Ординатура / Офтальмология / Английские материалы / Biomaterials and regenerative medicine in ophthalmology_Chirila_2010

(b)

17.6 A smooth non-porous implant with the rectus muscles imbricated over its anterior surface can slip between the muscles. In this photograph of a child (a), the implant has migrated supero-temporally (between the lateral and superior rectus). The corresponding diagram (b) shows the imbricated muscles, the direction of migration (arrow) and part of the spherical implant (cross-hatched area).

from Germany led to development of the acrylic poly(methyl-methacrylate) (PMMA) ocular prosthesis (Murphey et al. 1945). Pink acrylic for dentures had already been used to make post-operative conformers (Kelley 1970).

The acrylic ocular prosthesis permitted custom fitting of the prosthetic to the implant and was critical to development of the motility implants. Evolution of the exposed and buried integrated implants was concurrent.

17.4.1 Exposed integrated implants

Ruedemann’s PMMA implant (1945) was a combined motility implant and ocular prosthesis. The posterior portion of the implant was covered with tantalum mesh to allow attachment of the rectus muscles, Tenon’s capsule and conjunctiva (Ruedemann 1945). The inability to remove the ocular prosthesis for cleaning, the need for custom prefabrication prior to

444 Biomaterials and regenerative medicine in ophthalmology

implantation and problems with positional deviations led to its abandonment (Durham 1949).

In 1947, Cutler described a PMMA ‘ball and ring’ implant. The exposed face had a square female receptacle, into which a 14 kt gold square male peg of the ocular prosthesis would fit. The rectus muscles were looped around and sutured to the ring (Cutler 1947), Other similar pegtype implants were produced. The Hughes enucleation implant was similar to Cutler’s in shape, but was made of vitalium (an inert lightweight alloy of cobalt, chromium and molybdenum) (Hughes 1948). Whitney coupled the extraocular muscles to an implant by incorporating tantalum gauze around an acrylic sphere. Stone designed implants with metallic prongs on to which the rectus muscles were impaled. The Rolf implant incorporated a ring for muscle attachment and tantalum mesh anteriorly for conjunctival attachment (Gougelmann 1970).

Direct coupling of the ocular prosthesis to the implant significantly improved motility, apparently by eliminating slippage between the implant and ocular prosthesis. The supporting peg also helped to create a fuller upper lid sulcus and reduced the weight placed on the lower eyelid. Unfortunately, excessive secretions, recurrent granulations and chronic infection were common complications of exposed integrated implants (Drucker 1951; Perry 1991). In a review of the outcomes of 91 exposed integrated implants (74 hollow tantalum, 17 gold ring and cylinder), 50% of the hollow tantalum implants survived at 2 years while 60% of the gold ring implants survived after 3 years. Infection was the reason for extrusion/removal in 80% of cases (Choyce 1952). In retrospect this was the expected outcome of a chronically disrupted epithelial lining. This unifying failure led to the adoption of buried integrated implants.

17.4.2 Buried integrated implants

In theory a buried integrated implant combines the advantages of the smooth buried and exposed integrated implants in providing: (a) an uninterrupted conjunctival lining to minimize discharge and infection related to exposures, and (b) an irregular surface to translate implant movement to ocular prosthesis movement and to partially support the weight of the ocular prosthesis – reducing pressure on the lower lid and permitting a fuller upper lid sulcus.

In 1942, Dimitry described and patented the creation of an elevated stump on the implant surface, meant to fit into a ocular prosthesis with a posterior concavity – but he did not report any actual results (Guyton 1948). Cutler introduced a basket implant in 1945. This implant had four openings through which the rectus muscles were pulled and sutured together with conjunctiva closed over it. The ocular prosthesis had a knob on its posterior surface that

fitted into the concavity of the implant (Cutler 1946). As with the exposed integrated implants, others developed similar types of implants. The King implant consisted of a pear-shaped tantalum mesh. The rectus muscles were attached to the base of the mesh and conjunctiva closed over it (Gougelmann 1970).

Among the better known of the buried integrated implants is the Allen implant. The story of this implant nicely captures the progression of orbital implant design/philosophy from the mid 1940s to the mid 1980s. What eventually became known as the Allen implant initially began life in 1946 as an exposed integrated implant (Allen 1950). In contrast to the Cutler design (Cutler 1947) (female implant), the Allen design incorporated the peg into the implant (male). Each rectus muscle was passed through a peripheral tunnel, split lengthwise to straddle the gold peg and sutured to its antagonist. Most such implants were retained only a few months before they extruded or were removed because of secondary infection. Consequently, the peg was removed and muscles were sutured together (i.e. imbricated) through a central 6 mm opening. Tenon’s capsule and conjunctiva were completely closed over the flat surface of the implant. This design also turned out to be problematic: repeated exposures over the flat anterior plastic surface. It was thought that the exposures were related to an inadequate subconjunctival tissue bed. Thus the central anterior opening was enlarged from 6 to 15 mm. Imbrication of muscles within the ring created a broad, flat surface, permitting excellent translation of movement (since flat surfaces do not slip past each other as easily as curved surfaces) (Allen 1950).

Translation of movements was perhaps too good, as prosthetic edge show on extreme gaze and torsional end-point movements were particular problems of the Allen implant (Jordan et al. 1987). A possible explanation is that the flat apposition between the implant and the ocular prosthesis prevents the implant from slipping underneath the ocular prosthesis when the ocular prosthesis is restricted by the conjunctival fornices on side gaze. Any imbalance between the superior and inferior forces acting on the ocular prosthesis will create a rotational movement. Adding more to the peripheral edges to decrease edge show was impractical as it often created discomfort with opposite gaze. Late exposures over the outer ring were long-term complications (Fan & Robertson 1995) (Fig. 17.7). Since the flat surface did not support the weight of the ocular prosthesis well against gravity, lower lid droop and exaggeration of the upper lid sulcus were noticeable in some patients. In 1979, Jahrling reported a 19% incidence of extrusion among 168 Allen implants (40% in 43 Allen implants placed after severe trauma) (Jahrling 1979). Other investigators, however, have found much lower extrusion rates (2/186 = 1%) (Fan & Robertson 1995).

The successor to the Allen enucleation implant was the Iowa enucleation implant (Spivey 1970). The Iowa I implant was first reported in 1959. The

446 Biomaterials and regenerative medicine in ophthalmology

Iowa II implant was similar in shape but nearly one-third larger in volume (Allen et al. 1969). Like the Allen implant this was a buried integrated implant, originally reported as a ‘quasi-integrated’ implant (Allen et al. 1960). The Iowa implant was made of methyl-methacrylate resin and had four peripheral mounds (of 5 mm height) on its anterior surface designed to integrate with four depressions on the back of the ocular prosthesis. The rectus muscles were brought together through the valleys between the mounds, overlapped (5.0–5.5 mm) and tied together at a central anterior depression. Holes were made through parts of the implant in the hope of permitting fibrovascular tissue growth into the implant. This implant addressed many of the problems with the Allen implant. The four surface mounds supported the ocular prosthesis and reduced the gravitational effect on the lower lid (Jordan et al. 1987). Gaze-dependent ocular prosthesis edge show and torsional movements were also corrected (Allen et al. 1969; Spivey et al. 1969; Spivey 1970). When Iowa implants exposed, it was often over the surface of mounds (Spivey et al. 1969). This was likely due to localized pressure necrosis. As a result the Universal implant was introduced in 1987, with lower and more rounded mounds (Fig 17.7). (Jordan et al. 1987; Jordan 2000) Experience with the Universal implant was limited by the introduction of a new generation of porous implants.

17.4.3 Magnetic implants

Magnetic implants involved coaptation of the implant and ocular prosthesis by use of magnets, with conjunctiva sandwiched in between. There were a number of variations on this premise. (Tomb & Gearhart 1954; Young 1954; Ellis & Levy 1956; Roper-Hall 1956; Gougelmann 1970). These implants had adequate movement, but if the magnet was too strong or misaligned, conjunctiva and Tenon’s capsule could become compressed between the implant and the ocular prosthesis, leading to breakdown and exposure along the outer edges (Soll 1986). In 2007, magnetic coupling of implant and ocular prosthesis was reported again, but this time with porous polyethylene (Miller et al. 2007). It will be interesting to see, in long-term follow-up, if these new porous magnetic implants will have similar complications.

An important cause of possible tissue breakdown and late exposure (Murray et al. 2000) related to magnetic implants, not well described in the literature, may be ferrous toxicity and associated tissue necrosis. Over the long term the magnets rust with exposures developing over the central anterior surface, as opposed to the peripheral edges which are prone to pressure necrosis. The levels of iron in the conjunctiva have been found to be 3–5 times normal (Dr Mark A Baskin, Kaiser Permanente, Oakland, CA, personal communication, 2003).

448 Biomaterials and regenerative medicine in ophthalmology

17.5Porous implants

In the absence of a vascular base, repairing or patch grafting of exposures is difficult. This led to the development of porous orbital implants using hydroxyapatite, porous polyethylene (Medpor®) and alumina. In theory, by permitting vascular ingrowth these implants should:

(a)increase the success rate of surgical repair/patch grafting when exposures develop (although porous implants appear to have a higher exposure rate overall, when compared with smooth synthetic implants) (Custer et al. 2003);

(b)reduce the incidence and severity of infection (since a vascular supply permits immune surveillance and defense);

(c)reduce the incidence of implant migration and extrusion. (Trichopoulos & Augsburger 2005).

17.5.1 Hydroxyapatite implants

Hydroxyapatite had been used as a bone substitute since 1975, but received US Food and Drug Administration (FDA) approval for use as an orbital implant in 1989 (Bio-Eye®: Integrated Orbital Implants, San Diego, CA) (Dutton 1991). Guist had inserted charred bone spheres into the muscle cone over 70 years earlier which produced ‘considerable tissue reaction’ and some reabsorption (Molteno 1980; Molteno & Elder 1991). However, hydroxyapatite implants to do not appear to absorb over time (Holmes & Hagler 1987; Sires et al. 1998).

Porous hydroxyapatite, Ca10(PO4)6(OH)2, is made by a specific genus of reef-building coral. The porous form has a micro-architecture similar to human cancellous (spongy) bone with interconnecting channels. Hydroxyapatite is the primary inorganic portion of human bone. The process by which hydroxyapatite implants are created from sea coral involves intense heat that denatures the proteins, to reduce immune response. When implanted next to bone, new bone growth occurs within its pores. When implanted within soft tissues, fibrovascular tissue grows into the pores (Perry 1991). Reports suggest that unwrapped hydroxyapatite does not become encapsulated as do silicone and PMMA spheres. (Holmes 1979; Dutton 1991; Perry 1991).

Hydroxyapatite incites a foreign-body giant cell reaction. (Rosner et al. 1992). In the animal model a foreign-body reaction may persist up to a year after implantation of a synthetic hydroxyapatite sphere (Sires et al. 1995; Saitoh et al. 1996). In addition, the rough surface of hydroxyapatite can produce exposures where implant and ocular prosthesis come into contact (Perry 1991; Buettner & Bartley 1992; Goldberg et al. 1992; Remulla et al. 1995) (Fig. 17.8). Surgeons began to wrap hydroxyapatite in banked human sclera (which adds about 1–1.5 mm to the final diameter of the implant) to

decrease early exposure risk (40% exposure for unwrapped hydroxyapatite vs 7% exposure for wrapped hydroxyapatite in one study) (Remulla et al. 1995) and to facilitate suturing of muscles to the brittle implant (Custer et al. 2003). Cutting windows into the scleral wrap at the attachment site of the foure rectus muscles and associated ciliary vessels was advocated to promote faster vascularization of wrapped implants (Perry1991; Shields et al. 1992a; Gayre et al. 2002). Pore size also appears to have an effect on the rate of vascularization. Vascular ingrowth occurs more rapidly in hydroxyapatite implants with 200 mm pores than in hydroxyapatite implants with 500 mm pores (Bigham et al. 1999).

The scleral wrap incites a foreign-body reaction and becomes partially absorbed over time (Soll 1986; Rosner et al. 1992), with associated late exposure (Fig. 17.8). Thus the use of autologous grafts (e.g. temporalis fascia or fascia lata from the thigh) (Wiggs & Becker 1992; Jordan & Klapper 1999) was suggested, on the basis that a homologous graft incites less inflammation and is less likely to be absorbed. Other options for autologous wrapping materials include the rectus abdominus sheath (Kao & Chen 1999) and posterior auricular muscle (Naugle et al. 1999). The disadvantage is scarring, inflammation and infectious risk associated with a second surgical site. Processed wraps are also available. These include human donor fascia lata, human donor pericardium and expanded polytetrafluoroethylene (e-PTFE)

(Choo et al. 1999; Kao 2000). The use of bovine pericardium as a wrapping material appears to be more inflammatory than scleral wrapping on histology

(DeBacker et al. 1999) and clinically may have higher exposure rates, (Arat et al. 2003), although there is disagreement between authors (Gayre et al. 2001). Another alternative is polyglactin mesh (Jordan et al. 1995). A fibrovascular capsule of variable thickness forms external to the polyglactin mesh and replaces it by 12 weeks (Jordan et al. 2003b). Polyglactin meshwrapped hydroxyapatite implants must be placed deep into the orbit to prevent exposure risk (Jordan & Klapper 1999; Custer et al. 2003; Jordan et al. 2003b). More recently, the use of acellular dermis (AlloDerm – human cadaveric dermis) has been advocated as a wrapping material. Histological studies suggest that it permits vascularization of porous implants and does not incite significant inflammation. (Thakker et al. 2004). Although there is a theoretical risk, no case of disease transmission has been reported to date (Kadyan & Sandramouli 2008)

Polymer-coated hydroxyapatite implants became available in 2003. The coating consists of two different color-coded polymers. The anterior amber portion absorbs over 18 months and the posterior purple portion absorbs over 6 weeks. The idea is to avoid the need for a tissue wrap, with protection against anterior exposure, while promoting fibrovascular ingrowth posteriorly

(Shields et al. 2007). Overall, the use of implant wrapping material appears to be declining among surgeons in North America (Su & Yen 2004).

(d)

17.8Continued

(H+Y Comprehensive Technologies, Philadelphia, PA) has been reported to contain some calcium oxide impurities. When hydrated in tissues, calcium hydroxide may form, which is caustic (Jordan et al. 1999c). The Brazilian hydroxyapatite implant (available in Brazil only) (Jordan et al. 2000b), has higher weight and lower porosity when compared with the Bio-Eye® implant (Jordan et al. 2000b). The Molteno M-Sphere® (IOP Inc., Costa Mesa, CA) is an antigen-free cancellous bone implant. It is comparatively fragile and may not be able to support a peg as well as the Bio-Eye® and FCI3® implants (Jordan et al. 2000a).

In some ways the early porous implants represented a regression in design (to the simple buried implants) in that the hydroxyapatite sphere may not translate movement to the implant as well as the irregular anterior surface of buried integrated implants (Fig. 17.9). In order to overcome this, the placement of a peg to translate movement better was advocated (Perry 1991).

17.9 50 years of ‘progress’ – 1950 to 2000 (shown from left to right). In some respects, spherical porous implants represented a regression in design (to the simple buried implants) in that they may not translate movement to the implant as well as the irregular anterior surface of buried integrated implants. The irony is that after 50 years of development we returned to plastic balls with channels to permit vascular ingrowth!

Vascularization of an unwrapped hydroxyapatite implant takes approximately 6 months. However, when wrapped in sclera, the speed of vascularization is more variable. Thus it is recommended to wait at least 6 months before considering peg placement, and to confirm complete implant vascularization with gadolinium-enhanced magnetic resonance imaging (Dutton 1991; Hamilton et al. 1992; Spirnak et al. 1995; Klapper et al. 2003; Park et al. 2003). Technetium-99m bone scintillography may also be used to verify implant vascularization (Ferrone & Dutton 1992; Numerow et al. 1994; Leitha et al. 1995). Drilling 1 mm holes to the center of the implant at the time of surgery appears to accelerate implant vascularization (Perry 1991; Ferrone & Dutton 1992; Jordan et al. 1998).

The ocularist can assist the surgeon in drilling the implant by creating a conformer with a central opening to guide the drill. If the initial drilled tunnel is not perpendicular and central to the implant surface, re-drilling can be problematic, especially if the drill shafts need to be juxtaposed. Once adequate vascularization is confirmed, a hole (about 2–3 mm wide and 10 mm deep) is drilled into the implant and a temporary peg is placed. Follow- up with an ocularist is needed in 4–6 weeks, at which time the flat-head peg is removed (Perry 1991).