Mock Circulation Loop For Biomedical Device Testing Physical Education Essay

As biomedical devices are become ever more high-tech, research in them is spreading wider across various engineering disciplines. In order to allow the advancement of technologies in this area to continue, it is necessary that companies and research organisations have access to their own low cost, pulsatile Mock Circulation Loop (MCL) for developmental testing. A Mock Circulation Loop simulates pressure-flow response of the human circulatory system, for different physiological states. This study proposes a low cost MCL, designed from “off the shelf” commodity components.


A Mock Circulation Loop simulates pressure-flow response of the human circulatory system, while also being able to replicate this system for different physiological states. Previous authors have suggested that an effective MCL should have a least three benchmark states; healthy person in sleep, rest and mild physical activity. MCLs are used to investigate the effectiveness of biomedical devices across a wide range of applications, but are predominantly used to test artificial heart valves, vascular prostheses and stents. They have also been found to be a very useful educational tool as they have the advantage over various other educational medium, simply by providing the student with a visual platform. To be able to observe the phenomena in operation makes it easier to interpret many different physiological aspects [1].

MCLs that closely mimic human parameters are an extremely important tool nowadays in bringing devices to the biomedical market from a cost, but also an ethical perspective. Unnecessary in vivo animal trials can be alleviated by carrying out comprehensive testing of devices in an in vitro environment under various physiological conditions. This means only the best and most promising designs will go forward for animal trials, thus reducing needless in vivo trial and error design which can be extremely costly and unethical [2].

Recent design and research into MCLs has come predominantly from those interested in producing accurate in vitro testing systems for aiding the development of ventricular assist devices [3-6]. However there is little or no reason to believe that the systems designed for these purposes would not be equally applied to beneficial effect in the development of new, innovative devices such as BioSensors or BioMEMS (biomedical micro-electro-mechanical devices). The origins of these components are usually not from the well established areas of biomedical research and design, but from historically new up and coming areas. Without the know how or experience in place, it is inevitable that the requirement for easily attainable, low cost “out of body” testing equipment will become more in demand as this new area of biomedical device design continues to grow with demand. It is hoped that this thesis can go some way in bridging this knowledge gap.

This thesis presents a research project for the design and construction of a mock circulation loop in a low cost manner, employing the best aspects of currently available MCLs and where possible, using off the shelf components. The mock loop will consist of two major elements: (1) passively filling left pulsatile artifical heart ventricle; and (2) air/water vessels to simulate the venous and arterial compliances. These elements will be coupled together using appropriate tubing.

The performance level of the completed MCL will be made accessible to engineers in the form of: (1) End systolic pressure-volume relationship plots, (2) Ventricular pressure-time relationship plots, (3) Systemic & Pulmonic Pressure Distribution versus time plots. These can then be compared against equivalent data available for the human physiology and also other current MCLs. The key contributions expected from this work are: (1) that W.I.T. will have a fully functioning MCL test rig for “in-house” development of microelectronic biomedical device components, (2) the MCL will present further opportunities within W.I.T. for projects in the area of cardiovascular research (3) the complete design and construction data will be available to any institution or company wishing to build their own low cost MCL testing rig.

Existing MCL Designs

Mock circulation loops for the in vitro replication of the human cardiovascular system have been developed since the 1970s [7-8]. The more modern versions vary quite a lot in how they attempt to replicate these human parameters, from the pumping systems they use, to how they achieve compliance. Some of the methods used for imitating the contraction of the heart are shown in Figures 1-3. These are; 1) diaphragm pumps controlled by pneumatic compression and vacuum [2, 4], 2) motor controlled piston pump [1], and 3) pneumatic supply directly into artificial ventricle [9-10]. In terms of producing a low cost model then it would appear the third method would be the most beneficial, while also providing an accurate model based on its demonstration to closely replicate the key elements of any MCL.

The disadvantage to using such a system would be the ability to closely replicate the time-dependent flow waveform supplied by the ventricle, both in physiological and pathological conditions. It has been reported that the hydraulic volumetric pumping systems inherently provide a finer control of the flow waveform [2]. However, improved control of ventricle contractility in order to improve waveform replication can be achieved through the use of an electro-pneumatic regulator [10]. The downside to this solution is that these regulators are very costly.

Cardiovascular Physiology
The cardiac cycle, the systemic compliance, and the systemic vascular resistance are three very important parameters to understand and replicate, if an accurate mock circulation loop is to be designed and built. These three areas are described herein.
Cardiac Cycle

The atria and ventricles contract in sequence, resulting in a cycle of pressure and volume changes. The cardiac cycle has four phases, over a combined time of 0.8 to 0.9sec for 70 beats per minute. The four phases and their time duration are outlined briefly.

Ventricular Filling, Duration 0.5s

Ventricular diastole lasts for nearly two-thirds of the cycle at rest, providing adequate time for refilling the chamber. There is an initial phase of rapid filling, lasting about 0.15s, as shown by the cardio-meter volume trace in Figure 4. As the ventricle reaches its natural volume, the rate of filling slows down and further filling requires distension of the ventricle by the pressure of the venous blood; ventricular pressure now begins to rise. In the final third of the filling phase, the atria contract and force some additional blood into the ventricle. The volume of blood in the ventricle at the end of the filling phase is called the end-diastolic volume (EDV) and is typically around 120ml in an adult human. The corresponding end-diastolic pressure (EDP) is a few mmHg. EDP is a little higher on the left side of the heart than on the right, because the left ventricle wall is thicker and therefore needs a higher pressure to distend it.

Isovolumetric Contraction, Duration 0.05s

As atrial systole begins to wane, ventricular systole commences. It lasts 0.35s and is divided into a brief isovolumetric phase and a longer ejection phase. As soon as ventricular pressure rises fractionally above atrial pressure, the atrioventricular valves are forced shut by the reversed pressure gradient. The ventricle is now a closed chamber and the growing wall tension causes a steep rise in the pressure of the trapped blood; indeed the maximum rate of rise of pressure (dP/dt)max, is frequently used as an index of cardiac contractility.

Ejection, Duration 0.3s

When ventricular pressure exceeds arterial pressure, the outflow valves are forced open and ejection begins. Three quarters of the stroke volume is ejected in the first half of the ejection phase and at first blood is ejected faster than it can escape out of the arterial tree. As a result, much of it has to be accommodated by distension of the large elastic arteries, and this drives arterial pressure up to its maximum or ‘systolic’ level. As systole weakens and the rate of ejection slows down, the rate of which blood flows away through the arterial system begins to exceed the ejection rate, so pressure begins to fall. As the ventricle begins to relax, ventricular pressure falls below arterial pressure by 2-3mmHg (see stippled zone in Figure 4) but the outward momentum of the blood prevents immediate valve closure. The reversed pressure gradient however, progressively decelerates the outflow, as shown in the bottom trace of Figure 4, until finally a brief backflow closes the outflow valve. Backflow is less than 5% of stroke volume normally, but is greatly increased if the aortic valve is leaky. It must be emphasised that the ventricle does not empty completely but only by about two-thirds. The average ejection fraction in man is 0.67, corresponding to a stroke volume of 70-80ml in adults. The residual end-systolic volume of about 50ml acts as a reserve which can be utilised to increase stroke volume in exercise.

Isovolumetric Relaxation, Duration 0.08s

With closure of the aortic and pulmonary valves, each ventricle once again becomes a closed chamber. Ventricle pressure falls very rapidly owing to mechanical recoil of collagen fibres within the myocardium, which were tensed and deformed by the contracting myocytes. When ventricular pressure has fallen just below atrial pressure, the atrioventricular vales open. Blood then floods into the atria, which has been refilling during ventricular systole [11].


Compliance is related to the ability of a vessel to distend when encountering a change in blood volume [12]. It is defined as a change in volume for a given change in pressure and can be described by Equation 1 as follows:


The distension of the elastic arteries raises the blood pressure and the amount by which pressure rises depends partly on the distensibilty of the arterial system [11]. The compliance of the veins is approximately 24 times greater than that of arteries. This gives the veins the ability to hold large amounts of blood in comparison to arteries. Ventricular compliance influences the ventricle’s pressure volume curve. If the compliance of the ventricle is decreased, this increases the end diastolic pressure for any given end diastolic volume [10]. The EDPVR provides a boundary on which the PV loop falls at the end of the cardiac cycle [13].

Systemic Vascular Resistance

The systemic vascular resistance (SVR) or total peripheral resistance (TPR) is the ratio between the mean pressure drop across the arterial system [which is equal to the mean aortic pressure (MAP) minus the central venous pressure (CVP)] and mean flow into the arterial system [which is equal to the cardiac output (CO)]. Unlike aortic pressure by itself, this measure is independent of the functioning of the ventricle. Therefore, it is an index which describes arterial properties. According to its mathematical definition, it can only be used to relate mean flows and pressures through the arterial system.

Design of MCL System

The most recent circulation loop designs are complete mock circulation systems i.e. they include both the systemic and pulmonary circuits into their design. This is not desirable in a lost cost mock circulation model as it doubles the cost, without necessarily adding equivalent value, and in terms of testing microelectronics biomedical devices, a simple pulsatile single loop is sufficient. The mock circulation loop in this study will be a single loop, replicating the parameters of the systemic system.

Mock Circulation

Based predominantly on one half of the Timms [9] mock circulation rig. The rig consists of three main systems (Figure 5). These are 1) compressed air supply, 2) mock circulation loop, and 3) data acquisition system.

Air Supply

The air supply system is responsible for heart contraction, contractility, heart rate, and systolic time. The system is made up of the following components in series; air compressor (24litre Hobby), precision pressure regulator (SMC NIR201), electro pneumatic regulator (SMC ITV2030-31F2BL3-Q), and 3/2 solenoid valve (SMC EVT307-5DZ-02F-Q).

Air is supplied via the air compressor at 7bar pressure to the precision regulator, where the pressure is reduced to a desired lower pressure, sufficient to pump a required amount of fluid contained in the mock ventricle tube. Contraction of the mock ventricle is initiated when the solenoid is in the open position, and air is able to flow through it and into the top of the ventricle chamber. The contractility of contraction can be varied by the electro pneumatic regulator, which increases or decreases the amount of air supplied. When the solenoid changes to the closed position, air is exhausted out of the ventricle chamber, mimicking diastole. The rate at which this happens can be varied by changing the exit port size. The period of systole and diastole for a given cycle can be varied by changing the time the solenoid is in the open and closed positions.

Mock Circulation Loop

The hydraulic circulation loop itself consists of atrium, ventricle, and systemic and coronary vasculature components (Figure 6). The atrium is open to atmosphere and is passively filling. It is constructed from 40mm diameter pvc pipe. The ventricle is downstream of this and is constructed similarly, except that it is capped. The capped ventricle is tapped with the pneumatic line from the air supply system. For the heart valves, a 40mm brass check valve (Cimberio C80-40) is used as the mitral valve and a 32mm brass check valve (Cimberio C80-32) as the aortic valve. The cross-sectional opening areas of these valves are similar to that of the human heart valves. The swing gate flow resistance on these brass valves is satisfactorily low, yet also prevents fluid backflow during cardiac pumping.

Vasculature parameters of compliance and resistance were replicated through the use of windkessel chambers and a pinch valve. Compliance is varied by altering the vertical position of the test plug. In doing so, the amount of air contained above the fluid is altered, and in turn the compliance level is varied.

Resistance is increased by tightening of the pinch valve, which increases pipe occlusion. Inherent resistance values are calculated by taking the required pressure drop across a component and dividing it by the maximum flow rate through the component (Eq. 1). Max flow rate is calculated by first using Bernoulli’s equation (Eq. 2) to determine maximum fluid velocity (v2), assuming initial and final water heights are equal (z1=z2) and initial velocity (v1) is zero. Multiplying final velocity (v2) by each component’s maximum pipe cross sectional area (A) reveals maximum flow rate (Eq. 3) [12].

Data Acquisition System

Flow rate in the systemic path will be measured by an electromagnetic flow meter (Omega FMG 3002-PP) as represented in Figure 6. Pressures at three locations in the circulatory loop, as identified by the number 1 in Figure 6, will be measured by pressure sensors (WIKA Pressure Transmitter A-10.). All measured signals are taken to I/O connector block (NI SCC-68 National Instruments Inc.) and then sent to a data acquisition board (NI-6040e National Instruments Inc.) and processed in Labview on desktop PC.

Test Protocol

There will initially be two different activity levels simulated with the mock circulation system. These will be a healthy person at rest, and in exercise (equivalent to ascending stairs). It is envisaged that further conditions of activity and pathology will be simulated, should the primary two activity levels run successfully. The parameters for the aforementioned states are set out by Liu [14] in Table 1.

Experimental Procedure

National Instrument’s LabVIEW will be used to control heart rate and ventricle pressure, while also being able to record in realtime, the feedback pressures and flowrates from the system.

Compliance values as described earlier, are achieved by filling the compliance chambers with a set amount of water and adjusting the test plugs to the required height. Once in the correct position, the bleed valve can then be closed, thereby trapping a volume of air above the water level. The bleed valve and pressure sensor are fitted into two small holes drilled into the top of the plug. With the compliance chambers now set, a small volume of water is added to the ‘open to atmosphere’ atrial chamber. By adding water, the compliance pressures can be finely tuned to suit the desired physiological state. The air compressor is then charged until its 24litre reservoir has been filled. The output regulator for the compressor is tuned, followed by the ventricle input precision regulator. Heart beat, which is controlled by the 3/2 solenoid, is instigated in accordance with Table 1 (40% Systolic) using the LabVIEW controller. A manual tuning clamp is used to adjust the vascular resistance.

Results Expected

The results of simulated tests for the various physiological conditions will be presented in order to compare with natural cardiovascular hemodynamics. There will be an individual graphical plot for each condition. The plot should be contain at least four continuous cardiac cycles, in order that repeatability of each physiological state can be verified. The graphical plots will have time as the X-axis. From this, ventricle pressure, atrial pressure and flow-rate over time shall be presented. Results obtained for all scenarios will be tabulated into one table, similar to Table 2 [14], for easy comparison.

The test results for simulated healthy at rest conditions for two very similar mock circulation loops [9, 14] are displayed in Figures 7 and 8. For a normal, healthy individual, the heart rate and systolic ratio set by these authors are 60bpm and 40% respectively. Values obtained for peak left ventricular pressure are both in the region of 120mmHg and ventricle end diastolic pressure is the 5-10mmHg region. Cardiac cycle times are very close at 0.8-0.9 seconds.

Intense pressure fluctuations in the atrial, and ventricle chambers has been observed in the results of both authors. This has been attributed to the rigid nature of the valves, causing a water-hammer effect during brisk closing.

No results are provided by either author in relation to whether their systems display the ‘Frank Starling Effect.’ However, Timms [9] states that this effect has been visually observed through the clear PVC ventricle chamber as changes in fluid level prior to systole.


The results for the above two similar mock circulation loops demonstrate that they can successfully replicate the conditions of the human physiological state of a healthy person at rest. They also demonstrated success for other conditions, the results of which have not been presented here. Of particular interest was the success of the compressed air system used by Timms [9] to produce the contracting ventricle. This method is directly similar to that proposed for the MCL in this study.

Liu [14] has demonstrated that LabVIEW, as will be employed in this project, can be used to promising effect in the control and recording of various elements contained in the mock circulation loop.

Proposed methods for the reduction of the intense fluctuations in pressure due to water-hammer include suppression by imposing a digital filter on the recorded pressure data or the introduction of an accumulator near the valves to physically reduce these transients [9],

It is a desire within this study to record ventricle volume data, which has not been attempted in the MCLs of Liu or Timms [9, 14]. By recording the ventricle volume, it is possible to demonstrate the ‘Frank Starling Effect’ during the operation of the MCL. It is envisioned that this can be achieved through the use of a capacitance coil placed within the ventricle chamber.


The system proposed in this study very closely matches those of Timms and Liu [9, 14], except that is expected to be produced for a lower cost, due to the commodity components it will use and also that it will only simulate the systemic loop, which is deemed sufficient for development of microelectronic biomedical device components. Due to the successful results obtained by the aforementioned MCLs, and the similarity of the proposed design to these MCLs, suggests that equally positive replication of the human cardiovascular physiology can also be achieved.


The author would like to acknowledge the financial support provided by the Department of Mechanical Engineering, Waterford Institute of Technology, together with the invaluable guidance and technical support afforded by Dr. Austin Coffey and Mr. Philip Walsh.

Key Words: Cardiovascular system, pulsatile mock circulation loop, MCL, PV loops, biomedical device testing