Methods for reducing biosensor membrane biofouling

Colloids and Surfaces B: Biointerfaces 18 (2000) 197 – 219 www.elsevier.nl/locate/colsurfb

Methods for reducing biosensor membrane biofouling Natalie Wisniewski, Monty Reichert * Department of Biomedical Engineering, Duke Uni6ersity, Durham, NC 27708, USA

Abstract The deleterious effect that biofouling has on sensor stability is a serious impediment to the development of long term implanted biosensors. This paper reviews the surface modification strategies currently employed to minimize membrane biofouling of in vivo sensors. Nine sensor modifications are discussed herein: hydrogels, phospholipidbased biomimicry, flow-based systems, Nafion, surfactants, naturally derived materials, covalent attachments, diamond-like carbons, and topology. © 2000 Elsevier Science B.V. All rights reserved. Keywords: Membrane biofouling; Biosensors; Protein adsorption; Cell adhesion; Biocompatibility

1. Introduction Biofouling is the accumulation of proteins, cells and other biological materials on a surface. It is one of several causes for the failure of in vivo biosensors. Fig. 1 illustrates several potential sources of sensor malfunction for a model needletype electrochemical subcutaneous glucose sensor. Failure modes may be divided into two main categories: (1) component-based failures such as lead detachment, electrical shorts, and membrane delamination; and (2) biocompatibility-based failures such as membrane biofouling, electrode passivation, fibrous encapsulation, and membrane biodegradation. Although no study has shown a convincing rank ordering of the issues leading to in vivo sensor failure, membrane biofouling is clearly detrimental to sensor function [1 –12]. * Corresponding author. Tel.: +1-919-6605151; fax: +1919-6605356. E-mail address: [email protected] (M. Reichert).

In order to be detected by a sensor, a model analyte such as glucose must diffuse out of a capillary, through the surrounding tissue, to the sensor surface without being consumed by nearby cells. The analyte must then penetrate the outer sensor membrane, and diffuse into a chamber where detection takes place. Anything that raises or lowers capillary perfusion or analyte diffusion can affect sensor performance [13]. Normal skin wound healing has four stages: homeostasis; inflammation; repair; and scar formation. The body responds to a newly implanted sensor with essentially the same wound healing regimen, except the type of material on the exterior of the sensor will influence the length and characteristics of each wound healing stage [14,15]. Table 1 outlines the effect of implantation on the performance of subcutaneous sensors for small molecular weight analytes (e.g. glucose). The membrane biofouling layer and scar-like capsule that develop around sensors are generally avascular and fibrous, lead-

0927-7765/00/$ - see front matter © 2000 Elsevier Science B.V. All rights reserved. PII: S 0 9 2 7 - 7 7 6 5 ( 9 9 ) 0 0 1 4 8 - 4

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ing to reduced analyte diffusion and perfusion to implanted sensors which ultimately causes a decrease in sensor response. Fig. 2 shows two examples of declining glucose sensor response. In one case the sensors were immersed in heparinized blood (Fig. 2(A)); in the other case, the sensors were implanted subcutaneously (Fig. 2(B)). In both instances the level of sensitivity loss due to exposure to a biological medium ranged from minimal to extensive. The decay of the sensor signals seen in Fig. 2(A and B) is typical for tissue contacting or blood-contacting sensors. Membrane biofouling is phenomenologically distinct from electrode fouling [16], often called electrode passivation. Membrane biofouling is driven by adsorptive and adhesive interactions of proteins and cells at the outer sensor surface, whereas electrode passivation occurs when molecules are able to penetrate the sensor and come in contact with the electrodes. Passivation is caused by small molecules [17,18], whereas membrane biofouling can be caused by large and small molecules as well as by cells. The majority of membrane fouling studies have involved materials employed in bio-processes such as membranes for microbial suspension [19], hormone separation [20], protein fractionation [21], cell separation [22], waste water treatment [23],

oligosaccharide bioreactors [24], protein ultrafiltration [25], and dairy processing [26]. There are however a number of studies specifically concerning implantable sensor membrane biofouling. Unfortunately, the biomedical sensor community is dispersed throughout several disciplines, and many of the pertinent papers on sensor fouling are deceptively difficult to identify. For example, if one performs a Medline search of the literature combining ‘sensor,’ ‘membrane’ and ‘fouling’ only two reports will arise [12,22]. Even worse, a Medline search combining ‘biosensor’ and ‘fouling’ yields no papers. Although a Medline search on sensors and biosensors yields over 6000 papers, the vast majority do not address biofouling at all. Considering the enormous number of sensor studies reported in the literature, one finds that relatively few studies have systematically compared two or more sensor membranes and that membrane materials are generally chosen out of convenience. Biocompatibility is a vague concept for which a variety of definitions exist. In the context of implanted sensors, biocompatibility encompasses the body’s reaction to the implanted sensor as well as the sensor’s reaction to the body. The latter phenomenon is coined sensocompatibility and is the subject of two recent reviews [14,15]. Currently,

Fig. 1. Schematic illustration of glucose molecules exiting a capillary and diffusing to a subcutaneously implanted needle-type glucose biosensor. In addition to normal component failure such as electrical failure, enzyme degradation, and membrane delimination, the sensor can fail from several physiologically related causes, such as membrane biodegradation, electrode passivation, fibrous encapsulation, and membrane biofouling.

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Fig. 2. Two examples of biosensor sensitivity decay caused by exposure to the biological environment. In both cases the time scale of sensitivity decay is on the order of minutes, which is typical of protein adsorption and cell adhesion. (A) Decay of glucose sensor response as a function of residence time in heparinzed blood spiked with glucose from Reddy et al. (1997) [11]. The upper to lower curves are sensors with outer membranes of Cuprophan, polycarbonate, and polyvinylchloride, respectively. The ordinate (‘% original response’) is the sensor signal measured at a given time relative to the to the initial steady-state response achieved within 2 min of blood exposure. (B) Decay of glucose sensor response as a function of implantation time in subcutaneous tissue from Rebrin et al. (1992) [1]. The open circles in the upper portion of the graph show the relatively constant plasma glucose (PG) concentration as indicated on right hand ordinate. The lower portion of the graph shows the response of ten identically prepared and implanted sensors. Sensors 1–7 were implanted in non-diabetic, healthy dogs. Sensors 8 – 10 were implanted in diabetic dogs. The left hand ordinate (I/Io) is the ratio of the sensor signal measured at a given time relative to the sensor signal recorded 5 min after implantation.

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Table 1 Effects of wound healing on sensor detection of small molecular weight blood borne analytes

Cause

Effect

Hemostasis

Inflammation

Repair

Encapsulation

Minutes to hours

Days

Days to weeks

Weeks to months

Blood borne proteins and platelets adhere to sensor surface Membrane biofouling restricts analyte diffusion to sensing layer

Proteins and cells of immune system adhere to sensor surface Membrane biofouling, and enzymatic degradation of sensor membrane

Vascularized and gelatinous granulation tissue formation Increased analyte perfusion and diffusion to the sensor surface

Increasing avascularity and fibrocity of surrounding tissue Decreased analyte perfusion and diffusion to the sensor surface

there are four noteworthy reviews that discuss sensor membrane design as it relates to biocompatibility. Gerritsen et al. [27] reviews the failure mechanisms that lead to the poor in vivo performance of subcutaneously implanted glucose sensors, highlighting current glucose sensor configurations and animal models used to assess in vivo performance. Kryolainen and coworkers [12] characterize the problem of glucose sensor membrane biofouling as being caused by increased barriers to analyte diffusion at the outer membrane surface. Schlosser and Zeigler [28] describe biomaterials – host interactions pertinent to glucose sensors and review glucose sensor membrane modifications such as heparin, phospholipid, and diamond-like carbon surface treatments. Reichert and Saavedra [29] review the material science of implantable sensor encapsulants, including a section on physiochemical and biological modifications that may lead to improved sensor membrane biocompatibility. All of these papers devote significant attention to the broader category of biocompatibility, one aspect of which is biofouling. Recently, we reviewed the various analytical techniques used to characterize biofouling of biosensor membranes [30]. In the current paper, the diverse set of membrane modifications used to combat the biofouling of biosensors is reviewed. Nine sensor modifications are discussed herein: hydrogel overlays, phospholipid-based biomimicry, flow-based systems, Nafion, surfactants, naturally derived materials, covalent attachments, diamond-like carbons, and topology treatments. Table 2 summarizes studies employing

these modifications. Because biosensor designs and test methods vary so widely from group to group, it is virtually impossible to make valid comparisons between the various approaches. For this reason, only studies that test at least two sensor membranes under the same conditions are listed in Table 2. Unfortunately none of these nine approaches completely eliminate biofouling, nor does any one stand out as being most effective. However, strategies based on biomimetics and surface perfusion show strong potential to substantially improve sensor longevity considering the relative length of in vivo functionality of sensors employing these modifications (Table 2).

2. Hydrogel overlays (Table 2A) Hydrogels have been incorporated into a broad range of biomaterial and pharmaceutical applications [31]. The most widely used hydrogels are cross-linked polymers of either poly(hydroxyethyl methacrylate) (PHEMA) or poly(ethylene glycol) (PEG). Coatings of these polar, uncharged, water swellable, flexible materials mask the underlying surfaces by producing a hydrophilic interface between the solid surface and aqueous bulk. Both PHEMA and PEG are attractive outer membrane coatings for sensors because water-soluble analytes can diffuse readily through the water-swollen polymer gel layer. The degree of analyte diffusion is readily modulated by controlling crosslink density of the gel, which in turn controls the gel water content and the openness of the polymer network. Several studies have focused on optimizing hy-

Table 2 Summary of studies which compare biosensor membrane performances Year

Inner membrane

Outer membrane

Biological media

Results

A. Hydrogel o7erlays Quinn [54]

1997

None

1. PellathaneTM (PU) 2. SPA-3400-APEG

Rat SQ for 7 days

1. PU was covered with neutrophils, macrophages and FBGCs 2. SPA-340-APEG was covered with few neutrophils, but no macrophages or FBGCs on hydrogel Did not compare sensor responses in vivo

Quinn [147]

1995

Polyallylamine

1. Cross-linked horseradish peroxidase derivatives 2. Copolymer of HEMA and PEG-derivative

Rat SQ for 3 days

1. Completely encapsulated by fibrous tissue 2. Very little fibrous tissue Did not compare sensor responses in vivo

Espadas-Torre [106]

1995

None

1. PU 2. 18.5 k PEO hydrogel 3. 200 k PEO hydrogel

Platelet rich plasma for 3 days

Drastic reduction in platelet adhesion on hydrogel compared to PU control Did not describe sensor response in biological medium

B. Phospholipid-based biomimicry Ishihara [59] 1998

None

1. Cuprophan 2. MGC

14 days in human SQ

1. Cuprophan gave stable sensor response over 4 days 2. MGC gave stable sensor response over 7 days

30–70 g l−1 BSA

1. Cuprophan had 57–90% drop in recovery over 14 days 2. MGC 15–25% drop in membrane permeability over 14 days

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Reference

201

202

Table 2 (Continued) Year

Inner membrane

Outer membrane

Biological media

Results

Shichiri [61]

1998

None

1. Cuprophan 2. MPC-co-BMA

50 g l−1 BSA

1. Stable over 3 days; on day 7 sensor had 28.5% of initial response; on day 14, sensor had 2.8% of initial response 2. Stable over 7 days; on day 14 had 26.8% of initial response TEM after human implantation showed nothing adhering to MPC-co-BMA, but Cuprophan was completely biofouled

Zhang [58]

1996

None

1. MPC-co-BMA 2. Heparin

1 h in whole blood and 3 h 1. MPC-co-BMA sensors were in rat artery fairly stable under both conditions 2. Heparin coated sensors had 40–50% decrease in response under both conditions

Zhang [148]

1996

None

1. PU 2. PVC

Several days in serum

3. MPC-co-BMA modified PU 4. MPC-co-BMA Nishida [8]

1995

None

1. PVA 2. MPC-co-BMA

1. Significant decrease in response 2. Significant decreases in response Stable response Stable response

Human SQ for 14 days 70 g l−1 BSA

1. PVA maintained 71% of initial response on day 7 2. MPC-co-BMA sensor maintained 96% of initial response on day 7 Protein adhesion results: compared to PVA, MPC-co-BMA had five times less protein adhering on Day 1 and 2.5 less on Day 14

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Reference

Table 2 (Continued) Year

C. Flow-based systems See Ishihara above — these sensors employ both phospholipid biomimicry and microdialysis perfusion technology Towe [149] 1997

Inner membrane

Outer membrane

Biological media

Results

None

1. Hi Flux (Fresenius)

Dog vein for 5 days

1. High Flux membranes were too fragile to use 2. Time to reach 70% equilibrium doubled between day 1 and day 5 — attributed to biofouling

6 day in vivo

Nafion and MPC maintained better glucose permeability than uncoated cellulose based membranes

2. Cuprophan

Sudoh [83]

1998

None

1. 2. 3. 4. 5. 6. 7.

Regenerated cellulose Cellulose acetate Cellulose triacetate PMMA EVA Nafion MPC

Rigby [73]

1996

PES

Rigby [74]

1995

PES

1. PU with perfused saline Whole blood and rat SQ for All three conditions yielded 2. PU with perfused heparin 4 h same degree of protection against biofouling Pre and post calibrations show microperfused sensors were within 3% or original sensitivity 3. PU with perfused glycerol 1. PU Rat SQ and blood for 4 h Microperfused sensor followed 2. PU with saline blood glucose adequately and microperfusion had significantly less fouling

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Reference

203

204

Table 2 (Continued) Year

Inner membrane

Outer membrane

Biological media

Results

Rigby [75]

1999

PES

1. PU with perfused saline 2. PU with perfused EDTA 3. PU with epinephrine vasoconstrictor

Human SQ for 4 h

EDTA improved blood:tissue glucose correlation compared to saline; vasoconstrictor caused delayed glucose tissue response and reduced blood:tissue glucose correlation; post explant response was within 5% of pre-implant response

D. Nafion membranes Valdes [102]

1999

None

1. Nafion

Culture media for 4 weeks

Soaking Nafion in FeCl3 significantly reduced Nafion calcification

2. FeCl3 soaked Nafion Harrison [97]

1998

None

Nafion films of various thickness

6 days in whole blood

Coating above 0.8 mm thick gave satisfactory electrode stability — 6% loss of glucose calibration slope

Park [103]

1998

None

1. 2. 3. 4.

5-6 h in rats

Did not report individual membrane performances, but after explantation, composite membrane sensors (Nafion, 1,3-phenylenadiamine, and 1,3-benzenediol) retained 88% of pre implant sensitivity

Moussy [5]

1994

Poly(o-phenylenediamine

1. Uncoated AgCl 2. Uncured Nafion 3. Cured Nafion

Rat SQ for 2 weeks

Pre implant and post explant potentials dropped (mV):

Nafion 1,2-phenylenadiamine 1,3-phenylenadiamine 1,3-benzenediol

1. 21 to −158 2. 24 to −141

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Reference

Table 2 (Continued) Reference

Year

Inner membrane

Outer membrane

Biological media

4. PU

Results 3. 30 to 15 4. 21 to 16

1994

Poly(o-phenylenediamine

1. Uncured Nafion 2. Cured Nafion

Dog SQ for 10 days

1. Uncured Nafion failed after 1 week 2. Cured Nafion had constant response over 10 days

Moussy [6]

1994

Poly(o-phenylenediamine

1. Nafion cured at 120°C 2. Nafion cured at 170°C

Dog SQ for 10 days

1. 50% loss of sensitivity upon implantation, but was then stable 2. Lost all sensitivity due to curing

E. Surfactant modifications Reddy [11]

1997

Not reported

1. 2. 3. 4. 5.

60 min in whole blood; Post-exposure response of exposed membranes reapplied rinsed membranes on clean to fresh electrode for post electrodes compared to testing pre-exposure response 1. 100% 2. 48% 3. 68% 4. 90% 5. 100% Reduced also fouling visualized with SEM

Kryolainen [150]

1997

Sulphonated polyether-ether sulfone/PES

1. PVC 2. PVC with non-ionic surfactant

Kryolainen [17]

1997

Sulphonated polyether-ether sulfone/PES and CA

1. PVC — low MW nonionic 90 min in whole blood surfactant 2. PVC — high MW nonionic surfactant Both 1 and 2 were cast and spin coated

CUP PVC–unmodified PVC–cationic (Aliquat) PVC-lipid (IPM) PC-unmodified

3 h in 50% diluted blood

Surfactant reduced fouling and broadened linear range, but quantitative comparison was not reported; 3% difference in calibration before and after 3 h in blood with surfactant-modified PVC

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Moussy [99]

1. When cast was stable, but when spun, lost up to 40% of response 2. When cast or spun was stable 205

206

Table 2 (Continued) Year

Inner membrane

Outer membrane

Lindner [110]

1994

None

1. PVC Rat SQ for 14 days 2. Aliphatic polyurethane (1:2 polymer:plasticizer in each)

Both had similar acute inflammatory responses, but PVC had a reduced chronic imflammatory response compared with Tecoflex; lower plasticizer content caused less of a foreign body reaction

F. Naturally deri7ed materials Ammon [64] 1995

1. PU

1. CUP (wood cellulose)

1:10 diluted blood

and Eisele [10]

2. CUP

2. Bacterial cellulose (known to possess a hydrogel character)

1. 50% loss of activity after 30 h in diluted blood 2. 40% loss after 200 h in diluted blood

Whole blood

1. 50% loss after 3 h in whole blood 2. 50% loss after 25 h in whole blood

1994

Biological media

Results

G. Co7alently modified membranes See Park in Nafion section — this study employs both covalent modifications and Nafion Yang [121] 1997

PTFE

1. Control 2. 1,3-phenylenediamine

Whole blood

1,3-phenylenediamine coating improved sensor performance in whole blood compared to control

Myler [123]

1997

PC

1. PC 2. Poly(o-phenylenediamine) on a gold, sputter-coated porous PC

2 h in whole blood

PC stability not reported, but 30% loss in sensitivity; in extended tests sensor was stable after the initial 30% decrease Microscopy shows reduced biological build-up compared to plain PC

Reynolds [4]

1994

None

1. No membrane 2. Poly 1,3-DAB

6 h in 3% BSA

1. 58% decrease in sensitivity 2. 13% decrease in sensitivity

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Reference

Table 2 (Continued) Year

Inner membrane

Outer membrane

Biological media

Guo [124]

1996

None

1. 2. 3. 4. 5. 6.

0.4% HSA and serum for 1 h 1–4. Reduced fouling compared to a glassy carbon electrode 5–6. Fouled more than the glassy carbon electodes Conclusion: diols are effective modifiers, alcohols with hydrophobic terminal groups are not

Keedy [125]

1991

Cellulose acetate

1. Dimethytldichlorosilane treated PC 2. Cellulose nitrate

Whole blood — unspecified 1. Signal was constant for an amount of time unreported amount of time 2. 1–7% loss/5 min in blood and 1–2% loss/5 min in plasma

H. Diamond like carbons Higson [129]

1993

PC

DLC coatings of various thicknesses applied to PC (0.0075–0.053 mm)

30 min in whole blood

Uncoated PC fouled more than coated PC; thicker DLC films resisted biofouling more than thinner ones; in best case sensor lost 6% of sensitivity after 30 min

Higson [7]

1995

PC

1. PC 2. DLC on PC 3. Unspecified dialysis material 4. DLC on unspecified dialysis material

30 min whole blood

1. 43% of original pre-exposure response retained 2. 60% of original pre-exposure response retained 3. 51% o f original pre-exposure response retained 4. 66% of original pre-exposure response retained Reduced fouling was visualized with SEM

Ethyleneglycol Diethyleneglycol 1,2-propanediol 1,3-propanediol Ethanol 2-methoxyethanol

Results

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Reference

207

208

Reference

Year

Inner membrane

Outer membrane

Biological media

Results

I. Surface topology Higson [129]

1993

PC

1. PC pore-0.02 mma 2. PC pore — 0.1 mma 3. PC pore — 0.2 mma

30 min in whole blood

1. 70% or original response retained 2. 67% or original response retained 3. 54% of original response retained

Reddy [11]

1997

Not reported

1. PC 0.1 mm poresa 2. PC 0.01 mm poresa

60 min in whole blood

1. 10% loss of response 2. No loss of response Subsequent data showed the differences were due to passivation of the electrode, not biofouling of the membrane

a

Pore diameter.

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Table 2 (Continued)

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drogel permeability for various applications [32– 34]. Many studies claim favorable biocompatibility for PEG [35,36] and PHEMA [37–39]. Impurities in the polymer [40] and the molecular weight of the precursors [35,41] have been shown to be critical factors that affect the interaction of hydrogels with biological media. One can also chemically modify hydrogels to enhance their biocompatibility. For instance, recalcification times (RCT) of plasma exposed to heparin-modified polyvinyl alcohol (PVA) hydrogels were significantly higher than RCT values for unmodified PVA hydrogels [42]. Relative to their extensive use as biomaterials in other applications, comparatively few sensor studies have applied hydrogels to the exterior of sensors and tested their in vivo response over time [43,44]. A more common application of hydrogels to sensors has been their use as a gel matrix to imbed bio-recognition proteins, such as enzymes, in the interior of the sensors [39,45 – 50]. Several problems have been encountered in applying hydrogels to the sensor surface: (1) hydrogels can poorly adhere to the underlying substrate; (2) hydrogels can have less than acceptable mechanical stability to withstand forces of implantation; (3) the monomer, solvent and crosslinking agent (like UV light) can have detrimental effects on the enzyme and other components; and (4) hydrogels can limit analyte diffusion into the sensor. Hydrogel coatings applied to sensors should be optimized to minimize these problems. Studies of successfully applied PHEMA to sensors have shown reduced clotting, reduced protein adsorption, and enhanced electrode stability in dogs and monkeys [51 – 53]. Quinn et al. [54] showed a substantially decreased immune response around the PEG-derivatized samples compared to Pellethane™ samples after 7 days in rat SQ tissue, although the in vivo sensor response was not characterized. The above sensor results and the vast number of non-sensor biocompatibility studies lead one to believe that hydrogels have the potential to be an effective anti-fouling sensor component. More optimization of the application processes and material characteristics is needed before hydrogels become widely used as biosensor coatings.

209

3. Phospholipid-based biomimicry (Table 2B) One strategy for altering biosensors is to cover them with coatings that mimic the body. Similar to a cell membrane, sensor surfaces should ideally be able to selectively and specifically bind certain molecules while resisting undesirable, non-specific interactions. Currently, the most successful strategies for improving sensor biocompatibility by biomimetic modifications involve phospholipid, phospholipid-containing, or phospholipid-like biomaterials [55]. Phospholipids are the major component of cell membranes. By coating sensors with phospholipids, their surfaces resemble a cell’s own membrane and perhaps seem less foreign to the body’s immune system. Because phospholipid membranes are fragile and are difficult to deposit, the majority of these materials are phospholipid-modified polymers [56]. The water content of the phospholipid polymers is extremely high (nearly twice the free water fraction of PHEMA). Further, ultraviolet and circular dichroism spectroscopic measurements have shown that proteins absorbing on PHEMA undergo considerable changes in conformation, whereas proteins adsorbing on phospholipid polymers are in a similar confirmation to the native state [56]. It is suggested that the high water content of the phospholipid polymers is responsible for their good antifouling properties [56]. Originally developed for improving the blood compatibility of hemodialysis membranes [56], Ishihara and coworkers showed that coating needle-type glucose sensors with a copolymer of 2-methacryloyloxyethylphosphorylcholine and nbutylmethacrylate (MPC-co-BMA) improves their longevity up to 14 days [57]. This group has also applied MPC-co-BMA coatings to improve the performance of pO2 sensors in vitro and in vivo [58]. Subcutaneous glucose monitoring in humans using MPC-grafted and MPC-co-BMA coated cellulose microdialysis membranes also yielded close correlation to blood glucose for up to 7–14 days [59–61]. Besides the biomimetic character of MPC-co-BMA, other reasons for improved sensor longevity include covalent attachment of the polymer to the underlying sensor membrane, hydrogel characteristics of MPC-co-BMA, and application

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of perfusions systems to reduce biofouling (microdialysis). Similar coatings tested in vivo showed significant reductions in platelet adhesion, platelet activation, and fibrinogen adsorption [62,63]. In general, phospholipid-derived coatings represent a promising short-term solution to biofouling. However, longer term biomimicry approaches will likely require some sort of surface replenishment.

4. Flow-based systems (Table 2C) Combining sensing technology with flow systems has gained considerable interest recently. Flowing fluid over the material-tissue interface seems to allow implanted surfaces to resist biofouling and remain functional for a longer period of time [64 – 74]. Four different types of sensors have been developed that rely on flowing fluid. Fig. 3 compares differences in the implanted portion of these four systems.

4.1. Manchester open flow microperfusion sensor In the Manchester Open Flow Microperfusion System, a thin stream of phosphate buffered saline is constantly perfused over the surface of a glucose sensor that protrudes from the tip of a

fluid filled syringe (Fig. 3(A)). Surface-perfused sensors retained 100% sensitivity in blood, whereas unperfused sensors retained only 33% of their sensitivity under the same conditions [74]. During a 4 h implantation in a rat, surface-perfused sensors successfully tracked blood glucose levels [72] and retained 97% of their original sensitivity [73]. The stable sensor response indicated that little or no biofouling took place in vivo. The constantly replenished thin liquid film appears to physically deter accumulation of undesirable molecules and cells at the sensor surface, and hence, reduce the amount of biofouling. Other purported reasons for sensor enhancement include improved diffusion through hydrated local tissue, facilitated local convection, increased sample pick-up area and reduced surface coagulation [72]. Further studies with the Manchester System show that perfusing with a chelatating buffer (EDTA) yields more stable results than perfusing with isotonic saline. Adding a vasoconstrictor to the perfusion solution caused a delayed glucose detection and worsened correlation to blood values [75]. Future studies to determine the long-term effectiveness will be of interest. Addition of heparin or angiogenic factors to the perfusion stream could perhaps extend sensor longevity. For long-term application, the total amount of perfusion fluid released into the body could be substantial. The effect of the fluid release on the body has not been assessed; however, it is believed that because the perfusion rate is low (60 ml h − 1) [74], the effects would be negligible.

4.2. Microdialysis sensors

Fig. 3. Four types of sensors that employ flow strategies. The illustrations shown are the portions of the sensors that reside in the tissue. (A) Manchester Open Flow Microperfusion (B) Microdialysis (C) Graz Open Flow Microperfusion (D) Ultrafiltration. In systems B-D, the fluid is transported to an exterior sensor, whereas in system A, the sensor is implanted in the tissue within the flow stream.

The most widely studied flow-based sensor combines microdialysis with traditional sensing technology [59,61,65–70,76–79]. In contrast to the Manchester set-up, microdialysis sensors do not permit convection of fluid into the tissue. A microdialysis probe consists of a semi-permeable hollow fiber membrane (a single dialysis fiber) through which fluid is continuously perfused (Fig. 3(B)). Analyte sampling occurs by diffusional movement of analyte across the implanted fiber membrane [80]. It has been shown that a microdialysis probe configured with a downstream

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biosensor can continually monitor blood glucose levels while suffering less sensitivity loss than conventional subcutaneous sensors [65,67 – 70]. Stable microdialysis detection of glucose for three weeks has been reported in humans [81]; however, this study relied on off-line glucose analysis rather than a coupled glucoanalyzer. A novel configuration of the microdialysisbased sensors is The Ulm Sugar Watch System [82]. It combines an implanted subcutaneous microdialysis probe with an in-line enzymatic glucose meter which telometrically signals a digital read-out wristwatch. Another unique application of microdialysis is the Mini Shunt System [69]. This device employs a microdialysis fiber fed through the lumen of an ex vivo blood catheter that both withdraws and returns blood. The 160 mm long fiber allows small blood-born analytes like glucose to come to equilibrium with the perfusate. A downstream biosensor is used to assay glucose in the dialysate. This system has been successfully evaluated over 4 h in dogs, but more extensive tests are needed to show its longer-term potential. The choice of dialysis membranes for the majority of microdialysis studies seems to be based primarily on availability with little attention paid to membrane optimization. For the short implantation times reported (less than a week), most membranes performed adequately. Ishihara et al. [59] and Sudoh et al. [83] have attempted to optimize microdialysis membranes for longer implantation times. Application of a phospholipid based coating [59] and Nafion [83] significantly improved the in vivo and in vivo longevity of cellulose microdialysis fibers, and these microdialysis sensors have functioned in vivo for up to 14 days. Microdialysis based sensors seem to resist fouling in the short term and several researchers have attested to their biocompatibility [64–71]. However, long-term reliable implantation requires more extensive work on membrane modifications or the selection of new membrane materials.

Graz system is more similar to microdialysis. However, the Graz system does not possess any size exclusion properties because the semi-permeable outer membrane has macroscopic holes punched in it. In the Graz system, a perforated, open-ended double lumen catheter inserted in either a vein [84] or in subcutaneous tissue [85,86] is perfused with a heparin solution or isotonic fluid, which flows out of the inner lumen. At the same time, heparinized blood or ‘interstitial fluid’ is continually drawn out of the vein or tissue through the open-ended outer lumen by means of an applied negative pressure (Fig. 3(C)). The sampled fluid passes over an exterior glucose electrode for continuous analysis. The system has been shown to function successfully for 6 h in blood and 24 h in subcutaneous tissue [84–86].

4.3. Graz open flow microperfusion sensor

For ultrafiltration sensors and all other flowbased sensors, the reliance on moving fluid necessitates the use of pumps, vacuums or fluid reservoirs. Although flowing fluid seems to extend

While bearing the same name as the Manchester open flow microperfusion system, the

4.4. Ultrafiltration sensors Another flow-based sensor combines ultrafiltration with conventional sensors. This technique is similar to microdialysis except that it uses an applied negative pressure to draw interstitial fluid through an implanted microdialysis fiber, and it does not utilize perfusion for sample uptake (Fig. 3(D)). The collected outflow (ultrafiltrate) is believed to represent 100% of the concentration of the interstitial fluid for small analytes (i.e. the interstitial fluid is drawn directly into the fiber through the semi-permeable fiber wall and not diluted with any perfusion fluid). Discontinuous ultrafiltration without an online biosensor has been used for one month in humans to monitor subcutaneous glucose [87], and no significant tissue reactions were noted. An online analysis sensor has been developed for use with ultrafiltrations and was tested for up to 24 h in conscious rats [88,89]. The system currently used needs to be miniaturized for it to be appropriate for continuous monitoring in humans.

4.5. Concluding remarks about flow-based systems

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sensor lifetime, it also adds complexity, making perfusion-based sensors difficult to implant fully. Since insulin infusion devices with exterior pumps have been used for many years by diabetics, it seems that as long as reliable sensor output could be attained, extracorporeal flow-based sensors could also be adopted. Despite the seeming difficulties, Towe has proposed a novel fully implantable prototype microdialysis-based sensor that would have to be regenerated approximately every two weeks (reagent refilled and waste removed) [90]. Extremely small volume perfusion systems and single molecule sensing devices (nanoliter and picoliter quantities compared to the microliter systems commonly used now) show that the potential for such an implantable device exists [91]. In the near future however, flow-based sensors show the most potential as extracorporeal devices.

5. Nafion membranes (Table 2D) Nafion is a perfluorosulfonic acid polymer that has gained considerable popularity as a coating for glucose sensors [48,92 – 96]. It is a chemically inert, anionic polymer with both hydrophobic and hydrophilic properties [97,98]. Nafion films applied to biosensors at a functional film thickness of approximately 1 – 2 mm promote relatively little adsorption of molecules from solution [98] and prolong the life of the sensors more than cellulose membranes [97]. Moussy et al. have characterized this material in numerous studies where the coated sensor output and in vivo tissue biocompatibility were evaluated during long-term implantation [5,6,95,97,99 – 101]. The stress cracking of Nafion films on implanted sensors, thought to be caused by calcification [101], was minimized by pre-incubation in a FeCl3 solution [102] and high temperature curing [5,99]. Stable sensor response and limited inflammatory reaction was attained in dogs for 10 days. In comparison, uncured Nafion sensors or uncoated sensors failed much sooner and elicited a stronger inflammatory reaction [99]. One reason Nafion is widely used is that it is easy to apply to sensors using a simple dip-coating procedure. Although several coats may be

required [103], no specialized reaction chambers (nitrogen chambers, UV radiation, etc.) are needed as with other polymers. Another reason Nafion has been widely used is its commercial availability. Several other anti-fouling modifications require in-house synthesis, making them more labor-intensive than Nafion. A third reason for Nafion’s wide usage is its biocompatibility and its exclusion of interferents believed to be caused by the anionic nature of this material. Though the long term stability (weeks) is still an issue, Nafion is a good candidate for short-term sensor needs.

6. Surfactant modified sensors (Table 2E) A surfactant, or surface active agent, is a molecule that preferentially seeks the interface between two phases. Surfactant molecules, such as lipids with polar head groups attached to hydrocarbon tails, possess both hydrophobic and hydrophilic character. The head group can be anionic, cationic or uncharged. Surfaces adsorbed with Pluronic surfactants (PEO-PPO-PEO triblock copolymers) have been shown to possess enhanced resistance to protein adsorption and cell adhesion [104–106]. Many membranes used in electrochemical sensors contain surfactant plasticizing molecules that affect species mobility and selectivity [107,108]. Vadgama and coworkers [11,17,109] have examined whether charged and uncharged surfactant plasticizers could be used to reduce the fouling of blood contacting enzyme electrodes. Adding surfactants to sensor outer membranes has improved sensor functionality by decreasing protein adsorption and cellular adhesion in blood, and by increasing the stability of glucose sensors. It was shown that anionic surfactants improve sensor longevity more than non-ionic surfactants, and much more than cationic surfactants [11,17]. In a thorough study by Reddy and Vadgama [11], the effects of surfactants on membrane biofouling were clearly separated from their effects on electrode passivation, and it was shown that surfactant modified membranes, in most cases, reduce both biofouling and passivation. It was suggested that the anionic surfactants may inhibit adhesion

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of negatively charged blood borne cells, and that the increased fluidity of the platicized membrane may reduce adhesion as well [11,17]. However, no clear explanation is evident. Using surfactants to reduce biofouling of sensors is still in the early developmental stage compared to MPC or Nafion. Perhaps leaching antifouling surfactants, such as calcium chelating plasticizers that limit Factor X activation (an enzyme in the coagulation cascade), could be a means of continuously renewing the sensor surface. However, Lindner et al. recently showed that plasticizers leaching from implanted sensor membranes can lead to enhanced inflammation [110], suggesting that membrane plasticizer content should be minimized. To date only short exposure times have been evaluated (see Table 2E). Whether or not surfactantplasticized membranes are appropriate for longer-term, in vivo applications remains to be seen.

7. Naturally derived membranes (Table 2F) Membranes of biological origin such as silk fibroin [111], cellulose [112], and chitosan [113] have been applied to biosensors. Cellulose and cellulose derivatives are commonly used as biosensor membranes [93,97,112,114 – 116]. It has been shown that when hydroxyl groups of cellulose are converted into acetates or amines, complement activation decreases significantly [117,118]. Ammon and Eisele [10,64] compared the biological response and functional longevity of sensors modified with an experimental bacterial cellulose (BC), and a commonly used cellulose of wood origin (Cuprophan). They found BC coated sensors operated six to seven times longer than Cuprophan coated sensors in undiluted blood [10,64]. BC coated sensors also performed reliably when placed intravascularly in rats for three days [64]. The improved function of BC sensors is attributed to the hydrogel-like qualities of bacterial cellulose. Sensors with silk fibroin [111] and chitosan [113] outer membranes have been produced, but their in vivo stability has not been tested.

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Albumin [119,120] and heparin [121,122] are two other naturally derived substances that have been employed to improve the biocompatibility of biosensors. Albumin is widely used on the interior of sensors (in combination with gluteraldehyde) as a means of immobilizing glucose oxidase or other analyte recognition elements. It has also occasionally been employed as a coating on the outside of biosensors to reduce non-specific binding and improve surface biocompatibility [119,120].

8. Covalently modified membranes (Table 2G) Some modifications discussed in other sections (hydrogels, phospholipids, naturally-derived modifications) may be covalently or non-covalently attached to a sensor surface depending on the surface chemistry. The primary advantage of covalently attaching anti-fouling agents is improved longevity of the surface modification. Some other covalent modifications that have shown promise on biosensors include phenylenediamine [4,121,123], various diols, [124] glycols [124], and silanes [125]. It is proposed that the surface N-H groups of phenylenediamine might be responsible for its bicompatible characteristics [123], since it has previously been shown that N-H groups on polyacrylonitrile (PAN) facilitate absorption of certain factors that contribute to PAN’s favorable biocompatibility [126]. Glycols and silanes have been used to modify sensor surface charge and wettability; however, none of these modifications have been tested in vivo, and their effectiveness as an anti-biofouling strategy remains to be seen. Phenylenediamine has been widely used in sensors as an inner coating, but a few groups have also investigated its use as an outer protective coating. As an outer coating, phenylenediamine has been found to increase sensor longevity in blood and protein solutions [4,121,123].

9. Diamond-like carbons (Table 2H) Deposition of carbon coatings are a common means of enhancing the biocompatibility of a material [127], but few researchers have investi-

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gated their use in biosensors [7,128]. Diamond like carbons (DLC) are chemically inert hydrocarbons that may be deposited on many different types of substrates with thicknesses on the order of 10–50 nm [129]. They are deposited on surfaces by fast atom bombardment in an argon atmosphere with a hydrocarbon introduced into the argon beam [7,128]. These coatings are generated at room temperature, and the process of application is not destructive to the underlying sensing materials as may be the case with the application of other coatings that make use of solvents, UV light, or chemically active agents. In vivo testing for the enzymatic output of macrophage and fibroblasts cells exposed to DLC coatings showed no indication of an increased inflammatory potential compared to control tissue culture plates [130]. Though the application process is relatively easy, it requires specialized atom bombardment equipment. Therefore, DLCs are not as easily obtained compared to the commercially available Nafion. DLCs also have had limited characterization on sensors. Higson and Vadgama have applied DLC coatings to sensors and have shown significant improvements in sensor longevity in whole blood testing over time periods of less than 1 hour [7,129,131]. It has yet to be shown whether or not DLC coatings will improve in vivo stability.

10. Surface topology (Table 2I) Membrane composition, including micro-architecture of the surface, has long been linked to membrane performance in vivo [132]. Most in vivo sensors are smooth-surfaced implants that become encapsulated with a relatively dense, avascular fibrous sheath of tissue [14]. Recently it has been suggested that a textured rather than a smooth surface could lead to improved long term sensor performance by increasing the vascularity around the implant [133 – 136]. Updike and coworkers now incorporate a textured ‘angiogenic layer’ onto the surface of their fully implantable glucose sensor [137]. Although there is a clear link between encapsulation, vascularity and topology, it is unknown how topology affects biofouling.

Two studies systematically examined the effect of pore size of a sensing membrane on a sensor [11,129]. No in vivo results were generated, but it was shown that PVC and PC membranes with smaller pore sizes had greater stability in blood than the same membranes with larger pore sizes. This may simply be due to proteins being more likely to enter larger pores and thus clog the membrane. Biofouling seems to be minimized by small pore sizes [11,129], but encapsulation seems to be minimized by intermediate pore sizes [133– 136]. In spite of compelling evidence linking implant topology and tissue response, the link between surface texturing and sensor performance, particularly biofouling, remains poorly understood.

11. Future Considerations Of the artificial materials that have been placed in the body over many years, those best tolerated are either completely inert, such as hip implants, or completely degradable, such as resorbable sutures. Now there is considerable interest in developing implants that are both biologically interactive and that reside in vivo for long periods, such as biosensors and encapsulated cells. A primary problem with interactive implants is that current artificial materials are non-renewable and the body will relentlessly attempt to either encapsulate or consume them. Even materials that are treated with ‘stealth technologies’ such as PEG grafting lose their ability to resist protein adsorption and cell adhesion and eventually fail in the body. The keys to success with long-term implantable, interactive implants may lie in developing biomimetic surface treatments that are renewable or that actively direct the behavior of the tissues that come in contact with them. Understanding how the body regenerates various membranes may give insight into possible means of creating renewable biosensor membranes. The kidney’s glomerular basement membrane (GBM) acts as a naturally occurring parallel to a biosensor outer membrane. The GBM provides a transport barrier excluding proteins and other large molecules from passing into the urine, yet it al-

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lows passage of smaller waste molecules [138]. A functional glucose biosensor membrane allows passage of glucose, oxygen and other small molecules, but excludes interferents, proteins, and other large molecules. Just as biofouling corrodes implanted artificial membranes, the natural GBM also becomes biofouled [138]! It is believed that the high level of acid hydrolase activity found in the GBM is responsible for clearance of biofouling debris and turnover of the extracellular membrane [138]. A key difference between the GBM and an artificial membrane is that the GBM has mechanisms to degrade and replenish the membrane, whereas biosensors do not. Beyond the Manchester surface perfusion system (Section 4.1), which is a physical regeneration method, incorporating some sort of stimuli-responsive regeneration system may lead to membranes of greater longevity. The surfactants discussed in Section 6 continually diffuse to the surface and out of the membrane until they are depleted. This is one simple form of surface renewal. Other novel means of sensor surface renewal include the use of osmotic gradients to continually strip heparin out of a membrane [139] and periodic flow injection of modified conducting and non-conducting beads to serve as renewable electrodes and renewable enzyme layers [140]. Although these methods are simplistic compared to the bodies own membrane renewal system, they form the basis for future work on renewable biosensor surfaces. Other biomimetic efforts focus more on composition rather than renewal. The phospholipid work discussed in Section 2 is the first step in the development of a more advanced mimicry of the cell membrane that can not only repel undesirable adhesion, but at the same time, also attract desired interactions. Efforts concerning phospholipids in sensors have involved the development of receptor-based devices that mimic the ability of cell membranes to selectively and specifically bind certain molecules while resisting deleterious and non-specific interactions. In general, these devices consist of bilayer phospholipid membranes doped with ligand binding proteins placed over the sensing region [141]. Binding events are detected by changes in the optical [142] or electrical [143]

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properties of the membrane. Another type of biomimetic surface called mimetic peptide surfaces (MPS) employs specific peptide sequences like RGD for binding of cells and FHRRIKA for binding heparin [144]. Immobilization of these molecules in specific patterns can direct cellular and molecular interactions. Currently, no reliable sensors based on these more advanced biomimicry concepts have emerged, and so their ability to resist non-specific binding and to promote specific binding in vivo has not been measured. Another potential technology that could lead to reduced biofouling and directed material-tissue interactions is surface protein templates. There is evidence that proteins can be used as templates to build structures with differing responses to protein interactions. Ratner and co-workers recently reported the synthesis of surfaces imprinted with the ‘footprints’ of adsorbed proteins [145]. Exposure of these protein-imprinted surfaces to protein solutions showed enhanced binding affinity for the desired target (those proteins from which imprints were produced). In order to apply this nano-architecture to current membrane technology, one may need to apply nano-texturing (to influence protein interactions) onto the microlevel pore architecture (to influence cellular interactions). An alternative would be to build the membrane structure itself out of proteins. Such protein synthesis has already been used to produce membranes with nano-pores and channels with potential for application to sensors [146].

12. Conclusions Biofouling remains a significant challenge in implantable biosensor research. The nature of the biofouling layer is primarily dependent on the surface of the biosensor and the wound healing state of the surrounding tissue. The sensor modifications described in this review are designed to reduce membrane biofouling either by preventing non-specific interactions or encouraging specific desirable interactions. Although none of these approaches eliminate biofouling, the biomimicry and perfusion technologies appear to be making the most progress, at least for short-term applica-

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tions. Designing non-fouling biocompatible sensor surfaces is an extremely complex task that will require much more fundamental research on material-tissue interactions. Further advances in biomimicry beyond single component phospholipid-like coatings seem to be the most attractive solution for long-term implantable sensor surfaces.

Acknowledgements This work was funded by a graduate student fellowship from the National Science Foundation and a predoctoral traineeship from the National Institutes of Health (T32 GM08555), as well as by research funds from the National Institutes of Health (RO1 DK54932). We appreciate helpful discussion with Erno Lindner of the University of Memphis and Francis Moussy of the University of Connecticut Health Sciences Center.

MPS PAN PC PEG PEO PES PHEMA PPO PU PVA PVC RCT RGD SPA-3400PEG

SQ TASP TEM

Appendix A. List of acronyms MW BC BSA CUP DAB DLC EDTA EVA FBGC FeCl3 FHRRIKA GBM HEMA HAS IPM MCG MPC MPC-coBMA

bacterial cellulose bovine serum albumin Cuprophan diaminobenzene diamond-like carbons ethlyenediaminetetraacetic acid ethylene vinyl alcohol foreign Body Giant Cell ferric chloride phenylalanine-histidine-argininearginine-isoleucine-lysine-alanine glomerular basement membrane hydroxyethyl methacrylate human serum albumin isopropyl myristate lipid poly[2-methacryloyl-oxyethyl phosphorylcholine (MPC)] methacryloyloxyethylphosphorylc holine 2-methacryloyloxyethyl phosphorylcholine-co-n-butyl methacrylate

mimetic peptide surfaces polyacrylonitrile poly carbonate poly(ethylene glycol)= (polyethylene oxide) poly(ethylene oxide)= poly(ethylene glycol) polyethersulfone poly(hydroxyethyl methacrylate) poly(propylene oxide) polyurethane poly(vinyl alcohol) poly(vinyl chloride) recalcification times arginine–glycine–aspartic acid copolymer of an eight-armed amine-terminated poly(ethylene glycol), and a di-succinimidyl ester of poly(ethylene glycol) dipropiomic acid subcutaneous tissue template-assisted synthetic protein transmission electron microscopy molecular weight

References [1] K. Rebrin, U. Fischer, H. von Dorsche, T. von Woetke, P. Abel, E. Brunstein, J. Biomed. Engr. 14 (1992) 33–40. [2] B.J. Gilligan, M.C. Shults, R.K. Rhodes, S.J. Updike, Diabetes Care 17 (1994) 882 – 887. [3] A. Silber, N. Hampp, W. Schuhmann, Biosens. Bioelectron. 11 (1996) 215 – 223. [4] E.R. Reynolds, A.M. Yacynych, Biosens. Bioelectron. 9 (1994) 283 – 293. [5] F. Moussy, D.J. Harrison, Anal. Chem. 66 (1994) 674– 679. [6] F. Moussy, S. Jakeway, D. Harrison, R. Rajotte, Anal. Chem. 66 (1994) 3882 – 3888. [7] S.P.J. Higson, P.M. Vadgama, Anal. Chim. Acta 300 (1995) 77 – 83. [8] K. Nishida, M. Sakakida, K. Ichinose, T. Uemura, M. Uehara, K. Kajiwara, T. Miyata, M. Shichiri, K. Ishihara, N. Nakabayashi, Med. Prog. Technol. 21 (1995) 91 – 103. [9] N. Labat-Allietta, D. The´venot, Biosens. Bioelectron. 13 (1998) 19 – 29.

N. Wisniewski, M. Reichert / Colloids and Surfaces B: Biointerfaces 18 (2000) 197–219 [10] S. Eisele, H. Ammon, Biosens. Bioelectron. 9 (1994) 119 – 124. [11] S.M. Reddy, P.M. Vadgama, Anal. Chim. Acta 350 (1997) 77–89. [12] K. Kyrolainen, P. Rigsby, S. Eddy, P. Vadgama, Acta Anaesthesiol. Scand. 39 (1995) 55–60. [13] S. Woodward, Diabetes Care 5 (1982) 278–281. [14] W. Reichert, A. Sharkawy, in: A. von Recum (Ed.), Handbook of Biomaterials Evaluation, vol. 2, Taylor & Francis, Washington, DC, 1999, pp. 439–460. [15] A. Sharkawy, M. Neuman, W. Reichert, in: K. Park (Ed.), Controlled Drug Delivery: The Next Generation, ACS, Washington, DC, 1997, pp. 163–203. [16] R.B. Beard, B.N. Hung, R. Schmukler, Ann. Biomed. Eng. 20 (1992) 395–410. [17] M. Kyrolainen, S.M. Reddy, P.M. Vadgama, Anal. Chim. Acta 353 (1997) 281–289. [18] W. Kerner, M. Kiwit, B. Linke, F.S. Keck, H. Zier, E.F. Pfeiffer, Biosens. Bioelectron. 8 (1993) 473–482. [19] V. Kuberkar, P. Czekaj, R. Davis, Biotechnol. Bioeng. 60 (1998) 77–87. [20] Y. Maa, C. Hsu, J. Pharm. Sci. 87 (1998) 808 – 812. [21] L. Millesime, J. Dulieu, B. Chaufer, Bioseparation 6 (1996) 135–145. [22] H. Buntemeyer, C. Bohme, J. Lehmann, Cytotechnology 15 (1994) 243–251. [23] H. Ridgway, A. Kelly, C. Justice, B. Olson, Appl. Environ. Microbiol. 45 (1983) 1066–1084. [24] N. Torto, J. Bang, S. Richardson, G. Nilsson, L. Gorton, T. Laurell, M. Marko-Varga, J. Chromatogr. A 806 (1998) 265–278. [25] R. Boyd, A. Zydney, Biotechnol. Bioeng. 59 (1998) 451 – 460. [26] D. Lee, R. Merson, J. Dairy Sci. 58 (1975) 1423 – 1432. [27] M. Gerritsen, J.A. Jansen, A. Kros, R.J. Nolte, J.A. Lutterman, J. Investig. Surg. 11 (1998) 163–174. [28] M. Schlosser, M. Ziegler, in: D. Fraser (Ed.), Biosensors in the Body, John Wiley & Sons, New York, 1997, pp. 138 – 170. [29] W. Reichert, S. Saavedra, in: D. Williams (Ed.), Medical and Dental Materials, VCH Publishers, Inc, New York, 1992, pp. 303–343. [30] Wisniewski N, Moussy F, Reichert W, Fresenius J Anal Chem 2000 (in press). [31] N. Peppas, in: B. Ratner, A. Hoffman, F. Schoen, J. Lemons (Eds.), Biomaterials Science. An Introduction to Materials in Medicine, Academic Press, New York, 1996, pp. 60–64. [32] S. Murphy, C. Hamilton, M. Davies, B. Tighe, Biomaterials 13 (1992) 979–990. [33] Uragami T, Furukawa T, Sugihara M, Polym. Commun., 1984 (1984) 30–32. [34] S. Shamlou, J. Kennedy, R. Levy, Biomed. Mater. Res. 35 (1997) 157–163. [35] N.P. Desai, J.A. Hubbell, Biomaterials 12 (1991) 144 – 153.

217

[36] C.P. Pathak, A.S. Sawhney, J.A. Hubbell, J. Am. Chem. Soc. 114 (1992) 8311 – 8312. [37] N.A. Peppas, H.I. Moynihan, L.M. Lucht, J. Biomed. Mater. Res. 19 (1985) 397 – 411. [38] I. Gursel, V.N. Hasirci, Biomaterials 13 (1992) 150–155. [39] L. Doretti, D. Ferrara, Biosens. Bioelectron. 11 (1996) 365 – 373. [40] P. De Vos, B.J. De Haan, G.H. Wolters, J.H. Strubbe, R. Van Schilfgaarde, Diabetologia 40 (1997) 262 –270. [41] S.F. Hossainy, J.A. Hubbell, Biomaterials 15 (1994) 921 – 925. [42] E. Brinkman, L. van der Does, A. Bantjes, Biomaterials 12 (1991) 63 – 70. [43] M. Ishikawa, D. Schmidtke, P. Raskin, C. Quinn, J. Diabetes Complicat. 12 (1998) 295 – 301. [44] E. Csoregi, C. Quinn, D. Schmidtke, S. Lindquist, M. Pishko, L. Ye, I. Katakis, J. Hubbell, A. Heller, Anal. Chem. 66 (1994) 3131 – 3138. [45] L. Doretti, D. Ferrara, G. Barison, S. Lora, Appl. Biochem. Biotech. 49 (1994) 191 – 202. [46] K. Sikar, M.V. Pishko, Anal. Chem. 40 (1998) 2888– 2894. [47] G.V. Tiutrina, A.V. Zharikova, V.A. Zanin, T.T. Berezov, I. Mozer, G. Jobst, E. Aschauer, P. Svasek, M. Varahram, G. Urban, Vopr. Med. Khimii 43 (1997) 20 – 28. [48] D.W. Schmidtke, A. Heller, Anal. Chem. 70 (1998) 2149 – 2155. [49] C. Eggenstein, M. Borchardt, C. Diekmann, B. Grundig, C. Dumschat, K. Cammann, M. Knoll, F. Spener, Biosens. Bioelectron. 14 (1999) 33 – 41. [50] S. Yang, P. Atanasov, E. Wilkins, Biomed. Instrum. Technol. 30 (1996) 55 – 61. [51] B. McKinley, K. Wong, J. Janata, W. Jordan, D. Westenskow, Crit. Care Med. 9 (1981) 333 – 339. [52] G. Margules, C. Hunter, D. MacGregor, Med. Biol. Eng. Comput. 21 (1983) 1 – 8. [53] K. Shimada, M. Yano, K. Shibatani, Y. Komoto, M. Esashi, T. Matsuo, Med. Biol. Eng. Comput. 18 (1980) 741 – 745. [54] C. Quinn, R. Connor, A. Heller, Biomaterials 18 (1997) 1665 – 1670. [55] Lewis A, Colloids Surf. B 18 (2000) 261 – 275. [56] K. Ishihara, H. Nomura, T. Mihara, K. Kurita, Y. Iwasaki, N. Nakabayashi, J. Biomed. Mat. Res. 39 (1997) 323 – 330. [57] K. Ishihara, N. Nakabayashi, K. Nishida, M. Sakakida, M. Shichiri, New Biocompatible Polymer: Application for Implantable Glucose Sensor, American Chemical Society, Washington, DC, 1994, pp. 194 – 210. [58] S. Zhang, G. Wright, M. Kingston, P. Rolfe, Med. Biol. Eng. Comput. 34 (1996) 313 – 315. [59] Ishihara K, Nakabayashi N, Sakakida M, Kenro N, Shichiri M. American Chemical Society Annual Meeting, Biocompatible microdialysis hollow – fiber probes for long-term in vivo glucose monitoring, Orlando, FL, USA, 1998 p. 24 – 33.

218

N. Wisniewski, M. Reichert / Colloids and Surfaces B: Biointerfaces 18 (2000) 197–219

[60] M. Shichiri, Y. Yamasaki, K. Nao, M. Sekiya, N. Ueda, Horm. Metab. Res. (Suppl) 20 (1988) 17–20. [61] M. Shichiri, M. Sakakida, K. Nishida, S. Shimoda, Artif. Organs 22 (1998) 32–42. [62] E.J. Campbell, in: A. Ellerton (Ed.), Phosphorylcholine Surfaces: their Relative Biocompatibility and Evaluation, Advanstar Communications, Chester, UK, 1994, pp. 105 –124. [63] E.J. Campbell, Mater. Res. Soc. Symp. Proc. 252 (1992) 229 – 236. [64] H.P. Ammon, W. Ege, M. Oppermann, W. Gopel, S. Eisele, Anal. Chem. 67 (1995) 466–471. [65] C. Meyerhoff, F. Bischof, F. Sternberg, H. Zier, E. Pfeiffer, Diabetologia 35 (1992) 1087–1092. [66] A. Aalders, F. Schmidt, A. Schoonen, I. Broek, A. Maessen, H. Doorenbos, Int. J. Artif. Organs 14 (1991) 102 – 108. [67] E. Dempsey, D. Diamond, M.R. Smyth, M.A. Malone, K. Rabenstein, A. McShane, M. McKenna, T.V. Keaveny, R. Freany, Analyst 122 (1997) 185–189. [68] J. De Boer, H. Korf, H. Plijter-Groendijk, Int. J. Artif. Organs 17 (1994) 163–170. [69] R. Freaney, A. McShane, T.V. Keaveny, M. McKenna, K. Rabenstein, F.W. Scheller, D. Pfeiffer, G. Urban, I. Moser, G. Jobst, A. Manz, E. Verpoorte, M.W. Widmer, D. Diamond, E. Dempsey, F.J. Saez de Viteri, M. Smyth, Ann. Clin. Biochem. 34 (1997) 291–302. [70] J. Bolinder, U. Ungerstedt, P. Arner, Lancet 342 (1993) 1080 – 1085. [71] R. Lenigk, X. Zhu, T. Lo, R. Renneberg, Fresenius J. Anal. Chem. 364 (1999) 66–71. [72] P. Vadgama, Sens. Actuators B1 (1990) 1–7. [73] G. Rigby, P. Crump, P. Vadgama, Analyst 121 (1996) 871 – 875. [74] G. Rigby, P. Crump, P. Vadgama, Med. Biol. Eng. Comput. 33 (1995) 231–234. [75] G. Rigby, S. Ahmed, G. Horseman, P. Vadgama, Anal. Chim. Acta 385 (1999) 23–32. [76] Y. Hashiguchi, M. Sakakida, K. Nishida, T. Uemura, K. Kajiwara, M. Shichiri, Diabetes Care 17 (1994) 387 – 396. [77] F. Palmisano, D. Centonze, M. Quinto, P. Zambonin, Biosens. Bioelectron. 11 (1996) 419–425. [78] F.J. Schmidt, W.J. Sluiter, A.J. Schoonen, Diabetes Care 16 (1993) 695–700. [79] F.J. Schmidt, A.L. Aalders, A.J. Schoonen, H. Doorenbos, Int. J. Artif. Organs 15 (1992) 55–61. [80] U. Ungerstedt, J. Inter. Med. 230 (1991) 365– 373. [81] K. Wientjes, P. Vonk, Y. Vonk-van Klei, A. Schoonen, N. Kossen, Diabetes Care 21 (1998) 1481–1488. [82] E. Pfeiffer, Horm. Metab. Res. 26 (1994) 510 – 514. [83] M. Sudoh, T. Okouchi, K. Ohashi, M. Kimura, Kagaku Kogaku Ronbunshu 24 (1998) 190–194. [84] Z. Trajanoski, P. Wach, R. Gfrerer, G. Jobst, G. Urban, P. Kotanko, F. Skrabal, Biosens. Bioelectron. 11 (1996) 479 – 487.

[85] L. Schaupp, M. Ellmerer, G.A. Brunner, A. Wutte, G. Sendlhofer, Z. Trajanoski, F. Skrabal, T.R. Pieber, P. Wach, Am. J. Physiol. 276 (1999) E401 – 408. [86] Z. Trajanoski, G. Brunner, L. Schaupp, M. Ellmerer, P. Wach, T. Peiber, P. Kotanko, F. Skrabal, Diabetes Care 20 (1997) 1114 – 1121. [87] Ash S, Rainier J, Zopp W, Truitt R, Janle E, Kissinger P, Poulos J, ASAIO J., 1993 (1993) M699 – 705. [88] W. Kaptein, R. Kemper, M. Ruiters, K. Venema, R. Tiessen, J. Korf, Biosens. Bioelectron. 12 (1997) 967– 976. [89] D. Moscone, K. Venema, J. Korf, Med. Biol. Eng. Comput. 34 (1996) 290 – 294. [90] Towe B, http://enuxsa.eas.asu.edu/btowe/research/ glucose/rti-gluc.html, Arizona State University, web site accessed September 1999. [91] P. Fernandes, Curr Opin Chem. Biol. 2 (1998) 597–600. [92] M. Madaras, R. Buck, Anal. Chem. 68 (1996) 3832– 3839. [93] Y. Zhang, Y. Hu, G. Wilson, D. Moatti-Sirat, V. Poitout, Anal. Chem. 66 (1994) 1183 – 1188. [94] E. Wilkins, P. Atanasov, B. Muggenburg, Biosens. Bioelectron. 10 (1995) 484 – 494. [95] F. Moussy, D.J. Harrison, D.W. O’Brien, R.V. Rajotte, Anal. Chem. 65 (1993) 2072 – 2077. [96] H. Liu, J. Deng, Anal. Chim. Acta 300 (1995) 65 –70. [97] D.J. Harrison, R.F. Turner, H.P. Baltes, Anal. Chem. 60 (1988) 2002 – 2007. [98] S. Bruck, Blood Compatible Synthetic Polymers, Thomas, Springfield, IL, 1974. [99] F. Moussy, D.J. Harrison, R.V. Rajotte, Int. J. Artif. Organs 17 (1994) 88 – 94. [100] Moussy F, Mercado RC, Biomedical Engineering Society Annual Fall Meeting, Calcification induced degradation of Nafion membrane used for biosensor, San Diego, CA, USA, 1997, p. S-1. [101] R.C. Mercado, F. Moussy, Biosens. Bioelectron. 13 (1998) 133 – 145. [102] T. Valdes, F. Moussy, Biosens. Bioelectron. 000 (1999) 1 – 7. [103] J. Park, P. Tran, J. Chao, R. Ghodadra, R. Rangarajan, N. Thakor, Biosens. Bioelectron. 13 (1998) 1187 – 1195. [104] J. Neff, K. Caldwell, P. Tresco, J. Biomed. Mater. Res. 40 (1998) 511 – 519. [105] J. Lee, J. Kopecek, J. Andrade, J. Biomed. Mater. Res. 23 (1989) 351 – 368. [106] C. Espadas-Torre, M. Meyerhoff, Anal. Chem. 67 (1995) 3108 – 3114. [107] P. Bu¨hlmann, E. Pretsch, E. Bakker, Chem. Rev. 98 (1998) 1593 – 1687. [108] E. Bakker, P. Buhlmann, E. Pretsch, Chem. Rev. 97 (1997) 3083 – 3132. [109] Y. Benmakroha, I. Christie, M. Desai, P. Vadgama, Analyst 121 (1996) 521 – 526. [110] E. Lindner, V.V. Cosofret, S. Ufer, R.P. Buck, W.J. Kao, N.R. Neuman, J.M. Anderson, J. Biomed. Mater. Res. 28 (1994) 591 – 601.

N. Wisniewski, M. Reichert / Colloids and Surfaces B: Biointerfaces 18 (2000) 197–219 [111] H. Liu, Y. Liu, J. Qian, T. Yu, J. Deng, Microchem. J. 53 (1996) 241–252. [112] S. Yang, P. Atanasov, E. Wilkins, Biomed. Instrum. Technol. 29 (1995) 125–133. [113] J. Xu, B. Liu, Analyst 119 (1994) 1599–1601. [114] S.M. Reddy, P.J. Higson, I.M. Christie, P.M. Vadgama, Analyst 119 (1994) 949–952. [115] S.P.J. Higson, M.D. Desai, S. Ghosh, I. Christie, P. Vadgama, J Chem. Soc. Faraday Trans. 89 (1993) 2847 – 2851. [116] M. Shichiri, N. Asakawa, Y. Yamasaki, R. Kawamori, A. Hiroshi, Diabetes Care 9 (1986) 298–301. [117] G. Malsch, D. Paul, S. Kramer, Clin. Mater. 8 (1991) 63 – 69. [118] L. Lucchi, D. Bonucchi, M.A. Acerbi, G. Cappelli, A. Spattini, M. Innocenti, A. Castellani, E. Lusvarghi, Artif. Organs 13 (1989) 417–421. [119] J. Armour, J. Lucisano, B. McKean, D. Gough, Diabetes 39 (1990) 1519–1526. [120] J. Barisci, D. Hughes, A. Minett, G. Wallace, Anal. Chim. Acta 371 (1998) 39–48. [121] Q. Yang, P. Atanasov, E. Wilkins, Biomed. Instrum. Technol. 31 (1997) 54–62. [122] Y. Nakayama, T. Matsuda, ASAIO J. 38 (1992) M421 – 424. [123] S. Myler, S. Eaton, S.P.J. Higson, Anal. Chim. Acta 357 (1997) 1997. [124] B. Guo, J. Anzai, T. Osa, Pharm. Soc. Jpn. 44 (1996) 860 – 862. [125] F. Keedy, P. Vadgama, Biosens. Bioelectron. 6 (1991) 491 – 499. [126] P. Vadgama, G. Reach, Advances in Biosensors: Chemical Sensors for in vivo Monitoring, JAI Press, London, 1993. [127] B. Joon, R. Lakes, Biomaterials: an Introduction, Plenum Press, New York, 1992. [128] P. Treloar, I. Christie, P. Vadgama, Biosens. Bioelectron. 10 (1995) 195–201. [129] S.P.J. Higson, P.M. Vadgama, Anal. Chim. Acta 271 (1993) 125–133. [130] L. Tomson, F. Law, N. Rushton, Biomaterials 12 (1991) 37 – 40.

.

219

[131] S.P.J. Higson, P.M. Vadgama, Anal. Chim. Acta 300 (1995) 85 – 90. [132] Hopfenberg H, in K.R. et al., editors. Polymers in Medicine and Surgery. New York: Plenum Press, 1974, p. 99 – 107. [133] J.H. Brauker, V.E. Carr-Brendel, L.A. Martinson, J. Crudele, W.D. Johnston, R.C. Johnson, J. Biomed. Mater. Res. 29 (1995) 1517 – 1524. [134] A.A. Sharkawy, B. Klitzman, G.A. Truskey, W.M. Reichert, J. Biomed. Mater. Res. 37 (1997) 401 – 412. [135] A.A. Sharkawy, B. Klitzman, G.A. Truskey, W.M. Reichert, J. Biomed. Mater. Res. 40 (1998) 598 – 605. [136] A.A. Sharkawy, B. Klitzman, G.A. Truskey, W.M. Reichert, J. Biomed. Mater. Res. 40 (1998) 586 – 597. [137] S. Updike, M. Shults, R. Rhodes, in: D. Fraser (Ed.), Biosensors in the Body, John Wiley & Sons, New York, 1997, pp. 117 – 138. [138] A. Singh, Histochem. J. 25 (1993) 562 – 568. [139] S. Mathison, E. Bakker, J. Pharm. Biomed. Anal. 19 (1999) 163 – 173. [140] M. Mayer, J. Ruzicka, Anal. Chem. 68 (1996) 3808– 3814. [141] H. Tien, Z. Salamon, A. Ottova, Crit. Rev. Biomed. Eng. 18 (1991) 323 – 340. [142] J. Ramsden, G. Bachmanova, A. Archakov, Biosens. Bioelectron. 11 (1996) 523 – 528. [143] K. Toko, M.T.K. Yamafuji, K. Hayashi, H. Ikezaki, K. Sato, R. Toukubo, S. Kawarai, Biosens. Bioelectron. 9 (1994) 359 – 364. [144] K. Healy, A. Rezania, R. Stile, Ann. NY Acad. Sci 875 (1999) 24 – 35. [145] H. Shi, W.B. Tsai, M.D. Garrison, S. Ferrari, B.D. Ratner, Nature 398 (1999) 593 – 597. [146] H. Bayley, Curr. Opin. Biotechnol. 10 (1999) 94 –103. [147] C.P. Quinn, C.P. Pathak, A. Heller, J.A. Hubbell, Biomaterials 16 (1995) 389 – 396. [148] S. Zhang, Y. Benmakroha, P. Rolfe, S. Tanaka, K. Ishihara, Biosens. Bioelectron. 11 (1996) 1019 – 1029. [149] B. Towe, V. Pizziconi, Biosens. Bioelectron. 12 (1997) 893 – 899. [150] M. Kyrolainen, H. Hakanson, B. Mattiasson, P. Vadgama, Biosens. Bioelectron. 12 (1997) 1073 – 1081.