bio mechanics of human common carotid artery and design

8/6/2019 Bio Mechanics of Human Common Carotid Artery and Design

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Biomechanics of human common carotid artery and designof novel hybrid textile compliant vascular grafts

B. S. Gupta1,* and V. A. Kasyanov1,2

1College of Textiles, Department of Textile Engineering, Chemistry, and Science, North Carolina State University,

Raleigh, N orth Carolina 27695-8301; 2 Laboratory of Biomechanics, Riga Technical University, Riga, LV-1658, Latvia

The mechanical properties and structure of a human com-mon carotid artery were stud ied in order to d evelop criteriafor designing and manu facturing comp liant textile vasculargrafts. The arterial w all comprised a comp osite of elastinand collagen fibers w ith the collagen fibers crimp ed. Thisstructure led to a unique pressurecircumferential stretchratio curve, the slope of which increased with an increase instrain. The increase in slope was particularly rapid at a

stretch ratio above 1.4 or pressure above 120 mmHg. Basedon the knowledge gained, a criteria for the design of biome-chanically compliant arterial grafts was developed. An elas-tomeric prestretched p olyurethane monofilament yarn witha low modulus of elasticity and a bulked polyester multi-

filament yarn with a high modulus of elasticity were com-bined and used as threads in the manufacture of grafts. Tu-bular structures of diameters in the range 46 mm weremad e by weaving. Transverse compliance and morphologi-cal and permeability properties of these grafts were deter-mined and compared with those of a currently availablewoven commercial grafts and hum an carotid arteries. Re-sults indicated that the compliance values of the hybrid

grafts were comparable with those of the human carotidartery. Preliminary in vivo studies in dogs showed promis-ing results: a thin, stable neointima developed within 6months of implantation on the flow surface. 1997 JohnWiley & Sons, Inc.

INTRODUCTION

The design and fabrication of synthetic vascular

grafts has been a challenging area in vascular surgicalresearch during the past 30 years. Large diameter

grafts (diameter 6 m m) u sed for bypassing arteries in

high flow regions such as the thoracic and abdom inal

aorta have generally performed well. However, the

replacement with grafts of small diameter arteries,

such as the coronary, renal, and carotid, has not yet

been successful and continues to be a problem in re-

constructive surgery. Efforts to develop vascular

grafts of diameters less than 6 mm with potential for

long-term patency have not yet met with success. A

major cause for poor performance of such grafts has

been shown to be the lack of comp liance.14

Replacement of small arteries by rigid woven or

knitted prostheses, which have little or no compliance

in the circum ferential direction, causes th e dam pen ing

out of the higher harmonic in the pulse wave5 that

leads to an increase in the pulse wave velocity and

therefore to an increase in wave reflection and energy

loss. The extent to which the pulse amplitude is damp-

ened depends upon the length of the rigid part. For

examp le, in a stud y by Womersly,5 a 15-cm long rigid

section inserted in the femoral artery of a dog showed

a redu ction of the amp litud e by about 13% in the first

and 42% in the fourth harmonic. This problem was

also noted and discussed by How and Clarke6 (1984).

Baird and Abbott7 and Rittgers et al.8 showed that

hemodynamic forces play an important role in the for-

mation of thrombus and hyperp lastic intima. Doo and

colleagues9 determined theoretically and experimen-

tally the differences in the behaviors of an elastic and

rigid tube used as a model for an aortic arch. The

resulting flow distributions examined showed a dif-

ference in the flow behaviors of the rigid and the elas-

tic mod els of the arterial system. The arterial w all elas-ticity had an effect on the blood flow distribution; a

lack of elasticity led to high turbulence. The work by

Stein et al.10 showed that for a given Reynolds nu mber

the intensity of turbulence was significantly lower in

compliant tubes than in rigid ones. In the latter, unfa-

vorable flow conditions led to the formation of anas-

tomotic aneurysm, development of hyperplastic neo-

intima, and failure of sutures or tearing of the host

artery.*To whom correspondence should be addressed.

Journal of Biomedical Materials Research, Vol. 34, 341349 (1997) 1997 John Wiley & Sons, Inc. CCC 0021-9304/ 97/ 030341-09


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Clearly, successful d evelopm ent of a small diameter

vascular graft will depend not only on the use of new

biocomp atible m aterials but also on ideas of n ew con-

structions. Special attention has been devoted in re-

cent years to the use of polyurethanes, which have

good biocomp atibility11,12 and deform ational andstrength p roperties.6,1316 However, problems have

been encountered with the u se of sm all diam eter

grafts made exclusively of this material. The polymer

is found to creep, which leads to the development of

aneurysm. Thus, in spite of the availability of conve-

nient spray technology for manufacturing grafts with

such elastomers, interest in woven and knitted textile

grafts continues because the latter have been success-

ful in m edium and large caliber app lications, and su b-

stantial experience has been gained in th eir design and

construction du ring the past four decades.

1721

The literature makes it clear that mechanical char-

acteristics of vascular grafts play an important role in

governing long-term patency. Specifically, the most

important consideration to be given in d esigning a

graft is to match the m echanical prop erties of the pr os-

thesis with th ose of the host artery. Information exists

in the literature on the mechanical properties of the

arteries of animals, especially the dog.2226 The work

of Hayashi et al. and that of one of the present auth ors

(V.A.K.) and his associates2830 shed some light on the

properties of human blood vessels; however, more in-

sight is needed on the structure and properties of ar-teries before small diameter grafts with maximum po-

tential success can be engineered.

The objectives in the p resent work w ere to study the

structure and mechanical behavior of a human com-

mon carotid artery (CCA), and to use the information

in designing and constructing a compliant textile vas-

cular graft. The gr aft so constructed was composite in

structure an d characterized by nonlinear elasticity and

large transverse deformation. Animal studies con-

ducted with the grafts showed highly promising re-

sults.

MATERIALS AND METHODS

Mechanical properties and structure ofhuman CCA

Seven hu man CCAs, retrieved at the autopsies of

persons aged 2135 years, were used as experimental

materials. The vessels were marked with gentian vio-

let stain before resetting to identify the in situ axialextension ratio. After resection, the specimens were

stored in physiological salt solution until the mechani-

cal tests, conducted within 2 h, were performed. The

device used for these tests is shown in Figure 1. An

artery was cannulated at both ends. The sample was

placed in the chamber with the physiological salt so-

lution maintained at the temperature of 37 1C. One

end of the tube was clamped to a support to which a

pressure transdu cer (Micron Inst. M-15) and a spe-

cially designed inductive force transducer were con-

nected. The other end was clamped to a support to

which a pressure bottle containing fluid was con-

nected. The force transd ucer recorded the force neces-

sary to maintain the vessel at its in situ length. Axial

stretch w as introdu ced by a slide m echanism to w hich

the balance arm s w ere fixed. The axial deformation of

the artery was measured with a specially designed

inductive strain transdu cer connected to one of the

arms of the balance. Diameter changes in the specimen

were sensed optically with a video-dimensional ana-

lyzer coupled with a suitable lighting system for high

contrast. The changes in diameter with pressure were

tracked and recorded continuously.

An arterial samp le was gradu ally loaded by internal

pressure from 0 to 200 mmHg while maintaining thelength of the sample constant at L0, the length in situ.

The pressure was elevated in 20-mmHg steps with

pressure held constant in each step for 1 m in. The

initial external diameter a t inner p ressure p = 0 mmH g

a n d a t in situ axial length L0 was noted as D0. The

diameter D was recorded at each pressure level. The

value of wall thickness h was calculated as follows:

h = h0 3, (1)

where

3 =1

1 2, (2)

2 = (D/D0), (3)

an d

Figure 1. The schematic of the experimental device. 1,sample in chamber with a physiological solution; 2, force

transd ucer; 3, pressure tr ansd ucer; 4, TV camera; 5, displace-ment transd ucer; 6, stepper motor; 7, system for liquid feed-ing.

342 GUPTA AN D KASYAN OV


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1 = (L/L0) = 1.0. (4)

In these equations, h0 is the initial thickness of the

specimen wall and 1, 2, and 3 are, respectively, the

stretch ratios in the axial, circumferential, and radial

directions. Because the length of the artery was main-tained constant at L0, the value of1 (=L/L0) was 1.

The initial wall thickness h0 was measured with a

cathetometer to 0.001 mm accuracy. The artery was

preconditioned before the tests by subjecting it to cy-

clic loading to bring it to a stable state, which would

give a more reproducible mechanical response. Dur-

ing this process the vessel was p ressurized from 0 to

200 mmH g in 20 mm Hg steps five times with pr essure

held constant for 1 m in at each step. The initial curves

were m arkedly hysteretic, but the th ird or fourth cycle

gave reproducible curves with minimal hysteresis.

The structure of the inner layers was studied after

sectioning under a JEM-100C scanning electron micro-

scope (SEM). For this the p rep ared section w as fixed in

3% glutaraldehyde for 8 h, postfixed in 1% osmium

tetroxide for 5 h, dehydrated in ethanol of increasing

concentration, and dried in a Japanese Eiko-Dh-I criti-

cal point apparatus. The dried section was coated in a

vacuum chamber with 5060 nm of gold and exam-

ined under an SEM with a ASJD-4D scanning attach-

ment at 40-kV accelerating voltage and 80030,000

magnification. Light microscopy (LM) w as used to

gain preliminary information about the constitutiveelements of the connective tissue. For the latter, the

samples were fixed in 10% formaldehyde (pH 7.0) and

then embedd ed in p araffin wax, sectioned in various

directions, and stained using hematoxylin-eosin. The

co ll ag e n fi be r s w e r e d y e d u s in g t h e v a n G is on

method, and the elastin fibers were dyed using the

Weigert method.

Manufacturing, mechanical properties, and

structure of composite textile grafts

The performances of textile comp osite vascular

grafts designed and constructed in this work and of

available commercial w oven grafts (CVG)1 (6-mm di-

ameter, woven structure m ade of 7 tex linear d ensity

polyester, North, St. Petersburg, Russia) were evalu-

ated. In ma nufacturing the former, textile threads w ith

two widely different deformative characteristics, one

nearly matching those of the elastin and the other of

the collagen fibers, were selected. The m aterials u sed

were a polyurethane monofilament yarn (VolgogradChemical Thread Plant, Volgograd, Russia) with a low

modulus of elasticity (0.8 MPa) and a bulked polyester

multifilament yarn (Mogilev Chemical Thread Plant,

Mogilev, Russia) with a high modulus of elasticity (1.4

102 MPa). Tubular grafts of diameters 46 mm were

made by a weaving process utilizing a foil ribbon

loom.

Two types of grafts were made. In the first type

(HVG-1), polyester threads of 9 tex linear density were

used as the warp (Fig. 2), and the same polyester and

a prestretched polyurethane (7.8 tex linear density)were used as the w eft. The stretch in the elastic thread

caused the crim p to d evelop in the weft threads,

which made the graft stretchable in the transverse di-

rection. In the second variant (HVG-2), prestretched

polyurethane thread combined with p olyester w as

used as both the warp and the weft. This combination

in the warp threads is novel in that the crimp, usually

introdu ced in the longitud inal direction by the tedious

process of crimping and heat setting in commercial

grafts, developed autom atically in this gra ft by d iffer-

ential shrinkage. The gr afts obtained were stretchableand thus compliant in both the transverse and the lon-

gitudinal directions.

The mechanical prop erties of the grafts so prod uced

were determined at pressures ranging from 0 to 200

mmHg by following the procedure described for the

carotid arteries. Five specimens of each type were

tested. A thin latex tube of diameter larger than those

of the grafts was inserted into the sp ecimen before the

fluid was passed and the graft pressurized. For deter-

mining hydraulic permeability, the procedure of Gui-

doin et al.31 was used. In this method, the volume of

water passing through the wall under a fixed hydro-static pressure of 120 mmHg was collected for 5 min

and expressed as milliliters per minute p er square cen-

timeter of water. The grafts were also examined u nd er

an SEM for their surface and pore characteristics.

Eight HVG-1 hybrid grafts were also implanted in

the carotid and femoral regions of mongrel dogs1 for

periods of up to 1 year. (In performing the in vivo

studies, the guidelines of the Scientific Councils of the

Latvian Academy of Sciences and Academy of Medi-

cine w ere followed.) The form of the pulse w ave

found in the graft, the healing characteristics of the

Figure 2. The structure of the hybr id textile vascular grafts:(a) the p olyester yarn and (b) the p olyurethane yarn.

343CAROTID ARTERY/ BIOMECHANICS AND NOVEL VASCULAR GRAFTS


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surface, and the changes in the mechanical properties

as a result of implantation were examined.

Compliance and stiffness parameter

The flexibility and stiffness of arteries or grafts have

been frequently characterized by the values of com-

pliance,32 Cv, pressurestrain elastic modu lus,33 Ep ,

and stiffness parameter,27 . These parameters are de-

fined as follows:

Cv = (2/ De) (De/P); (5)

Ep = De (P/De); (6)

ln (P/Ps) = (D/Ds 1). (7)

In these De is the external diameter at 80 mm Hg, P

is the pressure difference (12080 mm Hg) over wh ich

measurements are made, Ps is the mean systemic pres-

sure (100 mmHg), and Ds is the corresponding exter-

nal diameter. Compliance Cv is thus the fractional

change in external d iameter, De, w ith change in

physiological pressure, P, from 80 to 120 mmHg (orfrom 10.67 to 16 kPa), and the coefficient is the slope

of the natural logarithm (ln) pressurediameter curve

and thus represents the stiffness of the vascular wall.

Because the pressurediameter relation of an arte-

rial wall is generally nonlinear, even within the ph ysi-ological pressure range, the parameters Cv an d Ep ar e

not usually material constants but change with the

internal pr essure. The stiffness param eter , however,

is independent of the pressure and has been used to

characterize the elastic properties of polyurethane

grafts.6,13

RESULTS

A typical pressure p to circumferential stretch ratio

2 relationship for hum an CCA, which differs fromthe relationship generally found on traditional p oly-

mer materials, is shown in Figure 3. The p2 curves

for the arteries are concave upward as expected.27,30

The slope gradually increases with stretch ratio and

becomes exceptionally high when 2 reaches a value

of about 1.42 0.12 (or when pressure exceeds 120

mmH g). The u nderlying structure of the arterial wall

must be responsible for the noted mechanical behav-

ior. Studies by Wolinsky and Glagov34 and Lange-

wou ters et al.35

show tha t the wall tissue is mad e up ofat least two major fibrous materials, an elastin fiber

with a low modulus of elasticity (25 105 Pa) and a

collagen fiber w ith a high mod ulus of elasticity (510

107 Pa). In the relaxed state, the collagen fibers are

slack. At low strain or pressure, most of the load is

borne by the elastin fibers, and therefore, the artery is

highly distensible. However, as the pressure is raised

a n d t h e s t r ai n i s i n c r ea s ed , t h e c ol la g e n f ib e r s

straighten out and start to bear load. This process

causes an increase in the slope.

Results from LM and the SEM show that carotid

arteries have a sp ecific wavy structure an d the tissue is

a biocomposite (Figs. 4 and 5). The wavy membranes

consist of the elastic fibers plaited w ith collagen fibers.

With an increase of internal pressure the diameter of

the artery increases and the degree of waviness de-

creases [Figs. 4(b) and 5(b)]. At a pressure of 120

mmH g and the longitudinal stretch ratio correspond-

ing to the in situ length, the waviness of the wall ele-

ments practically disappears. With a further increase

in pressure, the collagen fibers begin to resist the cir-

cumferential strain in the artery and thus give rise to

the behavior noted in Figure 3.

The behavior of grafts H VG-1 and HVG-2, also

shown in Figure 3, is similar in character to that of the

c ar o t id a r t er y . T h e p o l yu r e t h a n e a n d p o l ye s te r

threads in the former seem to play about the sameroles as the elastin and the collagen fibers, respec-

tively, in the latter. Some differences are noted in the

values and the shapes of the curves. A pressure of 160

mmHg or greater is needed in the grafts to straighten

out the higher modulus fibers; in arteries the corre-

Figure 3. Pressure p to circumferential stretch ratio 2 re -lationship. 1, commercial vascular graft (CVG); 2, h ybridtextile vascular graft (HVG-1, after 3 months of imp lanta-tion); 3, hybrid textile vascular graft (HVG-2); 4, hybrid tex-tile vascular graft (HVG-1); and 5, human common carotidartery.



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sponding pressure was lower. For the pressure range

80120 mm Hg, the increase in circumferential stretch

ratio w as 7.14% for the carotid arteries, 9.02% for the

HVG-1 graft, 6.06% for the HVG-2 graft, and only

0.32% for the commercial graft. These results show

that the commercial grafts used are the least disten-

sible of all materials examined and are n ot biocomp li-

ant. Comp liance values of the hybrid textile grafts

compare favorably with those of the carotid arteries

(Table I).

Comparison of the Cv an d Ep values shows that the

deformability of the HVG-1 graft is greater than that of

the carotid ar tery. Based on the valu es of the stiffness

parameter , one can conclude that graft HVG-1 is

more compliant than graft HVG-2, the stiffness coef-

ficients being 4.87 1.56 and 7.81 1.62, respectively.

Clearly, the differences in the values of the two graftsarise from the differences in their structures, indicat-

ing that a graft with the d esired m echanical properties

could be constructed if the structure was carefully

controlled.

The results of the water perm eability obtained show

that HVG-1 grafts are less w ater p erm eable than

HVG-2 grafts, the values being 2.06 0.16 and 2.42

0.18 mL/ m in/ cm 2. These values are about the same as

obtained by others on grafts.17,31

HVG-1 grafts of 4-mm internal diameter were im-

planted in eight mongrel dogs for periods of 1 monthto 1 year. Four were implanted in the carotid and the

other four in the femoral regions. Six grafts functioned

satisfactorily, but the remaining two (femoral) devel-

oped a stenose of d istal anastomoses. Six months a fter

implantation, the circumferential stretch ratio (for the

ran ge 80120 mm Hg ) d ecreased from 9.02 to 2.62 (Fig.

3), due obviously to some inward growth of the sur-

rounding tissues.

The structure of the grafts was examined with LM

and an SEM before implantation and after retrieval

following implantation. Figure 6 shows the structureof the outer and inner walls of HVG-1 before implan-

tation. The loops noted are those of the bulked poly-

e st e r y a r n s f or m e d b y t h e r e c ov e ry o f t h e p o ly -

urethan e yarns. Figures 7 and 8 show typical morp hol-

ogy found in the six p atent grafts after implantation.

These figures indicate that the healing process pro-

Figure 5. SEM of the circumferential slice of the carotidartery wall: (a) at zero internal p ressure and (b) at internalpressur e of 120 mmH g (sample 9, man 34 years old; originalma gnification 5000).

Figure 4. Circumferential histological slice of the carotidarterial wall: (a) at zero internal pressure and (b) at internalpressure of 120 mmHg (sample 9, man 34 years old; originalmagnification 50).



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gressed norm ally. At retrieval, none of the grafts

showed any aneurysm formation, p erigraft h emato-

mas, or rupture. All anastomoses were intact, and ex-

ternal tissue reaction was minimal. The inner surface

was lined with a smooth, thin layer of transparent

glistening tissue, with occasional small foci of yellow-

tan staining, but there was no evidence of adherent

thrombus (Fig. 7). SEM examination showed that the

endothelial-like cells app eared flattened with elon-

gated nuclei (Fig. 8). The results of the pulse wave

measured on the dogs showed that the pulse wave

obtained with compliant grafts was practically of the

same form as found with the carotid artery (Fig. 9).

DISCUSSION

Although many factors affect the success of a vas-

cular graft in surgery, the two most important are the

mechanical characteristics of the graft and the ability

of the graft to heal. The hope for achieving long-term

patency in 6-mm grafts lies in matching the compli-

ance of a vascular graft with that of the artery and

developing a thromboresistant surface at the luminal

wall.

Numerous studies have shown that a positive cor-

relation exists between the matching of compliance

and the p atency of the graft.1,2,12,14,36,37 Preliminary

studies show15 that compliance is a particularly im-

portant factor during the first six or so weeks of the

operation; long-term compliance may be a negligible

factor in determining the overall patency of small

grafts. In an ideal graft, the velocity of flow is high

while the stresses at the suture lines and the reflected

energy losses are small. For achieving minimum stress

concentration and energy losses in the existing com-

mercial grafts, the diameter of the graft chosen has to

be 1.41.5 times the diameter of the host artery.38

An impedance mismatch between the host and thegraft leads to development of eddy currents and wave

reflections. These contribute to structural fatigue, false

aneurysm, thrombus formation, hyperplastic tissue

growth, an d atherosclerotic changes in th e host artery,

which can lead to early occlusion of the graft.14,39 The

amount of reflected energy can be reduced to zero if a

perfect match exists between the fluid impedances of

the host artery and the graft. In a study by Scott and

Wilson40 in wh ich blood flow behavior in the hu man

leg was sim ulated with a com puter m odel, it was

shown that a m atch of both the diam eters and the

compliances of the prosthesis and the host vessel were

needed to maximize flow velocity while minimizing

reflected energy losses and stresses at th e sutu re lines.

The present work shows that in the human CCA,

the slope of the pressure (p)circumferential stretch

Figure 6. SEM of th e h ybrid textile vascular gr afts: (a) out-side structure an d (b) inside structure.

TABLE IMean Value SD of Human Common Carotid Artery (CCA), Textile H ybrid Vascular Grafts (HVG-1 and HVG-2),

and Commercial Graft (CVG)

Artery or

Graft

Specimen

Wall

Thickness

(mm)

Diameter

(mm)

Compliance, Cv

(kPa1

)

Elastic

Modulus,

Ep (kPa)

Stiffness

Parameter,

CCA 1.54 0.12 5.92 0.47 0.0238 0.0132 83.86 22.38 5.18 1.94

H VG-1 0.42 0.04 4.02 0.04 0.0324 0.0083 61.81 16.92 4.87 1.56

H VG-2 1.23 0.06 4.05 0.06 0.0227 0.0078 87.78 19.04 7.81 1.62

CVG 0.27 0.02 6.01 0.04 0.00186 0.0005 1074.64 136.18 102.91 16.84



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ratio (2) curve increases w ith pressure and becomes

exponential at pressures greater than 120 mmHg (Fig.

3). The LM and SEM analyses show that the arterial

wall is a biocomposite of elastin and collagen fibers

and has a specific crimped character. This structure, as

explained earlier, leads to the unu sual m echanical be-

havior noted above, which differs from that of the

traditional polymer materials. The value of compli-

ance measured on hum an carotid arteries coincides

with the values obtained by Hayashi et al.27 w ho

found that the stiffness param eter had a value of 5.25 for the hum an CCA and 19.84 for the femoral

artery. Our r esults show a v alue of 5.18 for the h um an

carotid artery (see Table I).

The novel hybrid textile vascular grafts prod uced in

this work have elastic properties that match those of

the hu man carotid artery. The materials used in their

construction were polyurethane and polyester yarns

that allowed the hybrid textile grafts to have a com-

pliance value in the circumferential direction 10 or

more times those of the currently available commer-

cial grafts. The w ater p ermeability value of the h ybrid

grafts is at the same level as found in the commercial

prostheses. All commercial grafts are crimped by one

or two heat-setting processes. These grafts rapidly lose

crimp and longitud inal compliance up on implanta-

tion.17 The procedure used in manufacturing the hy-

brid textile prostheses, on the other hand, renders the

heat setting process for crimping unnecessary. The

combination of prestretched polyurethane yarns with

polyester in the warp threads led to the d evelopment

of a stable crimp in the longitudinal direction. More-

over, by varying the sizes and properties of the indi-

vidual yarns, the structure and properties of vasculargrafts could be effectively and conveniently engi-

neered to suit the application.

However, the mechanical compliance alone does

not guarantee success in vivo. Kambic et al.41 showed

that no one material satisfies the requirements for by-

pass of all small caliber arteries. Future research mu st

additionally focus on development of improved inner

surface coatings, which may minimize tissue reaction

and undesirable cellular events at the anastomoses. A

g e la t in co a tin g w a s u s e d t o p r o v id e a b lo od -

compatible graft with a smooth nonpseudoneointimagenerating surface that does not prom ote cell in-

growth at the anastomoses.15

A difficult problem faced in matching compliance is

a change in the mechanical properties of the graft that

is induced by tissue ingrowth. The in vivo test results

on the grafts of this study are encouraging. Examina-

tion of the HVG-1 grafts after 6 months of implanta-

tion indicated th at the h ealing p rocess progressed n or-

mally. Some loss of compliance was noted due to tis-

sue ingrowth into the wall. Hasegava and Azuma,42

working with woven Dacron or Teflon grafts, foundthat 3 weeks postimplantation, the longitudinal stiff-

ness h ad increased but the circumferential stiffness

had not changed. On the other hand, the w ork of Lee

and Wilson37 showed a marked increase in circumfer-

ential stiffness after 3 months of implantation, indicat-

ing that at 3 months the connective tissue ingrowth

was orga nized an d played a significant role in increas-

ing wall stiffness.

It was also found43 that although both neointima

and adventitia cells were closely attached to the poly-

urethane fibers near the surface, there was no trans-

mural or through growth of the tissues. After 9months of implantation in m ini-pigs, the grafts w ere

still fairly compliant. Studies on spandex prostheses in

dogs indicated 18 that while these grafts lost some cir-

cumferential stretchability after becoming infiltrated

with unyielding collagenous tissue, they remained

compliant d uring the first w eek after insertion and

adapted their diameter to the flow conditions of the

arteries that were bypassed. The spandex grafts usu-

ally adjust their diameter in response to changing flow

Figure 7. Hybrid vascular graft HVG-1 after 6 months ofimplantation. The interior is lined with a thin layer of trans-parent glistening tissue w ithout thrombu s.

Figure 8. Endothelium cells on the flowing surface of theHVG-1 graft.



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and pressure conditions, particularly the former, and

develop a thin, almost transparent, neointima on the

flow surface.

Another important parameter that determines the

success of a graft in bypass application is porosity.Tissue incorporation and healing of synthetic grafts

are related to this property.44 The pore size affects the

type of tissue grown through the wall.4 Working with

polyurethane grafts, White4 showed that if the pore

size was less than 15 m, minimal tissue ingrowth

took place; if it was greater than 15 m but less than

about 45 m, fibrohistiocytic tissues grew; and if the

size w as greater than about 50 m, the structure was

infiltrated with organized fibrous tissue.

The above results thus indicate that by choosing

appropriate materials and controlling the structure, atextile graft can be engineered that h as the desired

compliance and the potential for attaining the needed

tissue ingrowth. An important question still remains;

however: Should th e properties of the host be matched

by those of the virgin prostheses or those of the pros-

theses after they have been at the site for some length

of time?45 To address this question, it is necessary to

know what changes take place in the properties of a

graft with time in situ and how these affect the func-

tion of the product. Clearly, this knowledge is needed

before an optimally useful graft can be designed.

The benefits accruing to the au thors from a commercial or

industrial party will be applied to a research fund, nonprofit

institution, or other organization with which the authors are

associated.

References

1. I. G. Kidson and W. M. Abbot, Low compliance andarterial graft occlusion, Trans. Am. Soc. Artif. Intern.Organs, 58, 1.11.4 (1978).

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Figure 9. Shape of the pulse waves: (a) dog carotid artery,(b) commercial graft, and (c) hybrid vascular graft HVG-1.



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Received February 6, 1995Accepted March 8, 1996


bio mechanics of human common carotid artery and design

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