. But what a waste of energy would result! So the variability of rate and strength of contraction have only ecological value: that of conserving world food supply.

Two mechanisms cause the heart to roughly parallel its energy expenditure with cardiac output, preventing failure during high output states and saving energy during low output: (1) The neural and humeral stimuli that increase mean cardiovascular pressure are also those that increase heart rate and strength of contraction. (2) Increased stretching of the ventricles at diastole causes some increased strength of contraction at systole (Starling's law of the heart). While this observation of Starling is true and contributes to the paralleling of cardiac output to strength of contraction, it is probably not a major determinant of energy expenditure, as strong cardiac contractions continue even when the heart is completely empty, as seen during cardiac bypass surgery.

The reciprocal relationship of heart rate and stroke volume can be demonstrated in patients with complete heart block by varying the rate of firing of their pacemakers. Variations in rate, within physiological levels, above that required to prevent complete diastolic ventricular filling, are associated with no change in cardiac output. Therefore, a compromise rate of 78 is a commonly used setting that allows a wide range of activity acceptable to most patients.

Two factors cause the increase in cardiac output during exercise: Exercise causes the cardiovascular impedance to be decreased and the mean cardiovascular pressure to be increased. The intermittent skeletal muscle contractions around venous beds, which contain one-way valves, act as a peripheral pump which overcomes significant impedance to flow. Neuro-humeral reflexes speed the heart rate, which slightly lowers the inlet impedance to the heart, by lowering the end-diastolic pressure over what it would otherwise be. The increased heart rate guarantees excess energy expenditure, thus preventing cardiac power failure at the higher circulation rate. The neuro-humeral response to exercise also throws the vascular system into spasm, thereby increasing the mean cardiovascular pressure from the "G-suit" effect of tensing the body in general. These are temporary changes, lasting only during the exercise period, which do not effect the long term homeostatic control of circulation.

Heart failure is the only situation where the non-sucking heart determines its output. In this state, the heart is pumping at its maximum output and therefore, in effect, by limiting the output, it determines the output.

Normally, the heart rate and pump energy are in excess of that needed to eject enough blood at systole, so that the ventricles are empty enough at diastole to allow unobstructed passive filling. Normally ventricles are not maximally filled so there is reserve compliance which allows cardiac output to be determined by the extra-cardiac factors, rather than being limited by the heart. If the heart fails, it pumps less than the normally controlling extra-cardiac factors would dictate. During pump energy failure the heart's output is being limited by whatever amount it is able to pump. The heart in such a situation of failure, by limiting output, becomes the sole regulator of cardiac output.

Definition: Heart failure exists whenever heart function is inadequate to produce the output which would be allowed by the mean cardiovascular pressure and inlet impedance. Heart failure is present when cardiac function is limiting and, therefore, determining cardiac output. Normally, extra-cardiac factors being constant, if the heart contracts more strongly, efficiently, or rapidly, no increase in output occurs. Normally, the heart is exerting excess energy over that used to produce cardiac output and thus a reserve is present, so that increases in circulation rate are not impeded by the heart. Normally, at a given circulation rate, an increase in heart rate or increase in strength of cardiac contraction does not increase the output. However, during failure where the heart is pumping out the maximal circulation rate that it can produce at the moment, increases in the rate and strength of contraction do increase cardiac output, up until venous pressure is reduced to normal, at which time the heart is no longer in failure.

When heart failure occurs, blood volume equilibrium shifts to the system behind the failing ventricle with an increase in the venous pressure at the inlet of the failing ventricle. The other ventricle, in the presence of the slowed circulation rate, is able to maintain a low inlet pressure, with the resulting shift in blood volume equilibrium toward the circuit behind the failing ventricle. With left ventricular failure, for instance, the circulation slows to the rate limited by the left ventricle and a new equilibrium occurs with a shift of blood volume from the systemic circuit to the pulmonary vascular bed. Pulmonary edema occurs if this shift in equilibrium is severe.

The high end-diastolic ventricular pressure and high end systolic volume, from a low ejection fraction in heart failure, would also be factors limiting flow, if the heart weren't already determining a lower output.

A secondary change occurs when the left ventricle fails. The lowered cardiac output decreases renal blood flow, thereby causing retention of fluid and elevation of the mean-vascular pressure. Fluid equilibrium, which exists between the cardiovascular system and extra-vascular spaces, then results in edema, liver engorgement, and the other fluid abnormalities of congestive heart failure. Heart failure can intensify in a descending spiral if the blood volume progressively expands. The over-stretching of the heart puts it at a poor mechanical advantage, thus further increasing its failure (the descending limb of Starling's curve).

An extreme example of right heart failure is seen after a "Glen" right atrial-pulmonary artery anastomosis procedure for palliation of tricuspid atresia. Here, the person has virtually complete failure of a non-functioning right heart, with a surgical bypass of the right ventricle so that, in effect, the circulation is by the left ventricle alone. Circulation is slowed by the increase in impediment that normally would be overcome by the right ventricle. As a consequence of the slowed circulation, the kidneys receive slow flow, causing retention of fluid, resulting in a permanent elevation in the homeostatic level of mean cardiovascular pressure, resulting in some compensatory increase in cardiac output at the new equilibrium.

Causes of Heart Failure:

Cardiac causes (low myocardial power):

• Myocardiopathies � viral, arteriosclerotic, toxic, rheumatic
• Incompetent leaky valves, e.g., mitral or aortic
• Abnormally slow heart rate, e.g., complete heart block

Peripheral vascular causes:

• Abnormally high mean cardiovascular pressure causing potentially higher circulation rate than a normal heart is able to deliver, e.g., renal shutdown resulting in high output failure
• Abnormally low impedance to the heart, e.g., arteriovenous fistula

Signs and Characteristics of Heart Failure:

1. High venous pressure is the major pathognomonic finding in cardiac failure. If the heart were not in failure, the pressure would be normally low.

    

2. A new equilibrium occurs between the two vascular systems with a shift in volume to the field behind the failing ventricle. I.e., left ventricular failure shifts the equilibrium with a greater volume in the pulmonary circuit, resulting in x-ray evidence of pulmonary vascular engorgement and high pulmonary venous pressure. With right heart failure, the neck vein pressure is seen to be elevated. However, the systemic system initially is large enough to accommodate the shift in blood volume from the lungs without significant venous engorgement until secondary changes (as seen below) occur.

    

3. With either left or right heart failure, increased fluid accumulation in the body progresses in a descending spiral caused by slowed cardiac output —› less than optimum renal blood flow —› low renal output —› retained fluid in the vascular system —› increased mean cardiovascular pressure —› edema —› liver engorgement and ascites.

Pathologic Conditions Resulting in Heart Failure:

• Complete heart block with slow pulse
• Extremely rapid pulse
• Intermittent, pulsatile venous flow
• Low ejection fraction
• Myocardiopathy with weak myocardial contraction
• Myocardial infarction to the point of low ejection volume
• Valvular heart disease with stenosis or insufficiency
• Myocardial failure
• Pericardial tamponade
• Constrictive pericarditis

Goals in the Treatment of Heart Failure:

1. Make the heart contract stronger and faster.
2. Lower the resistance to ejection of blood from the ventricles.
3. Lower the mean cardiovascular pressure by salt and water restriction and diuretics
4. Restore the atrial effect if it is compromised by atrial fibrillation or nodal rhythym.

Shock is a low cardiac output state that results from low mean cardiovascular pressure. The low mean cardiovascular pressure can result from altering of one or the other of the two determinants of that pressure: (1) Increasing vascular compliance, and (2) Loss of blood volume.

1. Low output with resulting hypotension from sudden loss of vascular tone is frequently seen after inducing spinal anesthetics. A dramatic example was seen following total central nervous system anesthesia, induced when marcaine was inadvertently injected into the spinal canal during an attempt to do an intercostal nerve block. In this case, sudden total relaxation of the entire vascular system resulted in severe hypotension. Loss of vascular tone can be reversed by the use of vasopressors and the addition of fluid.

    

2. Shock from loss of blood volume can be temporarily compensated for by decreasing vascular compliance by vasopressor drugs, but ultimately needs to be corrected by blood volume restoration.

A combination of the two mechanisms of shock is seen in anaphylaxis, where there is both relaxation of the vascular system plus a shift in blood volume to the extra-vascular space, caused by increased capillary porosity.

Whatever the cause, the resulting low mean cardiovascular pressure in shock can be restored to normal by rapid infusion of fluid and electrolytes, aided by the use of vasopressor drugs.

During both of the shock mechanisms, the heart is already spending excess energy, so increasing the heart rate or strength of contraction will not increase the cardiac output. The obvious increase in output that follows the use of epinephrine, or other vasopressors, in such states is not caused by the inotropic effect on the heart, but by their effect of increasing the mean cardiovascular pressure by decreasing vascular compliance. Thus another temporary therapeutic measure, which increases mean cardiovascular pressure by making the system less compliant, is the surrounding of the vascular system by a "G-suit" such as those used by emergency medical personnel. It matters little how fast or how strongly the heart contracts: If there is no pressure in the cardiovascular system, there can be no circulation.

Arterial hypertension results from two radically different mechanisms. One is caused by increased arteriolar resistance in a system with normal mean cardiovascular pressure. The other results, in the presence of normal arteriolar resistance, from elevated mean cardiovascular pressure.

1. In arterial hypertension from increased arteriolar resistance, there is no slowing of the circulation rate from the resistance. The heart, which normally expends excess energy, just forces whatever blood that comes to it right past the resistance point. The only change in circulation is the elevated arterial pressure. Such arterial abnormal resistance may have its origin from transient neuro-humeral stimuli. However, with longstanding arterial hypertension, changes eventually occur, resulting in more permanent structural resistance increase. If, eventually, the vascular changes effect blood flow to the kidneys, the second type of hypertension may develop.

    

2. High mean cardiovascular pressure hypertension: The most graphic example of this state is seen during renal shutdown, from whatever cause. Here, if fluid and electrolyte intake continues, and there is reduced fluid leaving the body, the vascular volume progressively rises, with a corresponding increase in mean cardiovascular pressure. If the heart is strong enough that it continues to exert excess energy over that necessary to maintain a low ventricular inlet pressure, cardiac output increases. With the increased flow, even in the presence of unaltered arteriolar state, the resulting arterial pressure is elevated. If the heart is normal, but the pressure becomes so high that venous pressure rises, high output heart failure occurs. Severe high output failure has been seen where the systemic arterial pressure was 350/200, the systemic venous pressure was elevated with the neck veins at 30 cm., the pulmonary artery pressure was doubled, and the pulmonary capillary pressure was so high as to cause pulmonary edema with frothing of blood from the mouth and nose. This is in marked contrast to the distribution hypertension caused by arteriolar resistance increase.

The spectrum of hypertension states varies greatly between these two extremes and combinations of the two. Determination of the mean cardiovascular pressure helps delineate the significance of each in diagnosis. (See Appendix.)

Therapy for the two types of hypertension should be aimed at the etiology. For high mean cardiovascular hypertension, improving renal blood flow by removing renal arterial obstruction, diuretics, low salt diet, and water restriction to lower mean cardiovascular pressure or even renal transplant may be indicated. For hypertension caused by high arteriolar resistance, arteriolar vaso-relaxing drugs and changing the patient's response to stress may be effective.

An arteriovenous fistula, with blood flow going directly from an artery to a vein without going through arterioles, capillaries, and venules, bypasses most of the resistance that has upstream compliance. Therefore, depending on size, a fistula can cause tremendous increase in cardiac output. If a shunt is great enough, high output failure can result. If a shunt is in an extremity, where it can be occluded by manual pressure, alternating decrease and increase of cardiac output can be observed, as the shunt is intermittently occluded and opened. This maneuver is a good demonstration of the peripheral vascular control of circulation rate.

Arteriovenous fistula shunts occur in a number of disease states. For example, blood flow is shunted from arteries to veins in thyrotoxic goiters. Rather than there being one large arteriovenous channel, there are many small shunting vessels. The blood may be flowing so rapidly through the many small shunts that a loud, continuous murmur may be heard over the thyroid gland. This arteriovenous shunting of blood flow accounts, in part, for the high cardiac output associated with toxic goiter.

Because of the peripheral control of cardiac output, changing the size of the vascular bed, permanently (such as by a leg amputation) or by short term decrease (such as by cross-clamping the aorta and vena cava during surgery), causes an instantaneous decrease in cardiac output, mechanically, without any involvement or need of neurologic or humeral reflexes.

Patients with heart block, with a very slow pulse rate, have maximal ventricular filling before the completion of diastole, thereby restricting venous flow. This restriction in venous flow causes high venous pressure and low cardiac output. As the rate is incrementally increased by the use of a pacemaker, pressure gradually falls, until the ventricles are no longer maximally filled at diastole. Further increase in pacemaker rate above that level causes no more increase in output and no further drop in the venous pressure, as ventricular capacity is no longer restrictive to flow. A rate of 60 produces this plateau in most adults at rest and in a recumbent position. If the patient is active, the pacemaker rate may need to go to 80 or even 100 before further increases in pacemaker rate cause no corresponding increase in cardiac output. When the rate is fixed at 80 and the patient is resting, the output goes down normally with a waste of energy expended by the heart, the excess not being used for circulation. The rate of 80 allows a great variation in activity without exceeding a need for pulse increase to prevent failure. If the pacemaker is set at 120, such a rate would probably accommodate cardiac output for any violent activity. However, at rest, even though the output goes down to normal, the 120 pulse tachycardia is uncomfortable, the excess myocardial energy waste is great, and myocardial oxygen availability may be taxed. Therefore, fixed pacemaker rates are usually set at about 76 to 80 as a compromise between excess myocardial energy spent for circulation at rest and needed cardiac energy for strenuous activity.

From pacemaker experience we conclude that: (1) The heart, when not in failure, is always expending excess energy over that necessary to produce cardiac output. (2) The ability to lower pulse rate below maximum during periods of low cardiac output has implications for energy conservation rather than cardiac output.

Myocardial energy failure is a frequent finding during acute myocardial infarction and for short periods of time after coming off of bypass after open heart surgery. In both of these situations, left ventricular booster pumping has proved useful during a recovery period.

A competent, passive filling, pulsatile outflow, continuous inflow assist pump, placed between the left atrium and aorta will automatically lower the left atrial pressure and restore circulation rate to normal. A passive filling booster pump, which expends excess energy, put in parallel with the failing left ventricle, will make the combined output of the two normal. The combined output of two non-sucking pumps in parallel will act as one pump which produces excess to that needed for normal cardiac output. Their combined output will be controlled by the extra-cardiac factors that normally control circulation rate. Whatever amount the ventricle is unable to pump will run automatically into the booster ventricle and be ejected into the aorta. When the heart has recovered, the left atrial and arterial blood pressure will remain the same whether the pump is on or off. One method of weaning the patient off of the booster pumping is to raise the pump a few centimeters above the left atrium. Blood will then run preferentially into the heart's ventricle and be pumped out by it, if the heart is no longer in power failure. If the ventricle is still in failure, the venous pressure will rise only those few centimeters. Then whatever the heart does not pump the booster pump will.

From cardiac booster pumping, we conclude that the non-failing heart is expending excess energy to that needed at any moment to produce its output. Expenditure of energy by a parallel passive filling booster pump cannot increase the circulation rate above that dictated by extra-cardiac determinants.

The pumps have the four characteristics in common with the heart that allow them automatically, without pump regulation, to reproduce circulation under control of normal physiologic mechanisms: (1) The pumps fill passively and don't suck at their inlets. (2) The output of the pumps is pulsatile. (3) The pumps have atria that allow uninterrupted inflow to the intermittent outflow pumps. (4) The capacity of the pump ventricles is greater than any anticipated diastolic filling.

Many designs of pumps with the above characteristics could be made. One of the six variations that the author has used is shown in Figure 4. The pump consists of a flat atrial-ventricular silicone rubber tube with a reinforcing flat cotton cover (Fig. 5). The flat configuration, which prevents rebound to a round cross-section after being compressed, is responsible for the passive filling characteristic. The tube is mechanically compressed, sequentially, by four plates activated by cams and lifters (Fig. 6). Two narrow plates act as inlet and outlet valves to the ventricular portion of the tube. Two wide plates act as atrial and ventricular impellers (Fig. 4). The atrial impeller compression and sequence are critical in preventing interruption of venous inflow when the ventricles are being emptied. Therefore, atrial compression is incomplete, leaving the atria with a 3/16" channel at maximum compression. The slope of the atrial cams is gentle, to prevent any sharp rise in pump atrial pressure that might interrupt venous flow. The cam's timing is such that the "compression let go" occurs just before the inlet valve closes.

Cardiac surgery using both right and left heart bypass pumps has one significant advantage over conventional cardiopulmonary bypass: No oxygenator is used, as the lungs are left in the perfusion circuit and are provided with normal circulation. With normal pulmonary blood flow during the operation, post-operative pulmonary insufficiency is less of a problem than when the lungs are bypassed using an oxygenator.

By connecting the unique pumps in parallel to the heart's ventricles and with the pumps at heart level, cardiac function can be temporarily interrupted during surgical procedures without any interruption in normal circulation. When both the heart and bypass pumps are functioning simultaneously, the combined output remains the same as when the heart alone is pumping. Because both fill passively, the combined output remains regulated by the mean cardiovascular pressure and inlet impedance. Any amount that goes to the pumps doesn't go to the heart, and vice versa. If the heart function is then interrupted, by induced fibrillation or diversion of all the venous flow from one or both ventricles, the flow automatically goes to the pumps and circulation is uninterrupted and continues at the same rate. Subsequently, when the heart is defibrillated, or when the flow to the heart is no longer interrupted, the heart output takes over part of the circulation, with the total circulation rate still remaining the same. Weaning off of bypass is done simply by sequentially elevating the pumps a few centimeters at a time, thereby diverting more and more of the venous flow to the heart. Myocardial competence to take over the entire circulation becomes evident if the circulation rate and blood pressure remain unchanged after the incremental elevations are made. By appropriate elevation of the pumps above heart level, booster pumping can be done to maintain any desired atrial pressure in the heart during a recovery period.

The advent of open-heart surgery has provided a wealth of data and observations regarding cardiovascular physiology. Furthermore, our expanding ability to correct cardiovascular defects has been a stimulus to better understand how the system works. Delineation of the essential mechanical features of the heart has allowed construction of heart replacement pumps that: (1) automatically provide normal circulation during heart bypass, with the flow rate remaining under control of normal physiological mechanisms; (2) allow research, in animals, by isolating the peripheral vascular versus cardiac effect on circulation rate from responses to various drug, humeral, and neural stimuli; and (3) provide a mechanical model having the ten unique characteristics of the cardiovascular system, for teaching and the study of circulatory phonomena.

The pumps have the four characteristics in common with the heart that allow them automatically, without pump regulation, to reproduce circulation under control of normal physiologic mechanisms: (1) The pumps fill passively and don't suck at their inlets. (2) The output of the pumps is pulsatile. (3) The pumps have atria that allow uninterrupted inflow to the intermittent outflow pumps. (4) The capacity of the pump ventricles is greater than any anticipated diastolic filling.

Many designs of pumps with the above characteristics could be made. One of the six variations that the author has used is shown in Figure 4. The pump consists of a flat atrial-ventricular silicone rubber tube with a reinforcing flat cotton cover (Fig. 5). The flat configuration, which prevents rebound to a round cross-section after being compressed, is responsible for the passive filling characteristic. The tube is mechanically compressed, sequentially, by four plates activated by cams and lifters (Fig. 6). Two narrow plates act as inlet and outlet valves to the ventricular portion of the tube. Two wide plates act as atrial and ventricular impellers (Fig. 4). The atrial impeller compression and sequence are critical in preventing interruption of venous inflow when the ventricles are being emptied. Therefore, atrial compression is incomplete, leaving the atria with a 3/16" channel at maximum compression. The slope of the atrial cams is gentle, to prevent any sharp rise in pump atrial pressure that might interrupt venous flow. The cam's timing is such that the "compression let go" occurs just before the inlet valve closes.

Coronary Bypass Surgery:

Cardiac bypass during coronary surgery is used to provide a non-moving target with decompressed ventricles and the security that adequate circulation is being maintained during manipulation of the heart. Coronary artery surgery is not an open heart procedure, as the coronary arteries are on the surface of the heart. Therefore, closed bypass is applicable. Four cannulae are used. One pump is connected from the left atrium to the aorta, and the other from the right atrium to the pulmonary artery (Fig. 7).

It is important to start the left-sided bypass before the right. If the right one is started and the heart happens to fibrillate before the left bypass is started, blood will be pumped into the pulmonary circuit while no blood is leaving it. The result would be acute pulmonary engorgement ("liver lungs") with a fatal outcome. Likewise, the left-sided bypass should be discontinued last.

The heart is electrically fibrillated while the coronary-graft anastomoses are made. Then the heart can be defibrillated before the aorta-graft anastomoses are made. If, during the coronary-graft anastomoses, greater decompression of the ventricles is desired, the atrial cannulae can be advanced into the ventricles to drain them as well as the atria. Figure 8 shows arterial blood pressures staying relatively constant before and during a coronary bypass procedure, using right and left heart bypass.

During cardiac bypass with the two pumps, with the lungs functioning and without the use of an oxygenator, the anesthesiologist maintains ventilation and support of circulation just as he would in other non-cardiac procedures. Circulation rate reacts to blood loss, fluid infusion, and vasoactive drugs, just as when the heart is functioning.

Pulmonary Stenosis:

While it might appear that pulmonary stenosis could be repaired by using a right-sided bypass alone, one laboratory experience has demonstrated this to be too hazardous for clinical application. During the course of a right heart bypass, with the left ventricle functioning, the heart unexpectedly fibrillated. With the left ventricle suddenly not pumping, and before the right heart bypass pump could be turned off and the heart defibrillated, the lungs became irreparably overloaded with blood. The resulting "liver lungs" were not reversible, causing the death of the animal. Therefore, for safety, total heart bypass is always used, even when a right heart bypass alone could allow adequate access for correction of the lesion.

Use of a Passive Filling Pump with a Bubble Oxygenator:

During open-heart surgery, two pump bypass is precluded because of the difficulty in bypassing the left atrium, with its many pulmonary veins. Therefore, cardiopulmonary bypass with an oxygenator is necessary whenever the cardiac chambers need to be opened.

The passive filling, pulsatile output, continuous inflow pump has advantages over other types of pumps when used in a pump-oxygenator circuit. It can automatically produce normal circulation rate with a normal pulse wave, without any control adjustments, and without the hazard of oxygenator blood level fluctuations or air emboli.

If the inlet of the pump is placed at the priming level in the oxygenator (Fig. 9), because the pump is non-sucking, the blood will never fall below that level. Any blood that runs into the oxygenator that would tend to raise the level is automatically pumped back into the patient. The oxygenator is positioned at such a level that normal venous pressure is maintained by gravity drainage during bypass.

With this pump-oxygenator setup, circulation rate is determined by the patient's mean cardiovascular pressure and inlet impedance, just as it does in the intact body. During the bypass, circulation rate is modified by using vasoactive drugs, blood and fluid replacement to change the mean cardiovascular pressure, not by pump alteration.

This technique — beginning partial bypass in parallel with the heart — results in normal circulation rate, as both the heart and the extra-corporeal system fill passively. When complete bypass is produced, by occluding the vena cava around the caval catheters, the circulation rate remains the same. Terminating bypass is very simple. The occluding tapes around the vena caval tubes are released and the venous drainage tubing is occluded in increments. As more and more blood is diverted to the heart, the venous pressure will indicate whether or not further decrease in extra-corporeal pumping will be tolerated. In this way, the pump-oxygenator can be used as an automatic booster device during the myocardial reovery period, following the cardiac repair procedure. Figure 10 shows the blood pressure before, during, and after bypass, including the parallel bypass at the end of the procedure, in a patient with wide-open aortic insufficiency. The wide, abnormal pulse wave of aortic insufficiency is followed with a normal pulse wave while on bypass and after valve replacement. Figure 11 shows typical pulse waves during bypass and Figure 12 shows superimposed waves during circulation produced by parallel pumping of the heart and extra-corporeal pump.

Figure 13 illustrates a safety feature of the passive filling pump used with an oxygenator. Bypass was just underway, during a mitral valve replacement, when the inferior vena cava cannula slipped back into the right atrium. The inferior vena cava being temporarily occluded by the circumferential tape, caused the venous drainage to the pump to drop to one liter per minute. The oxygenator did not run out of blood and no air was pumped into the patient during this low output episode. After the tube was re-advanced properly into the inferior vena cava, the blood flow to and from the pump automatically returned to normal.

The Non-Sucking Pump Used with a Membrane Oxygenator:

The non-sucking heart replacement pump is ideal for use with a membrane oxygenator. It automatically allows a constant blood volume in the system with this closed-volume oxygenator. There is no need for a reservoir for volume monitoring of changing blood levels, which would occur and require alteration of pump rate with other types of pumps. In effect, the passive filling, continuous inflow pump used with a membrane oxygenator provides a closed automatic bypass system.*

Pulsatile Blood Flow During Bypass Surgery:

Cardiac bypass during heart surgery has given an opportunity to observe the benefit of pulsatile blood flow, by comparing it with techniques using non-pulsatile flow. Pulsatile flow ensures diffuse normal distribution of blood flow to all the organs and tissue of the body, while non-pulsatile flow results in reduced flow to certain vascular beds and excessive flow to others. With non-pulsatile flow, the brain and kidneys receive reduced blood supply. Also, incomplete metabolism from islands of under-perfused tissue results in acidosis, not found when pulsatile flow is provided. When hypothermia is used, the need of pulsatile flow to insure diffuse distribution of circulation is increased. While the adverse effects of non-pulsatile flow can be offset by the use of vaso-relaxers, blood dilution, and other methods, the value of pulsatile flow in normal function is illustrated by comparing findings from the two types of pumping systems.

Ventricular Compliance — an Inlet Impedance Factor:

Studying the model has an advantage over studying in vivo preparations in that the effect of altering one variable at a time can be directly observed. Furthermore, the model eliminates the possibility that some simultaneous, hidden, undetected change might have occurred that would invalidate cause and effect conclusions (Fig. 37). The model has proven to be a very effective study and teaching tool.

The Model:

Figure 38 shows the layout of the model pictured in Figure 39.

The Pumps:

The model pumps share the three unique characteristics of the heart: They are non-sucking, and therefore fill passively at their inlets, they have a pulsatile outflow, and they have atria which allow uninterrupted inflow to the intermittent, pulsatile outflow ventricles.

Figure 40 shows one of the two assembled pumps. Figure 41 shows the unassembled parts of a pump.

The atrial and ventricular pump chambers consist of a rigid side (Fig. 41 at A) and an opposing pliable silicone rubber side (Fig. 41 at B). The ventricular half of the rigid side has both an inlet and an outlet port containing inlet and outlet valves, respectively. The atrial half of the rigid side has a single port with no valve. There is a "Y" connection between the atrial port, inlet port to the ventricle and the venous line. This "Y" inlet allows venous fluid to run both into the atrium, when the inlet ventricular valve is closed during systole, and into the ventricular chamber during both ventricular diastole and atrial systole. This arrangement reproduces the "atrial effect" by allowing continuous uninterrupted flow from the veins to the intermittent outflow ventricles, thus preventing the need of overcoming inertia by starting stopped flow after each pump beat.

The impeller part of the pump also consists of a rigid side (Fig. 41 at E) and an opposing pliable rubber side (Fig. 41 at D). Each half of the rigid part has a port through which air can be forced into or sucked out of the impeller chambers. When the impeller chambers are inflated, their rubber side exerts pressure on the silicone rubber side of the ventricle or atrial chambers, thus emptying the ventricle or atrium.

A spacer (Fig. 41 at C), placed between the opposing impeller and pumping chambers, has many side holes which allow air to pass freely to and fro between the ambient air and the space between the impeller and pump ventricles. This free communication prevents any negative pressure, caused by suction used to deflate the impeller spaces, from being transmitted to the pump ventricles. Thus, the ventricles fill passively from venous pressure.

The power supply to the pumps is from compressed air and vacuum sources, delivered alternately to the impeller spaces timed in the desired sequence. An electronic regulator triggers solenoid valves to control pulse rate and the ratio of systole to diastole. The pulse rate is variable from 1 to 160 strokes per minute at an impeller pressure of one to thirty p.s.i. Reducer valves regulate peak pressure and onset to peak time. Impeller deflating suction is strong enough that passive ventricular filling is not impeded.

Simulated Vascular Network:

The major vessels are silicone tubing, while the capillary beds are made of more compliant Penrose drain material. Variable resistance points are at various sites in both arterial and venous lines. Pressure transducers connect to a direct writing recorder and pressure tubes connect to an electromagnetic flow meter. A transfusion reservoir provides opportunity to change the fluid volume within the system. Different weights placed on a plate resting on the "capillary beds" allow changes in the system compliance to be made. The system was filled with water. Hundreds of pieces of #0 silk suture material, 2 millimeters in length (found to be isobaric with water), were suspended in the water to allow visualization of the circulation.

With the pumps turned off, the compliance remaining unchanged, and starting with the system full of water, at a volume of 2000 cc. at a pressure of zero, pressure was recorded as fluid was added by way of the transfusion reservoir. Figure 43 shows the non-linear relationship between the mean system pressure and the volume.

Starting with a mean system pressure of zero, and with arbitrary settings of the resistance points, the pumps were started. Appropriate amounts of fluid were added to increase the mean system pressure by increments of 2 cm. water pressure (Fig. 44)

Findings:

1. Figure 44 shows a direct correlation between pump output and mean system pressure over the range of 0 to 20 cm. of water pressure. The greater the mean system pressure, the greater the pumps' output.

    

2. At no time, over the range of pressure studied, did the ventricles fill to capacity at diastole. There was always a "pump capacity excess."

    

3. At no time was there any interruption of venous flow by the intermittent closing of the inlet valves of the ventricles at systole, as the atria had been emptied during ventricular diastole.

The pump rate in the model was varied from 0 to 170 beats per minute with other variables remaining constant (Fig. 45). The mean system pressure was kept at 12 cm. water pressure, with the impeller pressure at 14 p.s.i. and the resistance kept constant.

Findings:

1. At pump rates between 0 and 20/min.:
   • The pump output correlated in a linear way with the pump rate (Fig. 45, from A to halfway to B).
   • The ventricles filled to capacity at each ventricular diastole.
   • The atria filled to capacity during each ventricular systole.

    

2. At pump rates between 20 and 40/min.:
   • The pump rate correlated in a non-linear way with the pump output (Fig. 45, from halfway in between A and B, to B).
   • The ventricles filled slightly less than capacity at ventricular diastole.
   • The atria were filled to capacity at the end of ventricular systole.
   • Venous flow was slowed but not completely interrupted before the end of ventricular systole.

    

3. Increasing the pump rate from 50 to 110:
   • Caused no increase in pump output (Fig. 45, B to C).
   • Was associated with progressive decrease in ventricular end-diastolic volume.
   • Caused progressive decrease in atrial diastolic volume.
   • Was unaccompanied by any interruption in venous flow.
   • The pump rate and stroke volume were reciprocals of one another.

    

4. Increasing the pump rate from 110 to 170 beats/min. (Fig. 45, C to D) was accompanied by a progressive increase in residual air pressure in the ventricular impeller after ventricular systole which:
   • Progressively decreased pump output.
   • Impeded ventricular filling at ventricular diastole.
   • Caused venous flow interruption at ventricular systole.

The ventricular impeller pressure was varied from 0 to 20 p.s.i., in increments of 2 p.s.i., while keeping other variables constant (Fig. 46). The pump rate was 80, the mean system pressure was 12 cm. water pressure and all resistances remained unchanged.

Findings:

1. At impeller pressures between 0 and 10 p.s.i. there was:
   • Linear correlation between impeller pressure and pump output (Fig. 46, from A to B).
   • The ventricles were never completely emptied at ventricular systole.
   • The ventricles were completely filled at ventricular diastole.
   • Venous flow to the pumps was interrupted at each ventricular systole.

    

2. At impeller pressures between 10 and 18 p.s.i. there was:
   • No increase in the output as the pressure increased (Fig. 46, from B to C).
   • Maximal ventricular emptying at systole.
   • Sub-maximal filling of the ventricles at diastole
   • No venous flow interruption at any time.

    

3. As the impeller pressure was progressively increased above 18 p.s.i. (Fig. 46, from C to D) there was progressively incomplete evacuation of compressed air in the ventricle during diastole which:
   • Progressively decreased pump output.
   • Impeded ventricular filling at diastole.
   • Caused venous flow interruption at ventricular systole.

Resistance at a variety of sites was changed in the model's vascular network while maintaining other factors constant, with the mean system pressure at 12 cm. water pressure, pump rate at 80 beats per minute, and an impeller pressure of 16 p.s.i. There was a marked difference in response to a given resistance depending on where the resistance was placed in the model. Responses fell into two categories which became more obvious the closer the resistance was to either the inlet or outlet of a pump. Therefore, resistance was studied in two situations: near a pump inlet, where there was a large compliant bed upstream; and near a pump outlet, with no compliant bed upstream.

Relationship of Resistance Near a Pump Outlet to Pump Output: (Fig. 47)

A variable resistance clamp was placed 20 cm. from the left heart homologue pump outlet (Fig. 38, #9), across which a pressure gradient could be monitored by transducers (Fig. 38, #8).

Findings:

1. As the resistance is progressively increased (Fig. 38, #9), the superimposed arterial pressure tracings, which demonstrated no gradient initially (Fig. 47 at A), separated as resistance was added (Fig. 47, from A to C).
2. With increase in outflow resistance up to 100 mm. Hg. gradient, there was no drop in pump output (Fig. 47 at C).
3. Further increase in gradient from 100 to 250 mm. Hg did cause a corresponding drop in flow (Fig. 47, from C to D).
4. In the range where there was no drop in flow (Fig. 47, A to C), the left ventricle did not fill maximally, but did empty maximally, and there was no interruption of venous flow at any time during the pumping cycle.
5. However, when the gradient became great enough that the flow slowed (Fig. 47, C to D), the left ventricle progressively emptied less completely, and the residual volume resulted in complete left ventricular filling at diastole and interruption of venous inflow at each systole.
6. Whenever the left ventricle was being filled completely at diastole, increasing the mean system pressure caused no concomitant increase in pump output.
7. As slowing of flow occurred (Fig. 47, C to D), there was gradual increase in a greater than previous volume in the pulmonary circuit at the expense of the systemic circuit. The greater the slowing, the more the volume equilibrium shifted to the simulated pulmonary circuit.

The Effect of Resistance Near a Pump Inlet on Pump Output: (Fig. 48)

An adjustable resistance clamp was placed 20 cm. from the left pump's inlet (Fig. 38, #7), across which a pressure gradient could be monitored using the transducers (Fig. 38, #6). The clamp was progressively closed. The resulting inflow resistance magnitude is represented by the pressure gradient between the transducers. Figure 48 shows the two superimposed pressure lines (Fig. 48, A to B) in the absence of resistance. When resistance is progressively increased, starting at B, a pressure gradient becomes evident (Fig. 48, B to C).

Findings:

1. There was a direct inverse correlation between pump output and inflow resistance to the pump (Fig. 48, B to C) over the whole range studied. This is in marked contrast to outflow resistance of the previous experiment (Fig. 47, B to C), where decrease in flow did not occur until there was pump failure.
2. As inlet resistance increased there was progressively less filling of the ventricles at diastole. Maximal emptying did occur at each ventricular systole, and venous flow remained uninterrupted at all times.
3. Progressive increase in left pump inflow resistance shifted the fluid equilibrium between the pulmonary and systemic circuits toward the pulmonary bed.
4. The circuit with the greatest inflow resistance was the major determinant of the output of both.

In the model, to eliminate the atrial effect, the air pressure in the atria can be left continuously obliterating the atrial chambers, while not impinging on the connection of the venous tubes to the ventricles. By intermittently producing circulation with and without the atrial effect, its contribution to pump output was determined. At pump rates between 60 and 90 there was four times the pump output (2400 cc./min.) with the atria functioning as with no atrial effect (600 cc./min.). The output without atria progressively decreased as the rate was increased. At 130 beats per minute, flow stopped completely. It is demonstrated that atria markedly increase flow when used with pulsatile, passively filling pumps. They prevent inertia, which would otherwise occur if the inlet valves were allowed to stop venous flow at each beat.

A branching hydraulic system was made with twenty, 3/16 inch diameter, transparent tubes, six to twelve inches in length, linked together with "Y" connectors in the configuration of a simulated vascular bed.

Findings:

1. With pulsatile flow there was flow in all parts of the system. It fairly danced with the pulsation. There were areas where flow was in one direction during most of the cycle and reversed in the other direction for the rest of it. With pulsatile flow, the distribution remained stable for a long period of time.

    

2. With continuous non-pulsatile flow, the distribution started out in a diffuse manner. Very shortly, some channels had higher flow than others. Yet there was some flow in all of them. However, within a short time the 2 mm. silk particles began clustering at "Y" junctions, eventually totally blocking some and partially blocking others. Before long, almost all of the flow was in just a few channels and very little was in the others. When pulsatile flow then replaced the steady flow, the aggregates of silk were broken up, the whole bed began dancing with the pulse, and flow was again established in a diffuse manner, with flow fairly equal in all channels.

Conclusion:

There is obviously a difference in the mechanism of aggregation of silk particles in this system and the mechanism that caused poor distribution with non-pulsatile flow in the animal in the previous chapter. However, this experiment does show that the simplest way to guarantee diffuse, equal distribution of fluid in a hydraulic system is to make flow pulsatile.

We would have sealed the woman's fate if we had corrected only one of the valves. She would have remained in low output failure if we had corrected the aortic valve alone; and she would have gone into acute left ventricular failure if we had relieved the mitral stenosis alone. Therefore, we replaced both valves and the patient made an excellent recovery.

1. The cardiovascular system is a circle containing a mean cardiovascular pressure dependent upon volume and compliance.

    

2. The heart is a passively filling pump.

    

3. Atria increase cardiac output by causing continuous venous flow to the intermittently contracting ventricles, not by pumping up the ventricles.

    

4. Circulation rate is normally determined by the extra-cardiac factors: mean cardiovascular pressure and inlet impedance.

    

5. It is only during heart failure that the heart is the regulator of cardiac output.

Equipment:

A narrow pneumatic blood pressure cuff (approximately one inch in width) is used so that during inflation it will not displace any blood volume distally into the arm.

A one liter air pressure reservoir, pressurized to 300 mm. Hg, is connected to the pressure cuff with a valve interposed to allow instantaneous inflation of the cuff. The reservoir is pressurized by a regular blood pressure inflating bulb and valve, interposed with a "Y" connecter between the reservoir and the valve.

Two pressure transducers and recorders are needed, to allow pressure recording without any loss of fluid from the vascular bed, which would occur with a simple manometer system.

Technique:

 

1. The patient should be lying perfectly horizontal.

    

2. Both an artery and a vein are cannulated at the antecubital area of the arm and connected to the pressure transducers.

    

3. The blood pressure cuff is applied above the biceps bulge of the arm. It must be applied loose enough that it does not cause any venous obstruction, as evidenced by the observation that it produces no elevation of venous pressure above that seen before its application.

    

4. The cuff must be applied tight enough that its inflation, to 300 mm. Hg pressure, completely interrupts the arterial flow to the arm. Complete interruption can be assumed if arterial and venous pressure approach equalization after 30 seconds, and do not continue to rise thereafter.

    

5. The arm should be abducted sufficiently from the side of the body, so as to obtain the lowest venous pressure reading, thereby being certain that any abduction is not mechanically interfering with venous flow in the axilla.

    

6. The arm should be horizontal, with the anterior surface of the antecubital skin at mid- chest position.

    

7. The thickness of the chest should be accurately measured. The pressures are small, so careful standardization is necessary in order to get meaningful data.

    

8. The patient is instructed to leave the arm perfectly relaxed, and not to contract any muscle.

    

9. The valve or switch in the pressure line is turned, thereby inflating the cuff, which simultaneously interrupts venous and arterial flow to and from the arm.

Pressure readings are made 30 seconds after the occlusion. As the equalization process progresses, the pressure gradient becomes so small between the arteries and veins that the pressures may never completely equalize. However, the average of the arterial and venous pressures recorded at 30 seconds can be interpreted as the mean-cardiovascular pressure.