Multi-Disciplinary Engineering Design Conference
Kate Gleason College of Engineering
Rochester Institute of Technology
Rochester, New York 14623
Project Number: 07041
EXPANDABLE ALVEOLAR SAC
Christopher Salvato/Project Manager
Timothy Brackbill/Manufacturing Lead
Nicholas Daniel/Imaging Lead
William Dapolito/Pumping Lead
Jessica Oakes/Air Sac Model Lead
Matthew Waldron/Controls Lead
ABSTRACT
An expandable alveolar sac was created for
research applications. The system will be used to
profile alveolar fluid flow using PIV and to explore
fluid mixing dynamics during respiration. The system
was designed and implemented to be easily modified
as necessary for applications using breathing curves
for different diseases and different subject age groups.
This is a requirement to allow for a variety of other
possible research applications. System was consistent
with fluid flow properties in the human lung and
compatible with Particle Image Velocimetry (PIV)
technology.
INTRODUCTION
The objective of this design is to create an in vitro
alveolar sac that creates fluid flow and pressure
parameters similar to those parameters that occur in
vivo. The main goal of creating these parameters in
vitro is to map velocity field profiles using PIV
technology.
Secondary goals include using the
system to map fluid velocity profiles through layering
fluids of varying densities as well as analyzing fluid
mixing within an alveolar sac.
A previous attempt was made at this design and did
not succeed. There were several reasons for this
failure. Firstly, the pump access into the negative
pressure chamber did not allow for adequate fluid flow
to create pressure changes quickly enough within the
bottom cavity. Secondly, the material that was used
was neither optically clear nor was it flexible with
elastic properties. Finally, the method of fabrication
of various sized sac models was costly as well as
ineffective. A triumph of the previous design,
however, was the rough creation of the air sac model.
This model, constituted of complex geometries,
allowed the current group to focus on more practical
problems of the design rather than focusing on
creating a complicated air sac model. This model was
then redesigned and improved during the course of the
current project. These improvements allow for a more
accurate in vitro representation of in vivo geometries.
Overall, the current project was a major success.
All primary customer needs were met. Additionally,
most secondary and tertiary customer needs were met
and exceeded as well. While the project was a great
success, it has also been designed to easily facilitate
change and improvement for possible future research
applications.
NOMENCLATURE
casting - to form into a particular shape creating a
mold in a fluid state and letting it harden
elastomer - a polymer having the elastic properties of
natural rubber
in vivo – within a living organism
in vitro – in an artificial environment outside of the
living organism
molding – creating the mold that will be used to cast
alveolar sacs
plasticizer – a substances added to plastics or other
materials to make or keep them soft or pliable
rapid prototyping - an additive manufacturing
process that creates a model of an object directly from
a CAD model by building it in layers.
syringe pump - a device consisting of a glass, metal,
or hard rubber tube, narrowed at its outlet, and fitted
© 2007 Rochester Institute of Technology
Proceedings of the Multi-Disciplinary Engineering Design Conference
with either a piston or a rubber bulb for drawing in a
quantity of fluid or for ejecting fluid in a stream.
velocity field profiling – discerning vector fields of
velocity based on fluid flow observations.
DESIGN METHODOLOGY
Overview
The design of the system was broken down into 5
logical subsystems.
Each subsystem had a set of
needs that they must satisfy. These subsystems
include Imaging, Air Sac, Manufacturing, Pumping,
and Controls.
Imaging
One of the highest level needs for this project was
that the system be compatible with PIV. This is an
optical method used to measure the velocity profile in
a given fluid.
PIV uses a laser to illuminate
fluorescent particles that are immersed in the fluid.
Particle settling is undesirable, for the particles need to
move with the motion of the fluid, which will allow
for accurate fluid velocity measurements using PIV.
Three different forces act on a particle suspended in
fluid. These include buoyancy force, FB, force due to
gravity, Fg, and force due to drag, FD. The buoyancy
force, shown in Eq. (1), where ρf is the density of the
fluid, g is gravity and Vp is the volume of the particle.
To calculate the drag force Eq. (2) was used where dp
is the diameter of the particle and urel is the relative
velocity of the particle. The force due to gravity can
be calculated with Eq. (3), where ρp is the density of
the particle. It was assumed that the particle will have
reached terminal velocity when the fluid is in motion.
This assumption allows for the summation of forces to
equal zero as shown in Eq. (4).
FB = V p *g * ρ f
(1)
FD = 3 * π * d p * u rel
(2)
Fg = V p * g * ρ p
(3)
ΣFy = 0 = FD + FB − Fg
(4)
The relative velocity was solved for, which is the
difference between the velocity of the particle and the
velocity of the fluid, and it was determined to be
8.06E-9m/s. With a time period of twenty seconds the
particle will separate from the fluid a distance of 1.6E5m, which was determined to be negligible.
To ensure that the PIV will collect data accurately, it
was essential that the index of refraction of the sac
material and the surrounding fluid be matched as
closely as possible. Refraction occurs when light
waves travel between mediums of different refractive
indexes, causing the light waves to bend at a certain
angle. The index of refraction of the sac material was
measured with a refractometer to be 1.482, and the
Page 2
fluid selected, glycerin, had an index of refraction of
1.473. Using Snell's Law, a relationship which relates
the angle of incidence and angle of refraction to the
refractive indexes of the mediums, it can be found that
any disparities resulting from the slight difference in
index of refractions between the sac material and
glycerin was be negligible. Aside from this major
concern, most other concerns of the imaging
subsystem were handled through design and
implementation of the other subsystems. As a result,
all needs for this subsystem were fulfilled.
Air Sac
The primary needs addressed by the air sac
subsystem involve creating accurate geometries to
properly replicate an in vivo alveolar sac. This
requires analysis of the fluid flow parameters and how
they relate to geometry.
Fluid
The alveolar sac in the human lung has a volume of
3.7E-11 m3, which was too small to model in a
laboratory setting; therefore a larger model had to be
designed in order to allow for it to be manufactured.
The Navier-Stokes partial differential equation, as
shown in Eq (5), describes fluid flow, where ρ is the
fluid density, u is the speed of the fluid, t is time, P is
the pressure, ∇ is the gradient in the x, y, and z
direction, μ is the fluid viscosity, and g is gravity. The
Navier-Stokes equation must be non-dimensional in
order to properly scale the model. This is shown in
Eq. (6), where ω is the frequency, D is the
characteristic diameter, and ν is the kinematic
viscosity.
⎡ δu
⎤
+ u * ∇u ⎥ = −∇P + μ∇ 2 u + ρg
⎣ δt
⎦
ρ⎢
⎡ p * g * D2 ⎤
⎡ ρ * D 2 * ω ⎤ δ u ⎡ ρ * D * vo ⎤
⎡ Ps * D ⎤
2
+⎢
⎥
⎢
⎥
⎥u * ∇u = − ⎢
⎥∇ P + ∇ u + ⎢
μ
μ
⎦
⎣ μ *v ⎦
⎣
⎦ δt ⎣
⎣ μ * vo ⎦
(5)
(6)
Two dimensionless parameters that describe inertial
flow were characterized from Eq. (6): Reynolds (Re)
and Womersley (Wo) numbers. These are shown in
Eq. (7) and Eq. (8), respectively.
ρ * D * vo
μ
ρ * D2 *ω
Wo =
μ
Re =
(7)
(8)
Both the Reynolds and Womersley numbers were
calculated for in vivo conditions. The dimensionless
parameters were determined to be much less than one,
characterizing the fluid to be lack inertia. This is
beneficial since it is necessary for the fluid to lack
Paper Number 07041
Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
Page 3
inertia in order to create a model for the laboratory
setting.
Data was generated based on the size of the model
and the time required to complete a cycle for a model
that expands thirty percent in a pure glycerin fluid.
Pure glycerin was used to match the index of
refraction of the material and thirty percent expansion
was determined from literature [4]. The laboratory
model controls are able to be modified based on the
data determined from the dimensionless parameters.
Geometry
The previous senior design group had used idealized
alveolar sac geometry was used to create the original
alveolar sac model. Three different alveolar sac
geometries were created. The previous senior design
team created the first model, shown in Fig. 1, using
75% Total Lung Capacity (TLC) Model. These
geometries were obtained from literature [1,4].
Because the dimensions from the literature were
measured from a lung cast that was created at 75%
TLC, the alveoli had to be scaled down to breathing
volume which is 50% TLC. This value is referred to
as the Functional Residual Capacity (FRC). Each one
of the alveoli were scaled to FRC using Eq. (5), where
dVm is the diameter of the alveoli at the measured
volume, dVFRC is the diameter at FRC, Vm is the
volume at which the diameter was measured and VFRC
is the volume at FRC.
⎛ V
d Vm
= ⎜⎜ m
d VFRC ⎝ VFRC
⎞
⎟⎟
⎠
1
3
(9)
The alveolar sac created was compared to geometry
dimensions published in literature [2, 3]. Values from
literature were produced by measuring the shape of a
alveoli by measuring a dimensionless parameter. This
parameter is the ratio of alveoli depth and alveoli
mouth diameter. The 75% TLC model and the 50%
TLC model dimensionless parameter was much
smaller than the data measured in literature. In order
to create a model that represented dimensions found in
literature, a model was created with fewer alveoli.
Each of the alveoli were more defined in the model,
therefore the dimensionless number was comparable
to that found in the literature [2, 3].
Manufacturing
The primary concerns that the manufacturing
subsystem had to address involved creating a physical
model of the alveolar sac. The process had to be
efficient and easily repeated so that many different
models of various sizes could be used in the system.
Also, the sac model had several material requirements.
The material had to be optically clear to satisfy the
Fig. 1. Original alveolar sac model with measurement locations
corresponding to duct diameter, sac diameter and sac length
labeled.
requirements for PIV. Also, elastic properties had to
exist that were similar to the elastic properties of
alveolar sacs in vivo.
Material
In order to satisfy the material requirements several
different materials were researched. Samples were
obtained and their properties were analyzed. At this
point, the chosen method of fabrication involved
dipping a mold into the molten material and letting the
resulting model cure. This meant that the material had
to have a reasonable melting point, must meet the
requirements of PIV, maintain its physical and optical
properties after being melted and would cure quickly.
An elastomer material called UltraFlex was found
through a supplier. After testing, this material was
shown to have all desired material properties,
including an index of refraction that matched that of
glycerin. A process was then developed to create
alveolar sac models from this material.
Molding
The first attempt was to dip a rapid prototyped male
mold of our alveolar sac coated with epoxy. These
attempts had shown that most organic chemicals, such
as epoxy and rapid prototyping material, reacted with
the elastomer. These reactions caused changes to the
optical and physical properties of the material. This
method, therefore, ruined the cast. We found however
that when metals or glass were used to create the cast
no reactions occurred. This process also yielded
adequate surface finishes. From this realization it was
decided to pursue options in creating aluminum molds.
The RIT School for American Crafts had provided a
means of creating aluminum molds in an efficient and
cost effective manner.
The process of creating aluminum molds started
with an 8 piece, rapid prototyped female mold of the
Copyright © 2007 by Rochester Institute of Technology
Proceedings of the Multi-Disciplinary Engineering Design Conference
alveolar sac. The mold was filled with molten wax
and set to cool to create a male model from the mold.
Next, this male wax model was removed and cleaned
of any unintended effects. From here, male ducts were
attached to the wax mold to allow aluminum to enter.
This created a sand cast of the wax model. When a
sand cast was made, it was taken to the kiln and the
aluminum is melted. Molten aluminum was then
skimmed to remove defects from the metal that float to
the surface. The molten aluminum was then poured
into the sand cast and allowed to cool. When the
aluminum had thoroughly cooled, the sand cast was
opened and the metal casting was then removed. The
additional ducting was sawed off of the top of the
completed aluminum casts so that only the alveolar
sac remains. The top duct of the aluminum cast was
then lathed to the desired diameter and the cast of the
mold was sanded to a smooth surface finish.
Casting
The procedure to cast the alveolar sac molds was
devised iteratively until a process that produced
desired results consistently was produced. The main
element of the process involved the dipping of the
aluminum mold into molten material and then curing.
After nine different processes were developed and
tested a final process was decided upon based on the
results of the testing.
First, scissors were used to dice the required amount
of elastomer into cubes that were approximately 0.2”
per side. Next, a beaker was placed on a hot plate in a
fume hood and the hot plate was turned to the
maximum setting. At this point, approximately ten to
twelve cubes of the elastomer were placed into the
beaker and heated from the top with a heat gun. When
the elastomer started to melt, it was stirred while still
applying heat from the hot plate and heat gun. When
the elastomer in the beaker had mostly melted, more
elastomer was added until the necessary amount of
elastomer was melted to dip the mold. When the
required amount of elastomer was melted, heat was
continually applied with both the hot plate and heat
gun while stirring until the elastomer lost most of its
viscous properties. The molten material was then left
idle under the heat until all bubbles evacuated from the
fluid. The top duct of the metal mold was then placed
into vice grips and heated under the heat gun until the
mold achieved a temperature between 120 ºC and
175ºC. This was done to prevent the elastomer from
rapidly cooling and solidifying when dipped. At this
point, the heated aluminum mold was slowly dipped
into the beaker of molten elastomer. This was done
extremely slowly so that no bubbles formed in the
sharp corners of the mold. When the mold had been
fully immersed in the molten elastomer it was
removed from the beaker. Immediately upon removal
from the beaker, the mold was rotated along the axis
Page 4
Fig. 2. Original concept art of system design. The top cavity
holds fluid to flow into the sac. The bottom cavity creates a
negative pressure system driven by a syringe pump.
of the vice grips while the elastomer was still molten
in order to even out any discontinuities from dipping.
Immediately after the outside of the elastomer had
hardened, the mold was placed with elastomer film
under slow-running cold water tap for approximately
five minutes. This was to prevent plasticizers from
evaporating and to decrease the cooling time. Once
the mold had completely cooled, the elastomer film
was rolled off of the mold. This was done while
ensuring that the film did not rip on the way off. The
molds were then sealed in a container for storage to
assure that no dirt or dust from the environment would
adhere to them.
Pumping
The pumping subsystems had to meet primary needs
that were associated with creating the negative
pressure environment that would drive the expansion
of the alveolar sac. This consisted of two main parts;
the system housing and pump design.
Housing
There were many designs for the housing that the
team produced in the first few weeks of the project.
Some ideas included a rigid-bottomed container, a
piston-bottomed container, and a diaphragm flexible
bottom. There were many different methodologies that
were assessed to change the volume with each of these
ideas. Some of these included using one or more
syringe pumps, an eccentric crank, and a magnetic
actuator. It was decided that the final design was to be
a rigid-bottomed container with a single syringe pump
that would evacuate fluid from a bottom cavity. This
bottom cavity would house the alveolar sac. The
alveolar sac would be attached to the bottom of a lid
that would also create a top cavity. The top cavity
would hold fluid that would flow into the alveolar sac
and be used to measure volume within the sac. This
original concept design is shown in Fig. 2.
Paper Number 07041
Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
A 3D model was then created based on concept art
using the design software SolidWorks. Many inherent
problems with the first design iteration were apparent.
Firstly, the upper portion of the housing was too large
to adequately monitor volume change. To alleviate
this problem, the upper cavity was replaced with a
graduated cylinder that allowed for fluid access into
the alveolar sac model. Secondly, the through-wall
fitting in the first design was much too small. Since it
was assumed that the fluid within the cavities would
be glycerin, the head loss on such a small through wall
fitting would have been much too great unnecessarily
using a small through wall fitting. Hence, the second
design iteration implemented a larger through wall
fitting. The first design included a looking glass in
order to assure PIV compatibility. This looking glass
was to provide the PIV laser and camera optical access
to the inside of the bottom cavity. Lastly, the second
design iteration eliminated this looking glass and used
lexan for the housing walls so that the entire housing
was optically clear.
A review of the housing design yielded another
design iteration resulting in more refinements. The
graduated cylinder on the upper portion was made to
have a smaller diameter thereby giving volume
measurements more resolution. An even larger
through-wall fitting was also added to allow for an
easier accommodation of pressure changes in the
bottom cavity. Washers and bolts were added to the
detail of the housing model. Lastly, a pressure sensor
was added to the assembly.
A final review of the housing design was done
creating the final iteration of the housing design. Due
to leaks around the seal between the housing lid and
the bottom cavity an aluminum collar was made. This
collar more evenly distributed the pressure across the
sealing gasket. Threaded holes were also added for a
bleed valve and a mechanical pressure gauge. This
final design is shown in Fig 2.
Pumping
A syringe pump system was chosen to create the
negative pressure system. By attaching the syringe
pump to the housing’s through wall fitting it was
possible to have a syringe pump evacuate fluid from
the bottom cavity causing negative pressure.
To meet all of the required needs associated with
pumping a pump had to be decided upon without
knowing the operating pressures and the volumetric
flow rate ranges. This is due to the amount of
unknown parameters that existed when the design
process began. To find a starting point for the pump
selection, many assumptions were made. First, it was
assumed that the maximum alveolar sac size would be
the same as the previously attempted Senior Design
Project. With this information it was possible to find
the maximum flow rate. Secondly, it was assumed that
the working fluid would be glycerin. From this
Page 5
Fig. 3. Final system design. The top cavity had been reduced to
a single graduated cylinder. The bottom cavity creates a negative
pressure system driven by a syringe pump with access for
multiple sensors and is sealed by a gasket and aluminum collars.
assumption, viscous losses were calculated. Then, the
necessary additional pressure due to viscous losses
alone was calculated. Lastly, the pressure to expand
the alveolar sac was grossly over estimated to ensure
that the chosen pump would work. The operating
pressure for the pump was then calculated by adding
the pressure due to viscous losses to the estimated
pressure required to expand the alveolar sac model.
The previous Senior Design group had purchased an
expensive pump that they had attempted to use. This
pump, a NE-500 Syringe Pump System from New Era
Pump Systems, allowed for a syringe pump to
interface with a motor that is controlled by a
microcontroller. This would allow for control with a
computer, as desired by the customer as well as
possibly providing the required specifications that had
been calculated.
Testing and analysis had shown that simple
modifications could be made to yield the necessary
output parameters from the motor. The first and most
simplistic modification was to change the syringe size.
More specifically, the diameter of the syringe was
greatly increased. By increasing the syringe diameter
and keeping the linear velocity constant, the
volumetric flow rate increases. The original syringe
had a 60 ml capacity. This was replaced with a 140 ml
syringe. The larger syringe also had a 50% increase in
diameter over the original. Secondly, the gear ratio
between the stepper motor and the driving lead screw
was modified. The original syringe pump had a gear
reducer setup. This means that the stepper motor’s
shaft speed was greater than the lead screw’s shaft
speed. This setup was changed to a speed multiplier,
which. This means that the stepper motor’s shaft
speed was less than the lead screw’s shaft speed. This
was accomplished by replacing the belt and pulley
system with a chain and sprocket system. With this
modified system many gearing ratios were possible.
This makes power transmission within the system very
flexible for different volumetric flow rates. Since this
Copyright © 2007 by Rochester Institute of Technology
Proceedings of the Multi-Disciplinary Engineering Design Conference
made the lead screw rotate much faster, the flow rate
increased.
The trade-off associated with this
modification was between speed and torque. In this
instance, torque was sacrificed where a speed increase
occurred. To compensate for this, the stepper motor
was upgraded to a more powerful model with higher
overall torque and speed. In order to perform this
upgrade, the circuitry of the microcontroller board had
to be modified by changing several resistors as well as
other electronics.
Other various enhancements to the pump were made
in order to improve overall functionality. When the
base that holds the lead screw sprocket was modified,
it structurally weakened the pump’s base. To correct
this, a new base was constructed out of aluminum.
Other improvements include two thrust bearings and a
shaft collar that were added to the lead screw for better
stability.
The summation of all of these modifications allowed
the syringe pump meets all customer specifications.
Controls
The two primary needs that were met by the controls
subsystem include driving the model alveolar sac to
expand and contract similar to in vivo breathing
patterns and to monitor model volume and pressure.
The latter need had to be done such that the
experimental model data can be compared with in vivo
data.
The sac expansion and contraction was
controlled by the negative pressure system created by
the pumping subsystem. The rate of fluid fluid flow
was controlled by a LabVIEWTM program that
continually changed the pumping rate to match a
desired transient breathing curve.
An example of the in vivo tidal volume curve that
the LabVIEWTM program matched is shown in Fig. 4.
After this curve was converted to a digital and discrete
waveform, the volume and time scales were
normalized to match the shape of the desired breathing
curve. Within LabVIEWTM, the volume and time
scales were denormalized to the model and experiment
parameters put in by the user. These parameters
included the unexpanded model volume, peak percent
expansion, and breath period.
The NE-500 syringe pump was capable of receiving
commands that dictate the rate of either withdrawing
or infusing fluid. To make the pump action match the
breathing curve, the LabVIEWTM program sampled
the curve every several milliseconds, and calculated
the desired change in volume to compute the flow rate.
The command to change the rate to the calculated
value was sent to the pump, and this process was
repeated approximately 20 times per second to
generate a smooth change in volume that very closely
matches the desired curve.
To accommodate different activity levels and
breathing pathologies, the LabVIEWTM program was
designed to be flexible enough so that any shape curve
can be loaded into the system. This allowed for
Page 6
Fig. 4. Sample Tidal Volume vs. Time Breathing Curve. The
Labview program must change the pumping rate to make the
model volume displacement match in vivo patterns. Image
adapted from Slonim, N. B., Hamilton, L. H. Respiratory
Physiology. C.V. Mosby 1971
operation over a range of breath periods and expansion
volumes. So long as the desired operation did not
exceed the limits of the pump, the system performed
well and matched the true change in volume closely to
the desired change.
In order to monitor how well the model activity
actually followed the desired operation, two pressure
sensors were used: one was located in the wall of the
housing to measure the pressure produced by the
pump to displace the sac volume, the other was
located at the duct of the model to measure the model
volume change via the hydrostatic pressure of the fluid
outside of the model. Since the hydrostatic pressure of
the fluid outside of the model would always be
positive (with respect to atmospheric pressure) the
sensor was chosen to be sensitive enough so that there
was a large change in output voltage over the expected
range of fluid heights. The pressure caused by the
pump was expected to be both positive and negative,
depending on either exhalation or inhalation of the
model. Since the sensor will only read pressure
positive with respect to atmospheric pressure, the
sensor is placed at the very bottom of the housing
enclosure. This way, the hydrostatic pressure of the
fluid had positively offset pressure fluctuations created
by the pump. This allowed the measured pressure to
always be within the range of the sensor.
The output of the two sensors was continuously
measured with LabVIEWTM. The changes in sensor
output voltage were correlated to the sac volume and
pressure outside of the sac. The results were recorded.
To measure the flow rate of fluid into and out of the
sac, the derivative of the sac volume was be computed
within LabVIEWTM. After an experiment was run,
transient plots of volume, flow rate, and pressure can
be obtained, which provided a good indication of how
well the true model behavior matched the desired
behavior. Literature from experiments on in vivo
human respiration often present graphical relationships
between pressure vs. volume, and flow vs. volume, so
these plots can also generated from the recorded
sensor readings.
After testing the LabVIEWTM program with the
optimum model size and with various percent
expansions and breath periods, the actual model
moved very similarly to the model behavior desired by
the input waveform. It was evident that there is a
maximum and a minimum pumping rate at which the
Paper Number 07041
Proceedings of the KGCOE Multi-Disciplinary Engineering Design Conference
pump could operate. This problem emerges due to the
nature of the stepper motor driving the pump. These
minimum and maximum boundaries can be overcome
by modifications to the pump as explained in the
pumping subsystem.
Also, another method of
alleviating this problem was to change speed and
sampling constants within the LabVIEWTM source
code.
This improvement, however, comes at the
expense of accuracy matching the desired breathing
curve.
For ideal specifications however, the
LabVIEWTM control system meets all the necessary
needs.
ACKNOWLEDGMENTS
The team would like to acknowledge and thank Dr.
Risa Robinson who was the team's guide, sponsor and
customer. Secondly, the team acknowledges the
Provost Learning Initiative Grant provided by the
office of the Provost at RIT. Additionally, the team
would like to thank all consultants who provided
prompt and relevant assistance as well as consultative
advice when necessary. These consultants include Dr.
Steven Day, Dr. Steve Weinstein, Dr. Elizabeth
DeBartolo, Dr. David Borkholder, Dr. Daniel Phillips,
Dr. Kathleen Lamkin-Kennard and Ms. Jackie Russo.
Page 7
The team thanks the School for American Crafts
(SAC) at RIT that has contributed both their time and
effort in creating the aluminum alveolar sac molds.
Also, the team thanks Mr. Jeffrey G. Lonneville of the
Surface Mount Technology Lab at RIT for his
assistance in replacing surface mount resistors.
Finally, the team thanks William Leonard of the
Mechanical Engineering Technology department at
RIT for his assistance in creating the rapid prototypes
of the molds.
REFERENCES
[1] Haefeli-Bleuer B and Weibel ER, 1988,
Morphometry of the human pulmonary acinus.
Anat.Rec. 220: 4: pp.401-414.
[2] Klingele, T G Staub,N C, 1970, Alveolar shape
changes with volume in isolated, air-filled lobes of cat
lung. 28: pp.411.
[3] Mercer, R R Laco, J M Crapo,J D,1987, Threedimensional reconstruction of alveoli in the rat lung
for pressure-volume relationships. 62: 1480
[4] Weibel ER, 1964, Morphometrics of the lung.
Handbook of applied physiology. Respiration.
[5] Slonim, N. B., Hamilton, L. H. Respiratory
Physiology. C.V. Mosby 1971
Copyright © 2007 by Rochester Institute of Technology