In conjunction with Foresight Update 42
Will Serum Proteins Stick to Nanorobot Surfaces?
By Robert A. Freitas Jr.
Research Scientist, Zyvex Corp.
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| Robert A. Freitas Jr. |
When an artificial nanoorgan is implanted in the body, it may be desirable to promote rapid adhesion to cells and tissues. But for medical nanorobots floating or swimming in the circulation such adhesion will not normally be desirable. Thus one preliminary question is whether sticky biological "gunk" will adhere to the surface of a diamondoid nanorobot when it is placed in the bloodstream, and if so, what can be done about it? As the famous physiological chemist Leo Vroman once hyperbolized [1]: "Facing a hail of miscellaneous eggs, we cannot expect to come away clean. Unless they are hard-boiled ones, we are most likely to become coated rapidly with a relatively thin film of matter from the most numerous and most fragile eggs. Similarly, no interfaces may exist that, facing blood plasma, can escape being coated with the most abundant and fragile plasma proteins."
When a foreign material is implanted into a host tissue, the first event to occur at the tissue-material interface which dictates biocompatibility is the noncovalent adsorption of plasma proteins from blood onto the surface [2, 3]. Protein adsorption is much more rapid than the transport of host cells to foreign surfaces. Once proteins have adsorbed to the surface of the foreign material, host cells no longer see the underlying material, but only the protein-coated surface overlayer. This adsorbed protein overlayer — rather than the foreign material itself – then mediates the types of cells that may adhere to the surface, which ultimately can determine the type of tissue that forms in the vicinity [2]. Thus the type and state of adsorbed proteins, including their conformational changes, will be critical determinants of biocompatibility [3-5], including nanorobot-cell interactions and nanorobot surface fouling.
Even by the late 1960s, Vroman and Adams [1], and Baier and Dutton [6] had found that within 10 seconds of exposure to blood or plasma, a uniform ~6 nm layer of fibrinogen formed on surfaces of Ge, Pt, Si, and Ta. After 60 sec, the layer was less uniform and averaged ~12.5 nm thick, but was still dominated by fibrinogen. Rudee and Price [7] determined that human serum albumin (molecular dimensions 8 nm x 3.8 nm) formed a continuous film on amorphous carbon surface in only 1.3 sec of exposure. Fibrinogen required 2.5 sec to form films.
The first direct study of protein adsorption on diamond, done by Tang et al. [3] in 1995, focused on fibrinogen. Fibrinogen, a 340-kilodalton soluble plasma glycoprotein ~47.5 nm in length, is the major surface protein to initiate coagulation [5] (via platelet adhesion to fibrinogen) and inflammation including fibrosis [4] around implanted biomaterials. The adsorption and deformation ("denaturation") of adsorbed fibrinogen molecules is commonly used as a biocompatibility indicator. The amounts of denatured fibrinogen accumulated on surfaces correlates closely with the extent of biomaterial-mediated inflammation [8].
Accordingly, Tang and colleagues [3] measured ~3.7 mg/m2 (~6600 molecules/micron2) gross surface adsorption of human fibrinogen on chemical-vapor-deposited (CVD) diamond surfaces, after incubation of the plasma-coated diamond surface in a 20 microgram/cm3 fibrinogen solution (~0.1% of blood concentration [9]) for 8 hours at room temperature. Much of this adsorbed fibrinogen was only loosely bound, however. A solution of sodium dodecyl sulfate (an anionic detergent or surfactant commonly used to solubilize proteins) was rinsed over the incubated CVD surface to remove the loosely-bound or elutable (non-denatured) fibrinogen and ~48% of the molecules detached, leaving ~2.1 mg/m2 (~3700 molecules/micron2) of spontaneously denatured fibrinogen still present on the CVD diamond surface. These adsorbed amounts are comparable to the amounts adsorbed on titanium and stainless steel materials widely used as surgical implants that are regarded as "biocompatible."
However, CVD diamond might not accurately represent the atomically-smooth flawless diamond surfaces which may characterize the typical MNT-manufactured medical nanorobot exterior. Far from being atomically smooth, CVD diamond films are amorphous and polycrystalline [10], often with grain sizes up to 1-10 microns. In Tang’s experiment, diamond wafers with two distinct sides were tested, as follows:
The nucleation side of the diamond wafers was grown in contact with a flat silicon substrate, which was then dissolved away by acid. The formation of SiC on such a substrate allows silicon to bond well with carbon during the growth process [10]. However, the presence of small amounts of surviving carbide in the nucleation diamond surface, or of concave nanoscale surface features recording the removal of SiC by etchant, could markedly alter the protein adsorbent characteristics of the diamond surface at the molecular level. Also, SiC is tolerated by cells up to 0.1 mg/cm3 concentration but is cytotoxic at 1 mg/cm3 [11]. Furthermore, a contact profilometer measured the nucleation surface as having a rugosity of up to 250 nm, a roughness 100-1000 times greater than that which may be expected at the surface of the typical diamondoid medical nanodevice.
The growth side of the diamond wafers used in Tang’s experiment was much rougher than the nucleation surfaces, so this surface was ground and polished but only to a rugosity of ~1 micron peak-to-valley – roughly the diameter of an entire bloodborne medical nanorobot and clearly not representative of an atomically-precise engineered medical nanodevice surface. There is no indication whether the grinding and polishing of the growth surface was done under oil (thus preserving a predominantly hydrogen-terminated, hence strongly hydrophobic, surface [12]), and there was no evaluation of whether subsequent etching with H2SO4 and H2O2 might have produced carbonyl and hydroxyl conversions at the surface (thus possibly creating regions of hydrophilicity).
Furthermore, diamond crystals are believed to polish by successive repeated microcleavage along {111} planes, which is why polishing is much easier in some directions than others [12]. Non-{111} surfaces, when mechanically polished, will always be rough and will consist of small domains of {111} surface canted at appropriate angles to the macroscopic orientation [12] – residual asperities of ~5 nm have been reported even for extremely carefully polished surfaces. The general conclusion is that the chemical and mechanical processes used in Tang’s experiment seem unlikely to have produced a surface that is well-characterized at the molecular level. Protein adhesion to near-atomically smooth diamond surfaces remains to be investigated experimentally, and may be quite low.
Once the precise molecular mechanisms of protein adhesion are fully understood (this knowledge was still incomplete in 2000), nanodevice surfaces can be altered for maximum proteinophobicity if this is required for a particular medical mission. It is almost certain that such designs are possible because numerous semi-proteinophobic artificial molecular surfaces are already available.
For instance, a polyethylene glycol (PEG) coating on a 200-nm poly(lactic acid) (PLA) nanosphere surface creates a brushlike steric barrier, hindering its opsonization and uptake by the mononuclear phagocyte system, thus increasing its bloodstream half-life [13]. These "pegylated" nanospheres have been investigated as an injectable blood-persistent system for controlled drug release, for site-specific drug delivery, and for medical imaging [14]. Adsorption of human serum albumin (MW = 66 kilodaltons) on pegylated nanosphere surfaces at pH 7.4 at equilibrium (i.e., after 5 days) is 0.15 mg/m2 (~1400 molecules/micron2) compared to 2.2 mg/m2 (~20,000 molecules/micron2) for unpegylated polymer [14], which is of the same order of magnitude as that observed for other unmodified hydrophobic surfaces [15]. Under in vitro conditions at 37o C and pH 7.4, about one-third of the adsorbed PEG detaches from the PLA nanospheres after 2 weeks, at a near-linear detachment rate [16]. Tetraethylene glycol dimethyl ether glow-discharge plasma-deposited surfaces can reduce fibrinogen adsorption to ~0.2 mg/m2 (~350 molecules/micron2) on many different substrates [17].
Another even more effective way to create nonadhesive nanorobot surfaces may be the biomimetic approach. For example, the external region of a cell membrane, known as the glycocalyx, is dominated by glycosylated molecules, which direct specific interactions such as cell-cell recognition and contribute to the steric repulsion that prevents undesirable non-specific adhesion of other molecules and cells. Holland and colleagues [18] have modified a pyrolytic graphite surface by attaching oligosaccharide surfactant polymers which, like a glycocalyx, provides a dense and confluent layer of oligosaccharides that mimics the non-adhesive properties of a glycocalyx. The surfactant polymers consist of a flexible poly(vinyl amine) backbone (MW ~ 6000 daltons, diameter 0.25 nm) with multiple randomly-spaced dextran (MW ~ 1600 daltons, diameter ~0.9 nm) and alkanoyl (hexanoyl or lauroyl) side chains which constrain the polymer backbone to lie parallel to the substrate. Solvated dextran side chains protrude into the aqueous phase, creating a glycocalyx-like monolayer coating 0.7-1.2 nm thick as measured by tapping-mode AFM [18]. Dextran has a stable helical structure; steric repulsion between adjacent dextrans is believed to produce a brush-like conformation. In vitro experiments show that the resulting biomimetic surface, which the authors assert undergoes spontaneous adsorption on diverse hydrophobic surfaces, is effective in suppressing at least ~90% of all plasma protein adsorption from human plasma protein solution [18]. According to the authors:
"The steric barrier provided by the highly hydrated dextrans is designed to suppress non-specific adsorption of plasma proteins [19], whereas the high energy of desorption and low water solubility of the adsorbed surfactant polymer is designed to minimize possible displacement or exchange reactions with highly surface-active plasma proteins."
Diamond particles have already been encapsulated inside stealth liposomes. Stealth liposomes are relatively nonadhesive polyhydroxylated species that exhibit reduced recognition and uptake by the body’s reticuloendothelial system, and give a circulation half-life of ~1 day [20]. In one experiment to test a possible artificial oxygen carrier [21], hemoglobin molecules were irreversibly adsorbed onto carbohydrate-coated diamond particles measuring ~75 nm in diameter, then were encapsulated in a standard mixture of phospholipids, yielding preparations of spherical liposomes which were stable for >48 hours with bound-Hb concentrations near 100 gm/liter with as little as 1% free Hb. Evaluation of oxygen lability showed normal sigmoidal O2 binding behavior with p50 from 12 mmHg up to 37 mmHg under control of an allosteric effector [21].
Whether pure atomically-smooth diamondoid materials will give us sufficiently nonadhesive surfaces, or if instead thin engineered coatings or active semaphoric surfaces [9] will be necessary to ensure adequate biocompatibility of medical nanorobots, is an outstanding research issue that can best be resolved by future experiments. This is a very critical topic because, unlike the materials used in a joint prosthesis, nanorobots will be present in the microvasculature of critical organs. The adhesiveness of many hundreds of serum proteins to the artificial nanorobot calyx must be assayed, and the relative serum concentration of these proteins changes according to the time of day or the physiological state of the individual (e.g., TNF, IL-1, IL-2, and transferrin rise dramatically in the acute phase response to a pathogen). Relevant investigations are to be encouraged at the earliest possible opportunity. The author welcomes and invites further dialogue on this subject by interested specialists.
© 2000 by Robert A. Freitas Jr.
Acknowledgments
Thanks to Stephen S. Flitman, M.D., and C. Christopher Hook, M.D., for
helpful comments on an earlier version of this paper.
References
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