Table of contents :

metallic prostheses : cobalt and chromium alloys, stainless steel, titanium, polyvinyl, polyuretane or polyesters (for soft tissues), tantalum or acrylic resins (for bony structures)

organic prostheses : polyethylene, teflon, dacron, nylon, polytef / polytetrafluoroethylene (PTFE) (a polymer of tetrafluoroethylene, used as a surgical implant material for prostheses, such as artificial vessels and orbital floor implants and for many applications in skeletal augmentation and skeletal fixation. Also used widely in industry, e.g., as an antistick coating for cooking utensils), silicone (for breast, testicular, and outer ear)

biological prostheses :

• autologous :
• saphena vein for tibial and coronary revascularization
• left internal mammary artery (LIMA) for coronary revascularization
• splenic artery and hypogastric artery for visceral revascularization
• homologous : human umbilical vein (Dardik prosthesis) covered by a dacron mesh
• heterologous : bovine carotid and mammary arteries, ovine collagen prostheses

Thrombotic threshold velocity (TTV) : the minimum velocity of blood flow through a vascular graft or prosthesis below which thrombogenicity would begin along the intimal surface; variables affecting it include lumen diameter and type of material the graft or prosthesis is made of

• alloplastic prostheses : 150-200 ml/min
• biological prostheses : 50-150 ml/min
• autologous saphena vein (without valves or with valves in opposite orientation) (in femoral and below knee distal bypass) : 30-50 ml/min

• implantology : the science dealing with the study and practice of inserting implants (an object or material, such as an alloplastic or radioactive material or tissue, partially or totally inserted or grafted into the body for prosthetic, therapeutic, diagnostic, or experimental purposes)
• engineering : the application of physical, mathematical, and mechanical principles to practical purposes.
• biomedical engineering / bioengineering : the use of engineering in biomedical technology such as the analysis of movement of body parts or prosthetics
• tissue engineering : application of the methods and principles of engineering to gain an understanding of the relationship between structure and function in both normal and pathological tissue and to develop possible substitutes for pathological tissue

• hydrophilic surfaces
• containing negatively charged chemical groups

• bone prostheses
• microinterlock : the growth of mineralized bone into the porous surface of a specially constructed prosthesis, anchoring the prosthesis into place.
• vertebroplastic for vertebral fractures : polymethylmethacrylate (PMMA) solidifies and the exothermic reaction burns periosteal sensitive fibers.
• cranioplasty (skull prostheses) : plastic operation on the skull; surgical correction of defects of the skull. Porous hydroxyapatite which can be colonized by osteoblasts within 5-6 months.
• durapatite : a crystalline form of the compound (Ca₃(PO₄)2)3·Ca(OH)2 (see hydroxyapatite), used as a prosthetic aid
• duraplasty : a plastic operation on the dura mater; graft of the dura.
• arthroplasty / joint replacement : plastic surgery of a joint or of joints; the formation of movable joints
• arthroprosthesis / total joint arthroplasty / total joint replacement : arthroplasty in which both sides of a joint are removed and replaced by artificial implants anchored to the bones
• TMJ prostheses :
• interposition arthroplasty : surgical correction of ankylosis of the temporomandibular joint by separating the immobile fragment from the mobilized fragment and interposing a substance, such as fascia, cartilage, metal, or plastic, between them.
• intracapsular temporomandibular joint arthroplasty : operative recontouring of the articular surface of the mandibular condyle without the removal of the articular disk.
• hip prostheses :
• total hip arthroplasty (THA)
• Austin Moore arthroplasty : hip reconstruction surgery using an Austin Moore prosthesis (metallic implant)
• Charnley's hip arthroplasty : a hip replacement operation involving insertion of Charnley's prosthesis (acetabular cup and a relatively small femoral head component that form a low-friction joint)
• Thompson arthroplasty : hip reconstruction surgery using a Thompson prosthesis (a Vitallium implant)
Complications :
• dislocation (2.4%)^ref
• periprosthetic fracture of the femur : Vancouver classification^ref :
• type A fractures involve the trochanteric region and are subclassified into AG fractures involving the greater trochanter, and AL fractures involving the lesser trochanter
• type B fractures occur around the stem of the femoral component, or extend slightly distal to it. They are further subclassified based on the stability of the implant and the quality of the bone stock.
• subtype B1 fractures : the implant is stable and the bone stock around the femoral component is adequate
• subtype B2 : the implant is unstable and the bone stock around the femoral component is adequate
• subtype B3 : the implant is unstable and the bone stock around the femoral component is not adequate
• type C fractures are located distal to the stem.
In a review of 1,442 THAs, Löwenhielm et al.^ref found the accumulated risk of postoperative fractures to be 25 per 1000 over a 15-year period, while Beals and Tower^ref estimated the incidence of such fractures over the lifespan of a prosthesis to be less than 1%. A recent review from the Mayo Clinic Joint Registry identified 262 (1%) postoperative periprosthetic femoral fractures following 23 980 primary THAs, and 252 (4%) such fractures following 6,349 revision THAs^ref
• 63% remained stable after one or more closed reductions : closed treatment of postoperative periprosthetic femoral fractures is associated with high nonunion and malunion rates, and frequently results in subsequent revision^{ref1, ref2, ref3}
• 37% ultimately required revision surgery : in cases of stable prostheses, open reduction and internal fixation employs either :
• cortical onlay allografts^ref
• dynamic compression plates (DCPs)^ref
• special plates such as the Ogden, the Dall-Miles cable and plate fixation system or the Mennen plates^{ref1, ref2, ref3, ref4, ref5, ref6}
• combined medial and lateral allograft struts with cerclage cables or lateral plating with a single anterior or medial allograft strut may be used^{ref1, ref2}
• furthermore, Haddad et al.^ref reported good results following the combination of a Dall-Miles plate with two strut grafts and cancellous autograft in four patients. In cases of a peri-prosthetic femoral fracture necessitating cemented revision, DCPs may also be used for reduction during the prosthesis insertion; they may be fixed with unicortical screws around the prosthesis and bicortical screws below it.
• Charnley-D'Aubigne-Postel score : Charnley's modification of the Merle D’Aubigne and Postel hip scoring system^ref
• pain
• 0 intense and permanent;
• 1 severe, even at night;
• 2 severe on walking, prevents all activity
• 3 tolerable, limiting activity
• 4 mild with activity, rest relieves
• 5 slight pain, less with activity
• 6 no pain
• function
• 0 cannot walk
• 1 only with crutches
• 2 time and distance very limited without sticks
• 3 1 hour with stick, very difficult without;
• 4 long time with stick: short without + limp
• 5 without stick but slight limp
• 6 normal
• movement : total range of movement of hip joint = flexion +abduction + adduction + internal rotation + externalrotation – any fixed deformity)
• 1 0–30°
• 2 31–60°
• 3 61–100°
• 4 101–160°
• 5 161–210°
• 6 >210°
• category :
• A patient fit for age except for affected hip
• B both hips affected but patient otherwise fit for age
• C patient has inbuilt limitations other than from hips
• knee prostheses :
• total knee arthroplasty : arthroplasty of both sides of the knee joint.
• foot prostheses :
• Flex-Foot : trademark for a type of ankle-foot prosthesis that includes a flexible rod replacing the lower leg and foot as well as a heel replacement component allowing push-off.
• solid ankle cushion heel (SACH) foot : a type of prosthetic foot that allows flexion, extension, inversion, and eversion
• Seattle foot : a type of prosthetic foot that allows flexion, extension, inversion, and eversion, and allows some push-off when walking; it is made of plastic and resembles a normal human foot.
Bacteria in prosthetic joint infections^ref :
• coagulase negative staphylococci, 30-43%
• Staphylococcus aureus, 12-23%
• Streptococci, 9-10%
• Enterococci, 3-7%
• Gram negative bacilli, 3-6%
• anaerobes, 2-4%
• polymicrobial, 10-12%
• unknown, 10-11%
Treatment of large defects requires the harvest of fresh living bone from the iliac crest. Harvest of this limited supply of bone is accompanied by extreme pain and morbidity. This has prompted the exploration of other alternatives to generate new bone using traditional principles of tissue engineering, wherein harvested cells are combined with porous scaffolds and stimulated with exogenous mitogens and morphogens in vitro and/or in vivo. Large volumes of bone can be engineered in a predictable manner, without the need for cell transplantation and growth factor administration. The crux of the approach lies in the deliberate creation and manipulation of an artificial space (bioreactor) between the tibia and the periosteum, a mesenchymal layer rich in pluripotent cells, in such a way that the body's healing mechanism is leveraged in the engineering of neotissue. Using the "in vivo bioreactor" in New Zealand White rabbits, bone that is biomechanically identical to native bone has been engineered. The neo-bone formation followed predominantly an intramembraneous path, with woven bone matrix subsequently maturing into fully mineralized compact bone exhibiting all of the histological markers and mechanical properties of native bone. The bone was harvested after 6 weeks and transplanted it into contralateral tibial defects, resulting in complete integration after 6 weeks with no apparent morbidity at the donor site. Furthermore, in a proof-of-principle study, by inhibiting angiogenesis and promoting a more hypoxic environment within the "in vivo bioreactor space," cartilage formation can be exclusively promoted^ref.
• distraction osteogenesis : a process whereby attachment of the device across a surgical bone cut separates the bone ends in a precise, continuous manner that results in the formation of healthy bone tissue at a rate of 1 mm per day. The bone generator device (GenerOs SB, made by Orthonetx, Inc.) corrects bone deficiencies and deformities in small bones of the hands, wrists, feet, and ankles, and pediatric forearms. Previously approved GenerOs devices include a craniofacial bone generator (GenerOs CF), a limb strengthener (GenerOs EX), and a dental implant distractor (GenerOs DI).
• cineplastic surgery : creation of a skin-lined tunnel through a muscle adjacent to the stump of an amputated limb, to permit use of the muscle in operating a prosthesis.
• antireflux prosthesis : a ring-shaped device that is placed around the esophagus above the stomach and below the diaphragm for treatment of gastroesophageal reflux and hiatal hernia
• Angelchik prosthesis : a C-shaped silicone device that wraps around the distal esophagus for treatment of reflux esophagitis or during laparotomy.
• oral prostheses
• cleft palate prosthesis : a prosthetic device, such as an obturator, used to correct cleft palate
• dental prosthesis : a replacement for one or more of the teeth or other oral structure, ranging from a single tooth to a complete denture
• implantodontics / dental or oral implantology / implantodontology : the branch of dentistry dealing with the implantation of artificial devices and materials into the oral hard and soft tissues for prosthetic, therapeutic, or diagnostic purposes
• dental implant : a prosthetic device of alloplastic material implanted into the oral tissues beneath the mucosal or periosteal layer or within the bone to provide support and retention to a partial or complete denture.
• endosseous or endosteal implant : a dental implant made of metal or sometimes ceramic or polymeric material, consisting of a blade, screw, pin, or vent, inserted into the jaw bone through the alveolar or basal bone, either directly or through the root canal and apex of a tooth, with a post protruding through the mucoperiosteum into the oral cavity to serve as an abutment for dentures or orthodontic appliances, or to serve in fracture fixation.
• osseointegrated implant : an endosseous implant containing pores into which osteoblasts and supporting connective tissue can migrate. Metallic, ceramic, and polymeric materials have been used.
• Brânemark implant : a type of osseointegrated implant consisting of a two-stage system of titanium screws.
• intraperiosteal implant : a frame that conforms to the shape of the jawbone and is implanted beneath the outer or fibrous layer of the periosteum so that it rests firmly on the bone, with a post protruding into the oral cavity to serve as an abutment for dentures; used most commonly for upper fixed partial dentures and in the treatment of cleft palate.
• subperiosteal implant : a dental implant consisting of a metal frame implanted under the periosteum and firmly bound by the mucoperiosteum, resting on the jaw bone, with a post protruding into the oral cavity; used most commonly as an abutment for upper fixed partial dentures and in the treatment of cleft palate.
• transmandibular implant : a dental implant consisting of baseplate, cortical screws, threaded posts, and a superstructure that attaches to a denture, inserted via submandibular incision and fixed to the symphyseal border. The implant transverses the mandible, bears the denture directly, and is designed for patients with severe mandibular alveolar atrophy.
• Bosker implant : a common type of transmandibular implant
• transosteal implant : a dental implant consisting of a bolt that crosses the mandible and is secured in place by a nut; a post protrudes into the oral cavity and serves as an abutment for dentures or orthodontic appliances.
• endodontic implant : a metallic implant extending through the root canal of a tooth into the periapical bone structure, thereby lengthening the root of a pulpless tooth.
• magnet implant / denture magnet : a magnet made of a nonreactogenic platinum-cobalt alloy or a rare earth element, used for additional retention of dentures. One magnet is implanted into the mandible under the periosteum and the other is attached to the denture, its poles being opposite of those in the mandible
• the conventional treatment of dental caries involves mechanical removal of the affected part and filling of the hole with a resin or metal alloy. But this method is not ideal for tiny early lesions because a disproportionate amount of healthy tooth must be removed to make the alloy or resin stick. A dental paste of synthetic enamel rapidly and seamlessly repairs early caries lesions by nanocrystalline growth, with minimal wastage of the natural enamel^ref.
• maxillofacial prosthesis : a prosthetic replacement for those regions in the maxilla, mandible, and face that are missing or defective because of surgical intervention, trauma, pathology, or development malformations
• a major goal of research in bone transplantation is the ability to avoid creation of secondary bone defects.
• heterotopic bone induction : custom-built, bioengineered jaw transplant was implanted in a 56-year-old man in June 2004, who had part of his jawbone removed after a tumour spread through the area. A 3-D CT scan was taken of the patient's jaw (upper) and used to generate a virtual jawbone (lower), then used to create a scale model made of Teflon. They then fitted a lightweight titanium mesh around the replica, and removed the Teflon to leave a perfectly sculpted, hollow, U-shaped tube. The cylinder was filled with a natural bone mineral called hydroxyapatite, 20 mL of blood from the patient's own bone marrow, and a dash of  7 mg bone morphogenetic protein 7 (BMP7). The bone mineral offered structural support as BMP7 coaxed stem cells inside the blood to form new bone. Next, the replacement jaw needed to develop its own blood supply, so the team grafted it into a region rich in blood vessels inside the man's right latissimus dorsi muscle. 7 weeks later, medics removed the mandible and slotted it into the skull. Microsurgery was used to connect the jaw's newly developed blood supply with that of the head. In-vivo skeletal scintigraphy showed bone remodelling and mineralisation inside the mandibular transplant both before and after transplantation. CT provided radiological evidence of new bone formation. Postoperatively, the patient had an improved degree of mastication and was satisfied with the aesthetic outcome of the procedure. By the 4th week post-transplantation the patient enjoyed his first dinner in 9 years (bread and sausages); before reconstruction he had only been able to eat soft food and soup^ref. The procedure is an advance on alternative methods, where surgeons remove bone from hip or leg bones and carve it into a rough jaw shape before grafting it into the face. The jaw is rarely a perfect fit and it injures the hip or leg : this can cause pain and joint problems. Before the operation, the patient had a cosmetic, titanium jawbone that was covered with skin. Although this looked reasonably normal, the patient was unable to chew and dined solely on soft food and soup. Around 10% of surgery centres worldwide are set up to perform operations like this
• speech-aid prosthesis / artificial palate or velum / obturator / prosthetic speech aid : a device designed to close a cleft of the hard or soft palate and to substitute for structures used for normal speech
• palatal lift prosthesis : a type of speech-aid prosthesis sometimes used to aid in closure in cases of velopharyngeal insufficiency.
• artificial larynx / electrolarynx : an electromechanical device that enables a totally laryngectomized person to speak. When it is placed against the region of the laryngectomy a buzzing sound is produced, which is converted into simulated speech by movements of the organs of articulation (lips, tongue, glottis)
• penile implants or prosthesis : a semirigid rod or inflatable device implanted in the penis to provide an erection in men with organic impotence. Much has been written about sildenafil, which has revolutionised the treatment of erectile dysfunction. However, in some countries, including the United Kingdom, the cost of sildenafil is not reimbursed by the state for most men with erectile dysfunction. General practitioners in the United Kingdom can now prescribe the drug privately, and several guidelines have been produced to inform and educate general practitioners about the investigation and treatment of erectile dysfunction. Insertion of penile prostheses is still required for men in whom sildenafil is ineffective and who do not respond to or dislike intracavernosal injections or vacuum devices, particularly men with diabetes, cardiovascular disease, or neurological disease. Surgical treatment may also be required in some patients with Peyronie's disease. In penile prostheses this can be devastating and, frequently, results in removal of the device despite aggressive antibiotic therapy^ref. The incidence of infection associated with primary implantation ranges from 1 to 3%^{ref1, ref2, ref3}. In reoperation for revision or reimplantation the infection rate is higher, with reported rates ranging from 10 % to 13 %^{ref1, ref2, ref3}. In recent years, new strategies have been developed in an attempt to minimize this risk, expecially the use of an antibiotic coating on the device^ref
• 3-piece inflatable implants (75%) : inflatable cylinders in the corpora cavernosa, small pump in the scrotum and reservoir filled with saline solution placed deep in the abdomen. Gently squeezing the concealed pump in your scrotum several times moves the saline solution from the reservoir into the cylinders. As the cylinders fill, the penis becomes erect and firm. To end the erection, simply press a "deflation site" on the pump. Deflating the cylinders transfers the fluid back to the reservoir and the penis becomes flaccid
• 2-piece inflatable implants (15%) : inflatable cylinders in the corpora cavernosa and a small pump in the scrotum. The device is completely concealed and unobtrusive. The pump inflates the cylinders by transferring fluid within the system. A simple deflation technique then transfers fluid to return the penis to a flaccid state
• semi-rigid malleable (positionable) rods (10%) : 2 bendable, "positionable" rods into the penis. The rods have an outer coating of silicone and inner stainless steel core or interlocking plastic joints. For intercourse, the man places his penis in the erect position. As he does so, the rods inside his penis bend. To conceal the penile implant, the man bends his penis down. He can bend the device in more than one place to create the desired erection.
The device is completely hidden and unobtrusive. It commonly takes about 4-6 weeks to regain sexual function. Complications can occur, including infection and device malfunction
• Hydroflex penile prostheses
• AMS hydraulic penile prostheses
• vacuum erection device has 3 main parts: a plastic cylinder that the man places his penis into, a hand or battery operated pump that draws air from the cylinder and creates a vacuum that allows increased blood flow to the penis, causing an erection, and 1 or more tension rings placed on the base of the penis after vacuum erection is created to help maintain the erection by reducing blood flow out of the penis. A vacuum erection device will generate an erection that lasts for up to 30 minutes aided by the tension ring. Some side effects are possible: reddish pinpoint-size dots on the surface of the penis when negative pressure is applied too rapidly, some bruising if the penis is held under pressure too long, uncomfortable cold glans as the device and tension rings limit circulation, impaired ejaculation, and pain ata application of the tension rings. The vacuum erection device can be cumbersome to use and interfere with romantic spontaneity : for that reason, a vacuum erection device probably is best suited to a man in a stable relationship with a supportive sexual partner.
• neural prosthetics
• auditory brainstem implant : the past 15 years has seen much progress in restoring hearing to deaf patients. Cochlear implants were successfully introduced for patients with inner ear deafness. However, in rare cases when both auditory nerves are destroyed, such as in neurofibromatosis type 2 or after traumatic lesions, a cochlear implant cannot be used. Auditory brainstem prostheses have been developed for these patients that work by electrical stimulation of the second auditory neurone in the brainstem. The implant has been developed from single channel electrodes positioned in the brainstem^ref, to the 21 channel model currently used in Europe. Like patients with cochlear implants, patients with auditory brainstem implants wear a body level speech processor and a transmission coil on the scalp. The electrodes are placed in the lateral recess of the fourth ventricle, adjacent to the cochlear nuclei. Hearing results are not as good as with cochlear implants^ref. Only a few patients with auditory brainstem implants can understand closed set speech without lip reading, but many patients report that the device improves their conversational ability significantly because it helps lip reading^{ref1, ref2}. Experience with multichannel auditory brainstem implants is limited because of the low incidence of patients with neurofibromatosis, but the implants are a promising way of improving hearing in these patients. Continuing research and refinement of hardware and software will help to improve the hearing of patients with neurofibromatosis further.
• hearing aids : amplify sound
• bone anchored hearing aids (BAHA)
• implantable hearing aid
• SOUNDTEC semi-implantable hearing aid
The recent development of digital hearing aids has been a big step forward in overcoming hearing impairment^{ref1, ref2}. The hearing aids work by converting an incoming analogue signal into a digital one, which is then processed by a microprocessor and then changed back into an analogue signal before output^ref. Schematic illustration of a digital hearing aid (modified according to Steffens^{ref1, ref2})

The main advantages of the digital technique are that the signal can be processed to suppress noise or to amplify specific frequencies or sounds^ref. The latest generation of digital hearing aids can recognise speech, which allows them to suppress other noise and make hearing clearer. The multiprogram devices can adapt to specific situations in the patients' everyday lives, especially with noise suppression, which further improves the ability to hear. The enhanced sound quality of the digital amplifying system also make voice and sound more similar to reality, increasing the acceptability to patients. Other advantages include greatly reduced feedback and easier handling. The development of "real" binaural interactive signal processing may provide even better hearing in noisy environments^ref.
Hearing loss of almost any extent can be ameliorated with a hearing aid. For conductive hearing loss this is simply a matter of amplification, although for sensorineural hearing loss the mechanism is more complex. However, there are many factors which will interfere with a patient's satisfaction with, and benefit from, a hearing aid. Lack of motivation because of fear of stigmatisation, low expectations of benefit, or failure to accept there is a problem remain significant obstacles. Stephens et al showed that despite the fact that 50% of those aged 50-65 in two villages in South Wales had a hearing disability only 7% had a hearing aid^ref. Clinicians have an important role in identifying those who would benefit from a hearing aid and emphasising the benefits of its use. Hearing aids should have electroacoustic characteristics which make speech audible but comfortable. All consist of a microphone, which converts acoustic signals to electrical signals, an amplifier which selectively processes the output signals, a receiver which converts the electrical signal back to an acoustic signal, and an earmold and tubing to deliver this to the patient's ear. All hearing aids available on the NHS are behind-the-ear types (although it is common practice to supply some other aids to war veterans whose hearing impairment is due to bomb blasts).In behind-the-ear aids the microphone, amplifier, and receiver are in a crescent shaped plastic case that rests behind the ear. A small tube connects this to the earmold. This style remains popular with older people as it can provide higher gain and the larger controls are easier to manipulate. For those with severe hearing impairment or manual dexterity problems larger devices are needed, but these can be worn quite unobtrusively attached to clothing. Also available on the market from registered dispensers are in-the-ear and in-the-canal styles
• behind-the-ear (BTE)



• in-the-ear (ITE)
• in-the-canal (ITC)
• body worn aids


On most hearing aids there are 3 switch positions: O, T, and M. At the O position the hearing aid is off, and M denotes the microphone is on. At the T position (telecoil on) the aid can pick up signals from transmitting devicesfor example, certain telephones and loop systems which are now incorporated into the majority of concert halls, theatres, and customer service points at post offices and banks. A sympathetic hearing scheme symbol (ear with a line through it) denotes a loop system is in operation. A frequently encountered problem is a constantly whistling hearing aid. The clinician must exclude a canal blocked with wax, and ensure the hearing aid has not become unseated before referral for audiological reassessment. Like dentures, hearing aids may become loose when patients loose weight and new earmolds may need fitting. Quality of life improvements with hearing aids occur after 6 weeks and can be shown to be sustained after 1 year^ref. Such benefits are comparable in younger and older adults^ref
• prosthetic stapes
• Robinson prosthesis : a type of stainless steel prosthetic stapes.
• cochlear implants
• Clarion cochlear implant
• Iowa/Nucleus 10 mm electrode
• endosteal electrode : residual hearing might better be preserved if the implanted electrode permits the fluid-filled inner ear space to remain intact^ref
A cochlear implant is a small, complex electronic device that can help to provide a sense of sound to a person who is profoundly deaf or severely hard of hearing. The implant is surgically placed under the skin behind the ear. An implant has 4 basic parts :
• microphone, which picks up sound from the environment;
• speech processor, which selects and arranges sounds picked up by the microphone;
• transmitter and receiver/stimulator, which receive signals from the speech processor and convert them into electric impulses;
• electrodes, which collect the impulses from the stimulator and send them to the brain.
An implant does not restore or create normal hearing. Instead, under the appropriate conditions, it can give a deaf person a useful auditory understanding of the environment and help him or her to understand speech. Cochlear implants compensate for damaged or non-working parts of the inner ear. When hearing is functioning normally, complicated parts of the inner ear convert sound waves in the air into electrical impulses. These impulses are then sent to the brain, where a hearing person recognizes them as sound. A cochlear implant works in a similar manner. It electronically finds useful sounds and then sends them to the brain. Hearing through an implant may sound different from normal hearing, but it allows many people to communicate fully with oral communication in person and over the phone. Different types of deaf and severely hard of hearing people choose cochlear implants. Both children and adults can be candidates for implants. According to the Food and Drug Administration 2002 data, approximately 59,000 people worldwide have received implants. In the USA, about 13,000 adults have cochlear implants and nearly 10,000 children have received them. Adults who have lost all or most of their hearing later in life can often benefit from cochlear implants. These older candidates can often associate the sounds made through an implant with sounds they remember. This may help them to understand speech without visual cues or systems such as lipreading or sign language. Young children can also be candidates for implants. Cochlear implants, coupled with intensive post-implantation therapy, can help young children to acquire speech, language, developmental, and social skills. The best age for implantation is still being debated, but most children who receive implants are between 2 and 6 years old. Earlier implantation seems to perform better. A cochlear implant is a surgical procedure. The decision to receive an implant should involve discussions with many medical specialists and an experienced surgeon. The process is expensive. Some may choose not to have a cochlear implant for a variety of personal reasons. Also, though surgical implantation is almost always safe, complications are a risk factor, just as with any kind of surgery. An additional consideration is learning to interpret the sounds created by an implant. This process takes time and practice. Speech-language pathologists and audiologists are the professionals frequently involved in this learning process. Not everyone performs at the same level with a cochlear implant. Prior to implantation, all of these factors need to be discussed. The technology behind cochlear implants is changing rapidly. With advancements in technology and continued follow-up research with people who have already received implants, researchers are evaluating new opportunities and additional possible candidates for cochlear implants.

Our ability to learn language is already on the wane by our third year of life, according to a study of profoundly deaf children given cochlear implants to restore some of their hearing. The research supports the widely held belief that there is a 'sensitive period' for language learning, during which the capacity to acquire vocabulary and grammar is heightened.  The implants, which are surgically inserted in the ear, convert sound into electrical signals that the brain can interpret, allowing many deaf people to hear. The rate of language learning measured with the widely used Reynell Developmental Language Scales, which test a child's vocabulary and understanding of grammar, was greatest for those given implants before they turned 2 (still relatively rare). Children given implants at 3 or 4 years of age acquired language skills more slowly, although these children still benefited from the devices. Some evidence for a sensitive period of language formation comes from the tragic cases of feral or isolated children, who were cut off from any language stimulus in their formative years and never learned to speak using proper syntax. But the physical and emotional deprivation that these children suffer almost certainly affects their ability to interact with people, having a knock-on effect on their language skills. It's hard to tease out those effects. The ultimate test would be what Svirsky calls the 'forbidden experiment': deliberately depriving children of all language input, which is obviously impossible. Instead, many researchers in this field observe the progress of profoundly deaf children given implants. It's not a perfect way of studying language development, but it's a new and exciting source of evidence. The Food and Drug Administration approves cochlear implantation in children as young as 12 months old, and even earlier in special cases, but < 33% of patients receive it before age 2. Cochlear implantation is deeply controversial. The procedure has sometimes been billed as a 'cure' for deafness, but a lot of deaf people don't feel the need to be 'cured'. Parents should not be rushed into making decisions about surgery on their child : early communication isn't just about speech, pointing out that the child's sensitive period is just as useful for developing sign language. Cochlear implants are not a miracle for most children : the best solution would be for children with implants to learn sign language and speech simultaneously, as a bilingual approach seems to boost general language skills.
Congenital deafness results in abnormal synaptic structure in endings of the auditory nerve. If these abnormalities persist after restoration of auditory nerve activity by a cochlear implant, the processing of time-varying signals such as speech would likely be impaired. Congenitally deaf cats were stimulated for 3 months with a 6-channel cochlear implant. The device used human speech-processing programs, and cats responded to environmental sounds. Auditory nerve fibers exhibited a recovery of normal synaptic structure in these cats. This rescue of synapses is attributed to a return of spike activity in the auditory nerve and may help explain cochlear implant benefits in childhood deafness^ref
Web resources :
• Cochlear implants at NIDCD
• American Academy of Otolaryngology-Head and Neck Surgery (AAO-HNS)
• Alexander Graham Bell Association for the Deaf and Hard of Hearing (AG Bell)
• Cochlear Implant Association, Inc. (CIAI) Executive
• House Ear Institute (HEI)
• League for the Hard of Hearing (LHH), New York
• Combined Health Information Database (CHID)
• bionic restoration of movement : many paralysed people control computers with their eyes or tongue. But muscle function limits these techniques, and they require a lot of training.

bionic approach (with electrodes implanted or external)	candidate users	actuator types
brain-machine interface (BMI)	completely paralyzed persons who can benefit from mechanical assist devices	motors, i.e., robotic arms controlled by monkeysref
brain-computer interface (BCI) invasive noninvasive	completely paralyzed persons who wish to simply communicate	computer screen, i.e., moving a cursorref Braingate (source : Cyberkinetics Inc) recognizes brain signals for movement still generated by immobile patients and converts them into action on a computer screen. A 16-mm2 chip containing 100 electrodes is implanted into the brain during a 3-hour procedure to record cell activity while an external PC sized processor converts the signals from the simultaneous firings of 100 individual neurons into computer instructions. FDA gave approval to implant chips in the motor cortex region of 5 quadriplegic patients to give them mouse control and computer access : in June 2004, surgeons implanted it into the motor cortex of a 24-year-old quadriplegic, allowing the patient to control a computer or television using his mind, even when doing other things at the same time. Macaque monkeys bearing implants to the parietal and primary motor cortices can learn, via visual feedback, to make a robotic arm reach, grasp, and gripref. In an algorithm called "continuous shared control", the robots contain tactile sensors that provide local intelligence – akin to a cat's whiskers : controlling the robot via a combination of local intelligence and monkey brainpower seems to make the tasks a lot more reliable than having solely brain-based control. BrainGate could be coupled with functional electrode stimulation (FES), in which partially paralyzed patients achieve some control of paralyzed limbs by stimulating electrodes placed in the muscles of those limbs. Current FES systems rely on external signals, but now imagine if we could hook up the sensor directly to this FES system : now by thought alone these people could be controlling their arm musclesref Brain Communicator (source : Neural Signals) : requires 10-hr surgery Implanting chips in the parietal cortex might yield unexpected side-effects. Suppose you planned to shake your boss's hand, but thought transiently about slapping him in the face. The slap could happen. Training would soon rule out unwanted responses. And the ideal brain-chip would tap into many different brain regions, coordinating planned actions with instructions for movement. But there are many hurdles to overcome before neural prosthetics become a reality. Current chips function for up to a year and are connected to a computer by wires. The next generation will need to be longer lasting and wireless. Normal and spinal-injured human patients could move a cursor on a computer screen by controlling their own mu and beta rhythms in EEGref. The level of control displayed in this experiment far surpasses that of a similar 1994 studyref : this advance could be attributed in part to improvements in signal processing, but mostly to a new adaptive algorithm that adjusts the EEG features on which it focuses based on the user's past performance. The P300 component of event-related brain potentials (ERP)ref can be harnessed to help a BCI to identify which of 36 characters and symbols displayed on a screen a disabled user has selectedref : this system generates just 7.8 characters per minute with 80% accuracyref. And, like their implantable counterparts, EEG-based methods can be cumbersome, so widespread adoption will likely require wireless technologies : P300 BCI can be implemented in a device the size of a Palm Pilot. Proponents of invasive technologies argue that noninvasive scalp recordings are surrogates that do not actually read motion intent from neurons. Syed has grown snail neurons on a silicon chip and induced them to form a network whose activity was manipulated at the single-cell level through the chipref. A capacitor located underneath one cell induces activity in that cell, while a transistor attached to the chip detects electrical activity in a second cell, providing evidence that the chip can "talk" to the cell and vice versa. That is a critical step for a successful interfacing of electronic devices with the brain cells, whereby the machine talks to the brain and the brain can talk back to the machine : Syed currently is developing chips interfaced with 10,000 to 16,000 neurons that can be used to study how various stimuli affect neural circuits. At the University of Florida, biomedical engineer Thomas DeMarse cultures rat neurons on 60-channel multielectrode arrays that, like Syed's, can be used both to record from and stimulate the cells. DeMarse recently garnered considerable media attention by interfacing his system with an F-22 fighter jet simulator and "training" the cells to manipulate the plane. At the University of Washington in Seattle, Tom Daniel's team implants small computer chips in the hawk moth, Manduca sexta, to study how the insect's nervous system controls flight. Computer scientists have created a hat that can read your thoughts. It allows you to stroll down a virtual street. All you have to do is think about walking. Called a brain-computer interface, the device detects activity in certain brain areas linked to movement, and uses the signals to mimic that movement in a virtual world. The technology could one day help paralysed patients to move robotic arms, or help sufferers of motor neuron disease to type out words on a virtual keyboard. By picking up on these bursts of nerve activity, the computer can decide whether you are thinking about moving your hands or feet, and react accordingly. The technology detects brain waves by using electrodes placed at strategic points on the scalp; they are positioned over brain areas known to be involved in moving specific body parts. The computer can then distinguish between signals corresponding to different types of movement. Previously, accurate detection of local brain activity has required electrodes to be implanted in the brain. This technique has allowed recipients to control robots and even send e-mails. The new device, presented at the International Society for Presence Research 2005 technology meeting in Londonref, achieves a similar feat using non-invasive methods. The team tested their creation by asking participants to navigate a virtual-reality studio called the Virtual Cave. Test subjects sit in a square studio wearing 3D goggles, which project a scene such as a street, complete with pedestrians and buildings. The computer then chooses a task for the participant: either walking forwards or moving their hands. It tells the user what to do through sound cues. If the person is asked to think about walking, and they do so in a way that can be picked up by the cap, the virtual character steps forwards. If they fail, the character stays still. When asked to think about moving their hands, successful volunteers are rewarded by staying still. Failure leads to punishment: their character takes a step backwards. One of the world's best-trained users is Doris Zimmerman, a student who has worked extensively with the team in Austria. She was flown in to help with demonstrations at the conference. As she sits in front of the Virtual Cave's 2D screen, we watch her effortlessly glide down a high street. However, it's not as easy as it looks and 5 hours can be required to learn it. It could even enable stroke patients to regain movement by allowing them to 'exercise' their brain's motor centres. And strengthening the mind might lead to better motor controlref. Web resources : UCL computer science Guger Technologies Technology news
peripheral-machine interface (PMI)	amputees and persons with intact CNS but weak muscles	motors, i.e., prosthetic handref
hybrid brain-machine interface (HBMI)	spinal cord-injured persons with intact limb muscles	muscles, i.e., direct brain control of FreeHandref
computer-brain interface (CBI)	Parkinson's disease	musclesref

Neuroscientists have created a computer game based on table tennis that people can play using nothing more than the power of their minds. It is hoped that the technology will one day train people to generate neural signals that could control a wheelchair or communication device. Each 'brain pong' player lies in a functional magnetic resonance imaging (fMRI) machine, usually reserved for medical brain scans. After a short period of training, the players are able to make their ping-pong bat move up and down the screen by concentrating on specific thoughts. The technique can measure specific areas of the brain with millimetric accuracy, and most subjects proved adept at switching thoughts 'on' and 'off' after just three 45-minute training sessions. Sophisticated data analysis software makes the system responsive enough for two players to compete in real time. The specific method of training varies from subject to subject : one has to find the area of the brain that the subject can control best
Web resources :
• Brain-machine interfaces Web Focus by Nature
• Rainer Goebel
• The Story of Pong
• Play 'Radial Pong'
• Human Subjects play Mind Games
• blood substitutes / artificial blood : although transfusion of donor blood is a routine and sage procedure, there are several reasons for developing blood substitutes :
• a greater need for blood can be predicted as growing numbers of elderly patients needing blood-intensive surgery such as hip repairs or replacements strain the existing donor supply
• human RBCs have exacting storage requirements designed to prolong clinical effectiveness and to reduce the risk of bacterial contamination. This substantially limits the availability of blood at disaster sites or on the battlefield.
• platelet substitutes :
• fibrinogen adsorbed on heat stabilised albumin microcapsules of defined size (Synthocytes™) (source : Quadrant Healthcare plc) for the prevention and treatment of thrombocytopenia^{ref1, ref2, ref3}
• infusible platelet membranes (IPM) (Cyplex^®) (source : Cypress Bioscience) prepared from outdated human platelets^ref
• artificial plasma
•  (solurce : Genzyme)
• replacement of erythrocytes : donor blood shortages make blood substitutes attractive for the short-term replacement during surgery. An ideal blood subsitute should :
• be able to transport oxygen
• require no cross-matching or compatibility testing
• be suitable for long-term storage (preferably at room temperature)
• be able to survive in the circulation for several weeks (the intravascular "dwell" time) before being cleared by the kidney : usually half-life is dose-dependent
• be free of side effects (usually are reversible)
• be more amenable to sterilization to remove infectious pathoges
• not only transport but also effectively deliver oxygen to tissues
Types of blood substitutes :
• hemoglobin-based oxygen carriers (HBOCs) : a cell-free solution of hemoglobin (so that compatibility testing is not required) can be used as a blood substitute because hemoglobin maintains its ability to transport oxygen outside of the RBC. It can be sterilized by ultrafiltration and low heat to inactivate infectious agents, a strategy that is not possible with RBCs.
• human hemoglobin : the source could be outdated donor blood (RBCs that have exceeded the approved storage period), but the limited supply of human blodd has compounded difficulties in developing a human-derived HBOCs. Genetically engineered bacteria (e.g. Escherichia coli) to produce a recombinant source of human hemoglobin would also eliminate the concerns about diseases transmission.
• Hemolink^® (source : Hemosol) : used for heart surgery, half-life 18-20 hours, shelf-life 1 year refrigerated, side effects : vasoconstriction, jaundice
• rHb 1.1 (Optro^®; source :  Somatogen and Eli Lilly) is a genetically engineered variant of hemoglobin Presbyterian, with modifications to decrease its oxygen affinity, produced by Escherichia coli. Intramolecular cross-linking of the tetramer made rHb 1.1 similar in structure to DCLHb
• pyridoxalated hemoglobin polyoxyethylene (PHP)
• source : Curacyte
• polymerized stroma-free hemoglobin (poly-SFH-P) (PolyHeme^®; source : Northfield Laboratories) : used for trauma and urgent blood loss, half-life 24 hours, shelf-life 1 year refrigerated, no side effect reported
• HemoZyme^® (source : Synzyme)
• bovine hemoglobin (Bos taurus) : more readily available (and cheaper) source. It does not have 2,3-DPG, resulting in a P₅₀ of about 30 mmHg both inside and outside the RBC (similar to that of the intracellular native human hemoglobin). In addition, once stripped of all other bovine proteins, bovine hemoglobin is not recognized by the human immune system as foreign. One potential obstacle is the fear that it could harbor the bovine BSE PrP^Sc.
• HbOC-201 (Hemopure^®; source : Biopure), used for elective surgery, half-life 24-36 hours, shelf-life 2 years at RT, side effects : vasoconstriction, transient increase in liver and pancreas enzymes
• worm hemoglobin : hemoglobin from Arenicola marina (a.k.a. lugworm) is 50 times larger thann human hemoglobin and cannot damage kidneys
Problems with these HBOCs included :
• survival times in the circulation < 6 hours (in solution the hemoglobin tetramer readily dissociates into monomers and dimers that are quickly eliminated from the blood stream by glomerular filtration and uptake by the reticuloendothelial system in the kidneys). Integrity of the hemoglobin tetramer can be achieved by chemically modifying hemoglobin :
• specific covalent chemical cross-bonds between polypeptide chains prolong intravascular dwell time of up to 12 hours
• treatment with 3,5-dibromosalicyl fumarate => diaspirin cross-linked haemoglobin (DCLHb) (Hemassist^®; source : Baxter Healthcare) has been used in cardiac and orthopaedic surgery, and for resuscitation of traffic accident victims. It is a DCLHb tetramer made from outdated human blood : it has the advantages of a shelf life of approximately 9 months frozen and 24 hours refrigerated. Its intravascular half-life is limited to 2-12 hours and is dose-dependent.
• treatment with bifunctional cross-linking agents, such as o-raffinose or glutaraldehyde, that target specific amino groups and polymerize the hemoglobin molecule. These reagents produce polyhemoglobin, composed of 4 to 5 hemoglobin molecules, which varies in molecular size and configuration and has intravascular dwell time of up to 24 hours
• conjugation to a variety of larger molecules such as dextran, polyethylen glycol (PEG) (e.g. Hemospan^®; source : Sangart, Inc.), or polyoxyethylene, retarding the rate at which it is cleared from the circulation. In this instance, the intravascular dwell time can be extended to 48 hours
• liposome-encapsulated hemoglobin (LEH) has been found to be an effective oxygen carrier, without the adverse effects of vasoconstriction. The liposome encapsulation appears to increase plasma retention time; however, adverse immune interactions occur with the liposome. This product has the advantage of being freeze-dried and stored at room temperature.
• abnormally high oxygen affinity (without 2,3 bisphosphoglycerate (2,3 BPG) stroma-free hemoglobin has a P-50 of 12-14 mm Hg instead of 26.5 mm Hg, not allowing for the adequate release of oxygen to the tissues. Aeveral to decrease the avidity of hemoglobin to oxygen include the addition of organic phosphate to serve the function of 2,3 BPG and adenosine triphosphate
• hemoglobin (like other plasma proteins) has a colloidal osmotic effect and hence cell-free hemoglobin acts as a plasma expander because the colloidal osmotic pressure it exerts alters the intravascular volume by more than the volume transfused (on the other hand RBCs exert little if any colloidal osmotic pressure) => clinical side effects such as malaise, abdominal pain, hemoglobinuria, and renal toxicity
• some hemoglobin solutions also have vasopressor effects, that is, they increase blood pressure and decrease cardiac output. Several studies have indicated that this vasopressor effect is partly attributable to the ability of cell-free hemoglobin to penetrate the lining of the blood vessels and to scavenge NO^., a cellular chemical messenger which stimulates blood vessel relaxation. Although this vasopressor effect is undesirable for most applications (it can limit circulation to oxygen-starved tissue), it may turn out to be clinically advantageous for treating a small subpopulation of patients with septic shock who suffer from an uncontrollable decrease in blood pressure. Baxter is working on rHuHGB that doesn't scavenge NO^..
New HBOCs under development include polyhemoglobin linked to antioxidant enzymes such as SOD and catalase in an effort to minimize ischemia-reperfusion injury (IRI). In addition, some groups are trying to package human hemoglobin in artificial RBCs made from biodegradable polymer nanocapsules or lipid vesicles.
• perfluorocarbon (PFC) emulsions (PFCE) ("blue blood") are emulsions of synthetic fluorinated hydrocarbons : for religious reasons, certain groups of people (e.g. Jehovah's Witnesses) are unable to accept transfusions of either donor blood or human and animal proteins such as hemoglobin. For these patients, fluorocarbons - molecules that are structurally similar to hydrocarbons except that hydrogen atoms are replaced with fluorine atoms - may be their only option if a blood transfusion is required. Perfluorocarbon liquids have an excellent capacity for carrying O₂ and CO₂ without actually binding to these gases : O₂ readily dissolves in the chemically inert perfluorocarbon liquid and can be easily extracted by O₂-deprived tissues (also used to overcome tumor hypoxia during cancer radiotherapy and cancer chemotherapy). Perfluorocarbons are not miscible with aqueous solutions and must be prepared as emulsions (containing particles with a diameter of approximately 0.2 mm, capable of entering the microcirculation) before they can be used as blood substitutes. The O₂-loading capacity of perfluorocarbons is linearly related to P_O2. in equilibrium with the emulsion. Thus, at a given P_O2, hemoglobin binds significantly more O₂ than can be dissolved in the perfluorocarbons. One perfluorocarbon, Fluosol DA, has been approved by the FDA as a blood substitute for heart surgery: the fully oxygenated perfluorocarbon is infused at the time of percutaneous transluminal coronary angioplasty (PTCA) / percutaneous coronary intervention (PCI) / percutaenous coronary revascularization to provide oxygen delivery during the surgical procedure. Even though the perfluorocarbon reduced cardiac trauma and pain during surgery, its use has never been fully embraced by physicians. Through partial liquid ventilation of the lungs, perfluorocarbons have been successfully used to provide O₂ to premature infants with severe respiratory distress syndrome
• mixture of perfluorodecalin (the carbohydrate decalin (C₁₀H₁₈) in which fluorine atoms replace hydrogen atoms) and perfluorotripropylamine emulsified with Pluronic F-68 (Fluosol-DA^® ; source : Green Cross Corporation of Japan), 20% infusion. Ultimately, this product was ineffective at significantly improving oxygen delivery in the setting of acute hemorrhage. Other shortfalls of Fluosol-DA included a short effective intravascular half-life, temperature instability, low oxygen-carrying capacity, and a poor shelf life. Its major adverse effects included acute complement activation, uptake by the reticuloendothelial system, and disruption of normal pulmonary surfactant. Many of these effects were determined to be caused by Pluronic F-68. Fluosol-DA was approved in the United States for use only during coronary angioplasty; however, because of improvements in angioplasty techniques and the shortfalls of Fluosol-DA, the Green Cross Corporation ceased manufacturing this product in 1994.
• mixture of perfluorooctyl bromide (PFOB) / perflubron and perfluorodecyl bromide emulsified with egg yolk lecithins (Oxygent^® ; source : Alliance Pharmaceutical Corp.) is in stage II/III clinical trials in the USA for general surgery, intravascular half-life 24-48 hours (dose-dependent), shelf-life 2 years refrigerated at 2-8° C, side effects : slight fever, thrombocytopenia. This compound is more stable at room temperature than Fluosol-DA. The oxygen-carrying capacity of Oxygent is 4-5 times greater than that of Fluosol-DA, making this compound more clinically useful in the setting of acute hemorrhage. However, a high fraction of inspired oxygen is required for the compound to be effective
• Neo-PFC
• PHER-O2 (source : Sanguine Corporation) is a second-generation formulation of Fluosol^® designed to overcome some of the deficiencies of the original product; i.e., need for frozen storage, low (20%) PFC content and short intra-vascular residency time
Web resources :
• American Association of Blood Banks (AABB)
• Bloodless Medicine Research (BMR) by  Lelio Mario Sarteschi M.D.
• Bloodless Medicine & Surgery Institute (BMSI)
• NoBlood.com - The Bloodless Medicine and Surgery Network
• Hemo Concepts
• Physicians and Nurses for Blood Conservation (PNBC)
• Artificial Cells & Organs Research Centre at McGill University
• International Society for Artificial Cells, Blood Substitutes &  Biotechnology (ISABI)
• Other companies :
• Perftoran
• Synthetic Blood International, Inc.
• Sangui BioTech International, Inc.
• Journals :
• Artificial Cells, Blood Substitutes & Biotechnology
• artificial saliva / saliva substitute : a solution in distilled water, used to lubricate the oral tissues in xerostomia
• ipromellose
• polyvinylic alcohol
• Optimoist
• Saliment
• hydroxy-propyl-methyl-cellulose (HPMC) containing artificial saliva
• carboxy-methyl-cellulose (CMC) based saliva substitute
• mucin-containing saliva substitute
• xanthan gum-based saliva substitute
• artificial tears / aqueous tear substitute / viscosity agents to treat xerophthalmia
• lubricating ointments / lubricants
• white petrolatum
• mineral oil
• liquid or alcohol lanolin
• cellulose polymers
• carboxymethylcellulose
• hydroxyethyl cellulose
• hydroxypropyl cellulose
• hydroxypropyl methylcellulose
• methylcellulose
• polyvinyl alcohol
• polyethylene glycol
• mineral oil
• glycerin
• dextran
• sodium hyaluronate 0.1%
• ocular prosthesis : an artificial eyeball
• Snellen's reform eye : an artificial eye composed of 2 concavoconvex plates with an empty space between.
• odontokeratoplasty
• intraocular lenses (IOL)
• silicone IOL
• SI-40NB
• sharp-edged
• ClariFlex(R), AMO
• sharp-edged hydrophilic acrylic IOL
• Sensar(R) 40e, AMO
• SC60B-0UV
• AcrySof(R), Alcon Laboratories
Side effect : posterior capsule opacification (PCO).
• vitreous substitutes
• nonexpansile gases
• air
• Ar
• CO₂
• He
• Kr
• N₂
• O₂
• Xe
• expansile gases
• sulfur hexafluoride (SF₆)
• octafluorocyclobutane (C₄F₈)
• perfluoromethane (CF₄)
• perfluoroethane (C₂F₆)
• perfluoropropane (C₃F₈)
• perfluoro-n-butane (C₄F₁₀)
• perfluoropentane (C₅F₁₂)
• silicone oil / polydimethylsiloxane : a long-chain silicone polymer with a viscosity of 5000 to 5400 centistokes, injected into the vitreous to maintain retinal tamponade in the management of complicated retinal detachment
• nonfluorinated silicone oils [ (CH₃)₃SiO[(CH₃)₂SiO]_nSi(CH₃)₃ ]
• fluorosilicone [ (CH₃)₃SiO[C₃H₄F₃)(CH₃)SiO]_nSi(CH₃)₃ ]
• "high-tech" silicone oils [ (CH₃)₃SiO[(C₆H₅)(CH₃)SiO]_nSi(CH₃)₃ ]
• retinal implants : a bionic eye that allows blind people to see has now got a protective coat of diamond that should significantly improve its performance. The silicon chip retinal implant is being developed by Second Sight, a company based in Sylmar, California, along with a consortium of university researchers. The device needs a hermetic case to prevent it from reacting with fluids in the eye. The approach until now has been to lock it in a big waterproof can, but it's very big and bulky, so researchers have developed an ultrananocrystalline diamond (UNCD) film that is guaranteed to be safe, long-lasting, electrically insulating and extremely tough. The coating can also be applied at low temperatures that do not melt the chip's microscopic circuits. The UNCD film is the first coating to meet all the necessary criteria for the implant. The tiny diamond grains that make up the film are about 5 nm across. They grow from a mixture of methane, argon and hydrogen passing over the surface of the 5-mm² chip at about 400°C. Xiao and his colleagues have already tested the implants in rabbits' eyes, and saw no adverse reaction after 6 months : the results were presented on 1 April 2005 at the Materials Research Society meeting in San Francisco, California. A healthy retina contains rod and cone cells that convert light into electrical impulses, which fly to the brain through the optic nerve. But for millions of people with diseases such as retinitis pigmentosa or macular degeneration, these cells do not work properly. The retinal implant bypasses these diseased cells by electrically stimulating healthy cells that sit beneath them at the back of the eyeball. The patient 'sees' using a pair of glasses carrying a tiny video camera that sends digitized images to the implant using radio waves. The first human trial of Second Sight's artificial retina has been running since 2002, and it has enabled a formerly blind patient to distinguish between objects such as cups and plates, and even to make out large letters. But with only 16 electrodes, the device does not allow the patient to see a clear picture. For that, thousands of electrodes are needed on the same size of chip, making it even more delicate. In the first trial, the electronics package was separated from the electrodes and implanted behind the patient's ear because its casing was so bulky, explains Greenberg. The team's goal is to integrate everything into a single unit that fits neatly into the eye. A UNCD coating could be equally useful for other implantable devices, such as biosensors to monitor a patient's health. They've managed to create a diamond film that is of considerably higher quality than other methods have managed, but the ultimate potential of these devices will only be told in long-term trials. Second Sight is planning clinical trials of a 60-electrode device this year, but it may be several more years before the diamond-coated chip is used in humans
• heart prostheses
• pace-makers
• ventricular reshaping^ref : lessons learned from ventricular remodeling procedures have shown that Laplace mechanics, whereby mural tension is related to ventricular diameter, may be beneficially applied to the left ventricle (LV). Recently developed devices, predicated on the Law of Laplace, are directed at interrupting the heart-failure cascade without removing functioning myocardium.
• partial left ventriculectomy / left ventricular chamber reduction surgery
• endovascular patchplasty
• device-based external ventricular reshaping / passive containment surgery ^ref
• dynamic cardiomyoplasty (CDM) using the latissimus dorsi muscle implanted all around the heart and stimulated by electrodes placed among muscle bundles, becoming spontaneous within 6 weeks via a transformation of type II (fatigueable) muscle fibers into type I (not fatigueable) muscle fibers. Dynamic cardiomyoplasty has been performed in over 1000 patients worldwide but due to limited success the procedure was never been adopted as an alternative approach for the surgical therapy of heart failure. However, observations in these patients showed that the nonstimulated or fibrotic transformed latissimus dorsi by itself led to an improvement of heart failure symptoms. These findings stimulated animal experiments with so-called passive cardiomyoplasty devices. In several animal models, the progression of heart failure could be stopped, and even reversed remodeling could be demonstrated. Several different devices have been developed and tested in animal models. Diabetes-induced alterations in latissimus dorsi muscle properties impair effectiveness of dynamic cardiomyoplasty in rats^ref.
• Myosplint^® (source : Myocor, Inc., Maple Grove, MN) : passive implantable device that uses transmural tension members to bisect the dilated LV to create a bilobed LV with decreased mural tension in each cavity^ref.


• CorCap^® cardiac support device (CSD) (source : Acorn Cardiovascular Inc., St. Paul, Minn) is a polyester mesh-like device that is surgically implanted around the heart. The device is designed to provide circumferential myocardial wall end-diastolic support,  reduce ventricular wall stress and myocyte stretch, improve function, and reverse cardiac remodeling. The CorCap device has been implanted worldwide in > 130 patients with dilated cardiomyopathy (idiopathic or ischemic), with or without concomitant cardiac surgery^{ref1, ref2}. It has already passed a successful clinical feasibility study and already has EMEA approval. However, the final evaluation of two multicenter trials has to be awaited to assess the future role of this device in the treatment of heart failure^ref.


• mechanical circulatory support (MCS) in patients presenting with cardiogenic shock waiting for heart transplantation. In general, 2 categories of blood pumps can be distinguished:
• displacement pumps move a certain amount of blood by the movement of a flexible membrane. This movement requires an extensive installation: artificial ventricles, tubing, motors, pneumatic systems and driving consoles
• pneumatically driven
• electrically driven
• rotary blood pumps are well suited for miniaturisation and endovascular implantation. They can be either
• axial : Hemopump, MicroMed DeBakey VAD, Jarvik 2000, HeartMate II, Streamliner, Impella, Berlin INCOR I, Valvo pump, and IVAP^ref
• diagonal
• radial
• ventricular assist device (VAD) / auxiliary ventricle / booster heart : an implanted pumping mechanism designed to assist a ventricle of the heart in maintaining normal output, rate, and blood pressure. While "resting" during VAD support, damaged cardiac muscle starts to repair itself and to contract with more vigor. The benefit of a VAD versus a TAH is that the natural heart remains in place-if the VAD is stopped or removed, the patient's heart can still support the circulation. Clinical experience has demonstrated that once a damaged heart is "rested" through VAD support and eventally treated for several months with b₁-agonists, the heart can repair itself and its function improves. In a few cases, the VAD was removed and the patient no longer needed a transplant. The goal is for wider application of VAD support so that heart failure patients are supported for weeks or months until cardiac damage has been repaired sufficiently such that a transplant is not necessary. Design of the VAD began as early as 1963 but did not progress rapidly until the 1990s. Since then, thousands of patients have been successfully supported, sometimes for years
• right ventricular assist device (RVAD) for right ventricular failure (RVF)
• left ventricular assist device (LVAD) or system (LVAS) supports the left ventricle, the locus of the most critical heart damage in left ventricular failure. Blood flows from the left atrium through the left ventricle and the pump inflow valve into the LVAD, which then pumps the blood into the ascending aorta. With a LVAD, the right ventricle benefits from reduction of left atrial and pulmonary arterial pressures and the heart can start to repair itself
• biventricular assist device (BiVAD / BVAD) or system (BVS) for congestive heart failure (CHF)
• extracorporeal or paracorporeal VAS are available for RVAS, LVAS, or BVAS
• Toyobo VAS^®
• Zeon VAS^®
• Medos-HIA system^®
• Berlin Heart^®
• Medtronic Bio-Medicus Pump^®
• Pierce-Donachy pediatric system^®
• percutaneous applicable cardio-pulmonary support (pCPS)^®
• Thoratec Ventricular Assist Device (VAD) System (VAS) (Thoratec Laboratories, Pleasanton, CA) is a paracorporeal pump that can provide univentricular or biventricular assistance for patients with heart failure. The system consists of a prosthetic ventricle that has a blood-pumping chamber of Thoralon / BPS-215M^® (Thoratec Laboratories) polyurethane ...


..., cannulas for univentricular or biventricular support ...

..., and either a hospital-based pneumatic drive console ...

... or a portable battery-powered drive unit^ref.

For biventricular assistance, 2 pumps are used.

The Thoratec voluntary registry indicates that, as of May 2000, this system had been implanted in 1,376 patients, mainly for bridging to transplantation (828 patients) or postcardiotomy support (195 patients); the remaining 353 patients received a hybrid configuration of the device or had incomplete information, so they are not included in this analysis. In the 828 bridge-to-transplant patients, the Thoratec system provided biventricular assistance in 472 cases, left ventricular assistance in 326 cases, and right ventricular assistance in 30 cases for up to 515 days. During the support period, the cardiac index increased significantly from 1.4 +/- 0.8 L/min/m² to 3.0 +/- 0.5 L/min/m² (with biventricular assistance and left ventricular cannulation). 60 percent of the 828 patients underwent transplantation, and the posttransplant survival rate was 86%. In the 195 patients who needed postcardiotomy support, VADs were used for up to 80 days for cardiac recovery. Thirty-eight percent of the patients were weaned from the VAD, and 59% of the weaned group were discharged from the hospital. In addition, 49 postcardiotomy patients were considered for transplantation; of these, 32 received a transplant and 23 were discharged. Patient mobility is being improved by the use of a portable driver. The Thoratec VAD is suitable for a wide range of applications, and efforts are underway to facilitate patient mobility and allow hospital discharge. An intracorporeal version of the VAD, which is currently under development, will help achieve these goals^ref.
• implantable or intracorporeal ventricular assist device (IVAD)
• Novacor LVAS^® (source : Baxter) : pulsatile, long-term VAD commonly used as a bridge to transplantation


• HeartMate™ LVAS (Thermo Cardiosystems Inc., Woburn, MA) (the most commonly used VAD) : starting initially with a pneumatic actuated pusher plate pump, clinical use of the HeartMate I began in 1986. In 1990, electric motor-actuated versions of the HeartMate I began to be used clinically. Presently, the HeartMate I has been implanted in some 2,300 patients worldwide, and this LVAD is a standard by which all others are currently measured.
• Thoratec HeartMate^® Implantable Pneumatic Left Ventricular Assist System (IP LVAS) [1000 IP]
• HeartMate LVAS IP^®
• TCI HeartMate vented-electric system (VE)^®



• Novacor N100^® wearable left ventricular assist system)
It is a rotating torque motor and cam system that displaces 75 cc of blood from the ventricle into the aorta. The control system can increase or decrease the pump rate to match the demand. The surfaces that make contact with blood are textured so that an adherent layer of blood cells settles on the surface, enabling blood to flow smoothly. Blood is directed through the pump by inflow and outflow valves made of pig tissue. In the Cleveland Clinic experience, the incidence of emboli or stroke is about 2%, using only aspirin as an anticoagulant.
• a new type of IVAD in which rotary dynamic pumps produce a continuous nonpulsatile blood flow and no artificial valves are required is currently in clinical trials. The entire mechanism sits in the ventricle immersed in blood, an environment hostile to bacteria. In some situations, patients with this new VAD have a normal cardiac output of 5 to 6 liters/min, but no palpable pulse of measurable blood pressure! There was concern that this would cause physiological abnormalities, but experience with animal models and early clinical results suggest that the difficulties may have been overestimated. These continuous-flow devices are much smaller and simpler than the pulsatile pumps and thus fits the majority of adults and children. They may also be less expensive. Most current designs use highly polished titanium surfaces that appear to be compatible with blood, at least for the short resistance times during which blood is in contact with the pump. As a partial assist VAD that elevates the toal cardiac output of a weak heart elevates the total cardiac output of a weak heart closer to the normal range, these devices show tremendous potential. Prototypes of the continuous flow VAD include
• Hemopump^® (Medtronic, Inc, Minneapolis, MN) was conceived in 1975 and designed in 1982 as a temporary, extracorporeal cardiac assist system. Although it has been used clinically in Europe, it is not currently available in the United States. In vitro and in vivo testing of the Hemopump began in 1983. Clinical investigations have included studies of patients in cardiogenic shock, Hemopump-supported coronary artery bypass operations in Sweden, and European studies of percutaneous transluminal coronary angioplasty (PTCA) with Hemopump support. The Hemopump has demonstrated positive hemodynamic effects in patients. Laboratory and clinical studies have shown that the nonpulsatile axial flow generates flows of up to 4.5 L/min while maintaining adequate perfusion of other organs. In Europe, hemopumps have been used successfully to support coronary bypass and PTCA. The Hemopump system is simple, inexpensive, and well tolerated by the blood elements. Moreover, its design allows flexibility in supporting patients during cardiopulmonary bypass (in lieu of conventional techniques) and high risk angioplasty, as well as in rescuing patients with low cardiac output^ref
• Jarvik 2000 FlowMaker LVAD^® (Jarvik Heart Inc., New York, NY) is a larger, longer-term, electrically powered, axial-flow blood pump that provides continuous flow from the left ventricle to the descending thoracic aorta. Because of its small size, this device can be placed within the left ventricle, and no inlet cannula is needed. The continuous blood flow eliminates the need for valves, an internal compliance chamber, or an externalized vent. The system consists of a blood pump, 16-mm outflow graft, percutaneous power cable, pump-speed controller, and direct-current power supply. The blood pump is simple in design, its single movable component being an impeller located in the center of the titanium housing. A brushless direct-current motor contained within the housing creates the electromagnetic force necessary to rotate the impeller. The impeller consists of a neodymium-iron-boron magnet suspended by 2 ceramic bearings. On the impeller’s outer surface are 2 titanium hydrodynamic blades. All the blood-contacting surfaces within the pump are made of smooth titanium. Blood flow is directed through the outflow graft by stator blades located near the pump outlet. The controller and battery are worn on a belt or are carried in a bag. The vaned impeller rotates in response to the electromagnetic force created by the motor, which is contained within the pump housing. The impeller is suspended by ceramic bearings (not shown). Outflow stator blades (not shown) direct blood flow through the pump’s outflow conduit.



The motor receives power from the external controller through a series of wires that are contained in a silicone tube covered with Dacron. The motor speed is manually adjusted and is regulated by a pulse-width-modulated, speed-control circuit. Lead acid or lithium ion batteries provide the system’s 12-volt power supply. The pump’s operating range is 8000 to 12 000 rpm; under optimal physiological conditions, these speeds can generate flows of up to 6 L/min. The normal power consumption is 3 to 7 watts. The pump’s speed is adjusted in 1000-rpm increments until minimal flow is ejected through the aortic valve. In case of a system malfunction, the controller provides both an aural and visual alert. Ambulatory patients can maintain the system’s external equipment and change the batteries as needed. The pump is implanted through a thoracotomy incision with the aid of partial cardiopulmonary bypass (CPB). The femoral artery and vein are exposed for CPB cannulation. A left thoracotomy is then made in the fifth or sixth intercostal space, and the heart and descending thoracic aorta are exposed. The aorta is partially occluded while the outflow graft is anastomosed to the aorta. The femoral vessels are then cannulated and CPB is begun. Ventricular fibrillation is induced, and a circular knife is used to create an opening in the left ventricle. A silicone/polyester sewing cuff is sewn to the apex of the heart with pledgeted sutures. The power cable is tunneled subcutaneously across the abdomen from left to right just below the diaphragm and exits the skin at the right anterior portion of the abdomen. The pump is inserted into the left ventricle and is secured with cotton tape around the cuff and pump housing. The pump’s outflow graft is anastomosed to the graft previously placed on the aorta. The heart is then defibrillated, and air is removed from the left ventricle, pump, and graft. As CPB is gradually discontinued, the pump is turned on. Pulmonary vasodilators and inotropic agents are used as necessary to support right-sided cardiac function. Intraoperatively, hemodynamic assessment is performed with transesophageal echocardiography, a pulmonary artery catheter, an arterial pressure line, a pressure line to the left ventricle, and an ultrasonic flow probe on the pump’s outflow graft. Hemodynamic values are recorded at each pump speed, and the optimal speed is determined. Flow through the pump can be calculated accurately against varying mean pressures^ref. The ultrasonic flow probe is removed before the thoracotomy incision is closed. In the early postoperative period, hemodynamic function is assessed with an arterial pressure line, a pulmonary artery catheter, and transthoracic echocardiography. Once all the monitoring lines have been removed, hemodynamic assessment is based on echocardiography and routine evaluation of vital signs. Postoperatively, the goal is to maintain pump flow at the lowest level compatible with a normal cardiac index and NYHA class I status. Hemodynamic and echocardiographic studies are performed routinely to determine the proper pump-speed setting. During these tests, the pump speed is increased in 1000-rpm increments over the usual operating range of 8000 to 12 000 rpm, after which the pump is turned off for up to 5 minutes. Serial echocardiographic studies are performed at 24 hours, 1 week, 1 month, and every 3 months thereafter until support is terminated^ref.
• LionHeart LVD 2000^® (Arrow International, Reading, PA) left ventricular assist device is the first fully implantable system designed for destination therapy (ie. no heart transplantation required)^{ref1, ref2}. It was first implanted on October 26 1999 in Germany. In 2001 the first LionHeart was implanted in Italy at Pavia Hospital. As for AbioCor TAH, there are no cables connected trough the skin and the internal battery is continuously recharged via a battery waist which conducts energy to an internal serpentine


• Kantrowitz CardioVAD (KCV)^® (LVAD Technology, Detroit, MI) is an implantable IABP designed to provide permanent mechanical assistance that can be activated intermittently, require no anticoagulation, allow for treatment of CHF and potential myocardial recovery, and not pose an increased risk of cerebrovascular accidents (CVA). It is an electrically powered, pneumatically driven circulatory assist device which provides diastolic augmentation and systolic unloading to the failing heart. It consists of a 60cc-pumping chamber (a), a percutaneous access device (PAD) (b), and an external controller (c).


The pumping chamber, is surgically implanted in the descending thoracic aorta with the patient on cardiopulmonary bypass. Its physiologic function is analogous to that of the intra-aortic balloon pump (IABP)^ref. A 1x5 cm full thickness segment of skin was obtained from the forearm, under local anesthesia at the bedside. Autologous fibroblasts were cultured from the biopsy specimen for 10 to 14 days and used to coat the percutaneous access device (PAD). Once this procedure was completed, patients were admitted for KCV implantation. Under general anesthesia, a left thoracotomy was performed through the 5th interspace. The aorta was encircled 1, 8, and 16 cm distal to the subclavian (points A, B, C, respectively). Patients were completely heparinized and placed on warm, heart beating cardiopulmonary bypass (CPB) using femoral arterial and venous cannulae. Another arterial cannula was placed at the aortic arch for cerebral perfusion. The aorta was crossclamped at points A and B and incised longitudinally. Removed from the edges were 3-mm strips of the aorta. The superior portion of the KCV was sutured in place with a single layer of 2 to 0 polypropylene running mattress suture. A felt strip was used on the aortic side as a buttress. The crossclamp at B was released and reapplied at point C. Intercostals were controlled with Fogarty catheters. The distal portion of the KCV was sewn to the aorta in a similar technique to the proximal end. The point C clamp was removed and a second layer of sutures used to further secure the KCV. After obtaining hemostasis, 3 epicardial screw-in electrodes (St. Jude Pacesetter, St. Paul, MN) were placed in a triangular pattern 1 cm apart on to the lateral portion of the heart. The device (bladder integrity and electrode sensing) was tested as mechanical support was initiated. This assisted in weaning from CPB. The thoracotomy incision was closed and the patient placed in the supine position. The PAD was then inserted into a subcutaneous pocket in the right upper quadrant and the device connected. Patients were extubated, weaned from inotropes and ambulated as soon as possible. All patients were run with maximum diastolic augmentation with the device set to inflate immediately after the dicrotic notch. Anti-coagulation was not used for the device—1 patient received coumadin for a mechanical prosthetic aortic valve^ref.
• DeBakey VAD^® system (MicroMed Technology, Inc.; Houston) is a miniaturized, electromagnetically actuated, fully implantable titanium axial flow pump weighing 93 g^ref and consisting of 4 parts: a pump, a controller, a clinical data acquisition system (a laptop computer), and a patient home support system (wall power and battery charger). The pump consists of a titanium inflow cannula and apical ring, the housing unit containing the impeller (only moving part) and motor, a Vascutec gelweave outflow vascular graft that exits the pump as an outflow conduit and connects to the aorta, and a percutaneous cable. An integral flow probe placed around the outflow graft measures real-time cardiac output through the graft. The cable from the flow probe and the impeller are bundled into a single cable assembly, which exits the skin of the anterior abdominal wall at a point superior to the right iliac crest. The cable is then connected to the controller. The pump is designed to achieve 5 L/min against 100 mm Hg pressure with a speed of 10 000 rpm. The flow curves, speed, current and power are displayed in a bedside monitor unit (the clinical data acquisition system (CDAS)) that is accessed to make adjustments to the pump speed. A pump motor cable along with the flow probe’s wire exit transcutaneously and connect to the external controller. The latter, consisting of the controller module, battery packs and battery charger, is designed to operate the pump. To facilitate mobility and eventual hospital discharge, the pump can be actuated by 2 12-volt DC batteries for 4 to 6 hours^ref. 2 batteries can be connected to the controller at the same time. The controller can also use the power supplied by the CDAS or the patient home support system. The CDAS is used by the clinician to start or stop the pump, modify the pump speed, and set alarm limits. It displays and also stores the data and serves as a primary power source for the pump. The patient home support system is a small unit that charges 4 batteries at a time and also provides wall power to the controller when the patient is at home. European trials of the DeBakey VAD began in November 1998, and CE Mark certification of the device as a bridge to transplantation was awarded in May 2001. In the United States, clinical trials began in June 2000, and a 20-site Phase III Pivotal clinical trial is expected to conclude early in 2004. As of 4 March 2003, 179 patients at 14 heart centers in 7 countries had been implanted with the device. The longest surviving patient had lived for over 443 days with the VAD in place. Surgical procedure :
• in general, the heart is exposed via standard median sternotomy and full cardiopulmonary bypass support is established with atrio-aortic cannulation. In most instances, implantation proceeds without cardioplegic arrest. Transesophageal echocardiography (TEE) is routinely used to rule out the presence of aortic insufficiency, intracardiac thrombus and/or atrial septal defect (or patent foramen ovale) that may require further intervention. In addition, TEE guides intracardiac deairing, and provides superior imaging of inflow cannula placement, right ventricular function, and left ventricular decompression. Without the need to extend the median sternotomy incision, a small preperitoneal pocket is created to house the pump under the left costal margin. The left ventricular apex is cored out and an apical ring is sewn to the apex. The inflow cannula is inserted into the left ventricular cavity and secured to the apical ring. A small incision is made in the left lower quadrant and the thin driveline is routed transcutaneously and connected to the controller. Next, the outflow graft is tailored to reach the ascending aorta. A partial occluding clamp is placed along the greater curvature of the ascending aorta and the graft is anastomosed in end-to-side fashion using standard suture technique. In difficult reoperative cases, the device can be implanted to the descending aorta via a left thoracotomy. The graft is deaired and the patient is weaned from cardiopulmonary bypass as the device is actuated. Usually, inotropic support for the right heart is used. In some instances, intra-aortic balloon pump support and inhaled nitric oxide are employed^ref.
• a small, extrapericardial, subdiaphragmatic pocket was formed below the rectus muscle. The pericardium was opened. The ascending aorta was cannulated, and then bicaval cannulation from the right atrium was performed. Under cardiopulmonary bypass, while the heart was beating, its left ventricular apex was elevated, and an apical fixation ring was sewn to the apex. The inflow cannula was inserted through the fixation ring into the left ventricle in such a manner that it was angled toward the aortic valve but not the septum or the free wall of the myocardium. The pump cable was threaded out of the body through the anterior abdominal wall, at a point above the right iliac crest. The VAD was placed into the pocket, and the outflow graft was anastomosed to the ascending aorta. Using transesophageal echocardiography, we checked the location of the inflow cannula. After removing the remaining air in the system, we started the pump at a speed of 7,500 rpm. When pump flow was adequate, the patients were weaned from cardiopulmonary bypass^ref.
• HeartMate II (designed to run for 5 to 8 years) (Thermo Cardiosystems Inc., Woburn, MA) is a rotary-pump-based LVAD that uses an axial flow blood pump having blood immersed mechanical bearings. Clinical trials of the HeartMate II were initiated in 2000. It has evolved from 1991 when a partnership was struck between the McGowan Center of the University of Pittsburgh and Nimbus Company. Early iterations were conceptually based on axial-flow mini-pumps (Hemopump) and began with purge bearings. As the project developed, so did the understanding of new bearings, computational fluid design and flow visualization, and speed control algorithms. The acquisition of Nimbus by ThermoCardiosystems, Inc (TCI) sped developments of cannulas, controller, and power/monitor units. The system has been successfully tested in more than 40 calves since 1997 and the first human implant occurred in July 2000. Multicenter safety and feasibility trials are planned for Europe and soon thereafter a trial will be started in the United States to test 6-month survival in end-stage heart failure^ref.
• HeartMate III / HeartMate XVE™ (designed to run for up to 10 years) (Thermo Cardiosystems Inc., Woburn, MA) is structured around a centrifugal blood pump that uses a magnetically levitated rotating assembly. Compared to the HeartMate II, the HeartMate III has the potential for higher overall efficiency. The pump's operating life is not dependent on bearing wear^ref. The pump has been optimized by in vitro testing to achieve a design point of 7 L/min against 135 mm Hg at high hydrodynamic efficiency (30%) and to be capable of up to 10 L/min under such a load. Furthermore, the pump has demonstrated no mechanical failures, low hemolysis (4-10 mg/dl plasma free Hb), and low thrombogenicity during 6 (40, 27, 59, 42, 27, and 49-day) in vivo bovine studies^{ref1, ref2}. Electronics based on a rigid-flex satellite printed circuit board (PCB) arrangement that could be folded into a very compact, dense package were designed, fabricated, and tested. The pump's lower housing was redesigned to accommodate these PCBs without increasing any dimension of the pump except the height, and that by only 5 mm. The interconnect cable was reduced from 22 wires to 10 (two fully redundant sets of 5). An ongoing test of the assembled pump in vitro has demonstrated no problems in 5 months. In addition, a 20-day in vivo test showed only 1°C temperature rises, equivalent to pumps without incorporated electronics at similar operating conditions^ref.
• Thoratec IVAD^® is controlled with the Thoratec TLC-II Portable VAD Driver, which is a small briefcase sized, battery powered, pneumatic control unit. Intracorporeal LVADs and/or RVADs are implanted in a preperitoneal position, with a single small (9 mm OD) percutaneous pneumatic driveline for each VAD. The major advantages of the new IVAD design are size and simplicity. The IVAD weight (339 g) and implanted volume (252 ml) are substantially smaller than current implantable electromechanical LVAD systems. Only the small blood pump is implanted, leaving the more complex control unit external, where it can be serviced and replaced. The versatile design is intended for left and/or right heart support in large or small patients. The IVAD in combination with the TLC-II portable driver will be a viable and attractive alternative to large, implanted electromechanical systems^ref.
• Magnevad LVAD^® : a miniature axial flow LVAD whose first animal implant is scheduled for the fourth quarter of 2003. This is being performed by the George E. Reed Heart Center of Westchester Medical Center and by New York Medical College; both are in Valhalla, NY. This LVAD has new innovations not found in any other turbo pump to minimize thrombus, blood turbulence, flow separation, and the generation of microemboli. The Magnevad is only 25 mL in volume, similar in size to the Micromed & Jarvik 2000 axial flow turbo pumps that have contacting bearings. US Patent 6 527 699 was issued to Gold Medical on March 4, 2003 and World Wide PCT patents are pending. The improvements (patents pending) are designed to minimize flow separation and turbulence, the precursors of microemboli that lodge in end organs. This problem has been largely ignored in the published literature. A new long-term stable miniature ultrasonic position sensor is used for bearing control. It measures the axial position of the rotor to obtain LVAD differential pressure. Differential pressure is used to obtain pulsating flow and automatic physiologic control. The term "fourth generation pump" is being coined for the Magnevad because in addition to being noncontacting, it inherently measures pump differential pressure on which physiologic control can be based^ref.
• Intracardiac Pump LV^® for left ventricular support and the Intracardiac pump RV^® for right ventricular support (source : Impella CardioSystems AG, Aachen, Germany). The Impella pumps are rotary blood pumps of the axial flow type and produce 4.2 L/min at physiological pressure differences and a rotational speed of 32,500 rotations/min. These micropumps can widen the indications of beating heart surgery by sustaining hemodynamic stability and protecting the heart from warm ischemia. The current concept is aimed at bridging a procedure. Therefore, the proof of safe duration of usage has not been extended beyond 6 hours



• MYO-VAD^® (source : Biophan Technologies, Inc., a developer of next-generation medical technology, in conjunction with Myotech, LLC)
They are all now in clinical trials. Each of these devices has an axial flow impellers supported on blood-lubricated pivot points. So far, the devices have been implanted with a percutaneous wire that connects the device to an external controller and battery. Nothing prevents the use of a transcutaneous energy transmission system (TETS) system in future versions, although this would add hardware to the implant. The next generation of VAD blood pumps (still being tested in the laboratory) use 3D magnetic suspension of the pump rotor. These pumps address residual concens regarding friction, wear, and the crevice-like geometry of the pivot supports (which trap blood cells), and are predicted to remain reliable for > 10 years. This technology, the prototype of which is hopefully to be launched this year, brings us much closer to the day when small reliable pumps with few complications become as common as pacemakers and heart valves.
In USA LVAS are a destination therapy in patients with end-stage heart failure who, despite receiving optimal treatment, have a life expectancy of < 2 years and are not viable candidates for heart transplantation
Complications : significant infection rates have been reported among LVAD recipients. However, few reports have focused specifically on device infection, its treatment, and the impact of LVAD-related infection on clinical outcome. 46 LVAD-related infections were diagnosed in 38 (50%) of 76 patients who underwent LVAD implantation as a bridge to transplantation. 29 episodes of LVAD-related bloodstream infection (BSI) (including 5 that were cases of LVAD endocarditis) and 17 episodes of local LVAD infection were identified. Diabetes mellitus appeared to increase the risk of BSI among patients with LVAD infection. LVAD-related infection delayed transplantation, as reflected by longer device-support times (a mean duration +/- SEM of 182.8+/-31.1 days, compared with 66.3+/-8.8 days; P<or=.001). Continuous antimicrobial treatment before, during, and after transplantation was associated with fewer relapses than was a limited course of antibiotics (P<.001). A trend for longer hospital stays after receipt of a transplant and increased early mortality was observed in the cohort with LVAD-related infection, although long-term survival was similar to that associated with patients without LVAD-related infection. Posttransplantation invasive vancomycin-resistant Enterococcus faecium (VREF) infection was diagnosed in 6 patients with LVAD-related infection; 4 of these patients died. No VREF infections were identified in patients without LVAD-related infection. LVAD-related infection is common and may require antimicrobial therapy before, during, and after transplantation, but that it does not prevent successful transplantation. However, patients with LVAD-related infection appear to be at increased risk for invasive VREF infection, which may contribute to early mortality after transplantation^ref.
• Abiomed BVS 5000i^®


• Baylor Gyro BVAD^®
• Impella Recover 100 (IR100)^® (source : Impella CardioSystems AG, Aachen, Germany) : an intravascular microaxial flow BiVAD having a diameter of 4 mm, which is implanted, inclusive the electrical motor, in the failing heart and delivers an output of 2 to 10 liter/minute^ref. It is for use as short-term mechanical support (for a maximum of 7 days) for cases of acutely reduced left ventricular function^ref.
• total artificial heart (TAH) : bilateral atrial and arterial anastomoses are practiced. Current design goals are 90% reliability after 5 years of operation, which is comparable to the 70% 5-year survival for current heart transplant patients. In 1969 Cooley implanted the first TAH to a 47-years old patient with severe ischemic heart disease waiting for heart transplatation : he lived for 64 hours, was transplanted, but died after 32 hours due to Pseudomonas pneumonia. In 1981 a second Cooley patient died from Candida sepsis the 8th day after transplant. Currently 65-80% of patients with implanted TAH survive until heart transplantation, with posttransplantation 1-month OS = 50-70%, ie. 40% of patients requiring TAH has a good prognosis : most would have died without TAH.
• Jarvik-7^® (named for its designer Robert Jarvik) : Barney Clark, a 61-years old Seattle dentist, was the first patient to receive it in 1982 at Utah University hospital by surgeon De Vries : he lived for 112 days after the operation tethered to an external power unit the size of a washing machine, never able to leave the hospital room. Even though another patient with a Jarvik-7 lived for 620 days, the artificial heart became linked in the public mind with Barney Clark ordeal. The ventricles are connected to external pneumatic machine through transcutaneous cables.



• AbioCor^® TAH (source : Abiomed Corp. of Danvers, Massachussetts) consists of a whole network of machines : a 1-kg pump, a computer, a battery, a power converter, and cables to link them all together. It is a titanium and polymer construct that uses and electrohydraulic actuator system to shuttle blood alternately between the right and left pumps. Unlike with Jarvik-7, the diseased heart is completely removed and the TAH allows long-term survival without need for heart transplantation. It was received in June 2001 by the 59-years old man Robert Tools in Louisville (Kentucky, USA). After 4 months of relative well-being the patient died due to severe relapsing gastrointestinal hemorrhages due to the anticoagulant prophylaxis


• Liotta^®
• Akatsu III^®
• Symbion^®
• CardioWest (CardioWest Technologies Inc.)
• Currently under development are electromechanical TAH devices that directly convert electrical energy into mechanical action. A recent improvement has been the use of a transcutaneous energy transmission system (TETS) that transmits radio-frequency energy through intact skin, eliminating the risk of infection entering the body along the percutaneous power line that was used with earlier implanted blood pumps. A rechargeable implanted battery provides brief periods of freedom from the external battery and electronics associated with the TETS, allowing the patient to shower or, theoretically, even to swim. A TAH presents many design challenges. In all probability the patient will die very rapdily should the machine malfunction. Therefore, it must perform flawlessly. A major potential problem are blood clots that form around the 4 artificial plastic valves and in response to the synthetic material with which the blood is constantly in contact. Pulmonary problems may develop from a mismatch between right and left blood flows. Existing TAHs are large and do not fit many male candidates, most females, or any children. Should the TAH become infected, even intravenous antibiotics may bot be effective. Despite these problems, the TAH is a desirable device for certain patients with severe heart muscle fdamage, or for those where the continued presence of the native heart is a liability. However, because of the inherent limitations in current TAH technology, there is wide agreement that only about 10% of the patients with end-stage heart disease (approximately 5000 to 10,000 US candidates per year) should be treated with this radical approach.
Complications : hemorrhages (patients are treated with anticogulants to prevent thromboses), infections, hemolysis, thromboembolisms and immunological alterations.
Companies involved in MCS : Abiomed, Arrow International, Baxter Healthcare Corp., CardioWest Technologies Inc., Datascope Corporation, McGowan Center for Artificial Organs, University of Pittsburgh Medical Center, Thermo Cardiosystems Inc., Thoratec Corporation (formerly Nimbus Inc.)
• artificial cardiac valve / heart valve prosthesis : a substitute, mechanical or composed of tissue, for a cardiac valve affected by valvulopathies



• mechanical heart valves consist of a titanium ring in which one or more leaflets made of pyrolite carbon assure the opening and closing mechanism.
• bileaflet valve (fig. e) : a type of artificial cardiac valve consisting of a sewing ring surrounding a circular valve seat, to which are attached 2 semicircular occluding disks that swing open and closed to regulate blood flow.
• St. Jude Medical valve : a bileaflet valve with a Dacron sewing ring and pyrolytic carbon leaflets and housing, the leaflets opening to 85°
• caged-ball valve (fig. d) : a type of artificial cardiac valve consisting of a sewing ring to which is attached a cage made of curved struts; a ball within the cage floats freely, allowing passage of blood or occluding the orifice to prevent reflux.
• Cooley-Cutter valve : a caged-ball valve consisting of an open cage with four metal struts forming foot-like projections into the orifice, a cloth-covered sewing ring, and a disk-shaped or biconical poppet.
• Smeloff-Cutter valve : a caged-ball valve with 2 open titanium cages, one on each side of the valve ring, a barium-impregnated silicone rubber ball, and a Teflon sewing ring.
• Starr-Edwards valve : a caged-ball heart valve prosthesis consisting of a Stellite retaining cage containing a Silastic ball and a Teflon and polypropylene cloth–covered sewing ring
• Ionescu-Shiley valve : a type of artificial cardiac valve consisting of glutaraldehyde-fixed bovine pericardium constructed as a 3-cusp valve mounted on a Dacron-covered titanium frame
• tilting-disk valve (fig. e) : a type of artificial cardiac valve consisting of a sewing ring and a valve housing containing a suspended disk that swings between closed and open positions
• Lillehei-Kaster valve : a tilting-disk valve consisting of a titanium valve housing, Teflon sewing ring, and flat, free-floating, pivoting pyrolytic carbon disk that opens to 80°
• Omniscience valve : a modification of the Lillehei-Kaster tilting-disk valve, having a curvilinear pyrolytic carbon disk suspended in a one piece titanium frame with fin-like projections and a Teflon sewing ring.
• Omnicarbon valve : a tilting-disk valve similar to the Omniscience valve but entirely coated with pyrolytic carbon, including the sewing ring.
• Medtronic-Hall valve : a tilting-disk valve comprising a titanium valve housing an S-shaped disk guide strut, a Teflon cloth sewing ring, and a centrally perforated pyrolytic carbon–coated graphite disk that opens to an angle of 75°
• Björk-Shiley valve : a tilting-disk valve consisting of a cobalt alloy cage, a Teflon fabric–covered sewing ring, and a disk made of pyrolytic carbon that opens to an angle of 60°
Advantages : unlimited duration => used in young patients
Disadvantages : extrinsic nonstructural dysfunctions (mechanical noise, thrombogenic => need for anticoagulant chemoprophylaxis (prothrombin activity = 25-30) => hemorrhages, mechanical block, post-pericarditis periprosthetic detachment, hemolysis, mismatch)
• biological heart valve : an artificial cardiac valve composed of sterilized biological tissue
• homograft valve (fig. a) : a transplanted cardiac valve obtained from a human donor and antibiotic-sterilized or cryopreserved, then sewn into the native anulus.
• heterotopic homotransplant or transplant from living donor
• fascia lata
• dura mater
• homotopic transplant from cadaver
• heterograft or xenograft valve (fig. b and c) : a transplanted cardiac valve obtained from other animals
• porcine valve : a type of artificial cardiac valve made from a pig (Sus scrofa) aortic valve cured in glutaraldehyde and mounted on a supporting structure.
• Carpentier-Edwards valve : a porcine valve mounted on an Elgiloy alloy stent with a Teflon cloth–covered sewing ring
• Hancock valve : a porcine valve mounted on a semiflexible stent made of a Stellite ring and flexible struts of polypropylene
• stentless (replace also aorta root)
• decellularization of porcine valve => replacement with saphena vein-ex vivo-expanded endothelial cells => replacement of pulmonary valve (which is used to autologously replace the affected aortic valve) (Ross procedure / aortic pulmonary autograft implant). The Shelhigh No-React pulmonic valve conduit is a new porcine conduit that is glutaraldehyde-treated and detoxified using a proprietary heparin process^{ref1, ref2}.


• fixed bovine (Bos taurus) pericardium
Both tissues can be mounted on a metal frame (so-called 'stented valves') or they can lack this structure ('stentless valves').
Advantages : inaudible, low thrombogenicity => does not require anticoagulation
Disadvantages : the problem with these biological valves is the durability: during the years they start degenerating or calcifying => used in elderlies. To prevent this, recent biological heart valves are treated with an anti-mineralisation procedure
Complications : infectious endocarditis => need for antimicrobial chemoprophylaxis
Mortality rates after valve replacement and repair :
• aortic valve replacement (AVR) (isolated) : 4.3%
• mitral valve replacement (MVR) (siolated) : 6.4%
• multiple valve replacement : 9.6%
• AVR + coronary artery bypass (CAB) : 8.0%
• MVR + CAB : 15.3%
• multiple valve replacement + CAB : 18.8%
• AVR + any valve repair : 7.4%
• MVR + any valve repair : 12.5%
• aortic valve repair : 5.9%
• mitral valve repair : 3.0%
• tricuspid valve repair : 13.9%
• percutaneous mitral valve (MV) repair system (source : Evalve Inc., Redwood City, CA) : a tiny clip device that can be delivered by a catheter and deployed in the heart to repair a malfunctioning and leaking mitral valve is building a favorable safety and feasibility profile as it nears completion of a nationwide Phase I trial. To date, a total of 24 patients who suffered from severe mitral valve regurgitation (MR) have received the clip as part of a clinical study under an FDA-approved investigational device exemption (IDE). Based on these promising results, a larger, Phase II safety and efficacy trial may be initiated later this year. According to Evalve's web site, this device is being tested on people with moderate to severe mitral regurgitation, originating from the central two-thirds of the valve. Pts with rheumatic heart disease are excluded. There is no information provided about what type of Carpentier classified MR is being included in the testing.


• Coapsys^® (source : Myocor, Inc., Maple Grove, MN) annuloplasty device, which consists of anterior and posterior epicardial pads connected by a sub-valvular chord, aims to restore leaflet apposition and valve competency off-pump in functional mitral regurgitation (MR)


• devices for closure of patent foramen ovale: the protocol of most devices require general anesthesia and TEE monitoring during device implantation. Complete cardiac catheterization is performed to confirm the clinical and echocardiographic diagnosis and to exclude any associated anomalies, especially partial anomalous pulmonary venous return. Cineangiography from the right upper pulmonary vein at its junction with the left atrium in a left axial oblique view (30° LAO and 30° cranial) to serve as a landmark for use during device implantation is usually performed. Balloon sizing of the ASD by withdrawal of progressively decreasing sized balloons across the ASD has been used in the past. More recently, static sizing of the ASD with a compliant balloon has been described and is used by some investigators. In dynamic sizing in vitro measurement of the balloon diameter is used to measure stretched diameter. In both methods, balloon stretched diameter may be measured both by TEE and by cineradiography. It is also important to examine the atrial septum by TEE during balloon occlusion to ensure that there are no additional defects in the atrial septum. The recommended device/defect ratio varies with each device varying from 1.6 to slightly above 2. The method of implantation is different from one device to the other, although there are some similarities. The left atrial disc is initially delivered into the left atrium and under fluoroscopic and TEE guidance, withdrawn to position it against the left side of the atrial septum occluding the ASD. The tip of the device delivery sheath is withdrawn further across the ASD into the right atrium to effect release of the right atrial component of the device (Amplatzer^®, Angel Wings II^®, CardioSeal^®, StarFlex^® and Helex^®) or an additional right atrial disc (counter-occluder) delivered (Buttoned device^®). Following TEE verification of optimal position of the device without interference with other critical structures (pulmonary vein and mitral valve) the device is disconnected. A final TEE, angiography or hemodynamic evaluation are performed, as per the protocol of that particular device.
• CardioSEAL^®/STARFlex^® devices : these are re-designed versions of clamshell device. Complete and effective occlusion rates are slightly lower than other devices. Arm fractures continue to exist despite design modifications and is a cause for concern. Long-term follow-up results are needed for further evaluation of these devices. The first atrial septal defects (ASDs) were closed percutaneously in 1974 and reports about their closure published in 1976^ref. In the 1980s the Rashkind occluder was introduced and revived the interest in the topic^ref, focusing for the first time on the patent foramen ovale (PFO)^ref. This device is still available in two modified versions. Its initial name, Rashkind Clamshell occluder^®, has been changed, first into CardioSEAL^® and more recently into STARFlex^® (Nitinol Medical Technologies, Boston, Massachusetts, USA).
• Sideris Buttoned Device^®, 4th generation (Custom Medical Devices, Amarillo, Texas, USA) : this is the oldest of the currently used devices and has the largest published clinical experience, thus far. Unbuttoning problem seen with early generations of the device has been resolved by introduction of double button modification and the effective occlusion rate remained unchanged. But, the size of the device has to be at least 2 times as large as the stretched ASD diameter. To address this problem, a centering modification was undertaken.
• COD Buttoned device^® : this is a modified rounded 4th generation device with a centering mechanism. Reduction of size of the device without adversely effecting the device dislodgment or occlusion rate has been achieved. Experience in a larger cohort with longer-term follow-up are needed to confirm the early encouraging results.
• Angel Wings II^® device (Microvena Corp, Whitebear Lake, Minnesota, USA), more recently called the Guardian Angel device: clinical trials with this redesigned version of Angel Wings device have not yet started
 The ASDOS device^® and the Monodisk^® device are no longer on the market. All these devices had one major problem in common pertaining to the closure of ASDs. They were not self-centering. Although some efforts were made to remedy this with intricate nitinol wire constructions or interlinking filaments, the only self-centering device was introduced in the 1990s—the Amplatzer ASD occluder^ref
• Amplatzer^® septal occluder : this is a relatively new double-disc self-centering device, which can be implanted via small delivery sheath. Rapidly accumulating immediate and short-term data show good occlusion rates. It is easily retrievable and re-positionable prior to disconnecting the device. Thick profile of the implanted device and nickel toxicity have been pointed out as potential problems with this device. Long-term results are necessary to accept its short-term success. A derivative of the Clamshell occluder was dedicated to PFO closure and named PFO STAR (Cardia, Burnsville, Minnesota, USA); the Amplatzer occluder was modified for the PFO indication and first implanted on 10 September 1997 by the author in the presence of Kurt Amplatz, the inventor. Like all Amplatzer occluder devices, it consists of a nitinol mesh double disk containing polyester fabric inside the two disks. The disks are connected by a thin neck formed by the woven wires forming the disks. The neck is twisted around its long axis and hence is extendable. The device has to be screwed on to a pusher/puller cable and pulled into an 8 or 9 French introducer sheath. When pushed out of the sheath, it resumes its disk shape instantly. The whole process is fully reversible, as many times as required, up to the moment the device is unscrewed from the pusher/puller cable. 3 sizes of Amplatzer PFO closure devices are available and named after the diameter of the right sided disk. The most commonly used 25 mm Amplatzer PFO occluder features a right sided disk of 25 mm diameter and a left sided disk of 18 mm diameter. The 18 mm PFO occluder comprises two 18 mm disks and is meant for small PFOs with a stable septum primum. The 35 mm Amplatzer PFO occluder is destined for large PFOs with an extremely redundant and flimsy septum primum (atrial septal aneurysm) and features a 35 mm disk on the right side and a 28 mm disk on the left side. It requires a 9 French sheath in contrast to the two smaller devices fitting through an 8 French sheath. The PFO STAR device is made by square Ivalon patches fixed to 2 nitinol cross bars connected by a 2 mm metallic axis. The pusher cable is derived from a biotome gripping a ball tip at the right side end of the axis. The 2 squares are attached in a way that the blunted nitinol wire tips alternate to avoid pressure necrosis of the tissue between two opposed wire ends. The device comes in several sizes. The minimum sheath size is 10 French, the maximum 12 French. More recent generations of these devices feature a hexagon shape.
• Helex Septal Occluder^® (WL Gore and Associates, Flagstaff, Arizona, USA)  : this is a new device, appears to be useful in occluding small defects. Immediate results appear to be comparable to other devices. Device implantation technique, though complex, can be mastered. It is retrievable until final release. Follow-up results and experience in larger cohorts are necessary prior to making predictions of its success. The Helex septal occluder is the most recently developed of the devices. It is a hybrid device recommended for PFO and ASD closure. It has a distensible neck in its final shape but also some self-centering qualities. It consists of a nitinol wire with a polytetrafluoroethylene patch attached in a banner-like fashion. Within the delivery sheath it is fully elongated. Once implanted, it forms a double disk (15–35 mm in diameter), one on each side of the septum. The sheath size required for implantation is 9 French.
• transcatheter patch : all the currently used devices are double disc devices with wire components and are therefore subject to limitations and complications associated with such designs. Transcatheter delivered patches also are useful in occluding defects that cannot be closed with conventional double-disc devices. However, there is only a limited clinical experience with these devices. Results of larger international clinical trials and FDA approved (with IDE) human trials are awaited.

The most widely used devices so far for PFO closure are the CardioSEAL device in the USA (2000 to 2002 the only one available) and the Amplatzer PFO occluder in the rest of the world (available in the USA since 2002). The CardioSEAL device was the model for the PFO STAR and differs from it in two major aspects. The wires spreading the tissue have joints made of spring coils shaped into the wires, and the release mechanism employs the principle of a ball tip end on the device and a ball tip end on a cable within the pusher. Inserting the ball tip end of the device first into a cylindrical case at the end of the pusher and pulling back the ball tip end of the pusher cable, the ball tip of the device is enclosed in the case until the pusher cable is pushed out of the case first, thereby releasing the ball tip end of the device. The introducer sheaths required for the different sizes of the CardioSEAL or STARFlex devices vary from 10 to 14 French. Many aspects of the implantation technique pertain to all devices. It is essential that the diagnosis of a PFO is made unequivocally. The gold standard for this is a transoesophageal echocardiogram with a bubble test after a sustained Valsalva manoeuvre. Immediately after the release of the Valsalva manoeuvre, the blood kept out of the thorax during Valsalva rushes back and fills the right atrium before filling the left atrium, thereby opening the PFO. Provided the correct plane is visualised, bubbles can be seen passing through the gaping foramen ovale. The septum secundum positioned towards the aorta is usually a robust, wedge-like structure. The caudad septum primum may be paper thin and extremely mobile. An atrial septal aneurysm is usually diagnosed when the septum primum moves more than 10 mm into each atrium during certain phases of the cardiac cycle. Aneurysm is a misnomer as it merely describes the mobility of the non-muscular part of the septum primum. Nevertheless, the presence of a highly mobile septum primum has been identified as an important risk factor for paradoxical embolism. Initially, this mobility had been considered a risk factor on its own, even in the absence of a PFO. Today this can no longer be supported^ref. It is now generally accepted that closing the hole is all it takes. Stenting the septum primum with a large device is no longer deemed necessary. A PFO clearly demonstrated by this transoesophageal echocardiographic technique can be found during catheterisation virtually always by sliding a curved catheter along the septum primum towards the head. This opens the foramen and provides passage. If the passage cannot be found easily (funnel shaped PFO) the manoeuvre is repeated, torquing the catheter as soon as it engages the entrance of the PFO. A wire may be required for further probing in difficult cases. Not infrequently, the initial wire pushed up into the right atrium through the puncturing needle may already find the left atrium in the presence of a widely patent PFO with an atrial septal aneurysm. Rarely, a PFO cannot be found in spite of echocardiographic proof of bubbles passing in that region. There may be two reasons for this. First, the foramen ovale may be fused over almost the entire width of its mouth. If only one corner remains open, the gap may be in an awkward direction for the catheter probing through a funnel shaped entrance of the PFO. An injection of contrast medium through the probing catheter can usually demonstrate the location of the PFO, but it will require an investment of time and material to negotiate such a passage. Moreover it is doubtful that such a small, difficult to find passage would be a significant risk for paradoxical embolism. Second, instead of a PFO a small ASD may be located in the vicinity of a completely fused foramen ovale. If such a small hole is located in a flimsy part of a volatile septum primum it may be as difficult to find as a needle in a haystack and probably irrelevant for paradoxical embolism as well. The fitting sheath for the selected device is placed over the wire positioned in the left atrium. The obturator of the sheath is withdrawn slowly to avoid aspiration of air during this process. If a sheath without a sidearm is utilised (currently only with the Amplatzer method), the risk of air embolisation during the procedure is small as the sheath can be flushed just like any regular cardiac catheter. If sheaths with side arms are used, purging is much more difficult and incomplete and a certain percentage of clinically manifest air embolism cannot be avoided. This becomes usually apparent by transient ST elevation in the inferior leads. Given the supine position, air bubbles will travel along the ventral side of the left ventricle and the sinus of Valsalva, thereby finding the take-off of the right coronary artery most easily. Only large quantities of air are clinically relevant as they may block flow in both coronary arteries or lead to cerebral symptoms. Persistent sequelae of air embolism during PFO closure have not been described^ref. The left sided portion of the device is opened in the left atrium, and sheath and device are pulled back as a unit until the open left sided device lies parallel against the interatrial septum, thereby closing the PFO. Under continued tension on the device the sheath is withdrawn further and the right sided part of the device released, embracing the atrial septum. The correct position is assured and the device is released with a final assessment of the device position. Even if the PFO is initially not tight, it may become completely closed, once endocardium has overgrown the device. This may take several weeks to months. Debate is ongoing whether echocardiographic guidance is required for percutaneous PFO closure. Transoesophageal echocardiography yields the best picture but it is uncomfortable for the patient lying on his back on the catheterisation table. Only when it is undertaken in combination with intubation and general anaesthesia does it provide a comfortable working environment for the operator without having the patient suffer unnecessarily. However, it turns a relatively simple 15 minute procedure into a demanding intervention for the patient with a prolonged recovery period, not to mention a sore throat. A more convenient alternative is intracardiac ultrasound. However, this increases the cost of the already expensive device implantation. In addition it requires a second large (12 French) venous access, which increases the risk of an arteriovenous fistula at the puncture site, an already not too rare complication of the procedure, occurring in about 1% of cases. In our own experience with over 500 PFO closures without echocardiographic guidance, the need to summon echocardiography for assistance in the middle of the procedure has never arisen. Several technical advantages favour the Amplatzer PFO occluder. Although rarely an issue, even after deploying the right sided disk, the device can be easily recaptured into the sheath and placed a second time or removed. This also holds true for the Helex septal occluder which, however, is by far the most difficult to use device requiring a long learning curve. All other devices have to be completely removed outside the body and often replaced by a new one, if the position is unacceptable after releasing the right sided disk. The fact that they are folded like two umbrellas connected at their tip within the delivery sheath explains that pulling the right sided disk back will fold it the wrong way. This must be corrected outside of the body, and the same device can only be re-used if it has not been distorted during the process. The Amplatzer PFO occluder and the Helex septal occluder are also the only devices that conform to long funnel shaped PFOs. The other devices have a fixed length of the connector between the disks. This usually results in an incomplete opening of the right sided umbrella, which remains partially stuck in a tunnel-like PFO. This may still provide a complete occlusion of the PFO, but the device arms sticking out into the right atrium like a half opened umbrella are not a nice sight for the echocardiographer. For procedures done without echocardiography, a few anatomical landmarks have to be observed during implantation. Contrast medium injection once the device is deployed but still connected to the pusher cable is possible with all techniques. In a usually left anterior oblique projection with a cranial tilt, the two disks can be seen in profile. They form two lines almost parallel in the inferior part where the membranous part of the septum primum barely divides them. In the superior part they should be gaping at a certain angle as they embrace the wedge-like septum secundum. This has been alluded to by us as the Pacman sign, making reference to the well known arcade figure gobbling up a dot^ref. The contrast medium outlining the right atrium should clearly depict that all of the right sided disk and none of the left sided disk is in the right atrium. A residual shunt into the left atrium seen at that moment is of little importance for the reasons stated above. A final check before release is usually done by wiggling on the pusher cable. On an extremely rare occasion, the device may be dislocated by such wiggling. In that case it should be replaced by a larger device. Observing these precautions, embolisation of the device during or after implantation has not occurred in our experience of several hundred consecutive cases. Embolisation remains, however, one of the possible complications^ref. Other serious complications occurring in < 1% of cases are infection, erosion into the pericardium or aorta at the rim of the device, a new ASD caused by the lower rim tearing the thin septum primum, or paroxysmal or persistent atrial fibrillation. Some palpitations are not uncommon and are reported by about 10% of patients in the early period after implantation. They usually subside spontaneously.  Thrombosis on the device has been found in about 6% of devices used for PFO closures at the one month transoesophageal echocardiographic control in 1000 patients, except for the Amplatzer PFO occluder where it was found in < 1%^ref. A recent study reported thrombosis on five of 23 CardioSEAL devices checked at one month but in none of 20 Amplatzer PFO occluders^ref. It can be assumed that clinical results will be better with devices providing a higher complete closure rate^ref. An experience with 170 patients treated with 5 different devices and a follow up of about 7 years showed an incidence of recurrent embolic events such as transient ischaemic attacks, cerebral vascular attacks, or peripheral embolism of 2% per year. This contrasts favourably with the literature on the natural course of patients with a PFO and a cryptogenic stroke with an incidence of recurrences of 5% per year^ref. It contrasts even more favourably with 42 patients on warfarin and 56 patients on aspirin during a two year follow up after a cryptogenic stroke in the presence of a PFO with a recurrence rate of death or stroke (transient ischaemic attacks not included) of 5% and 9% per year, respectively^ref. Pooling the results of the published studies on percutaneous PFO closure, it appears that the intervention yields favourable clinical results over a conservative treatment.6 Randomised trials, however, are not yet available. Three of them are on the way. The PC trial, an international multicentre trial with the Amplatzer PFO occluder, started recruiting patients about three years ago but has yet to reach the target of 425 patients. The RESPECT trial, which is also using the Amplatzer PFO occluder exclusively, has started recruiting the planned 300 patients late in 2003, as has the CLOSURE trial using the CardioSEAL device with an inclusion target of 1600 patients. These two trials are multicentre trials in the USA. What comes closest to a randomised trial is a contemporary matched control follow up study in about 300 patients, of which half were arbitrarily sent for PFO closure by neurologists and half were treated conservatively by the same physicians. Already at a follow up of four years, there was a trend in favour of device closure with an average of 5% events per year (counting all neurological or peripheral symptoms) compared with 7% in the conservative group. This advantage was significant in terms of major strokes which only occurred in the conservative group. It was also significant in the subgroup of patients who had had more than two events before treatment allocation as well as in those who had complete closure at the six month follow up transoesophageal echocardiography in the device group^ref. Comparing clinical results of device closure with surgical results, a certain similarity can be observed^ref. Considering the inconvenience of heart surgery (general anaesthesia, scar, prolonged hospital stay, wound infection, etc), and the fact that not a single emergency operation following a failed attempt of device closure for PFO has been described even though all patients could have undergone uncompromised surgery in case of failure, it appears imperative that device closure is the therapy of first choice. The clinical justification for PFO closure after paradoxical embolism is still lacking. Other indications are even less supported by clinical data. Nonetheless, growing interest has been observed recently in the connection between the PFO and diving accidents^ref or migraine^ref. The indication for PFO closure in the face of paradoxical embolism has been widely accepted in the presence of an atrial septal aneurysm or after several embolic events. It has been adopted at our centre for any suspicion of paradoxical embolism. This has introduced a frequency of PFO closure at our catheterisation laboratory that already amounts to > 10% of the entire interventional activity. This is bound to increase further, particularly if divers and patients with migraine are included. Finally, a scenario can be envisioned where the PFO is considered a threat for longevity. A pathological study on about 1000 necropsies of all ages showed that the prevalence of PFO decreased from 33% in those younger than 30 years of age to 20% in those over 80^ref. Late spontaneous closure of a PFO is a likely explanation, but not less likely is a selective mortality of PFO carriers. Percutaneous PFO closure is the most simple procedure in interventional cardiology. Further simplified and ubiquitously available it could be advocated as a kind of mechanical vaccination against the late hazards of a PFO. It would have to be carried out in teenagers screened for a PFO. Before that age, venous thrombosis is virtually unheard of and the PFO has no potential for harm^ref.
• vascular grafts : production in infected grafts appears to be organism specific and is significantly higher with P. aeruginosa than common Gram-positive organisms. Increased neointimal hyperplasia from subclinical infection may be a significant factor contributing to late graft failures^ref.
• a plastic that releases nitric oxide could be used to make artificial blood vessels that do not clog. Plastic tubes can be used to replace blood vessels, for example in coronary-artery bypass surgery. But tubes smaller than about 6 mm across are often gummed up by platelets, the sticky blood cells that make clots. If the clots break free, they can damage the brain and other organs. Nitric oxide stops platelets sticking together, and stimulates growth in the cells lining blood vessels. Even better, it stunts the growth of smooth muscle that could squeeze an artificial vessel closed. Anticoagulant drugs are often used to keep conventional grafts clear, but this can lead to bleeding elsewhere in the body. Releasing nitric oxide directly from the graft localizes the effect to where it is needed. To combat these issues, a diazeniumdiolate-modified nitric oxide (NO)-producing peptide (that decomposes in water) was incorporated into a polyurethane to improve the thromboresistance of this biocompatible polymer. NO production by polyurethane films occurred in a burst over the first 2 days, and then more slowly for approximately 2 months under physiological conditions, and mechanical properties of the material were suitable for vascular graft applications. Platelet adhesion to NO-releasing polyurethane was dramatically decreased compared to control polyurethane. Furthermore, endothelial cell growth was stimulated in the presence of the NO-releasing polyurethane, while smooth muscle cell growth was greatly inhibited. The ability of this bioactive material to inhibit platelet adhesion and smooth muscle cell proliferation while encouraging endothelialization suggests that this NO-generating polyurethane may be suitable as a candidate material for small-diameter vascular grafts. Nitric oxide, and not the polyurethane, is behind the effect, because they see an increase in cGMP in the cell cultures. Nitric oxide triggers the release of cGMP, which then affects enzyme activity and gene expression. Plastic that had shed all its nitric oxide had no effect on the cells. Polyurethane is more flexible than other plastics used in grafts, so it should also integrate with surrounding tissue better^ref
• stent : a slender rodlike or threadlike device used to provide support for tubular structures that are being anastomosed, or to induce or maintain their patency
• endoprosthesis : a hollow stent
• ureteral stents
• endotracheal stents
• the Dumon stent used to treat tracheobronchial stenosis^ref requires use of a technically difficult rigid bronchoscope and rigid introducer. (Both the rigid fiberscope and introducer are manufactured by Efer; La Ciotat, France.). Umemoto et al reported positioning a Dumon stent using a unique endotracheal tube for introduction and a slender tube for placement. Their method was modified in this study by using a conventional endotracheal tube for introducing the stent and a cylindrical-tipped stainless steel wire to push and place the stent. The present method has the following advantages over a conventional rigid bronchoscope: (1) the use of an endotracheal tube and flexible bronchoscope makes introduction easy for the practitioner and less stressful for the patient; (2) even though the diameter of the endotracheal tube introducer is somewhat larger than that of the original rigid introducer, the endotracheal tube is introduced more easily and safely than the rigid introducer because of its flexibility; and (3) the procedure is used more easily by the practitioner and less stressfully for problem cases such as in patients who cannot easily stretch their necks or who suffer from left main bronchial stenosis or tumor compression with tracheobronchial displacement that makes use of a rigid system difficult. The present method does, however, have the disadvantage of difficulty in removing or repositioning the stent because of weakness of the grasp of the flexible forceps via a flexible bronchoscope. One may therefore conclude that Dumon stent placement using a conventional endotracheal tube is useful, particularly in malignant stenosis, which usually does not require stent removal. A rigid fiberscope or tracheostomy should, however, be used for removing or repositioning the stent.
• Aero™ Hybrid Tracheobronchial Stent System (source : Alveolus, Inc)
• esophageal stents
• biliary tract stents : a hollow metal stent in a bile duct, allowing biliary drainage across an obstruction

Indications :
• cholangiocarcinoma (41.5%)
• pancreatic adenocarcinoma (40.5%)
• nodal metastases at the porta hepatis (14.6%)
• gallbladder adenocarcinoma (3.4%).
• intravascular stents : a metal wire or tube introduced into a stenotic blood vessel to create and maintain luminal patency. Ferromagnetic stents are considered MRI-compatible 6 weeks after implantation. They play an increasingly important role in the treatment of iliac artery occlusive disease and their use has expanded the indications for percutaneous endoluminal therapies. The past several years have seen a sharp increase in the number of commercially available covered and uncovered stents. Knowledge of their design and mechanical properties is crucial for selecting the appropriate stent for a particular type of lesion. In this article, the indications for and results of iliac artery stent placement are reviewed and the various characteristics of the currently available stents that may influence operator choice for use in specific lesions are discussed. The past 25 years have seen the evolution of percutaneous transluminal angioplasty (PTA) from an experimental procedure to a widely used, effective, minimally invasive method of treating peripheral arterial occlusive disease. Although technical success rates are reportedly high, in a defined percentage of patients, suboptimal results caused by intimal dissection or elastic recoil lead to early failures. Most of these failures can be salvaged by placement of a stent, which serves as an intravascular scaffold of thin spiral or mesh-shaped metal struts buttressing the arterial wall from the inside, thereby maintaining luminal patency. Indeed, the advent of stents has brought with it an expanded application of percutaneous endovascular therapies, particularly for the treatment of aortoiliac inflow disease, in which patency rates approach those of surgical bypass grafts without the inherent risks associated with surgical revascularization^{ref1, ref2, ref3, ref4, ref5, ref6, ref7, ref8, ref9, ref10, ref11, ref12, ref13}. The Food and Drug Administration (FDA) has defined the indications for iliac artery stent placement as follows: (i) a suboptimal result of PTA caused by an extensive dissection or residual pressure gradient, (ii) total occlusions, and (iii) recurrence after PTA. However, in clinical practice, iliac artery stent placement is more widespread, with many practitioners electing to place stents as the primary treatment for many lesions. Currently, there are only two stents approved by the FDA for placement in the iliac arteries: the large Palmaz 308 stent (Cordis, Warren, NJ) and the Wallstent (Boston Scientific/Medi-tech, Natick, MA). However, numerous other stents that are FDA-approved for biliary or tracheobronchial applications have been used in an "off-label" fashion in the iliac arteries. Indeed, the multitude and variability of the available stents has made stent selection a confusing issue for many interventionalists. Unfortunately, there is little objective data comparing the properties of different stents, and the information that exists is unavoidably dated relative to the rapidly evolving market of peripheral stents. Nonetheless, this article is a review of the various characteristics of the currently available stents as well as the indications and results of iliac artery stent placement. It is not intended as an endorsement or a criticism of any particular product, but rather as a commentary on the properties of the various stents, which may influence operator choice for use in specific lesions.

Classification of stents : the evolution of stents has occurred along 2 fundamental design philosophies: balloon-expanded stents and self-expanding stents. The prototypical balloon-expanded stent is the Palmaz stent; that of the self-expanding variety is the Wallstent. Beyond these, there is a myriad of commercially available uncovered and covered stents with varying materials, designs, and mechanical properties. The metals used for stents include stainless steel, nitinol, tantalum, platinum, and various metal alloys. Construction of stents is quite variable and includes laser-cut or etched, woven, knitted, coiled, or welded constructions. Stent properties, such as flexibility, radial strength, hoop strength, radiopacity, and foreshortening characteristics, differ among the stents. In addition, various characteristics of a stent such as metal thickness, surface charge, method of stent cleaning and polishing, source of the metal, corrosion resistance, durability, the amount of open area–to–metal surface ratio, "kinkability" of the stent, and the sharpness of the ends of the stent may all impact on the biocompatibility of the stent, and ultimately, long-term patency rates. All stents are not alike, and no stent exhibits all the ideal properties:
• high radial force/hoop strength to resist recoil
• minimal or no induction of intimal hyperplasia or restenosis
• longitudinal flexibility to negotiate tortuosity
• high radiopacity for visualization
• radial elasticity or crush resistance
• ability conform to the vessel
• low profile and high expansion ratio
• minimal or no foreshortening for precise placement
• easy deployment system
• maintenance of side branch patency
• MRI compatibility
• durability
• low price
Knowledge of the design characteristics and properties of the various devices may help guide selection of the appropriate stent for a particular lesion. At the risk of being outdated by the time of publication, the following is a list (in alphabetical order) of currently available stents that may be appropriate for iliac artery indications. The stents are divided into 4 categories: balloon-expanded stents, self-expanding stents, and covered stents.
• uncovered stents (used to treat stenoses)
• short balloon-expanded stents that are primarily indicated for use in aortic branch vessels (eg, renal arteries) have been omitted because they are rarely used in the iliac arteries.
• balloon-expanded stents : slotted tubes of malleable metal which are expanded by a coaxial balloon to the size of the target vessel. When deployed, the metal tube maintains the nominal balloon diameter and fixation to the arterial wall is enabled by friction and subsequent endothelialization. The plastic nature of these stents provides resistance against elastic recoil of the vessel, but also renders them susceptible to irreversible deformation when subjected to a powerful external compressive force.

Stent  Vendor  Material  Construction  Diameter (mm)  Introducer Size (F)  Key Features  List Price ($)  AVE Bridge Flexible Medtronic AVE 316L stainless steel Welded 6–10 7–8 High hoop strength, flexible in one plane 1,395–1,595 Assurant Medtronic AVE 316L stainless steel Welded 6–10 6 Low profile, flexible, good hoop strength, minimal shortening 1,345–1,495 Express Boston Scientific 316L stainless steel Laser-cut 6–10 6–7 Low profile, flexible, good hoop strength, minimal shortening 1,295–1,495 Intrastent Sultzer Intratherapeutics 316L stainless steel Laser-cut 5–10 6–7 Open cell design, flexible, minimal shortening 950–1,570 Intrastent DS (Double Strut) Sultzer Intratherapeutics 316L stainless steel Laser-cut 5–10 6–7 Same as Intrastent but with double parallel struts for added hoop strength 1,150–1,750 Intrastent LP (Low Profile) Sultzer Intratherapeutics 316L stainless steel Laser-cut 4–7 5–6 Low-profile version of Intrastent for 0.018-inch compatible balloon catheter 950 Intrastent LD (Large Diameter) Sultzter Intratherapeutics 316L stainless steel Laser-cut 8–12 8–10 Same design as Intrastent DS for good hoop strength; large diameters 1,350 Omnilink/Megalink Guidant 316L stainless steel Laser-cut 5–10 6–7 Low profile, flexible, good hoop strength, minimal shortening, 0.018- and 0.035-inch guide wire compatible 1,000–1,595 Palmaz Genesis Cordis 316L stainless steel Laser-cut 4–11 5–7 Low-profile, flexible, good hoop strength, minimal shortening, 0.018- and 0.035-inch guide wire compatible 1,050–1,795 Palmaz Large Cordis 316L stainless steel Laser-cut 8–12 8–10 Large diameters, rigid, less hoop strength than Palmaz medium, FDA-approved for iliac use 895 Palmaz Medium Cordis 316L stainless steel Laser-cut 4–9 6–7 Rigid, high hoop strength 895 Palmaz-Schatz Long-Medium Cordis 316L stainless steel Laser-cut 6–10 6–7 Articulated, less hoop strength than Palmaz medium 1,200 Strecker Boston Scientific/Medi-tech Tantalum Knitted 4–12 7–8 Flexible, good vessel conformability, very radiopaque; not available in US NA Vistaflex Angiodynamics Platinum Woven 5–10 7–8 Flexible, good vessel conformability, very radiopaque, MR compatible 1,195

Most balloon-expanded stents are made of stainless-steel alloy. Stainless steel is a suitable material for this purpose because it combines malleability and strength. For this reason, balloon-expanded stents tend to have greater radial and hoop strength than self-expanding stents. They also tend to be more radiopaque. Because of their mode of deployment and the fact that there is a minimal amount of predictable foreshortening, balloon-expanded stents can be placed with a high degree of precision. The first available and most widely used balloon-expanded stent to be produced was the rigid Palmaz stent, the efficacy of which has been validated in clinical trials^{ref1, ref2, ref3}. Recently, there has been a drive toward manufacturing low-profile, flexible, balloon-expanded stents with complex cell geometry and open or hinged designs. These stents derive their flexibility from the ability of individual cell rings to track independently. However, there is very little scientific data comparing the clinical impact of different stent designs and cell configurations. Nowadays, many balloon-expanded stents are available premounted on balloon catheters. If nonmounted stents are employed, manufacturers generally recommend use of scratch-resistant balloons to avoid balloon rupture during stent deployment. In addition, the compliance of the angioplasty balloon used for stent deployment may also be important. A noncompliant balloon can be useful for expanding the stent in areas of increased resistance because a compliant balloon tends to expand most outside the lesion. However, the mild overexpansion of a semicompliant balloon at higher inflation pressures may result in some molding of the stent to the contour of the vessel. Palmaz^ref has suggested that the embedding of the stent members into the arterial wall as a result of mild overexpansion is a crucial factor in the promotion of rapid multicentric reendothelialization. Because most balloon-expanded stents are made of ferromagnetic stainless steel, they cause large susceptibility artifacts (signal loss) on MRI. Notable exceptions to this rule are the Strecker and Vistaflex stents, which are made of tantalum and platinum, respectively.
• AVE Bridge Flexible and Assurant Stents (Medtronic AVE, Santa Rosa, CA) has replaced the old AVE Iliac Bridge stent, which was a rigid, balloon-expanded stent marketed outside the United States. Unlike most balloon-expanded stents, the Bridge Flexible stent is a modular device, which is not laser-cut from a single tube of 316L stainless steel. Instead, sinusoidal ringed elements of 316L stainless steel with six crowns are laser-fused or welded to form a tubular structure with simple rectangular cell geometry. Like the rigid Palmaz stent, the Bridge stent is characterized by good hoop strength. However, the pattern of welding (two of six crowns at opposite sides of the ring) imparts flexibility to the stent in one plane (perpendicular to the plane of the welded crowns). Practically, the stent rotates when negotiating tortuous anatomy. The old AVE Iliac Bridge stent was welded at all six opposing crowns and, therefore, was more rigid. The AVE Bridge Flexible stent is available premounted in lengths ranging from 28 mm to 100 mm and diameters ranging from 6 mm to 10 mm. Stents with diameters as large as 8 mm are compatible with a 7-F introducer system, whereas the 9- and 10-mm stents require an 8-F introducer sheath. Current plans at Medtronic AVE are to replace the Bridge Flexible stent with a low-profile stent called the Bridge Assurant. Apart from a lower profile, this stent has greater flexibility in all planes as a result of a variation in the design, with eight crowned sinusoidal rings and a staggered welding pattern. The Assurant is available in the same diameters as the Bridge Flexible stent but is compatible with 6–7-F introducer sheaths.
• Express Stent (Boston Scientific Corporation, Natick, MA) is the newest stent produced by Boston Scientific and the first balloon-expanded stent the company has offered to the U.S. market. Like many balloon-expanded stents, the Express stent is laser cut from a tube of 316L stainless steel. The stent consists of alternating rows of large and small sinusoidal rings with five interconnecting struts between the rings. This cell architecture is designed to provide flexibility and vessel conformability while maintaining hoop strength and radiopacity. The stent is available premounted on 0.035-inch guidewire compatible balloon catheters in diameters ranging from 6 mm to 10 mm and lengths ranging from 17 mm to 57 mm.
• Intrastent Stents (Sulzer Intratherapeutics, St. Paul, MN) is a balloon-expanded stent, laser-cut from a single tube of 316L stainless steel. The stent has a unique open cell design, which affords a high degree of flexibility and virtually no foreshortening on deployment. It is available in lengths as long as 76 mm and diameters as large as 10 mm. Mounted on 5-F balloon catheters, the stent can be advanced through 6–7-F introducer sheaths. For situations in which high hoop strength is required, there is a stronger version of the Intrastent called the Intrastent DS (Double Strut), which has paired parallel stent struts. Other versions of the Intrastent include the Intrastent LP (Low Profile), designed for use in small vessels over a 0.018-inch guide wire system, and the Intrastent LD (Large Diameter), for vessels as large as 12 mm in diameter. All models of the Intrastent have the same cell design and exhibit minimal foreshortening.
• Omnilink (premounted on a balloon catheter) and Megalink (unmounted) stents (Guidant, Santa Clara, CA) have the same multiple linked corrugated ring design (4–4–4 pattern) as the Multilink coronary stent, which is produced by the same company. The stents are cut from a single 316L stainless-steel tube. The Omnilink stent belongs to the newer generation of low-profile flexible balloon-expanded stents with good radial force. The stent is available in both 0.018-inch and 0.035-inch guide wire systems, in diameters ranging from 5 mm to 10 mm and lengths from 18 mm to 58 mm, and is compatible with 6–7-F introducer sheaths.
• Palmaz and Palmaz-Schatz Stents : the Palmaz stent (Cordis) is one of two metallic stents FDA-approved for use in the iliac arteries. It is a rigid stent laser-cut from a single tube of malleable 316L stainless steel with staggered rows of diamond-shaped fenestrations or cells. The stent foreshortens slightly when expanded, to an extent that is proportional to the amount of expansion. This predictable and minimal foreshortening, combined with its good radiopacity, allows for precise placement of the Palmaz stent. The medium-sized Palmaz stents (which are not FDA-approved for iliac use) are available in diameters ranging from 4 mm to 9 mm and lengths ranging from 10 mm to 39 mm. These can be mounted on 5-F balloon catheters and deployed through 6- or 7-F introducer sheaths. The large Palmaz stents (which are FDA-approved for iliac use) have diameters ranging from 8 mm to 12 mm in lengths as long as 30 mm. However, these stents require a 5.8–7-F balloon catheter and a 7–10-F introducer sheath. The Palmaz-Schatz long medium Stents are long, articulated versions of the Palmaz stents, available in diameters ranging from 6 mm to 10 mm and lengths from 41.8 mm to 77.8 mm. Unlike the regular Palmaz stents, these stents can be easily advanced across the aortic bifurcation for contralateral iliac artery stent placement.
• Palmaz Genesis Stent : the Genesis stent (Cordis) is the newest stent manufactured by Cordis and represents improved technology in terms of flexibility and foreshortening characteristics. Like the Palmaz stent, it is a laser-cut 316L stainless-steel slotted tube. The Genesis stent is low-profile and consists of seven rows of staggered cells connected by an S-shaped hinge. This design affords the Genesis stent far greater flexibility and less foreshortening than the Palmaz stent while maintaining good hoop strength. For sizing purposes, the Genesis stents are divided into two groups: large (5–11 mm) and medium (4–9 mm). The stents are identical in terms of design and construction and differ only with regard to length. The Genesis medium stents range in length from 12 mm to 24 mm and are therefore more suitable for treatment of aortic branch vessels (eg, renal artery stent placement). However, they can also be used for very focal lesions in the iliac arteries. The Genesis large stent is available in lengths ranging from 29 mm to 79 mm. All sizes are available unmounted or premounted on a balloon catheter. The stent can be mounted on balloon catheters designed for use on a 0.018- or 0.035-inch guide wire platform.
• Strecker Stent (Boston Scientific/Medi-tech) is constructed from a single tantalum filament, knitted into a tubular wire mesh and premounted onto a variety of balloon sizes. Tantalum, an inert, nonferromagnetic metal, provides excellent radiopacity with fluoroscopy and is MR-imaging–compatible. The Strecker stent is highly flexible and conforms well to tortuous anatomy. However, it has less hoop strength than the Palmaz stent. The Strecker stent has been used in Europe for more than a decade with good clinical results^{ref1, ref2}, but it is not available in the United States.
• Vistaflex Stent : like the Strecker stent, the Vistaflex stent (Angiodynamics, Queensbury, NY) is not made of stainless steel, but rather of an alloy consisting of 90% platinum and 10% iridium. The stent consists of a single wire woven in a sinusoidal fashion, radially coiled, with a longitudinal wire connecting the ends of the stent. The advantages of platinum are essentially the same as those of tantalum—improved radiopacity and MR imaging compatibility as well as a high degree of flexibility and ability to conform to tortuous anatomy. The Vistaflex stent is available in diameters as large as 10 mm and lengths as long as 55 mm and is compatible with 7-F and 8-F (9–10-mm) introducer sheaths. A unique construction characteristic of the Vistaflex stent is its beveled ends, a theoretic advantage in terms of minimizing circumferential hyperplastic tissue narrowing the functional vessel luminal diameter.
• self-expanding stents at transition from room temperature (RT) to 37°C of human blood within 24-48 hours : the mode of action of self-expanding stents is based on the property of elasticity. These stents self-expand to their nominal diameter when released from a constrained state within the delivery system. Typically, a stent with a nominal diameter greater than that of the target vessel is chosen so that it exerts an outward expansile force, resulting in good vessel wall apposition. The ultimate diameter of the vessel with the stent is the result of equilibrium between the elastic recoil of the vessel and radial expansion force of the stent. The elasticity of self-expanding stents is derived either from the design of the stent mesh, such as with the Wallstent, or from the elastic properties of the metal, such as with nitinol stents
 

stent  vendor  material construction  diameter (mm)  introducer size (F)  key features  list price ($)  AVE Bridge SE Medtronic AVE Nitinol Laser-cut 5–10 7 Flexible, minimal shortening, vessel conformability, single-handed deployment 1,450–1,550 Dynalink (0.035-inch) Guidant Nitinol Laser-cut 5–14 6–7 Flexible, minimal shortening, vessel conformability, single-handed deployment, 6-F introducer (up to 10 mm). Large diameters 1,595–1,795 Dynalink 0.018 Guidant Nitinol Laser-cut 5–10 6 Same as Dynalink 0.035 but compatible with 0.018-inch guide wire 1,645–1,795 Intracoil Sulzer Intratherapeutics Nitinol Coiled 4–8 7–9 Highly flexible, low metal surface–to–open area ratio, significant shortening 1,400–1,550 Luminexx Bard Nitinol Laser-cut 6–12 6 Flexible, minimal shortening, vessel conformability, tantalum markers, tapered inner catheter prevents "jumping," 6-F introducer 1,400–2,100 Protégé Sulzer Intratherapeutics Nitinol Laser-cut 6–12 7 Flexible, minimal shortening, vessel conformability, contrast injection around stent, retaining ring prevents "jumping" 1,300–1,750 6-F Protégé Sulzer Intratherapeutics Nitinol Laser-cut 6–9 6 Same as Protégé but compatible with 0.018-inch guide wire and 6-F introducer 1,300–1,750 SMART Cordis Nitinol Laser-cut 5–14 7 Flexible, minimal shortening, vessel conformability, tantalum markers, large diameters 1,540–2,495 Precise (0.018-inch SMART) Cordis Nitinol Laser-cut 5–10 6 Same as SMART but compatible with 0.018-inch guide wire and 6-F sheath 1,795–1,950 SMART Control Cordis Nintinol Laser-cut 6–10 6 Same as SMART but compatible 6-F introducer 1,350–1,795 Symphony Boston Scientific/Medi-tech Nitinol Welded 6–14 7 More rigid then other nitinol stents, platinum markers, single-handed deployment, large diameters 1,212.64–1,321.84 Wallstent RP (Reduced Profile) Boston Scientific/Medi-tech Elgiloy Woven 5–24 6–12 Flexible, reconstrainable, large diameters, significant shortening 1,174.16–2,072.30 Wallstent Iliac RP Boston Scientific/Medi-tech Elgiloy Woven 6–10 6 Similar to Wallstent RP, 6-F introducer, greater radial force but more shortening than Wallstent RP, FDA-approved for iliac use 1,174.16–2,072.30 Monorail Wallstent (0.014-inch) Boston Scientific/Medi-tech Elgiloy Woven 6–10 5–6 Similar to Wallstent Iliac RP, very low profile, 5–6-F introducer, 0.014-inch guide wire monorail system 1,174.16–2,072.30 Zilver Cook Nitinol Laser-cut 6–10 7 Flexible, minimal shortening, vessel conformability, gold markers 1,250

Perhaps the most significant recent development in stent technology has been the dependable application of nitinol in the manufacture of stents. Although placement of a nitinol coil stent was first described by Dotter in 1983^ref, not until recently has nitinol been widely embraced as a stent material. Nitinol, an alloy of approximately 50% nickel and 50% titanium, is named after its composition and place of origin. Ni and Ti are the symbols for nickel and titanium and NOL stands for the Naval Ordinance Laboratories, where nitinol was developed. Whereas other self-expanding stents rely on their design to produce their self-expanding nature, nitinol stents exploit the unique properties of the metal alloy. The properties of nitinol that serve this purpose are "superelasticity" and "thermal shape memory," which are based on the ability of nitinol to exist in different crystal forms or phases^ref. Superelasticity causes a constrained stent to expand to its nominal shape when deployed and thermal shape memory allows a plastically deformed stent to return to its original shape when heated to its transition temperature. These properties also afford nitinol stents with the properties of crush resistance and vessel conformability. In general, self-expanding stents are more flexible than balloon-expanded stents and they tend to track better when advanced through tortuous anatomy. The self-expanding nature of these stents also provides a certain amount of crush resistance not possessed by most balloon-expanded stents. Because most self-expanding stents are made of inert, nonferromagnetic metals, they generally cause less susceptibility artifacts than stainless-steel stents on MR images. However, self-expanding stents tend to be less radiopaque than balloon-expanded stents.
• AVE Bridge SE (self-expanding) stent (Medtronic AVE) is a recent addition to the Medtronic AVE biliary-approved stents. It is one of a number of self-expanding stents laser-cut from a single tube of nitinol. As such, it benefits from the properties inherent to the metal alloy, including flexibility, good vessel wall apposition, ability to conform to tortuous anatomy, crush resistance, and minimal foreshortening. The design of the stent resembles that of the balloon-expanded modular AVE Bridge Flexible stent, although the construction is fundamentally different. The Bridge SE has a 12-crown design to the sinusoidal rings, with four of the 12 crowns connected to the adjacent sinusoid. The Bridge SE has a unique single-handed delivery mechanism, which involves pushing a deployment knob with the thumb and retracting the outer restraining catheter. The stent is compatible with 0.035-inch guide wires and 7-F introducer sheaths and is available in diameters as large as 10 mm and lengths as long as 60 mm.
• Dynalink Stent (Guidant) : like the AVE Bridge SE stent, the Dynalink stent is a laser-cut nitinol stent that demonstrates good flexibility, vessel wall conformability, crush resistance, and minimal foreshortening. It has similar cell geometry to the Multilink coronary stent and the Omnilink balloon-expanded stent, which are produced by the same company. A notable advantage of the Dynalink stent is that stents as large as 10 mm in diameter are compatible with a 6-F introducer sheath. A 7-F sheath is recommended for the 12-mm and 14-mm stents. All stents are available in lengths ranging from 28 mm to 100 mm on both 0.018- and 0.035-inch guide wire platforms. Like the AVE Bridge SE stent, the Dynalink stent has a unique deployment handle designed for single-handed use.
• Intracoil (Sulzer Intratherapeutics) : formerly the Vascucoil stent from Medtronic, the Intracoil has a unique design, which is essentially the same as the nitinol coil first described by Dotter (16). It is made from a single strand of nitinol wire wound into the shape of a simple coil. This design affords the stent extreme flexibility and a low metal surface–to–open area ratio. The stent coil is constrained onto its delivery catheter by sutures at the proximal and distal ends. The coil is released at the trailing end first and experiences significant foreshortening. Radiopaque markers on the delivery catheter indicate the ultimate expanded length and position of the stent. The Intracoil is available in diameters as large as 8 mm and lengths as long as 60 mm and requires 7–9-F introducer sheaths. The Intracoil has recently obtained FDA-approval for use in the superficial femoral and popliteal arteries.
• Luminexx Stent : Bard (C.R. Bard, Murray Hill, NJ) was the first company to produce a flexible, self-expanding nitinol stent, called the Memotherm stent. Subsequent improved models include the Flexx stent, followed by the Luminexx stent. The fundamental design of the stent has remained unchanged, consisting of a zigzag pattern. However, the original stent struts have undergone electropolishing to render more rounded edges. Like the Wallstent, the Bard stents have flared ends to optimize anchoring of the stent to the vessel wall. The only change made to the stent in its transition from the Flexx model to the Luminexx model was the placement of four spoon-shaped tantalum markers at each end of the stent to improve radiopacity. The mode of deployment is relatively simple and involves retraction of the outer membrane while holding the inner catheter in place ("unsheathing"). Improvements in the delivery system of the Luminexx model include a tapered inner catheter on which the stent is mounted that is designed to prevent imprecise placement related to the stent "squirting forward" during the final stages of deployment. The Luminexx stent is available in a wide variety of sizes with diameters ranging from 6 mm to 12 mm and lengths ranging from 20 mm to 120 mm. The newest version of the delivery catheter is compatible with a 6-F introducer sheath in all sizes.
• Protegé Stent (Sulzer Intratherapeutics) is a flexible nitinol stent with a similar open cell design to the balloon-expanded Intrastent produced by the same company. It possesses the qualities of flexibility and vessel conformability inherent to nitinol. As with most nitinol stents, there is minimal foreshortening on deployment. Although the mechanism of deployment is similar to that of many other self-expanding stents ("unsheathing"), the Protegé stent does feature a retaining ring at the trailing end designed to prevent longitudinal motion ("squirting forward") of the stent during its deployment, similar to the Luminexx stent. The delivery catheter also has longitudinal rails to reduce the stored longitudinal forces that may accumulate during the stent deployment process. A side port on the delivery catheter is designed to allow contrast material injection around the stent to check its position before its deployment. The Protegé stent is available in sizes as large as 12 mm in diameter by 80 mm in length. The 6–8-mm stents are also available in a 150 mm length. All sizes require a 0.035-inch guide wire and 7-F delivery system. However, a recently developed 6-F version of the Protegé is available in diameters ranging from 6 mm to 9 mm and lengths as long as 80 mm. This stent is identical to the original Protegé but is compatible with a 0.018-inch guidewire and 6-F introducer sheath.
• Shape Memory Alloy Recoverable Technology (SMART) Stent (Cordis) : like the AVE Bridge SE, Dynalink, Luminexx, and Protegé stents, the SMART stent is a laser-cut nitinol stent with good flexibility, vessel wall conformability, and crush resistance. The design of the stent results in only minimal foreshortening on deployment. However, like most nitinol stents, it has poor radiopacity. To counteract this disadvantage, small tantalum markers have been placed on the ends of the stent to aid precise placement. The deployment mechanism of the stent is similar to those of most self-expanding stents and involves an "unsheathing" maneuver. However, although the name of the stent may suggest otherwise, the SMART stent is not reconstrainable. The SMART stent is available in sizes as large as 14 mm in diameter and 80 mm in length and can be advanced through a 7-F introducer sheath over a 0.035-inch guide wire regardless of size. Newer versions of the SMART stent include the Precise and the SMART Control, both of which are compatible with a 6-F introducer sheath. The Precise stent is designed for a 0.018-inch or 0.014-inch guidewire system and is available in sizes as large as 10 mm (diameter) by 40 mm (length). The SMART Control will likely replace the original SMART stent and is currently available in 6–10 mm diameters and lengths ranging from 20 mm to 120 mm. Like the AVE Bridge SE and Dynalink, the SMART Control has a deployment handle designed for single-handed use.
• Symphony Stent (Boston Scientific/Medi-tech) is a self-expanding stent made from a single strand of nitinol wire arranged in a hexagonal cell design and welded at the contact points. This design gives the Symphony stent a very low metal surface–to–open area ratio but also renders it somewhat more rigid than most nitinol stents, such that it has the tendency to straighten tortuous vessels. The stent foreshortens less than the Wallstent on deployment, but it is also significantly less radiopaque. To aid visualization during placement, three platinum markers have been welded to either end of the stent. The Symphony stent is available in a wide variety of diameters ranging from 6 mm to 14 mm and in lengths as long as 60 mm. A pistol-like delivery system allows single-handed deployment of the stent.
• Wallstent Endoprosthesis (Boston Scientific/Medi-tech) is the only self-expanding stent—and one of only two stents—FDA-approved for use in the iliac arteries. It is made of Elgiloy, a "superalloy" combining cobalt, chromium, nickel, and other metals. Elgiloy contains a relatively small amount of iron and is therefore nonferromagnetic and MR-imaging–compatible. A platinum core renders the stent struts radiopaque. The woven mesh design of the struts imparts flexibility as well as an outward self-expanding force to the stent. Like the Luminexx stent, there is a subtle increase in the diameter at the ends of the stent, giving it a flared appearance. This design feature is intended to improve anchoring of the stent to the vessel wall. Because the stent has a woven design, it is considerably longer when constrained within the delivery system than it is in its expanded state. Consequently, the stent foreshortens significantly on deployment, making precise placement of the trailing end somewhat challenging. However, it is the only stent that can be completely recaptured or reconstrained into its housing after more than 80% of the stent has been deployed. The Wallstent is available in a wide range of diameters from 5 mm to 24 mm and lengths from 20 mm to 94 mm and requires introducer sheaths ranging from 6-F to 11-F. For iliac sizes (6–10 mm), there is a lower-profile version of the Wallstent (Iliac Wallstent), which can be advanced through a 6-F introducer sheath. The lower profile is achieved by increasing the angle between the stent struts from 110° on the regular Wallstent to 140° on the Iliac Wallstent. As a result, the iliac Wallstent has greater radial force and hoop strength than the regular Wallstent, but it also foreshortens to a larger degree when deployed. A new version of the Iliac Wallstent has an even lower profile and is housed within a 5–5.9-F catheter. This stent, which is available in diameters ranging from 6 mm to 10 mm and lengths as long as 37 mm, is designed for use over a 0.014-inch guide wire in monorail fashion.
• Zilver Stent (Cook, Bloomington, IN) is the most recent addition to the increasing number of nitinol stents. The design of the Zilver stent is similar to the large stainless-steel Z-stent manufactured by the same company and consists of interconnected rows of "Z"s. This design results in no appreciable foreshortening on deployment of the stent. As with other laser-cut nitinol stents, the Zilver stent exhibits flexibility, vessel conformability, and crush resistance. The stent has four gold markers attached to either end to improve visibility, similar to the SMART, Symphony, and Luminexx stents. The Zilver stent is available in lengths ranging from 20 mm to 80 mm and diameters ranging from 6 mm to 10 mm, all of which are delivered through a 7-F introducer sheath.
• covered stents are composite devices consisting of a metallic stent skeleton covered with synthetic graft material. These devices have the ability to exclude plaque and endothelium, thereby offering the potential of mitigating late luminal loss by halting migration and proliferation of cells through open stent struts. No covered stents are currently FDA-approved for application in the iliac arteries for occlusive disease. However, "homemade" covered stents, iliac limbs from modular aortic stent grafts, and commercially manufactured covered stents approved for other nonvascular indications have been placed in the iliac arteries. The most common indications include treatment of aneurysms, ruptures, and arteriovenous fistulas. The graft material of a covered stent is attached to either the internal or external surface of the stent or, alternatively, it can be "sandwiched" between two stents. Most covered stents are of the self-expanding variety, but covered balloon-expanded stents do exist. Because of the bulk of the additional graft material, covered stents require larger delivery systems than uncovered stents. This may expose the patient to a higher risk of groin complications. Covered stents are used to treat aneurysms, hemorrhages, and for TIPS :
• internally covered stents
• externally covered stents
• sandwich covered stents
• polytetrafluoroethylene (PTFE)-covered stents
• Dacron-covered nitinol stents
• polyurethane
• macroporous polyester-covered stents

stent  vendor  stent type  stent material  graft material  diameter (mm)  introducer size (F)  key features  list price ($)  aSpire Vascular Architects self-expanding nitinol PTFE 6–12 8 only 50% covered, maintains branch vessel patency, reconstrainable, approved for tracheobronchial use 1,950–2,050 Cragg Endopro System 1 Boston Scientific/Medi-tech self-expanding nitinol Dacron 6–12 8.5–10 first commercial covered stent, not available in US NA Hemobahn (Viabahn) WL Gore and Associates self-expanding nitinol PTFE 6–13 8–14 flexible, minimal shortening, vessel conformability, Viabahn approved for tracheobronchial use 1,995–2,759 (Viabahn) Jostent Jomed balloon-expanded 316L stainless steel PTFE 4–12 7–8 flexible balloon-expanded stent, precise placement, hoop strength, only coronary version available in US NA Wallgraft Boston Scientific/Medi-tech self-expanding Elgiloy Polyethylene terephthalate 6–14 9–11 similar to Wallstent, reconstrainable, approved for tracheobronchial use 1,950–2,250

• aSpire Covered Stent (Vascular Architects, San Jose, CA) has a unique design. It is comprised of a double-helical nitinol frame molded in a cylindrical coiled fashion and covered with polytetrafluoroethylene (PTFE). The helical turns are 5 mm wide, with 5-mm bare gaps between turns. Therefore, the stent is not a classic covered stent insofar as it cannot be used to seal arteriovenous fistulas or vessel ruptures. Rather, the goal of this unique construction is to allow for variable vessel wall coverage to promote endothelial ingrowth while potentially maintaining branch or collateral vessel patency. The stent skeleton is etched from a sheet of nitinol and subsequently annealed into the shape of a helical coil. An ultrathin PTFE tube is then pulled over the skeleton along the length of the coil so the final product has no exposed metal. The aSpire design results in good flexibility, radial force, and hoop strength. The delivery mechanism is also unique and involves rotating the outer catheter of the delivery system over the inner catheter, each of which are attached to either end of the stent by nitinol cables. The stent foreshortens when released, but is not fully deployed until the cables are pulled. Therefore, the stent can be reconstrained by rotating the catheters in the opposite direction and repositioned if necessary. The stent is approved in the United States for tracheobronchial use and is available in diameters ranging from 6 mm to 12 mm and lengths of 25 mm, 50 mm, and 100 mm. The 25-mm stents have a coil width of 2.5 mm without any gaps between the helical turns. All stents are compatible with a 0.035-inch guide wire system and an 8-F introducer sheath.


• Cragg Endopro System 1/Passager Stent Graft (Boston Scientific/Medi-tech) was the first covered stent to be marketed. It is a self-expanding stent made of nitinol, which is covered with thin, woven polyester fabric (Dacron). The stent lattice consists of nitinol wire fashioned into a zigzag configuration with 7–0 polypropylene ligatures. Although his stent has been marketed in Europe for several years (18), it is not available in the United States. In Europe, the stent is available in lengths as long as 100 mm and diameters as large as 12 mm.
• Hemobahn Endoprosthesis (WL Gore and Associates, Flagstaff, AZ) is a self-expanding covered stent consisting of a nitinol skeleton covered on the luminal surface with ultrathin PTFE. The stent skeleton is constructed from a single strand of nitinol wire formed into a sinusoidal shape and wound in helical fashion around the PTFE tube. The device is deployed from hub to tip by releasing a single restraining suture line (19). There is minimal foreshortening of the stent on deployment. The Hemobahn Endoprosthesis is available in diameters ranging from 6 mm to 13 mm and lengths of 50 mm, 100 mm, and 150 mm. There is no 12-mm stent and the 10-, 11-, and 13-mm stents are not available in the 150-mm length. Recommended introducer sheath sizes range from 8 F to 14 F and the system is compatible with a 0.025-inch guide wire. In the US, the Hemobahn stent is being marketed as the Viabahn stent, which has recently received FDA approval for tracheobronchial use.
• Jostent Stent-Graft (Jomed, Helsinborg, Sweden) is the only commercially manufactured balloon-expanded covered stent. As such, it exhibits the advantages of balloon-expanded technology, including radial strength and ease of precise placement. The device consists of ultrathin PTFE graft material sandwiched between two low-profile flexible 316L stainless-steel stents. There is minimal foreshortening on deployment, the mode of which is the same as that of other balloon-expanded stents. Peripheral and coronary versions of the stent graft are manufactured. However, only the coronary version is available in the United States and is approved for the treatment of coronary artery perforations. The recommended maximum expansion diameter of the coronary stent graft is 5 mm. The peripheral covered stent is available in Europe in diameters ranging from 4 mm to 12 mm and lengths as long as 58 mm.
• Wallgraft Endoprosthesis (Boston Scientific/Medi-tech) consists of a Wallstent covered with polyethylene terephthalate. The polyethylene terephthalate is attached to the extraluminal surface of the stent with polycarbonate urethane adhesive. A spiral platinum tracer wire enables differentiation of the Wallgraft from the Wallstent fluoroscopically. Otherwise, the only structural difference between the metallic portion of the Wallgraft and the Wallstent is the lack of flaring at the ends of the Wallgraft. It is therefore important to oversize the Wallgraft by 1–2 mm relative to the diameter of the target vessel. The mechanism of deployment is identical to that of the Wallstent, including the ability to reconstrain the stent when as much as 80% of the stent has been released. The Wallgraft is approved in the United States for tracheobronchial use and is available in diameters as large as 14 mm and lengths as long as 70 mm. These devices require 9–11-F introducer sheaths.
Properties of stents :a variety of specific mechanical properties germane to stent technology bear consideration when selecting a stent for a particular lesion. Although these factors may be considered crucial for the ultimate success of endovascular procedures, there are surprisingly few objective data comparing the properties of the various stents, and differences in experimental technique render comparisons between studies difficult. Moreover, with continuous developments and improvements in the stent marketplace, research based on specific devices is at risk of becoming outdated and obsolete. Nonetheless, several investigators have performed comparative in vitro experiments^{ref1, ref2, ref3, ref4} on a variety of available stents in an effort to characterize their features. These studies form the basis of the following synopsis.
• radial force/hoop strength : the radial force of a stent represents its ability to efface the arterial lumen despite the compressive force exerted by a stenotic lesion on the stent. Radial force is dependent on a variety of factors, including stent construction, design, material, mass, and size. Hoop strength is the ability of a stent to resist radial compression by external forces. Balloon-expanded stents and self-expanding stents respond in fundamentally different ways to external compression^ref. When an increasing compressive force is applied to a balloon-expanded stent, it initially demonstrates mild elastic deformation. During this phase, the stent will return to its nominal diameter when the external force is removed. However, when the compressive force reaches the yield point, the stent will undergo irreversible plastic deformation and become crushed. Self-expanding stents, conversely, demonstrate a near-linear stress–strain relationship. Therefore, they behave in an elastic manner through a large range of external forces. However, as demonstrated by Lossef et al^ref, the force required to deform self-expanding stents is generally lower than that required to deform balloon-expanded stents. This finding was corroborated by Dyet et al^ref, who compared the radial force of three balloon-expandable stents (including Palmaz and AVE Bridge Flexible stents) and four self-expanding stents (Wallstent, Instent Vascucoil [equivalent to the Intracoil], Memotherm [old version of the current Luminexx], and Symphony). The Palmaz and AVE Bridge stents exhibited high degrees of hoop strength, which were comparable and significantly higher than those of the self-expanding stents. Of the evaluated self-expanding stents, the Wallstent demonstrated the greatest hoop strength, followed by the Symphony stent, the Instent Vascucoil, and the Memotherm stent. Interestingly, both Lossef et al^ref and Dyet et al^ref found that the Wallstent had significantly greater hoop strength when the ends of the stent were constrained. Lossef et al^ref also discovered that the medium Palmaz stent had 30% greater hoop strength than the large Palmaz stent when both were expanded to a diameter of 8 mm. This can be explained by the slightly thicker stent strut of the medium Palmaz stent or by the fact that the struts of the medium Palmaz stent are oriented more radially than those of the large Palmaz stent at a diameter of 8 mm. Another interesting finding in that study^refwas that overlapping two stents, or placing a stent within a stent, effectively doubles the hoop strength in the overlapped segment. This finding applied to the Palmaz stent as well as the Wallstent^ref. A balloon-expanded stent will experience permanent plastic deformation if the external force exceeds its maximum hoop strength. Duda et al^ref compared the hoop strengths of various balloon-expanded stents (Palmaz medium, Palmaz large, Palmaz-Schatz long-medium, AVE Bridge Flexible, and rigid AVE Iliac Bridge stents). The stent with the greatest hoop strength was the AVE Iliac Bridge stent. The hoop strength of the Palmaz medium stent was found to be comparable to that of the AVE Bridge Flexible stent and significantly greater than those of the Palmaz large and Palmaz-Schatz long-medium stents. When comparing hoop strength per unit mass, the Palmaz medium stent was superior to the AVE Bridge Flexible stent and comparable to the rigid AVE Iliac Bridge stent. To better evaluate the radial force of self-expanding stents, Duda et al^ref compared the "chronic outward force" and "radial resistive force" of a variety of self-expanding stents, including the Wallstent, SMART stent, Symphony stent, and Memotherm stent (old version of the current Luminexx stent). Chronic outward force refers to the force exerted on a vessel by a self-expanding stent as it expands to its nominal diameter. Radial resistive force is a measure of the force exerted by the stent as it resists external constriction. Contrary to the results of Dyet et al^ref, Duda et al^ref found the Wallstent to have a lower radial resistive force (hoop strength) and chronic outward force (radial force) than the nitinol stents. Of the evaluated nitinol stents, the SMART stent had the greatest force in both categories, followed by the Symphony and Memotherm stents^ref. The discrepancy in results between the two studies is most probably explained by the different radial strength exhibited by the Wallstent when its ends are constrained compared with when it is allowed to lengthen with extrinsic compression.
• flexibility : Dyet et al^ref compared the flexibility of various stents by fixing one end of each stent and measuring the force required to bend the stents 10°. The Palmaz stent was found to be the most rigid stent. Surprisingly, the AVE Bridge Flexible stent was more flexible than the Symphony and Memotherm stents, albeit in only one plane. Unlike the self-expanding stents, the AVE Bridge Flexible stent did not return to the original straight position when the bending force was removed. The Instent Vascucoil was the stent that required the least force to bend 10°, followed by the Wallstent.
• "trackability" or "pushability" refers to the ease with which a stent, on its delivery system, can negotiate an angle such as the aortic bifurcation. Apart from the flexibility of the stent, this feature is also dependent on the properties of the delivery catheter. Therefore, the trackability of an unmounted balloon-expanded stent can be improved by mounting the stent on a more flexible balloon catheter. Dyet et al^ref and Duda et al^ref analyzed the trackability of standard stent delivery systems by assessing their ability to be advanced across an in vitro aortic bifurcation model. In the study by Dyet et al^ref, the Palmaz stent could not be advanced across 90° or 60° angles. Surprisingly, the Memotherm stent was unable to cross the 60° angle, although it did successfully negotiate the 90° angle. All other evaluated stents successfully crossed both angles. Duda et al^ref performed a quantitative analysis of the force required to advance stents on delivery catheters across an aortic bifurcation model. The SMART stent demonstrated the greatest trackability. Surprisingly, the AVE Bridge Flexible stent demonstrated better trackability than all evaluated self-expanding stents. The Palmaz-Schatz Long-Medium stent displayed comparable trackability to the Symphony stent. The Memotherm stent required the greatest force to negotiate the bifurcation, followed by the Wallstent. It must be noted that the delivery catheter of the new Luminexx stent has improved trackability compared with the delivery system of the old Memotherm stent.
• radiopacity : again, the studies by Duda^ref, Dyet^ref, and their coworkers represent the only systematic analyses of the radiopacity of various stents. Duda et al^ref placed the stents into a water-filled plexiglas phantom and acquired radiographic images with use of a digital subtraction angiography unit. Stent attenuation was assessed digitally as a gray-scale value. The results of this study are somewhat perplexing. The stents with the greatest radiopacity were the Palmaz large stent, the Palmaz-Schatz long-medium stent, the rigid AVE Iliac Bridge stent, and the Symphony stent. Although the three balloon-expanded stents are the stents with the greatest mass of stainless steel, the inclusion in this group of the Symphony stent, which is made of nitinol, is unexpected. The SMART stent and the Wallstent were found to have intermediate radiopacity. The group of stents with the lowest radiopacity included the Palmaz medium, AVE Bridge Flexible, and Memotherm stents. The poor performance of the Palmaz and AVE Bridge Flexible stents in this experiment is somewhat surprising. In the study by Dyet et al^ref, radiopacity was measured as the amount of aluminum required to render a stent invisible on fluoroscopy. The Palmaz stent, AVE Bridge Flexible stent, and Instent Vascucoil required the thickest slabs of aluminum and were therefore considered the most radiopaque. The Wallstent was slightly less radiopaque, followed by the Memotherm stent. Contrary to the findings of Duda et al^ref, the Symphony stent was significantly less radiopaque then all other evaluated stents. The results of the report of Dyet et al^ref are more in keeping with the experience of the authors with regard to the radiopacity of these stents. Although not included in any systematic analysis, the Strecker and Vistaflex stents, made of tantalum and platinum, respectively, exhibit the greatest radiopacity of all available stents.
• MRI compatibility : with MRA evolving into a widely used noninvasive imaging modality for the assessment of vascular disease, the ability to assess the patency of metallic stents with MR is becoming increasingly important. Many metallic stents are touted as being MRI–compatible because they are made of inert nonferromagnetic material. These include the nitinol stents as well as other non–stainless-steel stents such as the Wallstent (Elgiloy), the Vistaflex stent (platinum), and the Strecker stent (tantalum). Because these stents do not experience significant torque when placed in a strong magnetic field, patients can undergo MRI examinations in the immediate postimplantation period. MRI examinations in patients with recently placed stainless-steel stents that have not undergone endothelialization are generally not recommended because of the risk of stent translation and/or torque. Nonetheless, patients with recently implanted stainless-steel stents have undergone successful MRI without undue complications^ref. Nonferromagnetic stents cause significantly less artifact on MRA images than stainless-steel stents. However, this does not necessarily translate into the ability to evaluate for stent patency or in-stent restenosis. The artifactual signal loss on MRI in and around metallic stents is the result of 2 different mechanisms. First, susceptibility artifacts brought about by metal-induced local magnetic field inhomogeneities cause signal loss around the stent struts. Second, radiofrequency artifacts related to the induction of eddy currents in the conductive metallic material cause a reduction of signal within the stent lumen. The extent of the artifact produced by a stent is further dependent on a host of other factors, including the type of scan sequence used, the orientation of the stent with respect to the magnetic field, the strength of the magnetic field, the imaging bandwidth, and the material and geometry of the stent. A detailed discussion of these factors goes beyond the scope of this review. Nevertheless, the results of several studies evaluating MRI artifacts of various stents do bear mentioning. Many investigators have performed in vitro tests comparing the appearances of a multitude of stents and stent-grafts on gadolinium-enhanced 3D-MRA^{ref1, ref2, ref3, ref4, ref5, ref6}. Not surprisingly, stainless-steel stents were generally associated with the largest artifacts. The artifacts produced by all stents were heavily dependent on factors such as orientation to and strength of the magnetic field^ref. Nonferromagnetic stents did not perform uniformly across the experimental studies and their artifacts were found to be influenced by stent geometry and strut orientation^ref. The performance of nitinol stents (and nitinol-based stent-grafts) was variable but generally good, with most authors concluding that stent patency was likely assessable despite some luminal signal loss. The notable exception was the Symphony stent, which consistently produced greater artifacts than other nitinol stents. The tantalum Strecker stent was found to be well-suited to MRI because it induces minimal artifact^{ref1, ref2}. Although the Wallstent contains only a very small amount of iron, most studies found that it caused significant artifacts, thereby precluding evaluation of luminal patency. Good results have been reported with respect to the MR characteristics of some nitinol covered stents including the Cragg Endopro^ref and the Hemobahn^ref. However, like in most studies, stents were evaluated for luminal patency rather than in-stent restenosis. Ultimately, further clinical trials are needed to determine which stents are appropriate for MRA follow-up examination. Although preliminary findings with regard to the performance of the platinum Vistaflex stent are promising, its level of MRI compatibility also remains to be seen.
Indications for iliac artery stent placement : treatment of stenotic or occluded iliac arteries is primarily indicated in symptomatic patients. The main clinical indications for percutaneous endovascular treatment are as follows: lifestyle-limiting claudication, rest pain, limb-threatening ischemia or tissue loss, insufficient inflow into an extremity before bypass grafting, amputation (to lower the level of amputation or promote healing) or renal transplantation, and vasculogenic impotence. The best results of PTA are usually obtained in patients with focal or short lesions, ie, Category 1 and Category 2 lesions as determined by the American Heart Association Guidelines^ref. Stent placement has its greatest benefit in patients with long lesions and occlusions (Categories 3 and 4). A stenosis is considered hemodynamically significant if the diameter of the vessel lumen is reduced by 50%–60% (Berne FA, Lawrence WP, Carlton WH. Roentgenographic measurements of arterial narrowing. AJR Am J Roentgenol 1970; 110: 757–759; May AG, Van de Berg L, DeWeese J, Rob CG. Critical arterial stenosis. Surgery 1963; 54: 250–259). Occasionally, it is difficult to determine the significance of a lesion based on its angiographic appearance. In these situations, a translesion pressure gradient can help determine the hemodynamic significance of a lesion^ref. A peak-to-peak systolic pressure gradient of 10 mm Hg or greater at rest is considered hemodynamically significant. A translesion pressure gradient greater than 15 mm Hg after administration of a vasodilator (100 µg nitroglycerin or 25 mg tolazoline intraarterially) is also considered significant^{ref1, ref2}. The latter test is performed to simulate exercise and can be positive even in the absence of a significant gradient at rest^ref. Other definitions of hemodynamic significance include a resting mean translesion gradient greater then 5 mm Hg, which is the value adopted by the FDA, and a mean translesion gradient greater than 10 mm Hg after augmentation with a vasodilator^{ref1, ref2}. Guidelines for iliac PTA established by the joint council of the American Heart Association^ref
• category 1 : noncalcified, concentric stenosis < 3 cm in length
• category 2 : stenosis 3–5 cm in length that are calcified, or eccentric stenosis < 3 cm in length
• category 3 : stenosis 5–10 cm in length, or chronic occlusions < 5 cm in length after thrombolytic therapy
• category 4 :
• stenosis > 10 cm in length
•  chronic occlusions over 5 cm in length after thrombolytic therapy
•  extensive aortoiliac disease
•  iliac stenosis in a patient with an abdominal aortic aneurysm or other lesion requiring aortic or iliac surgery
The main indication for iliac artery stent placement in which it has been shown to significantly improve results is in the initial period after failed or inadequate balloon angioplasty as a result of elastic recoil or a flow-limiting dissection. Significant residual stenosis after PTA is defined as greater than 30% residual diameter narrowing and/or residual peak-to-peak systolic pressure gradient of > 5 mm Hg^{ref1, ref2}. Primary stent placement is also generally viewed to be indicated in patients with chronic iliac artery occlusions and in patients with recurrent iliac stenosis after previous PTA, although there is little evidence to support the latter approach. The efficacy of the use of stents to treat severely ulcerated plaques and lesions deemed at high risk for distal embolization is also as yet unproven. The term "primary stenting" has been used in the medical literature to describe both the practice of placing a stent without predilation and that of placing a stent after predilation of the lesion regardless of the outcome of PTA. The approach without predilation, also referred to as "direct stenting," has been advocated as a way to reduce the risk of peripheral embolization when treating chronic occlusions^{ref1, ref2}. However, these findings have not been corroborated by other studies in which occlusions were primarily treated with stent placement after predilation^{ref1, ref2, ref3}. Nonetheless, because of the significantly higher rate of embolization when treating occlusive lesions than when treating stenotic lesions, many consider direct stent placement without predilation to be the method of choice for the treatment of iliac artery occlusions. Other potential advantages of direct stent placement include the shorter duration of the procedure and the reduced radiation exposure. Conversely, predilation facilitates proper stent selection in terms of providing important information about the size of the vessel as well as the nature of the lesion (soft vs fibrous). Irrespective of the method of primary stent placement, some authors advocate its routine use as a method to treat all iliac artery lesions. This may be an appealing approach because stent placement is almost always associated with a good angiographic result with restoration of laminar flow. However, like angioplasty, stents induce intimal hyperplasia. Therefore, the acute luminal gain achieved by stent placement is not necessarily maintained in the long term. Indeed, no published study has definitively demonstrated a reduced rate of restenosis or superior long-term patency rates of iliac stent placement compared to PTA. Bosch and Hunink performed a meta-analysis of 14 studies that include more than 2,000 patients, comparing the results of iliac artery stent placement to balloon angioplasty alone^ref. The initial technical success rate was higher for stent placement (96%) than for balloon angioplasty (91%), but the difference was not statistically significant. Complication and mortality rates were also not significantly different. Four-year patency rates, stratified for stenoses versus occlusions and claudication versus critical ischemia, were all higher for the stent group compared with the angioplasty group. The authors concluded that the risk of long-term failure was reduced by 39% after stent placement compared to PTA. However, in a prospective randomized multicenter trial of 279 patients who underwent iliac artery intervention, Tetteroo and coworkers^ref failed to demonstrate a significant difference between primary stent placement and PTA with selective stent placement. In the group that underwent PTA, 43% of the patients ultimately received a stent as a result of a persistent mean translesion pressure gradient of 10 mm Hg or more. Two-year patency rates by duplex sonography were 71.3% in the primary stent group and 69.9% in the selective stent group. Moreover, a cost-analysis study performed by Bosch and coworkers^ref comparing PTA alone, PTA with selective stent placement, and primary stent placement found selective stent placement to be the most cost-effective approach. However, the majority of patients in this series were treated for claudication. Therefore, in patients with critical ischemia, this conclusion may or may not be applicable. In addition, cost savings applicable to shorter procedural time with primary stent placement were not taken into account. In summary, the patency rates of iliac artery stent placement are comparable if not better than those of PTA alone. However, evidence suggests that the added cost and potential risk of primary stent placement is probably not justified and that angioplasty followed by stent placement for failures may be the preferred approach for stenotic iliac artery lesions. Nonetheless, stent placement has expanded the range of lesions amenable to percutaneous endovascular therapy by including occlusions, long lesions, and lesions that are diffuse, eccentric, or ulcerated. Primary stent placement in these cases may even reduce the rate of complications.
Results of iliac artery stent placement : numerous studies have demonstrated the safety and efficacy of iliac artery stent placement^{ref1, ref2, ref3, ref4, ref5, ref6, ref7, ref8, ref9, ref10, ref11, ref12, ref13}. Although direct comparison of the studies is difficult because of variations in patient demographics, indications for treatment and extent of disease, several studies have adhered to the uniform reporting standards set forth by Rutherford and Becker^ref. Murphy^ref (45) summarized the results of 18 published studies of iliac artery stent placement for occlusive disease in a total of 2,058 limbs, including seven series which had reported patency according to accepted reporting standards. Stents used included the Palmaz stent, the Wallstent, and the Strecker stent. The weighted average technical success rate was 97%, with a weighted average complication rate of 6%. Primary and secondary 5-year patency rates ranged from 58% to 82% and 78% to 91%, respectively, with weighted averages for primary and secondary 5-year patency of 73% and 85%, respectively. These results approach those of surgical aortoiliac bypass grafting. No randomized prospective trial has been published comparing the results of different stent types in the iliac arteries, and the short and long-term results of the existent studies involving individual stent types appear roughly equivalent. In the meta-analysis performed by Bosch and Hunink^ref (41), 4-year primary patency rates including technical failures were 80% for the Palmaz stent, 70% for the Wallstent, and 64% for the Strecker stent, suggesting the possible superiority of the Palmaz stent. However, in a nonrandomized study with use of a logistic regression analysis, Ballard et al^ref (46) failed to demonstrate a significant difference between the Palmaz stent and the Wallstent, despite the fact that the Wallstent had been preferentially placed in longer lesions. Several investigators have identified prognostic factors that may influence the outcome of iliac artery stent placement. Unlike for PTA, most investigators have found no difference in the long-term patency of stents in the common iliac artery versus the external iliac artery^{ref1, ref2}. However, a recent study of 247 iliac stent placement procedures with use of a logistic regression analysis of a host of variables did identify external iliac artery stent placement as an independent predictor of a significantly reduced primary patency in women but not in men, even though the average size of the external iliac artery was not significantly different between the sexes^ref. In general, the patency of combined common and external iliac artery stent placement is lower than that of single-artery stent placement^ref. Laborde et al^ref demonstrated poorer results in patients with limb-threatening ischemia than in those with claudication. However, contradictory findings were reported by Murphy et al^ref, who demonstrated better results in patients with critical ischemia. Patency rates stratified according to the status of runoff vessels (patency of the superficial femoral artery) have also generated conflicting results, but most authors have found the presence of a diseased or occluded runoff vessel to be a risk factor for an increased rate of stent failure^{ref1, ref2}. Overall, very few factors have been identified as consistent predictors of the outcome of iliac artery stents. The use of covered stents to treat occlusive disease of the iliac arteries has been proposed as a method to reduce intimal hyperplasia and improve patency rates^ref. Although the use of stent-grafts to treat aortoiliac aneurysms is gaining acceptance, data on the treatment of iliac occlusive disease with covered stents is limited. Henry et al^ref placed the Cragg EndoPro System 1 covered stent into the iliac arteries of 38 patients, achieving a technical success rate of 97% and primary and secondary patency rates of 97% and 100%, respectively^ref. The results of a prospective multicenter trial of the Hemobahn covered stent for the treatment of occlusive disease of the iliac and femoral arteries were recently published^ref. The iliac group comprised 61 limbs in 53 patients with symptomatic stenotic or occlusive lesions. Placement of the Hemobahn endoprosthesis was successful in all cases, yielding an initial technical success rate of 100%. One covered stent underwent early thrombosis within 30 days of implantation. A further 5 stents had late restenosis or reocclusion, resulting in 1-year primary and secondary patency rates of 91% and 95%, respectively. Although these results are promising, there is, at present, insufficient evidence to support widespread application of covered stents for occlusive iliac artery disease, and their use should be reserved for the treatment of aneurysms, ruptures, and arteriovenous fistulas^ref.
Selection of stents : there are several important considerations when selecting a stent for an iliac artery lesion. First, as mentioned earlier, only 2 stents, the Iliac Wallstent and the Palmaz 308, are approved for iliac artery indications. The implantation of any other stent into iliac arteries represents an "off-label" use of the device. Second, the only stents for which long-term patency data exists are the Wallstent, Palmaz stent, and Strecker stent. Although these three devices represent the most diverse stents in terms of design and mechanical properties, their long-term patency rates are comparable. The equivalence of these results suggests that the choice of stent may not be critical for the success of the procedure. Ultimately, only a prospective randomized trial can definitively show the superiority of any particular stent. As a result of the rapidly changing technology related to stents, it is unlikely such a trial could be completed before the stent technology used in the study is viewed as outdated by the interventional community. Nonetheless, some generalizations can be made regarding the suitability of specific stents for particular types of iliac artery lesions. Apart from lesion characteristics (ie, calcified, fibrous, soft, eccentric, concentric, focal, diffuse, etc), factors that influence the choice of stent include access site, introducer size, vessel tortuosity, and lesion location. Unavoidably, operator experience, familiarity with particular products, ease of deployment, availability, and cost also play into the selection of stents. Short focal lesions can be treated with balloon-expanded or self-expanding stents. Occlusions or long lesions are best treated with self-expanding or long balloon-expanded stents. When treating a calcified, resistant, fibrous, or eccentric lesion, which is prone to elastic recoil, a balloon-expanded stent would be more appropriate than a self-expanding stent because of its better hoop strength. The rigid Palmaz stent is extremely well suited for this type of lesion. Based on the available evidence, the AVE Bridge Flexible stent may have comparable hoop strength and also offers some flexibility. Whether the newer generation of low-profile, flexible balloon-expanded stents retains the hoop strength of the more robust models remains to be seen. However, this newer generation of stents, including the Genesis, Intrastent, Omnilink, and Assurant (future replacement of the AVE Bridge Flexible), does offer advantages with respect to introducer size, flexibility, crossing profile, expansion ratio, and foreshortening characteristics. Calcified, resistant lesions are typically seen in the area of the aortic bifurcation. Such lesions are best treated by using the "kissing" stent technique^ref. Even if flow to only one limb is obstructed, "kissing" stents are usually placed to prevent compromise of the contralateral iliac artery ostium by plaque shifting. The disadvantage of this technique is that uncovered metallic stent struts protruding into the distal abdominal aorta may cause hemolysis or act as a nidus for thrombus formation^{ref1, ref2}. This effect can be somewhat reduced by placing self-expanding stents, which better appose the aortic wall. However, resistant plaque lesions are prone to elastic recoil and may require the added hoop strength of balloon-expanded stents^ref. In any event, the hemodynamics of "kissing" stents is poorly understood and it remains unclear what kind of stent would be most appropriate for this purpose. If access site and lesion location require traversal of a tortuous arterial segment, or crossing the aortic bifurcation, a flexible balloon-expanded stent or a self-expanding stent should be used. If stent placement across a joint is unavoidable, a self-expanding stent should be employed because it is more crush-resistant and is less likely to experience plastic deformation or fracture. Regardless of flexibility, placement of any stainless-steel balloon-expanded stent will result in some degree of straightening when placed into a tortuous vessel. Depending on the severity of tortuosity and the length of the segment with the stent, this can result in angulation or kinking of the vessel at the end of the stent. To avoid such an occurrence, a stent with good flexibility and vessel conformability should be employed. Self-expanding segmented nitinol stents such as the SMART, Luminexx, Protegé, Dynalink, Bridge SE, and Zilver stents, which are laser-cut from a single nitinol tube, are appropriate for this purpose because they expand in a segmental fashion, thereby apposing all aspects of the vessel wall and conforming to preexisting anatomy. This feature also renders this group of stents suitable for placement into vascular segments that transition abruptly in size; for example, from a common iliac artery measuring 8 mm to an external iliac artery measuring 6 mm. Because of its nonsegmented construction, the Wallstent does not conform to all aspects of an irregularly shaped vessel, resulting in areas in which the metal may not come into contact with the vessel wall. Conversely, the woven construction of the Wallstent will render a smooth curve to a vessel with an abrupt angulation, whereas a segmented nitinol stent may appear "kinked" because of its ability to conform to the vessel. The Symphony stent, which is also a self-expanding nitinol stent, does not possess the same ability to conform to the vessel as the segmented group of nitinol stents and tends to straighten arterial segments in which it is placed. This behavior is probably related to the construction of the stent, which consists of a welded nitinol filament. Presumably, the welded elements render the stent relatively rigid compared with segmented nitinol stents. The Vistaflex stent, conversely, is a woven, balloon-expanded stent that exhibits similar characteristics to the laser-cut nitinol stents. It is extremely flexible and demonstrates good vessel conformability. As such, it serves to bridge the gap between features associated with balloon-expanded stents and self-expanding stents. Whether these features come at the cost of reduced hoop strength compared with other balloon-expanded stents is as yet unclear. Balloon-expanded stents offer some advantage over self-expanding stents in terms of precision of placement. This is related to the mode of deployment as well as the foreshortening characteristics, which tend to be minimal and predictable with balloon-expanded stents. However, self-expanding nitinol stents do narrow the gap between balloon- and self-expandable stents in this regard because they demonstrate minimal foreshortening. Some of the nitinol stents, such as the Symphony, SMART, Luminnex, and Zilver stents, have radiopaque markers attached to the stent ends to improve stent visibility and thereby facilitate precise deployment.
Conclusion : the advent of intravascular stents has expanded the indications for percutaneous endoluminal therapies in the treatment of iliac artery occlusive disease. Although a multitude of stents exists, with varying designs and properties, no study has definitively shown the superiority of any particular stent. Reduction of in-stent restenosis is the subject of continued research. Whether promising new concepts have a durable impact on the long-term results of iliac stent placement remains to be seen. At present, there is insufficient data to advocate primary stent placement of nonocclusive iliac obstructions. Therefore, indications for stent placement remain largely those approved by the FDA. Covered stents have a role in sealing ruptures, aneurysms, and arteriovenous fistulas, but there is insufficient evidence to advocate their routine use for occlusive disease. Many factors deserve consideration during stent selection in different clinical situations, and knowledge of the individual characteristics of various stent types is crucial. Unfortunately, there is limited objective data on the comparative properties of the various stents. As the number of new devices increases, investigators must continue to objectively evaluate new stent technology so that unbiased performance data, along with clinical trials, help guide operator choices.
• self-expanding nitinol stent system with a balloon catheter (Wingspan with Gateway PTA, made by Boston Scientific) for improving cerebral artery lumen diameter in patients with intracranial atherosclerotic disease (ICAD) refractory to medical therapy. The device is intended for use in intracranial vessels with 50% or more stenosis that are accessible to the system.

Pathogenesis :
• neointima : in a blood vessel graft or vascular prosthesis, a new layer of endothelial cells on the intimal surface
• pseudointima : in a blood vessel graft or vascular prosthesis, a new layer on the intimal surface that consists of cells other than endothelial cells, such as plasma proteins or collagen


Prevention : drug-eluting stents (DESs) have rapidly become the standard of care in treating coronary artery disease through percutaneous intervention. The classification considers 4 broad criteria:
• performance :
• drug options : sirolimus and paclitaxel are very different drugs with distinct mechanisms of action. Sirolimus is a macrolide antibiotic, whereas paclitaxel is the most widely used chemotherapeutic agent in the world. Each drug affects a different stage of the cell cycle. Sirolimus arrests the cell before it enters the dividing cycle, whereas paclitaxel interrupts cell division during the mitosis phase. Therefore, given their characteristics, sirolimus is regarded as a cytostatic drug and paclitaxel as cytotoxic. Furthermore, sirolimus has a prominent anti-inflammatory effect (ie, it decreases levels of IL-6 and MCP-1 locally), which may account for a significant portion of its differential effect versus paclitaxel. Another important consideration in evaluating these 2 drugs is their action in terms of local and systemic toxicity. Regarding local toxicity, sirolimus has a large therapeutic window. Even at increased doses (as seen with placement of multiple and overlapping stents), the drug does not exhibit local toxicity, having no untoward effects on the vascular bed. Systemically, the dosing range of sirolimus is also very safe. Even if 17 stents were simultaneously placed in a patient’s coronary tree, the plasma levels of sirolimus would be within the therapeutic range for oral dosing of the drug. On the other hand, paclitaxel may exhibit significant local toxicity. Three to 4 times the dose present on a single stent may result in medial necrosis and hemorrhage. Thus, paclitaxel is associated with a narrower therapeutic window, and the amount of drug that can safely be placed on the stent is limited. This may have implications for efficacy, given that both the amount of neointimal proliferation inhibition and local toxicity are directly related to the drug dose. Rifampicin-medicated stents can prevent intimal hyperplasia. Currently, there are 2 DES platforms:
• sirolimus / rapamycin-eluting stent (CYPHER^®, Cordis Cardiology, Miami Lakes, FL)
• The XIENCE™ V everolimus-Eluting Coronary Stent System is manufactured by Abbott and distributed on a private label basis by Boston Scientific as the PROMUS Everolimus-Eluting Coronary Stent System. Boston Scientific is the first company to offer 2 distinct drug-eluting stent platforms.
• paclitaxel-eluting stent : at a dose of 2.7 µg per mm² it has been shown in clinical studies to inhibit restenosis by decreasing smooth muscle cell proliferation and migration, secretion of extracellular matrix, and inflammation.
• TAXUS™ Express2™ Paclitaxel-Eluting Coronary Stent System, Boston Scientific, Natick, MA
• Axxion^TM, Occam International BV, a subsidiary of Biosensors International Group Ltd. : the drug is applied over a layer of semisynthetic glycocalix, a substrate designed to improve biocompatibility of the metal stent surface after the drug is released.
• delivery platform comprises the balloon catheter, the stent, and the drug carrier vehicle. The sirolimus-eluting Bx VELOCITY^TM stent (Cordis Cardiology) has a closed cell design and is available in 2 configurations (6 cells and 7 cells). As a consequence of the closed cell design, drug distribution is theoretically more uniform, but the stent is less flexible. The advantage of having a larger stent for larger vessels is that a more constant drug-to-artery ratio is achieved. The paclitaxel-eluting Express 2^TM stent (Boston Scientific), although exhibiting somewhat greater flexibility because of its open cell design, elutes a fixed dose of drug for any given length of stent (regardless of the diameter of the underlying vessel). This implies a variable drug-to-artery ratio. Compared to smaller vessels, larger ones may be relatively underdosed, so the drug may have a diminished effect in large-diameter vessels. The other, and perhaps more important, aspect of the delivery platform is the drug carrier vehicle. There are 2 important reasons for having a drug carrier vehicle associated with a DES. The first is to offer controlled release of the drug, in regard to both rate and total duration of release. The second is to enhance drug distribution. Ideally, the drug is distributed into the wall of the blood vessel in a uniform manner. Interestingly, by using different polymers and polymer designs, the direction of distribution can also be manipulated. For ideal application with a DES, the drug carrier vehicle should have several important characteristics: it should be noninflammatory, nonthrombogenic, and sterilizable, with a durable shelf life. Also, the vehicle should neither affect the structure of the drug nor hinder delivery of the stent itself. The drug carrier vehicle may be present in the form of a polymer, which may serve only as a diffusion barrier or may be a matrix for drug loading :


This matrix may be durable or degradable :

The CYPHER stent has both a polymer matrix loaded with the drug and a polymer topcoat (without a drug) that serves as a diffusion barrier. The polymer, poly(ethyleneco-vinyl acetate) (PEVA)/poly(butyl methacrylate) (PBMA), elutes 100% of the sirolimus, most of which has eluted after approximately 1 month. The TAXUS stent has a monolayer polymer matrix (polyisobutylene), which elutes the drug directly in a biphasic manner. There is an initial burst release of the drug on the more superficial portion of the polymer. Subsequently, up to 10% of the paclitaxel is released over the next 2 months. However, 90% of the drug remains sequestered in the polymer indefinitely. The long-term implications of the presence of the sequestered drug are unknown. Instead of being associated with a polymer, a structural modification of the stent itself may serve as the drug delivery vehicle. The drug may be placed directly in a groove or well on the surface of the stent. In this case, a polymer may be used as mentioned above to further fine tune delivery of the drug. The Conor stent (Conor Medsystems, Inc., Menlo Park, CA) has a series of built-in wells along its struts that serve as drug reservoirs. These wells may contain 1 or more drugs in layers, and the opening of the well may be “capped” with a polymer that can serve as a partial or complete diffusion barrier. In this manner, elution of 1 or more drugs in sequence can be engineered to take place in a unidirectional fashion. The drug carrier vehicle may also be formed by a treated stent surface (i.e., in association with a nanoporous material). Clearly, use of a drug carrier vehicle adds to the precision and performance of a DES
• practical, clinical applications :
• efficacy : in comparing the CYPHER and TAXUS stents, efficacy may be the most important consideration. In evaluating stents, research protocols can obtain widely different restenosis rates, depending on the clinical variables involved (vessel size, stent design, clotting status, the presence of unstable angina, and the overall clinical situation). Therefore, any valid comparison should rely on the variable that changes the least, which in this case is late luminal loss. Stent placement causes an arterial injury, which commonly includes vessel medial dissections. The smooth muscle cells within the artery become active, proliferate, migrate, and produce extracellular matrix. The end result is neointima formation, and the thickness of the neointima is directly proportional to late loss. The lower the level of this variable, the better the clinical outcome. As late loss increases, there is more neointimal volume, or intimal hyperplasia obstructing the vessel, and patients are more likely to need target-lesion revascularization (TLR). Because late loss is minimally influenced by other clinical variables, it can be used to determine which stent is more effective at suppressing neointimal proliferation. In 7 different trials late loss with a bare metal stent stayed relatively constant, ranging from 0.8 to 1.0 or 1.1 mm. A variety of bare metal stents were assessed: the MULTI-LINK stent (Guidant Corporation, Indianapolis, IN); the NIR stent (Scimed, Boston Scientific, Maple Grove, MN); the MULTI-LINK PENTA^TM stent (Guidant); the MULTILINK VISION^TM stent (Guidant), which is made of cobalt chromium; the Bx VELOCITY stent (Cordis); and the BiodivYsio™ stent (Abbott Laboratories and Biocompatibles International PLC, Abbott Park, IL). The Randomized Study With the Sirolimus-Eluting Bx VELOCITY Balloon-Expandable Stent (RAVEL) and the TAXUS II trial, which involved benign lesions, had a late luminal loss of 0 and 0.31 mm, respectively. The relative reduction in TLR was 100% in RAVEL and 61% in TAXUS II. More clinical trial results are now available, including those from the SIRIUS and New SIRIUS studies, the latter representing pooled data from the C-SIRIUS (Canadian) and E-SIRIUS (European) trials, which set a new benchmark for evaluating sirolimus-eluting stents in real-world conditions. Respective late losses of 0.17 and 0.18 mm, which are typical with these stents, can be compared to a late loss of 0.39 mm in the TAXUS IV trial, which represented real-world clinical results with the TAXUS stent. The relative reduction in TLR was lower with the TAXUS (58%) than with the CYPHER stent (83%). With respect to volume obstruction (the ability of the drug to inhibit neointimal proliferation), the TAXUS II trial showed less reduction in TLR than did the RAVEL trial. This finding held true in more complex patient subsets that, again, had a smaller decrease in volume obstruction in TAXUS IV than in SIRIUS and New SIRIUS. The occurrence of clinical events seems to parallel the reduction in late luminal loss. The relative reduction in TLR was 100% in RAVEL, 75% in SIRIUS, 80% in New SIRIUS, 63% in TAXUS II, and 73% in TAXUS IV. In the odds ratio plot for the SIRIUS trial, based on results obtained at 9 months, the overall TLR rate was 4.1%. The most striking result was the overwhelming effect that sirolimus had in the different subsets of patients in the trial, including those with diabetes mellitus, left anterior descending coronary artery disease, small vessels, larger vessels, overlapping stents, and nonoverlapping stents. At 9 months, the difference between the TAXUS IV and SIRIUS data was less dramatic, and 2 subgroups showed no significant improvement. One of these subgroups consisted of insulin-treated diabetic patients, but this group was so small that firm  conclusions were hard to reach. The other subgroup consisted of patients with larger vessels (> 3.0 mm), who did not fare significantly better with a TAXUS stent than with a bare metal one. With respect to the 12-month TLR rate in the TAXUS IV and New SIRIUS trials, a significant improvement was seen with both devices. However, the New SIRIUS patients had a much more dramatic reduction in TLR rate than did the TAXUS IV patients (1.3% versus 3.0%, respectively). There is a question regarding the best DES for treating diabetic patients. In the different randomized clinical trials, the results have generally been similar with both sirolimus- and paclitaxel-eluting stents. The e-CYPHER results (with adjudicated registry data) were recently presented at EuroPCR (the Paris Course on Revascularization) with a favorable TLR rate of 1.4% in 2716 patients. These results are significant, given the sheer number of patients involved. For insulin-treated diabetic patients, the data are, again, fairly similar with both stents. It is reassuring that the e-CYPHER group included 884 of these patients, which is one of the largest groups of diabetic patients analyzed in this regard. For patients with small vessels, late luminal loss ranges from 0.01 mm to 0.22 mm in the sirolimus trials versus 0.35 mm in the TAXUS IV trial (Figure 10). In assessing the effects of these stents in challenging subsets of patients, it is important to note that whereas specific data are available for the CYPHER stent regarding the different clinical correlates of TLR, from the European and Canadian SIRIUS trials, the Sirolimus-eluting stent in Chronic Total Occlusion (SICTO) Trial, the Rotterdam trial, and the TROPICAL Study recently presented at EuroPCR, similar data are not available for the TAXUS stent at this time. With regard to in-stent restenosis, a different profile of restenosis patterns has emerged between sirolimus- and paclitaxel-eluting stents as shown in the TAXUS IV, TAXUS VI, and SIRIUS trials. When in-stent restenosis patterns were analyzed in the SIRIUS trial, with the CYPHER stent, most of the restenosis (87%) was focal as opposed to nonfocal (13%). In TAXUS IV and TAXUS VI, however, focal lesions accounted for 63% and 62% of the restenoses, respectively, the remainder being nonfocal. Total occlusion as a result of restenosis occurred twice as often with TAXUS stents as with CYPHER stents. These findings are significant in terms of clinical outcomes because the TLR associated with the treatment of nonfocal restenosis is higher than that associated with focal restenosis. With respect to long-term followup, the insight gained from followup of the GAMMA I Trial may raise important concerns. At 1 year, brachytherapy treatment was considered to have worked exceptionally well. After 3 to 4 years, however, there was no significant difference between the study group and the control group. This experience highlights the importance of longterm follow-up data for any new therapy in interventional cardiology. Follow-up results from the First-in-Man trial, performed in Brazil, are very reassuring: 4 years after placement of the CYPHER stent, the Control bare metal stent results are similar to those obtained immediately after the procedure.
• safety : the fourth broad category for comparison of the CYPHER and TAXUS stents is safety. In this regard, the most important consideration is the rate of major adverse cardiac events (MACE). The 3-year follow-up data from the RAVEL study show favorable MACE curves for the CYPHER stent. The 12-month MACE rate was 10.6% in the TAXUS IV study and 8.0% in the New SIRIUS study. Moreover, stent thrombosis, which was initially believed to be a cause for concern, has not been a problem with either stent system. The recent presentation of the US e-CYPHER registry demonstrates a real-world pattern of usage of CYPHER stents in the USA. It showed an overall MACE rate of 3.9% and a total stent thrombosis rate of 0.7% at 6-months follow-up. The TAXUS VI trial, recently completed in Europe, involved patients with long lesions. When the TAXUS VI cohort is compared to 2 SIRIUS and New SIRIUS subgroups that had comparable lesion lengths, the TLR rate is very similar; rather surprisingly, however, the MACE rate was significantly elevated (16.4%) in the TAXUS VI group, which was not different from the control group in this respect. Few data are available for the TAXUS stent concerning overlapping of stents, but data have recently become available from the TAXUS VI trial. The comparison might not be clinically relevant, as the TAXUS moderate-release stent (not available in Europe or the USA) was used in TAXUS VI. Nonetheless, some lessons can be gained from this trial. In the TAXUS VI cohort, 27.8% of the patients received overlapping stents. They were followed up for 9 months, and their rate of subacute stent thrombosis was comparable to that of the control group. Thus, with respect to TLR, overlapping with the moderate-release TAXUS stent was not associated with increased hazardous events. Data regarding overlapping CYPHER stents are significantly more robust, including 2-year followup results from SIRIUS and 1-year results from New SIRIUS. In these 2 trials, the fraction of patients who received overlapping stents was 28% and 35%, respectively. The stents significantly reduced the need for TLR and did not significantly increase the rate of subacute thrombosis. Speaking at the EuroPCR, Virmani directly compared overlapping of the TAXUS slow-release stent (which is available in the United States and Europe) to overlapping of the CYPHER stent in a rabbit model. The latter device resulted in a significantly higher percentage of surface endothelialization, which is an important indicator of stent safety.
With respect to efficacy, the CYPHER stent has a lower late luminal loss rate than the TAXUS stent. The significance of this finding and its impact in clinical practice remains unclear and is still subject to the individual physician’s judgment. However, all of the
above-mentioned trials have shown that less intimal proliferation and obstruction occurs with the sirolimus platform than with the paclitaxel platform. Both stents offer excellent short-term safety. In addition, compared to bare metal stents, the CYPHER device reduces late loss in challenging subsets of patients. No parallel data are available for the TAXUS stent. Further knowledge, from a head-to-head comparison (reality trial) of these 2 stents, should improve our understanding of their varying applications.
Complications :
• arterial dissection (treated with a new stenting procedure)
• arterial rupture
• atheroembolism
• injury due to catheterism
• hematomas or pseudoaneurysm at entry site
Indications :
• renal artery stenosis (RAS) when...
• clinical indications :
• uncontrolled systemic arterial hypertension
• initial stage renal failure
• transplanted kidney
• aortorenal bypass
• radiological indications
• stenosis of renal artery > 60%
• kidney diameter > 8 cm (risk of ischemia reperfusion injury (IRI) is higher in small kidneys)
• peripheral arteries occluding disease (PAOD)
• lower limb obliterating arteriopathy since Leriche-Fontaine stage IIb on
• epiaortic vessels
• subclavian steal syndrome (stenosis of vertebral artery > 60%)
• hemodialysis access
• endoprosthesis (stent-endograft) for aneurysms (endovascular aneurysm repair (EVAR)) : prosthesis lacks support to resist bloodstream
• thoracic aorta aneurysm (TAA) : occlusion of intercostal artery could lead to tetraplegia or paraplegia
• abdominal aorta aneurysm (AAA) (endovascular AAA repair (EVAR) ) : treatment is evaluated on the basis of diameter, calcifications, thrombosis, relationships with iliax axes, length > 5 cm.
• modular flexible bifurcated endovascular stent-graft
• Aorfix^®
• AneuRx^® (source : Medtronic; Minneapolis, MN)
• Excluder^®, considered a 3rd generation device
• Ancure^® endografting system (Guidant Cardiac and Vascular Division, Menlo Park, Calif) features a unibody, nonsupported woven polyester graft
• Zerith^®
• Talent^® from Medtronic
• Endologix^®
• Lifepath^®
• Vanguard^®
• Fortron^®
• Zenith^® from Cook

• nonmodular endografts



The proximal attachment should be > 15-20 mm long and have healthy walls (calcifications for < 30% of diameter, no thrombi, diameter < 3 cm) : if the subrenal proximal neck is not ideal, a cranial attachment can be used with uncovered stent
Complications in placement arise from calcifications, tortuosity or small-diameter of iliac arteries. Patent inferior mesenteric artery (IMA) (75%) may also cause intestinal ischemia.
Follow-up of endograft by ECD, angio-CT (at dimission, month 3, 6, 12, 18 and 24) to evaluate reduction of aneurysmal diameter and perigraft flow, and complications :
• endoleaks^ref (persistent or recurrent flow into the aneurysm) :
• type I endoleak / attachment leak (3.2%) : mechanical separation of the device from the native vascular system. Blood flows around the stent-graft at an attachment site, which results in filling of the aneurysm sac, and outflows from lumbar artery. Because systemic arterial pressure is transmitted to the aortic wall, the aneurysm remains at risk of rupture and rquires immediate treatment. Creation of a tighter seal at these attachment sites by using balloon catheters, additional stents, and stent-graft extensions usually can be used to treat this leak.
• type II endoleak / branch flow (11.5%) : flow enters the aneurysm sac through the branch vessels—the lumbar arteries, the inferior mesenteric artery (IMA), the hypogastric arteries, and so on—that have reversed their normal direction of flow. As is the case in type I endoleaks, in type II endoleaks, arterial pressure is transmitted into the aneurysm sac, and the patient remains at risk for aneurysm rupture. Requires treatment with embolization only when impairs reduction in aneurysmal volume.
• type III endoleaks (0.6%) are due to fabric tears, graft disconnection (modular disconnection in multimodular endograft), or disintegration and requires immediate treatment
• type IV endoleak (1.2%) : due to graft wall porosity. This kind of endoleak disappears with time
Endoleaks can also be classified according to when they first appear. Leaks identified at final postcompletion intraoperative angiography are called primary, while those identified later are called secondary. This classification system is problematic, as there is the potential for small leaks to be unrecognized at intraoperative digital subtraction arteriography and identified at postoperative CT
• graft limb occlusion (2.5%)
• graft miigration (10.9%)
• risk of AAA rupture after endograft placement = 0.4% within 1-yr, 1.2% within 2-yr, 1.6% within 3 years, 2.9% within 4 years
• freedom from secondary intervention : 85% after 1 year, 75% after 2 years, 65% after 3 years, 58% after 4 years, 45% after 5 years. Secondary intervention is performed by endovascular route (12.9%) or open surgery (6.8%).
• infection of prosthesis (0.4%) => mortality = 27.4%
• in the early postoperative period there is a small, yet significant quality of life (QoL) advantage for EVAR compared to open repair (OR). At 6 months and beyond, patients reported better QoL after OR than after EVAR.
• iliac arteries
• popliteal arteries
If the prosthesis become infected, it can be replaced with a cadaver homograft.
• transjugular intrahepatic portosystemic shunt (TIPS) : percutaneous creation of a shunt between the right hepatic vein and right branch of portal vein within the liver followed by placement of an expandable metallic covered stent in the tract created, performed by a transjugular route under radiologic guidance; done for the treatment of bleeding esophageal varices unresponsive to sclerotherapy. PTFE-covered stents are able to solve pseudointimal hyperplasia within the stent tract, but have a high incidence of hepatic or portal vein stenosis


• vena cava filters for the prevention of venous thrombembolism, expecially pulmonary thrombembolism (PTE), in patients who have contraindications to anticoagulation
• superior vena cava filters : in 1985, Langham et al^ref reported that placement of the Greenfield filter in the SVC of 11 dogs was well tolerated at 3-month follow-up, protected against PE, and allowed thrombus resolution while caval patency was maintained. Despite this, there have been only 11 cases of SVC filter placement in humans reported in the English-language literature^{ref1, ref2, ref3, ref4, ref5, ref6}. Although in each of these reports, the SVC filter was effective in preventing PE, the maximum patient follow-up was only 14 months^ref. Percutaneous filter insertion in the SVC is technically more demanding than insertion in the IVC because of the relatively small area for filter deployment. This could theoretically result in a higher complication rate for SVC filter insertion. Furthermore, recognized complications of IVC filter placement could potentially be more severe for SVC filter placement. Filter migration is more likely to result in an intracardiac position. Caval perforation may result in cardiac or aortic injury. SVC occlusion is more likely to result in substantial morbidity because of the reduction in potential collateral vessel pathways. In addition, the safety of central venous or Swan-Ganz catheter placement after SVC filter placement has yet to be determined.


• retrievable
• Bard
• Gunther Tulip caval filter (William Cook, Europe)
• permanent Kim-Ray Greenfield filters
• inferior vena cava (IVC) filters : in order to trap thrombus from pelvic and lower limb origins. They are surpassing surgical methods of preventing pulmonary embolus, e.g. femoral vein ligation, because they have a similar efficacy but are associated with a lesser morbidity. They are particularly indicated in patients who have a contraindication to anticoagulation or who have ongoing pulmonary embolism despite full anticoagulation. The filter is inserted percutaneously with only local anaesthesia via jugular or femoral routes. The filters are commonly sited below the renal vein. Even with a filter, there is a 5% risk of recurrent pulmonary embolus. Similarly, the complication of leg swelling can occur. Hence, anticoagulation is continued for several months. Surgical interruption of the IVC to prevent pulmonary embolization was first suggested by Trousseau in 1868 and subsequently performed by Bottini in 1893. In 1959, Moretz first described external clipping of the inferior vena caval by means of a clip that created a single slitlike channel. This development was followed by the introduction of the serrated Miles and Adams and DeWeese devices that created 4 channels within the IVC. Transvenous interruption of the vena cava, eliminating the need for general anesthesia and a laparotomy, became clinically feasible in 1967 with the introduction of the Mobin-Uddin filter. Due to the unacceptably high incidence of IVC occlusion associated with use of this device, the Kimray-Greenfield filter, introduced into clinical practice soon thereafter, became the preferred device for transvenous caval interruption. Nearly 50,000 patients underwent IVC filter placement in the USA in 1999. Since the introduction of the Greenfield filter, a number of other percutaneous IVC filters have been developed and introduced into clinical practice, with the intent of optimizing flow dynamics, maximizing clot-trapping capacity, and expediting ease of insertion. Currently, several IVC filters are available for use in the USA :
• umbrella filter : the most common kind of vena cava filter for prevention of pulmonary embolism; it is inserted in a folded position and springs open like an umbrella to engage the caval wall. Common types are ...
• Kimray-Greenfield stainless steel or titanium filter (Boston Scientific): an umbrella filter consisting of 6 stainless steel struts; small hooks on the ends of the struts anchor the filter in the vena cava when it is opened
• Mobin-Uddin filter : an umbrella filter consisting of 6 stainless steel spokes connected to a hub and covered by a perforated, heparin-impregnated Silastic membrane
• Vena-Tech LGN or LP (Braun)
• Simon-Nitinol (Bard)
• Recovery (Bard)
• bird's nest filter (Cook) : an inferior vena cava filter consisting of 4 long, thin stainless steel wires with many bends, attached to struts that are fixed into place
• Gunther Tulip Vena Cava MReye™ filter and retrieval system (Cook); design allows for intravascular retrieval and removal of the filter up to ten days after placement. Currently marketed filters are usually permanently placed and not removed.
• TrapEase (Cordis / Johnson + Johnson)
• OptEase (Cordis / Johnson + Johnson)
Beyond differences in design and material, each differs in terms of the outer diameter of the introducer system, ranging from 29F for the older Greenfield stainless steel (24F), to approximately 12F for the Greenfield stainless steel (12F) and the Greenfield titanium, Vena-Tech, and Bird's nest, 9F for the Simon-Nitinol, to 6F for the TrapEase.
• synthetic "pumps"
• artificial tendons : gazing at an aquarium tank almost 20 years ago, Thomas Koob made an observation that would eventually change the course of his research. While reflecting on the remarkable properties of a skate egg case, he realized it might hold the secret to an artificial tendon. Tendons are the ropes that tie muscle to bone, enabling the former to maneuver the latter quickly and efficiently. Some, such as the Achilles' tendon, act as springs. Others get the body moving and, for example, help turn the open hand into a clenched fist. But when they break, our ability to repair tendons is limited : surgeons try to fix some 300,000 tendons a year--and that doesn't include those in the legs and feet. Recovery can take months and at best the restored tendon is about 60% functional. In their first attempts to create replacement tendons, researchers tried polyesters or other synthetic materials. But more recently, several research teams have gone after biologically based materials that might be less prone to immune rejection. A substitute must be strong, stiff, and resilient, and it should be able to change its properties--such as its stiffness--as real tendons do. Contrary to the common view of them as elastic cables, tendons are very much alive : viewed one way, the cells within a tendon look like boxcars along tracks of collagen. But it's really more like a spider web. The cells are highly interactive and intertwined with the collagen. This network includes molecular bridges that reach across and set up cross-links between nearby collagen strands, thereby preventing the collagen from slipping as the tendon is pulled by a contracting muscle. Collagen helps make the tissue strong, but the tendon cells are equally important, because they react to loads, activate genes, and change the properties of the tendon itself. Any replacement tendon must either provide new cells or recruit new ones into the artificial tissue if it is to function as well as the original. Others are trying to make a biological material strong enough to temporarily take a tendon's place as it heals : skate eggs emerge white and pliable from a large gland in the reproductive tract; over the next few hours, they darken and become leathery. Made of collagenlike materials, the egg cases are tough enough to persist in the ocean for months, and so are its enzymes. These enzymes form bonds that are the cross-links between the case's collagenlike fibers. A chemical called NDGA, derived from a shrub called creosote, can do just that in the lab. The collagen-NDGA combination is very compatible with rabbit tissue and does not trigger immune reactions (the Food and Drug Administration requires this test for biomaterials to be used in humans). For that reason, this natural tendon material might be safer than a synthetic. Not that this new tendon material is ready for use. Researchers still need to find a way to glue in the artificial tendon--something that can already be done with a woven polyester tendon material. The ideal tendon material would need to contain or recruit cells that could help the tendon develop the right set of properties and repair small tears as needed. Because his material accumulates cells only on the outside, Koob is now working on making it less dense and more like a real tendon. Meanwhile, Banes and those working on gel-based substitutes have materials that already come with their own tendon cells or, in one case, stem cells. These groups plan to add genes to these cells that would enhance their ability to work like natural tendon cells. At this stage, however, no one can predict which approach will yield a tendon that is almost as good as the real thing^ref.
• real-time continuous glucose monitoring (CGM) system (Guardian RT, made by Medtronic, Inc.) uses a disposable sensor that is inserted subcutaneously in the abdominal area to measure interstitial fluid glucose levels. Each sensor lasts up to 3 days and takes about 288 glucose readings per day. Sensor readings are relayed via radiofrequency transmitter to a monitor and displayed in real-time every 5 minutes. Preset alarm thresholds alert patients to potentially dangerous glucose levels, and trend reports/charts can be analyzed after data are downloaded to a computer using therapy management software and a docking system. Fingerstick measurements are required to calibrate the system every 12 hours and prior to adjusting insulin delivery. Unlike glucose meters that are used only at discrete intervals and yield information gaps, the system provides continuous data regarding the effects of diet, exercise, medication, and lifestyle on glucose levels for improved disease management. The real-time display also enables prompt intervention to reduce the frequency and severity of hypoglycemic and hyperglycemic episodes for improved control.
• biofuel cells : they convert directly into electricity the energy produced when glucose reacts with oxygen during normal metabolism. It could be inserted in contact with glucose-containing body fluids : to be medically useful, such cells need to be tiny, they must work at the temperature, acidity and salt concentration of blood, and must produce sufficient power and voltage. An experimental model consists of 2 carbon fibres, 2 cm long and 7mm wide, each coated with a catalyst that assists the chemical reactions of glucose burning. This reaction takes place in 2 halves, one at each electrode. One electrode is coated with glucose oxidase from Aspergillus niger :
2 a-D-glucose + 2 O₂ + 2 H₂O => 2 gluconolactone + 2 H₂O₂ ... immobilized and wired to hydrophilic (plasma-treated) graphite particles of ~ 100 µm diameter. At the other electrode, another polymer-wired enzyme (microperoxidase-11, cytochrome b oxidase, ... ) adds electrons to dissolved oxygen. As the reaction proceeds, electrons are fed in a current around an electrical circuit. The device works at a temperature and alkalinity close to those of normal blood: at 37 ºC, and at pH 7.2. It produces about the same amount of power as a wristwatch battery: about 1.9 mW. It could drive a miniature glucose sensor for monitoring diabetics, but it couldn't power an artificial heart. The device is too weak and too short-lived - currently it loses around 6% of its power per day. Rather such biofuel cells could be used as research tools, for example to power tracking devices attached to insects or animals over the course of a few days. Moreover, real body fluids, like those between our cells, are more complicated than the model fluids that the researchers have used so far, and may introduce complications.

Dr Lee Ki Bang of the Singapore state-funded Institute of Bioengineering and Nanotechnology said a drop of urine placed on the paper battery will generate enough electricity to power a "biochip device" that can analyse the urine sample for disease "biomarkers".
• artificial muscles : feel the pumping of your heart or leap to witness the wonder of some of nature's muscles. Skeletal muscles self-repair, provide billions of work cycles involving contractions of more than 20%, increase strength and change stiffness in response to need, generate stresses of ~0.35 MPa, contract at 50% per second, and can even transform to fuel for the starved body. They convert the energy of a safe, energetic fuel (adenosine triphosphate) to mechanical energy with higher maximum efficiency (~40%) than that achieved by a typical car engine. Artificial muscles offering even higher performance are being sought for artificial and damaged hearts, artificial limbs, humanoid robots, and bird- or insect-like air vehicles that fly by flapping wings. How well do such artificial muscles compare with natural muscle, and what are the prospects for future advances? This line focuses on artificial muscles that generate large strains (fractional changes of muscle length) of about 20%, rather than high force, high response rate, and/or high output power at low strain. Instead of mimicking nature by creating large macroscopic strains by the combined effects of trillions of molecular actuators, the artificial muscles use material deformations. Electronically conducting polymers such as polyaniline and polypyrrole provide one type of high-strain actuator. In what is basically a battery, these muscles actuate by using dimensional changes produced by electrochemically inserting solvated dopant ions into a conducting-polymer electrode. Although first described almost 2 decades ago^ref, conducting-polymer muscles with strains matching those of natural muscles have only recently been realized. These advances are likely to accelerate commercialization efforts^ref. Arrays of micrometer-scale conducting-polymer muscles have been made that undergo complex, coordinated motions. However, these devices^{ref1, ref2} typically operate at a strain < 1% by using the cantilever effect to amplify displacements.


A proposed high-stroke artificial muscle. (Top) Cross section of the proposed actuator. A highly twisted nanofiber yarn electrode (black) is filled with a conducting polymer or gel electrolyte (red). The conducting polymer or electrolyte does the actuation. The nanofibers provide nanometer-scale electronic connections and giant surface area for electrochemical charge injection. A counterelectrode and a porous electrode separator are also included. A liquid electrolyte (yellow) provides the electrochemically inserted solvated ions that cause actuation when a voltage is applied between the electrodes. (Bottom) Scanning electron micrographs of highly twisted four-ply and single-ply multiwalled carbon nanotube yarns^ref. The oblique nanotube alignment (red lines) enables large extension in the yarn direction. State-of-the-art conducting-polymer artificial muscles operate at voltages of a few volts and can generate high strains (26%), high strain rates (11% per second), and large stresses (7 to 34 MPa). However, actuator performance typically degrades after a few thousand cycles when the actuator is operated at high actuator strains and high applied mechanical loads. Energy conversion efficiencies are low (typically <1%) but could be improved drastically by avoiding electrolyte electrolysis, harvesting stored electrical energy, and using doping-induced conformation changes^{ref1, ref2}. Another type of artificial muscle is based on dielectric elastomers, which resemble the silicone rubbers used as sealants. These systems can generate higher actuator strains than do conducting-polymer artificial muscles, and they constitute capacitors rather than batteries^ref. Actuation is mainly caused by "Maxwell stress," which results from the attraction between charges on opposite capacitor electrodes and the repulsion between like charges. Artificial Muscle Inc.^ref was recently founded to exploit the impressive performance of elastomeric actuators, which can generate strains of 120%, stresses of 3.2 MPa, and a peak strain rate of 34,000% per second for 12% strain. When elastomer stiffness is decreased, maximum stress generation decreases, but maximum actuator stroke and work-per-cycle increase. Operating voltages of several thousand volts have thus far limited application, but strategies for avoiding this problem are evolving^ref. A third type of artificial muscle uses the volume change of an electrolyte and electrostatic repulsion. One version, which amplifies low strains using the cantilever effect, is called the ionic polymer/metal composite actuator^ref. These actuators, sold by Environmental Robots Inc.^ref, consist of 2 metal-nanoparticle electrodes that are both filled with a solid electrolyte and separated by this electrolyte. The actuators act as supercapacitors: An applied interelectrode potential injects electronic charge into the high-surface-area electrodes. To balance this charge, solvated ions migrate between the electrodes, causing one electrode to expand relative to the other, thereby bending the cantilever actuator strip. Tensile strains of 40% have been produced by a related device in which an electrode-wrapped polymer gel yarn is separated from a counter electrode by a liquid electrolyte^ref. Actuation results from a local pH change caused by electrolysis of the electrolyte and transport of hydrated ions and water into the yarn, but actuation rate and energy conversion efficiency are low. Other high-strain artificial muscles include shape-memory alloys, which are commercially important and generate strains of up to 8%, but require conversion of electrical energy to thermal energy and the inefficient conversion of the latter to mechanical energy. Actuators that use electrochemically generated gases confined in carbon nanotubes provide strains of 300%, but energy conversion efficiency and cycle life are low^ref. All of these high-strain artificial muscles suffer from challenging problems that have delayed their commercial application. How might these problems be overcome for applications such as autonomous biomimetic robots? A possible clue is provided by low-voltage, low-strain actuators that use electrochemical charge injection into a nanostructured electrode, resulting in electrostatically driven electrode expansion^ref. One such device type, carbon nanotube artificial muscles, can generate 100 times the stress of natural muscle and provides comparable actuation rates (20% per second), but the actuator strain is at best 2%. These performance characteristics could be improved by replacing the carbon nanotubes with, for instance, a nanostructured elastomeric conducting polymer that has a comparable gravimetric surface area but is more easily deformable (by a factor of 100). Another approach is to radically improve the ionic polymer/metal composite actuators by first replacing cantilever actuators that operate by bending with actuators that operate in tension. This could be done by separating the opposing electrodes by a liquid electrolyte, which would provide the solvated ions that cause actuation. The gravimetric capacitance of today's ionic polymer/metal composite actuators is less than one-tenth that of other supercapacitors. Increasing this capacitance by increasing the electrode surface area could provide a corresponding increase in actuator strains, as long as the electrode does not restrict movement in the actuator stroke direction. Highly twisted carbon nanotube yarns, with their high electrical conductivities and high surface area^ref, seem ideal for use as this type of electrode. Nature's scheme of converting the chemical energy of a high-energy-density fuel to mechanical work has advantages, which can be partially achieved by driving electrical actuators with fuel cells. Because a fuel cell and fuel source can deliver more than three times as much electrical energy as a battery of the same weight, fuel cells are preferred for autonomous robots that use electrically powered artificial muscles. It would be very useful to have muscles for artificial hearts that use the same fuel and arterial fuel delivery system as natural muscle, or autonomous humanoid robots for hazardous environments that drink a shot of alcohol or hydrazine to continue work. Gel actuators and conducting-polymer actuators can be driven to provide high actuator strains by means of chemical energy, but either the fuel energy density or the chemical-mechanical conversion efficiency is low. If an artificial muscle could simultaneously serve as a fuel cell, then weight, volume, and cost reductions could result. Inactive muscles could be used as fuel cells to provide other electrical needs for the robot, or could be joined together in series to power high-rate actuation for active fuel-cell muscles. The concept of making a carbon nanotube-based fuel cell muscle is patented^ref and has been demonstrated in a primitive prototype. Yet humankind is still far behind in much of nature's game of providing high-strain muscles^ref.

• 1504 : iron prosthetic hand with flexible finger joints
• 1888 : first reports on use of contact lenses to correct vision
• 1905 : early attemp at an artificial hip replacement
• 1928 : iron lung developed for treatment of polio victims
• 1939 : hard (plastic) contact lenses introduced
• 1943 : kidney dialysis machine developed
• 1951 : first artificial valve implanted
• 1953 : development of the hear-lung machine
• 1957 : first cochlear implant developed
• 1958 : early attempts at developing an implantable pacemaker
• 1969 : total artificial heart implanted in a human as temporary measure
• 1972 : testing of modern design steel/polymer hip joint
• 1979 : first auditory brainstem implant
• 1980 : first successful single-channel cochlear implant in a child
• 1982 : implantation of the Jarvik-7, a permanent total artificial heart
• 1987 : first clinical use of a bioartificial liver device
• 1993 : FDA approval of left ventricular assist device as a bridge to heart transplantation
• 2000 : implantation of a prototype artificial pancreas
• 2001 : implantation of the AbioCor, a permanent self-contained heart replament

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