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Multi-Disciplinary Engineering Design Conference
Kate Gleason College of Engineering
Rochester Institute of Technology
Rochester, New York 14623
Project Number: P09026
HEMODYNAMIC SIMULATOR II
Alex Baxter
|
Data Acquisition Team
Liliane Pereira
Joseph Featherall |
Lead Engineer
Mark Frisicano |
Pump Design Team
Clarissa Gore
Control System Team
|
ABSTRACT
The Hemodynamic Simulator is a modular system
that in intended to reproduce the hemodynamic flows
and pressures associated with a circulatory system.
The simulator will enable performance of experiments
and associated with the fundamental properties
associated with typical cardiovascular circulatory
system. The system will incorporate control elements
that will allow generation and measurement of
arbitrary dynamic flow rates, volumes and pressures as
well as the evaluation of the impact of variations of the
characteristics of system components such as vascular
compliance, fluid composition and overall topology.
INTRODUCTION
Biomedical engineering involves the application
of engineering principles and techniques to help
improve the quality of lives and overall patient
healthcare and consists of research and development in
areas such as bioinformatics, medical imaging, image
processing, and biomechanics.
In this project, a Hemodynamic Simulator was
developed. The system allows for simulation of fluid
flow through a mock circulatory loop. The simulator
will be useful in research related to cardiovascular
|
Data Acquisition Team
Jonathan Peyton |
Pump Design Team
Gaurav Zirath
Team Leader
|
diseases, as it will allow study of blood flow dynamics
as a function of varying anatomical parameters such
as the heart rate and systolic ejection period.
NOMENCLATURE
HR – heart rate, measured in beats per minute is the
number of times a human heart contracts, to provide
continuous supply of oxygenated blood throughout the
body.
SEP – systolic ejection period, the amount of time the
ventricles spend in systole per minute
CO – cardiac output, is the amount of blood pumped
by the heart, a ventricle in particular, in a minute.
Systole: contraction of heart chambers, pumping blood
out of the chambers
Diastole: rhythmical relaxation and dilation of heart
chambers to allow the filling of blood
Stroke Volume: amount of blood pumped out of
the ventricular chambers, every beat
VI: LabVIEW programs and subroutines are called
virtual instruments (VI).
DAQ: data acquisition
THK: leading manufacturer of precision linear motion
guides, ball spines, ball screws etc.
Yaskawa: world’s largest manufacturer of ac drives
and motion control products.
EDGE: The Engineering Design Guide and
Environment -- is an open source integrated design
environment to foster collaboration within design
Copyright © 2008 Rochester Institute of Technology
Proceedings of the Multi-Disciplinary Engineering Design Conference
Page 2
project teams, and across design teams working on
families of closely related projects.
OVERVIEW
In order to produce a functional and aesthetically
pleasing Hemodynamic Simulator that meets the
project’s needs and requirements, the team was
divided into three distinct groups and each group was
assigned tasks specific to their role.
The Pump Design group was responsible for
designing and manufacturing of all components
associated with the pumping mechanism for the
simulator. Some of the tasks for the group were to
identify an actuator and servo motor drive, design and
manufacture a separate compliance and atrial reservoir
chambers and redesign of a pneumatic cylinder. The
Data Acquisition group was responsible for
identifying suitable sensors and transducers to acquire
pressures and flow data including any additional signal
conditioning circuitry. The Control System group
developed a software based user interface and control
system based on LabVIEW 8.6 (National Instruments,
Austin, Texas) running under the Windows XP
operating system (Microsoft, Redmond, Washington)
on a personal computer.
BACKGROUND
This project, P09026 was a continuation of project
P08026 carried out during 2007-08 academic year.
During the course of project P08026, a working model
of a hemodynamic simulator was developed. The
simulator successfully proved the concept of moving
fluid in a controlled fashion by transferring generated
pneumatic pressure onto to a fluid column through a
diaphragm.
Figure 1: Full Assembly of Hemodynamic Simulator
from project P08026
http://edge.rit.edu/content/P08026/public/Home
A linear actuator was used to generate pneumatic
pressure and was applied to the diaphragm, attached to
the end on buffering chamber, via a metal rod. As seen
in Figure 1, the buffering chamber is directly
connected to the heart chamber to pressurize it, which
created contraction and relaxation in left ventricular,
characterized by a plastic bag. The system allowed for
various motion profiles stored on the servo controller
used to operate the actuator.
Although, a modular hemodynamic simulator
model was presented at the end of project P08026’s
term, the system was far from being portable. Also, the
amount of control to servo controller was limited to
the motion profiles that were programmed onto the
programmable memory. Consequently, project P09026
was created with following objectives:



Enhancement of pump design
Enhancement of instrumentation and
data acquisition software
Research and development of computer
control of all system parameters
In this project, the redesigning phase of pressure
generation system (mechanical pump), instrumentation
and the control system was carried out to ensure that
the final product is portable, robust and easily
operatable. Consequent sections of this paper discuss
in detail the design approach of various segments of
the project. All documents related to design and
manufacturing for project P08026 and P09026 are
located on following webpages, respectively.
http://edge.rit.edu/content/P08026/public/Home
Hemodynamic Simulator II
Project P09026
Fall ’08 – Winter ‘08
Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
https://edge.rit.edu/content/P09026/public/Home
REDESIGNING PHASE
As mentioned earlier, the main focus of the
redesigning phase was to manufacture a system which
will rather be easily transportable from one classroom
to another and is user friendly. The redesigning of the
system began with slight modifications to the
pneumatic pressure transfer mechanism.
Earlier
project utilized a rubber diaphragm to transfer
pneumatic pressure to hydraulic pressure, directly
through a metal rod. The rod was directly connected to
the diaphragm with no safe interface; and therefore
constant beating of the rubber with a metal rod caused
a potential risk of the rubber rupturing and then a
water leak, all during operation. Hence, a safer
mechanism was designed and later successfully
implemented. As shown in Figure 3, the new design
replaced the metal rod with a air cylinder, which was
directly connected to the buffering chambers’ solid top
with a tube. This setup ensured that there was no
failure in the pump mechanism, even during extreme
pressures.
In order to make the system smaller in size, the
compliance chamber and the atrial reservoir chambers
were also redesigned. Many fluid dynamic
Page 3
through and out of the ventricle. Previous system
didn’t allow for any ultrasound measurements. The
new design, as shown in figure 2, the heart was housed
on a stainless steel stand, which allowed for mounting
of ultrasound probes on adjustable clamps.
Furthermore, a new actuator and servo with relatively
smaller dimensions and comparatively better
performance were identified as replacement to the
inherited equipment.
Figure 3: Sketch of the redesigned overall system
From the control and data acquisition segment of
the project, the industrial based Omega pressure
transducers. A NI-USB6008 DAQ is utilized to
measure and record pressure and flow data to
LabVIEW. To enhance the control of the system, a
LabVIEW graphical user interface (GUI) was
developed to display pressures and flow meter
readings and provide user with the control to the
actuator via a PCI board. Subsequent sections discuss
in detail, the design approach for all segments of the
project.
Figure 2: Heart/Ventricle Chamber
characteristics were studied before finalizing the
design of the chambers. The heart chamber used in the
project was originally constructed with three windows
on the base of the structure (as seen in figure 2) to
allow for ultrasound measurements of the ventricular
contractions and more importantly the fluid dynamics
Copyright © 2008 Rochester Institute of Technology
Proceedings of the Multi-Disciplinary Engineering Design Conference
Page 4
of the aortic chamber resulting with a cylinder with a
six inch OD by 5.75 inch ID was chosen for the atrial
chamber. The inlet and outlet also incorporate one
inch NPT threads and were placed five inches from the
bottom of the chamber to allow for the incoming flow
to disperse energy into the body of water evenly as the
water level in the chamber was aimed at being ten to
twelve inches above the base of the chamber.
Figure 4: Overall System Architecture
MECHANICAL CHAMBER’S DESGIN
Previously the cardiovascular loop and pressure
generation loop were viewed to be overly complicated
and unable to meet the set performance specifications
for the system. One main design goal was to greatly
simplify the system and eliminate any un-needed
components. It was decided to utilize materials that
would most closely match the buffering and
ventricular chamber from the previous system. Cast
acrylic was chosen for its mechanical and chemical
properties, availability and ease of fabrication.
The new components that were designed are
an aortic compliance chamber and an atrial reservoir.
The aortic compliance chamber was sized to
incorporate a hydraulic column as seen in the
buffering chamber in the pressure generation loop.
The closest stock size of acrylic tubing was four inch
outer diameter (OD) and 3.75 inch inside diameter
(ID). To reduce the risk of a fountain like effect at the
inlet of the compliance chamber a taper was placed
into the riser inlet from the ventricular chamber. Also
the chamber was sized to have at least a six inch tall
water column and a four inch (maximum) air buffer.
The outlet was kept to one inch NPT thread, which
was chosen to be the common thread in the overall
system flow loop. The outlet was placed as low as
possible in the chamber to keep pressure changes from
the inlet due to its elevation to a minimum as
preliminary calculations pointed that elevation change
was the main factor in the head loss seen thru the
system.
Critical aspects of the design of the chambers
were ease of manufacture and ease of the assembly pf
the overall system. Common fasteners were chosen
and a common top design was made to ensure proper
sealing while allowing easy emptying and filling of
the chambers. The chamber and atrial reservoir tops
are easily modified- they are essentially disks into
which any fitting style can be incorporated. The top
disk bolts down to a ring that is turned on a lathe and
then welded to the body. This ring holds the nuts for
the bolts to allow for one hand tightening. The top
ring also has a groove cut into it to allow for the
placement of an o-ring to seal between the top ring and
top. The chambers were design overall to be easily
modified to optimize the performance of the system.
Figure 5: Display of the Taper inlet from the Ventricular
Chamber
Determining the diameter for the atrial
reservoir took into account the fact that the reservoir’s
air pocket would be pressurized to a much lower
pressure than in the aortic chamber. As such, it was
designed to have a surface area at least twice the size
Hemodynamic Simulator II
Project P09026
Fall ’08 – Winter ‘08
Page 5
Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
machining. Some of these include: ”crazing” or
miniscule cracking during welding, overheating during
machining, and solvent selection and surface
preparation.
Solvent Welding Acrylic
Figure 6: Cutaway of top ring and top lid fastener
interface.
MANUFACTURING
Acrylic Manufacturing Study
During the design process, acrylic fabrication
was selected as the most effective process for
manufacturing the compliance chambers. Acrylic
material is easily allocated, inexpensive, easily
machined, and can be welded using solvent bonding
techniques. These material properties would allow the
team to use preformed tubing and sheet stock to
construct the final geometry of the compliance
chambers. This process cut down drastically on
material costs, machining time, and waste material.
Surface Prep.
400 grit sandpaper
Surface Cleaning
Denatured Alcohol
Tolerancing
Slip fit .003”-.007” clearance
Solvent Adhesive
Weld-On #1802
Applicator
Hypo. 65 needle
Work time
15 sec.
Supplier
www.rplastics.com
Cure time
24 hrs. (1 hr. till handlable)
Table 2: Developed solvent bonding guidelines
To avoid overheating, surface melting and
creating thermally induced residual stress, all cutting
tools need to be razor sharp. The use of dull or nicked
cutting tools will increase the risk of crazing especially
in thin cross sections and when removing large
amounts of material. If crazing does not occur
immediately, applying solvent to the surface for
welding after cutting can weaken the material enough
to allow crazing to propagate. Table 1 lists the cutting
speeds and feeds used as guidelines for machining
operations.
Machining Acrylic ( w/ HSS)
Sfm
Chip load (in)
Milling
315
0.002
Turning
600
0.01
Boring
150-200
0.0015
Table 1: SFM values for various machining types and
their appropriate chip loads
Figure 7: Acrylic based connector
Although the acrylic is soft and easily machined, the
builders needed to experiment with different cutting
and fabrication methods to optimize the process and
avoid some the negative idiosyncrasies of clear plastic
In addition to these speeds, consideration should
be taken to reduce chip load and therefore the cutting
force near the edge of the part to prevent chipping.
Lubricant is not necessary except during high friction
operations such as tapping.
During the welding process a number of
different fitment tolerances, solvent adhesives, and
surface preparations were tested. Thinner solvents
required extensive surface preparation, and extremely
tight tolerances, while medium bodied adhesives
provide excellent bonding, sealing and product clarity
with minimal surface preparation time. It is also noted
as critical that the solvent bond be held in place for
curing with light pressure so as not to squeeze the
Copyright © 2008 Rochester Institute of Technology
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Proceedings of the Multi-Disciplinary Engineering Design Conference
solvent out of the joint creating a “dry” bond. Table 2
provides the developed solvent welding guidelines.
ACTUATOR & SERVO SELECTION
One major area for mechanical design was the actuator
and pump that were to be used in the system. The first
step in selecting the actuator was to calculate the
needed travel force the actuator provide. Based on
fluid calculations, three feasible scenarios were
determined and travel and force values were derived.
It was assumed the buffer chamber would be entirely
filled with air. This would be the worst case scenario
for the system because the system would have the
most air to compress in this scenario. If the buffering
chamber were completely filled with air, the travel of
the actuator needed would be 200mm in order to
create an assumed pressure of 200mm Hg. The next
calculation assumed the buffering chamber would be
half-filled with water. With this assumption the travel
needed to create 200mm Hg of assumed pressure
would be 125mm. The third scenario assumed the
chamber would be filled entirely with fluid, this means
there would be minimal air to compress. This yielded
a travel of 88mm to gain the 200mm Hg of assumed
pressure. Detailed calculations can be found on
following documents located on EDGE:




https://edge.rit.edu/content/P09026/public/Ac
tuator%20Specifications_1_JPEG
https://edge.rit.edu/content/P09026/public/Ac
tuator%20Specifications_2_JPEG
https://edge.rit.edu/content/P09026/public/Ac
tuator%20Specifications_3_JPEG
https://edge.rit.edu/content/P09026/public/Ac
tuator%20Specifications_4_JPEG
The next property that was necessary to calculate
was how much force was required to compress the air
cylinder. Using a spring gauge we were able to
measure the resistance force from the seals in the air
cylinder. Based on 10 trials, the average force was
found to be 6.5 lbs. Next it was necessary to calculate
the force required to compress the air in the cylinder.
Using the pressure and the piston area along with
the assumed pressures it was shown that 11.5lbs were
needed to compress the air. This means the total force
required is about 18lbs. The last factor was the
velocity of the actuator. Given the heart rate we
wanted to achieve we determined that the required
would be a factor of four larger than the travel that
was calculated.
THK was offering the most
reasonable prices for actuators and had one product
line that would meet the required specifications. It
was determined the system would use the VLA-ST60-12-0250 actuator; which has a travel of 250mm.
Hemodynamic Simulator II
The longer travel was selected so there was extra
travel for expansion of the project as needed.
The
actuator has a maximum speed of 1000mm/s which
yielded a factor of safety of 2, this was the only
actuator found that could meet our specifications in the
velocity category. The actuator was rated for 45lbs of
force which is much larger than the required 18lbs
giving us a factor of safety of 2.5.
𝐵𝑜𝑟𝑒 𝑜𝑓 𝐴𝑖𝑟 𝐶𝑦𝑙𝑖𝑛𝑑𝑒𝑟 =
1.970′′
= 0.985 ′′
2
𝑉𝑜𝑙. 𝑜𝑓 𝐵𝑢𝑓𝑓𝑒𝑟𝑖𝑛𝑔 = (2.985′′ )2 𝑥 𝜋 𝑥 (758 ′′ )
𝑉𝑜𝑙. 𝑜𝑓 𝐵𝑢𝑓𝑓𝑒𝑟𝑖𝑛𝑔 = 53.36 𝑖𝑛3
Figure 8: THK Linear Actuator
https://tech.thk.com/upload/catalog_claim/pdf/320E_V
LA.pdf
Through research it was determined that a
Yaskawa motor to power our linear actuator would be
the best option. This specific motor comes with a PCI
mechatrolink car that was specifically designed to
accurately control a linear actuator using Labview
VI’s created by Yaskawa. The motor was a 100W
motor with a torque output of 1.15 N/m. The motor
was then coupled to the actuator using a coupling
supplied from THK that fit both the motor and
actuator shafts.
Project P09026
Fall ’08 – Winter ‘08
Page 7
Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
interface was designed to be user-friendly so that the
user can easily define the stroke volume and heart rate
without needed to alter the parameters in the setup
utility.
Figure 9: Yaskawa Servo Motor
http://www.acsindustrial.com/servo-motorrepair/skins/common/pics/yaskawa-servo-motorrepair.jpg
Once the actuator and motor were selected
(selected actuator and motor are shown in Figure 8, 9),
a suitable mounting design for this part of the system
was developed. The main goal of the mounting
included withstanding the continuous operation of the
system and the associated vibrations that were
produced. The mounting was also designed or ease of
manufacture and serviceability of the connected parts.
An important design consideration was low cost. To
dampen the shaking of the overall system duethe
actuator motion, rubber feet were included on the
bottom of the mounting plate. Urethane bushings
were made from urethane stock and press fit into
counter-bores, this would allow for any misalignment
in the mounting system to be dispersed into the
urethane.
CONTROL SYSTEM
The Controls Team’s specific goals for this
project included the development of a system that
provided the user with accurate and precise control
over the actuator parameters. This was achieved by
integrating LabVIEW Vis provided with the Yaskawa
Mecatrolink II control electronics..
The speed and position of the servo motor can be
controlled both by software settings and by altering
the device parameters. Changing the device
parameters is done in the setup utility that is obtained
as part of the Mechatrolink software. The acceleration
of the linear actuator is controlled by the 1st and 2nd
linear acceleration constants and the switching speed.
Each “move” of the motor progresses through the
sequence seen in Figure 11 which defines to
pumpmotion. The acceleration/deceleration constant
and switching speed are used to adjust the amount of
the displacements during flexion. The goal of the
program is to enable the user to control the motion of
the actuator without having to change the setup of the
motor. A LabVIEW© (National Instruments, TX)
Figure 10: Actuator Motion Profile
SIGNAL CONDITIONING
Edwards Lifesciences TruWave and Merit
Medical MeriTrans disposable pressure transducers
were used to measure the Atrial, aortic and venous
pressures.
These transducers utilize strain gauge mechanism
to convert mechanical stress into an electrical signal.
According to authors Webster and Pallas- Arney in
Sensors and Signal Conditioning, strain gages are
based on the variation of resistance of a conductor or
semiconductor when subjected to mechanical stress.
The electrical resistance of a wire with length, l and
cross sectional A, and resistivity ρ can be defined as:
𝑅= 𝜌
𝑙
𝐴
(1)
and for small variations, the resistance for the metallic
wire can be expressed as:
𝑅 = 𝑅0 (1 + 𝐺𝜖) = 𝑅0 (1 + 𝑥)
(2)
where Ro is the resistance when there is no applied
stress, G is the gage factor, which is a constant for any
specific metal.
The bridge circuit, as shown in Figure 12, is based
on a feedback, in order to adjust the value of the
standard until the current through the current meter
indicates zero [1].
Copyright © 2008 Rochester Institute of Technology
Page 8
Proceedings of the Multi-Disciplinary Engineering Design Conference
These transducers utilize strain gauge mechanism
to convert mechanical stress into an electrical signal.
According to authors Webster and Pallas- Arney in
Sensors and Signal Conditioning, strain gages are
based on the variation of resistance of a conductor or
semiconductor when subjected to mechanical stress.
The electrical resistance of a wire with length, l and
cross sectional A, and resistivity ρ can be defined as:
Figure 11: Wheatstone bridge configuration
http://zone.ni.com/cms/images/devzone/tut/a/83a1fe69
766.gif
Signal Conditioning Gain Computations
𝑉max 𝑖𝑛 = 5𝑚𝑉
When the bridge circuit is in balanced condition,
resistance R3 is:
𝑅3 = 𝑅2
𝑅4
𝑅1
(3)
In the equation above, R3 is directly proportional to
corresponding changes in R2 in order to balance the
circuit. This condition is achieved independent to the
supply voltage or the current, and any possible
variations.
The sensitivity of the pressure transducers used
for the system is stated as a voltage:
𝑉𝑜 =
5 𝑢𝑉
𝑚𝑚𝐻𝑔
𝑉
(4)
From the specification above, the change in output
voltage generated by the balanced bridge circuit is
5 𝑢𝑉for a change in one mmHg of pressure and
5 𝑢𝑉for increase for every volt provided to the circuit
in form of excitation.
The system was designed for pressures 0 – 140
mmHg, and therefore according to the sensitivity of
the sensors, the output voltage would range from 0 –
5mV, given a chosen excitation voltage of 5V. Any
variations in the output voltage will be very small to
be captured by the NI USB 6008 data acquisition
board. Hence, signal conditioning must be designed in
order to amplify the signal to amplitudes that are
captureable by the DAQ. Also, the signal conditioning
must eliminate any noise propagating through to the
output, before it detected and transmitted to the
computer. Since the system is going to be used to
measure pressure from 0 -140mmHg, it was logical to
design the signal conditioning to allow minimum of 0200mmHg.
Edwards Lifesciences TruWave and Merit
Medical MeriTrans disposable pressure transducers
were used to measure the Atrial, aortic and venous
pressures.
Hemodynamic Simulator II
𝑉max 𝑛𝑒𝑒𝑑𝑒𝑑 = 2.5𝑉
𝐺𝑎𝑖𝑛 =
2.5𝑉
= 500𝑉/𝑉
5𝑚𝑉
For our purposes, a two stage signal conditioning is
chosen. The first stage is the DC gain stage and the
second stage is identified as the low pass filter stage.
Both the stages were built and tested on a breadboard.
The resistive components deployed in the circuit carry
5% tolerance and therefore small discrepancy was
observed between the simulated and the hardware
gains. The physical gain for all three signal
conditioning was calculated to be 475V/V, 487V/V,
and 481V/V.
Stage 1: DC Gain Stage
This stage is implemented with an Burr Brown
INA128 instrumentation amplifier. This stage provides
a buffer for the input circuitry and more importantly
reduces the common mode noise to great extent. This
stage was designed for a gain of 50V/V.
𝐺𝑎𝑖𝑛 = (1 +
50 = (1 +
50 𝑘𝛺
)
𝑅𝐺
(5)
50 𝑘𝛺
) → 𝑹𝑮 = 𝟏. 𝟎𝟐 𝒌𝜴
𝑅𝐺
Stage 2: Low Pass Filter w/ small gain
The output from the first stage is treated as an input to
the second stage of the signal conditioning circuitry.
Since the first stage is providing 50V/V of the overall
500V/V gain, only 10V/V is needed out of this stage.
The low filter is implemented with a negative
feedback, based non-inverting op-amp, as shown in
Figure 2.
Project P09026
Fall ’08 – Winter ‘08
Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
Page 9
to generate motion profile based on the user entered
value for stroke volume. The flow meter was setup to
measure volume of water pumped though the
circulatory loop every minute, which is defined as the
cardiac output of the heart. Given the cardiac output
and stroke volume, equivalent heart rate can be
computed using following relationship:
Figure 12: Low Pass Filter Design with a single supply
𝐺𝑎𝑖𝑛 = (1 +
10 = (1 +
𝑅3
𝑅2
)
(6)
𝐻𝑒𝑎𝑟𝑡 𝑅𝑎𝑡𝑒 (𝐻𝑅) =
𝐶𝑎𝑟𝑑𝑖𝑎𝑐 𝑂𝑢𝑡𝑝𝑢𝑡 (𝑄)
𝑆𝑡𝑟𝑜𝑘𝑒 𝑉𝑜𝑙𝑢𝑚𝑒
Figure 14, 15 & 16 are the measured aortic, arterial
and ventricular pressures respectively.
𝑅3
𝑅3
)→
=9
𝑅2
𝑅2
𝑹𝟐 = 𝟏 𝒌𝜴, 𝑹𝟐 = 𝟗 𝒌𝜴
The pressures associated with human cardiovascular
system are known to be around 20Hz. And therefore
the low pass filtering is designed with a roll frequency
of 50Hz.
𝐹0 =
𝐹0 =
1
2𝜋 𝑅1 𝐶1
(7)
1
1
→ 50 =
2𝜋 𝑅1 𝐶1
2𝜋 𝑅1 𝐶1
C1 is arbitrarily selected to be 0.33 μF.
50 =
Figure 13: Aoritc Pressure Output for Cardiac Output of
70mL
1
→ 𝑹𝟏 ≈ 𝟏𝟎𝒌𝜴
2𝜋 𝑥 𝑅1 𝑥 0.33
The schematic of the overall signal conditioning
circuitry is shown in Figure 4.
Development Procedures
Software Development
1. Actuator control
2. Flow and Pressure Measurements
TESTING THE FINAL DESIGN
The system was tested by measuring and
recording the aortic, atrial, and ventricular pressures as
well as the flow through the circulatory loop for
different stroke volumes. The stroke volume was used
as parameter to control the velocity of the actuator. In
a healthy human, the stroke volume is measured to be
around 70mL.The control software was programmed
Figure 14: Arterial Pressure Output for Caridac Output
of 70mL
For the first five trials, all three pressures were
measured with disposable pressure transducers and
Omega industrial grade transducers, to confirm proper
functioning of signal conditioning circuitry. The
pressure generatd by medical grade pressure
Copyright © 2008 Rochester Institute of Technology
Proceedings of the Multi-Disciplinary Engineering Design Conference
10
transducers was measured to be within ±10mmHg of
the pressure measured by the Omega transducer.
Page
Refinement of flow compliance and resistance
mechanisms as well as final system integration, testing
and validation are other suggested areas of future
enhancements.
Currently, the control system can simulate for the
system for one motion profile and arbitrary motion
input is not possible. Pressure and flow readings are
accessible to the user graphically and numerically.
With the numerical data, a variety of manipulations
can be carried out to monitor different hemodynamic
activities in real time.
CONCLUSIONS
Figure 15: Ventricular Pressure Output for Cardiac
Output of 70Ml
The shape and amplitudes of the pressure
waveforms are relatively similar to an actual human
pressure waveforms. Further tuning of the actuator
parameters would results in a more accurate pressure
readings. From figure 13, 14 & 15, the concept of a
hemodynamic simulator was successfully proved.
HIGHLIGHTS OF FINAL DESIGN
The current Hemodynamic Simulator is intended
to be a mechanically and electrically robust system,
which is capable for generating controlled motion of
the actuator, which further provides a controlled fluid
flow through the circulatory loop.
Further enhancements made to the simulator
allows for easy transportability from one classroom to
another. Easy filling and draining procedures have
been implanted. The new Mechatrolink II PCI board
provides not only provides a direct electronic link
from the servo to computer but also to LabVIEW,
which allows for accurate and real time control of the
system.
The final design of Hemodynamic Simulator
successfully mimics the activity of left ventricular of a
human heart. Although, the not all parameters are fully
controllable, the simulator has gone through multiple
folds reduction in size, making the system portable.
Reduction in number of hose clamps in the circulatory
loop is another major customer spec that the final
design complies with.
ACKNOWLEDGMENTS
The team would like to express its sincerest
gratitude to those who have made invaluable
contributions to this project. Many thanks to the
advisors, Dr. Daniel Phillips and Dr. Karl Schwarz for
their guidance and support. Additionally, the team
would like to express thanks to consultants, Mr. John
Wellin, Dr. Steven Day, Dr. Mark Kempski, Dr.
Jeffrey Kozak, who provided prompt and important
assistance, when necessary.
REFERENCES
[1] Pall, Ram, and John G. Webster. Sensors and
Signal Conditioning. New York: Wiley-Interscience,
2000.
FUTURE WORK
In future senior design projects, the primary focus
should be to enhance the simulator’s controllability.
Hemodynamic Simulator II
Project P09026
Fall ’08 – Winter ‘08