U.S. patent application number 11/273888 was filed with the patent office on 2006-07-13 for ex vivo remodeling of excised blood vessels for vascular grafts.
This patent application is currently assigned to The Trustees of the University of Pennsylvania. Invention is credited to Valerie Clerin, Keith Gooch, Rebecca Gusic.
Application Number | 20060155164 11/273888 |
Document ID | / |
Family ID | 26861410 |
Filed Date | 2006-07-13 |
United States Patent
Application |
20060155164 |
Kind Code |
A1 |
Clerin; Valerie ; et
al. |
July 13, 2006 |
Ex vivo remodeling of excised blood vessels for vascular grafts
Abstract
The present invention provides an ex vivo vascular remodeling
methods and system by which an excised, small diameter blood vessel
can be harvested and expanded to provide viable vascular grafts, as
demonstrated at the physical and molecular levels, and as optimized
in vivo. The tissue-engineered vessels generated by the present
invention closely resemble native vessels in terms of structure,
histologically, including endothelial coverage and intricate
structural components such as the internal elastic lamina,
viability (as measured with the MTT assay and TUNEL analysis), and
function (vasoactivity, mechanical and biomechanical properties).
Thus, the resulting vascular grafts behave in a manner similar to
native arteries in terms of mechanical integrity, and provide
clinically relevant patency rates when implanted in vivo. Moreover,
the ex vivo methods and system permit the precise control of the
mechanical environment involving the excised vessel, while at the
same time permitting carefully monitoring of the resulting
growth/remodeling, thereby opening new avenues of research
regarding the mechanical stimuli responsible for specific aspects
of remodeling in vivo.
Inventors: |
Clerin; Valerie;
(Philadelphia, PA) ; Gusic; Rebecca;
(Philadelphia, PA) ; Gooch; Keith; (Media,
PA) |
Correspondence
Address: |
DRINKER BIDDLE & REATH;ATTN: INTELLECTUAL PROPERTY GROUP
ONE LOGAN SQUARE
18TH AND CHERRY STREETS
PHILADELPHIA
PA
19103-6996
US
|
Assignee: |
The Trustees of the University of
Pennsylvania
Philadelphia
PA
|
Family ID: |
26861410 |
Appl. No.: |
11/273888 |
Filed: |
November 14, 2005 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
10165461 |
Jun 7, 2002 |
7011623 |
|
|
11273888 |
Nov 14, 2005 |
|
|
|
60297203 |
Jun 8, 2001 |
|
|
|
Current U.S.
Class: |
600/36 ;
623/916 |
Current CPC
Class: |
A61L 27/3695 20130101;
A61F 2/062 20130101; A61L 27/3691 20130101; A61L 27/3625 20130101;
Y10S 623/916 20130101; C12N 5/0691 20130101 |
Class at
Publication: |
600/036 ;
623/916 |
International
Class: |
A61F 2/04 20060101
A61F002/04 |
Goverment Interests
GOVERNMENT INTERESTS
[0002] This invention was supported in part by the National
Institutes of Health Grant No. R01 HL64388-01A1. The Government may
have certain rights in this invention.
Claims
1. A method of physically remodeling a small blood vessel, while
maintaining the viability of the vessel, comprising the steps of:
excising the blood vessel from its native site, and subjecting the
excised vessel to a controlled, ex vivo mechanical environment for
a time sufficient to remodel the vessel by increasing the diameter,
length, or wall thickness of the vessel, or any combination
thereof.
2. The method of claim 1, wherein the excised vessel is a small
artery or a vein.
3. The method of claim 1, further comprising applying pressure,
shear, and strain to the vessel under controlled conditions within
the mechanical environment, wherein transmural pressure drop
regulates wall thickness, longitudinal tension regulates length,
and flow-induced shear stress regulates inner diameter of the
remodeled vessel.
4. The method of claim 3, wherein the mechanical environment is
controlled by an ex vivo perfusion system.
5. The method of claim 1, further comprising using the remodeled
vessel as an arterial graft in vivo.
6. The method of claim 1, wherein length of the remodeled vessel is
increased at least 100% over its native length when excised, and
wherein more than 50% of the increased length is retained after
recoil when the remodeled vessel is removed from the controlled
mechanical environment.
7. A method of physically remodeling a small blood vessel to be
used in vivo as a vessel graft in a patient in need of such a
graft, comprising the steps of: excising the blood vessel from its
native site; and subjecting the excised vessel to a controlled, ex
vivo mechanical environment for a time sufficient to increase
diameter, length, or wall thickness of the vessel, or any
combination thereof; removing the remodeled vessel from the ex vivo
mechanical environment; and surgically inserting the remodeled
vessel in vivo as a vessel graft (artery or vein) into the
patient.
8. The method of claim 7, wherein the excised vessel is a small
artery or a vein.
9. The method of claim 7, wherein the excised vessel is autologous
to the patient.
10. The method of claim 7, further comprising applying pressure,
shear, and strain to the vessel under controlled conditions within
the mechanical environment, wherein transmural pressure drop
regulates wall thickness, longitudinal tension regulates length,
and flow-induced shear stress regulates inner diameter of the
remodeled vessel.
11. The method of claim 7, wherein the mechanical environment is
controlled by an ex vivo perfusion system.
12. The method of claim 7, wherein length of the remodeled vessel
is increased at least 100% over its native length when excised, and
wherein more than 50% of the increased length is retained after
recoil when the remodeled vessel is removed from the controlled
mechanical environment.
13. An ex vivo perfusion system for exposing a viable, excised
blood vessel to precisely controlled flow and pressure regimes,
wherein the system comprises: a pump means, which when activated,
continuously pushes fluid through the system; a housing means,
comprising a medium-filled chamber, within which chamber the
excised vessel is housed, and the excised vessel is cannulated with
two sliding tubes, wherein when activated, the chamber housing the
vessel is perfused with cell culture medium supplemented with serum
and antibiotics, and wherein temperature, pH, pO.sub.2, pCO.sub.2,
and nutrients are maintained at levels sufficient to maintain the
viability of the vessel; a reservoir within which the culture
medium is pooled, having a gas exchange port, which permits gas
exchange within the medium; a controller means to control pressure
within the chamber housing the excised blood vessel; an in-line
probe means to measure and report pressure within the system; a
data measurement means attached to the in-line probe means for
digitizing the measured pressure data; and a computer node attached
to the data measurement means to record, analyze and store the
digital data.
14. The ex vivo perfusion system of claim 13, wherein the system
further comprises: as the pump means, a pulsatile blood pump, which
when activated, continuously pushes fluid through the system; as
the housing means, an enclosed Plexiglas cylinder, which forms the
housing comprising a medium-filled chamber, cannulated on each end,
within which chamber the excised vessel is cannulated with two
sliding stainless-steel tubes, wherein the chamber housing the
vessel is perfused with cell culture medium supplemented with serum
and antibiotics, and wherein temperature, pH, pO.sub.2, pCO.sub.2,
and nutrients are maintained at levels sufficient to maintain the
viability of the vessel; a reservoir within which the culture
medium is pooled, having a gas exchange port, which permits gas
exchange within the medium, before the medium is returned to the
pump for circulation within the system; as a controller, a needle
valve controller at either end of the chamber to control pressure
within the chamber housing the excised blood vessel; as an in-line
probe, at least one in-line probe to measure pressure within the
system at a rate of approximately 250 times per second, wherein the
data is reported in analog; as a data measurement means, a data
measurement module attached to the in-line probe(s) for digitizing
the analog pressure.
15. The system of claim 13, wherein the excised vessel is a small
artery or a vein.
16. The system of claim 13, comprising a single excised blood
vessel.
17. The system of claim 13, comprising multiple excised blood
vessels run in parallel, each vessel contained within its own
housing, corresponding chambers and needle valves.
18. The system of claim 13, wherein ports on the Plexiglas cylinder
allow the exchange of medium and nutrients, fluid overflow and
air/CO.sub.2 discharge.
19. The system of claim 13, wherein improved control of the
mechanical environment provides localized intravascular and
extravascular pressure measurement and control, which provides real
time monitoring of vessel remodeling.
20. The system of claim 14, wherein the two sliding stainless-steel
tubes slide independently of the rest of the unit to control vessel
strain.
Description
REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation of application Ser. No.
10/165,461, filed Jun. 7, 2002, herein incorporated by reference in
its entirety, which claims priority to 60/297,203, filed Jun. 8,
2001, herein incorporated by reference in its entirety.
FIELD OF THE INVENTION
[0003] This invention relates generally to the field of tissue
remodeling, specifically the ex vivo remodeling of blood vessels
for use as vascular grafts.
BACKGROUND OF THE INVENTION
[0004] More than a century ago, based on observations of the
microvasculature, Thoma proposed that longitudinal tension controls
vessel length. Since then, a number of studies have shown that
blood vessels can remodel either physiologically or pathologically
when exposed to altered mechanical environments. Arteries exposed
to elevated flow (such as arteries upstream of arteriovenous
fistulas (Holman, Surgery 26:889-917 (1949); Shenk et al., Surg.
Gynecol. Obstet. 110:44-50 (1960)), collateral arteries carrying
flow around an obstruction (Mulvihill, et al., N. Engl. J. Med.
104:1032 (1931)), and aortorenal bypass grafts (Stanley, et al.,
Surgery 74:931 (1973)) remodel (autoregulate) to increase their
luminal diameter in response to increased flow as the result of
vascular smooth muscle cell relaxation. In contrast, arteries
experiencing reduced flow decrease luminal diameter.
[0005] Animal studies substantiate these clinical observations and
suggest that vessels remodel so as to restore the wall shear stress
to initial levels (Fung et al., J. Appl. Physiol. 70(6):2455-2470
(1991); Kamiya et al., Am. J. Physiol. 239(1):H14-21 (1980); Zarins
J. Vasc. Surg. 5(3):413-420 (1987)). Inflation of a tissue expander
implanted within a rat hind limb over different periods of time
ranging from 2 to 21 days increased the length of adjacent blood
vessels 83.+-.43%. Relatively slow expansion-induced lengthening
(.ltoreq.10% per day) did not diminish vessel patency, though more
rapid expansion did substantially reduced patency (Stark, Plastic
and Reconstructive Surgery, 30(4):570-578 (1986)).
[0006] However, the complex interdependence between components of
the mechanical environment (e.g., pressure, shear, and strain) in
vivo has hindered the identification of the specific mechanical
stimuli responsible for remodeling. For example, by altering the
viscosity of the perfusing medium, Melkumyants and coworkers have
reported that by decoupling the effects associated with shear rate,
.differential.v.sub.z/.differential.r, (e.g., convection-enhanced
transport and streaming potentials) and the wall shear stress,
-.mu..differential.v.sub.z/.differential.r, that acute
autoregulation is a response to wall shear stress, not to flow rate
per se (Melkumyants et al., Cardiovasc Res. 24(2): 165-168 (1990)).
Several widely used systems that expose cultured endothelial and
smooth muscle cells to well-defined mechanical environments exist,
but extrapolating results from cell culture models to vascular
remodeling has proven to be problematic.
[0007] Traditional organ culture models employing excised vessels,
such as human saphenous veins under static conditions, provide a
well-defined chemical/biochemical environment and have been used to
study the effects of pre-existing intimal hyperplasia, surgical
preparations (Soyomo et al., Cardiovasc. Res. 27(11):1961-1967
(1993)), and specific biochemical factors, including bFGF (Soyomo
et al., 1993) and ET-1 (Porter et al., J. Vasc. Surg. 28(4):695-701
(1998); Masood, et al. Brit. J. Surg. 84(4):499-503 (1997) on
intimal hyperplasia. The inadequacy of these models is evidenced by
the fact that vessels maintained under static conditions, even in
the absence of known biochemical atherogenic stimuli, rapidly
undergo pathological remodeling, including substantial intimal
hyperplasia (Soyomo et al., 1993).
[0008] The atherogenic nature of traditional organ culture models
appears to be at least partially due to the absence of
physiologically relevant levels of mechanical forces. Porter and
coworkers developed a crude, first-generation flow system by
cutting an excised saphenous vein longitudinally and gluing the
adventitial surface of the vein to the inside of a perfused Tygon
tube (Porter et al., Cardiovasc. Res. 31(4):607-614 (1996)). The
application of venous levels of pressure and flow-induced shear
stress to excised human saphenous veins partially attenuated
intimal hyperplasia associated with traditional organ culture,
while arterial levels of pressure and shear stress completely
abolished intimal hyperplasia (Porter et al., 1996). These results
showed that, with a mechanically active environment, it was
possible to maintain blood vessels in organ culture for weeks
without pathological changes.
[0009] While the mechanical environments used in these studies were
intended to mimic aspects of the arterial or venous circulation,
they lacked many relevant mechanical features, including temporal
variations, cyclic strains, as well as pressure drops across the
vessel wall and the resulting transmural flow--each of which is a
potentially important mechanical stimulus to blood vessels as
summarized in reviews by, e.g., Gooch et al., Mechanical Forces:
Their Effects on Cells and Tissues, Berlin, Springer, 182 (1997),
and by Liu, Crit. Rev. Biomed. Eng. 27(1-2):75-148 (1999).
Perfusion systems have been developed and used to provide a
sophisticated mechanical environment by introducing pulsatile flow,
cyclic flexure (Vorp et al., Ann. Biomed. Eng. 27(3):366-371
(1999)) and transmural pressure (Chesler et al., Am. J. Physiol.
277(5 Pt 2):H2002-2009 (1999)). These have been used to study the
effects of the mechanical environment on gene expression (Vorp et
al, 1999), endothelial cytoskeleton (Herman et al., J. Cell Biol.
105(1):291-302 (1987), lipid transport across the endothelium
(Herman et al, 1987), and vasomotor responses (Labadie et al., Am.
J. Physiol. 270(2 Pt 2):H760-768 (1996)).
[0010] Perfusion systems have also been used to investigate the
effect of hydrodynamic forces on endothelial cells, with specific
focus on the mechanisms by which endothelial cells perceive a
mechanical stimulus and convert it to the initial biochemical
response (i.e., mechanotransduction) (Gooch et al., Am. J. Physiol.
270(2 Pt 1):C546-51 (1996)), as well as the effect of biochemical
pathways stimulated by fluid flow and mechanical forces on cellular
proliferation (Gooch et al., J. Cell Physiol. 171(3):252-258
(1997); Gooch et al., Mechanical Forces: Their Effects on Cells and
Tissues, 1997)) and susceptibility to viral infection. In addition,
the effect of a hydrodynamic environment on the development of
tissue-engineered cartilage has been investigated (Gooch, K., et
al., "Mechanical Forces and Growth Factors," in Frontiers in Tissue
Engineering, (C. Patrick, A. Mikos, and L. McIntire, editors.)
Pergamon, N.Y. p. 61-82 (1998)).
[0011] Vessel cultures have also been used to explore the molecular
biology of vascular remodeling, both under static (Porter et al.,
1998; Masood et al., 1997; Porter et al., Brit. J. Surg.
85(10):1373-1377 (1998); Porter et al., Eur. J. Vasc. Endovasc.
Surg. 17(5):404-412 (1999)), and mechanically active environments
(Chesler et al., 1999; Meng et al., 1999). One area in which the ex
vivo vessel models have been particularly insightful is mechanical
regulation of matrix metalloproteinases (MMPs), expression and
activity (Vorp et al, 1999; Chesler et al., 1999; Meng et al., Exp.
Mol. Pathol. 66(3):227-237 (1999); Mavromatis et al., Arterioscler.
Thromb. Vasc. Biol. 20(8): 1889-1895 (2000)), and the role of MMPs
in vascular remodeling (Porter et al., 1998; Porter et al., 1999;
Loftus et al., Ann. N Y Acad. Sci. 878:547-50 (1999)).
[0012] Tenascin-C (TN-C) is large (>1000 kDa), disulfide-linked,
hexameric extracellular matrix (ECM) glycoprotein that is
prominently expressed during embryonic development,
epithelial-mesenchymal interactions, wound healing, cancer, and
notably, vascular disease (Mackie, Int. J. Biochem. Cell Biol.
29(10): 1133-1137 (1997)), and is also subject to mechanical
regulation. TN-C expression has been shown to be increased in rats
and children suffering from pulmonary hypertension (Jones et al.,
J. Cell Sci. 1 12(Pt 4):435-445 (1999)), and under increased
mechanical loading regimes, TN-C expression co-localizes with
neointimal lesions expressing epidermal growth factor (EGF) and
proliferating cell nuclear antigen (PCNA) (Jones et al., J. Cell
Biol. 139(1):279-293 (1997); Jones et al., Circ. Res.
79(6):1131-1142 (1996)). The pro-proliferative role of TN-C is
supported by in vitro studies that show TN-C acts as a survival
factor for cultured smooth muscle cells (Cowan et al., Circ. Res.
84(10):1223-1233 (1999)). The majority of studies show that
soluble, extracellular, and matrix factors regulate TN-C at the
transcriptional level (Chiquet-Ehrismann et al., Bioessays
17(10):873-878 (1995)). In addition, targeted suppression of TN-C
arrests progressive pulmonary hypertrophy in organ culture (Cowan
et al., 1999). Taken together, these data strongly suggest that in
the vessel wall the expression of TN-C is regulated by the
mechanical environment, and the expression of this protein in turn
is a key regulator of SMC proliferation and vascular
remodeling.
[0013] Nevertheless, there is a sizable unmet demand for effective
small-diameter vascular prostheses for use in coronary bypass
surgery. Currently, the best replacements for occluded arteries are
autologous arteries, which have a cumulative patency rate of 93%
after 5 years (Lytle et al., J. Thorac. Cardiovasc. Surg.
89(2):248-258 (1985)). However, the number of expendable autologous
arteries of appropriate dimensions for bypass grafts is severely
limited, although there are numerous expendable arteries of smaller
dimensions.
[0014] In animal studies where autologous tissue-engineered
small-diameter vessels were evaluated in vivo, they performed much
worse than an autologous vein would have (e.g., about half of the
tissue-engineered vessels had decreased perfusion or loss of
patency within 1 month (Niklason et al., Science 284(5413):489-493
(1999); Campbell et al., Circ. Res. 85(12): 1173-1178 (1999)).
Donor veins of appropriate dimensions are more readily available
and are frequently used, but they have a substantially lower
patency. Human saphenous vein grafts have a patency of .about.90%
at early time points, and 81 % after 1 year (Fitzgibbon et al., J.
Am. Coll. Cardiol. 28(3):616-626 (1996)), but this has been
reported to diminish to 45% after 5 years (Lytle et al., 1985).
[0015] Thus, the limited availability of suitable autologous
arteries, coupled with the poor long-term patency of autologous
veins, has led researchers to explore a number of approaches to
create small-diameter vascular prostheses. These include using
natural (Sandusky et al., J. Surg. Res. 58(4):415-420 (1995)) and
synthetic polymeric materials (Smith et al., J. Med. Chem.
39(5):1148-1156 (1996); Uretzky et al., J. Thorac. Cardiovasc.
Surg. 100(5):769-776 (1990)), pre-endothelializing existing types
of polymer grafts in vitro (Stansby et al., Cardiovasc. Surg.
2(5):543-548 (1994); Stansby et al., Brit. J. Surg. 81(9):1286-1289
(1994)), and creating bioartificial or tissue-engineered blood
vessels from cells and various support structures (Weinberg et al.,
Science 231:397-400 (1986); L'Heureux et al., J. Vasc. Surg.
17(3):499-509 (1993); L'Heureux et al., FASEB J. 12(1):47-56
(1998); Tranquillo et al. Biomaterials 17(3):349-357 (1996);
Niklason et al., 1999); Shinoka et al., J. Thorac. Cardiovasc.
Surg. 1 15(3):536-545 (1998)). While there are a number of
different approaches to generating autologous tissue-engineered
vessels in vitro, they all follow the same general paradigm:
isolate specific cell types from blood vessels, expand these cells
in vitro, and reassemble these cells into a tissue-engineered blood
vessel--with the last step being the major challenge.
[0016] Many of these approaches yielded tissue-engineered arteries
that grossly resemble native vessels, but in animal studies where
tissue-engineered vessels generated in vitro were evaluated in
vivo, their performance was inferior to that of autologous veins
(Niklason et al., 1999; Campbell et al., 1999; Fitzgibbon et al.,
1996). However, it was generally found that the performance of
autologous blood vessels (whole vessels) was clearly superior to
that of tissue-engineered blood vessels (prepared from only cells
derived from the vessels).
[0017] There remains, however, a need in the art for a method or
system by which a blood vessel can be harvested and used to direct
the growth of an intact vessel ex vivo, wherein the newly formed
vessel would be of sufficient size to permit the formation of a
tissue-engineered vessel, which would be suitable for use as an
arterial graft in vivo. Criteria for assessing the remodeled
arteries relate both to the extent that the vessels grow ex vivo,
and the degree that the remodeled arteries resemble healthy
arteries of corresponding dimensions. Even modest increases in
vessel dimensions would be potentially useful. Based on rough
estimates using Poiseuille's law (i.e., vessels deform
iso-volumetrically (Milnor, Hemodynamics, 2.sup.nd, 1989)),
increasing the internal arterial diameter by 33% will increase the
ability of that artery to carry blood by more than 200%.
Poiseulle's law, Q = - .pi. .times. .times. .DELTA. .times. .times.
P 8 .times. .times. .mu. .times. .times. L r 4 , ##EQU1## relates
the volumetric flow rate, Q, to the radius of a straight
cylindrical tube of radius, r. Increasing the radius from 100% X to
133% X, increases flow from Y to 3.1 Y, a greater than 200%
increase in flow.
[0018] In addition, in light of the foregoing and because blood
vessels in vivo actively remodel (i.e., change size and/or
composition) in response to chronic changes in the mechanical
environment, the utilization of this ability of intact blood
vessels to remodel supports the use of the system and methods of
the present invention as a more effective and alternate approach to
generating tissue-engineered blood vessels.
SUMMARY OF THE INVENTION
[0019] The present invention provides a system and method by which
a small blood vessel is harvested with minimal morbidity of the
donor, and the diameter, length, and wall thickness of the excised
vessel are increased by subjecting the vessel to the appropriate
mechanical environment ex vivo over time. Thus, a tissue-engineered
vessel is produced, which is suitable for use as a blood vessel
graft in vivo.
[0020] Clinical observations and animal studies indicate that
vessels remodel in response to altered mechanical environments, but
the complex interdependence between components of the mechanical
environment (e.g., pressure, shear, and strain) in vivo has
hindered the identification of the specific mechanical stimuli
responsible for specific aspects of remodeling. To identify the
mechanical stimuli responsible for vascular remodeling, an ex vivo
perfusion system is provided for exposing viable, excised blood
vessels, cells and tissues to precisely controlled flow and
pressure regimes, while maintaining the viability of the vessel.
The excised vessels are housed in a medium-filled chamber,
cannulated on each end, and perfused with cell culture medium
supplemented with serum and antibiotics (FIG. 1). As in traditional
cell or organ culture systems, temperature, pH, pO.sub.2,
pCO.sub.2, and nutrient composition are regulated.
[0021] In addition, the system allows for the control of several
key aspects of the mechanical environment. It is an advantage of
this system over the prior art to offer improved control of the
mechanical environment and allow real time monitoring of vessel
remodeling. An understanding of which mechanical stimuli control
vascular remodeling is utilized to rationally direct the remodeling
of vessels ex vivo.
[0022] In a preferred embodiment of the invention the ex vivo
perfusion system was used to determine which aspects of the
mechanical environment direct the remodeling of arteries. Moreover,
experimental data have demonstrated the ability to control and
measure extravascular pressure in accordance with the provided
methods.
[0023] Furthermore, based on preliminary data, it appears
reasonable to expect increases of vessel length of at least 100%.
Unless the length and internal diameter of a vessel are greatly
increased, little to no medial thickening maybe required. For
example, rough estimates based on Laplace's Law suggest that if the
internal diameter of the vessel increase 33%, and thickness of the
vessel remains constant, the stresses in the arterial wall will
increase a proportional 33%, which is a relatively small increase.
This assumes that the wall thickness is small compared to the
vessel diameter, and that the average stress across the wall
thickness can be estimated with Laplace's equation, which states
that the transmural pressure, p.sub.i-p.sub.o=T(l/r). Therefore,
the hoop stress, T/h, is directly proportional to the radius.
[0024] Using carotid arteries as a model system, data is presented
demonstrating that mechanically induced, directed remodeling of
excised arteries is possible and that the structure and function of
the resulting arteries remain comparable to native arteries. At the
conclusion of each experiment, vessels were harvested and processed
for histology. From histological sections and subsequent
immunostaining, indices of vascular remodeling (intimal, medial,
and adventitial thickness) and injury (proliferation of,
extracellular matrix synthesis by, and phenotypic change of
vascular smooth cells, disruption of internal elastic lamina,
formation of a neointima, loss of endothelium) were quantified. By
subjecting excised porcine arteries to well-defined mechanical
environments, it is shown that, in arteries, transmural pressure
drop regulates wall thickness, longitudinal tension regulates
length, and flow-induced shear stress regulates inner diameter.
[0025] Thus, it is an object of this invention to provide a system
and method by which small blood vessels, such as arteries, or even
veins, can be harvested with minimal donor site morbidity and
remodeled ex vivo, thereby engineering blood vessels for autologous
small-caliber vascular grafts. In addition, the ex vivo system is
advantageously applied to better understand the molecular biology
of vascular remodeling by facilitating the testing of hypotheses
not amenable to study using in vivo or cell culture models.
[0026] It is a further object of this invention to explore the
extent to which arteries can be elongated ex vivo, to determine the
identity of the mechanical factors that regulate arterial lumenal
diameter and wall thickness, and to explore the extent to which the
lumenal diameter and wall thickness of arteries can be increased ex
vivo. In vivo studies will evaluate the efficacy of arteries
elongated ex vivo as autologous arterial graphs in model subjects,
to provide data for eventual human application.
[0027] It is also an object of the invention to utilize the ex vivo
perfusion system to explore the molecular regulation of
mechanically induced vascular remodeling by characterizing the
expression and regulation of key regulator factors. For example,
the spatial expression and distribution of TN-C mRNA and protein
resulting from various mechanical loads is monitored in the
cultured vessels to determine the region(s) of the TN-C promoter
responsible for mechanosensitivity.
[0028] Additional objects, advantages and novel features of the
invention will be set forth in part in the description, examples
and figures which follow, and in part will become apparent to those
skilled in the art on examination of the following, or may be
learned by practice of the invention.
DESCRIPTION OF THE DRAWINGS
[0029] The foregoing summary, as well as the following detailed
description of the invention, will be better understood when read
in conjunction with the appended drawings. For the purpose of
illustrating the invention, there are shown in the drawings,
certain embodiment(s) which are presently preferred. It should be
understood, however, that the invention is not limited to the
precise arrangements and instrumentalities shown.
[0030] FIGS. 1A and 1B diagrammatically depict the existing ex vivo
perfusion system. FIG. 1A is a schematic diagram of an embodiment
using only one vessel. FIG. 1B shows an enlargement of the chamber
housing the blood vessel 5 from FIG. 1A.
[0031] FIG. 2 graphically illustrates volumetric flow rate (upper
line) and pressure (lower red line) vs. time for an excised porcine
carotid artery exposed to a mechanical environment intended to
simulate its native arterial environment. Before harvesting the
vessel, the average volumetric flow rate was 320 ml/min as measured
using transit-time ultrasound.
[0032] FIG. 3 graphically compares artery length for neonatal
elongation experiments (n=5). Each artery length was normalized by
individual ex vivo unloaded length (i.e., all arteries unloaded
lengths are 1 on day 0); average data are shown with SEM. (**)
indicates p<0.005.
[0033] FIGS. 4A-4D photographically depict the immunohistochemistry
of paraffin sections prepared from porcine carotid arteries before
(FIGS. 4A and 4B) and after (FIGS. 4C and 4D) 10 days of perfusion.
Staining for smooth muscle .alpha. actin (FIGS. 4A and 4C) and for
elastin (FIGS. 4B and 4D) allows identification of the internal
elastic lamina (IEL) and the media, both of which are important
landmarks used to quantify vascular remodeling.
[0034] FIGS. 5A and 5B graphically depict as a function of time,
the effect of controlling extravascular pressure to atmospheric
pressure (FIG. 5A, bottom line, lightest gray), to a fixed amount
above atmospheric pressure (FIG. 5A, top two lines, black and dark
gray), or such that the calculated transmural pressure is a
constant (FIG. 5B, bottom curve). In FIG. 5B, intravascular
pressure is shown in the top curve (black line), and extravascular
pressure is shown in the middle curve (gray line).
[0035] FIGS. 6A-6F photographically depict a microscopic assessment
of arteries cultured ex vivo for 9 days under a mechanically active
environment. In FIG. 6A, hematoxylin and eosin staining reveal
healthy vessel. In FIG. 6B, immunostaining for the smooth muscle
cell specific isoform of .alpha.-actin strongly stains the media,
but not the adventia. In FIG. 6C, immunostaining is specific for
proliferating cell nuclear antigen (PCNA). FIG. 6D is a scanning
electron micrograph of the luminal face of the vessel. In FIG. 6E,
immunostaining of the extracellular matrix protein elastin reveals
the internal elastic lamina and underlying striations. In FIG. 6F,
terminal dUTP nick-end labeling reveals a very low rate of
apoptosis/necrosis.
[0036] FIGS. 7A and 7B graphically depict artery length for
juvenile elongation (n=6, FIG. 7A) and control (n=4, FIG. 7B)
experiments. Each artery length was normalized by individual ex
vivo unloaded length (i.e., all artery unloaded lengths are 1 on
day 0); average data are shown with SEM. Perfused refers to loaded
length of the arteries while in the perfusion system, while the
unloaded length refers to arteries out of the system, under no
load. (**) indicates p<0.005.
[0037] FIGS. 8A-8F provide a comparison of representative
histological sections of fresh and elongated arteries from juvenile
pigs taken at 10.times.. Sections were stained with hematoxylin and
eosin (H & E, FIGS. 8A and 8B), PCNA (FIGS. 8C and 8D), and the
TUNEL assay (FIGS. 8E and 8F). Arrows in FIGS. 8C, 8D, 8E and 8F
indicate approximate lumen location (lumen always faces right).
[0038] FIGS. 9A and 9B graphically depict an assessment of arterial
function of porcine carotid arteries cultured for 9 days ex vivo.
In FIG. 9A, the effect of adding a KCl solution to the medium
bathing the artery is recorded as the time average of the pressure
over 1 second. In FIG. 9B, the data shown represents the mechanical
testing of a strip cut from an artery on an Instron machine.
[0039] FIG. 10 graphically depicts the average longitudinal
stress-strain relationship for fresh (n=9), elongated (n=4), and
control (n=5) arteries from juvenile pigs. Data are shown until the
first point of failure for each group (e.g., the minimum ultimate
strain from the fresh group was 65%). Data were unavailable below
65% for control arteries.
[0040] FIG. 11 is a comparative line graph depicting the effect of
ex vivo remodeling on the longitudinal length of a vessel after 9
days. The length of the vessel shown by line A (the in vivo length)
was equal to that in line C (the initial length in the system).
Line B shows the freshly excised length showing elastic recoil from
the in vivo length. Line D was the length of the vessel after 9
days in the ex vivo system (showing a 100% increase over the in
vivo length. Line E shows the length of the vessel after removal
from the ex vivo system, which even after recoil was 70.+-.3%
greater than the initial length. The length of line D=2.0.times.
that of line C. The length of line E=1.7.times. that of line B.
Conditions B and E are unstressed (i.e., no applied longitudinal
loads); while all other conditions involve longitudinal
loading.
[0041] FIG. 12 graphically depicts outer vessel diameter as a
function of time of a porcine artery exposed to pulsatile flow.
[0042] FIG. 13 graphically depicts pressure harmonics for the
pressure-vs.-time curve shown in FIG. 2. The pressure harmonics are
derived from the coefficients for the Fourier series. Consistent
with the observations of others, the coefficients rapidly approach
zero with increasing harmonic number; and truncating the Fourier
series at n=10 accurately replicates the observed pressure versus
time.
DESCRIPTION OF PREFERRED EMBODIMENTS OF THE INVENTION
[0043] The invention provides a system and method by which
appropriate mechanical environments are applied ex vivo to direct
the remodeling of small, excised blood vessels to create
tissue-engineered vessels characterized by increased length,
internal diameter, and wall thickness. Thus, the small excised
vessels, arteries, or even veins, become tissue-engineered blood
vessels for use in vascular surgery. The invention further provides
an evaluation of the performance of these tissue-engineered blood
vessels in vivo.
[0044] The disclosed ex vivo system allows investigations of the
hypothesis that longitudinal stress or strain induces artery
elongation. In addition, while there are autologous donor arteries
with proper diameter and wall thickness for vascular grafts, they
often are of an insufficient length to meet the required need. For
example, the internal thoracic artery has excellent long-term
patency, but is of an adequate length for only a single bypass
graft. However, recognizing that if the artery could be elongated,
it could be used to bypass multiple occlusions, and the use of
vessels demonstrating inferior performance could be avoided, the
present invention advantageously provides reliable
tissue-engineered blood vessels of sufficient length to meet this
need. In addition, the ex vivo perfusion system is further used to
explore the molecular regulation of mechanically induced vascular
remodeling by characterizing the expression and regulation of key
regulatory factors, for which the spatial expression and
distribution of mRNA and protein are monitored as a result of
various mechanical loads.
[0045] Thus, the invention provides a protocol by which localized
intravascular and extravascular pressures are measured in real
time, and the measured pressures are compared with the calculated
pressure estimates. In a preferred embodiment of the invention the
ex vivo perfusion system was used to determine which aspects of the
mechanical environment direct the remodeling of arteries. Moreover,
experimental data have demonstrated the ability to control and
measure extravascular pressure in accordance with the provided
methods.
The Ex Vivo Perfusion System
[0046] The ex vivo perfusion system (FIG. 1A) was designed and
built with a capability of independent control specific aspects of
the mechanical environment (e.g., the magnitude and time rate of
change in intravascular pressure and flow). Turning to FIG. 1A, a
Harvard Apparatus pulsatile blood pump 1 pushes fluid clockwise
around the circuit. If desired, the initial pulsatile pressure/flow
profiles are dampened by the compliance chamber 2, wherein the
extent of dampening dependent on the volume of gas present.
Pressure in the chamber housing the excised blood vessel 5 is
controlled by needle valves 3, which are up and downstream of the
chamber. Pressure 4 and low 6 are measured 250 times per second
with in-line probes attached to the corresponding Triton
Cardiovascular Measurement Modules (models 200-204 and 200-206) (8
and 9, respectively). Analogue output from these modules is
digitized and sent to a personal computer for analysis and storage
using LabView.
[0047] Medium is pooled in a reservoir 7, which permits gas
exchange, before it is returned to the pump. The system is enclosed
in a 37.degree. C. environment.
[0048] For ease of presentation, only one vessel is shown in the
embodiment presented in FIG. 1A. However, multiple vessels can be
run in parallel, each having its own housing and, when necessary,
corresponding compliance chambers and needle valves.
[0049] FIG. 1B provides an enlarged diagram of the chamber housing
the blood vessel 5 from FIG. 1A. Vessels are cannulated with two
sliding stainless-steel tubes and the entire assembly is inserted
into the Plexiglas cylinder. The stainless steel tubes slide
independently of the rest of the unit to control the vessel strain.
Ports on the Plexiglas cylinder allow the entry and exit of bathing
medium and blood gas mixture.
[0050] The prototype perfusion system, however, was limited in its
ability to control other aspects of the mechanical environment that
may be important for vascular remodeling. Therefore, studies were
undertaken to expand the capabilities of the system for exposing
excised blood vessels to well-defined mechanical environments to
enable a) improved control of the mechanical environment and b)
real-time monitoring of vascular remodeling. The resulting
instrument system has been improved to now enable real-time
measurement of pressure and volumetric flow.
[0051] Elementary instrumentation principles were employed to
interface commercial flow and pressure meters with the PC used for
data analysis and storage. Basic fluid mechanics concepts (e.g.,
those embodied in the Navier-Stoke's equation and Poiseuille
equation) were used to guide the design of the existing system
capable of independent control of pressure. To ensure that the flow
through the vessel is fully developed, the length of straight
constant-diameter tubing immediately upstream of the vessel is
greater than that calculated for fully developed steady and
unsteady flow, 27 cm and 11 cm respectively.
[0052] The inlet length for steady flow, L.sub.s, (the length
required to establish a velocity profile with deviation of less
than 1% from parabolic) is given by the equation L.sub.s=0.16
rNR.sub.R, where N.sub.R is the Reynolds number. This relationship
holds when N.sub.R.gtoreq.50 (Fung, Biomechanics: Circulation, 2nd
ed. (1996), herein incorporated by reference). N.sub.R for this
system is 675, when the Reynolds number is defined as
N.sub.Re=.rho.rU/.mu., where .rho. is the density of the fluid and
.mu. is the dynamic viscosity. The viscosity of medium is
approximately that of water or about 1/5.sup.th of that of blood.
The average velocity, U, is calculated from
U=Q/.pi.r.sub.i.sup.2.
[0053] The unsteady entry length, L.sub.us, is approximated by the
equation, L.sub.us=2.64 U/.omega., where .omega. is the pulse rate
in radians (Fung, 1996). An understanding of the effects of
pulsatility on the velocity profile was aided by considerations of
the Womersley number. For example, for the hydrodynamic conditions
of the porcine carotid artery, the Womersley Number, .alpha., is
1.1. .alpha.=2r(.omega./v).sup.1/2 where .omega. is the pulse rate
in radians, r is internal radius of the vessel, and v is the
kinematic viscosity (Fung, 1996). The relatively low Womersley
number indicates that the transient inertial force is of the same
order as the shear force, suggesting that the pulsatile flow can be
crudely approximated as a parabolic velocity profile.
[0054] The system was validated by running excised vessels
(saphenous vein, jugular vein, and carotid artery) in the perfusion
system for up to .about.10 days. FIG. 2 illustrates representative
flow and pressure recordings from an excised porcine carotid artery
exposed to a mechanical environment intended to simulate its native
arterial environment. Vessels exposed to ex vivo culture were
subsequently characterized with immunohistochemistry performed on
paraffin sections (FIGS. 4A-4D).
Improved System Offers Enhanced Control of Mechanical
Environment.
[0055] Extravascular pressure: In the embodied perfusion system,
the extravascular pressure (i.e., the pressure inside the chamber
housing the excised blood vessel) is maintained at 0 mm Hg gauge
(i.e., ambient atmospheric pressure) by allowing the chamber to
vent to the atmosphere through a 0.22 .mu.m filter. This situation
is an appropriate model of in vivo conditions where extravascular
pressure is roughly atmospheric pressure, but it limits
investigations into the role of pressures in vascular remodeling.
As a result, the perfusion system was modified to allow for control
of extravascular pressure: 1) at a given constant level, or 2) to
provide constant transmural pressure.
[0056] Constant extravascular pressures are provided by attaching
to the chamber housing the excised blood vessel a side arm with a
fixed height of medium exposed to the atmosphere at the top
surface. The extravascular pressure, P.sub.o, is estimated from
hydrostatics as P.sub.o=.rho.gz, where .rho. is the density of the
fluid (1.03 g/cm.sup.3), g is the acceleration due to gravity, and
z is the height of the column of fluid. The validity of this
approach has been demonstrated in preliminary studies, the results
of which are shown in FIGS. 4A-4D.
[0057] An alternative to maintaining a given extravascular
pressure, involves fixing transmural pressure. To do so, the
intravascular and extravascular pressures are first set while under
no flow conditions (e.g., by using fixed heights of medium exposed
to the atmosphere at the top surface). The forces exerted by these
two pressures are balanced by the tension generated in the vessel
wall. To a first approximation, the circumferential tension
generated by the vessel, T, can be calculated from the law of
Laplace for thin walls, T=.DELTA.Pr, as the wall thickness h is
much less than the vessel diameter, r.
[0058] Radial strain: The outer diameter of the vessels is measured
in real time using a laser scanning system (Model # LX2-V10W from
Keyence, Woodcliff Lake, N.Y.) (FIG. 12). The embodied system is
capable of non-invasively measuring over 250 times per second in
vessel diameters up to 1 cm, with a repeatability of 5 .mu.m. Since
the measurement device is external to the perfusion system, with
only the scanning laser beam entering, it is easy to relocate the
device to measure diameters at different points and in different
regions along the vessel (e.g., near the end and at the endpoint).
Radial strain, .epsilon..sub..theta., is calculated as
.epsilon..sub..theta.=(D-Do)/Do, where D is the outer diameter of
the blood vessel at a given time and Do is the initial value.
[0059] Longitudinal strain: The embodied perfusion system allows
for changing the longitudinal strain of the vessel by sliding
(extending) the stainless steel tubes on which the vessel is
mounted (FIG. 1B). Since the length of the vessel is only being
changed very slowly (e.g., in several mm steps once per day), the
length is determined by manually measuring the position of the
stainless steel tubes entering the vessel chamber. Sonomicrometry
is also used to measure the distance between two points on the same
side of the outer surface of the vessel. These points will be
approximately 2 cm apart, and located along the middle portion of
the vessel. From these points, longitudinal strain,
.epsilon..sub.z, is calculated as .epsilon..sub.z=(L-Lo)/Lo, where
L is the length of the vessel at a given time, and Lo is the
initial length. The perfusion system allows for changing the
longitudinal strain of the vessel by sliding the stainless steel
tube on which the vessel is mounted.
[0060] The capabilities of the proposed system, wherein the
mechanical environment is controlled by the proposed perfusion
system device, are summarized in Table 1. TABLE-US-00001 TABLE 1
Key aspects of mechanical environment controlled by perfusion
system. Mechanical parameter Magnitude Time rate of change Flow
rate (Q) 0-500 ml/min Steady to .about.3 Hz cycles Pulse rate 0-200
beats/min Intravascular pressure (P.sub.i) 0-500 mm Hg Steady to
.about.3 Hz cycles Extravascular pressure (P.sub.o) 0-500 mm Hg
Steady or in-phase with intravascular pressure Transmural pressure
(P.sub.i-P.sub.o) 0-500 mm Hg Steady or in-phase with intravascular
pressure Longitudinal strain (.epsilon..sub.z) 0 to 100% Steady to
very slow changes (e.g. 10%/day max.) Radial strain
(.epsilon..sub..theta.) 0 to 10% Steady to .about.2 Hz cycles
[0061] All pressures are gauge pressures.
[0062] The extravascular pressure is controlled to atmospheric
pressure (FIG. 5A, light gray curve, lowest line), to a fixed
amount above atmospheric pressure (FIG. 5A, black and gray curves,
remaining lower lines), or such that the transmural pressure is a
constant (FIG. 5B, bottom curve).
[0063] Next the extravascular compartment is sealed to create a
constant volume extravascular system, thus fixing the vessel
radius. More precisely, .intg. 0 l .times. .pi. .times. .times. r 2
.times. .times. d r , ##EQU2## where l is the length of the vessel,
must be a constant volume at any time. The assumption that the
vessel radius is fixed at all times is one case, but definitely not
the only case that satisfies this integral. The velocity at which
the pressure wave travels along the length of the vessel (typically
several meters per second) is assumed to be rapid compared to the
radial motion of the vessel. However, fixing the radius (and
therefore the transmural pressure) for all times is a solution.
[0064] The validity of the analysis is confirmed by measuring
transmural pressure under the test conditions. Because the pressure
drop across the vessel wall depends on the vessel radius and the
material properties of the vessel wall, to the extent that neither
of these parameters changes, the transmural pressure is regulated.
Acknowledging that these assumptions are not trivial, the validity
of the conclusions were evaluated by measuring the transmural
pressure across ex vivo vessels in real time.
Real-Time Monitoring of Vascular Remodeling
[0065] Geometric Remodeling: In the original perfusion system,
vascular remodeling could only be assessed at the conclusion of the
study when the vessel was fixed and histological sections were
prepared, however by the present invention quantification of
vascular remodeling has been improved by allowing real-time
monitoring of vessel diameter. Although laser-scanning techniques
are used to measure radial strain, inner diameter (i.e., lumenal
diameter) cannot be measured by this method. Therefore,
echo-ultrasound may be used to estimate wall thickness, h. The
internal diameter, D.sub.i, is calculated as D.sub.i=D.sub.o-2
h.
[0066] Biomechanical Remodeling: Two independent procedures are
used to determine the viscoelastic biomechanical properties of the
arteries during ex vivo culture: pressure-wave propagation analysis
and pressure-diameter analysis (Milnor, Hemodynamics, 2nd ed.,
Baltimore, Williams & Wilkins (1989)). The primary benefit of
these two techniques is that they permit continuous real-time
evaluation of the mechanical properties of the vessel wall
throughout the ex vivo culture period.
[0067] "Pressure-wave propagation analysis" compares the pressure
harmonics resulting from Fourier transformation of the
pressure-vs-time profile generated at several points along the
vessel to determine the wave attenuation coefficient, a (which is
related to the viscous nature of the vessel) and the true phase
velocity, c (which can be used to calculate the dynamic elastic
modulus of the vessel). The three-pressure method of wave
propagation analysis, which mathematically removes the effects of
wave reflection (Gessener et al., IEEE Trans. Biomed. Eng. 13:2-10
(1996), herein incorporated by reference), is used, as implemented
by others to assess vascular remodeling in vivo (Wells et al., Am.
J. Physiol. (Heart Cir. Physiol.) 274:H1749-H1760 (1998)).
[0068] Consistent with the observations of others, the coefficients
rapidly approach zero with increasing harmonic number and
truncating the Fourier series at n=10 accurately replicates the
observed pressure versus time data and the use of additional terms
(e.g., n up to 30) did not noticeably improve accuracy (data not
shown).
[0069] P(t) is measured at 3 equally distant points within the
vessel using catheter pressure transducers. A Fourier analysis is
performed on each pressure profile and the resulting harmonics
expressed in complex form (FIG. 13). The coefficients for the
Fourier series are as follows: P .function. ( t ) = a 0 + n = 1
.infin. .times. ( a n .times. cos .function. ( n .times. .times.
.pi. .times. .times. t / T ) + b n .times. sin .function. ( n
.times. .times. .pi. .times. .times. t / T ) ) ##EQU3## were solved
by numerically integrating the integrals a n = 1 T / 2 .times.
.intg. - T / 2 T / 2 .times. P .function. ( t ) .times. cos
.function. ( n .times. .times. .pi. .times. .times. t / T ) .times.
.times. d t ##EQU4## for n 0, 1, 2, . . . and b n = 1 T / 2 .times.
.intg. - T / 2 T / 2 .times. P .function. ( t ) .times. sin
.function. ( n .times. .times. .pi. .times. .times. t / T ) .times.
.times. d t ##EQU5## for n=1, 2, 3 . . . . The pressure harmonic,
.DELTA.P, and the phase angle, .phi., were then obtained using the
relationships M= {square root over (a.sup.2+b.sup.2)} and
.phi.=arctan(b/a). Alternatively P(t) can be expressed as a complex
number in the form P .function. ( t ) = A 0 2 + 1 2 .times. n = 1
.infin. .times. ( A n - j .times. .times. B n ) .times. e j .times.
.times. n .times. .times. .omega. .times. .times. t + 1 2 .times. n
= 1 .infin. .times. ( A n = j .times. .times. B n ) .times. e - j
.times. .times. n .times. .times. .omega. .times. .times. t .
##EQU6##
[0070] Thus, the resulting harmonics are substituted into Bergel's
equation for the true wave propagation coefficient, .gamma. = 1
.DELTA. .times. .times. x .times. cosh - 1 .function. ( P 1 + P 3 2
.times. P 2 ) ##EQU7##
[0071] The true wave propagation coefficient describes the
transmission characteristics of each pressure harmonic as it
travels through an artery. It consists of a real portion, which is
the attenuation coefficient a, and an imaginary portion, which is
the angular frequency divided by the true phase velocity, c (i.e.,
.gamma.=a+(.omega./c). The dynamic elastic modulus, E.sub.dyn, is
related to true phase velocity by the equation E.sub.dyn=3
.rho.r.sub.o/(h(2-h/r.sub.o))c.sup.2 where .rho. is the density of
the cell culture medium, h is the arterial wall thickness, and
r.sub.o is the external diameter.
[0072] In addition, "pressure-diameter transient analysis"
harmonics resulting from Fourier transformation of the pressure and
external radius transients over the pulse cycle is used to
calculate the complex viscoelastic modulus (E*) using the equation
of Bergel (J. Physiol. (Lond) 156:458-469 (1961)), which is herein
incorporated by reference, E * = [ 3 .times. r i 2 .times. r o 2
.times. ( r o 2 - r i 2 ) M .DELTA. .times. .times. r o ] e ( j
.times. .times. .theta. ) ##EQU8## where r.sub.i and r.sub.o are
the internal and external radii, respectively, M is the amplitude
of the pressure harmonic, .DELTA.r.sub.o is the amplitude of the
radius harmonic, .theta. is the phase angle between the
corresponding pressure and radius harmonics, and j is {square root
over (-1)}. As a test for internal consistency, the real component
E* from the pressure-diameter transient analysis is compared to the
dynamic elastic modulus, E.sub.dyn, obtained form the pressure wave
propagation analysis.
[0073] To further assess the accuracy of the real-time measurements
of mechanical properties performed while vessels are in the ex vivo
perfusion system, static and dynamic stress-strain relationships
are measured from axial and longitudinal strips prepared from
select vessels. The static and dynamic stress-strain measurements
are made on, e.g., a fully digital Instron machine (model 5543)
with a positional accuracy of 0.156 .mu.m (FIG. 9B).
[0074] In addition to facilitating the determination of the applied
forces that modulate remodeling (e.g., absolute pressure or
transmural pressure), the ex vivo perfusion system provides insight
into the actual stresses to which the vessels actually respond. By
measuring the acute variations in vessel diameter in response to
cyclic changes in measured transmural pressure, it is possible to
estimate some of the stresses in the vessel wall. Following the
Berceli analysis of the biomechanics of excised arteries (Brant et
al., J. Biomechanics 21(2):107-113 (1988)) and veins (Berceli et
al., J. Biomech. 23(10):985-989 (1990)) (each of which are
incorporated by reference) exposed to various hemodynamic
conditions, the axial stress (T.sub.zz) and hoop stress
(T.sub..theta..theta.) are estimated.
[0075] Each of these parameters can be expressed as functions of
the incremental modulus (E.sub.inc), essentially the real component
of the complex viscoelastic modulus applied over a limited range of
strain. This is calculated as follows: E inc = TP max - TP min r o
, max - r o , min 2 .times. ( 1 - .sigma. 2 ) .times. r i , avg 2
.times. r o , avg r o , avg 2 - r i , avg 2 ; ##EQU9## T zz =
.sigma. .times. .times. E inc ( 1 - .sigma. 2 ) .eta. r i , min ;
##EQU9.2## T .theta. .times. .times. .theta. = E inc ( 1 - .sigma.
2 ) .eta. r i , min + E inc .times. h 2 .times. .eta. 12 .times. r
i , min 3 .function. ( 1 - .sigma. 2 ) ##EQU9.3## where TP is
transmural pressure, r is radius, .mu. is dynamic fluid viscosity,
h is wall thickness, and .eta. is the measured displacement of the
vessel wall, and the subscripts min, max, and avg refer to the
minimal (diastolic) value, the maximal (systolic) value, and
average values, respectively.
[0076] To a very close approximation, Poisson's ratio, .sigma., is
0.5 for blood vessels (i.e., vessels deform iso-volumetrically). In
this case, the vessel wall is considered elastic, axisymmetric,
semi-infinite in length, straight with circular cross-section,
constrained from motion longitudinally and the radial displacement
is small compared to the radius. These calculations provide
estimates of the mechanical stresses in the vessel wall and the
incremental modulus aid in the quantification of remodeling. The
calculated mechanical stresses are correlated with the observed
vascular remodeling.
EXAMPLES
[0077] The invention is further described by example. The examples,
however, are provided for purposes of illustration to those skilled
in the art, and are not intended to be limiting. Moreover, the
examples are not to be construed as limiting the scope of the
appended claims. Thus, the invention should in no way be construed
as being limited to the following examples, but rather, should be
construed to encompass any and all variations which become evident
as a result of the teaching provided herein.
[0078] Although the disclosed experiments were conducted using
porcine vessels as models to allow for the detailed in vivo
evaluation of the tissue-engineered vessels, the findings are
directly applicable to human vascular replacement and provide the
foundation for human tissue studies.
[0079] For all experiments, vessels were harvested from
anesthetized pigs prior to euthanization. Using aseptic technique,
an incision were made, the vessel were separated from surrounding
fascia and connective tissue, and the vessel was excised. The
vessel was briefly washed in buffer and submerged in cell culture
medium until placed in the perfusion system no more than 2 hours
later. Unless stated otherwise, the duration of each experiment was
4 weeks, which has been shown to be an adequate amount of time to
observe substantial vascular remodeling in vivo.
Example 1
The Perfusion System: Control of Mechanical Environment
[0080] The perfusion system consisted of a peristaltic pump,
compliance chamber, artery chamber, and reservoir, all connected
using Tygon laboratory tubing (Formula R-3603, Fisher Scientific,
Pittsburgh, Pa.), ports for injection into or sampling from the
perfusing medium, and pressure transducers (Model MER100, Triton
Technology, Inc., San Diego, Calif.) upstream and downstream from
the artery (FIGS. 1A and 1B). Steady flow was provided by a
Masterflex roller pump (1) (Model 7553-70, Cole-Parmer, Vernon
Hills, Ill.) with Masterflex Tygon LFL pump tubing (Formula
06429-25, Fisher Scientific, Pittsburgh, Pa.). Real-time pressure
data were acquired via an analog-digital board (Model PCI-6023E,
National Instruments, Austin, Tex.) connected to a Triton System 6
Twinpak Chassis (Active Redirection Transit-Time Flow Module, Model
200-206 and Dual Pressure Amplifier Module, Model 200-204, Triton,
San Diego, Calif.). Data were visualized and recorded using a
LabView-based routine (LabView Full Development System, National
Instruments, Austin, Tex.) on a PC. Gas exchange was provided to
both the artery chamber and reservoir via 5% CO.sub.2 bubbling
chambers. The entire system, except for the roller pump, was
maintained in a dark, 37.degree. C. environment. All components
were sterilized using ethylene oxide and assembled under sterile
conditions.
[0081] The pulsatile-flow pump forces medium through the ex vivo
perfusion system with a well-defined volumetric flow rate.
Controlling the compliance and the resistance of the system allows
for a wide range of mechanical environments (with respect to
magnitude and time rate of change) ranging from arterial to venous
conditions as well as supra- and sub-physiological conditions. As
shown in FIG. 2, the measured ex vivo pressure (lower line) and
volumetric flow profiles (upper line) are maintained in a
mechanical environment at values that simulate typical conditions
of a porcine carotid artery in vivo. Typical hemodynamic values for
pigs are a pulse of .about.80 beats per minute and arterial blood
pressure .about.100/60 mm Hg. Before harvesting the vessel, the
average volumetric flow rate was 320 ml/min as measured using
transit-time ultrasound. In addition to replicating in vivo
pressure and flow profile qualitatively, specific quantitative
features were also accurately reproduced (FIG. 2).
[0082] Table 2 summarizes key aspects of the mechanical environment
controlled by the existing perfusion system and the ranges over
which these parameters can be controlled. TABLE-US-00002 TABLE 2
Mechanical parameter Magnitude Time rate of change Flow rate 0-500
ml/min Steady to .about.3 Hz Intravascular pressure 0-500 mm Hg
Steady to .about.3 Hz Pulse rate 0-200 beats/min Pulse pressure
0-500 mm Hg Steady to .about.3 Hz Extravascular pressure 0 mm Hg
Steady Longitudinal strain 0 to 100% Steady to .about.10%/day
[0083] Note that while some of these parameters are independent of
one another (e.g., average intravascular pressure and average flow
can be independently controlled), other parameters are coupled
(temporal variations in pressure and flow are linked). Though it
would be ideal to have independent control of each mechanical
parameter, this is not always feasible. For example, radial strain
is dependent on parameters that can be directly controlled (e.g.,
intravascular and extravascular pressure), as well as other
parameters that cannot be directly controlled (e.g., wall thickness
and mechanical properties of the vessel, such as modulus).
Therefore, given the number of degrees of freedom in the system, it
is not possible to arbitrarily set Pi, Po and .epsilon..sub.o.
Example 2
Determining Which Mechanical Factors Regulate Remodeling of
Arteries
[0084] Artery Harvest, Preparation and Maintenance: Carotid
arteries from neonatal (.about.5-kg) and juvenile (.about.30-kg)
pigs were harvested by cardiothoracic surgeons at the Children's
Hospital of Philadelphia after the animals were euthanized. Carotid
arteries from adult pigs (.about.100-kg) were obtained from freshly
exsanguinated pigs at a local abattoir. Arteries were transported
in ice-cold culture medium (Dulbecco's Modified Eagle's Medium
(DMEM) supplemented with 10% fetal bovine serum, 100 U/mL
penicillin and 100 .mu.g/mL streptomycin, all from Life
Technologies, Inc., Rockville, Md.). Upon arrival, arteries were
prepared within a laminar flow hood using sterile instruments.
[0085] Arteries, measuring 3-6 cm in length, were individually
cleaned of excess adventitial and connective tissue. Sections were
taken for histology, methylthiazol tetrazolium (MTT) assay and, in
some cases, dry weight and/or mechanical testing. Dry weight
measurements were made after at least 8 hours in a Speedvac system
(SC100, Thermo-Savant, Holbrook, N.Y.).
[0086] Arteries were individually installed into the artery chamber
and pressurized with medium to locate leaks. Installation consisted
of cannulating the artery onto ribbed stainless steel rods
(stainless steel 8, 10 or 13 gauge microtubing, McMaster-Carr,
Dayton, N.J.) via silk sutures, where the outer diameter of the rod
roughly matched the inner diameter of the artery. Whole, leak-free
artery segments were installed at the approximated in vivo length
prior to perfusion (initial ex vivo loaded length) unless stated
otherwise. The initial extension ratio (ex vivo loaded to unloaded
length) was determined for each artery from neonatal and juvenile
animals by measuring the length of the artery before and after
harvest (unloaded). For arteries from adult animals where in vivo
length was not measurable, a ratio of 1.5 was used, since the
average ratio from neonatal and juvenile arteries was
1.47.+-.0.03.
[0087] After installation of the artery, the chamber was filled
with .about.200 mL of 37.degree. C. culture medium, completely
submerging the artery. The chamber was then connected to the
perfusion system containing .about.500 mL of 37.degree. C. culture
medium, wherein the desired volumetric flow rate had been
previously established (10-15 mL/min). Steady flow was then
diverted to the artery chamber from the bypass branch.
[0088] Carotid Arteries from Neonatal Pigs: Five carotid arteries
obtained from neonatal pigs were installed in the ex vivo perfusion
system at their physiological loaded length and elongated
1/6.sup.th of the initial loaded length (16.7%/day) on days 2 to 7
of a 9 experiment. Control arteries from neonatal pigs were
cultured under identical conditions at fixed length (n=6, separate
study by (Clerin et al., Ann. Biomed. Eng. 29(suppl):S-145 (2001)).
All arteries were perfused at the approximated in vivo volumetric
flow rate of 50 ml/min.
[0089] For the higher volumetric flow rates (50 mL/min, arteries
from neonatal pigs) the flow rate was increased slowly over a
2-hour period until the desired flow rate was achieved. The flow
rate for neonatal arteries was chosen to approximate the in vivo
flow rate for neonatal carotid arteries, however subsequent studies
determined that subphysiological flow rates were necessary to
abrogate de-endothelialization and massive cell death in neonatal
arteries (Clerin et al., 2001).
[0090] Upon removal from the perfusion system, control arteries
retained no increase in unloaded length, while elongated arteries
retained a 65.2.+-.4.5% increase in unloaded length (FIG. 3). Under
these flow conditions, histological evaluation revealed that both
elongated and control arteries were denuded of their endothelial
cells, had lost most of their cellularity, especially in the inner
medial region, and had high levels of cell death. The average MTT
index was 0.35.+-.0.13 (n=5) demonstrating low viability as
compared to control arteries which measured 0.87.+-.0.26 (n=6,
p=0.06).
[0091] Carotid Arteries from Juvenile Pigs: Because of these
findings, the flow rate for juvenile and adult arteries were both
chosen to be 10-15 mL/min, 5-10% of the approximated in vivo flow
rates for each artery. The artery chamber was maintained at
atmospheric pressure by venting the chamber to the atmosphere via a
0.22-.mu.m filter. The average time from harvest to installation in
the perfusion system was 60-90 minutes for neonatal and juvenile
and 2-3 hours for adult arteries.
[0092] A total of 18 carotid arteries from juvenile pigs were
perfused in the ex vivo perfusion system and either elongated
(n=12, "elongated arteries") or held at physiological loaded length
(n=6, "control arteries") for 9 days (Table 3). All juvenile
arteries were installed at their physiological loaded length and
perfused at a volumetric flow rate of 10-15 ml/min, previously
shown to be within the optimum range for retaining artery viability
(Clerin et al., 2001). TABLE-US-00003 TABLE 3 Summary of ex vivo
culture experiments. Age of Volumetric Donor Flow Rate Length Pigs
(ml/min) Longitudinal Strain Protocol n increase Rupture Neonatal
50 100% increase in physiological loaded 5 5** 0 length in 9 days
(16.7% on days 2 to 7) Juvenile 10-15 Fixed physiological loaded
length, 9 days 6 0 0 15 50% increase in physiological loaded 8 6**
2 length in 9 days (8.3% on days 2 to 7) 15 66% increase in
physiological loaded 1 0 1 length in 7 days (13.2% on days 2 to 6)
15 100% increase in physiological loaded 3 0 3 length in 9 days
(16.7% on days 2 to 7) Adult 10 100% increase in physiological
unloaded 1 0 0 length in 7 days (16.7% on days 2 to 7) 10 100%
increase in physiological unloaded 2 0 0 length in 27 days (5% on
days 4 to 23) 10 100% increase in physiological loaded 1 0 1 length
in 9 days (16.7% on days 2 to 7) 10 100% increase in physiological
loaded 2 0 2 length in 27 days (5% on days 4 to 23) A significant
increase in unloaded length (p < 0.005) is denoted by (**).
[0093] Carotid Arteries from Adult Pigs: All carotid arteries from
adult pigs were subjected to rapid protocols (stretched 16.7% on
days 2 to 7 of a 9 day experiment, n=1) or slow stretching
protocols (stretched 5% on days 4 to 23 of a 27 day experiment,
n=3), ruptured prior to completion on days 6, 15 (n=2), and 19
(Table 3). Rupture was avoided by installing arteries at ex vivo
unloaded length, and elongating 5% of the unloaded length on days 4
to 23 of a 27-day experiment (n=2), or elongating 16.7% of the
unloaded length on days 2 to 7 of a 9-day experiment (n=1, removed
on day 7 due to suture failure).
[0094] None of the arteries retained an unloaded length increase
upon removal from the perfusion system. Arteries installed at
unloaded length all increased their wet weight (41.6.+-.1.7%, n=3)
while those that ruptured showed no clear trend (8.5.+-.13.7%,
n=3). Viability of all arteries, as assessed by MTT index, was
similar to fresh arteries.
[0095] Application of Longitudinal Strain (Elongation Protocol): In
preliminary studies, it was demonstrated that by applying a
longitudinal strain, vessels could be elongated .about.100% over 9
days. Longitudinal strain was applied daily to the artery by manual
displacement of the steel rods. Arteries perfused for 9 days were
held at their physiological length (initial loaded length) on day
1, stretched at a rate of 1/6.sup.th or 1/12.sup.th of the
physiological length per day from day 2 to 7, held at the final
stretched length on day 8, and excised on day 9. Similarly,
arteries perfused for 27 days were held at constant length (initial
loaded or unloaded length) on days 1 to 3, stretched 1/20.sup.th of
the installed length on days 4 to 23, held at the final stretched
length on days 24 to 26, and excised on day 27. Control arteries
were cultured in the perfusion system under identical conditions,
but were held at their physiological length (initial loaded
length).
[0096] The in vivo length of the porcine arteries were noted prior
to excision. The length of the arteries without an applied load
were measured and the vessels were placed in the perfusion system
at their in vivo lengths. Vessels were randomly assigned to two
groups. Vessels in the first group (control) were arteries were
cultured in the perfusion system under identical conditions, but
were held at their physiological length (initial loaded length).
Vessels in the second group were subjected to different
longitudinal strain rates (.about.2 to 20% per day) for the various
indicated durations (1 week to 2 months). At the end of the
experiment, the lengths of the vessels without applied loads were
measured and compared. All unloaded lengths reported were measured
at least 15 minutes after removal from the artery chamber since no
significant change in artery length (>0.1 mm) was seen after
this time.
[0097] Wall thickness: Increased mean intravascular pressure (i.e.,
hypertension) results in remodeling of blood vessels characterized
by increased ratio thickness and alterations in the zero-stress
state of the vessel (Fung et al., 1991). In vivo, hypertension
results in increased transmural pressure that in turn results in
increased radial strain and transmural flow making it difficult to
identify which mechanical stimulus is responsible for the observed
remodeling. Ex vivo, it is possible to independently control
intravascular and extravascular pressure.
[0098] To determine whether intravascular pressure was affecting
remodeling directly, or whether it was acting by its effect on
transmural pressure, porcine carotid arteries were exposed to the
three sets of conditions summarized in Table 4. Vessels in the
control group were exposed to normal arterial pressures; vessels in
the "normal" hypertension group were subjected to elevated
intravascular pressure, but a normal extravascular pressure, as is
normally the case with hypertension. Both the intravascular and
extravascular pressures were increased by an equal amount so that
transmural pressure remains normal for vessels in the "corrected"
hypertension group. The fact that vessels in the normal
hypertension group, but not the other two groups, experience medial
thickening was a confirmation that intravascular pressure affects
remodeling by its effect on transmural pressure, as opposed to
directly. TABLE-US-00004 TABLE 4 A summary of experimental groups.
Number Intravascular Extravascular Transmural Condition of vessels
pressure pressure pressure 1) Control 3 100/60 0 Normal 2) Normal 3
200/160 0 Elevated hypertension 3) "Corrected" 3 200/160 100 Normal
hypertension
[0099] All pressures are in mm of Hg gauge. Flow is pulsatile with
.about.1 Hz cycle and mean volumetric flow of 300 ml/min.
[0100] Internal diameter: To evaluate the hypothesis that chronic
changes in luminal diameter resulting from vascular remodeling are
also dependent on the wall shear stress, chronic studies were
conducted, similar to the acute studies of Melkumyants et al.,
1990. The viscosity of the perfusion medium was varied from 1 to 10
cP by the addition of high molecular weight dextran, a compound
that is not harmful to excised vessels in chronic cultures (Chesler
et al., 1990). Excised porcine carotid arteries were perfused under
the conditions described in Table 4. A summary is presented in
Table 5 of the experimental groups used to investigate the relative
contribution of fluid flow and shear stress on vascular remodeling
leading to increases in internal diameter. TABLE-US-00005 TABLE 5
Number of Flow rate Viscosity Initial shear Condition vessels
(ml/min) (cP) stress (dyn/cm.sup.2) 1) Control 3 300 5 Normal 2)
High flow/ 3 1500 1 Normal normal shear 3) High flow/ 3 1500 5 5x
normal high shear 4) Normal flow/ 3 300 25 5x normal high shear
[0101] Po=100 mm Hg for all conditions.
[0102] In all conditions, the flow is steady. Therefore, flow can
be considered as fully developed laminar flow in a circular conduit
of constant cross section where the wall shear stress,
.tau..sub.rz, is calculated as: .tau..sub.rz=4
Q.mu./.pi.r.sub.i.sup.3
[0103] The internal diameter of the vessels was assessed throughout
the experiment. At the conclusion of the experiment, the arteries
were fixed at a pressure of 100 mm Hg and histological sections are
prepared. The fact that shear stress regulates chronic changes in
lumenal diameter is shown by the finding that groups 1 and 2 have
the same diameter as each other, but they have a smaller diameter
than that which was found in groups 3 and 4.
[0104] Results: The maximum elongation while retaining mechanical
integrity and viability was achieved by stretching 1/12.sup.th of
the physiological loaded length (8.3%) on days 2 to 7 of a 9-day
experiment. Six arteries were successfully lengthened in the
perfusion system 48.1.+-.2.8% from the initial physiological loaded
length (p<0.001). The corresponding increase from initial to
final ex vivo unloaded length upon removal from the system was
20.5.+-.3.3% (p<0.005)(FIG. 7A).
[0105] In contrast, none of the six control arteries perfused at
physiological loaded length for 9 days (n=6) retained a length
increase upon removal from the perfusion system (FIG. 7B).
[0106] The wall thickness of control arteries was significantly
lower than both elongated (p<0.005) and freshly harvested
arteries (p<0.005), whereas the wall thickness of elongated
arteries was similar to freshly harvested specimens (Table 6).
TABLE-US-00006 TABLE 6 Material properties for juvenile arteries.
Freshly Harvested Elongated Control Wall thickness (mm) 0.80 .+-.
0.04 (n = 9) 0.87 .+-. 0.08 (n = 6) 0.48 .+-. 0.02 (n = 5)**,++
Change in wet weight (%) N/A 39.9 .+-. 18.4 (n = 5) 21.5 .+-. 2.0
(n = 2) Dry/wet weight (%) 13.1 .+-. 0.9 (n = 6) 12.2 .+-. 0.4 (n =
4) 14.7 .+-. 2.8 (n = 2) Data are shown .+-. standard error of the
mean (SEM). Significant differences were found between fresh and
control arteries (**), and elongated and control arteries (++),
with p < 0.005.
[0107] As compared to freshly harvested arteries, the wet weight of
elongated arteries increased 39.9.+-.18.4% (n=5, p=0.07), whereas
the wet weight of control arteries increased 21.5.+-.2.0% (n=2,
p=0.06) (Table 6). The dry/wet weight ratio was not significantly
different between fresh, control or elongated arteries (Table
6).
[0108] Three arteries elongated 1/12.sup.th of their physiological
length on days 2 to 7 were removed before day 9. One was removed
from the perfusion system on day 7 (after an elongation of 50%) due
to a slow leak, but showed no other problems and was included in
the analysis.
[0109] More rapid elongation (i.e., >10%/day) always caused
arteries from juvenile animals to rupture (Table 3). Three arteries
ruptured when elongated 1/6.sup.th of the physiological loaded
length daily (16.7%/day) on days 2, 3 and 5 of 9, while one artery
failed on day 3 of 7 when elongated 1/8.sup.th of the physiological
loaded length daily (12.5%/day) on days 2 to 6. None of the
arteries that ruptured retained a permanent unloaded length
increase upon removal from the system.
[0110] While both elongation of the juvenile vessels in the
perfusion system (while under load) and the increased unloaded
length indicate arterial elongation, the more relevant parameter is
the increase in length at physiological longitudinal stress. Noting
that the average physiological longitudinal strain is 50% (bar 1
vs. bar 2 of FIG. 7A), the average physiological longitudinal
stress at 50% strain for fresh arteries is 0.40 MPa (mega
Pascals)(FIG. 8).
[0111] The longitudinal strain of elongated arteries at 0.40 MPa is
72%. Taken together (i.e., the product of the unstressed length and
the longitudinal strain at physiological stress), these data
indicate that arteries from juvenile pigs elongated for 9 days ex
vivo are 40% longer than equivalent fresh arteries at physiological
longitudinal stresses. Interestingly, the stress-strain curve for
elongated arteries was found to closely resemble the curve for
control arteries. Thus, the increase in longitudinal extensibility
at relatively low stress appears to be a result of ex vivo culture,
rather than the increase in the applied longitudinal stress or
strain.
[0112] Several non-exclusive mechanisms may contribute to the
observed elongation of juvenile arteries including plastic
deformation due to the applied longitudinal stress/strain (i.e.,
creep), mechanically-induced, biologically-mediated redistribution
of tissue components (i.e., remodeling without growth), and
mechanically-induced, biologically-mediated deposition of new
tissue components (i.e., growth). While both creep and remodeling
without growth could account for limited lengthening of the
arteries, substantial elongation of arteries without substantially
decreasing wall thickness or inner diameter would require growth as
well. As a result, the 40% increase in wet weight of the arteries
as the result of the 9-day elongation process shows that growth is
occurring during the elongation of the arteries. Associated with
this increase in wet weight, there is a small (1.07-fold) increase
in hydration of the arteries, but the majority of increase in wet
weight is due to the 29% increase in the dry weight of the
elongated arteries.
[0113] The greater ability of juvenile arteries to remodel as
compared to adult arteries is consistent with data from in vivo
studies showing that both adult and juvenile arteries can remodel
in response to changes in their mechanical environment, but that
juvenile respond more readily. Langille et al., Am. J. Physiol.
256:H931-939 (1989); Miyashiro et al., Circ. Res. 81:311-319
(1997)). By comparison, neonatal arteries elongated up to 100%
under load and 65% when unloaded within 9 days, though as reported
by Clerin et al., 2001, even control neonatal arteries had reduced
viability over 9 days in culture with physiological flow rates.
Example 3
Control of Extravascular Pressure
[0114] Since vessels are compliant viscoelastic materials, adequate
control of the extravascular pressure was essential to validate the
accuracy of the estimates for transmural pressure. Accordingly, to
critically contol extravascular pressure (i.e., the pressure of the
medium bathing the external surface of the vessel), pressures were
measured by placing a catheter pressure transducer close to the
external surface of a segment of compliant Penrose tubing used as a
surrogate of an artery for these preliminary studies and compared
measured values to the set point values.
[0115] A goal of this experiment was to reduce both the magnitude
and the variation in the transmural pressure across a compliant
tube (e.g., 30.+-.10 mm Hg controlled, as opposed to 120.+-.25 mm
Hg uncontrolled). As shown in FIG. 3B, these data indicate that
extravascular pressure was accurately controlled to a constant
value, and transmural pressure was controlled with a degree of
success which appeared to be sufficient to evaluate the relative
role of transmural pressure in vascular remodeling.
[0116] These findings are adaptable to studies with arteries
because when measured pressures are found to vary significantly
from predicted values, the measured values were used for subsequent
analyses. Therefore, the degree of control of transmural pressure,
which was obtained when supplemented with measurements of the
intravascular and extravascular pressures, enabled detailed study
of the effects of the relative contribution of absolute pressure,
transmural pressure, and cyclic strain on vascular remodeling, each
of which is an aspect of the mechanical environment affecting
vascular cells and blood vessels.
Example 4
Viability, Structure, and Function of Arteries after Ex Vivo
Culture
[0117] Porcine carotid arteries were harvested and cultured in the
ex vivo perfusion system under mechanical active environments for
up to 9 days as described in Example 2. Vessels were harvested at
select times (time zero, 1 hr and 1, 3, 5, and 9 days) and the
viability, structure, and function of the vessels were assessed
using various criteria summarized in Table 7. The following results
compare freshly harvested porcine carotid arteries to vessels
perfused ex vivo. TABLE-US-00007 TABLE 7 Assay Measure of Major
results/implications Viability MTT Mitochondrial activity Viability
not diminished after 9 days in culture PCNA # of proliferating
cells in histological Increased cell proliferation section
through-out full thickness of vessel wall TUNEL # of cells with
fragmented DNA in Normal levels of apoptosis and histological
section (indicating necrosis apoptosis or necrosis) Structure
H&E Histological/microscopic structure General arterial
structure preserved; No intimal hyperplasia Elastin Stains internal
elastic lamina (IEL) in Internal elastic lamina and histological
sections elastic layer in media intact Smooth Stains smooth muscle
cells in Strong staining in media similar muscle histological
sections to fresh isolated arteries .alpha. actin SEM
Microstructure of luminal surface Endothelium intact, but with some
cells rounded up Function Macroscopic Occlusion, aneurysms, and
Vessels intact and non-occluded Assessment "hemorrhage" Addition of
Vasoactive response of vessel Voltage-gated calcium channels KCl
exposed to contractile stimulus and contractile apparatus indicated
by altered pressure drop functional along vessel Abbreviations:
MTT--3(4,5-dimethylthia-zolyl-2)-2,5-dihenyl tetrazolium bromide;
PCNA--proliferating cell nuclear antigen; TUNEL--terminal pUTP
nick-end labeling; H&E--hematoxylin and eosin; SEM--scanning
electron microscope.
[0118] By all criteria employed so far, ex vivo cultured vessels
have been shown to be nearly identical to freshly harvested
vessels.
[0119] Histology and TUNEL Assay: Ring samples (.about.1 mm in
length) were taken from fresh and cultured arteries for TUNEL
assays and histological evaluation. Samples were fixed overnight in
either 70% ethanol or 10% formalin, dehydrated, embedded in
paraffin, and cut into 5 .mu.m thick sections, which were mounted
onto glass slides. Slides were deparaffinized and stained with
hematoxylin and eosin (H & E), the PC10 antibody recognizing
proliferating cell nuclear antigen (PCNA/HRP, DAKO, Carpinteria,
Calif.), and the in situ Cell Death Detection kit, POD (TUNEL,
Roche Molecular Biochemicals, Indianapolis, Ind.) according to
manufacturer's instructions or common protocols.
[0120] Mitochondrial activity was assessed using the methylthiazol
tetrazolium (MTT) assay (Sigma, St. Louis, Mo.). Artery ring
samples approximately 1-2 mm in length were incubated in 0.5 mL of
1 mg/mL MTT solution for 24 hours at 37.degree. C., rinsed with
0.9% saline solution, cut into 2-5 pieces, placed in covered
containers containing 5 ml of isopropanol and incubated at room
temperature for at least 24 hours. The absorbance of 1 mL of the
liquid was measured at 550 nm and normalized by the dry weight of
the sample. An MTT index was defined as the final normalized MTT
value divided by the initial normalized MTT value. An index value
near 1 indicated that mitochondrial activity was similar for fresh
and cultured specimens.
[0121] Viability index measured by MTT was 0.34.+-.0.03 units/mg
before perfusion and 0.30.+-.0.08 units/mg after 9 days of
perfusion (n=6, p=0.66); the frequency of cells containing
fragmented DNA, as measured by the TUNEL assay, were low in both
sets of vessel indicating little apoptosis or necrosis (FIG. 5F);
and the rate of cell division, indicated by the presence of
proliferating cell nuclear antigen (PCNA), was slightly higher in
the culture arteries than the freshly harvested arteries (FIG. 5C).
The elevated proliferation was not the result of intimal
hyperplasia. The gross macroscopic and microscopic structure of the
arteries was not changed by ex vivo culture (FIGS. 6A and 6D), nor
was the tissue-specific localization of ECM (FIG. 6E) or cells
altered (FIG. 6B). Cultured arteries continued to exhibit
vasoactive responsiveness (FIG. 6A).
[0122] The cellularity, structure, and viability of freshly
harvested, control, and successfully (i.e., not ruptured) elongated
arteries were similar, as assessed in histological sections stained
with H & E (FIGS. 8A and 8B), PCNA (FIGS. 8C and 8D) and TUNEL
(FIGS. 8E and 8F). There was no evidence of intimal hyperplasia or
fragmentation of the internal elastic lamina in any of the
elongated or fresh arteries (FIGS. 8A and 8B). Four of the
elongated arteries and five of the control arteries had good
endothelial coverage, while two elongated and one control artery
were denuded of their endothelial cells. One of the denuded
elongated arteries had been denuded (for unknown reasons) prior to
installation into the perfusion system, and thus had no endothelial
cells after elongation.
[0123] TUNEL staining of fresh, elongated and control samples
revealed minimal cell death (FIGS. 8E and 8F). An exception to this
finding was that the two elongated arteries and one control artery
which were denuded of endothelial cells stained strongly for TUNEL,
consistent with previous findings that denuded arteries experience
progressive cell death beginning in the inner lumen by day 9,
irrespective of elongation procedures (Clerin et al., 2001).
[0124] Test for Vasoactivity: Vasoactivity experiments were
performed on select arteries from juvenile animals upon completion
of elongation protocols. Pressure was measured upstream and
downstream of the artery in real time to yield the pressure drop
caused by the change in arterial inner diameter as endothelial
independent vasoactive agents were added to the artery chamber.
[0125] Porcine carotid arteries were cultured for 9 days ex vivo. A
KCl solution was added to the medium bathing the artery (FIG. 1B)
causing the vessel to contract as indicated by the increased
average pressure upstream (i.e., an increased pressure drop along
the length of the vessel). Though instantaneous pressure
fluctuated, as shown in FIG. 2, the time average of the pressure
over 1 second is displayed in FIG. 9A. The data represents the
response in one vessel, but similar responses were observed in the
other vessels investigated.
[0126] Vasoconstriction of cultured arteries in response to KCl
(FIG. 9A) is a salient observation, because tissue-engineered
vessels generated by the methods of L'Heureux et al. and Niklason
et al. do not respond to this stimuli (Nicholas L'Heureux, personal
communication), indicating that the smooth muscle cells have lost
important aspects of their basic function (e.g., their
voltage-gated calcium channels), perhaps as the result of their
expansion in two-dimensional cell culture prior to their use to
form engineered vessels.
[0127] Immediately prior to removal from the perfusion system on
day 9, three elongated arteries were tested for vasoactive response
to norepinephrine (NE) and sodium nitroprusside (SNP), a NO donor.
NE was added to the artery chamber to cause vasoconstriction. The
arteries were then monitored for roughly 60 minutes, at which time
SNP was added to the artery chamber to cause dilation. The pressure
drop across the arteries was measured over time to monitor the
constriction and dilation caused by these agents.
[0128] All arteries tested contracted in response to NE
(1.times.10.sup.-6 M (n=2) or 1.times.10.sup.-4 M (n=1)), which
caused a decrease in lumen diameter corresponding to a peak
pressure increase of 52.7.+-.30.3 mm Hg in an average time of
14.8.+-.4.1 minutes. Addition of SNP (1.times.10.sup.-4 M, n=3)
caused an increase in lumen diameter corresponding to an average
pressure decrease of 38.6.+-.16.0 mm Hg in an average time of
6.8.+-.0.4 minutes.
[0129] Evaluation of Mechanical Properties: Samples from select
freshly harvested and cultured arteries were evaluated for
mechanical properties as follows. Arterial sections approximately
1-2 cm in length were transported in room temperature medium and
cut into sheets by one longitudinal incision. Throughout all
mechanical evaluation, the specimens were kept at room temperature
and constantly hydrated with calcium-free phosphate buffered saline
(PBS). The thickness of the specimen was measured at 3 locations
using a near frictionless LVDT probe and platform apparatus, after
the probe was allowed to reach an equilibrium value (60
seconds).
[0130] A "dogbone" stamp was used to cut out a representative
sample from the sheet aligned in the longitudinal direction and,
when sufficient tissue was available, the circumferential
direction. The original test section width was measured with
digital calipers. The wide flaps of the samples were wrapped in 400
grit sandpaper and loaded into the tensile testing apparatus
(Instron 5543, Canton, Mass.) via spring-loaded grips.
[0131] Application of the testing protocol and acquisition of test
data were achieved using Instron's Merlin software. The testing
consisted of a slow ramp at 0.1 mm/sec, 10 precycles from 0.10 to
0.15 N, a 2-minute hold at constant length, then strain to failure
at 0.5 mm/sec. Engineering stress (load/initial cross-sectional
area), and engineering strain ((final--initial length)/initial
length) were used to determine the stress-strain relationship.
Ultimate stress and ultimate strain were defined as the stress or
strain at the point when the sample failed.
[0132] The average stress-strain relations for freshly harvested,
elongated, and control arteries are displayed in FIG. 10. Defining
the transition zone as the nonlinear region separating two
approximately linear regions of different slopes, the transition
zone for both the control and elongated arteries ended at about 65%
strain, whereas freshly harvested arteries ended at about 35%
strain. The average ultimate stress and strain in the longitudinal
direction was calculated for fresh, control and elongated arteries;
only the ultimate stress of control arteries was significantly
different than elongated and freshly harvested arteries (Table 8).
The ultimate stress and strain in the circumferential direction
were obtained for some fresh and elongated arteries (Table 8).
TABLE-US-00008 TABLE 8 Mechanical properties for arteries from
juvenile donors. Freshly Cultured harvested elongated Cultured
control arteries arteries arteries Longitudinal (n = 9) (n = 4) (n
= 5) Ultimate stress (MPa) 1.41 .+-. 0.13 1.39 .+-. 0.21 2.11 .+-.
0.10*,## Ultimate strain (%) 94.1 .+-. 7.67 121 .+-. 12.9 115 .+-.
9.53 Circumferential (n = 3) (n = 4) Ultimate stress (MPa) 1.98
.+-. 0.46 0.87 .+-. 0.09 N.A. Ultimate strain (%) 106 .+-. 3.70
89.9 .+-. 23.6 N.A. Data were determined from mechanical testing in
the axial and circumferential directions. Data are shown with the
standard error of the means (SEM). Significance differences were
denoted between fresh and control arteries (*), and elongated and
control arteries (+). One symbol equals p < 0.05, two equals p
< 0.005. N.A. indicates not analyzed.
[0133] Tests for Statistical Significance: In cases when the same
specimen could be tracked (such as the artery length before and
after culture), one-tailed, paired t-tests were used. Otherwise,
one-tailed, two-sample t-tests assuming unequal variance were
utilized, p<0.05 was considered significant. For figures and
tables, one symbol (*) denotes p<0.05, while two symbols (**)
denote p<0.005. The ultimate stress and strain in the
circumferential direction were obtained from fresh and elongated
arteries (Table 8). While the ultimate circumferential stress of
control arteries was 2.3 fold. greater than that of elongated
arteries, the difference was not significant (p=0.07). Taken
together, these data indicate that ex vivo cultured vessels
retained their viability, structure and function.
Example 5
Ex Vivo Vascular Remodeling
[0134] Of the three aspects of arterial remodeling to be
investigated (wall thickness, longitudinal length, and internal
diameter), the least is known from in vivo studies about the
mechanical factors that regulate the longitudinal length of a
vessel. Therefore, to test whether longitudinal stress or strain
stimulates vessels to elongate, and to further validate the ex vivo
perfusion system, four excised porcine carotid arteries were placed
in the ex vivo system and initially stretched to their in vivo
length. Each day the vessels were stretched an additional
.about.10% by sliding the stainless steel tubes shown in FIG. 2A.
After 9 days in culture, the length of the vessel had increased
100%. In contrast, acute stretching of the arteries resulted in
rupture after about 80% strain (FIG. 98B).
[0135] When removed from the system, arteries that had been
stretched for 9 days shortened (elastic recoil similar to what was
observed when vessels were initially excised from in vivo), but the
resulting length was 70.+-.3% greater than the initial length of
the freshly isolated arteries prior to stretching, as summarized in
FIG. 11. The length of the vessel in line A (the in vivo length)
was equal to that in line C (the initial length in the system).
Line B shows the freshly excised length showing elastic recoil from
the in vivo length. Line D was the length of the vessel after 9
days in the ex vivo system (showing a 100% increase over the in
vivo length). Line E shows the length of the vessel after removal
from the ex vivo system, which even after recoil was 70.+-.3%
greater than the initial length. The length in line D=2.0.times.
that of line C. The length in line E=1.7.times. that of line B.
Conditions B and E are unstressed (i.e., no applied longitudinal
loads); while all other conditions involve longitudinal loading
[0136] These data provide evidence that aspects of mechanically
induced vascular remodeling observed in vivo can be reproduced in
the ex vivo perfusion system of the invention.
Example 6
Evaluating Performance of Tissue-Engineered Blood Vessels In
Vivo
[0137] To rigorously evaluate the potential utility of ex vivo
remodeled arteries for bypass surgery, in vivo studies are being
conducted. Ex vivo cultured arteries are implanted as autologous
interposition left carotid artery grafts. The in vivo performance
of these grafts with respect to patency and resistance to intimal
hyperplasia are compared to autologous saphenous vein grafts placed
interpositionally in the right carotid arteries of the same test
pigs, and freshly harvested carotid arteries (i.e., no ex vivo
culture) placed back in their original donor set. This experimental
design allows comparison of the ex vivo remodeled vessels to a
positive control (the freshly harvested carotid artery, which is an
excellent vascular graft material) and a negative control (the
saphenous vein, which is a relatively poor vascular graft
material).
[0138] Several sets of in vivo studies are conducted, wherein the
major difference between the two sets being the conditioning of the
ex vivo remodeled vessels. The first set of studies is designed to
test the hypothesis that ex vivo culture of vessels under
mechanical conditions simulating normal physiological loading will
result in minimal vascular remodeling, and that the patency of
these arteries is approximate that of freshly harvested vessels. In
subsequent sets of experiments, the mechanical environment during
ex vivo culture is modified to direct the remodeling of the excised
vessels.
Example 7
Using the Ex Vivo Perfusion System to Explore the Molecular
Regulation of Mechanically Induced Vascular Remodeling
[0139] To evaluate the expression of Tenascin-C (TN-C) protein and
mRNA in arteries exposed to the different mechanical regimes of the
present invention, segments of arteries cultured ex vivo are
routinely fixed and sectioned to prepare histological sections.
Histological sections are immunostained for TN-C protein following
a procedure similar to the one used to stain for smooth muscle cell
.alpha.-actin and PCNA (FIGS. 6B, 6E).
[0140] The majority of studies show that soluble, extracellular,
and matrix factors regulate TN-C at the transcriptional level,
therefore, in situ hybridization studies with digoxigenin-labeled
TN-C riboprobes are used to ascertain the regulation of TN-C
expression at the mRNA level. If mechanically induced changes in
TN-C protein levels in the arterial wall are regulated on the mRNA
level, the region(s) of the promoter responsible for
mechano-sensitivity are determined using full length and a series
of 11 mutated TN-C promoters linked to a CAT reporter gene. These
constructs have been previously used to determine to the regions of
the TN-C promoter that regulate TN-C transcription by cultured
smooth muscle cells in response to remodeled type-I collagen (Jones
et al., J. Cell Sci. 112(Pt 4):435-445 (1999)).
[0141] The TN-C promoter--CAT reporter plasmids are individually
incorporated into a polylactic acid (PLA) (3 mg PLA/1000 ml
chloroform) to give a final DNA concentration of 14 .mu.g/ml.
DNA-polymer emulsions are applied to the surface of a Dacron mesh,
and then desiccated under a laminar flow hood. Plasmid DNA is
delivered from an adventitial position by wrapping meshes around
isolated arteries prior to their placement in the ex vivo organ
culture system. Jones and others have used this technique to
deliver DNA to the arterial wall in vivo.
[0142] After a period of ex vivo culture exposed to the desired
mechanical environment, the artery is retrieved, and a segment of
the vessel is fixed in paraformaldehyde, sectioned and
immunostained with antibodies that recognize the CAT protein. The
remaining segment of the vessel is analyzed for CAT activity using
established techniques. By coupling immunostaining of histological
sections and quantification of CAT enzyme activity, the spatial
distribution and amount of the reporter protein is determined. By
comparing the CAT expression driven by different promoters, the
salient region(s) for mechanosensitivity are indicated. Special
attention is given to the potential role of a putative shear stress
responsive element (GAGACC) 600 base pairs upstream from the
transcriptional start site.
[0143] In sum, the controlled, ex vivo vascular remodeling system
and method of the present invention has been shown to provide a
clinically significant tool for the tissue engineering of vascular
grafts from small excised vessels, as demonstrated at the physical
and molecular levels, and as optimized in vivo. Consistent with the
principle that tissue-engineered arteries generated by the present
invention more closely resemble the structure and function of
native arteries than arteries constructed from isolated cells,
arteries isolated from juvenile pigs and elongated ex vivo were
nearly identical to native arteries in terms of structure (both
macroscopically and histologically, including endothelial coverage
and intricate structural components such as the internal elastic
lamina), viability (as measured with the MTT assay and TUNEL
analysis), and function (vasoactivity and mechanical properties).
Aside from increased extensibility at low stress, the biomechanical
properties of fresh and elongated arteries, notably the ultimate
longitudinal and circumferential stresses and strains, were not
significantly different from fresh arteries demonstrating that when
the elongated arteries are used as vascular grafts, they are
expected to behave in a manner similar to native arteries in terms
of mechanical integrity, as well as to provide clinically relevant
patency rates when implanted in vivo. Moreover, ex vivo it is
possible to precisely control the mechanical environment while
carefully monitoring the resulting growth/remodeling, thereby
opening new avenues of research regarding the mechanical stimuli
responsible for specific aspects of remodeling in vivo.
[0144] Each and every patent, patent application and publication
that is cited in the foregoing specification is herein incorporated
by reference in its entirety.
[0145] While the foregoing specification has been described with
regard to certain preferred embodiments, and many details have been
set forth for the purpose of illustration, it will be apparent to
those skilled in the art that the invention may be subject to
various modifications and additional embodiments, and that certain
of the details described herein can be varied considerably without
departing from the spirit and scope of the invention. Such
modifications, equivalent variations and additional embodiments are
also intended to fall within the scope of the appended claims.
* * * * *