HIGHLIGHT

From Thermoplastic Elastomers to Designed Biomaterials

JOSEPH P. KENNEDYInstitute of Polymer Science, University of Akron, Akron, Ohio 44325-3909

Received 19 March 2005; accepted 23 March 2005DOI: 10.1002/pola.20844Published online in Wiley InterScience (www.interscience.wiley.com).

ABSTRACT: This highlight is

about my metamorphosis from a

cationic polymerization chemist to

a biomaterialist (no pun intended)

and some of the main events on the

road. My earlier career faded away

with the discovery of living cati-

onic polymerizations, chronicled in

my 1999 highlight, but it also put

me on the road to designed bioma-

terials. My new career started with,

and still focuses on, the creation of

new polymeric architectures,

mainly by cationic techniques, for

toughened bone cements, injectable

intervertebral discs, nonclogging

artificial blood vessels, and amphi-

philic networks for controlled drug

delivery and immunoisolatory

membranes. The enormous com-

plexities of immunoisolation of

pancreatic islets are now center

stage, and lately we have been

using all kinds of techniques to

make unique membranes to correct

type 1 diabetes. VVC 2005 Wiley Periodi-

cals, Inc. J Polym Sci Part A: Polym Chem

43: 2951–2963, 2005

Keywords: amphiphilic net-

works; biocompatibility; biomate-

rials; membranes; polyisobutylene

Joseph P. Kennedy started his university career in his native city,

Budapest. Just before graduating from the university, he was removed

by the communist administration because of his bourgeois origin. He

escaped to Vienna, where he received his Ph.D. in biochemistry in

1954, and subsequently he did postdoctoral work in Paris and Mon-

treal (1954–1957). He came to the United States in 1957 and became

an industrial polymer researcher, first with Celanese and then with

Exxon. In 1961, he received an M.B.A. at Rutgers. He resumed his

academic career at the University of Akron in 1970, where he is still

carrying out research as a Distinguished Professor of Polymer Science

and Chemistry. Kennedy’s main interest lies in ionic (particularly

cationic) polymerizations and, for the last 15 years, in designedJOSEPH P. KENNEDY

Correspondence to: J. P. Kennedy (E-mail: kennedy@polymer.uakron.edu)

Journal of Polymer Science: Part A: Polymer Chemistry, Vol. 43, 2951–2963 (2005)VVC 2005 Wiley Periodicals, Inc.

2951

PREAMBLE

Some 3 years ago, one of our editors (Virgil Percec) sug-

gested that I write a highlight about my current work. I

accepted his invitation in principle . . . but we did not set

a deadline. During the intervening months, I experienced

his gentle prodding to hunker down and write science. I

appreciated the nudging because otherwise I would have

done more pleasant things, such as playing with my pre-

cocious grandchildren. Then, an invitation to give a lec-

ture at a university gave me the needed activation

energy. As I was mulling over ideas for a lecture to a

sophisticated polymer audience, it occurred to me that

my scientific metamorphosis, which led from living cati-

onic polymerizations to thermoplastic elastomers (TPEs)

and thence to biomaterials, would be a rather interesting

tale to tell. Then it hit me: this was exactly what my edi-

tor wanted from me. Thus was born this highlight.

INTRODUCTION

Brief Background with a Fast Forwardto the Present

So where should I begin? I think I will start where I

left off late in 1998, when I completed my earlier

highlight, ‘‘Living Cationic Polymerization of Olefins:

How Did the Discovery Come About?’’, a they-said-it-

cannot-be-done-therefore-we-did-it kind of a piece.1

The experimentation for that discovery was carried out

in 1984, and the discovery was complete in 1985, but

we delayed publication because of wrangling with the

U.S. Patent Office (more on this in ref. 2). Ultimately,

this breakthrough led to the synthesis of a new TPE,

poly(styrene-b-isobutylene-b-styrene), by my group,

which included Gabor Kaszas, Judit Puskas, and Bill

Hager, whose Ph.D. thesis was on this subject.3,4

Toward the end of my highlight, in the section ‘‘A

Glimpse into the Future’’, among the many possibil-

ities that I foresaw for living cationic polymerizations,

I predicted a bright future for polyisobutylene (PIB)-

based TPEs. In view of their unique combination of

properties, such as good mechanical properties, soft-

ness, outstanding chemical resistance, and barrier prop-

erties, combined with excellent processing characteris-

tics and recyclability, I expected these TPEs to be suit-

able for new hot-melt adhesives, sealants, blending

agents, and so forth. I was, therefore, not surprised

when I learnt that Kaneka, Inc., of Japan had very

recently commercialized a series of brand-new TPEs

(named SIBS for styrene–isobutylene–styrene) based

on our triblock polymer. I was, however, astonished

when I read that at the same international trade show,

K2004 in Dusseldorf, Germany, where Kaneka had

brought SIBS to market, that BASF, the German chem-

ical giant, had also introduced an essentially identical

PIB-based triblock under the trade name Oppanol IBS.

Let me also highlight a surprising important new

application of these triblocks. Briefly, in March 2004,

the U.S. Food and Drug Administration (FDA) approved

the use of these triblocks as drug-eluting coatings of cor-

onary stents, and Boston Scientific Co. has started mar-

keting these devices in the United States under the trade

name Taxus Express Paclitaxel-Eluting Stents. These

drug-eluting stents were available in Europe a year

earlier.

Stents are small, expandable, tubular metal scaffolds

that, when inserted into dangerously occluded coronary

arteries, pry them open and restore blood flow. It was

observed, however, that after approximately 6 months of

stent placement, life-threatening restenosis could occur

in approximately 30% of patients. Therefore, stents

coated with a polymer carrying a restenosis-preventing

drug (e.g., paclitaxel) were a significant innovation. The

PIB-based triblock is eminently suitable as a drug-

eluting stent coating and satisfies the many requirements

of such a demanding application, including bio- and

hemocompatibility, controlled drug release, sterilization,

confluent stent coating, and satisfactory mechanical

properties of the coating that withstand the stresses dur-

ing stent insertion and expansion without the integrity of

the coating being compromised. Importantly, stents can

be placed with relative ease, and this obviates the risks

of major invasive bypass surgery. It is no wonder that

stents, particularly drug-eluting stents, are revolutioniz-

biomaterials. He has written three books and almost 700 publications

and has over 90 issued U.S. patents, some of them in commercial pro-

duction. He has received many awards, including the two premier

international polymer awards of the American Chemical Society (Pol-

ymer Chemistry and Applied Polymer Science). For obvious reasons,

he derives his greatest satisfaction from the honorary doctorate

awarded by the best science university in Hungary (1989) and by his

election as a member of the Hungarian Academy of Sciences (1993).

2952 J. POLYM. SCI. PART A: POLYM. CHEM.: VOL. 43 (2005)

ing coronary surgery, and because of their use, the num-

ber of bypass operations has plummeted some 85%. Cur-

rently, over 1 million stents are being implanted yearly

worldwide.

Start of the Journey

With living cationic polymerizations firmly in our hands

in 1985, we wondered: Which direction(s) among the

many exciting possibilities toward heretofore unattain-

able polymeric architectures should we take? Where lay

fame and fortune?

I saw only two choices: polymers for electronic

applications or biomaterials, that is, areas in which

function was critical and cost was relatively unimpor-

tant. I remember discussing these things with my group

of students, postdocs, and visiting scientists during

group meetings and individually in 1985 and 1986. In

hindsight, it was an easy call for me because of my

fascination with things biological; at heart, I remained

a biochemist (Ph.D. in enzymology, University of

Vienna, 1954). Thus, I decided to apply our new preci-

sion synthesis technique for the creation of polymeric

biomaterials, that is, polymers that can be implanted

into a living organism to assist or replace tissues or

organs.

In the late 1980s, biomaterials were developed

mostly by physicians (!) frustrated by the lack of suit-

able materials for clinical applications. These adventur-

ous M.D. innovators developed the first polymeric bio-

materials from polymers that they found on the shelf

put there by industrial scientists who used them in

more mundane industrial applications; thus arose surgi-

cal sutures from nylon yarns developed for ladies’

hose, implantable silicones from electrical insulators,

bone cements and ophthalmic materials (lenses, etc.)

from Plexiglas [poly(methyl methacrylate) (PMMA)],

indwelling polyurethane tubes and catheters from furni-

ture upholstery, vascular grafts and surgical meshes

from Dacron polyester no-iron fabric, orthopedic

implants from polyethylene electrical insulating cables,

and so forth.

I thought living cationic polymerizations could pro-

vide an excellent route to new molecular architectures

expressly designed for biomaterials. However, I had to

find the right need and match it with the right polymer.

So I started to buttonhole biologists and all kinds of

medical professionals willing to talk to with me, a poly-

mer scientist searching for clinical needs that could be

satisfied with polymers made with his newfangled tech-

nique of living cationic polymerization.

In the late 1980s, I thus became a biomaterialist.

DESIGNED BIOMATERIALS

Toughened Bone Cements

Among our first forays into designed biomaterials was

the synthesis of toughened bone cements. Contempo-

rary bone cements are essentially PMMA formulations

designed to secure orthopedic prostheses, such as hip

joints. The cements are prepared in the operating room

by the mixing of the contents of two vials, one con-

taining powdery PMMA, a radioopacifier (BaSO4), and

aninitiator (benzoyl peroxide), and the other containing

liquid methyl methacrylate (MMA) monomer, in which

are dissolved an accelerator (N,N-dimethyl-p-toluidine)and a stabilizer (hydroquinone). Upon the mixing of

these ingredients, the monomer dissolves the PMMA,

and a paste is formed; at the same time, the initiator

starts the complex polymerization/grafting of MMA.

While this ill-defined reaction is in progress and the

paste is still workable, the surgeon positions the paste

around the prosthesis in the patient. The considerable

heat of in situ MMA polymerization/grafting raises the

temperature of the surrounding tissue to 75–80 8C,and the heat kills adjacent living tissue (no wonder the

patient is under general anesthesia). Even more trou-

bling, approximately 10% of the MMA remains un-

reacted and is released into the patient, and this may

cause adverse reactions. (I doubt that the FDA would

have allowed PMMA in orthopedic surgery as it is

practiced today, had the FDA been in existence at the

time bone cements came to be used, but this is another

story.) After the polymerization is over, the paste sol-

idifies into a heterogeneous, porous, brittle glass, which

fixes the metal prosthesis to the bone. The residual

MMA may also act as a softener of the solid PMMA

that is formed. The polymer scientist immediately

understands that in view of the nature of the three

materials and two interfaces involved (bone/PMMA/

metal), true chemical bonding cannot occur because

PMMA is incompatible with both the bone and the

metal. Rather, I think fixation is due to mechanical

interlocking of the soft MMA/PMMA paste in the

microscopic cavities of the bone and metal surfaces.

(Essentially the same mechanism is at work when

icicles form on the roof.)

Other major shortcomings of PMMA bone cements

are brittleness and consequent debris formation. Revi-

sion surgery is needed in approximately 20% of the

cases 10 years after implantation,5 and cement fracture

is also frequently observed. Hip repair frequently fails

because of prosthesis loosening at the bone/cement

interface; because PMMA is a glass [glass-transition

temperature (Tg) � 105 8C], bone cements will sooner

or later crack under a constant and dynamic load.

HIGHLIGHT 2953

The toughening of brittle plastics has been thoroughly

studied by polymer engineers, and remedies for this

problem exist. One of the obvious things that a polymer

scientist can do to combat the brittleness of PMMA is to

covalently incorporate into the glassy matrix a dispersed

elastomer phase. We hypothesized that the toughening

of PMMA could be achieved by crosslinking with a tri-

arm star methacrylate (MA)–telechelic PIB (Tg ��70 8C) crosslinker. By this copolymerization, the PIB

rubber becomes covalently bound to the continuous

glassy PMMA phase, and because this nonpolar rubber

is incompatible with the polar PMMA, the rubber will

form a desirable dispersed phase in the continuous

glassy matrix. Figure 1 illuminates the concept and

shows the structure of the key ingredient, the rubbery

F-(PIB–MA)3. This crosslinking agent can be obtained

only by living cationic polymerization.6

We synthesized F-(PIB–MA)3’s of various molecular

weights and added them in various proportions to a com-

mercial bone cement formulation. Extensive synthesis

and characterization showed that cements containing

9.2% PIB with a number-average molecular weight (Mn)

of 18,000 g/mol exhibited particularly desirable overall

properties.7–10 Engineering tests indicated improved

flexural strength, maximum deflection, fracture tough-

ness, and fatigue crack propagation rate with respect to a

commercial product (Zimmer Regular Bone Cement). An

enterprising orthopedic surgeon tried this cement in dogs,

but he died and the experiment was discontinued before

data could be collected.

I do not know why this lead was not further devel-

oped by industrial researchers.

Cyanoacrylate (CNA)-Tipped PIB forIntervertebral Disc

The pain from a herniated (slipped) intervertebral disc

can be excruciating. To remove the source of the trauma,

the orthopedic surgeon excises the offending disc tissue

(by laminectomy, chemonucleolysis, etc.); however, the

tissue is not replaced, and the loads that were distributed

by the excised disc must be taken over by other tissues.

We hypothesized that this rather unsatisfactory sit-

uation could be remedied by the replacement of the

excised tissue with an elastomer whose viscoelastic

characteristics are similar to those of the disc. It

appeared feasible to inject a liquid prepolymer, which

would rapidly polymerize to a rubber with the needed

viscoelastic properties, into the cavity left behind by

surgery. I thought we had the right replacement mate-

rial: PIB carrying CNA groups. I knew we could pre-

pare linear or star-shaped PIB prepolymers and fit

them with reactive CNA groups. Figure 2 shows the

structures envisioned, together with Super Glue (in

which R is a small substituent, i.e., methyl, ethyl, or

butyl). We speculated that prepolymers of suitable

molecular weights and compositions could be injected

where they were needed by a prefilled sterilized

syringe containing the prepolymers. The principal con-

stituent, PIB, a hydrophobic, inert, biocompatible, bio-

stable, and oxidatively stable rubber, fitted with CNA

end groups, would rapidly polymerize upon contact with

living tissue (nucleic acids, proteins with ��NH2, ��OH

groups). Because the polymerization would be induced

by nucleophilic functional groups of the surrounding tis-

sue, the PIB would be covalently anchored to the tissue,

and the polymer would form only in the cavity where

the liquid prepolymer was injected. Leakage of the poly-

mer into the surrounding tissue could not occur. Further-Figure 2. Various PIB prepolymers fitted with CNA end

groups (note their similarity to Super Glue).

Figure 1. Toughening of PMMA by the covalent incorporation of PIB rubber.

2954 J. POLYM. SCI. PART A: POLYM. CHEM.: VOL. 43 (2005)

more, the long PIB chains would envelop and sequester

the ��CH2��C(CN)(COOPIB)�� main chain from the

hydrolytic and enzymatic degradation that occur with

small conventional CNAs.11

We demonstrated that the envisioned prepolymers could

be readily synthesized and that they polymerized upon con-

tact with living tissue (blood and egg yolk).11 The poly-

merization and crosslinking rates could be controlled by

the use of various molecular weight products and by the

copolymerization of linear and star-shaped prepolymers

(see Fig. 2). I still believe that these PIB-based macromo-

nomers would go a long way toward satisfying the require-

ments for disc replacement or other applications.

Smart Amphiphilic Hydrogels

This project started as a fundamental inquiry. It appeared

to me that amphiphilic networks, that is, networks con-

sisting of random hydrophilic and hydrophobic strands,

would exhibit some unexpected properties, and an explo-

ration of their synthesis and basic properties would be

worthy of modest experimentation. Little did we imagine

that this project would spawn drug delivery devices, arti-

ficial blood vessels, and immunoisolatory membranes.

However, first we should cover some basics.

Amphiphilic networks (aptly named conetworks by B.

Ivan) are random, bicontinuous assemblages of hydro-

philic/hydrophobic chain segments that swell in both

water and hydrocarbons. Because they swell in water,

they are hydrogels. Figure 3 helps to visualize amphi-

philic networks and emphasizes the conformational

changes that these networks undergo upon changes in the

medium (environmentally responsive networks). In tetra-

hydrofuran (an amphiphilic solvent), both chain elements

are solvated and expanded, and the entire network swells.

In water, the hydrophilic chains are extended, whereas

the hydrophobic chains collapse to coils. In hydrocarbons,

the opposite occurs: the hydrophilic chains are extended,

whereas the hydrophilic chains collapse. In the latter two

cases, the collapsed coils want to precipitate, but the

covalently bound solvated chains will not let them.12 Sur-

face analytical techniques by contact angle, atomic force

microscopy, X-ray photoelectron spectroscopy, and sur-

face atomic ratios (O/C and N/C) have indicated that the

surfaces of amphiphilic networks are highly mobile; spe-

cifically, dry surfaces rapidly reorganize upon exposure to

water.13 These networks are able to adopt different sur-

face conformations in different environments to increase

surface accommodation with the milieu and thereby mini-

mize the total free energy of the system. Thus, amphi-

philic networks are smart: they change their microstruc-

ture (morphology) with the medium. This chameleon-like

change may be important for accommodating complex

biological systems consisting of all kinds of molecules

and constituents (i.e., for biocompatibility).

PIB-Based Amphiphilic Networks: ControlledRelease, Artificial Arteries, and the FirstGeneration of Immunoisolatory Membranes

Our first generation of amphiphilic networks was pre-

pared by the free-radical solution copolymerization of

hydrophilic monomers [N,N-dimethyl acrylamide

(DMAAm), 2-hydroxyethyl methacrylate (HEMA), 2-

(dimethylamino)-ethyl methacrylate (DMAEMA), and

sulfoethyl methacrylate (SEMA)] with hydrophobic

Figure 3. Reorganization of amphiphilic networks in various media (THF

¼ tetrahydrofuran, HC ¼ hydrocarbon).

HIGHLIGHT 2955

crosslinkers (MA–telechelic PIBs).6,14–19 Figure 4

shows the formulas. The MA-capped PIB crosslin-

kers can be prepared only by living isobutylene poly-

merization followed by end-group functionalization.6

Figure 5 outlines the synthesis of one of the simplest

amphiphilic network by the copolymerization of

DMAAm with MA–PIB–MA.

By our technique, we can control the overall composi-

tion of the networks and the molecular weights between

crosslinking points (Mc’s). Because these networks have

hydrophilic and hydrophobic segments, they exhibit two

Mc’s: Mc,HI (the molecular weight of the hydrophilic seg-

ment between crosslinking points) and Mc,HO (the molec-

ular weight of the hydrophobic segment). Mc,HI controls

the pore (or mesh) dimensions of the networks in water

(i.e., when implanted in an organism). The pore dimen-

sions of the networks can be regulated by the concentra-

tion of the hydrophilic monomer with respect to the

hydrophobic crosslinking agent.

Because the chain elements in amphiphilic networks

are chemically very different and thermodynamically

incompatible, they are two-phase systems with two

Tg’s.20,21 Transmission electron microscopy of a typical

amphiphilic network has shown 20–50-A-diameter bicon-

tinuous domains with a salt-and-pepper morphology.

Controlled Release

We studied the kinetics of swelling of amphiphilic net-

works, using both water and n-heptane.15–18,21 We

found that, with an increase in the PIB content, the

rate of water uptake decreased, whereas that of n-hep-

tane increased; the opposite was found with an

increase in the content of the hydrophilic component

(i.e., the water uptake increased). Subsequently, we

studied the out-diffusion of select water-soluble model

drugs (e.g., folic acid) from drug-loaded amphiphilic

networks,14 and we found that the release rate changes

with the nature and concentration of the hydrophilic

constituent and the molecular weight of the PIB cross-

linker. Interestingly, the diffusion coefficient n(obtained from Mt/M? ¼ ktn, where Mt is the amount

of drug released at time t, M? is the amount of drug

loaded, and k is a constant), determined for several

networks, fell in the 0.7–0.8 range. If n ¼ 0, diffusion

is controlled by polymer relaxation (zero-order

release), whereas n ¼ 0.5 indicates conventional or

Fickian diffusion.22 The experimental values for sev-

eral of our systems suggest anomalous transport, that

is, the presence of another process besides passive dif-

fusion.

A series of amphiphilic networks containing poly(2-

sulfoethyl methacrylate) and MA–PIB–MA strands

were prepared and characterized by thermal, spectro-

scopic, mechanical, and swelling experiments18 Their

swelling followed non-Fickian kinetics in both water

and n-heptane. Networks with higher ionic contents

showed rapid and reversible swelling or deswelling

upon changes in the pH (in the 2–12 range) of the

medium.

The discovery that networks containing approxi-

mately 50/50 DMAEMA/PIB (Mn,PIB ¼ 10,000 g/mol)

exhibited excellent biocompatibility and biostability in

rats14 was of decisive importance for the future course

Figure 4. Typical starting materials for the synthesis of first-generation amphiphilic

networks.

2956 J. POLYM. SCI. PART A: POLYM. CHEM.: VOL. 43 (2005)

of our research. We found that certain well-defined

amphiphilic networks integrated well with tissue and

sowed minimal bacterial contamination, no edema, and

virtually no fibrosis and adhesion (even less than the

polyethylene negative control).14 Indeed some of our

materials exhibited better biocompatibility than the

negative controls. In cell culture and protein tests, the

numbers of cells and the total protein on the amphi-

philic networks were similar to those of negative con-

trols (polyethylene, silicone rubber, and glass), and this

indicated no toxic response. Cell adhesion and antiad-

hesion experiments with human monocytes showed

monocyte adhesion inhibition for various amphiphilic

networks and glass (negative control) with respect to

polystyrene (positive control). Further quantitative pro-

tein adsorption by radioimmunoassay showed that

amphiphilic networks made with DMAAm or HEMA

with 50% PIB adsorbed from human plasma less fibrino-

gen, Hageman factor, and albumin than glass, silicone

rubber, or polyethylene.23 These analyses, together with

blood cell counts, suggest that select amphiphilic net-

works are well accepted in vivo; that is, they are biocom-

patible.

Overall, these observations indicate reduced protein

adsorption, with a significant reduction of Hageman

factor and fibrinogen adsorption. Together with human

monocyte adsorption data, these studies indicate that

select amphiphilic networks are biologically compati-

ble at blood-contacting surfaces.

Artificial Arteries

Coronary artery obstruction (stenosis) is life-threatening.

A clinical solution to coronary stenosis is major surgery

in which the occluded artery is replaced by another of

the patient’s native blood vessels (typically, the saphe-

nous vein or mammary artery). Unfortunately, 3–

5-mm-diameter blood vessels, such as the coronary

artery, cannot be replaced by synthetic polymer tubes

[e.g., polyurethane, polyethylene, polyester, or poly(tetra-

fluoroethylene)] because after graft placement they

quickly (within a few hours) become occluded by platelet

and fibrin deposition.24

In view of the bio- and hemocompatibility of select

amphiphilic networks, we set out to explore whether a

narrow-caliber tube made of our materials could be

used for coronary artery replacement; specifically, we

wondered how many minutes per hour a 3-mm-caliber

amphiphilic tube would remain free of stenosis with

blood circulating through it.

We built an apparatus by which we circulated fresh

rat blood (no anticoagulant) at 37 8C through 4–5-cm-

long amphiphilic tube sections made by the copolymer-

ization of DMAEMA/MA–PIB–MA in a rotating glass

cylinder [the rotational copolymerization technique has

been described elsewhere].25 A peristaltic pump was

used to simulate the pumping action of the heart; that is,

the tube pulsated during blood circulation. Experiments

were carried out for 60 min, during which time no trace

of platelet deposition in the amphiphilic tube was

observed! Under similar conditions, the negative control,

a plasticized poly(vinyl chloride) (Tygon) tube, showed

significant platelet deposition.26 The results obtained by

this dynamic hemocompatibility test corroborated those

obtained by protein and cell adsorption experiments (as

previously discussed).

Encouraged by these findings, we carried out a

more demanding dynamic hemocompatibility test

jointly with researchers at the University of Wiscon-

sin.27 By the use of an artificial heart device, these

researchers circulated 60 mL of heparinized (3 units)

Figure 5. Synthesis of an amphiphilic network by the copolymerization of DMAAm

with MA–PIB–MA.

HIGHLIGHT 2957

fresh bovine blood containing 111In labeled platelets

through various 3-mm-caliber tubes for 120 min. We

compared three of our amphiphilic network tubes made

by the copolymerization of DMAAm, DMAEMA, or

HEMA plus the MA–PIB–MA crosslinker (as previ-

ously discussed), with a polyurethane tube as a nega-

tive control. Thrombosis was quantitated by g counting

of the platelets. Although the polyurethane control

gave more than 80,000 counts per minute (cpm), which

indicated relatively high platelet deposition, our amphi-

philic tubes showed far less platelet deposition, approx-

imately 1000 cpm, with the tube made with DMAAm

being the best (�200 cpm, average of three experi-

ments).

Jointly with Dutch biomaterials scientists, we

studied platelet adhesion and blood coagulation activa-

tion of select amphiphilic networks in reference to

polyethylene and poly(vinyl chloride) negative controls

in vitro, that is, polymers commonly used in blood-

contacting applications.28 Networks of DMAAm/PIB

exhibited lower thrombogenicity than polyethylene and

poly(vinyl chloride) and significantly lower platelet

adhesion than poly(vinyl chloride), the reference with

lowest thrombogenicity.

Although the results of these bio- and hemocompati-

bility experiments were encouraging, the mechanical

properties of our tubes were insufficient for vascular

implantation; microsuturing was deemed impossible

because of the poor puncture and tear strengths of our

water-swollen tubes. To overcome this problem, we

hypothesized that thromboresistant narrow-caliber

tubes could be obtained by the coating of 3-mm-diam-

eter expanded poly(tetrafluoroethylene) (ePTFE) tubes

with amphiphilic networks. Although large-caliber

(>6 mm) ePTFE tubes (Goretex and Impra) are exten-

sively used in vascular surgery, narrow-caliber tubes

cannot be used because they are thrombogenic. Thus,

in preparation for in vivo implantation experiments, we

coated the surfaces of 3-mm-diameter Goretex tubes

with one of our promising amphiphilic formulations

(50/50 DMAEMA/MA–PIB–MA, with Mn,PIB ¼ 4500

g/mol). Scanning electron microscopy (SEM) showed

smooth, featureless surfaces, whereas the surfaces of

uncoated Goretex tubes displayed the characteristic

striated, fibrillar morphology of Goretex. Interestingly,

water wetted the coated tubes even after their coating

was manually peeled off. Evidently, the amphiphilic

coating penetrated and remained in the interstices of

the porous ePTFE and rendered it amphiphilic.

We grafted approximately 2-cm sections of 3-mm-

caliber uncoated (control) and amphiphilic-polymer-

coated Goretex tubes into the aortas of two rabbits.29

Coating was carried out by the immersion of Goretex

tubes into polymerizing amphiphilic charges. Heparin

was administered to prevent blood clotting. After 7 days,

the animals were sacrificed, and the junctions of the

graft and native aortas were examined by SEM. The

luminal surface of the amphiphilic network-coated graft

showed a featureless confluent coating with no platelet

or fibrin deposition or cellular debris; in contrast, the

surfaces of the uncoated control were extensively cov-

ered with thrombus. Figure 6 shows representative SEM

pictures.

Significantly, we noted minor areas of discontinuities

in the homogeneous coats, which exhibited the charac-

teristic fibrillar morphology of Goretex (see Fig. 6,

right). These minor discontinuities were probably due to

imperfect coat deposition (entrapped air bubbles?). The

absence of cellular debris over these areas suggests that

even an ultrathin, SEM-invisible amphiphilic-polymer

layer deposited on the surface of Goretex is sufficient to

prevent thrombus formation. Recent follow-up experi-

Figure 6. Luminal surfaces of 3.0-mm Goretex tubes after 7 days of implantation in

a rabbit infrarenal aorta. The left image shows an uncoated tube (control), and the right

image shows a 3.00-mm Goretex tube coated with an amphiphilic network {�50/50

poly[2-(dimethylamino)-ethyl methacrylate]/PIB}. The magnification for both images is

600�.

2958 J. POLYM. SCI. PART A: POLYM. CHEM.: VOL. 43 (2005)

ments carried out jointly with Dr. K. Ouriel et al. of the

Cleveland Clinic Foundation have been focused on

improving the coating of ePTFE (Impra) by the use of a

rapidly rotating coating device (to be published).

In conclusion, these preliminary studies show that

3.0-mm-diameter amphiphilic-polymer-coated ePTFE

can be used as an aortic conduit in the rabbit model, that

the amphiphilic-polymer coat does not change the han-

dling characteristics of the ePTFE tube, and that amphi-

philic-network-coated ePTFE exhibit significantly less

thrombus deposition than uncoated tubes. Efforts are in

progress to follow up this lead.

First Generation of Immunoisolatory Membranes

Numerous discussions with clinicians, biomaterial

researchers, and immunologists have convinced me

that there is a need for novel immunoisolatory mem-

branes (the language barrier between our disciplines

sometimes hampered our discussions). Immunoisola-

tory membranes are used to encapsulate and transplant

living tissue from a donor to a host organism (xeno-

transplantation). Such membranes protect the trans-

plants from being destroyed by the host’s immune sys-

tem, thereby eliminating costly and dangerous immu-

nosuppressive drug therapy.30–32 Immunoisolation of,

for instance, living pig pancreatic islets (beta cells)

into diabetic humans may correct diabetes.

Islet encapsulation/transplantation has been investi-

gated by many researchers employing many kinds of

materials and methods. A thorough analysis of the

requirements gleaned from the scientific and patent lit-

erature has led us to conclude that an ideal immunoiso-

latory membrane, clinically useful for a bioartificial

pancreas, must simultaneously satisfy all the following

biological, chemical, physical, mechanical, surface, and

processing properties:

1. Biocompatibility with the host (human) and

guest (e.g., porcine islets).

2. Hemocompatibility.

3. Biostability.

4. Smooth, slippery, nonclogging, nonfouling,

avascular, and nonthrombogenic or, in other

words, immunologically invisible surfaces.

5. Controlled semipermeability: precisely

designed/defined pore dimensions (molecular

weight cutoff ranges) that allow the passage of

aqueous solutions of nutrients and biologically

active molecules (insulin) and the exit of meta-

bolic wastes but exclude antibodies and white

blood cells.

6. Physiologically satisfactory bidirectional fluxes

of glucose, insulin, nutrients, and metabolites.

7. Thin membrane walls (micrometer range) to

minimize diffusion paths.

8. Satisfactory mechanical properties (strength, modu-

lus, elongation, and fatigue) for the implantation

and explantation of large numbers (�8 � 105) of

islets.

9. Highest and rapid oxygen and water transport.

10. All the above properties to be maintained for

long times (6–12 months).

11. Simple and efficient membrane synthesis.

12. Easily manufactured into sealable and preferen-

tially transparent tubes of pouches of well-

defined volumes.

13. Ease of implantation and explantation.

14. Sterilizability.

15. All this for a reasonable cost.

Implantable membranes examined by others for the

correction of diabetes have several but not all of these

characteristics; for example, many researchers use algi-

nates or siloxane gels to microencapsulate individual

islets. One of the fundamental disadvantages of micro-

encapsulation by hydrogels is that the immunopro-

tected tissue cannot be reliably or completely retrieved.

The other is that hydrogels have very poor oxygen per-

meability (water is a barrier Io oxygen diffision).

With this analysis in mind, we set out to synthesize

immunoisolatory membranes from our amphiphilic pol-

ymers with the required biological and mechanical

properties. We thought to achieve semipermeability

control (see items 5–8) by regulating the length of

Mc,HI and by the overall hydrophilic/hydrophobic com-

position of the membranes. The molecular weight cut-

off range (pore size control) was to be achieved by the

regulation of the length of the hydrophilic and hydro-

phobic segments to allow the rapid countercurrent dif-

fusion of glucose and insulin, but the membranes were

to be impermeable to large proteins of the immune

system [e.g., immunoglobulin G (IgG)].

Systematic experimentation showed that amphiphilic

membranes containing approximately 50/50 poly(N,N-dimethyl acrylamide)/PIB with Mc,HI � 4500 g/mol

had semipermeability and diffusion rates suitable for

the immunoisolation of pancreatic islets.33 These mem-

branes allowed the countercurrent diffusion of glucose

and insulin (Mn ¼ 180 and 5700 g/mol, respectively)

but prevented the diffusion of albumin (Mn � 66,000

g/mol), and the diffusion rates (fluxes) of glucose and

insulin were deemed appropriate for islet immunoisola-

tion.33 Subsequently, we determined that pig islets

placed in such semipermeable amphiphilic-polymer tu-

bules could be kept viable in tissue culture for at least

4 months and that encapsulated islets produced insulin

upon glucose challenge.34 Importantly, we also demon-

HIGHLIGHT 2959

strated that a diabetic rat, fitted subcutaneously with

our bioartificial pancreas (i.e., an amphiphilic tubule

containing pig islets), started to produce insulin and

that the glucose concentration in the blood of the rat

decreased markedly. When the bioartificial pancreas

was removed, the rat became diabetic again. In these

experiments, the rat was its own control! Figure 7

summarizes the results of this experiment.35

Polydimethylsiloxane (PDMS)-Based AmphiphilicNetworks: The Second Generation ofImmunoisolatory Membranes

There were three compelling reasons that we decided

to advance from PIB-based immunoisolatory mem-

branes to PDMS-based immunoisolatory membranes:

(1) to simplify membrane synthesis, (2) to enhance

oxygen permeability, and (3) to obtain uniform pore

dimensions.

Let me explain:

1. Our first generation of amphiphilic membranes

obtained by the free-radical copolymerization of

acrylates yielded very promising products; how-

ever, the synthesis of the PIB-based crosslinking

agents needs considerable expertise, and the rad-

ical copolymerization mechanism is inherently

ill defined. I was therefore constantly searching

to find a less demanding synthetic procedure,

and I was pleased when we found that excellent

amphiphilic membranes could also be made by

simple hydrosilation/condensation (discussed later).

2. A thorough search of the scientific and patent

literature, including web pages of companies

engaged in immunoisolation, indicated that one

of the crucial requirements of immunoisolatory

membranes was oxygen permeability. Although

native pancreatic islets receive oxygen via

blood circulation, immunoisolated islets receive

oxygen (and eliminate waste) only via passive

diffusion through the encapsulating membrane.

Obviously, then, to ensure adequate oxygen

supply to the islets, the membranes must be as

thin as feasible and as friendly to oxygen as

possible. We reasoned that the oxygen perme-

ability of PIB-based amphiphilic membranes,

that is, membranes through which oxygen

transport can occur only via water in the

hydrophilic domains (PIB is a barrier to oxy-

gen diffusion), could be vastly enhanced by the

substitution of PDMS for PIB. Moreover, oxy-

gen diffusion only via the hydrophilic domains

was thought to be inadequate because of the

low solubility of oxygen in water (200–300

mg/mL). Thus, we turned to PDMS, the most

oxygen-permeable rubber, whose oxygen per-

meability is far superior to that of water. It is

true that PDMS is somewhat weaker than PIB,

but we had ideas how to increase the strength

of PDMS-based membranes, should this

become an issue.

3. Uniform pore dimensions cannot be obtained with

networks made by a random free-radical copoly-

merization of hydrophilic monomers with MA-

functionalized PIBs, andMc,HI will always exhibit a

rather broad length distribution [weight-average

molecular weight/number-average molecular

weight (Mw/Mn) � 2.0]. Such a broad distribution,

however, is dangerous because in the presence of

even a minute fraction of larger pores, some immu-

noproteins may traverse the membrane and thus

compromise the encapsulated living tissue. Thus, to

obtain uniform pore dimensions, we had to redesign

our membranes.

So how should we proceed?

To obtain the highest oxygen permeability and pore

uniformity, we decided to use precisely defined hydro-

philic poly(ethylene glycol) (PEG) and hydrophobic

PDMS starting materials, both with the narrowest possi-

ble molecular weight distributions (Mw/Mn � 1.0). Spe-

cifically, we thought that random combinations of pre-

cise-length (molecular weight) PEG and PDMS chains

would yield amphiphilic networks with uniform, pre-

cisely defined, and controllable pore dimensions. How-

ever, how can we combine the incompatible PEG and

PDMS segments into a network?

Figure 7. Graph showing that a bioartificial pancreas

[�2000 porcine islets immunoisolated by an amphi-

philic network (�50/50 DMAAm/PIB) subcutaneously

implanted into a diabetic (streptozocin) rat] improves

hyperglycemia.

2960 J. POLYM. SCI. PART A: POLYM. CHEM.: VOL. 43 (2005)

At this point, fate came to our aid: Fortuitously,

about this time, one of my students, P. Kurian, investi-

gated reactions of cyclosiloxanes for another project

and found that pentamethylcyclopentasiloxane (D5H)

rapidly polycondenses to poly(pentamethylcyclopentasi-

loxane) (PD5).36 It occurred to me that PD5 was exactly

what we needed because (1) PD5 domains could function

as efficient crosslinkers by the cohydrosilation of olefin–

ditelechelic PEG and –PDMS segments, (2) PD5 would

function as reinforcing domains, and (3) as a bonus the

oxyphilic PD5 domains would provide auxiliary oxygen

channels (in addition to those provided by PDMS). In

other words, the PD5 domains/phases in our networks

would perform triple duty by providing crosslinking,

reinforcement, and supplemental oxygen channels.36

Figure 8 helps to visualize the synthetic strategy and

micromorphology of an amphiphilic network consisting of

PEG and PDMS segments crosslinked and reinforced by

PD5 domains. The synthesis involves the cohydrosilation

of A–PEG–A and V–PDMS–V mixtures by D5H, followed

by simultaneous water-mediated oxidation of the excess

SiH groups to SiOH groups and in situ polycondensation

to PD5 domains.37 A–PEG–A is easily obtained from com-

mercially available HO–PEG–OH,37 and the other two

starting materials are inexpensive commercial products.

The relative concentrations of the three constituents (PEG/

PD5/PDMS) control the overall membrane composition,

which in turn controls the overall membrane properties.

The crosslink density and stiffness of the networks

increase with the D5H concentration, and the concen-

tration of water controls the rate and extent of cross-

linking. It did not take us long to develop suitable

cohydrosilation, oxidation, and polycondensation con-

ditions and to obtain membranes with excellent

mechanical properties (>5 MPa strength and 500%

elongation) appropriate for immunoisolation.38,39

Having established synthesis simplicity and versatil-

ity, we prepared and characterized series of amphi-

philic membranes with various Mc’s, PEGs (in the

4600–20,000 g/mol range), and various PEG/PD5/

Figure 9. Demonstration of the selective permeabil-

ity of a PEG/PD5/PDMS membrane. The positive con-

trol was prepared by the incubation of an aliquot of

rabbit IgG with protein A Sepharose beads followed by

a nonspecific IgG to block any unused IgG binding

sites and then with FITC-goat-anti-rabbit IgG. The

negative control was obtained by the incubation of pro-

tein A Sepharose beads with the blocking IgG and the

FITC-goat-anti-rabbit IgG. Fluorescence was obtained

after 120 h of diffusion. The donor chamber was

loaded with 0.2 mg/mL IgG, and samples were taken

after 120 h and incubated first with protein A Sephar-

ose beads and then with FITC-goat-anti-rabbit IgG.

Figure 8. Synthesis scheme of a tricomponent amphiphilic network.

HIGHLIGHT 2961

PDMS compositions.39 Surface analyses indicated the

expected rapid conformational rearrangements. The

oxygen permeability increased with the amount of

PDMS in the membrane, and the insulin permeability

was shown to be dependent on the length of Mc,PEG.

The permeability of insulin through a series of mem-

branes with various Mc,PEG’s was determined. Molecu-

lar weight cutoff studies showed controllable semiper-

meability, that is, the diffusion of small proteins

including insulin and the rejection of larger proteins

such as albumin (Mn � 66,000 g/mol). The implanta-

tion of representative PEG/PD5/PDMS membranes in

rats showed minimal response with respect to inflam-

mation, foreign body reaction, tissue ingrowth, and

fibrous capsule formation.39

To further enhance reinforcement, we prepared

amphiphilic networks by cocrosslinking PEG/PDMS

mixtures with polyhedral oligomeric silsesquioxane

(POSS) fitted with eight SiH groups (in lieu of D5H).40

Recently, we demonstrated, by a series of diffusion

experiments, the selective permeability of glucose and

insulin with the simultaneous exclusion of IgG.39 Subse-

quently, we also demonstrated that PEG/PD5/PDMS

membranes do not foul and remain permeable to both

glucose and insulin even after long (1-month) incubation

with IgG. Figure 9 provides details and shows our find-

ings.

In sum, we developed a simple and efficient mem-

brane synthesis method; the bicontinuous membrane

architecture and the oxyphilic PDMS/PD5 domains

provide superior oxygen transport; the precisely

defined PEG segments (in terms of the molecular

weights and molecular weight distributions) yield uni-

form pore dimensions (i.e., well-controlled Mc,HI); and

the membranes exhibit desirable semipermeability and

are biocompatible.

We are on the road to a clinically useful bioartificial

pancreas.

Although support has been received from many sources dur-

ing these investigations, the author is mainly indebted to the

National Science Foundation for its continuous financial help

(DMR-8920826 and DMR-0243314).

REFERENCES AND NOTES

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2. Kennedy, J. P.; Ivan, B.Designed Polymers by Car-bocationic Macromolecular Engineering: Theoryand Practice; Hanser: Munich, 1992; p 35.

3. Kennedy, J. P.; Puskas, J. E.; Kaszas, G.; Hager, W.G. (University of Akron). U.S. Patent 4,946,899,1990.

4. Kaszas, G.; Puskas, J. E.; Kennedy, J. P.; Chen, C.C.; Hager, W. G. J Polym Sci Part A: Polym Chem1991, 29, 427.

5. Sutherland, C. J.; Wilde, A. H.; Borden, L. S.;Marks, K. E. J Bone Joint Surg (Am) 1982, 64, 970.

6. Kennedy, J. P.; Hiza, M. J Polym Sci Polym ChemEd 1983, 21, 1033.

7. Kennedy, J. P.; Askew, M. J.; Richard, G. C. (Uni-versity of Akron). U.S. Patent 5,242,983, 1993.

8. Kennedy, J. P.; Richard, G. C. Macromolecules1993, 26, 567.

9. Kyu, T.; Kennedy, J. P.; Richard, C. G. Macromo-lecules 1993, 26, 572.

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11. Kennedy, J. P.; Midha, S.; Gadkari, A. J Macro-mol Sci Chem 1991, 28, 209.

12. Chen, D.; Kennedy, J. P.; Allen, A. J. J MacromolSci Chem 1988, 25, 389.

13. Park, D.; Keszler, B.; Galiatsatos, V.; Kennedy, J.P. Macromolecules 1995, 28, 2595.

14. Chen, D.; Kennedy, J. P.; Kory, M. M.; Ely, D. L.J Biomed Mater Res 1989, 23, 1327.

15. Ivan, B.; Kennedy, J. P.; Mackey, P. W. In Poly-meric Drugs and Delivery Systems; Dunn, R. L.;Ottenbrite, R. M., Eds.; ACS Symposium Series469; American Chemical Society: Washington,DC, 1991; p 194.

16. Ivan, B.; Kennedy, J. P.; Mackey, P. W. In PolymericDrugs and Delivery Systems; Dunn, R. L.; Otten-brite, R. M., Eds.; ACS Symposium Series 469; Amer-ican Chemical Society: Washington, DC, 1991; p 203.

17. Keszler, B.; Kennedy, J. P.; Mackey, P. W. J Con-trolled Release 1993, 25, 115.

18. Keszler, B.; Kennedy, J. P. J Polym Sci Part A:Polym Chem 1994, 32, 3153.

19. Isayeva, I. S.; Yankovski, S. A.; Kennedy, J. P.Polym Bull 2002, 48, 475.

20. Park, D.; Keszler, B.; Galiatsatos, V.; Kennedy, J.P. J Appl Polym Sci 1997, 66, 901.

21. Allen, A.; Kennedy, J. P. Des Mon Polym 1999, 2, 29.22. Pappas, N. A.; Korshmeyer, R. W. In Hydrogels in

Medicine and Pharmacy; Pappas, N. A., Ed.; CRC:Boca Raton, FL, 1987; Vol. 3, p 109.

23. Keszler, B.; Kennedy, J. P.; Ziats, N. P.; Brun-stedt, M. R.; Stack, S.; Yun, J. K.; Anderson, J.M. Polym Bull 1992, 29, 681.

24. Stanley, J. C.; Lindenauer, S. M. In Vascular Sur-gery, a Comprehensive Review, 3rd ed.; Moore, S.W., Ed.; Saunders: Philadelphia, 1991; pp 275–294.

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26. These experiments were carried out jointly withProfessor D. L. Ely, Biology Department, Univer-sity of Akron, 1992.

27. These experiment were carried out jointly withDr. F. Mohammad and Dr. R. J. Jaarsma, Univer-sity of Wisconsin, 1993.

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28. Blezer, R.; Lindhout, T.; Keszler, B.; Kennedy, J.P. Polym Bull 1995, 34, 101.

29. Lewis, R. D.; Wright, D.; Kennedy, J. P.; Keszler,B. Resident Essay Contest, Ohio Chapter, Ameri-can College of Surgeons, May 21, 1993.

30. Lim, F.; Sun, A. M. Science 1980, 210, 908.31. Sefton, M. V.; Stevenson, T. K. Adv Polym Sci

1993, 107, 143.32. Aebischer, P.; Goddard, M.; Signore, A. P.; Timp-

son, R. L. Exp Neurol 1994, 126, 151.33. Shamlou, S.; Kennedy, J. P.; Levy, R. P. J Biomed

Mater Res 1997, 35, 157.34. Isayeva, I. S.; Kashibhatla, B. T.; Rosenthal, K.

L.; Kennedy, J. P. Biomaterials 2003, 24, 3483.35. Kennedy, J. P.; Fenyvesi, G.; Levy, R. P.; Rosen-

thal, K. L.Designed Networks for Immunoisola-tion. Presented at Bioartificial Organs II, Engi-neering Foundation Banff, Alberta, Canada, July18–22, 1998.

36. Kurian, P.; Kennedy, J. P. J Polym Sci Part A:Polym Chem 2002, 40, 1209.

37. Kurian, P.; Kennedy, J. P. J Polym Sci Part A:Polym Chem 2002, 40, 3093.

38. Kurian, P.; Kashibhatla, B.; Daum, J.; Burns, C.A.; Moosa, M.; Rosenthal, K. L.; Kennedy, J. P.Biomaterials 2003, 24, 3493.

39. Kennedy, J. P.; Rosenthal, K. L.; Kashibhatla, B.Des Mon Polym 2004, 7, 485.

40. Isayeva, I. S.; Kennedy, J. P. J Polym Sci Part A:Polym Chem 2004, 42, 4337.

HIGHLIGHT 2963