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In conjunction with Foresight Update 47

Volumetric Cellular Intrusiveness of Medical Nanorobots

By Robert A. Freitas Jr.
Research Scientist, Zyvex Corp.

Robert A. Freitas Jr.

Medical nanorobots on cytotherapeutic missions will often need to enter the living cell to perform repairs. Such missions may require the participation of many cooperating nanorobots, or perhaps just a few but relatively large nanorobots, per cell. And so the question logically arises: How many nanorobots can safely be crammed into a single living cell? There are at least two issues here. First, how much new foreign material can be added to a cell? Second, how much of a cell’s existing volume can be replaced with foreign material with no change in total cell volume? Since we can’t yet do direct experiments with medical nanorobots and living cells, no precise answer is possible. But we can estimate the maximum volume of foreign material that the intracellular compartment might safely accommodate by examining analogous instances of cellular intrusion.

We can start by considering cell membrane elasticity. How much can the cellular membrane stretch before it breaks? The introduction of foreign material into a cell may cause intracellular volume to expand. Assuming a spherical cell shape, the change in cell volume Δ V_cell from the original cell volume V_cell is related to the change in plasma membrane area Δ A_cell of an unstretched membrane of area A_cell by the relations Δ V_cell /V_cell ~ ((1 + Δ A_cell/A_cell)3/2 – 1) and Δ A_cell/A_cell = T_memb / K_memb, where T_memb is the isotropic tension due to membrane expansion and the area compressibility modulus K_memb = 0.378 N/m for erythrocyte plasma membrane at 310 K and K_memb = 0.636 N/m for leukocyte plasma membrane ([1], Section 9.4.3.2.1). Taking a conservative lysis limit of T_memb ~ 4 x 10^-3 N/m for erythrocytes, then Δ V_cell /V_cell ~ 1.6% for red cells and ~0.9% for white cells. That’s not much stretch.

However, erythrocytes are not spheres but biconcave disks with a mean volume of 94 micron³ in isotonic solution (300 mosmol). Red cells absorb water in hypotonic solution, becoming spherical at 131 mosmol with a volume of 164 micron³, demonstrating a capacity for volumetric expansion of 74% without losing membrane integrity. Other cells may permit even more expansion. For example, taking T_memb = 1.7 N/m and K_memb = 1.3 N/m for TB/C3 hybridoma cells and T_memb = 1.8 N/m and K_memb = 1.2 N/m for NS1 myeloma cells [2], then Δ V_cell /V_cell ~ 250% for hybridomas and ~300% for myelomas. These estimates are crude at best because the lipid population of the plasma membrane is constantly changing and may enlarge or contract over time.

For more than four decades, microbiologists have routinely extracted from or inserted an entire nucleus into a cell using micropipettes without compromising cell viability. Such nuclear transplantation represents a volumetric change of Δ V_cell /V_cell ~ 3-4% for the typical 20-micron human tissue cell ([1], Table 8.17) but in the case of a human leukocyte would represent a volumetric change of Δ V_cell /V_cell ~ 18% for an eosinophil, 22% for a neutrophil, 26% for a monocyte, or 51% for a lymphocyte. Decades of laboratory practice have confirmed that at least ~10^-10 ml/cell or ~100 micron³/cell of foreign material (representing perhaps 1-3% of cell volume) can be safely injected into a somatic cell without any significant effect on cell viability [3]. Finally, we know that neutrophils increase volume by ~15% when stimulated in suspension, but rabbit neutrophils that migrate into the abdominal wall (cell size ~150 micron³) are +50% larger than those in the abdominal wall vasculature (~100 micron³) and human neutrophils induced by FMLP to migrate into collagen gels (~290 micron³) are 42% larger than those that did not migrate (~204 micron³) [4].

Another classic measure of tolerable volumetric intrusiveness in the context of medical nanorobots [5] is the amount of lipofuscin that can accumulate inside cells without ill effect. Lipofuscin consists of insoluble age-pigment lysosomal granules that collect in many of our cells, the accumulation starting as early in life as 11 years of age and rising with age, activity level and caloric intake, and varying with cell type. Clumps of these yellow-brown granules — typically 1-3 microns in diameter — may occupy up to 10% of the volume of heart muscle cells [6], and from 20% of brainstem neuron volume at age 20 to as much as 50% of cell volume by age 90 [7]. Lipofuscin concentrations as high as 75% have been reported in Purkinje neurons of rats subjected to protein malnutrition [8]. Elevated concentrations in heart cells appear not to increase the risk of heart attack and brain cell lipofuscin is not associated with any mental or motor abnormalities or other detrimental cellular function, although hereditary ceroid lipofuscinosis can lead to premature death. The fact that lipofuscin is an indigestible lipid peroxidation product that cannot be excreted but whose presence is usually not injurious to the cell argues strongly that cells can tolerate major volumetric replacements of protoplasm with artificial foreign bodies such as medical nanorobots while continuing to function normally. Other inert intracellular pigments are known along with a number of pathological intracellular storage diseases including Gaucher’s, Niemann-Pick, and Tay-Sachs. Noninert amyloid deposits average ~12% of pancreatic islet cell volume in patients with maturity onset diabetes [9].

Various particulate substances including microspheres and crystals have been introduced intracellularly to observe the effects on the cell. In one study [10], up to 500 polystyrene 0.26-micron beads were injected into a tissue cell and this 4.6 micron³ load did not affect the cell’s ability to transport the particles around inside as if they were tiny organelles or vesicles. A few micron³/cell of engineered nanoparticles are tolerated by living cells when employed as intracellular fluorescent reporters [11]. Cholesterol crystals have been induced to grow inside living J774 mouse peritoneal macrophages, reaching a concentration of ~120 mg cholesterol/mg protein or ~2.4% intracellular crystals by volume [12] without lethality, though excessive intracellular crystallization (e.g., of drug molecules) can lead to problems such as acute renal failure and intracellular crystals have been found inside chondrocytes in certain crystal deposition diseases. A useful and simple experiment that could be done today would be to microinject various cell types with progressively larger loads of chemically inert diamond microparticles, noting the effect (if any) on cell motility, behavior, and metabolic function.

Interestingly, Pseudomonas stutzeri AG259, a species of bacterium isolated from silver mines, protects itself from the usual bactericidal effect of silver ions by sequestering triangular and hexagonal insoluble nanocrystals of Ag⁰ and Ag₂S (believed to be acanthite, a stable crystalline form of silver sulfide) intracellularly in vacuole-like granules in the periplasmic space [13]. In one photomicrograph, several crystals ranging from 90-200 nm in diameter are visible inside a living bacterial cell ~800 nm in diameter, suggesting a total inert particulate ~13% volumetric intrusiveness.

We can obtain yet another perspective on cellular intrusiveness by considering the phagocytes — specialized cells optimized for ingestion of foreign particles [14]. Latex or polystyrene beads are among the most popular particles for phagocyte ingestion burden experiments. Guinea pig neutrophils can ingest up to 3.8% of cell volume in 3-micron polystyrene beads, but only 3.0% of cell volume of 0.3 micron beads. Peritoneal phagocytes from striped bass each ingested an average of four 3-micron latex beads during a 30-minute incubation time [15], giving a phagocytic capacity of ~64 micron³/phagocyte or ~4% of cell volume. Rabbit alveolar macrophages cultured in suspensions or on monolayers of latex particles internalized a maximum of 45 1-micron particles (45 micron³/cell or ~3% of cell volume) and 10 2-micron particles (~80 micron³/cell or ~5% of cell volume) at saturation [16]. Another study of rat alveolar macrophages confirmed particle burdens exceeding 15 2-micron microspheres (~63 micron³ or ~4% of cell volume) [17]. Murine bone-marrow macrophages that are only 13.8 microns in diameter can ingest IgG-opsonized beads up to 20 microns in diameter [18], representing an amazing ~200% of cell volume. Of course, phagocytes that eat too many latex microspheres develop impaired mobility, and some particles are highly toxic to phagocytes — just 0.05 µg of silica per 10⁶ macrophages [19], or 0.002% of cell volume assuming 1166 micron³ per rat alveolar macrophage, is cytotoxic.

What about inorganic particles? Rat alveolar macrophages ingested at least ~1 micron³/cell of iron oxide particles (~0.1% of cell volume) without ill effect in one experiment [20], but another experiment [21] found up to 72 spherical 2.6-micron iron oxide particles (~663 micron³) had been nonfatally ingested by human alveolar macrophages each of mean volume 4990 micron³, a cell burden of ~13% foreign particles by volume. Murine macrophages suffer only ~10% mortality from ingesting up to 2500 alumina ceramic (sapphire) 0.6-micron particles, or ~10% of cell volume, although mortality rises to ~30% from ingesting a similar volume concentration of 2-micron particles [22]. Micrographs of live mouse peritoneal macrophages [23] and human monocytes [24] that have been induced to ingest diamond dust particles up to 5 microns in diameter appear to have internalized particles amounting to 10-20% of their cell volume. A particle burden “overload criterion” (i.e., producing macrophage immobilization) of ~600 micron³ per rat alveolar macrophage (a ~50% cellular volumetric burden for 1166 micron³ cells) has been proposed by Oberdorster et al. [25].

Living cells are often seen swimming around inside larger living cells. Are there any obvious volumetric limits? Individual ~200 micron³ lymphocytes have been observed circulating for hours inside larger living cells (~3-5% volume fraction) with no evident ill effect, a phenomenon called emperipolesis ([1], Section 8.5.3.12). While neutrophils and macrophages are both found in mammalian lungs and neither cell phagocytoses the other in significant quantities, alveolar macrophages containing neutrophils have been observed. Neutrophils that have undergone apoptosis are taken up by macrophages, with mean uptake of 3 neutrophils per macrophage [26]. Taking nonmigratory human neutrophils as 204 micron³ and human alveolar macrophages as 4990 micron³, the uptake represents ~12% of macrophage cell volume.

Cells may also harbor smaller pathogens which are usually volumetrically harmless to the host. Perhaps the best-known example is the case of the bacteriophage T4. A single Escherichia coli bacterium injected with a single T4 phage virion at 37^oC in rich media lyses after 25-30 minutes, releasing 100-200 phage particles that have replicated themselves inside [27]. Taking E. coli volume as 0.6 micron³ ([1], Section 10.4.2.5) and phage T4 volume as ~200,000 nm³, then the bacteriophage particle load on E. coli at lysis is 3-7% of bacterial cell volume. In human cells, the tuberculosis bacterium enters the alveolar macrophage, which transports the intruder into the blood, the lymphatic system, and elsewhere. Each ~1 micron³ bacillus that hitches a ride in this manner represents a volumetric intrusion of 0.02% of macrophage volume. Other intracellular microorganisms such as Listeria (~0.25 micron³) and Shigella (~2 micron³), once free in the cytoplasm, are propelled “harmlessly” through the cytosol via continuous cytoskeleton-linked actin polymerization ([1], Section 9.4.6). Macrophages infected with Listeria have been observed with ~2% of their volume co-opted by the microbes (~100 organisms) [28]. While some motile intracellular parasites such as Tyzzer may cause disarrangement and depopulation of host cell organelles by the movement of their peritrichous flagella, other motile intracellular parasites such as the spotted fever-group rickettsiae [29] spread rapidly from cell to cell by actin-based movement but do not cause lysis of the host cell, and typhus-group rickettsiae [29] multiply in host cells to great numbers without profound damage (until cell lysis finally occurs) — providing a more positive intrusiveness benchmark for future medical nanorobots.

Copyright 2001 Robert A. Freitas Jr. All Rights Reserved

References

1. Robert A. Freitas Jr., Nanomedicine, Volume I: Basic Capabilities, Landes Bioscience, Georgetown, TX, 1999; see at: http://www.nanomedicine.com.

2. Z. Zhang, M. Al-Rubeai, C.R. Thomas, “Estimation of disruption of animal cells by turbulent capillary flow,” Biotechnology and Bioengineering 42(1993):987-993.

3. Julio E. Celis, “Microinjection of somatic cells with micropipettes: comparison with other transfer techniques,” Biochem. J. 223(1984):281-291.

4. G.S. Worthen, P.M. Henson, S. Rosengren, G.P. Downey, D.M. Hyde, “Neutrophils increase volume during migration in vivo and in vitro,” Am. J. Respir. Cell Mol. Biol. 10(January 1994):1-7.

5. K. Eric Drexler, Engines of Creation: The Coming Era of Nanotechnology, Anchor Press/Doubleday, New York, 1986.

6. Bernard L. Strehler, Donald D. Mark, Albert S. MilΔ Van, Malcolm V. Gee, “Rate and magnitude of age pigment accumulation in the human myocardium,” Journal of Gerontology 14(1959):430-439.

7. Christopher West, “A quantitative study of lipofuscin accumulation with age in normals and individuals with Down’s syndrome, phenylketonuria, progeria and transneuronal atrophy,” J. Comp. Neurol. 186(1 July 1979):109-116.

8. T.J. James, S.P. Sharma, “Regional and lobular variation in neuronal lipofuscinosis in rat cerebellum: influence of age and protein malnourishment,” Gerontology 41(1995):213-228 (Suppl 2).

9. P. Westermark, E. Wilander, “The influence of amyloid deposits on the islet volume in maturity onset diabetes mellitus,” Diabetologia 15(November 1978):417-421.

10. M.C. Beckerle, “Microinjected fluorescent polystyrene beads exhibit saltatory motion in tissue culture cells,” J. Cell Biol. 98(June 1984):2126-2132.

11. J. Lu, Z. Rosenzweig, “Nanoscale fluorescent sensors for intracellular analysis,” Fresenius J. Anal. Chem. 366(March-April 2000):569-575.

12. G. Kellner-Weibel, P.G. Yancey, W.G. Jerome, T. Walser, R.P. Mason, M.C. Phillips, G.H. Rothblat, “Crystallization of free cholesterol in model macrophage foam cells,” Arterioscler. Thromb. Vasc. Biol. 19(August 1999):1891-1898; see at: http://atvb.ahajournals.org/cgi/content/full/19/8/1891.

13. Tanja Klaus, Ralph Joerger, Eva Olsson, and Claes-Göran Granqvist, “Silver-based crystalline nanoparticles, microbially fabricated,” Proc. Natl. Acad. Sci. (USA) 96(23 November 1999):13611-13614; 96(7 December 1999):14183-14185 (comment).

14. Robert A. Freitas Jr., “How Nanorobots Can Avoid Phagocytosis by White Cells – Part I,” Foresight Update No. 45, 30 June 2001, pp. 10-12; http://www.imm.org/Reports/Rep027.php; and “Part II,” Foresight Update No. 46, 30 September 2001, pp. 18-20; http://www.imm.org/Reports/Rep028.php.

15. R.E. Polonio, R.E. Wolke, S.A. MacLean, J.F. Sperry, “An in vitro assay to measure phagocytosis in striped bass hybrids,” Fish Shellfish Immunol. 10(July 2000):405-418.

16. Y. Kubota, S. Takahashi, O. Matsuoka, “Dependence on particle size in the phagocytosis of latex particles by rabbit alveolar macrophages cultured in vitro,” J. Toxicol. Sci. 8(August 983):189-195

17. B.E. Lehnert, Y.E. Valdez, G.L. Tietjen, “Alveolar macrophage-particle relationships during lung clearance,” Am. J. Respir. Cell Mol. Biol. 1(August 1989):145-154.

18. G.J. Cannon, J.A. Swanson, “The macrophage capacity for phagocytosis,” J. Cell Sci. 101(April 1992):907-913.

19. E.D. Bateman, R.J. Emerson, P.J. Cole, “A study of macrophage-mediated initiation of fibrosis by asbestos and silica using a diffusion chamber technique,” Br. J. Exp. Pathol. 63(August 1982):414-425.

20. B.E. Lehnert, P.E. Morrow, “Association of 59iron oxide with alveolar macrophages during alveolar clearance,” Exp. Lung Res. 9(1985):1-16.

21. John C. Lay, William D. Bennett, Chong S. Kim, Robert B. Devlin, Philip A. Bromberg, “Retention and intracellular distribution of instilled iron oxide particles in human alveolar macrophages,” Am. J. Respir. Cell Mol. Biol. 18(May 1998):687-695.

22. Isabelle Catelas, Alain Petit, Richard Marchand, David J. Zukor, L’Hocine Yahia, Olga L. Huk, “Cytotoxicity and macrophage cytokine release induced by ceramic and polyethylene particles in vitro,” J. Bone Joint Surg. Br. 81(May 1999):516-521.

23. A.C. Allison, J.S. Harington, M. Birbeck, “An examination of the cytotoxic effects of silica on macrophages,” J. Exp. Med. 124(1966):141-154.

24. L. Nordsletten, A.K. Hogasen, Y.T. Konttinen, S. Santavirta, P. Aspenberg, A.O. Aasen, “Human monocytes stimulation by particles of hydroxyapatite, silicon carbide and diamond: in vitro studies of new prosthesis coatings,” Biomaterials 17(August 1996):1521-1527.

25. G. Oberdorster, J. Ferin, P.E. Morrow, ‘Volumetric loading of alveolar macrophages (AM): a possible basis for diminished AM-mediated particle clearance,” Exp. Lung Res. 18(January-March 1992):87-104.

26. L.C. Meagher, J.S. Savill, A. Baker, R.W. Fuller, C. Haslett, “Phagocytosis of apoptotic neutrophils does not induce macrophage release of thromoboxane B2,” J. Leukoc. Biol. 52(September 1992):269-273.

27. A.D. Doerman, “Lysis and lysis inhibition with Escherichia coli bacteriophage,” J. Bacteriol. 55(1948):257-275.

28. Amy L. Decatur, Daniel A. Portnoy, “A PEST-like sequence in listeriolysin O essential for Listeria monocytogenes pathogenicity,” Science 290(3 November 2000):992-995.

29. T. Hackstadt, “The biology of rickettsiae,” Infect. Agents Dis. 5(June 1996):127-143.

IMM would appreciate learning your thoughts on the above article.

In conjunction with Foresight Update 48

The Vasculoid Personal Appliance

By Robert A. Freitas Jr.
Research Scientist, Zyvex Corp.

Robert A. Freitas Jr.

Six years ago, while in the midst of writing the first volume of Nanomedicine [1], I was delighted to discover the germ of a fascinating idea for a very aggressive nanomedical augmentation, originally called “roboblood,” that had just been proposed online [2] at sci.nanotech by Foresight Institute Senior Associate Chris Phoenix. The initial concept elicited considerable enthusiastic discussion, but contained what I regarded as several crucial “show-stopper”-type technical flaws. I contacted Chris directly and we began an intermittent but constructive collaboration that extended through 1996-2002, finally culminating in a detailed technical paper [3] that runs ~100 pages with more than 500 references — which I can only briefly summarize here.

Before proceeding further, I must offer two notes of caution about this work to two distinct audiences of possible readers. First, conservative medical practitioners should be aware that the technical paper summarized in this article is not intended to represent an actual engineering design for a future nanomedical product. Rather, the purpose is merely to examine a set of appropriate design constraints, scaling issues, and reference designs as a purely theoretical exercise to investigate whether or not the basic idea of a blood replacement appliance might be feasible, and to determine key limitations of such designs.

Second, futurist- and transhumanist-oriented readers are warned that, in order to maintain a tight analytical focus, the technical paper necessarily ignores many possible future nanomedical augmentations to human cellular, tissue, and organ systems that would clearly be accessible to a molecular manufacturing nanotechnology capable of building the vasculoid appliance, and that might significantly influence vasculoid architecture, utility or the advisability of its use.

The idea of the vasculoid originated in the asking of a simple question: Once a mature molecular nanotechnology becomes available, could we replace blood with a single, complex robot? This robot would duplicate all essential thermal and biochemical transport functions of the blood, including circulation of respiratory gases, glucose, hormones, cytokines, waste products, and all necessary cellular components. The device would conform to the shape of existing blood vessels. Ideally, it would replace natural blood so thoroughly that the rest of the body would remain, at least physiochemically, essentially unaffected. It is, in effect, a mechanically engineered redesign of the human circulatory system that attempts to integrate itself as an intimate personal appliance with minimal adaptation on the part of the host human body.

A robotic device that replaces and extends the human vascular system is properly called a “vasculoid,” a vascular-like machine; but the vasculoid is more than just an artificial vascular system. Rather, it is a member of a class of space- or volume-filling nanomedical augmentation devices whose function applies to the human vascular tree.

The device is extremely complex, having ~500 trillion independent cooperating nanorobots. In simplest terms, the vasculoid is a watertight coating of nanomachinery distributed across the luminal surface of the entire human vascular tree. This nanomachinery uses a ciliary array to transport important nutrients and biological cells to the tissues, containerized either in “tankers” (for molecules) or “boxcars” (for cells). The basic device weighs ~2 kg and releases ~30 watts of waste heat at a basal activity level and a maximum of ~200 watts of power at peak activity level. The power dissipation of the human body ranges from ~100 watts (basal) to ~1600 watts (peak) [1], so the device presents no adverse thermogenic consequences to the user. Power is derived from native supplies of glucose and oxygen, both plentiful in the human body, as may be common in medical nanorobotic systems [1, 4-8].

		The idea of the vasculoid originated in the asking of a simple question: Once a mature molecular nanotechnology becomes available, could we replace blood with a single, complex robot?

The most important basic structural component of the exemplar vasculoid robot is a ~300 m² two-dimensional vascular-surface-conforming array of ~150 trillion “sapphiroid” (i.e., using sapphire-like building materials) basic plates. (Thermal conductivity favors sapphire over diamond in this application.) These square plates are nanorobots that cover the entire luminal surface of all blood vessels in the body, to one-plate thickness. Each basic plate is an individual self-contained nanorobot ~1 micron thick and ~2 micron² in surface area, a size small enough to allow adequate clearance even in the narrowest human capillaries.

Molecule-conveying “docking bays” comprise ~24 trillion, or 16%, of all vasculoid basic plates. Tankers containing molecules for distribution can dock at these bays and load or unload their cargo. Cell-conveying “cellulocks” are built on “cellulock plates” which span the area of 30 basic plates, or 60 micron² each. Boxcars containing biological cells for distribution can dock at these cellulocks and load or unload their cargo. With only 32.6 billion cellulock plates in the entire vasculoid design, cellulocks occupy the area of 0.978 trillion basic plates or only 0.65% of the entire vasculoid surface. The remaining ~125 trillion basic plates are reserved for special equipment and other as-yet undefined applications. All nanomachinery within each plate is of modular design, permitting easy replacement and repair by mobile repair nanorobots called vasculocytes [7].

Adjacent plates abut through flexible but watertight mechanical interfaces on metamorphic bumpers [1] along the entire perimeter of each plate. Each bumper has controllable variable volume, permitting the vasculoid surface: (1) to slightly expand or contract in area, or (2) to flex, either in response to macroscale body movements or in response to vascular surface corrugations or other irregularities to the same degree or better than the natural endothelium.

Thus, plated surfaces readily accommodate the natural cyclical volume changes of various organs such as lung, bladder, or spleen. Rigidity of the plate array is also subject to engineering control and to localized real-time control as well, via the bumpers; diamondoid or sapphire plating may be made substantially stiffer than natural endothelium, if desired.

Installation of the vasculoid involves complete exsanguination of a sedated patient, replacement of the natural circulatory fluid with various installation fluids, followed by mechanical vascular plating, defluidization, and finally activation of the vasculoid. Installation takes ~6.5 hours from start to finish and requires a peak ~200-watt power draw midway through the procedure. (By comparison, present-day kidney dialysis treatments require 4-12 hours and the equipment also draws a few hundred watts.) The hypothetical installation protocol was selected for maximum comfort, reversibility, and reliability according to contemporary medical standards.

The advantages of installing a vasculoid are potentially numerous. Many of these benefits theoretically could be provided on a temporary or more limited basis using terabot-dose injections of considerably less aggressive bloodborne nanomedical devices. However, the vasculoid appliance simultaneously provides all benefits on an essentially permanent and whole-body basis. Additionally, some benefits appear unique to the vasculoid and can be achieved in no other way.

Whether the entire package is sufficiently attractive to warrant installation will probably be a matter of personal taste rather than of medical necessity, since a molecular nanotechnology capable of building and deploying a complex vasculoid is likely to offer complete non-vasculoid cures for most circulatory and blood-related disorders that plague humanity today, and biological enhancements may also be available. And in nanomedicine, moving from an augmentation technology that works alongside a natural system (e.g., respirocytes [5]) to an augmentation technology that entirely replaces a natural system (e.g., vasculoid [3]) may involve significant safety, psychological, and even ethical considerations.

The most important benefits of vasculoid installation may include:

1. Exclusion of parasites, bacteria, viruses, and metastasizing cancer cells from the bloodflow, thus limiting the spread of bloodborne disease. Such microorganisms and cells are easily eliminated from the blood using ~cm³ doses of appropriately programmed nanobiotics [1, 4], but such individual nanorobotic devices might not normally be deployed on a permanent basis. Intracellular pathogens that can infect motile phagocytic cells (e.g., the tuberculosis Mycobacterium or the bacterium Listeria, both of which can reside inside macrophages [9]) cannot be directly excluded from the tissues when infected cells are transported by the vasculoid.

   However, cell surface markers will often reveal such infection, so vasculoid systems can check for the presence of such markers and thus deny these cells re-entry to human tissues.

   For example, the membrane surface of macrophages infected by Mycobacterium microti is antigenically different from that of uninfected macrophages [10]; Listeria-derived peptides are found acting as integral membrane proteins in the plasma membrane of infected macrophages [11], and other Listeria-infected antigen-presenting cells display hsp60 on their plasma membranes only when infected [12].

2. Faster and more reliable trafficking of lymphocytes throughout the secondary lymphoid organs, allowing them to survey for targeted antigens in minutes or hours, rather than days (because both white cells and antigenic sources can be efficiently concentrated), thus greatly speeding the natural immune system response to foreign antigen. This lymphocyte function might also be augmented or replaced using individual histomobile medical nanorobots [1] or biorobots.

   If biorobots are developed first, many vasculoid installations might take place in patients possessing largely artificial immune systems, thus obviating the need for much of the cellular component trafficking that would otherwise be mediated by boxcars and cellulocks.

3. Eradication of most serious circulatory-related pathological conditions including all vascular disease (e.g., aortic dissection, vessel blockages, spasms, aneurysms, phlebitis, varicose veins), heart disease, syncope (including orthostatic hypotension) and shock, stroke, and bleeding, due to the elimination of unconstrained metabolite and fluid circulation.

   Certain other conditions due to localized prevention of blood flow such as bedsores and subclinical paresthesias (e.g., “pins-and-needles” sensation) can also be ameliorated, since stiffened blood vessels will not be nearly so easy to close via external compression.

   Again, many of these conditions may already have adequate nanomedical treatments by the time the vasculoid can be built, but other conditions might not yet be readily or as conveniently treatable, such as the dangers of large-scale bleeding (both internal and external).

4. Reduced susceptibility to chemical, biochemical, and parasitic poisons of all kinds, including allergenic substances in food, air and water, although bloodborne nanotankers or pharmacytes [1] may be able to partially duplicate this function as well.
5. Faster metabolite transport and distribution, significantly improving physical endurance and stamina, including the ability to breathe at low O₂ partial pressures and the ability to flush out unwanted specific biochemicals from the body (a feature which might be duplicated using bloodborne respirocyte-class devices [1, 5]). The architecture would also permit convenient long-term storage of protein, or amino acid recovery and recycling, which could prove nutritionally useful.
6. Direct, rapid user control of many hormonal- and neurochemical-mediated, and all blood-mediated, physiological responses. It would be difficult (though not impossible) to provide equivalent comprehensive whole-body physiological control using individual micron-scale bloodborne nanorobots alone.
7. Voluntary control of capillary conductance and rigidity permitting conscious regulation of thermal energy exchange with the environment and at least limited control of whole-body morphological structure, rigidity (e.g., stiffness, bending modulus, etc.), and volume with ~millisecond response times.
8. At least partial protection from various accidents and other physical harm such as insect stings, animal bites, collisions, bullet or shrapnel penetrations, or falling from heights. This is perhaps the only specific benefit of the vasculoid appliance that could not be achieved by any less radical means: extreme trauma resistance, especially resistance to exsanguination and cushioning against mechanical shock.

Medically oriented readers might properly wonder why anyone would want to discuss replacing a perfectly functional natural fluid transport system with an untested, complex, artificial, dry system with which humans have no experience today. There are several answers to this very good question.

		It would seem that a somewhat more advanced and compact version of the proposed device could function independently of nearly all noncortical tissue. Thus the vasculoid is most fascinating because it may represent one last outpost of humanity at the final frontier of biological evolution.

First of all, medical skeptics should bear in mind that the vasculoid appliance is clearly a highly sophisticated medical nanosystem. It cannot be built without using a manufacturing system based on a mature molecular nanotechnology. Its use would come only after many decades of previous engineering experience in building, testing, and operating such highly complex systems inside the human body.

In the future nanomedicine-rich milieu in which it would be deployed, the vasculoid as a medical intervention may be closer to the typical than to the extreme (as it might appear today). It is as if we were looking forward from the limited vantage point of the 1950s — a technological era in which vacuum tubes still reigned supreme — to the year 2002, and estimating the future feasibility of a 1 GHz Pentium III laptop computer (a feat of prognostication actually achieved by Arthur C. Clarke [13]).

In the nanomedical era, it will be a matter of personal preference and choice for each patient, in consultation with their physician, whether the aforementioned benefits of the vasculoid appliance are worth the risks. The device described in this article would represent a most extreme intervention using a very advanced medical molecular nanotechnology.

The technical paper [3] concludes: “Ultimately, and from the standpoint of human-guided evolution, the body exists primarily to ensure the survival of the mind — not the replication of the genes, which was the ancient paradigm [14, 15]. It would seem that a somewhat more advanced and compact version of the proposed device could function independently of nearly all noncortical tissue. Thus the vasculoid is most fascinating because it may represent one last outpost of humanity at the final frontier of biological evolution.”

Acknowledgements

The author thanks Robert J. Bradbury, Ken Clements, J. Storrs Hall, Hugh Hixon, Tad Hogg, Markus Krummenacker, Jerry B. Lemler, M.D., James Logajan, Ralph C. Merkle, Rafal Smigrodzki, M.D., Tihamer Toth-Fejel, and Brian Wowk for their comments and review of an earlier version of the original technical paper upon which this article was based.

© 2002 Robert A. Freitas Jr. All Rights Reserved

References

1. Robert A. Freitas Jr., Nanomedicine, Volume I: Basic Capabilities, Landes Bioscience, Georgetown, TX, 1999; http://www.nanomedicine.com

2. Christopher J. Phoenix, “Early Nanotech Project: Replace Blood?” sci.nanotech posting on 14 June 1996; http://discuss.foresight.org/critmail/sci_nano/2273.html or http://crit.org/critmail/sci_nano/2273.html

3. Robert A. Freitas Jr., Christopher J. Phoenix, “Vasculoid: A personal
nanomedical appliance to replace human blood,” Journal of Evolution and
Technology, 11(April 2002);
http://www.transhumanist.com/volume11/vasculoid.html,
http://www.transhumanist.com/volume11/vasculoid.pdf

4. Robert A. Freitas Jr., “Microbivores: Artificial Mechanical Phagocytes using
Digest and Discharge Protocol,” Zyvex preprint, March 2001; http://www.rfreitas.com/Nano/Microbivores.htm.
See also: Robert A. Freitas Jr., “Microbivores: Artificial Mechanical Phagocytes,”
Foresight Update No. 44, 31 March 2001, pp. 11-13; http://www.imm.org/Reports/Rep025.php

5. Robert A. Freitas Jr., “Exploratory design in medical nanotechnology: A mechanical artificial red cell,” Artificial Cells, Blood Substitutes, and Immobil. Biotech. 26(1998):411-430; http://www.foresight.org/Nanomedicine/Respirocytes.html

6. K. Eric Drexler, Engines of Creation: The Coming Era of Nanotechnology, Anchor Press/Doubleday, New York, 1986; http://www.foresight.org/EOC/index.html

7. Robert A. Freitas Jr., “Vasculocytes,” unpublished document, 14 September 1996; http://www.foresight.org/Nanomedicine/Gallery/Species/Vasculocytes.html

8. Robert A. Freitas Jr., “Clottocytes: Artificial Mechanical Platelets,” Foresight Update No. 41, 30 June 2000, pp. 9-11; http://www.imm.org/Reports/Rep018.php

9. Amy L. Decatur, Daniel A. Portnoy, “A PEST-like sequence in listeriolysin O essential for Listeria monocytogenes pathogenicity,” Science 290(3 November 2000):992-995.

10. S. Majumdar, H. Kaur, H. Vohra, G.C. Varshney, “Membrane surface of Mycobacterium microti-infected macrophages antigenically differs from that of uninfected macrophages,” FEMS Immunol. Med. Microbiol. 28(May 2000):71-77.

11. P.M. Allen, D.I. Beller, J. Braun, E.R. Unanue, “The handling of Listeria monocytogenes by macrophages: the search for an immunogenic molecule in antigen presentation,” J. Immunol. 132(January 1984):323-331; P.M. Allen, E.R. Unanue, “Antigen processing and presentation by macrophages,” Am. J. Anat. 170(July 1984):483-490.

12. Cindy Belles, Alicia Kuhl, Rachel Nosheny, Simon R. Carding, “Plasma membrane expression of heat shock protein 60 in vivo in response to infection,” Infect. Immun. 67(August 1999):4191-4200; http://iai.asm.org/cgi/content/full/67/8/4191?view=full&pmid=10417191

13. Arthur C. Clarke, Profiles of the Future, Harper and Row Publishers, New York, 1962.

14. Charles Darwin, The Origin of Species by Means of Natural Selection, 1859; definitive 6^th London edition: http://www.literature.org/authors/darwin-charles/the-origin-of-species-6th-edition/

15. Edward O. Wilson, Sociobiology: The New Synthesis, Harvard University Press, Cambridge, MA, 1975.

More on the medical applications of nanotechnology at:
http://www.foresight.org/Nanomedicine/

IMM would appreciate learning your thoughts on the above article.

In conjunction with Foresight Update 49

Could Medical Nanorobots Be Carcinogenic?

By Robert A. Freitas Jr.
Research Scientist, Zyvex Corp.

Biocompatibility is an important property that must be carefully engineered into all medical nanorobots used in nanomedicine [1, 2]. One key aspect of biocompatibility is whether implanted nanoorgans, or in vivo medical nanorobots, might induce undesirable genetic changes as a side effect of their presence or activities inside the human body. Such undesirable changes might take many forms. For instance, mutagenicity [3] is the production of inheritable coding flaws in chromosomes that otherwise may retain much genetic functionality. (All carcinogens are mutagens but not vice versa – a mutation may be lethal to a cell, may prevent cellular replication, or may not affect metabolic or growth processes sufficiently to produce malignant behavior [4].) Genotoxicity is a more serious injury to the chromosomes of the cell, such that when the cell divides, fragments of chromosomes and micronuclei remain in the cytoplasm. Teratogenicity [5] is the ability of a foreign material (or a fetotoxic agent) to induce or increase the risk of developing abnormal structures in an embryo, or birth defects. Carcinogenicity is the ability to produce or increase the risk of developing cancer – materials may be directly carcinogenic or may potentiate other agents [4]. Tumorigenic materials tend to induce neoplastic transformations, especially malignant tumors.

Direct experimental exploration of the carcinogenicity of likely nanorobot building materials has barely begun, but information available to date appears guardedly optimistic. For example, diamond (DLC) coatings exhibit low mutagenicity toward human fibroblasts in vitro [6] and there are no reports of diamond carcinogenicity or tumorigenesis. Alumina (sapphire) produces no mutagenic or carcinogenic effects on cultured human osteoblasts [7] or when used as a blood-contacting material in a centrifugal blood pump [2]; while aluminum ion leached from sapphire at the highest plausible concentrations (~10^-5 M; [2]) might inhibit eukaryotic transcription, experiments suggest that the mutagenicity, carcinogenicity, and teratogenicity of aluminum is low [8]. Teflon particles appear to be noncarcinogenic [2], even though tetrafluoroethylene (a monomer used in Teflon manufacture) is hepatocarcinogenic after long-term inhalation by mice [9]. There are no reports of carcinogenicity from pyrolytic carbon, graphite, or pure India ink in humans [2]. In rodents, the inhalation of carbon black particles can produce pulmonary neoplasms and lung carcinoma [2], and particle-elicited macrophages and neutrophils can exert a mutagenic effect on in vitro rat epithelial cells [10].

The possible carcinogenicity of fullerenes was suggested a decade ago [11] but even by 1998 the risk was no longer considered serious [12]. Pure C₆₀ and C₇₀ molecules do not intercalate into DNA (which might promote cancer) when mouse skin is exposed to them [11], although water-miscible fullerene carboxylic acid can cleave G-selective DNA chains [13]. No mutagenicity or genotoxicity of C₆₀ as fullerol is observed in prokaryotic cells and only slight genotoxicity is seen in eukaryotic cells at the highest concentrations – even though C₆₀ dissolved in polyvinylpyrrolidone was found to be mutagenic for several Salmonella strains due to singlet oxygen formation, and pure C₆₀ is a known singlet oxygen generating agent, and singlet oxygen is known to be genotoxic [2]. Repeated epidermal administration of fullerenes for up to 24 weeks resulted in neither benign nor malignant tumor formation in mice, but promotion with a phorbol ester produced benign skin tumors [11]. Some C₆₀ derivatives have actually shown promise as anti-cancer or anti-tumor agents [2]. Carcinogenicity studies of rolled graphene sheets such as carbon nanotubes remain to be done.

There are four kinds of carcinogenesis [4] which may be relevant in medical nanorobotics:

(1) Chemical Carcinogenesis. Chemical carcinogenicity is actually a somewhat uncommon property of materials. An exhaustive literature search on 6000 of the most likely chemical candidates found only 1000 (17%) identified as possible carcinogens [4]. The classic study by Innes et al. [14] found that fewer than 10% of 120 pesticides and toxic industrial chemicals tested were carcinogenic, and even this study was criticized as being too pessimistic because testing toxic potential carcinogens at high dosages may artificially accentuate their activity by inducing increased rates of cell division [15]. Medical nanorobots normally will have chemically-inert nonleachable surfaces, but designers should ensure that all possible nanorobot effluents are noncarcinogenic. (Potential nanorobot effluents may be prescreened during design using existing computational toxicology techniques [3].)

(2) Nonspecific Carcinogenesis. Neoplasms can arise in response to chronic irritation, leading to chronic inflammation and granulomatous reaction to implants [2]. Chemicals, foreign bodies, infection and mechanical trauma [16] can lead to this type of neoplastic transformation which is characterized by replication infidelity – i.e., a cell that produces a daughter cell not identical to its parent, as in, for example, the formation of hyperplastic expansive scars known as nonmalignant keloids [4]. These benign lesions can occasionally, and apparently spontaneously, transform into malignant neoplasms such as fibrous histiocytomas.

(3) Ex Cyto Foreign Body Carcinogenesis. In the 1950s it was discovered that many agents not previously thought to be carcinogenic produced dramatic neoplasm incidence rates in rodents when implanted in solid form rather than injected or fed in soluble or dispersed form, an effect called foreign body (FB) carcinogenesis [17-20], solid state carcinogenesis, or the Oppenheimer effect. The induction of neoplasms increases with the size of the implant and with decreasing inflammatory response (e.g., well-tolerated materials are, in the long run, better FB carcinogens). The risk of transformation is reduced on surfaces with porosity of average diameter above 220 nm, and materials with distributed porosity of cellular dimensions are less carcinogenic in rodents than smooth nonporous material [4, 19]. Nonperforated polymer films induce subcutaneous sarcomas in mice and rats, but implanted foreign bodies with other shapes (e.g., perforated or minced films, or filters with 450-nm pores [19]) or with roughened surfaces are weakly or non-carcinogenic except when total foreign-body surface area exceeds ~1 mm² [20]. In vitro experiments by Boone et al. [18] and in vivo experiments by Brand [17] studied the effects of attachment of mouse fibroblasts to polycarbonate plates. Cells implanted after an in vitro exposure produced transplantable, undifferentiated sarcomas, leading these authors to conclude that the smooth surface of the plates acted as an FB carcinogen for at least initiation of tumorigenesis, independent of chemical composition. Brand [17] cited six possible mechanistic origins of FB carcinogenesis and concluded that: (1) disturbance of cellular growth regulation was most likely, based on the heritability of neoplastic behavior in the growing cell population, and (2) interruption of cellular contact or communication might also play a role in neoplasm expression and maturation (rather than neoplasm induction). It is now well-established that smooth-surfaced foreign bodies, regardless of their chemical composition, will produce sarcomas when transplanted subcutaneously into rodents [18], and foreign-body sarcomatous growth in mice appears resistant to substances which normally inhibit neoplastic growth.

Is there any evidence that humans are also susceptible to ex cyto FB carcinogenesis? There is no evidence that a single incident of mechanical trauma can cause cancer [21], but evidently there are 28 known cases of tumors in humans associated with partial or total joint replacements, which occurred either fairly soon after implantation or a very long time (10-15 years) after implantation, the latter primarily as malignant fibrous histiocytomas [4]. However, all of these tumors were associated with stainless steel or cobalt-based alloy devices, perhaps due to elevated tissue concentrations of metals near the implant [4] (e.g., metal-on-metal devices can produce a 10- to 15-fold rise in circulating serum chromium). There are a few additional reports of possible remote-site tumors [4], but other studies find such implant-related tumorigenicity to be very weak or nonexistent. Some investigators have therefore concluded that there is little clinical evidence for ex cyto FB carcinogenesis in humans, and that the Oppenheimer effect may be a consequence of the relatively primitive immune system of rodents in comparison to that of humans [4]. But Black [4] urges caution because, in rare cases, sarcomas appear to have arisen on unabsorbable foreign bodies in man [20] – a category of foreign bodies which would definitely include diamondoid medical nanorobots and nanoorgans. Polarizable foreign particles have also been associated with cutaneous granulomas in three cases of systemic sarcoidosis [22].

(4) In Cyto Foreign Body Carcinogenesis. Although FB carcinogenesis produced by materials external to cells appears to be rare in humans, solid materials in a form that can penetrate cells can be carcinogenic, a phenomenon originally known as the Stanton hypothesis [23]. The best-known example is chrysotile asbestos, first recognized as a human carcinogen only because it produced a relatively rare lung tumor. Subsequent studies of asbestos and related fibers in animal models revealed that mesothelioma could be induced by fibers <~0.25-1.5 microns in diameter and >~4-8 microns in length, regardless of fiber composition [23, 24]. Quantitatively, Stanton [23] found that ~10⁵ fibers of carcinogenic dimensions, embedded in the human body, yielded a ~10% probability of developing a tumor within 1 year; ~2 x 10⁷ fibers raised the probability to 50%; and 10⁹ fibers, 90%. In vitro fiber cytotoxicity correlates well with fiber dimensions [24], particularly the aspect ratio, with fiber durability, and not with fiber bulk composition but rather with the molecular nature of active surface properties which can also play a role in carcinogenic potency [2]. Stiff slender fibers such as mineral whiskers can penetrate cells and may produce mechanical or oxidoreduction damage to the nucleus and to chromosomes [25] regardless of the material of which they are comprised. The likely mechanism is oxyradical activity because antioxidant enzymes appear to protect cells against genotoxic damage induced by chrysotile fibers. This risk factor must be borne in mind when designing medical nanorobots (including all of their possible operational and failure-mode physical configurations) and any potentially detachable subsystems which may be of similar stiffness and size as the cytotoxic fibers.

© 2002 Robert A. Freitas Jr. All Rights Reserved

References

1. Robert A. Freitas Jr., Nanomedicine, Volume I: Basic Capabilities, Landes Bioscience, Georgetown, TX, 1999; http://www.nanomedicine.com

2. Robert A. Freitas Jr., Nanomedicine, Volume IIA: Biocompatibility, Landes Bioscience, Georgetown, TX, 2002.

3. A.M. Richard, “Structure-based methods for predicting mutagenicity and carcinogenicity: are we there yet?” Mutat. Res. 400(25 May 1998):493-507.

4. Jonathan Black, Biological Performance of Materials: Fundamentals of Biocompatibility, Third Edition, Marcel Dekker, New York, 1999.

5. M.R. Juchau, “Chemical teratogenesis in humans: biochemical and molecular mechanisms,” Prog. Drug Res. 49(1997):25-92.

6. D.P. Dowling et al. , “Evaluation of diamond-like carbon-coated orthopaedic implants,” Diam. Rel. Mat. 6(March 1997):390-393.

7. Y. Josset et al., “In vitro reactions of human osteoblasts in culture with zirconia and alumina ceramics,” J. Biomed. Mater. Res. 47 (15 Dec 1999):481-493.

8. A. Leonard, G.B. Gerber, “Mutagenicity, carcinogenicity and teratogenicity of aluminum,” Mutat. Res. 196(November 1988):247-257.

9. H.H. Hong et al. , “Frequency of ras mutations in liver neoplasms from B6C3F1 mice exposed to tetrafluoroethylene for two years,” Toxicol. Pathol. 26(September-October 1998):646-650.

10. K.E. Driscoll et al. , “Effects of particle exposure and particle-elicited inflammatory cells on mutation in rat alveolar epithelial cells,” Carcinogenesis 18(February 1997):423-430.

11. M.A. Nelson et al. , “Effects of acute and subchronic exposure of topically applied fullerene extracts on the mouse skin,” Toxicol. Ind. Health 9(July-August 1993):623-630.

12. Robert F. Service, “Nanotubes: The next asbestos?” Science 281(14 August 1998):941.

13. H. Tokuyama et al. , “Photoinduced biochemical activity of fullerene carboxylic acid,” J. Am. Chem. Soc. 115(1993):7918-7919.

14. J.R.M. Innes et al. , “Bioassay of pesticides and industrial chemicals for tumorigenicity in mice: a preliminary note,” J. Natl. Cancer Inst. 42(June 1969):1101-1114.

15. L.S. Gold et al. , “What do animal cancer tests tell us about human cancer risk? Overview of analyses of the carcinogenic potency database,” Drug Metab. Rev. 30(May 1998):359-404.

16. S. Vieira de Oliveira et al. , “Effects of uracil calculi on cell growth and apoptosis in the BBN-initiated Wistar rat urinary bladder mucosa,” Teratog. Carcinog. Mutagen. 19(1999):293-303.

17. K.G. Brand, in F.F. Becker, ed., Cancer: A Comprehensive Treatise, Volume 1, Plenum Press, New York, 1975, p. 485.

18. C.W. Boone et al. , “Spontaneous neoplastic transformation in vitro: a form of foreign body (smooth surface) tumorigenesis,” Science 204(13 April 1979):177-179.

19. T.G. Moizhess, J.M. Vasiliev, “Early and late stages of foreign-body carcinogenesis can be induced by implants of different shapes,” Int. J. Cancer 44(15 September 1989):449-453.

20. M. Mhic Iomhair, S.M. Lavelle, “Effect of film size on production of foreign body sarcoma by perforated film implants,” Technol. Health Care 5(October 1997):331-334.

21. L. Weiss, “Some effects of mechanical trauma on the development of primary cancers and their metastases,” J. Forensic Sci. 35(May 1990):614-627.

22. N.M. Walsh et al. , “Cutaneous sarcoidosis and foreign bodies,” Am. J. Dermatopathol. 15(June 1993):203-207.

23. Mearl F. Stanton et al. , “Relation of particle dimension to carcinogenicity in amphibole asbestoses and other fibrous minerals,” J. Natl. Cancer Inst. 67(November 1981):965-975.

24. L.E. Lipkin, “Cellular effects of asbestos and other fibers: correlations with in vivo induction of pleural sarcoma,” Environ. Health Perspect. 34(February 1980):91-102.

25. E. Dopp, D. Schiffmann, “Analysis of chromosomal alterations induced by asbestos and ceramic fibers,” Toxicol. Lett. 96-97(August 1998):155-162.

More on the medical applications of nanotechnology at:
http://www.foresight.org/Nanomedicine/

IMM would appreciate learning your thoughts on the above article.

In conjunction with Foresight Update 50

Is Sapphire Biocompatible With Living Cells?

By Robert A. Freitas Jr.
Research Scientist, Zyvex Corp.

Robert A. Freitas Jr.

Much of the discussion of medical nanorobotics building materials [1, 2] focuses on diamond because this form of carbon is expected to be convenient to synthesize mechanosynthetically and has high rigidity and strength, excellent chemical inertness, and good biocompatibility [1, 3]. However, in many proposed in vivo applications of medical nanorobots [4] and medical nanorobot aggregates [5], sapphire appears to be a useful material as well. Sapphire is a high-density monocrystalline alumina (Al₂O₃) already used in nanofluidics, dental implants ([1], Section 15.3.5.2), and other near-term nanomedical applications, and is being considered for use in the manufacture of prosthetic heart valves. Like diamond, sapphire [6, 7] is generally regarded as a nontoxic bioinert ceramic material that is likely noncarcinogenic [8].

The advantages of sapphire may be briefly summarized as follows. First, sapphire is almost as strong and hard as diamond ([2], Table 9.3), and only slightly more dense. Second, sapphire is already fully oxidized, so in particulate form (e.g., micron-size nanorobots) it cannot burn in air like diamond, and its crystalline structure remains stable to higher temperatures than diamond. Third, sapphire has more favorable bulk thermophysical characteristics. The thermal conductivity of sapphire is 100-1000 times less than for diamond, reducing the increase in the thermal conductivity of blood that is loaded with sapphire nanorobots, compared to blood containing pure-diamond nanorobots, especially at the highest loadings ([1], Section 15.3.8). Sapphire also has 60% greater heat capacity per unit volume than diamond. Fourth, sapphire offers designers an alternative hydrophilic surface chemistry as compared to hydrophobic diamond. Sapphire is amphoteric, absorbing H⁺ ions in very acidic environments (acquiring a positive charge) and absorbing OH^– ions in alkaline environments (acquiring a negative charge), while remaining isoelectric (electrically neutral) at intermediate pH values near human physiologic at ~7.4 ([1], Section 15.5.6.1). Fifth, sapphire can be manufactured in a full spectrum of colors ([2], Section 5.3.7) by replacing 0.01%-0.1% of the aluminum atoms with atoms of iron, titanium, or chromium, while producing only modest changes in the physical and chemical properties of the material.

Is sapphire, like diamond [3], a biocompatible material for living cells? A number of experiments have been performed to determine the response of fibroblasts to bulk alumina and sapphire surfaces. For example, alumina ceramic surface has shown excellent in vitro biocompatibility in a tissue culture of rabbit fibroblasts and cultured embryonic mouse fibroblasts [7]. Fibroblast-like mesenchymal cell populations cultured on solid alumina ceramic surfaces induce no cytotoxic or antiproliferative effects on monolayer populations in vitro. A scanning electron microscopic study [9] was conducted on the adhesion, spreading and formation of confluent cell monolayers from fibroblasts and epithelioid cells on Al₂O₃ ceramics. The study found that the cells adhered, spread, migrated and proliferated on the surfaces tested, leading to the conclusion that this implant material is compatible with cells [9]. In general, such cells adhere well to single-crystal or polycrystalline alumina. Other experiments found that human orbital fibroblasts grown on alumina bioceramic implant were free of debris and had the largest cell count, whereas cells grown on hydroxyapatite or porous polyethylene implants had cellular debris associated with them.

The response of bone cells to alumina ceramic has also been examined. For instance, the morphological responses of individual osteoblasts as they attached and spread on alumina surfaces in vitro have been examined by scanning electron microscopy [10]. The cells are round after 30 minutes, then spread radially during the next 1.5 hours until they are almost flat with a nuclear bulge on both rough and polished alumina [10]. Others have investigated the response of various oral cells to sapphire dental implant surfaces. In one study [6], the influence of single-crystal sapphire on the behavior of human epithelial cells and fibroblasts derived from biopsies of the oral mucosa was studied. Compared to control cultures, no effects on cell morphology and growth characteristics were observed. Another study [11] sought to elucidate the ultrastructure of peri-implant junctional epithelium (IJE) on single-crystal sapphire dental implants connected to adjacent teeth by a metal superstructure, by examining the peri-implant gingivae of ten monkeys using a transmission electron microscope at 3, 6 and 12 months after implant insertion. At the time of examination, the ultrastructural features of the IJE were almost identical to those of the natural junctional epithelium attached to natural teeth.

The biocompatibility of alumina and sapphire particles has also been investigated because it is important to understand the biological reaction to inhaled particles and to wear particles that might be produced by frictional forces in long-term prosthetic implants. Such studies generally involve micron-scale particles, roughly the size range of future medical nanorobots — which is why we’re interested in the results.

First, is alumina powder lethal? In one early experiment [12], massive administration of several bioactive <44-micron ceramic powders were lethal to Balb/c mice in 5 gm/kg doses when injected intraperitoneally, producing a swollen kidney with an ischemic color [12]. Alumina powder similarly injected as a control (equivalent to 90 trillion 1-micron³ nanorobots injected into the peritoneal cavity of a 70-kg human) was not lethal and elicited no changes in blood chemistry [12]. All powders had almost no systemic effects when injected intramuscularly or subcutaneously [12].

What about inflammation? Early studies [7] of 0.5-5 micron alumina particles implanted subcutaneously and intraarticularly (knee joint) for up to 5 months in mice revealed no persistent inflammatory or progressive fibrotic reactions around the powder deposits. After an initial acute (3-7 day) granulocytic inflammatory phase, the material was gradually contained within macrophages and deposited locally without significant fibrous tissue reaction. Some particles were transported via lymphatic vessels into regional lymph nodes [7]. Particles were found in the interstitium of the lung; in the reticuloendothelial cells of the liver, spleen, and bone marrow; and in one case in the meshwork of a renal glomerulum; but rarely in the bloodstream. Such crystal deposits caused no local cell necrosis, fibrosis, or granulomatous reaction in any of these organs [7].

Most cytocompatibility studies have examined the foreign-body reactions of macrophages to alumina particles. For example, it is known that rat alveolar macrophages readily ingest aluminum oxide particles [13]. Toxicity tests of alumina powder in vitro using rabbit alveolar macrophages and in vivo using direct intratracheal injection into rat lungs found that the powder had low toxicity for macrophages and minimal recruitment of airway cells and neutrophils in the rat lungs. Rat tissue responses to alumina powder administered at low doses were investigated by Di Silvestre et al. [14], who found that powdered alumina implantation in the subcutis, the muscle and the peritoneum of the rat produced the same intense acute inflammatory reaction in all implantation sites after 2 weeks. However, after 8 weeks the inflammatory reaction had regressed and there was a thin layer of connective tissue around the implanted material, completely isolating it from the surrounding tissues [14]. Examination of human biopsies from well-fixed human total hip prostheses showed that alumina particle deposits increase with time with only a low-grade macrophagic reaction. The amount of necrosis and fibrosis was lower for alumina implant wear debris than that associated with metal or polyethylene implants [15]. Intraperitoneal and intramuscular implantation of powdered alumina particles in rats showed an initial granulocytic reaction with some uptake by the reticuloendothelial system [13]. No antigenicity of alumina ceramic was found in another study that attempted induction of footpad swelling in ceramic-immunized mice.

Sapphire is generally biocompatible with macrophages. Harms and Mausle [13] tested the biocompatibility of alumina ceramic in macrophage cultures revealed no acute cytotoxicity. Pizzoferato et al. [16] found that saline-suspended 1-12.5 micron alumina particles were only slightly phagocytosed in vivo by mouse peritoneal macrophage cells lavaged 1 week post-injection. Christel [15] noted that an examination of human biopsies from failed total hip prostheses revealed a foreign-body reaction containing predominantly macrophages, loaded with alumina particles, that had no morphologic alteration and had not lost their chemotactic ability [15]. Nakashima et al. [17] reported that 1-, 100-, and 1000-micron alumina particles could induce the release of bone resorbing mediators (IL-6, TNF-α, IL1-α) by macrophages in a dose-dependent manner, but hydroxyapatite particles of equal size stimulated a greater release than the alumina. Alumina dust has no significant effect on the in vitro enzyme activity of alveolar macrophages in the rat.

Catelas et al. [18] measured the effects of size (0.6-4.5 micron), concentration (5-1250 particles/macrophage), and composition (e.g., alumina) of ceramic particles on phagocytosis and cell mortality in the J774 mouse macrophage cell line. Kinetic studies (from 5 min to 24 hours) revealed that phagocytosis of the particles begins very early after cell exposure, increasing with time and particle concentration and reaching a plateau after ~15 hours. Phagocytosis increases with concentration for particles up to 2 microns. For larger particles up to 4.5 microns, phagocytosis reaches a plateau independent of particle size and concentration, suggesting a saturation effect most likely dependent on the total volume ingested [18]. Cytotoxicity studies revealed that macrophage mortality increases with particle size and concentration for sizes greater than 2 microns (to >30% cell mortality). Smaller particles (0.6 microns) cause cell mortality only at higher concentrations, and the mortality is still very low (~10%) [18]. Related studies by Catelas using the same cell model investigated the induction of apoptotic cell death in macrophages by alumina ceramic and other powders of different sizes and concentrations. Of some concern, Catelas found that the apoptotic effect of ceramic particles on nuclear morphology was size- and concentration-dependent, but that alumina ceramic particles induce apoptosis more effectively than polyethylene particles at concentrations of 125-250 particles/macrophage for ~2 hours. More work is needed to resolve this issue.

The responses of a few other cell types to alumina ceramic powders have also been investigated. For example, cultured human fibroblasts exposed to 1-500 micrograms/cm³ of alumina powder showed no cytotoxic effects with cell viability at different exposure times measured by colony formation efficiency, neutral red uptake and colorimetric tetrazolium reduction [19]. No cytotoxic or antiproliferative effects were induced in fibroblast-like mesenchymal cell monolayer populations cultured in vitro on powdery alumina ceramic. Alumina powders generally induce no cytotoxicity in cell cultures [20] of human gingival fibroblasts or osteoblastlike cells. Rodrigo et al. [21] found some change in osteoblast function from 10-micron α-alumina particles in human bone cell cultures, and that while both polyethylene and a-alumina increase the expression and secretion of IL-6 in human osteoblastic cells, the stimulation is weaker from α-alumina at the same particle dose. Oonishi et al. [22] observed no inflammation or cell infiltration for 10- and 100-micron alumina particles implanted in holes drilled in the femoral condyles of rabbits. Dion et al. [23] found that the hemolysis eventually initiated in vitro by alumina powder is almost zero.

As for inhalation toxicity, human experience with alumina powders strongly suggests that they are not associated with major specific pulmonary hazards under typical 20th century conditions of occupational inhalation exposure, though rodent experiments suggest that clearance of alumina particles from the lung is slow [24]. OSHA occupational exposure limits for alumina dust are 10 mg/m³ (total fraction) and 5 mg/m³ (respirable fraction), respectively, according to the official Material Safety Data Sheets. (10 mg/m³ of sapphire particles equates to ~3 billion/m³ airborne 1-micron³ nanorobots.) α-alumina 100-700 nm particles have only minimal or no fibrogenic reactivity, and only at doses instilled intratracheally that are massive compared to the amount which could reasonably be inhaled in any one breath.

We can be guardedly optimistic that monocrystalline sapphire should prove to have very good biocompatibility with living cells.

Copyright 2002 Robert A. Freitas Jr. All Rights Reserved

References

1. Robert A. Freitas Jr., Nanomedicine, Volume IIA: Biocompatibility, Landes Bioscience, Georgetown, TX, 2003; http://www.nanomedicine.com

2. Robert A. Freitas Jr., Nanomedicine, Volume I: Basic Capabilities, Landes Bioscience, Georgetown, TX, 1999; http://www.nanomedicine.com

3. Robert A. Freitas Jr., “Is Diamond Biocompatible With Living Cells?” Foresight Update No. 39, 30 December 1999, pp. 7-9; http://www.imm.org/Reports/Rep012.php

4. Robert A. Freitas Jr., “Exploratory Design in Medical Nanotechnology: A
Mechanical Artificial Red Cell,” Artificial Cells, Blood Substitutes, and
Immobil. Biotech. 26(1998):411-430; http://www.foresight.org/Nanomedicine/Respirocytes.html;
Robert A. Freitas Jr., “Microbivores: Artificial Mechanical Phagocytes using
Digest and Discharge Protocol,” Zyvex preprint, March 2001; http://www.rfreitas.com/Nano/Microbivores.htm.

5. Robert A. Freitas Jr., Christopher J. Phoenix, “Vasculoid: A personal nanomedical appliance to replace human blood,” Journal of Evolution and Technology 11(April 2002); http://www.jetpress.org/volume11/vasculoid.html

6. K. Arvidson et al., “In vitro and in vivo experimental studies on single crystal sapphire dental implants,” Clin. Oral Implants Res. 2(April-June 1991):47-55.

7. Peter Griss, Gunther Heimke, “Biocompatibility of High Density Alumina and its Applications in Orthopedic Surgery,” in David F. Williams, ed., Biocompatibility of Clinical Implant Materials, Volume I, CRC Press, Boca Raton, FL, 1981, pp. 155-198.

8. Robert A. Freitas Jr., “Could Medical Nanorobots Be Carcinogenic?” Foresight Update No. 49, August 2002, pp. 8-9; http://www.imm.org/Reports/Rep033.php

9. G. Neupert et al., “Significance of scanning electron microscopy observations on cellular reactions in the biological evaluation of endoprosthetic materials,” Z. Exp. Chir. Transplant Kunstliche Organe 20(1987):139-143. In German.

10. M.A. Malik et al., “Osteoblasts on hydroxyapatite, alumina and bone surfaces in vitro: morphology during the first 2 h of attachment,” Biomaterials 13(1992):123-128.

11. M. Hashimoto et al., “Ultrastructure of the peri-implant junctional epithelium on single-crystal sapphire endosseous dental implant loaded with functional stress,” J. Oral Rehabil. 16(May 1989):261-270.

12. Keiichi Kawanabe et al., “Effects of injecting massive amounts of bioactive ceramics in mice,” J. Biomed. Mater. Res. 25(January 1991):117-128.

13. J. Harms, E. Mausle, “Tissue reaction to ceramic implant material,” J. Biomed. Mater. Res. 13(January 1979):67-87.

14. M. Di Silvestre et al., “Powdered alumina implants in the experimental animal: a histological study conducted in the rat,” Chir. degli Organi. di Mov. 76(April-June 1991):167-172.

15. P.S. Christel, “Biocompatibility of surgical-grade dense polycrystalline alumina,” Clin. Orthop. 282(September 1992):10-18.

16. A. Pizzoferrato et al., “The effect of injection of powdered biomaterials on mouse peritoneal cell populations,” J. Biomed. Mater. Res. 21(April 1987):419-428.

17. Y. Nakashima et al., “The stimulatory effects of ceramic particles on the production of the bone-resorbing mediators in vitro,” Trans. Orthop. Res. Soc. 20(1995):780.

18. Isabelle Catelas et al., “Cytotoxicity and macrophage cytokine release induced by ceramic and polyethylene particles in vitro,” J. Bone Joint Surg. Br. 81(May 1999):516-521.

19. J. Li et al., “Evaluation of biocompatibility of various ceramic powders with human fibroblasts in vitro,” Clin. Mater. 12(1993):197-201.

20. I. Dion et al., “Physicochemistry and Cytotoxicity of ceramics Part II: Cytotoxicity of ceramics,” J. Mater. Sci.: Materials in Medicine 5(1994):18-24.

21. A.M. Rodrigo et al., “Effects of polyethylene and alpha-alumina particles on IL-6 expression and secretion in primary cultures of human osteoblastic cells,” Biomaterials 23(February 2002):901-908.

22. H. Oonishi et al., “Comparative bone growth behavior in granules of bioceramic materials of various sizes, ” J. Biomed. Mater. Res. 44(January 1999):31-43.

23. I. Dion et al., “Blood haemolysis by ceramics,” Biomaterials 14(1993):107-110.

24. R.B. Schlesinger et al., “Clearance and translocation of aluminum oxide (alumina) from the lungs,” Inhal. Toxicol. 12(October 2000):927-939.

More on the medical applications of nanotechnology at:
http://www.foresight.org/Nanomedicine/index.html

IMM would appreciate learning your thoughts on the above article.

In conjunction with Foresight Update 51

Will Intracellular Medical Nanorobots Disrupt the Cytoskeleton (Part I)?

By Robert A. Freitas Jr.
Research Scientist, Zyvex Corp.

Robert A. Freitas Jr.

Active medical nanorobots [1, 2] maneuvering inside living cells could disturb or disrupt any of the many functions of the cytoskeleton. The cytoskeleton is the internal structural framework of a cell, consisting of several different types of filaments (see below). The functions of the cytoskeleton include: (1) mechanical support, such as cell movement, adhesive interaction with the extracellular matrix (ECM) and neighboring cells, and stabilization of cell shape including cellular “tensegrity”; (2) integration of diverse cellular activities, such as intracellular movement including transport and positioning to the appropriate locations of organelles, intracellular particles, RNA and proteins; and (3) intracellular signal transduction, which includes structural information regarding cellular origin and differentiation state [3]. In diverse cell types, microtubule and actin filament networks cooperate functionally during many processes, such as vesicle and organelle transport, cleavage furrow placement, directed cell migration, spindle rotation, and nuclear migration [4]. In principle, nanorobots could mechanically disrupt any or all of these functions during nanorobotic intracellular locomotion or manipulation of cell structures, especially if cytoskeletal/membrane links are disturbed.

Two classes of danger seem most significant. First is the risk of direct mechanical cytoskeletal reorganization, the subject of Part I of this article. The second is the risk of possible disruption of vesicular transport and related molecular motor diseases, the subject of Part II (next time). In both cases, it appears likely that potential problems can be avoided by conservative design.

Generalized disruption of the cytoskeleton can be very harmful to living cells. Disorganization of the cytoskeletal architecture has been associated with diseases as diverse as heart failure, rotavirus infection, sickle cell anemia, lissencephaly (smooth brain, lacking convolutions), and Alzheimer’s disease. A “collapse transition” of neurofilament sidearm domains may contribute to amyotrophic lateral sclerosis (ALS) and to Parkinson’s disease. Stress-related cytoskeletal fracture can be caused by 1-Hz cyclic forces imposed by a mechanical probe on isolated rat ventricular myocytes [5]. Cancer cells forced through 5- to 12-micron pores in polycarbonate membrane suffer traumatic spatial dissociation among components of the cell periphery, the cytoskeleton, and nucleus, inducing a ~1-week dormant state in the cells due to the mechanical trauma [6].

Nanorobots could induce various cell pathologies by mechanically disrupting specialized cytoskeletons consisting of cytoplasmic networks of ~6-nm diameter actin microfilaments, ~10-nm intermediate filaments, ~25-nm microtubules, or their many associated proteins ([1], Section 8.5.3.11), with effects similar to those of chemical disruption. Functions of these specialized cytoskeletons that could be disturbed include mechanical integrity and wound-healing in epidermal cells, cell polarity in simple epithelia, contraction in muscle cells, hearing and balance in the inner ear cells, axonal transport in neurons, and neuromuscular junction formation between muscle cells and motor neurons [7].

As a nanorobot enters a cell through the plasma membrane, the first risk is transmembrane linkage disruptions. Muscular dystrophy may be caused by disorganization of links between the intracellular cytoskeleton and the ECM [10], and the disruption of proper adhesive interactions with neighboring cells can lead to fatal defects in extracellular tissue architecture [11]. Epithelial cells subjected to mechanical strain may release in vivo proteases to cut intercellular adhesions [12]. Looking inward, cellular mechanoprotective adaptations involve a coordinated remodeling of the cell membrane and the associated cytoskeleton [13]. For example, the breakage of major cytoskeletal attachments between the plasma membrane and peripheral myofibers in cardiac myocytes predisposes the cell to further mechanical damage from cell swelling or from ischemic contracture [14]. As another example, attacking fungus cells (that are infecting parsley cells) extend a penetrating hypha through the cell membrane, eliciting a defensive cytoskeletal reorganization [15]. A local mechanical stimulus produced by a needle of the same diameter as the fungal hypha inserted through the host cell wall similarly induces the translocation of cytoplasm and nucleus to the site of stimulation, the generation of intracellular reactive oxygen intermediaries, and the expression of some elicitor-responsive genes. Without the mechanical stimulation, the morphological changes are not detected [15].

Mechanical disruption of cytoskeleton associated proteins by passing nanorobots could produce various cytopathologies. For instance, plectin is a 580 kD intracellular protein that links intermediate filaments with actin microfilaments, microtubules, and plasma membrane. Widespread disruption of plectin function results in severe skin blistering and muscular degeneration, consistent with plectin’s role in stabilizing cells against external mechanical forces [16] and as a regulator of intracellular actin dynamics [17]. Disturbance of centrosomes or other in cyto fixed polarity markers could result in developmental or morphogenetic defects during subsequent cell division. Mechanical disturbance of cytoskeleton associated proteins could also alter the mechanical properties of cells, such as the ability of the cytoskeleton to deform and flow.

Actin microfilaments might be disrupted by the mechanical activities of medical nanorobots. In the simplest case, endothelium exposed to shear stress undergoes cell shape change, alignment, and microfilament network remodeling in the direction of flow, though these changes can be blocked with nocodazole. Glomerular distention is also associated with cellular mechanical strain. A contractile cytoskeleton in mesangial cells, formed by F-actin-containing stress fibers, maintains structural integrity and modulates glomerular expansion. Mesangial cells have a cytoskeleton capable of contraction that is disorganized in long-term diabetes. Disorganization of stress fibers may be a cause of hyperfiltration in diabetes. Cutting the actin lattice may diminish both cell contractility [18] and mechanical signal transduction into the nucleus [19]. Care must also be taken to ensure that the surfaces or activities of intracellular nanorobots do not provide unplanned foci for actin polymerization, given that the kinetics of actin polymerization is autocatalytic and that the actin-based motility of functionalized microspheres can be reconstituted in vitro from only five pure proteins. Widespread actin disruption might produce symptoms analogous to elliptocytosis (abnormally large number of oval-shaped red cells in the blood) and other inherited hemolytic disorders that are caused by disorganization of the subsurface spectrin-actin cell cortex in the erythrocyte [20].

Intermediate filaments might also be disrupted mechanically. Perturbations in the architecture of the intermediate filament cytoskeleton in keratinocytes and in neurons can lead to degenerative diseases of the skin, muscle cells, and nervous system [7-9]. Knockout of the extensive keratin filament network jeopardizes the mechanical integrity of the epidermal cell, producing cell fragility and cytolysis manifesting as blistering skin disorders. Tissues lacking intermediate filaments fall apart, are mechanically unstable, and cannot resist physical stress, which leads to cell degeneration [21]. Perinuclear clumping of fragmented keratin intermediate filaments accompanies many keratin disorders of skin, hair, and nails. In active muscle, intermediate filaments play an important role in the organization and stabilization of myofibril-membrane attachment sites. Their disruption can eliminate the deep membrane invaginations that are normally present in the healthy sarcolemma (the membrane surrounding striated muscle fiber) [8]. Neuronal intermediate filaments are normally anchored to actin cytoskeleton. If this anchoring fails, the cell displays short, disorganized and unstable microtubules that are defective in axonal transport. Neuronal survival requires viable interconnects among all three cytoskeletal networks [9]. Impairment of normal axonal cytoskeletal organization in Charcot-Marie-Tooth disease results in distal axonal degeneration and fiber loss.

The microtubule cytoskeleton could become disorganized due to careless intracellular operations by nanorobots, possibly: (1) simulating congenital brain malformation; (2) giving results similar to treatment with vincristine, a microtubule depolymerizing drug that produces peripheral neuropathy in humans accompanied by painful paresthesias and dysesthesias (abnormal sensations of numbness, tingling, prickling, or burning); or (3) giving results similar to treatment with ethanol, leading to oxidative injury producing a loss of gastrointestinal barrier integrity. Mechanical disturbance of the microtubule cytoskeleton induces electrophysiological modification of cell-cycle-dependent EAG potassium channels in mammalian tissue cells, and mechanical strain can induce a major decline in tubulin production in osteoblasts [22]. Nanorobot mechanical operations could also induce buckling and loop formation of tubulin fibers, as has been observed [23] inside shrinking vesicles when the surface tension of the shrinking bubble overcomes the Euler buckling strength of the fibers — intracellular tubulin twisted into 5-micron tennis-racquet shapes has also been observed [24].

Microtubules allowed to form under microgravity conditions show almost no self-organization and are locally disordered, unlike microtubules formed in 1-g conditions. Nanorobotic manipulations of cytoskeletal elements that offset, reduce, or cancel the stimulative effects of normal gravity could produce the same sort of cellular architectural disorganization as observed under microgravity conditions ([1], Section 4.4.2) that alters the pattern of microtubular orientation.

A nanorobot with sharp edges that cuts a microtubule probably cleaves the hydrogen bonds between the alpha and beta monomers, rather than the covalent bonds within the monomers. This creates a new “plus” and “minus” end for the microtubule. In most cases this would not be fatal for the cell and in fact normally would have little impact because large-scale microtubule network patterns (e.g., asters, whorls, and interconnected pole networks) are self-assembling and are motor-molecule concentration-dependent [25]. Nevertheless, in cyto nanorobots should avoid physically severing cytoskeletal elements whenever possible. Simple estimates of mechanical strength ([1], Table 9.3) applied to typical fiber diameters suggest tensile failure strengths of ~170 pN for actin microfilaments, ~300-500 pN for microtubules and ~20,000 pN for intermediate filaments. Nanorobots should avoid applying local forces of these magnitudes or larger in the vicinity of such fibers.

Force thresholds for cellular activation ([2], Section 15.5.4.1) may be considerably less than the indicated tensile failure strengths. By the end of 2002, the absolute force thresholds for failure, the range of mechanical frequency responses, and the threshold fraction of disturbed cytoskeleton required to elicit cellular response all had yet to be precisely determined. For example, during mitosis a force of 15-20 pN is required to detach microtubule-bound chromosomes [26] but a tensile force of up to 210 pN is required to detach a microtubule from a kinetochore [27]. Moreover, a nanorobot presenting a 1-micron² forward surface during intracellular locomotion through a (20 micron)³ tissue cell intercepts only ~0.25% of the entire cytoskeleton during each 20-micron of transcellular travel. In cyto medical nanorobots may be restricted to speeds of ~10 microns/sec while traversing intracellular clear paths ([1], Section 8.5.3.12) and ~1 micron/sec during transfilamentary intracellular locomotion, with progressive resealing of cytoskeletal elements that must be temporarily severed to allow the nanorobot to pass ([1], Section 9.4.6). Intranuclear locomotion conservatively should progress no faster than natural chromosomal dragging rates during mitosis [26], or ~0.1 micron/sec, applying forces of at most ~50 pN ([1], Section 9.4.6).

Copyright 2003 Robert A. Freitas Jr. All Rights Reserved

References

1. R.A. Freitas Jr., Nanomedicine, Volume I: Basic Capabilities, Landes Bioscience, Georgetown, TX, 1999; http://www.nanomedicine.com

2. R.A. Freitas Jr., Nanomedicine, Volume IIA: Biocompatibility, Landes Bioscience, Georgetown, TX, 2003; http://www.nanomedicine.com

3. D. Broekaert, “Cytoskeletal polypeptides: cell-type specific markers useful in investigative otorhinolaryngology,” Int. J. Pediatr. Otorhinolaryngol. 27(May 1993):1-20.

4. B.L. Goode et al., “Functional cooperation between the microtubule and actin cytoskeletons,” Curr. Opin. Cell Biol. 12(February 2000):63-71.

5. G.C. Bett, F. Sachs, “Whole-cell mechanosensitive currents in rat ventricular myocytes activated by direct stimulation,” J. Membr. Biol. 173(1 February 2000):255-263.

6. H. Gabor, L. Weiss, “Mechanically induced trauma suffered by cancer cells in passing through pores in polycarbonate membranes,” Invasion Metastasis 5(1985):71-83.

7. E. Fuchs, “The cytoskeleton and disease: genetic disorders of intermediate filaments,” Annu. Rev. Genet. 30(1996):197-231.

8. R.B. Cary, M.W. Klymkowsky, “Disruption of intermediate filament organization leads to structural defects at the intersomite junction in Xenopus myotonal muscle,” Development 121(April 1995):1041-1052.

9. Y. Yang et al., “Integrators of the cytoskeleton that stabilize microtubules,” Cell 98(23 July 1999):229-238.

10. R.D. Cohn, K.P. Campbell, “Molecular basis of muscular dystrophies,” Muscle Nerve 23(October 2000):1456-1471.

11. C. Hagios et al., “Tissue architecture: the ultimate regulator of epithelial function?” Philos. Trans. R. Soc. Lond. B. Biol. Sci. 353(29 June 1998):857-870.

12. H.L. Kain, U. Reuter, “Release of lysosomal protease from retinal pigment epithelium and fibroblasts during mechanical stresses,” Graefes Arch. Clin. Exp. Ophthalmol. 233(April 1995):236-243.

13. K.S. Ko, C.A. McCulloch, “Partners in protection: interdependence of cytoskeleton and plasma membrane in adaptations to applied forces,” J. Membr. Biol. 174(15 March 2000):85-95.

14. M.D. Sage, R.B. Jennings, “Cytoskeletal injury and subsarcolemmal bleb formation in dog heart during in vitro total ischemia,” Am. J. Pathol. 133(November 1988):327-337.

15. S. Gus-Mayer et al., “Local mechanical stimulation induces components of the pathogen defense response in parsley,” Proc. Natl. Acad. Sci. (USA) 95(7 July 1998):8398-8403.

16. P.G. Allen, J.V. Shah, “Brains and brawn: plectin as regulator and reinforcer of the cytoskeleton,” Bioessays 21(June 1999):451-454.

17. K. Andra et al., “Not just scaffolding: plectin regulates actin dynamics in cultured cells,” Genes Dev. 12(1 November 1998):3442-3451; G. Wiche, “Role of plectin in cytoskeleton organization and dynamics,” J. Cell Sci. 111(September 1998):2477-2486.

18. J. Pourati et al., “Is cytoskeletal tension a major determinant of cell deformability in adherent endothelial cells?” Am. J. Physiol. 274(May 1998):C1283-C1289.

19. F. Guilak, “Compression-induced changes in the shape and volume of the chondrocyte nucleus,” J. Biomech. 28(December 1995):1529-1541.

20. Z. Zhang et al., “Dynamic molecular modeling of pathogenic mutations in the spectrin self-association domain,” Blood 98(15 September 2001):1645-1653.

21. M. Galou et al., “The importance of intermediate filaments in the adaptation of tissues to mechanical stress: evidence from gene knockout studies,” Biol. Cell 89(May 1997):85-97.

22. M.C. Meazzini et al., “Osteoblast cytoskeletal modulation in response to mechanical strain in vitro,” J. Orthop. Res. 16(March 1998):170-180.

23. M. Elbaum et al., “Buckling microtubules in vesicles,” Phys. Rev. Lett. 76(20 May 1996):4078-4081.

24. D. Kuchnir Fygenson et al., “Mechanics of microtubule-based membrane extension,” Phys. Rev. Lett. 79(1997):4497-4500.

25. T. Surrey et al., “Physical properties determining self-organization of motors and microtubules,” Science 292(11 May 2001):1167-1171.

26. A.J. Hunt, J.R. McIntosh, “The dynamic behavior of individual microtubules associated with chromosomes in vitro,” Mol. Biol. Cell 9(October 1998):2857-2871.

27. R.B. Nicklas, “Measurements of the force produced by the mitotic spindle in anaphase,” J. Cell Biol. 97(1983):542-548.

More on the medical applications of nanotechnology at:
http://www.foresight.org/Nanomedicine/index.html
Also, please check out http://www.nanomedicine.com/. Robert Freitas has acquired ownership of the domain, and has put up a cleaner and very extensively internally-linked version of Nanomedicine Vol. I at http://www.nanomedicine.com/NMI.htm

IMM would appreciate learning your thoughts on the above article.

In conjunction with Foresight Update 53

Will Intracellular Medical Nanorobots Disrupt the Cytoskeleton (Part II)?

By Robert A. Freitas Jr.
Research Scientist, Zyvex Corp.

Robert A. Freitas Jr.

Mechanical biocompatibility, or mechanocompatibility [1], is a critical issue in medical nanorobotics [1, 2]. For instance, active medical nanorobots maneuvering inside living cells could disturb or disrupt any of the many functions of the cytoskeleton. The cytoskeleton is the internal structural framework of a cell consisting of at least three types of filaments (microfilaments, microtubules, and intermediate filaments), forming a dynamic framework for maintaining cell shape and motion and allowing rapid changes in the three-dimensional structure of the cell ([2], Section 8.5.3.11).

Nanorobots could mechanically disrupt any of these functions during intracellular locomotion and manipulation of cell structures, if cytoskeletal/membrane links are disturbed. As already noted in Part I of this article, two classes of danger seem most significant. First, there is the risk of direct mechanical cytoskeletal reorganization, the subject of Part I (last time). Second, there is the risk of possible disruption of molecular motor-driven vesicular and organelle transport and related molecular motor diseases, the subject of Part II of this article. Three important examples of intracellular motor molecules are myosin (interdigitating myosin and actin filaments producing muscle contractions; [2], Section 6.3.4.2), dynein (drives motion of cilia; [2], Section 9.3.1.1), and kinesin (transport molecules carrying vesicles along microtubule tracks; [2], Section 9.4.6).

There are a variety of human disorders associated with dysfunction of cytoskeleton-based molecular motors, including, for instance, the motor-based diseases involving defective cellular myosin motors (e.g., implicated in hypertrophic cardiomyopathy [3]), spindle assembly- and function-related diseases [4], and other avenues for cellular malfunction. Cellular motors also participate in the self-organization of microtubule network structures [5]. But perhaps the most important function of molecular motors is intracellular vesicular transport, and most particularly axonal transport in neurons. Typically, organelles, vesicles and granules ~100 nm in diameter or larger are carried at a peak speed of up to ~2 microns/sec on the back of a 60-nm kinesin transport molecule ([2], Figure 9.32) that takes 8-nm ATP-powered steps along microtubule tracks running throughout the cell. (Mean unloaded kinesin motor speed is usually only 0.5-0.8 microns/sec.)

Conventional kinesin is a dimer of identical ~120 kD protein chains [6] with a diffusion coefficient D ~ 2 x 10^-11 m²/sec [5]. The vesicle-attached kinesin motor molecule steps toward the plus-end of microtubules by converting the energy of ATP hydrolysis to mechanical work.

Could a vesicle be dislodged from its microtubule track after being bumped by a passing intracellular nanorobot? A kinesin molecular mechanical detachment force of F_detach ~ 13 pN [8] requires that a 1-micron³ diamondoid nanorobot of mass m_nano ~ 2 x 10^-15 kg must impact and carry the cargo vesicle a distance S_vesicle ~ 1 nm to detach it at a constant velocity of v_nano >~ (2 S_vesicle F_detach / m_nano)^1/2 ~ 3500 micron/sec, well above the self-imposed 10 micron/sec intracellular locomotion speed limit ([1], Section 15.5.7.3.1) and even slightly exceeding the nanorobot instantaneous thermal velocity of ~2500 microns/sec in water at 37 °C ([2], Eqn. 3.3). The torsional stiffness of kinesin is so low that the molecule readily twists through more than 360° from its resting orientation [10], thus allowing the cargo to easily swivel out of the way of foreign impacting objects. And kinesin motors normally detach from a microtubule after a few seconds of travel anyway [6, 11]. Still, nanorobots should be able to exert mechanical forces well in excess of 13 pN, so care should be taken to minimize those motions and trajectories which are likely to produce kinesin detachment. Note that it has been found experimentally that intracellularly-placed microspheres (crude nanorobot analogues) experience enhanced diffusion over short time scales near the nucleus, possibly due to interactions with microtubule-associated motor proteins [12].

Could a nanorobot that has clamped or securely bound itself to one location on a microtubule prevent the passage of vesicular cargoes, causing vesicles to bunch up behind or to detach? And could kinesin motor molecules that encounter the obstruction be permanently damaged? Most likely a detached vesicle will reattach to a clear neighboring microtubule and continue its trek [13], or will reattach to the original microtubule downstream of the nanorobotic obstruction. In some cases organelles can attach to and move along multiple filaments simultaneously [14] using multiple motor molecules, potentially reducing the interference with forward motion from a single-filament blockage.

Coppin et al. [8] have carefully studied the behavior of kinesin molecules whose forward progress is mechanically constrained. Kinesin has a stall load of ~5 pN [7-9]. There is an increasing rate of dissociation with increasing load. Specifically, the dissociation rate is ~0.2/sec at 1 pN load, ~0.5/sec at 2 pN, ~1/sec at 3 pN, and ~2/sec at 5 pN load [8], rates which can be altered by the presence of microtubule-associated proteins [15]. However, super stall loads of 5-13 pN do not cause kinesin to walk backwards, “probably because of an irreversible transition in the mechanical cycle.” Rather, when super-stalled the kinesin motor most commonly takes a single backward movement and then dissociates (detaches) from the microtubule,* occasionally rebinding to the same microtubule (always at/below the stall load) and resuming its normal movement. This clearly demonstrates that the kinesin motor is still functional after being subjected to a dissociative induced stall. That is, a superload-induced detachment doesn’t “break” the biological motor. The conclusion is that similar mechanical interference by a nanorobot also should not damage a processive protein motor. Interestingly, a mutant form of kinesin with its ATP and microtubule binding sites decoupled has been found that binds so tightly to the microtubule that the motor cannot let go [16], crudely analogous to the case of a nanorobot that firmly grasps a microtubule for a period long in comparison to the timescale of kinesin procession.

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* Alternating back and forth movements also are observed at super-stall [8]. The dissociation rate increases with load as long as the motor is moving (up to 2 Hz at 5 pN), but then becomes independent of load once the motor stalls — e.g., the stall time is 0.57 sec, representing a dissociation rate of ~1.8 Hz, for either spontaneous (~5 pN) or induced (~12 pN) stalls. Interestingly, forward loads induce the kinesin motor molecule to step faster under a wide range of ATP concentrations [8]. Forward loads of 5 pN increase velocity by +200% if ATP concentration is rate-limiting (5-40 µM) or by +50% if ATP is saturated (1 mM), but forward loads >5 pN cause forward velocity to drop off sharply [8]. A small carboxyl domain acts as a switch that turns the motor off when the kinesin motor is not bound to cargo [17].

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A nanorobot ambulating along microtubules should endeavor to avoid applying lateral forces exceeding the kinesin detachment load of ~13 pN [8] which could have the effect of detaching associated vesicles as the nanorobot progresses — something like a tree limb being stripped of its leaves as it is pulled through a tight-fitting metal ring. Typically, the processive kinesin molecule only takes a few hundred steps before letting go [6, 11], so an occasional early detachment, by itself, should not induce cellular dysfunction. As long as the processive motor protein is not physically damaged, most of the detached vesicles should reattach and continue their journey after the nanorobot has passed by. The minimum spacing (maximum density) of kinesin motors along a microtubule is an axial repeat of 8 nm and experiments with isolated microtubules gliding on kinesin-coated glass find optimal motility at a 47 nm separation between kinesins [18]. But the vesicular transport of 100-2000 nm diameter organelles in cyto implies that propulsive kinesin contacts are normally spaced at least 0.1-2 microns apart along the microtubule tracks. This leaves plenty of unoccupied foothold space to allow nanorobots to avoid disrupting processive motor proteins already in transit.

Coordinated groups of in cyto medical nanorobots should avoid inadvertently corralling or bulldozing large numbers of vesicles or motor molecules into relatively small volumes within the cell, as such increases in motor molecule concentration could increase the local microtubule polymerization rate [5]. Analogously, in motor neuron diseases where vesicular transport is blocked by massive localized accumulations of kinesin molecules, the blockage can produce large axonal swellings in the motor neurons in human spinal cords and can disturb the machinery for anterograde fast axonal transport [19].

Intracellular traffic jams involving repositioned vesicles and organelles [20] appear to be initiated by the accumulation of stalled kinesin cargoes and are most commonly reported in neurons [21-23] where their effects are most serious. For example, axonal organelles transported by kinesin molecules that stall can cause organelle jams that disrupt retrograde as well as anterograde fast axonal transport, leading to defective action potentials, dystrophic terminals, reduced transmitter secretion and progressive distal paralysis that parallels the pathologies of motor diseases such as amyotrophic lateral sclerosis [23]. Stretch injury to axonal cytoskeleton resulting in major loss of microtubules disrupts fast axonal transport resulting in focal accumulation of membranous organelles and axonal swellings [22], and a chemically-created microtubule-free region can serve as a trap that causes axonally transported particles to accumulate into a swollen region [21]. Nanorobots should avoid creating such regions within the cell.

Similar considerations also apply during nanorobotic intranuclear operations, given the presence of myosin-based motors, RNA polymerase motors, and other motor molecules inside the cell nucleus.

Copyright 2003 Robert A. Freitas Jr. All Rights Reserved

References

1. R.A. Freitas Jr., Nanomedicine, Volume IIA: Biocompatibility, Landes Bioscience, Georgetown, TX, 2003; http://www.nanomedicine.com

2. R.A. Freitas Jr., Nanomedicine, Volume I: Basic Capabilities, Landes Bioscience, Georgetown, TX, 1999; http://www.nanomedicine.com

3. I. Rayment et al., “Structural interpretation of the mutations in the beta-cardiac myosin that have been implanted in familial hypertrophic cardiomyopathy,” Proc. Natl. Acad. Sci. (USA) 92(25 April 1995):3864-3868.

4. V. Mountain, D.A. Compton, “Dissecting the role of molecular motors in the mitotic spindle,” Anat. Rec. 261(15 February 2000):14-24.

5. T. Surrey et al., “Physical properties determining self-organization of motors and microtubules,” Science 292(11 May 2001):1167-1171.

6. S. Rice et al., “A structural change in the kinesin motor protein that drives motility,” Nature 402(16 December 1999):778-784.

7. K. Svoboda et al., “Direct observation of kinesin stepping by optical trapping interferometry,” Nature 365(21 October 1993):721-727.

8. C.M. Coppin et al., “The load dependence of kinesin’s mechanical cycle,” Proc. Natl. Acad. Sci. (USA) 94(5 August 1997):8539-8544.

9. M.E. Fisher, A.B. Kolomeisky, “The force exerted by a molecular motor,” Proc. Natl. Acad. Sci. (USA) 96(8 June 1999):6597-6602.

10. A.J. Hunt, J. Howard, “Kinesin swivels to permit microtubule movement in any direction,” Proc. Natl. Acad. Sci. (USA) 90(15 December 1993):11653-11657.

11. W.O. Hancock, J. Howard, “Kinesin’s processivity results from mechanical and chemical coordination between the ATP hydrolysis cycles of the two motor domains,” Proc. Natl. Acad. Sci. (USA) 96(9 November 1999):13147-13152.

12. A. Caspi et al., “Enhanced diffusion in active intracellular transport,” Phys. Rev. Lett. 85(25 December 2000):5655-5658.

13. Y. Wada et al., “Evidence for a novel affinity mechanism of motor-assisted transport along microtubules,” Mol. Biol. Cell 11(January 2000):161-169.

14. R.D. Vale et al., “Movement of organelles along filaments dissociated from the axoplasm of the squid giant axon,” Cell 40(February 1985):449-454.

15. B. Trinczek et al., “Tau regulates the attachment/detachment but not the speed of motors in microtubule-dependent transport of single vesicles and organelles,” J. Cell Sci. 112(July 1999):2355-2367.

16. H. Song, S.A. Endow, “Decoupling of nucleotide- and microtubule-binding sites in a kinesin mutant,” Nature 396(10 December 1998):587-590.

17. D.L. Coy et al., “Kinesin’s tail domain is an inhibitory regulator of the motor domain,” Nat. Cell Biol. 1(September 1999):288-292.

18. K.J. Bohm et al., “Speeding up kinesin-driven microtubule gliding in vitro by variation of cofactor composition and physicochemical parameters,” Cell. Biol. Int. 24(2000):335-341.

19. I. Toyoshima et al., “Kinesin and cytoplasmic dynein in spinal spheroids with motor neuron disease,” J. Neurol. Sci. 159(15 July 1998):38-44.

20. R.A. Spritz, “Multi-organelle disorders of pigmentation: intracellular traffic jams in mammals, flies and yeast,” Trends Genet. 15(September 1999):337-340.

21. H. Horie et al., “Effects of disruption of microtubules on translocation of particles and morphology in tissue cultured neurites,” Brain Res. 288(12 December 1983):85-93.

22. W.L. Maxwell, “Histopathological changes at central nodes of Ranvier after stretch-injury,” Microsc. Res. Tech. 34(15 August 1996):522-535.

23. D.D. Hurd, W.M. Saxton, “Kinesin mutations cause motor neuron disease phenotypes by disrupting fast axonal transport in Drosophila,” Genetics 144(November 1996):1075-1085.

More on the medical applications of nanotechnology at: http://www.foresight.org/Nanomedicine/index.html
Also, please check out http://www.nanomedicine.com/. Robert Freitas has acquired ownership of the domain, and has put up a cleaner and very extensively internally-linked version of Nanomedicine Vol. I at http://www.nanomedicine.com/NMI.htm

IMM would appreciate learning your thoughts on the above article.