An article of manufacture and method of making and implanting the article made of a polyolefin star or linear copolymer are disclosed in which the polyolefin copolymer is biostable and crack-resistant when implanted in vivo. The polyolefin copolymer is the reaction product of a rubbery component which when homopolymerized produces a polymer having a low level of hardness, and a hardening component which when homopolymerized produces a polymer having a high level of hardness. The polyolefin copolymer is elastomeric, has a hardness intermediate the low and high levels of hardness, and has a backbone in which the majority of polymer linkages along the copolymer chain are alternating quaternary and secondary carbon atoms.

BACKGROUND AND DESCRIPTION OF THE INVENTION 
The present invention generally relates to implantable prostheses and the 
like which are formed in a manner to substantially prevent cracking, 
crazing or degradation thereof when they are implanted or otherwise 
subjected to degradation conditions. A medical prosthesis or the like 
according to this invention includes a polyolefinic elastomeric triblock 
star or linear copolymer where the backbone comprises alternating units of 
quaternary and secondary carbons which will not crack or degrade when 
subjected to implantation for substantial time periods during which other 
types of polymers would crack or degrade. 
Several biocompatible materials which are quite suitable for use in making 
implantable medical devices that may be broadly characterized as 
implantable prostheses exhibit properties that are sought after in such 
devices, including one or more of exceptional biocompatibility, 
extrudability, moldability, good fiber forming properties, tensile 
strength, elasticity, durability and the like. However, many of these 
otherwise highly desirable materials exhibit a serious deficiency when 
implanted within the human body or otherwise subjected to harsh 
environments, such deficiency typically being manifested by the 
development of cracks or fissures. For example, surface fissuring or 
cracking occurs after exposure of on the order of one month or more, or 
shorter time periods depending upon the materials and the implant 
conditions, to body fluids and cells such as are encountered during in 
vivo implantation and use. 
It is desirable that long-term implantable elastomers, such as those used 
for vascular grafts, endoluminal grafts, intraocular lenses, finger 
joints, indwelling catheters, pacemaker lead insulators, breast implants, 
heart valves, knee and hip joints, vertebral disks, meniscuses, tooth 
liners, plastic surgery implants, tissue expanders, drug release 
membranes, subcutaneous ports, injection septums, etc., be stable for the 
duration of the life span of the recipient. 
Polymers that are not stable in the physiological environment tend to crack 
and degrade with time. There are many implant applications where this type 
of behavior cannot be tolerated. For example, pacemaker lead insulators 
can form current leaks thereby causing the wires to short out and the 
pacemaker to be rendered non-functional. It is therefore desirable to have 
a material for long-term use that is both elastomeric and does not degrade 
in the body. 
Several theories have been promulgated in attempting to define the cause of 
this undesirable cracking phenomenon. Proposed mechanisms include 
oxidative degradation, hydrolytic instability, enzymatic destruction, 
thermal and mechanical failure, immunochemical mechanisms, inhibition of 
lipids and combinations of the above. Prior attempts to control surface 
fissuring or cracking upon implantation or the like have included 
incorporating antioxidants within a biocompatible polymer and subjecting 
the biocompatible polymer to various different annealing conditions, 
typically including attempting to remove stresses within the polymer by 
application of various heating and cooling conditions. Attempts such as 
these have been largely unsuccessful. 
Other treatment approaches have been utilized, or attempted, to increase 
the structural stability of especially desirable materials. Included in 
the biocompatible materials which are desirable from many points of view, 
but which exhibit a marked tendency to crack or degrade over time, are the 
polyurethane materials and other biocompatible polymers that are of an 
elastomeric nature. It is particularly advantageous to use these types of 
materials for making products in which compliance and/or flexibility, high 
tensile strength and excellent fatigue life are desirable features. One 
basic approach which has been taken in the past in order to render these 
materials more suitable for implantation and other applications where 
material degradation can develop, has been to treat the material with 
so-called crack preventatives. Exemplary approaches in this regard are 
found in my U.S. Pat. Nos. 4,769,030, 4,851,009 and 4,882,148, the subject 
matter of which is incorporated by reference herein. Sulfonation of 
polyurethanes to prevent cracking is also described in my U.S. Pat. No. 
4,882,148, the subject matter which is also incorporated by reference 
herein. Such treatments, of course, require additional procedures, and 
post processing of the implantable article, thereby complicating 
manufacturing procedures, increasing expense and complexity and, if not 
coated or treated properly and entirely, are subject to delamination and 
failure. It would be advantageous if the material out of which the product 
is made would itself have the desired properties. It is also advantageous 
for the material to be compatible with other materials that are commonly 
used in the medical fields, such as with adhesives, surface coatings and 
the like. 
An especially difficult problem is experienced when attempting to form 
prostheses with procedures including the extrusion or spinning of 
polymeric fibers, such as are involved in winding fiber-forming polymers 
into porous vascular grafts or similar products, for example as described 
in U.S. Pat. No. 4,475,972 (Wong), the subject matter of which is also 
incorporated by reference herein. Such vascular grafts or the like include 
a plurality of strands that are of a somewhat fine diameter size such 
that, when cracking develops after implantation, this cracking often 
manifests itself in the form of complete severance of various strands of 
the device. Such strand severance cannot be tolerated to any substantial 
degree and still provide a device that can be successfully implanted or 
installed on a generally permanent basis whereby the device remains viable 
for a number of years. 
There is accordingly a need for a material which will not experience 
surface fissuring or cracking under implanted or in vivo conditions and 
which is otherwise desirable and advantageous as a material for medical 
devices or prostheses that must successfully delay, if not eliminate, the 
cracking phenomenon even after implantation for months and years, in many 
cases a substantial number of years. Exemplary medical devices or 
prostheses for which such a non-cracking material would be especially 
advantageous include those which have been previously discussed. 
The only elastomers that are currently implanted are polyurethanes, as 
previously discussed, and silicone rubbers. 
The silicone rubbers, most notably polydimethylsiloxane, are probably the 
most stable elastomers used in the body. However, there have been many 
reported instances where they do not perform well. For example, silicone 
rubber poppet valves for coronary valve replacement tend to swell and 
crack with time, heart valve leaflets tend to calcify, silicone gel-filled 
breast implant shells tend to plasticize with silicone oils and in many 
instances, rupture with time. The mechanism of biodegradation of silicones 
in the body is believed to involve oxidative pathways. 
Three families of polyurethanes have been used in long-term implant 
applications; i.e. the polyester urethanes which have been used as foamed 
coatings on some breast implants, the polyether urethanes which have been 
used as insulators on pacemaker leads, and the polycarbonate urethanes for 
use in vascular grafts. Polyether and polyester urethanes have repeatedly 
been shown to degrade with time in the body. L. Pinchuk, A Review of The 
Biostability and Carcinogenity of Polyurethanes in Medicine and the New 
Generation of "Biostable" Polyurethanes, J. Biomaterial Science, Polymer 
Ed., Vol 6, No. 3, pp 225-267 (1994). 
The more recent family of biostable elastomeric polyurethanes which contain 
polycarbonate groups, rather than other or ester groups, are described in 
my U.S. Pat. Nos. 5,133,742 and 5,229,431, the subject matter of which is 
incorporated by reference herein. A similar polycarbonate urethane, but of 
a lower modulus of elasticity, is disclosed in U.S. Pat. No. 5,254,662 
(Szycher et al). All of these polymers have demonstrated much improved 
biostability as compared to the polyether and polyester urethanes. 
However, as also described in my last mentioned review, some cracking and 
fiber breakage are observed on microfibers comprising a polycarbonate 
urethane vascular graft with time. 
Still another biostable polyurethane is described in U.S. Pat. No. 
4,873,308 (Coury et al). It is formed of all aliphatic soft segments of 
predominantly consecutive secondary carbon atoms. Two potential weaknesses 
of this polymer are that the secondary carbon atoms can oxidize with time, 
and the urethane linkages present on the backbone can hydrolyze with time. 
In addition, as reported in my aforementioned review, the polymer weakens 
with exposure to moisture and has a modulus with a yield point. 
Again referring to my aforementioned review, a number of investigators have 
demonstrated that biodegradation of materials is usually a result of 
oxidation. Cells, especially leukocytes, secrete superoxide and hydrogen 
ions which subject the material to high concentrations of oxidants (free 
radicals) and strong acids. It is therefore a principal purpose of this 
invention to formulate long-term implantable materials with molecules that 
are not readily susceptible to oxidation and to attack by acid. 
Two non-elastomeric polymers that have performed well in the body include 
polytetrafluoroethylene (Teflon) and polymethylmethacrylate with few, if 
any, reports of biodegradation. Other polymers in addition to those 
already discussed that have also enjoyed some measure success in the body, 
but do demonstrate some degree of biodegradation with time, include 
polypropylene, polyethylene and to some degree polyester terephthalate 
(PET). 
Examination of the chemistry of those polymers reveal that, except for 
polytetrafluoroethylene, which is inert due to the replacement of all 
hydrogens with fluorine, the most inert polymers are those with the most 
"quaternary" carbons, as defined below. The principal problem with all of 
these polymers is that they are non-elastomeric and, therefore, cannot be 
used in certain applications in the body, such as vascular or endoluminal 
grafts where elasticity is desired. 
Polymethylmethacrylate has repeating units of: 
##STR1## 
Polypropylene has repeating units of: 
##STR2## 
Polyethylene has repeating units of: 
EQU --(CH.sub.2).sub.n -- 
The number of repeating units is usually sufficiently large so that the 
molecular weight of the polymer is in excess of 60,000 Daltons. It will be 
seen that polyethylene is comprised only of "secondary" carbons, i.e. each 
carbon atom on its backbone is bonded to two other carbon atoms. 
Polypropylene has alternating "secondary" and "tertiary" carbons. A 
"tertiary carbon" is a carbon that is bonded to three other carbon atoms. 
A "quaternary" carbon is a carbon that is bonded to four other carbon 
atoms. The polymethylmethacrylate has a backbone of alternating 
"quaternary" carbons and "secondary" carbons. 
Further examination of polyethylene will reveal that, in the presence of 
free radicals and other oxidizing agents, the polyethylene molecule and 
its secondary carbons can undergo abstraction of hydrogens and the 
formulation of free radicals and double bonds, e.g. 
##STR3## 
Double bonds can also lead to intermolecular or intramolecular 
crosslinking. Once the double bond, unsaturation or crosslinking forms in 
the polymer, the polymer can become embrittled leading to cracking or 
degradation. For this reason polyethylene is hardly used anymore for 
pacemaker lead insulators, because it embrittles and then cracks and 
flakes with flexion in the body. 
Similarly, but not as frequently, polypropylene can oxidize to the 
formation of a double bond between the tertiary carbon and the secondary 
carbon, i.e. 
##STR4## 
Polypropylene, when loaded with antioxidants, is successfully used as a 
suture in the body, but does show some degradation with time as a haptic 
on intraocular lenses. 
On the other hand, polymethylmethacrylate, has quaternary and secondary 
carbons along its backbone. Therefore, it is not readily susceptible to 
oxidation. Formation of a double bond along the backbone of the polymer 
would require the cleavage of carbon to carbon bonds, i.e. 
##STR5## 
rather than carbon to hydrogen bonds as in the secondary and tertiary 
carbons. Extremely high energies are required to break carbon to carbon 
bonds. It is for this reason that polymers with alternating quaternary and 
secondary carbon bonds are very stable in the body. 
The problem with the polymethylmethacrylate, polypropylene and polyethylene 
polymers is that they are not elastomers. They are rigid engineering 
plastics. Therefore, they cannot satisfy a need in the medical industry 
for a flexible polymer with excellent oxidation resistance, such as one 
that has alternating units of quaternary and secondary carbons. 
The present invention achieves these objectives with a polymer which is a 
polyolefinic elastomer of a triblock star or linear copolymer backbone 
having alternating units of quaternary and secondary carbons. The polymer 
should have a resultant hardness which is between about Shore 20A-75D, and 
preferably between about Shore 40A and Shore 90A. 
Accordingly, a general object of the present invention is to provide 
improved crack-resistant devices and products. 
Another object of the present invention is to provide a polymeric material 
and products made therefrom which are particularly resistant to cracking 
and degradation, even under in vivo conditions. 
Another object of the present invention is to provide an improved 
polyolefin material which can be spun through a spinnerette or extruded 
through and/or into suitable molding devices into products which exhibit 
superior crack-resistant properties, and/or which can be injection or 
compression molded, solvent castable, or solvent sprayable into such 
products. 
Another object of the invention is to provide improved implantable devices 
and/or prostheses which exhibit an exceptional ability to prevent the 
formation of cracks and strand severance upon implantation for substantial 
time periods, such as those needed for generally permanent implantation 
procedures. 
Another object of the present invention is to provide an improved vascular 
graft and the like that is made from spun fibers of polymer and that 
exhibits exceptional stability with respect to crack formation and strand 
severance development under in vivo conditions. 
These and other objects, features and advantages of the present invention 
will be clearly understood through a consideration of the following 
detailed description.

DESCRIPTION OF THE PREFERRED EMBODIMENTS 
In the preferred embodiments of the present invention, a polymer is 
provided which is a polyolefinic elastomer having a backbone which 
comprises a triblock linear or star copolymer having alternating units of 
quaternary and secondary carbons. The term "secondary" carbons, as 
previously defined, means carbon atoms which are bonded to two other 
carbon atoms. The term "tertiary" carbons, as previously defined, means 
carbon atoms that are bonded to three other carbon atoms. The term 
"quaternary" carbons, as previously defined, means carbon atoms which are 
bonded to four other carbon atoms. 
The polyolefinic elastomer copolymer of the present invention contains at 
least two components I and II. 
Component I is a rubbery or soft segment component which is based upon a 
repeating unit of a quaternary carbon and a secondary carbon, having the 
following general formulation: 
##STR6## 
where R and R' are aliphatic moieties, such as methyl, ethyl, propyl, 
butyl, pentyl, hexyl or cyclic aliphatic groups. The preferred rubbery 
component I is polyisobutylene (PIB) with the following structure: 
##STR7## 
Pure polymers of PIB are commercially available. One of their principal 
uses is as the gum stock in chewing gum. Crosslinked or vulcanized PIB is 
used as inner tubes in tires, and simple low molecular weight PIB chains 
are used in high temperature lubricants. 
High molecular weight PIB is a soft material with a Shore hardness of 
approximately 10A to 30A. When combined with other block copolymers, it 
can be made at hardnesses ranging up to the hardness of the copolymer. For 
example, if it is copolymerized with polystyrene, and if the polystyrene 
has a Shore hardness of 100D, depending upon the relative amounts of 
styrene and isobutylene, the resultant polymer can have a range of 
hardnesses from as soft as Shore 10A to as hard as Shore 100D. 
Component II of the block copolymer is a hardening component. It may 
include anyone of a number of monomers or polymers as long as the majority 
of polymer linkages along the copolymer chain are comprised of alternating 
quaternary and secondary carbons, and the hardening component II, when 
combined with the rubbery or soft component I, is capable of altering or 
adjusting the hardness of the rubbery or soft component so that the 
ultimate polyolefin copolymer has the desired elastomeric and hardness 
qualities. Typical hardening component comonomers or copolymers used as 
copolymers or block copolymers, or more specifically triblock copolymers, 
with polyisobutylene (PIB) in the present invention are styrene, 
.alpha.-methylstyrene, methylmethacrylate, ethylmethacrylate, hydroxyethyl 
methacrylate and the like. 
Although PIB can be polymerized anionically, it is probably best 
polymerized under controlled means using carbocationic polymerization 
techniques such as those described in U.S. Pat. Nos. 4,276,394, 4,316,973, 
4,342,849, 4,910,321, 4,929,683, 4,946,899, 5,066,730, 5,122,572 and/or 
Re. 34,640, the subject matter of which is incorporated by reference 
herein. These materials may involve telechelic starting molecules, with 
block derived therefrom. Although the description to follow of the 
invention is set forth in terms of linear copolymers which are formed from 
ditelechelic starting materials, star copolymers are also contemplated in 
the invention. As chemists skilled in the art will appreciate, star 
copolymers are formed simply by using tri- rather than ditelechelic 
starting molecules and as disclosed in the above identified patents. 
The polyolefinic copolymer elastomer of the present invention has the 
general formulation: 
##STR8## 
and more preferably: 
##STR9## 
where R and R' are non-cyclic or cyclic aliphatic moieties, such as 
methyl, ethyl, propyl, butyl, pentyl, hexyl, heptyl and octyl groups with 
pendant aliphatic groups such as methyl, ethyl or propyl groups, 
preferably methyl groups, R" is a hydrogen, hydroxyl, methyl or ethyl 
group, and R'" is an aromatic (phenyl, benzyl or substituted benzyl) 
group, COOCH.sub.3, methoxy, ethoxy, aliphatic, cycloaliphatic, 
substituted aliphatic or other group, and such that homopolymers of 
##STR10## 
are high hardness materials. 
The rubbery component and hardening component I and II respectively are 
shown in the above copolymer formulations. The amount of hardening 
component II in the copolymer is preferably between about 20 wt % to 80 wt 
%, and more preferably between about 30 wt % to about 50 wt %. 
Repeating units n and m should range from about 250 to about 5000, and of p 
should range from about 1-10, with about 1 being typical. The combined 
molecular weight should be in excess of 60,000 Daltons, and preferably 
between about 90,000 to about 300,000 Daltons. Triblocks of PIB with 
polystyrene or polymethylstyrene are preferred polyolefin elastomeric 
copolymers of the invention and these may be made as described in U.S. 
Pat. No. 4,946,899, the subject matter of which is incorporated herein by 
reference. 
The copolymers of the present invention are preferably copolymerized in 
solvents. The solvents may be non-polar solvents, polar solvents or 
mixtures thereof. Suitable nonpolar solvents include hexane, cyclohexane, 
heptane, methylene chloride, toluene, Freon.RTM., low molecular weight 
silicones and the like. Suitable polar solvents include methanol, ethanol, 
propanol, tetrahydrofuran and the like. The solvent may also comprise a 
binary blend of the previously mentioned non-polar and polar solvents. For 
example, where the triblock copolymer is of isobutylene and styrene, the 
solvent is preferably a binary solvent of heptane in which isobutylene is 
soluble and methanol in which styrene is soluble. 
Other components such as antioxidants, extrusion agents and the like can be 
included, although typically there would be a tendency and preference to 
exclude such additional components a medical-grade polymer is being 
prepared. 
While no treatment of the polyolefin copolymer products according to this 
invention is required, suitable treatments can be conducted if desired. 
For example, they may be subjected to treatment with a crack preventative 
composition that includes an elastomeric silicone such as poly(dimethyl 
siloxane), as described in detail in my U.S. Pat. No. 4,851,009. 
EXAMPLE 1 
A potent in vitro screen for biostability is boiling the sample in 
concentrated (65%) nitric acid. Nitric acid is both a strong oxidant and a 
strong acid. 
Samples that have enjoyed some success in the body, such as the polyether 
urethanes, polyester urethanes, polycarbonate urethanes, Dacron (polyester 
terephthalate), Nylon 11, silicone rubber, natural rubber, PEEKEK 
(poly(ether-ether-ketone-ether-ketone)), polyethylene, polypropylene and 
polymethylmethacrylate were subjected to 65% boiling nitric acid for up to 
30 minutes. Also included was a sample of a carbocationically produced 
polyisobutylene and a copolymer according to the present invention which 
was a triblock polymer of styrene-PIB-styrene of 38 wt % styrene. The 
results are presented in Table 1. 
TABLE 1 
______________________________________ 
Tensile 
Time to Strength 
Sample dissolution Results Remaining 
______________________________________ 
Polyether &lt;3 seconds destroyed 0 
Urethane 
Polyester &lt;5 seconds destroyed 0 
Urethane 
Polycarbonate 
&lt;10 seconds destroyed 0 
Urethane 
Dacron Fiber 
&lt;10 seconds destroyed 0 
Nylon 11 &lt;30 seconds destroyed 0 
Natural Rubber 
&lt;50 seconds destroyed 0 
Crosslinked no dissolution 
very .about.0 
polycarbonate Brittle 
urethane 
PEEKEK Fiber 
no dissolution 
very .about.0 
brittle 
Silicone no dissolution 
very .about.10% 
rubber brittle 
Polyethylene 
no dissolution 
very .about.50% 
(low density) brittle 
PIB (homo- no dissolution 
plastic .about.60% 
polymer) deform- 
ation* 
polypropylene 
no dissolution 
no change .about.100% 
Polymethy- no dissolution 
no change .about.100% 
methacrylate 
Polystyrene no dissolution 
no change .about.100% 
Teflon no dissolution 
no change .about.100% 
Sty-PIB-Sty no dissolution 
no change .about.100% 
(38 wt % Styrene) 
______________________________________ 
*Melting point too low in sample tested. 
The polymers degraded in the same order as is observed in vivo; i.e. the 
polyether urethanes were destroyed first, followed by the polyester 
urethanes, followed by the polycarbonate urethanes, then Dacron, Nylon 11, 
then silicone rubber, etc. with no measurable degradation of 
polytetrafluoroethylene (Teflon) or polymethylmethacrylate. This validates 
this in vitro boiling nitric acid test of biostability. 
The PIB homopolymer was a mixture of different molecular weight 
polyisobutylenes and demonstrated plastic deformation due to partial 
melting of the polymer. 
The triblock polymer of styrene-PIB-styrene of the invention did not show 
any signs of degradation over the entire duration of the test and remained 
pliable and elastomeric. The fact that this triblock polymer did not 
degrade in concentrated boiling nitric acid is strong evidence that it is 
stable in the body for long durations. 
EXAMPLE 2 
A linear triblock copolymer resin of styrene-PIB-styrene of 34 wt % styrene 
according to the present invention was prepared. This copolymer resin was 
melt extruded from a cylinder through a heated orifice with a piston under 
pressure to form a continuous monofilament of the copolymer 1 mm in 
diameter. The monofilament was cut into two inch long samples. Some of the 
samples were sterilized in ethylene oxide, and some of the sterilized 
samples were implanted subcutaneously in the back of dogs for two and four 
weeks, respectively. Following explant of the samples which had been 
implanted, three of each of the following samples were tested under load 
to determine their breaking strength on an Instron Tensile Testing 
Machine, Model No. 1011: 
A--control, unsterilized, and not implanted. 
B--control, sterilized but not implanted. 
C--sterilized, implanted and explanted after 2 weeks. 
D--sterilized, implanted and explanted after 4 weeks. 
The results of the load tests are shown in FIG. 1. The "T" extensions shown 
on top of each of the bars in the FIG. 1 graph shows the high load value 
at which the filament being load tested broke. The top of each bar shows 
the average mean load value at breaking. Although not shown in FIG. 1, if 
the "T" extension of each bar was 180.degree. inverted to extend beneath 
the top of its bar, the minimum load value at which the filament being 
load tested broke would be depicted. Thus, by way of example the average 
mean load at which the control sample A broke was about 3 pounds force, 
the high load value was about 3.2 pounds force, and the minimum load value 
was about 2.8 pounds force. 
The differences in pounds force between samples A-D observed in the load 
tests and shown in FIG. 1 are not significant. Accordingly, FIG. 1 further 
shows that the triblock polymers of the invention are highly resistant 
against degradation under in vivo conditions. 
EXAMPLE 3 
The linear triblock polymer of the invention and as described in Example 2 
was formed into a porous non-woven membrane. To do so, the triblock 
polymer was dissolved in the amount of 6% solids in tetrahydrofuran. This 
solution was sprayed with an air brush onto a rotating mandrel. The 
environment was controlled during spraying so that the tetrahydrofuran 
evaporated between the sprayer and the mandrel and so that a porous mat 
was formed on the rotating mandrel. These samples were then fully dried in 
air and removed from the mandrel. 
A known polyether urethane was also formed into a non-woven membrane. The 
polyether urethane was of the kind which is regularly employed as 
implantable pacemaker leads and was Pellethane 2363-80A from Dow Chemical, 
Midland, Mich. Pellets of the polyether urethane were dissolved in the 
amount of 45% solids in dimethyl acetamide while agitating and heating. 
This solution was pumped through a spinneret with 30 orifices on a shuttle 
that reciprocated back and forth along a rotating mandrel so that the 
continuous fibers which issued from the spinneret orifices were wound on 
the mandrel. As with the polymer of the invention, the dimethyl acetamide 
environment was controlled during spraying so that the dimethyl acetamide 
evaporated between the spinneret and the mandrel. Each pass of the 
reciprocating spinneret deposited fibers at an angle on the preceding 
fibers to form a non-woven mat, as shown in FIG. 2, and in which the 
fibers were about 10-20 microns in diameter. These fibers were about 2-5 
times larger in diameter than the diameter of the fibers or strands of the 
triblock polymer previously described in this example. 
Both the triblock polymer and polyether urethane samples which had been so 
prepared were then cut into approximately rectangular samples which were 
about 1 cm by 2 cm and about 0.2 mm thick. Some of each of these 
respective samples were then implanted subcutaneously in the back of dogs. 
After one month in the case of the polyether urethane implants and three 
months in the case of the triblock polymer, these samples were removed 
from the dogs. These explants were soaked upon their respective removal in 
a sodium hydroxide (10%) and sodium hypochlorite (7%) solution for 24 
hours to remove ingrown tissue. They were then sputtered with gold and 
examined under a scanning electron microscope. 
FIGS. 2 and 3 are reproductions of photomicrographs of the previously known 
polyether urethane samples at 200.times. under the electron microscope. 
FIG. 2 shows one of the polyether urethane samples which was not 
implanted, and FIG. 3 shows one of the polyether urethane samples which 
was implanted and following the one month explant. It will be readily seen 
when comparing FIGS. 2 and 3, that substantial breaking, cracking and 
degradation of the fibers of the polyether urethane sample has occurred in 
as little as one month during the in vivo implant. 
FIGS. 4 and 5 are reproductions of photomicrographs of the samples of the 
triblock polymer of the invention also at 200.times. under the electron 
microscope. Although FIGS. 2 and 3 and 4 and 5, respectively, do appear to 
be somewhat structurally different from each other in the drawing, this 
difference is not the result of the presence or absence of degradation of 
the polymer. It is simply the difference in the fiber diameters of the 
respective membranes and also the difference in structure between the 
respective membranes and how they were formed on the rotating mandrel, 
i.e. the polyether urethane of FIGS. 2 and 3 having been first formed into 
fibers which are wound upon the rotating mandrel, and the triblock polymer 
of the invention and of FIGS. 4 and 5 having simply been sprayed with an 
air brush directly on the mandrel. 
FIG. 4 shows one of the triblock polymer samples of the invention which was 
not implanted, and FIG. 5 shows one of the triblock polymer samples of the 
invention which was implanted and following the three month explant. It 
will be seen when comparing FIGS. 4 and 5 that the explanted sample 
appears very similar to the non-implanted control sample, and no 
substantial breaking, cracking or degradation of the triblock polymer 
sample of the invention occurred in the three months during the in vivo 
implant, and in contrast to the explanted polyether urethane sample shown 
in FIG. 3. 
It will be understood that the embodiments of the present invention which 
have been described are illustrative of some of the applications of the 
principles of the present invention. Various modifications may be made by 
those skilled in the art without departing from the true spirit and scope 
of the invention.