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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
|
|
Data Acquisition Team
Jonathan Peyton |
Pump Design Team
Gaurav Zirath
Team Leader
|
ABSTRACT
The Hemodynamic Simulator is a modular system
that is 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 hrough a mock circulatory loop. The simulator
will be useful in research related to cardiovascular
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,
CO – cardiac output, is the amount of blood pumped
by the heart, a ventricle in particular, in a minute.
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.
Copyright © 2008 Rochester Institute of Technology
Proceedings of the Multi-Disciplinary Engineering Design Conference
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 dataincluding 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.
MECHANICAL CHAMBER’S DESGIN
Previously the cardiovascular loop and pressure
generation loop were viewed as being overly
complicated and unable to meet the desired
performance criteria for the simulator 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 components in the previous system. Cast
acrylic was chosen for its properties, availability and
ease of fabrication.
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designedto have a surface area at least twice the size of
the aortic chamber resulting in a cylinder with asix
inch OD by 5.75 inch ID was chosen for the atrial
chamber. The inlet and outlet also incorporateone 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..
Critical aspects of the design of the chambers
were ease of manufacture and ease of assembly of the
overall system. Common hose fasteners were chosen
and a common top design for both the aortic
compliance chamber and atrial reservoir was
implemented to ensure proper sealing while allowing
easy emptying and filling of the chambers. The
chamber and 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 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 inner 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 incorporates one inch NPT thread, which
was also chosen to be the common fitting 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.
Figure 1: Display of taper in the inlet from the
Ventricular chamber.
Determing the 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
Hemodynamic Simulator II
Project P09026
Fall ’08 – Winter ‘08
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Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
of clear plastic machining. Some of these include:
”crazing” or miniscule cracking during welding,
overheating during machining, and solvent selection
and surface preparation.
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 liststhe cutting
speeds and feeds used as guidelines for machining
operations.
Figure 2: Cutaway of top ring and top lid fastener
interface.
Machining Acrylic ( w/ HSS)
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.
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
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
solvent out of the joint creating a “dry” bond. Table 2
provides the developed solvent welding guidelines.
Solvent Welding Acrylic
Figure 3: 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
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.
Copyright © 2008 Rochester Institute of Technology
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Proceedings of the Multi-Disciplinary Engineering Design Conference
Supplier
www.rplastics.com
Cure time
24 hrs. (1 hr. till handlable)
Table 2: Developed solvent bonding 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.
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.
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.
Hemodynamic Simulator II
Figure 4: THK Linear Actuator
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.
Figure 5: Yaskawa Servo Motor
Figure 6: Yaskwawa Servo Controller
Once the actuator and motor were selected, 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
Project P09026
Fall ’08 – Winter ‘08
Figure 7: Pump Motion
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Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
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.
𝑅= 𝜌
𝑙
(1)
𝐴
and for small variations, the resistance for the metallic
wire can be expressed as:
𝑅 = 𝑅0 (1 + 𝐺𝜖) = 𝑅0 (1 + 𝑥)
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 6 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)
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. In order to achieve this goal the control team
where Ro is the resistance when there is no applied
stress, G is the gage factor, which is a constant for any
specific metal.
A Wheatstone bridge is known be an effective
method for measuring small resistances. This
technique was first proposed by S.H. Christie in 1833
and then reported by Sir Charles Wheatstone to the
Royal Society (London) in 1853. The bridge circuit, as
shown in Figure 7, is based on a feedback, in order to
adjust the value of the standard until the current
through the current meter indicates zero [1].
Figure 8: Wheatstone bridge configuration
http://zone.ni.com/cms/images/devzone/tut/a/83a1fe69
766.gif
When the bridge circuit is in balanced condition,
resistance R3 is:
𝑅3 = 𝑅2
SIGNAL CONDITIONING
The pressure transducers used to measure the
Atrial, aortic and venous pressures are disposable,
medical grade pressure transducers donated to the
team by Dr. Schwarz. The pressure transducers are not
normally commercially available for research or
industrial use.
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:
(2)
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.
Copyright © 2008 Rochester Institute of Technology
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Proceedings of the Multi-Disciplinary Engineering Design Conference
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. The signal conditioning for the pressure
transducers was designed for 0 – 200mmHg as
indicated in equations …
50 = (1 +
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.
Signal Conditioning Gain Computations
𝑉max 𝑖𝑛 = 5𝑚𝑉
𝑉max 𝑛𝑒𝑒𝑑𝑒𝑑 = 2.5𝑉
𝐺𝑎𝑖𝑛 =
2.5𝑉
= 500𝑉/𝑉
5𝑚𝑉
Figure 9: Low Pass Filter Design witha single supply
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.
Due to wide range of tolerances on the components
used, the physical gain was calculated to be 475V/V
and 487V/V.
Stage 1: DC Gain Stage
This stage is implemented with an 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 +
Hemodynamic Simulator II
50 𝑘𝛺
)
𝑅𝐺
𝐺𝑎𝑖𝑛 = (1 +
10 = (1 +
𝑅3
𝑅2
)
(6)
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.
(5)
Project P09026
𝐹0 =
𝐹0 =
1
2𝜋 𝑅1 𝐶1
(7)
1
1
→ 50 =
2𝜋 𝑅1 𝐶1
2𝜋 𝑅1 𝐶1
Fall ’08 – Winter ‘08
Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
C1 is arbitrarily selected to be 0.33 μF.
50 =
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
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.
FUTURE WORK
In future senior design projects, the primary focus
should be to enhance the simulator’s controllability.
Refinement of flow compliance and resistance
mechanisms as well as final system integration, testing
and validation are other suggested areas of future
enhancements.
Page 7
development of. It would be very useful if developa
theoretical model that would represent the dynamics of
the overall system were developed.
Lastly, future groups may also reduce the size of
the overall system significantly by replacing a regular
sized PC with a smaller single board computer, and
furthermore replace the LCD monitor with a touch
screen, which would eliminate the need for a dedicated
keyboard and mouse.
CONCLUSIONS
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.
The final product serves as a valuable teaching
tool, more importantly a research tool to dynamics of
blood flow. The simulator may also be used to test
operations of heart valve, LVAD and other assist
devices. The project may also be used to showcases
the newly developing bioengineering discipline at
Rochester Institute of Technology.
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
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. team couldn’t finish
[1] Pall, Ram, and John G. Webster. Sensors and
Signal Conditioning. New York: Wiley-Interscience,
2000.
Copyright © 2008 Rochester Institute of Technology