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Published: April 1, 2010
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Medical Coatings and Surfaces: Where Worlds Collide

This review of surface solutions describes some of the latest technologies available to improve device performance. These include a non-biofouling technology that offers mechanical durability and the inhibition of microbial adhesion, as well as technology that modifies rather than coats the polymer from which a device is fabricated.

By: J. Thies,
DSM Biomedical, Geleen, The Netherlands

Technology to manage the interface
Whenever a medical device comes into direct contact with living tissue or body fluids, the initial interactions between the device and the biological material occur at the interface: the surface. Thus, managing the interface where these “worlds collide” is the primary role of medical coatings and is arguably one of their most important performance attributes. When considering coatings for medical devices one can segment them into four major classes: passivating (compatibilising), physico–mechanical, antimicrobial and (drug) delivery. Because the area of drug delivery, particularly in relation to drug-eluting stents (DES), is enormous as well as complex, it is better treated separately in-depth and as such it has been omitted from this discussion. Readers with a particular interest in the DES coating field are directed to the excellent recent review of Pendyala et al.1
 
Passivating coatings and surfaces
Figure 1: Schematic representation of a novel non-biofouling coating approach: Step 1 Colloidal silica particles (diameter 15 nm) are covalently modified with acrylate functional groups; Step 2 Hydrophilic polymer chains are covalently grafted on the remaining available silica surface; Step 3 The liquid coating formulation (including other additives not shown) is deposited on the surface, the solvent (alcohol-water mixture) is evaporated and the coating crosslinked by UV irradiation. 
Artificial materials such as polymers, ceramics and metals broadly suffer from similar disadvantages when applied in medical devices. Chief among these is an inherent incompatibility with blood or tissue. For example, in blood, interaction with a polymeric surface leads to an activation of the cellular defence and humoral systems.2 Current understanding states that adsorption of plasma proteins onto the artificial surface is a critical contributor to subsequent thrombogenicity.3 Even in the absence of permanent adsorption to the surface, the conformation of the proteins can be altered by interaction with the surface and this may cause undesired adverse events downstream. Thus, the ability to control, delay or ideally inhibit the adsorption of proteins on the material substrate is a central concept in passivating coatings and surfaces.
 
In disposable plastic diagnostic and research devices such as blood collection tubes and micro-titre plates this has so far mainly been achieved by the adsorption of polymeric surfactants displaying hydrophilic surface groups. The resultant increased hydrophilicity minimises the hydrophobic–hydrophobic interactions between the substrate and the (segments of) proteins. Under relatively static application conditions these surfactants have been shown to be fairly effective at reducing surface mediated haemolysis and/or protein adsorption. However polymeric surfactants have the potential to be desorbed (leached) into the surrounding medium and this has lead to interference problems in subsequent clinical immunological assays performed on exposed serum.4,5,6
 
For ex vivo and in vivo devices such as extracorporeal (circulation) devices and intravascular stents, passivation has, for example, been achieved by application of heparin surfaces. Heparin is a highly sulphated glycosamino-glycan and is widely used as an injectable anticoagulant. Medical-grade heparin is derived from mucosal porcine and bovine tissue,7 thus it requires highly stringent quality control procedures, as recently highlighted by United States Food and Drug Administration concerns over contaminated Chinese sources of heparin (www.fda.gov/NewsEvents/Testimony/ucm115242.htm). The predominant method for fixing a heparin layer onto the device surface is by end-point immobilisation of the heparin (Carmeda, www.carmeda.se).8 The early hypothesis held that the reduced thrombogenicity of heparinised surfaces was exclusively due to its influence on antithrombin III activity, however, current thinking is that heparinised surfaces rely mainly on reduction, or selective adsorption, of blood plasma proteins for their performance.
 
Figure 2: Adsorption human blood plasma (100 × diluted) by stagnation point reflectometry26 on a bare silica surface of a silicon wafer (blue) and on novel coating (orange).
Other passivating coating systems include biomimetic approaches such as the use of phosphorylcholine containing surfaces (Biocompatibles International plc, www.biocompatibles.com). Chapman9 employed these phospholipids from natural membranes to coat substrates. Since then, these surfaces and other phosphorylcholine containing materials such as copolymers of 2-methacryloyl oxyethyl phosphorylcholine and alkyl methacrylates,10 have been shown to achieve haemocompatibility by reduced protein adsorption.
 
Other novel passivating approaches do not require a coating at all, but rather aim to manage the surface by the use of surface modifying additives (COBE Cardiovascular Inc. (www.soringroup-usa.com)11,12 surface modifying end (SME) groups, or the related self-assembling monolayer end groups (SAME technology) (DSM PTG) in the bulk polymer. In the case of the surface modifying additives, the modified surface comprises alternating micro-domains of hydrophilic and hydrophobic regions leading to a particular orientation of the adsorbed fibrinogen “mosaic,” which is hypothesised to reduce subsequent protein expression.
 
SAME technology modifies a structural polymer (for example, Bionate II PCU) by permanently binding special end groups onto the polymer molecules while it is being made. When the modified polymer is formed into a device, the end groups migrate and self assemble on the surface to form a molecular monolayer optimised for the application. These approaches have been shown to reduce (delay) protein adsorption13 and reduce thrombogenicity. For excellent reviews of surfaces for extracorporeal circulation see Wendel and Zieme,14 and for medical implants the reader is directed to Tanzi.15
 
Because all of the approaches discussed so far are comprised of thin hydrogel-like layers and/or adsorbed or surface grafted monolayers, the mechanical durability is not optimal. Coated device surfaces can be damaged during final assembly and/or packaging, which can potentially lead to localised loss of passivating (non-biofouling) behaviour.16 Recently a non-biofouling technology (VitroStealth, DSM) was unveiled. The coatings are based on colloidal silica particles (15 nm diameter), which are surface modified with acrylate groups and further surface modified with a hydrophilic polymer brush-like structure (Figure 1).17 After applying the coating to the substrate and subsequent ultraviolet crosslinking, the resultant surfaces demonstrate a dramatic reduction in protein adsorption (Figure 2) combined with remarkable mechanical durability.17 As a result, these coatings may be highly suitable as robust passivating coatings.
 
Physico–mechanical coatings
Another class of device coatings are those designed with physico–mechanical properties such as chemical, moisture and electrical barrier properties as well as tribological (lubricous) properties). Hydrophilic lubricous coatings such as ComfortCoat and PhotoLink (Hydromer Inc., www.hydromer.com) are nowadays applied to cardiovascular guidewires and guiding catheters, as well as urinary catheters.18 Hydrophilic coatings are generally known to have reduced protein adsorption compared with hydrophobic materials. They have also been shown to reduce the damage to the mucosal lining,19,20 leading to less haematuria as well as improved patient comfort.21
 
Antimicrobial coatings
The damage to the mucosal lining mentioned above has been cited as a focus of infections and has lead to the development of dual-function coatings with both physico–mechanical function (lubricity) and biological function (antimicrobial activity). These systems such as ComfortCoat are suitable for urological22 and (haemocompatible) vascular catheters.23
 
Figure 3: Adhesion of Staphylococcus epidermidis (HBH 276) under flow conditions;27 left: number of bacterial cells per square centimeter as a function of flow time for both uncoated glass (blue) and on the novel coated glass (green); right: micrograph of S. epidermidis adhering after four hours of flow.
Most current antimicrobial strategies rely on biocidal activity.24,25 In the formation of a microbial biofilm on the surface of a medical device, the first step is the rapid deposition of a conditioning layer of macromolecular species such as proteins, followed by bacterial cellular adhesion to this conditioning layer. Recent novel approaches have been proposed to hinder the formation of these biofilms by eliminating the formation of the conditioning layer thereby surpassing subsequent cellular adhesion events (Nerites Corp., www.nerites.com and DSM). The non-biofouling coating discussed above reduces the adsorption of proteins on the surfaces and as a result also effectively inhibits microbial adhesion (Figure 3). After approximately three hours of continual flow, uncoated glass shows approximately 20 million Staphylococcus epidermidis cells per square centimeter; the coated surface under the same conditions remains pristine. Current commercialisation of this coating technology is in the area of life science consumables and disposable devices for in vitro diagnostics used for example in pre-analytical blood collection and micro-fluidic point-of-care biosensors. However, given the general attractive properties of suppression of protein adsorption and cellular adhesion other applications in extracorporeal and in vivo medical devices are envisaged.
 
Take a tool box approach
“If the only tool you have is a hammer, you tend to see every problem as a nail,” said the psychologist Abraham Maslow. Given the growing number and complexity of medical devices, as well as the increase in various forms of treatments and fields of application, it seems logical that there can be no single “tool” for every surface “problem.” Consequently, a toolbox of biomaterials solutions is required to manage the interface at which the world of the device surface and the living system
collide. This will lead to better device performance and ultimately improved patient outcomes.
 
References
1. L. Pendyala et al., “Passive and Active Polymer Coatings for Intracoronary Stents: Novel Devices to Promote Arterial Healing,” J. of Interv. Cardio., 22, 1, 37 (2009).
2. K.M. Taylor, “SIRS: The Systemic Inflammatory Response Syndrome after Cardiac Operations,” Ann. Thorac. Surg., 61, 1608 (1996).
3. L. Vroman, “The Life of an Artificial Device in Contact with Blood: Initial Events and Their Effect on its Final State,” Bull NY Acad. Med. 64, 352–357 (1998).
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6. L.J. Krickaa and J.Y. Park, “Additive-Aggravated Assays: An Authoritative Answer,” Clin. Chem., 51, 1767 (2005).
7. R.J. Linhardt and N.S. Gunay, “Production and Chemical Processing of Low Molecular Weight Heparins,” Sem. Thromb. Hem., 3, 5–16 (1999).
8. O. Larm et al., “A New Non-Thrombogenic Surface Prepared by Selective Covalent Binding of Heparin via a Modified Reducing Terminal Residue,” Biomater. Med. Devices Artif. Organs, 11, 161–173 (1983).
9. D. Chapman, “Biomembranes and New Hemo Compatible Materials,” Langmuir, 9, 39–45 (1993).
10. K Ishihara et al., “Preparation of 2-Methacryloyl Oxyethyl Phosphorylcholine Copolymers with Alkyl Methacrylatyes and Their Blood Compatibility,” J. Polymer, 24, 1259 (1992).
11. Y.J. Gu et al., Cardiopulmonary Bypass Circuit Treated with Surface-Modifying Additives: A Clinical Evaluation of Blood Compatibility, Ann Thorac. Surg., 65, 1342–1347 (1998).
12. J.O. Defraigne et al., “SMA Circuits Reduce Platelet Consumption and Platelet Factor Release During Cardiac Surgery,” Ann. Thorac. Surg,. 70, 2075–2081 (2000).
13. Z. Chen et al., “Interaction of Fibrinogen with Surfaces of End-Group-Modified Polyurethanes: A Surface-Specific Sum-Frequency-Generation Vibrational Spectroscopy Study,” J. Bone and Mineral Research, Part A 62, 2, 254–264 (2002).
14. H.P. Wendel and G. Ziemer, “Coating-Techniques to Improve the Hemocompatibility of Artificial Devices Used for Extracorporeal Circulation,” Europ. J. Cardio-Thoracic Surgery, 16, 342–350 (1999).
15. M.C. Tamzi, “Bioactive Technologies for Hemocompatibility, Expert Review of Medical Devices, 2, 4, 473–494 (2005).
16. A. Roosjen et al., “ Stability and Effectiveness Against Bacterial Adhesion of Poly(Ethylene Oxide) Coatings in Biological Fluids,” J. Bone and Mineral Research, Part B, 73, 2, 347–354 (2005).
17. P.F. Holmes et al., “Surface-Modified Nanoparticles as a New Versatile, and Mechanically Robust Nonadhesive Coating: Suppression of Protein Adsorption and Bacterial Adhesion,” J. Bone and Mineral Research, Part A, 91A, 3, 824 (2009).
18. M.M.Tunney and S.P. Gorman, “Evaluation of a Poly(Vinyl Pyrollidone)-Coated Biomaterial for Urological Use,” Biomaterials, 23, 4601–4608 (2002).
19. S. Vaidyanathan et al., “Urethral Cytology in Spinal Cord Injury Patients Performing Intermittent Catheterisation,” Paraplegia, 32, 493–500 (1994).
20. J.M.Vapnek et al. “Prospective Randomised Trial Of The Lofric Hydrophilic Coated Catheter Versus Conventional Plastic Catheter For Clean Intermittent Catheterisation,” J. Urol., 169, 994–998 (2003).
21. J. Stensballe et al., “Hydrophilic-Coated Catheters for Intermittent Catheterisation Reduce Urethral Micro Trauma: A Prospective, Randomised, Participant-Blinded, Crossover Study of Three Different Types of Catheters,” European Urology, 48, 978–973 (2005).
22. A.A. Dias et al., DSM Biomedical Materials Presentation “Anti-Microbial Coatings for Urological Applications,” BioInterface 2007, San Mateo, California, USA (www.surfaces.org).
23. A.K Saxena and B.R. Panhotra, “Prevention of Catheter-Related Bloodstream Infections: An Appraisal of Developments in Designing an Infection-Resistant Dream Dialysis-Catheter,” Nephrology, 10, 240–248 (2005).
24. G.J. Gabriel et al.,“Infectious Disease: Connecting Innate Immunity to Biocidal Polymers,” Materials Science and Engineering, R 57, 28–64 (2007).
25. E.M. Hetrick and M.H. Schoenfisch, “Reducing Implant-Related Infections: Active Release Strategies,” Chem. Soc. Rev., 35, 780–789 (2006)
26. J. Buijs et al., “Adsorption Dynamics of IgG and its F(ab’)(2) and Fc Fragments Studied by Reflectometry,“ J. Coll. Interf. Sci., 178, 594–605 (1996).
27. H.J. Busscher and H.C. Van der Mei, “Microbial Adhesion in Flow Displacement Systems,” Clin. Microbiol. Rev. 19, 127–141 (2006).
 
Jens C. Thies
is Business Development Manager and Prinicpal Scientist at DSM Biomedical, Koestraat 1, NL-6167 RA Geleen,
The Netherlands, tel. +31 464 761 313
e-mail: jens.thies@dsm.com
www.dsmbiomedical.com 

 

Published in European Medical Device Technology, April 2010, Volume 1, No. 4



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