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TISSUE ENGINEERED HEART VALVES:
The Need, the Challenge, the Future
Kay B. Hsu
BE 553 - Final Paper
Dr. Gooch
April 27, 2001
Table of Contents
Introduction ........................................................................................................................................................ 1
Motivation for Tissue Engineered Heart Valves........................................................................................... 1
Strategies and Present Challenges for Tissue Engineered Heart Valves ................................................... 2
Overview of Recent Developments in Tissue Engineering Heart Valves ................................................ 3
Future Work ....................................................................................................................................................... 9
Conclusion ........................................................................................................................................................ 11
Tables................................................................................................................................................................. 12
Figures ............................................................................................................................................................... 16
Introduction
In 1991, a twelve-year comparative study between mechanical and bioprosthetic heart valves
found that 30-35% of heart valve replacement recipients had prosthesis-related problems within 10
years after surgery [1]. Since these heart valves replacements are still commonly used today, along
with heart valves from homograft, bovine, and porcine sources, the need for a more durable and
reliable heart valve substitute is clear. Tissue engineered heart valves (TEHVs) have the potential to
become ideal heart valve replacements because of their ability to grow and repair within the host,
minimize inflammatory and immunological responses, and limit thromboembolism1. Current efforts
to tissue engineer heart valves can be categorized into two strategies: using degradable polymeric
scaffolds or acellular biomatrices.
This paper summarizes the recent advances and ongoing
challenges for both of these TEHV approaches. Important areas for future research are also
identified as TEHVs move closer towards clinical and commercial realization.
Motivation for Tissue Engineered Heart Valves
The impact of valvular heart disease and the limitations of currently available heart valve
replacement options illustrate the need for TEHVs. According to the American Heart Association,
90,000 heart valve replacement procedures were performed in 1998, and valvular heart disease was
the primary cause of mortality for 18,520 individuals [2]. Heart valve disease occurs when a heart
valve can no longer perform its role as a gatekeeper in the cardiac system, and valvular deformities
can exist in two forms: stenosis or regurgitation [3]. In stenosis, the opening of the valve narrows,
and in regurgitation, the heart valve leaflets opens and closes incompletely, causing backflow of
blood through the valve. Both of these impairments can reduce the ejected stroke volume of the
heart and diminish blood perfusion to the body’s vascular system and organs. By disrupting the
1
The blocking of a blood vessel by a particle that has broken away from a blood clot at its site of formation.
(Merriam-Webster Online Dictionary, www.m-w.com)
1
flow and pressure dynamics of the entire cardiac cycle, valvular disease can ultimately cause
secondary heart failure.
Four types of heart valve replacements are currently used to treat valvular heart disease:
mechanical valves, homografts, xenografts, and bioprosthetic valves.
respective advantages and disadvantages of each class.
Table 1 summarizes the
Additionally, all current heart valve
substitutes in clinical use are unable to grow, repair, or remodel within an individual [4]. This
universal limitation is most detrimental to young patients needing heart valve replacements because
successive surgeries are needed to replace the implanted valves that cannot grow with the child.
In contrast, TEHVs include viable cells and can theoretically adapt to a growing and
changing environment like the native, biological structure [5]. Furthermore, TEHVs constructed
from autologous tissue have the potential to meet the following characteristics of an ideal valve
substitute [6]:






No inflammatory and/or foreign body response
No immunological response
Auto-repair potential and life-long durability
Endothelialized surface  not thromboembolic  no anticoagulation needed
Unlimited supply
Customized manufacturing
Therefore, TEHVs hold significant promise in overcoming the existing drawbacks of present day
heart valve replacements.
Strategies and Present Challenges for Tissue Engineered Heart Valves
Currently, there are two general approaches for tissue engineering heart valves. The first
strategy utilizes degradable polymeric scaffolds molded into heart valve geometries. Cells isolated
from donor tissue are cultured and then seeded onto these scaffolds, resulting in constructs that can
be implanted in vivo. The cells grow, develop, and produce extracellular matrix (ECM) as the
polymer degrades, ultimately leaving a natural tissue heart valve without any synthetic component
2
[7]. Ideally, autologous cells are used to eliminate immunological responses to the tissue engineered
construct and to facilitate the growth and remodeling processes [8]. The cell-polymer interaction is
also critical because the quality and extent of ECM formation determine the overall structure and
mechanical properties of the newly developed tissue structure [6].
Therefore, an appropriate
degradable polymer must be identified that 1) can be molded into the shape of a heart valve, 2)
promotes cellular attachment for ECM formation, 3) has sufficient mechanical strength to withstand
hemodynamic conditions, and 4) degrades at a rate comparable to natural tissue ingrowth.
The second TEHV strategy uses acellular, natural biomatrices. For example, porcine heart
valves are processed to remove their cellular antigens and reduce their immunogenicity [8]. These
constructs are then implanted in vivo and repopulated with host cells. This approach requires
decellularization techniques that do not adversely affect the mechanical properties of the biomatrices
or the reconstitution of the tissue in vivo. Issues involving the stability and resorption of the natural
biomatrices must also be resolved. Supporters of this strategy argue, however, that in contrast to
polymeric scaffold TEHVs, acellular biomatrices retain natural ligands and ECM constituents for
better promotion of cell attachment and endothelialization [8].
Overview of Recent Developments in Tissue Engineering Heart Valves
Many studies have been performed within the last four years to address the ongoing
challenges described above for both TEHV strategies. A summary of recent experiments following
the polymer scaffold TEHV approach is given in Table 2. Shinoka, et al. attempted to identify a
suitable degradable polymer for a fully functional and autologous TEHV.
This group first
successfully constructed a tissue engineered heart valve leaflet from woven and non-woven
polyglycolic acid (PGA) mesh sheets [7]. These researchers also fabricated leaflets from polyglactin
sheets sandwiched between PGA mesh sheets. These materials were too stiff, thick, and non-
3
pliable, however, to form non-stenotic, trileaflet heart valves. Additionally, fibrous PGA meshes
had insufficient strength to withstand in vivo flow conditions [9].
The group then investigated several naturally occurring thermoplastic polymers, including
polyhydroxyalkanoate (PHA), poly-4-hydroxybutyrate (P4HB), and polyhydroxyoctanoate (PHO),
for TEHV scaffold fabrication.
These materials are biocompatible, resorbable, and extremely
flexible, and they have high mechanical strength while inducing a minimal inflammatory response
[6]. The low melting point of these thermoplastics allows them to be readily molded into the
configuration of a trileaflet heart valve, and the salt leaching technique can be used to construct
porous polymer scaffolds that promote cellular ingrowth [10].
These polymers were used either alone or in combination to fabricate different TEHV
scaffolds for several experiments.
For example, one comparative study examined the in vitro
performances of TEHVs constructed from trileaflet-shaped polymer scaffolds of PHA versus P4HB
(Figures 1 and 2) [9]. Compared to cells seeded onto fibrous PGA sheets, these TEHVs had
moderate cellularity and collagen formation. Thus, more complex polymer scaffolds were developed
to combine the favorable cell-polymer interactions of PGA with the processability and strength of
thermoplastics. Specifically, one study fabricated TEHV scaffolds by molding PGA around a
softened PHA tube to form the conduit wall and then attaching leaflets constructed from PGAPHA-PGA sandwiches (Figure 3) [11]. Another approach coated nonwoven PGA meshes with a
thin layer of P4HB by dipping the meshes into a solution of P4HB [4]. After solvent evaporation,
the continuous coating of P4HB physically bonded adjacent PGA fibers (Figure 4). Finally,
TEHVs were constructed from a nonporous PHO/PGA scaffold after this combination was
successfully used in tissue engineered aortic segments. As shown in Figure 5, valve conduit walls
were fabricated from a nonporous PHO film sandwiched between nonwoven layers of PGA. The
leaflets were constructed from porous PHO and sutured to the conduit wall with polydioxanone [6].
4
This construct consisting of four different biomaterials needed modification, however, because of
different biomechanical, biochemical, and degradative properties among the various polymers.
Additionally, the nonporous PHO wall inhibited cell ingrowth, and extensive scar formation was
observed on the outside of the TEHV constructs. Based on these observations, a new thermal
processing technique was developed to replace leaflet suturing and to construct both the conduit
wall and leaflets from porous PHO [5]. The porosity permitted cells to grow into the polymer, form
viable tissue, and initiate polymer degradation.
Similar to the variety of polymers used in TEHV scaffold fabrication, a number of different
cell sources have been used to seed the various polymeric constructs. As indicated in Table 2, most
of the earlier experiments seeded mixed vascular cells extracted from ovine carotid arteries onto
PHA or PHA/PGA scaffolds described previously [10,11]. For the more sophisticated PHA and
PHO combination scaffolds, endothelial cells were first extracted from segments of lamb carotid
arteries, and then the remaining vessel segments were minced and cultured in petri dishes [4,6].
Myofibroblasts or medial cells migrated onto the dishes and were expanded separately. When
sufficient cell numbers were cultured, the myofibroblasts or medial cells were seeded onto the
polymeric scaffolds, and after four days of incubation, endothelial cells were seeded on the inside of
TEHV constructs.
In general, the cells attached to the polymer scaffolds, grew into the pores, and secreted
ECM, but the neotissue was immature, mechanically weak, and lacked organization. Researchers
then hypothesized that the hemodynamic function and performance of TEHVs could be improved
by exposing the developing tissue to physiological signals in vitro. A pulsatile fluid flow bioreactor
was developed to provide physiological pressure and fluid flow conditions developing TEHVs. The
system was designed to fit into a standard incubator and maintain sterile conditions for long-term
experimentation [12]. As shown in Figure 6, the bioreactor consists of two chambers separated by
5
a silicone diaphragm. The bottom chamber (1) is filled with air, and the upper chamber (2) is a dualcompartment fluid chamber. A silicone tube (4) connects the two compartments. A sterilized
TEHV construct is mounted onto a removable silicone tube (5), which is then slipped onto the fixed
silicone tube (4) in the bioreactor. The air chamber is connected to a respirator, and when air is
cyclically pumped into the lower chamber, the silicone diaphragm is periodically displaced, pushing
fluid through the TEHV and into the perfusion chamber (2a).
By adjusting the respirator pump
rate, pulsatile flows can range from 50 - 2,000 mL/min and systemic pressures can vary between 10 240 mmHg.
Using this device, various TEHVs were exposed to increasing levels of pulsatile flow and
pressure in vitro, and the cellular response to these constructs was evaluated using histological and
biochemical techniques and environmental scanning electron microscopy (ESEM). Specifically,
both PHA and PHA/PGA polymer scaffolds were seeded with mixed vascular ovine cells and
placed in the pulsatile flow bioreactor for 1, 4, or 8 days [10,11]. The bioreactor was started at a low
flow condition of 140 mL/min and a systolic pressure of 10 mmHg and increased over time to 350
mL/ min and a systolic pressure of 13 mmHg.
Under these testing conditions, the TEHVs were found to open and close synchronously
with pulsatile flow. From ESEM analysis, cells attached to the scaffold in a nearly confluent
manner, oriented themselves according to the direction of flow, and formed connective tissue within
the pores by day 4 in the bioreactor. Movat staining was performed to determine the composition
of the ECM, and collagen and glycosaminoglycans (GAGs) were detected, but not elastin.
According to DNA and 4-hydroxyproline assays, exposure to flow increased the number of cells and
the amount of collagen formed in the TEHVs compared to constructs cultured in static conditions.
Thus, pre-conditioning TEHVs at superphysiological flow conditions, but subphysiological
pressures appeared to strengthen the constructs. [10,11]
6
Another experiment was performed at more physiologically relevant pressures on TEHVs
fabricated from PGA scaffolds coated with P4HB. These constructs were placed in the bioreactor
for up to 21 days in flow conditions starting from 125 mL/min and a systolic pressure of 30 mmHg
and increasing gradually to 750 mL/min and a systolic pressure of 55 mmHg [4]. These pressures
approached the maximum pressures (56-70 mmHg) sustained by a normal aortic heart valve in the
circumferential direction at the points of leaflet attachment to the conduit wall [13]. In contrast to
fragile and disintegrating TEHVs after 14 days of culturing in static conditions, TEHVs cultured for
a similar time period in the bioreactor were intact, pliable, and competent in closure.
Following this pre-conditioning regimen, TEHVs were implanted into the supravalvular
position of pulmonary arteries in lambs. These TEHVs were autologous constructs because they
were implanted into the same lambs from which the cells were initially harvested [5]. Prior to
implantation, each TEHV passed a high-pressure test, withstanding pressures >150 mmHg for 60
minutes.
In vivo valve function was evaluated via a Doppler echocardiogram.
The TEHVs
functioned for up to 20 weeks without stenosis, thrombus, or aneurysm formation, but moderate
pulmonary regurgitation and inflammation were observed between 16 to 20 weeks following
implantation. Histological analyses revealed a patchy endothelial layer and incomplete polymer
degradation. The mechanical strength of the TEHV conduit walls was also determined using an
Instron mechanical testing device. As shown in Figure 7, the stress-strain curve of a TEHV
resembled that of a native pulmonary artery by 17 weeks [5]. In a similar in vivo study, autologous
TEHVs without bioreactor pre-conditioning were implanted into lambs, and their cell number and
ECM components were evaluated as percentages of a native pulmonary artery (Figure 8). While the
number of cells was less than a native artery, the amounts of collagen, elastin, and
proteglycans/GAGs exceeded their respective counterparts in a native artery.
7
Now switching to the decellularized biomatrices strategy for fabricating TEHVs, recent work
has focused on developing alternatives to the glutaraldehyde decellularization process.
Glutaraldehyde, an agent generally used in leather tanning, has been used since the 1960s to reduce
the immuogenicity of xenogenic tissue sources. Since glutaraldehyde cross-links collagen fibers to
minimize the xenogenic tissue solubility, the mechanical properties of the natural biomatrix can be
altered [14]. This chemical also increases the risk of heart valve calcification, and chemical residues
from the fixation process can invoke an inflammatory response or reduce the viability of cells
repopulating the graft [15].
Alternative decellularization solutions have been developed using Triton X-100, trypsin, and
an unspecified buffer with ribonuclease and deoxyribonuclease enzymes [8, 15, 16].
The
decellularization studies using these alternative processes are summarized in Table 3. Bader et al.
and Steinhoff et al. respectively decellularized porcine aortic valves and ovine pulmonary valves
following similar methods.
Specifically, both experiments removed the native cells from the
biomatrices and repopulated the scaffolds with cultured cells. Additionally, the Steinhoff group
implanted its TEHV constructs into lambs.
decellularization
processes
immunohistochemistry.
using
Both groups evaluated the efficiency of their
scanning
electron
microscopy,
histology,
and
The results suggested that the decellularization processes successfully
removed the majority of the cells while preserving the collagen and ECM organization of the
biomatrices.
Steinhoff et al. also evaluated the in vivo performance of their TEHVs using
echocardiography.
Pulmonary regurgitation was not observed among the tissue-engineered
constructs, although the leaflets did thicken and calcify without an apparent loss of function.
O’Brien et al. used a non-detergent-based decellularization solution when conditioning
porcine aortic valve leaflet, and this group did not repopulate the biomatrices with cells prior to
implantation into female sheep [4]. Nevertheless, this decellularization method also successfully
8
removed the majority of cells from the natural scaffold, and no pulmonary regurgitation,
calcification, or gross abnormality was observed among the TEHVs for up to 180 days following
implantation. Host sheep cells were observed to re-populate the porcine-derived graft, and the
performance of the tissue-engineered grafts was not statistically significant from cryopreserved,
cellularized, allogenic sheep aortic heart valves in a number of comparative tests.
This
decellularization method is, in fact, the basis for the commercially available SynerGraft
manufactured by CryoLife, Inc. [19]. In October 2000, CryoLife received a CE mark to distribute
this new tissue-engineered pulmonary heart valve throughout the European Union, and in February
2001, the company announced the first successful implant of SynerGraft into a 3-year-old male child
in Norway.
The company believes that its product is superior to other TEHVs in development
because it does not require pre-conditioning or pre-seeding like the polymeric scaffold-derived
TEHVs [15].
Future Work
Although both the polymeric scaffold and decellularized biomatrix strategies for tissue
engineering of heart valves have made great strides in recent years, it is still uncertain if one of these
two methods will yield the optimal solution. Additional research areas for both approaches include:
1) Cell source and culture: Most experiments to date have used vascular cells from the
carotid artery and myofibroblasts, but these cell populations are difficult and
impractical to obtain for widespread clinical applications. Other cell sources such as
the peripheral vein, mesenchymal stem cells, and arterial/dermal fibroblasts induced
along appropriate differentiation pathways should be considered for future TEHV
experiments [5].
2) Cell seeding techniques and labeling: High efficiencies are needed for cell seeding,
and reliable, long-term labeling techniques are desirable to confirm whether cells
populating the TEHV constructs are from initial seeding efforts or from ingrowth of
surrounding tissue in vivo.
9
3) Mechanical strength: TEHVs to date have only been implanted in the pulmonary
valve position because of lower pressures on the right side of the heart. As the
mechanical strength of TEHVs is continually improved, studies should be performed
in the aortic valve position on the left side of the heart.
4) Longer-term studies: Mechanical integrity and performance of TEHVs must be
examined over longer periods of time with larger experimental group sample sizes.
5) Proper EMC composition and organization: New tissue formation must be
synthesized in a physiological relevant manner. Improved understanding of growth
factors, metalloproteinases, and their inhibitors can lead to TEHVs with ECM
production and degradation that mimic native heart valves [6].
Other areas of improvement for the polymeric scaffold approach include material scaffold
optimization and determination of in vitro pre-conditioning conditions.
Important selection
parameters for choosing the ideal polymeric scaffold include porosity, thickness, degradation rate,
and shape. Appropriate porosity and scaffold thickness ranges are being identified, but scaffold
remnants of PHO and PHA beyond 24 weeks post-implantation [5,11] suggest that the degradation
rate of such polymers is still too slow. Additionally, the shape of the trileaflet heart valve structure
can be improved to include physiological intricacies. Currently, TEHV structures are perfectly
cylindrical, but the distal end of native aortic valve includes bulges called the sinuses of Valsalva [19].
Integrating such features into the scaffold configuration may improve the dynamic function of the
TEHVs [6]. Regurgitation in polymer-based TEHVs has also been attributed to increasing inner
diameters of implanted valves over time and leaflets that no longer touch at closure [4]. More
sophisticated leaflet designs with crimpled, expandable surfaces may be required for TEHVs to truly
grow with the patient.
Optimal bioreactor parameters, including magnitude, duration, and rates of change of flows
and pressures, must also be determined. These variables should be considered simultaneously with
cell culture medium characteristics, such as viscosity, concentrations, and the addition of growth
10
factors, to stimulate physiological ECM formation and tissue structure. For example, two present
concerns for some polymer-based TEHVs include a lack of elastin in the ECM [4] and a patchy,
incomplete endothelial cell lining [5].
While decellularized biomatrix TEHVs have demonstrated better endothelialization than
polymer based TEHVs in animal studies, clinical trials are needed to confirm whether endothelial
outgrowth onto acellular grafts is more potent in animal models than humans [8]. Additionally,
myofibroblasts are believed to migrate less than endothelial cells, such that unseeded acellular valves
may be deficient of myofibroblasts and muscle cells in humans [8]. Without adequate invasion and
ingrowth of these cells, complete graft integration of decellularized biomatrix TEHVs into the host
may be seriously compromised [16]. Furthermore, these TEHVs need rigorous mechanical testing
to confirm that the alternatives to glutaraldehyde processing remove antigens without impairing the
structural integrity of the biomatrices.
Finally, more research is needed among decellularized
biomatrix TEHVs to prove that retroviruses and other infectious risks are fully removed from these
xenogenic tissue sources [8].
Conclusion
Tissue engineered heart valves clearly have the potential to improve our ability to treat
valvular heart disease and to abolish many of the undesirable characteristics found in current heart
valve replacement options. While additional research and improvements are needed for TEHVs
constructed from both polymeric scaffolds and decellularized biomatrices, clinical use of TEHVs
within the next decade does not require a major stretch of the imagination.
11
Tables
Table 1: Advantages and Disadvantages of Existing Heart Valve Replacement Options
Device Type Example
Advantages
Disadvantages
Mechanical
 Ball-in –cage
 Long-term durability
 Thromboembolism
Prostheses
 Lifelong anticoagulant therapy
 Risk of infection
 Noisy operation
 Turbulent flow
 Catastrophic failure
Homografts
 Cadaver
 Less thromboembolic
 Limited supply
(Allografts)
valves
 Lower risk of immune
 Early failure in young people2
response, infection, or
 Degeneration & calcification
disease transmission
requires replacement in 5-15 years
Xenografts
 Porcine or
 Less thromboembolic
 Progressive degeneration & limited
bovine aortic  Less immunogenicity
durability
valves
after glutaraldehyde
 Early failure in young people2
treatment3
 Possible disease transmission
 Lower risk of infection
 Greater supply
Bioprosthetics  Titanium
 Less thromboembolic
 Catastrophic failure
stents
 Improved durability
 Risk of infection, foreign body
covered with
response
pericardium,
etc.
Compiled from [6,7,9,14,16]
2
3
This observation has been attributed to a more aggressive immune system in young people [14].
Refer to discussion of decellularized biomatrices in “Overview of Recent Developments in Tissue Engineering Heart
Valves” section.
12
Table 2: Summary of Recent Experiments using Polymer Scaffolds for TEHVs
Sodian, et al. (2000) [2]
Sodian, et al. (2000) [13]
Sodian, et al. (1999) [14]
 Mixed population of
fibroblasts, smooth muscle
cells, and endothelial cells
from ovine vascular artery
 Trileaflet scaffolds
constructed from porous
PGA, PHA, and P4HB (via
salt leaching)

Mixed cells from intima,
media, and adventitia of
ovine carotid artery


Trileaflet scaffold
constructed from salt
leached PHA

In vitro or In
vivo Model
 Bioreactor – 100 mL/min for
1 hour

Bioreactor – from 140
mL/min and 10 mmHg
systole to 350 mL/min and
13 mmHg systole
Observation
Points
Control Group
 1 day
 1, 4, and 8 days
 1, 4, and 8 days
 Scaffolds cultured in static
conditions
 Functionality: TEHV
function under sub- and
super-physiological flow
conditions; pliable
constructs
 ESEM: cells grow into
pores, confluent cell layer
by day 4, cells oriented in
direction of flow; control
group = disoriented cells
 ECM content: Movat stain
shows collagen & GAGs,
but no elastin
 Cell number & collagen
content: more cells and
collagen on constructs
exposed to flow and with
prolonged flow exposure
than static conditioning
 Scaffolds cultured in static
conditions
 Functionality: leaflets for all
TEHV open and close
synchronously with fluid
flow
 ESEM: nearly confluent cell
layer; leaflet cells oriented in
flow direction by day 4; cells
on conduit wall form bridges
in pores; control group =
disoriented cells
 ECG: no pulmonary
regurgitation, some
thickening of valve w/out
functional loss
 Cell number & collagen
content: more cells and
collagen on constructs
exposed to flow and with
prolonged flow exposure
than static conditioning
Cell source
Scaffold
Construct
 Scaffolds cultured in static
conditions
Evaluation
 Functionality: leaflets for all
Methods &
TEHV open and close
Results:
synchronously with fluid
flow
ESEM- cell
 ESEM: almost confluent cell
attachment,
layers by day 8 for PHA and
surface
P4HB constructs
morphology
 Mechanical strength:
supraphysiological strength
Tension testing  Cell number and collagen
– mech.
content: more cells on PGA
strength
sheets than on PHA and
P4HB constructs;
DNA & 4significantly more collagen
hydroxyproline
on PGA sheets than PHA &
assays: cell
P4HB
number &
collagen
content
Mixed vascular cells from
adult ovine artery
Trileaflet scaffold with
conduit wall constructed
from bilayer of PGA (outer
layer) and PHA (inner
layer); leaflets from PGAPHA-PGA sandwichese
 Bireoactor -- from 140
mL/min and 10 mmHg
systole to 350 mL/min and
13 mmHg systole
13
Table 2: Summary of Recent Experiments using Polymer Scaffolds for TEHVs (con’t)
Hoerstrup, et al. (2000) [4] Sodian, et al. (2000) [5]
Stock, et al. (2000) [6]
Cell source
Scaffold
Construct
In vitro or In
vivo Model
Observation
Points
Control Group
Evaluation
Method &
Results
ECG –Doppler
echocardiogram
for valve
function
ESEM –
Surface
morphology,
cell attachment
Histology –
H&E staining
Immunohistochemistry
Tension testingmechanical
strength
DNA and 4hydroxyproline
assays – cell
number and
collagen
content
 Endothelial cells extracted
from lamb carotid artery
segments; myofibroblasts
migrate out from minced
segments and then expanded
 Trileaflet scaffolds of PGA
coated w/ thin P4HB layer that
physically bonds PGA fibers;
myoblasts seeded first, then
endos seeded on inside
 Bioreactor pre-conditioning:
125 mL/min and 30 mmHg
systole to 750 mL/min and 55
mHg systole; then implant
autologous TEHV into same
lamb where cells initially
harvested
 In vitro bioreactor: 4, 7, 14, 21
days; in vivo study upto 20 wks
 Scaffolds cultured in static
conditions
 Functionality: valves open &
close synchronously under low
& high pressure; intact, pliable,
competent closure after 14 days
in bioreactor
 ECG: no thrombus, stenosis, or
aneurysm upto 20 wks; mild to
moderate pulmonary
regurgitation between 16-20
weeks
 Gross appearance: leaflets
thicker in earlier weeks; inner
diameter of valve construct
increases at leaflet attachment
 Histology: in vitro – dense outer
layer, maximum tissue
organization by 14 days; in vivo
loose layer on ventricular side
(inflow) & fibrous layer on
arterial (outflow side)
 ESEM: confluent surface by 7
days in vitro; smooth surface on
both sides of TE leaflets in vivo
 ECM content: GAGs & collagen
in vitro; elastin detected in vivo
on ventricular side
 Mechanical strength: decrease
over time until comparable to
native vessel (more pliable)
 Cell number & collagen content:
see Figure 8
 Mixed arterial vascular cells
from ovine carotid;
endothleial cells from jugular
vein
 No pre-conditioning
 Implant autologous TEHVs
into supravalvular position in
pulmonary artery of lambs
 Common carotid artery of Dover
lambs; separate endothelial cells
first, then medial cells migrate out
from minced segments &
expanded
 Trileaflet scaffold with conduit
wall of nonporous PHO film w/
layers of PGA felt on inside and
outside; leaflets of porous PHO
and sutured w/ dioxanone
 No pre-conditioning
 Implant autologous TEHVs into
supravalvular position in
pulmonary artery of lambs
 1,5, 13, 17 weeks
 1, 2, 4, 6, 8, 12, 24 weeks
 Acellular valved polymer
conduit
 Functionality: open & close
synchronously w/ flow; inner
diameter of valve increases
 ECG: mild stenosis & trivial
regurgitation; no thrombus
 Gross appearance: constructs
covered with tissue, porous
PHO still observed, mild
inflammation
 Histology: not confluent
endothelial layer, fibrous
capsule w/ moderate
vascularization
 ECM content: rich in
collagen, GAGs; no elastin
 Surface morphology: smooth
flow surfaces, control not
covered w/ tissue
 Mechanical strength:
supraphysiological strength
decreases over time until
stress-strain curve resembles
native artery (Figure 7)
 Cell number & collagen
content: comparable to native
artery in cell number (7080%); major increase in
collagen formation (46% at 1
week to 116% by 17 wks)
 Acellular valved polymer conduit
 Trileaflet scaffold from salt
leached PHO; first seeded
with arterial vascular cells,
then venous endothelial cells
 ECG: TEHVs function w/
exception of one w/ non-attached
leaflet
 Gross appearance: thickening of
TEHV wall & scar tissue; no
increase in diameter or thrombus
 Histology: laminated fibrous
tissue uniformly covering surface;
moderate polymer degradation
 ECM content: Movat staining
shows collagen, GAGs, but no
elastin; mild foreign body
response; elastin content 215% of
native vessel and PG-GAG
content 180% native vessel by 16
wks
 Polymer degradation: (by gelpermeation chromatography) – no
PGA, but PHO still evident in
conduit & leaflets at 24 weeks
 Cell number & collagen content:
increased over time, exceeded
native levels by 6 months
14
Table 3: Summary of Recent Experiments using Decellularized Biomatrices as TEHVs
Bader et al. (1998)
O’Brien et al. (1999)
Steinhoff et al. (2000)
Implantation Model
 Porcine aortic valve
 Solution 1:
 1% Triton X-100
 0.02% EDTA
 PBS w/out Ca2+, Mg2+
 Solution 2:
 RNase A (20 g/mL)
 DNase (0.2 mg/mL)
 Valves in HBSS at 4C
 Placed in solution 1 at 37C
for 24 hrs under continuous
shaking in 5%CO2/ 95% air
 Immersed in solution 2
 Washed in PBS
 Stored in HBSS, 4C
 Endothelial cells from
discarded segments of
humans saphenous vein for
coronary artery bypass
 None
Control Group
 Allogenic acellular grafts
Observation Points
Analysis & Results
 24 hrs, up to 3 days
 Light microscopy: mostly
cell-free structure after
treatment; loose, wavelike
collagen fibers still in tact
 SEM: porous network of
fibrillar ECM after Triton
treatment
 Immunohistochemistry:
confluent endothelial
monolayer after 3 days; cells
stained CD-31 positive to
mark human endothelial
cells, but not porcine
endothelial cells
Biomatrix
Decellularization
Solution
Decellularization
Process
Re-populating
Cell Source
List of abbreviations:
DNase Deoxyribonuclease, for enzymatic degradation
ECG
Echocardiography
EDTA Ethylenediamine Tetraacetic Acid
 Porcine aortic valve
 Unspecified buffer w/
RNase and DNase




 Cell lysis in sterile
water
 Equilibrated in buffer
w/ RNase and DNase
 Multi-day isotonic
washout period
 Valves in HBSS at 4C
 Placed in decellularization
solution at 37C for 24 hrs
under continuous shaking in
5%CO2/ 95% air
 Washed in PBS
 Stored in HBSS, 4C
 Ovine carotid artery segment
separated into endothelial &
myofibroblast cells
 None
Ovine pulmonary valve
0.05% trypsin
0.02% EDTA
PBS
 Right ventricular outflow  Lamb pulmonary artery after
tract of female sheep
pulmonary leaflets excised
 Cellularized,
 Allogenic acellular grafts
cryopreserved sheep
allograft aortic valve
 80, 150 days
 2, 4, 12 weeks
 Histology: cells
 Post-operative: 1 control
effectively removed from
animal died after 17 days from
leaflets & conduit;
endocarditis and thrombosis
fibrous collagen structure  Gross morphology: normal
remains
valve morphology in 5 of 6
 No gross disorders
TEHVs; some calcification
following implantation
observed
 Acellular porcine grafts
 ECG: no pulmonary
repopulated with host
regurgitation, some thickening
sheep cells
of valve w/out functional loss
 ECG: no difference in
 Histology: complete
mean pulmonary artery
endothelialization of TEHVs
flow velocity between
by week 4
experimental & control
 Immunohistochemistry: valve
groups
interstitium repopulated by
 Angiography: no
myofibroblasts from strong difference in pressure
actin staining; active matrix
gradient across valves
synthesis from pro-collagen I
between experimental &
staining
control groups
 Calcium determination:
no difference in calcium
levels  not calcified
HBSS Hank’s buffered saline solution
PBS
Phosphate buffered saline
RNase Ribonuclease, for enzymatic degradation
SEM
Scanning electron microscopy
15
Figures
Figure 1: Aluminum heart valve cast (left) and crown-like porous tissue
engineered heart valve construct fabricated from polyhydroxyalkanoates (PHA). [20]
Figure 2: Trileaflet tissue engineered heart valve scaffold
fabricated from porous poly-4 hydroxybutyrate (P4HB). [9]
Figure 3: Trileaflet tissue engineering heart valve scaffold. Conduit wall is fabricated from an outer layer of
PGA pressed around an inner layer of PHA. Leaflets are fabricated from PGA-PHA-PGA sandwiches. The
PGA was 1mm thick for the conduit wall and 500 m thick for the leaflets, with 95% porosity, 15 m fiber
diameters, and 70.0 mg/cm3 density). The PHA was 500 m thick for the conduit wall and 250 m thick for
the leaflets. [11]
16
Figure 4: Tissue engineered heart valve fabricated from PGA with thin
coating of P4HB after 14 days in bioreactor. [4]
Figure 5: (A) Frontal view of trileaflet tissue engineered heart valve scaffold measuring 18mm in diameter
and 20mm in length. (B) Schematic of the TEHV: conduit wall fabricated from 240 m nonporous
polyhydroxyocatonate (PHO) film with 1mm nonwoven PGA felts on both inside and outside of PHO layer.
Three leaflets of porous PHO are sutured to the conduit wall with polydioxanone. [6]
17
A
B
Figure 6: (A) Photograph of the pulsatile flow bioreactor. (B) Schematic of the device: the bottom chamber
(1) is filled with air, and the upper chamber (2) is a dual-compartment fluid chamber. A silicone tube (4)
connects the two compartments, and a sterilized TEHV construct is mounted onto a removable silicone tube
(5). The air chamber is connected to a respirator, and when air is cyclically pumped into the lower chamber,
the silicone diaphragm is periodically displaced, pushing fluid through the TEHV into the perfusion chamber
(2a). By adjusting the respirator pump rate, pulsatile flows can range from 50 - 2,000 mL/min and systemic
pressures can vary between 10 - 240 mmHg. [12]
18
A
B
C
D
Figure 7: Biomechanical testing of a tissue engineered heart valve conduit wall using an Instron (model
#5542). (A) Stress-strain curve of an unseeded, porous PHA sample. (B) Stress-strain curve of a tissue
engineered conduit wall after five weeks in vivo. (C) Stress-strain curve of a tissue engineered conduit wall
after 17 weeks in vivo. (D) Stress-strain curve of a native pulmonary artery. Notice the similarity between C
and D. [5]
19
Figure 8: (A) DNA content (quantifying number of cells) of tissue engineered heart valve conduit wall in
percentage of native pulmonary artery. (B) 4-hydroxyproline content (quantifying amount of collagen) in
TEHV conduit wall in percentage of native pulmonary artery. (C) Elastin content of TEHV conduit wall in
percentage of native pulmonary artery. (D) Proteoglycan-glycosaminoglycan content of conduit in percentage
of native pulmonary artery.
20
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22